Ionizing radiation in medical diagnostics, material analysis and radiotherapy of cancer

AstroNuclPhysics ® Nuclear Physics - Astrophysics - Cosmology - Philosophy Physics and nuclear medicine

3. Applications of ionizing radiation
- nuclear and radiation methods -
3.1. Nuclear and radiation methods
3.2. X-ray diagnostics
3.3. Radiation measurement of mechanical properties of materials
3.4. Radiation analytical methods of materials
3.5. Radioisotope tracking methods
3.6. Radiotherapy
3.7. Technological use of radiation

3.1. Nuclear and radiation methods - general properties
In this chapter we will try to give a brief overview of radioisotope measurement methods and applications of ionizing radiation in various fields of science and technology, health care, industry, ecology, etc. Before we discuss specific radiation methods, we will mention some common characteristics of these methods.
Note + apology: The application methods of ionizing radiation are discussed here from a physical point of view, without details of technical solutions, rather than from the point of view of individual special fields of application; the exceptions are methods of X-ray diagnostics, radiotherapy and especially nuclear medicine (where reference is made to a detailed and complete explanation - Chapter 4 " Radioisotope scintigraphy "). I therefore ask for the leniency of experts for specific methods, when they do not find a some technical or medical aspects for practical use in their field; I also apologize for any inaccuracies and excessive simplifications in these aspects. I focus here mainly on the interpretation of the physical nature.
  Radioisotope and radiation methods have some important advantages :

In addition to higher technical and cost demands, a certain disadvantage of radiation methods may be the risk of harmful effects of ionizing radiation on materials and human health; however, this risk can be eliminated or minimized by ensuring appropriate radiation protection - see Chapter 5 "Biological effects of radiation - radiation protection".

Does the material glow or not after the application of radiation?
This is a frequently discussed issue, especially in the general public. It is argued that "During irradiation, a given object (including possibly the human body) absorbed radiant energy - and this energy should then be gradually radiated back ! ". In the vast majority of common applications of radiation, this seemingly logical argument is flawed. Photon radiation, X and gamma, passes through a substance at the speed of light (corpuscular radiation only a little slower) and from the moment it leaves the substance it no longer occurs in it. The radiation "does its job" in the substance and then immediately disappears. Only physico-chemical (or later biological) effects of radiation can persist :
¨ When irradiated with X, g, b radiation with energies less than about 10MeV, excitations and ionizations of the atoms of the irradiated substance occur, accompanied by secondary radiation and possibly chemical radiation effects. Thus, during exposure, the irradiated object emits secondary radiation, the intensity of which represents a fraction of a percentage of the intensity of the primary beam. After the end of the radiation flow, deexcitation and recombination of atoms occur almost immediately (within about 10-8 sec.) and the substance then does not radiate at all. This radiation behaves like light to a certain extent: when we stop the irradiation ("go out"), the radiation immediately disappears (it is "dark"). Thus, the patient does not shine after X-ray examination or after normal radiotherapeutic irradiation of gamma or X, does not shine objects after X-ray fluorescence analysis or defectoscopy, does not shine materials after radiation sterilization.
¨ A more complex situation can occur with irradiation with neutron radiation (even at low energies - slow neutrons) and in general with high-energy radiation, the quantum of which has an energy higher than about 10MeV. In this case, the radiation can cause nuclear reactions, in which radionuclides can be formed in originally non-radioactive materials. Such a substance may "glow" for some time after irradiation. Not because "accumulated energy" is emitted, but because nuclear activation has taken place in material. Thus, the samples glow after neutron activation analysis, targets irradiated in nuclear reactors and accelerators glow strongly, weakly and for a short time also patients after radiotherapy with radiation higher than 10MeV, more significantly after hadron radiotherapy (see §3.6 "Radiotherapy", part "Hadron radiotherapy"). And, of course, patients shine after the application of a radioactive substance to the body in nuclear medicine - not by radiating some absorbed energy, but by the lingering radioactivity accumulated in the organism. The intensity of this radiation decreases exponentially with the rate given by the half-life of the used radionuclide and the rate of its excretion from the body.

Types of radiation methods
For applications of ionizing radiation, both closed emitters are used - X-ray, closed radioisotope, particle accelerators, as well as open emitters - radioactive liquids, gases or aerosols. All applications of ionizing radiation can be divided into two basic groups :

1. Radiation measuring, analytical and detection methods
This large group of methods uses the properties of ionizing radiation to measure certain physical and technical quantities, to analyze the properties of substances and to study and detect certain processes in natural and industrial systems or in living organisms
(see also section "Introskopy" below).
  In terms of the nature of primary and secondary radiation, as well as the relative position of the radiation source, the analyzed object and the detector, these methods can be further divided into four groups :

Fig.3.1.1. Geometric arrangement of the radiation source, the analyzed or irradiated object and the detector in various applications of ionizing radiation.
a) Transmission measurements of radiation absorption. b) Scattering and fluorescence measurements. c) Emission radiation measurement. d) Measurement of radioactive samples. e) Radiation irradiation of objects.

Radiation measuring, analytical and detection methods belong to a wider field, sometimes called introscopy
(Latin intro = inside , Greek scopeo = observation ; literally "looking inward") - non-destructive examination of the internal structure of objects and the processes taking place in them, using physical methods: sound waves (including ultrasound), electromagnetic field and electromagnetic waves (light - eg classical endoscopy in medicine, radio waves, UV, X and g-radiation - nuclear medicine), fluxes of elementary particles (accelerated electrons, protons, neutrons, heavier ions). These methods are used mainly in medicine (from classical stethoscope, through optical endoscopy to ultrasound sonography, X-ray diagnostics and gammagraphy), but also in a number of scientific and technical and industrial applications (defectoscopy, activation analysis, X-ray fluorescence analysis and more). All of these methods, when using ionizing radiation or nuclear physics methods, will be described in more detail below.

2. Radiation irradiation and technological methods
Here, the energy transferred to the substance during irradiation is used - Fig.3.1.1e, ionization of substances and subsequent physical, chemical and biological effects of ionizing radiation in the irradiated object. In the field of medical applications, this includes radiotherapy, industrial applications include some radiation-technological processes in chemistry (such as polymerization), sterilization of materials, production of radionuclides.

  The following paragraphs (§3.2-§3.7) will describe individual specific methods of ionizing radiation application, some briefly (industrial applications), others in detail (X-ray diagnostics, radiotherapy; in a special reference to a separate chapter 4 "Radionuclide scintigraphy" dedicated to methods of nuclear medicine).

Collimation of ionizing radiation
In the vast majority of processes of ionizing radiation, this radiation is emitted almost isotropically in all directions *).
*) Exceptions are the interactions of high-energy particles, where due to the relativistic laws of conservation of momentum, the resulting particles and radiation are kinematically directed (collimated) in the direction of motion of primary high-energy particles.
  However, we often need to direct the radiation to a certain angle, or to concentrate it in a certain place; radiation in other directions can be undesirable - disruptive or even harmful and dangerous. This routing, or collimation of radiation, can be performed in two basic ways :
¨ Electromagnetic collimation of charged particles
In the case of corpuscular radiation of charged particles, suitable direction - collimation - can be achieved by the action of electric and magnetic fields, which exert a force on the charged particles. This deflects the direction of movement of the particles (beam), which can be directed to the desired location.

Mechanical absorption collimation of radiation
However, a simpler way, which works both for charged particles and for
g and X radiation, is to use collimators.  A collimator is a mechanical and geometric arrangement of materials absorbing a given type of radiation, that transmits only radiation from certain desired directions (angles), while absorbing and retaining radiation from other directions *).
*) However, such absolutely sharp collimations cannot always be achieved in practice. For the case of penetrating high-energy radiation gamma, partial cross-radiation trought the shielding occurs at the peripheral edges of the collimator, in which creates a kind of "half-shadow" ("penumbra") in the edge parts of the collimated beam.
  Collimators are used in virtually all applications of ionizing radiation. Most of them are simple collimators in the shape of various tubes or orifices (as shown in a simplified way, for example, in Fig.3.1.1). Intricately configured collimators then play a key role especially in scintigraphy
(imaging collimators with a large number of holes - §4.2 "Scintillation cameras", part "Collimators"), in X-ray diagnostics(§3.2 "X-ray diagnostics") and in radiotherapy (eg multi-lamellar MLC collimators - §3.6 "Radiotherapy", passage "Modulation of irradiation beams").
Electronic collimation of radiation

In some special detection and imaging systems, another method of directional radiation selection, so-called electronic collimation, is used without the use of a mechanical collimator. It is based on the specific behavior of quantum ionizing radiation in the detection system - the propagation of pairs (or more) of quantums in certain precisely given directions, their coincident detection by a system of a number of detectors and subsequent positional and angular reconstruction of the direction of quantum propagation. This analysis makes it possible to select for further processing only those quanta of radiation that have the desired direction - to perform electronic collimation and display the distribution of radiation in a given field. The electronic collimation method is used in positron emission tomography PET (see §4.3 "Tomographic cameras, part "
PET cameras") and in some complex detection systems such as ring imaging Cherenkov RICH detectors (see ....), trackers and muon detection systems for accelerators (see §2.1, section "Arrangement and configuration of radiation detectors").

Imaging using radiation - radiography
The very concept of imaging is based on the ability of our eyes to perceive light, its intensity, wavelength and spatial distribution, from which we create basic ideas about the shapes, size and placement of objects in space. If we want to get an objective idea of an object, its structure, changes and processes taking place in it, the most clear is to obtain the relevant data in pictorial form. This applies to an inanimate object, a living organism, the human body, or perhaps a distant galaxy in universe. This imaging is performed by visualizing the physical fields with which the examinated object interacts, or which brodcasts. That is, by means of various types of radiation, with which we irradiate the object, or which the object itself emits. The transmitted, reflected, scattered or emitted radiation is detected by suitable position-sensitive detectors, which display the spatial distribution of the radiation field (or its planar projections) and possibly also its other properties (especially the energy of quantum radiation)
- see §2.1 "Methodology of ionizing radiation detection".
  Radiography is the collective name for measuring quantity and displaying distribution radiation from studied objects that emit radiation either primarily, or secondarily when they are irradiated from external radiation sources. This imaging is performed using photochemical manifestations in photographic emulsions, fluorescence of luminescence of screens and especially physical processes in electronic imaging detectors. This includes a number of methods from the fields of X-ray diagnostics, radiation defectoscopy, gammagraphy (scintigraphy) using radiopharmaceuticals. Imaging methods using different types of radiation are discussed below.
   About the X-ray image in the following §3.2 "
X-rays - X-ray diagnostics" (including the appendix "X-ray telescopes"). Autoradiography - photographic imaging of the distribution of the beta-radioindicator in the examined preparations in close contact of the photographic emulsion with the sample is described in §2.2 "Photographic detection of ionizing radiation", passage "Autoradiography". Gamma-ray imaging is discussed in detail in Chapter 4 "Radionuclide Scintigraphy", especially for applications in nuclear medicine (however, there are also brief methods for g- imaging from space - gamma-telescopes, "High-energy gamma cameras").
  In addition to the visual observing and evaluating the thus obtained image is often also important mathematical analysis of the images, either static (filtering, comparing data from different locations of images or between various images) or dynamic (evaluation and quantification of temporal changes in different parts of the image reflecting the dynamics of the respective processes in displayed object); these aspects are discussed in detail for the field of scintigraphy in §4.7 "
Mathematical analysis and computer evaluation in nuclear medicine".

3.2. X-radiation , X-ray diagnostics
The oldest, most widespread and still probably the most important application of ionizing radiation is X-ray diagnostics
(rtg diagnostics, often also called radiodiagnostics, popularly called "x-raying"). From a physical point of view, we will here discuss the instrumentation and methods of X-ray diagnostics :

X-radiation and X-ray imaging X-ray sources - X-rays tubes, construction, power supply
Visualization of X-ray images, electronic detectors X-ray planar imaging, sciascopy, subtraction radiography
X-ray tomography - CT Bone densitometry , mammography , dental X-ray

Discovery of X-radiation
In the last decades of the 19th century, a number of researchers have experimented with high-voltage electric discharges in dilute gases. The so-called cathode rays were discovered , which were later found to be fast-moving electrons (see also §1.1, section "
Structure of atoms"). These experiments with discharges in the cathode ray tube were also performed in 1895 by W.C.Röntgen in a laboratory in Würtzburg. In the darkroom, he observed the fluorescence caused by cathode rays on luminescent screens. He tried to cover the cathode ray tube with black paper and found that the luminescent screen glowed as it approached even the tube thus covered; even when he inserted a thick book between the tube and the screen. Only when he placed a metal object between the tube and the screen, did a shadow appear on the screen. And as he accidentally placed his hand between the cathode ray tube and the screen, faint outlines of bones appeared on the screen. It was clear that unknown penetrating rays is emitted from the cathode ray tube - the "X- rays", as Roenrgen called them (letter X as a symbol for something unknown - an unknown variable in mathematics, an unknown person in a detective story). This radiation can penetrate trought paper and fleshy tissue, but metal objects and bones are "opaque" to this radiation. Furthermore, Roentgen found that this radiation caused the blackening of the photographic plate.

Discovery of X - radiation .
Left: Laboratoy of W.C.Röntgen in Würtzburg. Middle: Röntgen shows off its X-rays. Right: X-rays were independently of Röntgen at the same time discovered also by H.Jackson and A.A.Campbell-Swinton, but they did not deal with medical applications.

Immediately after his discovery of penetrating radiation emanating from the cathode ray tube in 1895, Roentgen himself took the historically first X-ray image on a photographic plate, namely his wife's hand (Fig.3.2.1 on the right, even with a ring). Both Roentgen and other physicians have been aware from the beginning of the great importance of newly discovered radiation for medicine. Roentgen thus became the first radiologist...
  Roentgen and other researchers initially thought, that penetrating radiation originated in the diluted gas of the cathode ray tube. In further experiments it was shown, that the source of X-rays is not the discharge in the gas itself; this ionization only supplies the electrons, that are accelerated and their impact on the anode excite the braking X-rays. The removal (exhaustion) of gas and the use of a hot cathode emitting of electrons will increase the efficiency of X-rays - vacuum X-ray tubes have developed over the course of several decades
(described in detail below).
Note: A brief reflection on the extent to which the discovery of X-rays was the result of chance or methodological procedure, is given in §1.0 "Physics - fundamental natural science", passage "Significant scientific discoveries - chance or method?".

Fig.3.2.1. The principle of X-ray diagnostics.
Left: Basic principal scheme of X-ray imaging. Middle: X-ray spectrum emitted from the the X-ray tube (filtered).
Right: The first X-ray image taken by Roentgen himself - his wife's hands even with the ring
(according to other sources, it was perhaps the hand of his friend Prof. of anatomy A.Koelliker..?..).

Origin and properties of the X-ray image
When using X-rays for imaging (especially in medicine), its basic properties of penetrating even materials opaque to light are used. The basic principal scheme of X-ray transmission imaging is in the left part of Fig.3.2.1. The penetrating electromagnetic X-rays with a photon energy of about 20
-150 keV (wavelengths of about 5 to 50 picometers), generated in the X-ray tube, pass through the examined object (organism tissue), while part of the radiation is absorbed depending on the thickness and density of the tissue, while the remainder portion passes through the tissue and is displayed either photographically or on a luminescent screen, more recently using electronic detectors. In the body, X-rays are most absorbed by bones, less by soft tissues, least by body cavities and by air. When exposed to X-rays, an X- ray image of the examined tissue is created, which is a projection shadow image of density, showing differences in density of tissues *). In other words, an X-ray image is created by projecting X-rays from the focus of the anode, through tissue structures within the organism with different absorption coefficients and different thicknesses, onto a film or imaging detector. Different absorptions of X-rays in different tissues are assigned different intensities in gray scale in the image; this assignment is realized either in an analog manner (film blackening) or digitally (electronic imaging detectors + computer, see below). This creates an image reflecting the size, shapes and arrangement of tissues and organs in the body, including possible changes induced by pathological processes.
*) Differentiated absorption of X-rays are the basis for formation of X-ray image. This absorption depends on the layer thickness, density and proton number of the irradiated substance. Soft tissues have a lower density and lower absorption of X-rays - more radiation is transmitted through these places, we get a clearer image or greater blackening of photographic material. The bones with calcium content are denser and absorb more X-rays - less passes through it, we get a less intense image or less blackening of the photographic film in these places. In Fig.3.2.1 on the right is an X-ray image on a photographic film.
   X-rays interact with tissue atoms mainly through two processes, discussed in more detail in §1.6, section "Interaction of gamma and X-rays": photoeffect and Compton scattering (formation of electron-positron pairs does not occur here due to the low energy of photons used in X-ray diagnostics; an insignificant exception may be portal and tomo-therapeutic images on radiotherapy irradiators, see §3.6 "Radiotherapy"). Both of these processes are involved in the different absorption of radiation in individual tissues (and also in the different absorption in normal and pathological districts within the same tissue), depending on the thickness, density of the substance and the proton number of the atoms. X-ray diagnostics is based on this different absorption of X-rays in different tissues, as well as differences of absorbtion in their physiological or pathological conditions.
Chemical (atomic-elemental) composition of tissues and organs?
Different tissues and organs differ in their chemical composition, which may or may not be reflected in their different densities. If two adjacent structures in the body have the same or close absorption coefficient
(linear attenuation coefficient) for the X-rays used, they will be indistinguishable from each other on X-ray images - they will appear identical, even if their material (chemical, elemental) composition is significantly different. Differentiation or classification of different tissue types by standard X-ray imaging is therefore very difficult and often impossible.
  A certain possibility of at least partial resolution of the material composition of the displayed structures is measurement - imaging - at different X-rays energies - X-rays spectrometry. We will deal with these possibilities below in the sections "
Electronic X-ray imaging detectors" - "Spectrometric Photon-counting X-ray imaging", "X-ray detectors for CT" and "CT with 2 X-rays - DSCT: Dual Source and Dual Energy CT".

X-ray image quality
Three parameters are important for high-quality X-ray imaging and recognition of fine structures and anomalies :
¨ Sharpness and resolution ability of imaging
For the projection image sharpness is important the small size of the impact focus, from which the X-radiation is emitted
(see below, "The design of the X-ray tube"). For classical X-ray diagnostics, the focus is about 0.5¸2 mm, but for X-ray microscopy, an almost point focus with a diameter of the order of micrometers is required. Closely related to sharpness is the spatial resolution of the image *). Sharpness and resolution can also be affected by the properties of the imaging medium - photographic film, amplifying films, electronic imaging detectors. The resolution of the X-ray image is around 0.5-2 mm, depending on the size of the focus (at X-ray microscopy can be a thousand times better!).
*) Resolution is defined as the smallest distance between two "point" objects, at which they still displayed as two separate structures; or equivalently as the half-width of the point object image profile. At shorter distances, both objects already appear as one, they are not distinguished. As in photography, resolution is often measured in the maximum number of lines per millimeter [lp/mm] that can still be distinguished; in practice, the real X-resolution is around 2-5 lp/mm. The quality of X-ray imaging in terms of real resolution is sometimes quantified in detail using the so-called modulation transfer function MTF, indicating using Fourier harmonic analysis, what details of the examined object can be displayed with the given contrast. The issue of resolution, contrast and recognizability of lesions on X-ray images is largely similar to scintigraphic imaging - it is discussed in detail in §4.2, section "Scintigraphic image quality and detectability of lesions".
   Significant deterioration in sharpness and resolution occurs especially when the image is blurred due to patient movement during exposure - motion blur. With modern devices, this risk is minimized by shortening the exposure time, thanks to the simultaneous increase the intensity of X-rays. Also, the movements of certain structures inside the body - heart beating or breathing movements - lead to image blur. This adverse effect can be eliminated by gating (trigration) and image synchronization in certain phases of cardiac pulsation or respiration - ECG-gating, respiratory-gating.
¨ The contrast of the imaging ,
which expresses the gradient of displaying differences in X-rays absorption using a gray scale, is given by two factors. First of all, it is the ratio of absorption coefficients for different types of displayed tissue. It depends mainly on the differences in the density of individual areas of tissue; where this difference is negligible or non-existent, we can sometimes increase it by applying contrast agents (see below). The contrast in absorption further depends on the energy of the X-rays. For thinner layers of soft tissue, soft X-rays (approx. 20 keV) are more suitable, which interact mainly with a photoeffect with a steeper difference in absorption depending on density
(the greatest contrast is achieved for X-rays close to the binding energy of electrons on K or L shells). Harder X-radiation (approx. 80-100 keV) is required to display thicker layers and denser materials (eg skeletal structure). Contrast in image is negatively affected by Compton scattered radiation (see "secondary diaphragms" below).
   An important geometrical-anatomical factor, significantly worsening the contrast of the X-ray image and the overall recognizability of the lesions, is the cross-radiation and superposition of X-rays from individual layers of tissues and organs at different depths, generally with different densities. This adverse effect is largely eliminated in CT imaging.
   For digital devices, the contrast can be additionally increased by computer processing ( post-processing ) - a suitable brightness modulation of the image. In such processing, the so-called bit depth is important - the number of bits in which the image is created in the process of analog-digital conversion.(ADC) from an electronic X-ray detector to an image matrix in a computer. When displayed, the bit depth indicates the maximum number of shades of gray that we are able to display in the image - the larger this number of shades of gray, the more depicted we show particularly small differences in density and fine detail. A higher number of bits in the image allows you to emphasize the details in the image using suitable display windows for brightness modulation - stretching a certain small range of brightness values in the image to the full range.
  The relationship between the most commonly used bit depth b and the maximum number of shades of gray is as follows (given by the power of 2 b ) :
2 bits = 2 shades (white and black only); 4 bits = 16 shades; 8 bits = 256 shades; 12 bits = 4096 shades; 14 bits = 16384 shades; 16 bits = 65536 shades; 24 bits = 16777216 shades.
Although a large number of shades (tens and hundreds of thousands) are no longer directly distinguishable by the eye, this allows by the use of narrow display windows to emphasize density gradients.

¨ Number of photons in the image - statistical noise
To obtain a quality well-exposed image, a certain optimal number of X-ray photons is needed. In films and luminescent screens, this number of photons is mainly determined by sensitivity the material used, so that the image is not underexposed or overexposed. With digital imaging detectors, we can additionally adjust the brightness of the image, but the image quality is still determined by the following factor
: X-ray emission, its interaction and imaging detection is subject to stochastic quantum laws, leading to quantum statistical fluctuations in photon flux. With insufficient X-ray photons, the image is "noisy", composed of disturbing artificial brighter and darker spots and clusters of dots, where fine structures and details can disappear. If we have the registered number of N photons of X-rays in a given element (pixel) of the image, then the local statistical fluctuations - scattering, relative error - are SD = 1/ÖN. To obtain a well-drawn image with statistical fluctuations of less than 3%, more than 1000 photons must be recorded in each element of the image, for 1% of the fluctuation there must be more than 10,000 pulses/element.
   For digital imaging detectors - flat panels
(described below) - the quality of the X-ray image in terms of noise depends on the sensitivity of the sensor: this is given by the detection quantum efficiency DQE (Detection Quantum Efficiency), which is the percentage of photons X-rays incident on the detector, that are actually recorded by the detector and used to create the image (the rest is uselessly absorbed by the input window or detector material without scintillation or electrical response). For digital X-ray images, especially CT, the statistical noise of the image is expressed in Hounsfield units HU (introduced below in the section "X-ray tomography -CT", passage "Origin of the density image"); in a good picture, the SD noise should not exceed about 20-30 HU. The total number of photons for the exposure of a given image is set by the product of the X-ray current and the exposure time (see below "X-raytube", section "Braking X-rays") - "milliampere-seconds" [mAs]; it can also be electronically controlled using automatic exposure - see below "Setting X-ray parameters".
¨ Artifacts on an X-ray image
Under certain circumstances, some structures that do nothave their origin in the displayed object, may appear on X-ray images - they are false artifacts . They can be caused by inhomogeneities, defects or impurities on the photographic film or reforcing foils, inhomogeneities in the flat-panel detectors, unwanted objects (eg metal) in the X-ray beam. In CT imaging, so-called structural artifacts may occurs, arising during the reconstruction of transverse sections in places with sharp differences in density, especially at the transition of bone and soft tissue.

X-ray tube
Source of X-rays for X-imaging is a special vacuum tube called X-ray tube or X-ray lamp. From an electronic point of view, the X-ray tube is simply a classic diode connected in a circuit with a high voltage of approx. 20-200 kV. So it has two metal electrodes :

--> Cathode
formed by a thin metal wire wound into a narrow spiral. A metal that is very resistant to temperature is suitable for the heated filament of the cathode - it has a high melting point, is strong and mechanically stable, and has a low output work of electron emission. Tungsten is most often used, which has a high melting point of 3300 °C. It is basically similar to the tungsten filaments of classical light bulbs
(but where it is the emission of light, not electrons).
   An electric current
(several Amperes) is applied to this metal wire, which heats up the fiber to a temperature of approx. 2000 °C. Thermoemission then releases electrons from the metal - the heated cathode emits electrons. The release of electrons occurs when, during their thermal movement, the electrons acquire a kinetic energy higher than the output work of the electrons from the given metal. As the temperature of the metal increases, the electron thermoemission density increases significantly. The dependence of the thermoemission intensity J on the metal temperature T is described by Richardson's formula :
J = F . T2. e-w/(k.T) ,
where J is the current density of emitted electrons [A/cm2] - current per unit area of the emitting surface of the metal, T is the absolute temperature of the metal [°K], w is the output work of electrons [eV], k=8.62.10-5 eV/° K is Boltzman's constant. Electron emission has the character of a quantum tunneling effect (§1.1, passage "Tunneling effect").
F is a material-dependent constant [A/(cm2.°K2)], for tungsten it has a value of F~ 60 A/(cm2.°K2). The output work of electrons from tungsten is w=4.5 eV and it starts emitting electrons when heated to a temperature higher than 2000 °C, but effective emission only occurs at temperatures of 2300-2500 °C.
For the X-ray cathode, instead of pure tungsten, a thorium-coated tungsten filament is sometimes used, which has a lower electron work function of only 2.6 eV. The cathode from this thoriated tungsten effectively emits electrons already at a temperature of 1700-1900 °C. The lower operating temperature extends the life of the cathode by about three times.A metal that is very resistant to temperature is suitable for the heated filament of the cathode - it has a high melting point, is strong and mechanically stable, and has a low output work of electron emission. Tungsten is most often used, which has a high melting point of 3300 °C. It is basically similar to the tungsten filaments of classical light bulbs (but where it is the emission of light, not electrons).
An electric current (several Amperes) is applied to this metal wire, which heats up the fiber to a temperature of approx. 2000 °C. Thermoemission then releases electrons from the metal - the heated cathode emits electrons. The release of electrons occurs when, during their thermal movement, the electrons acquire a kinetic energy higher than the output work of the electrons from the given metal. As the temperature of the metal increases, the electron thermoemission density increases significantly. The dependence of the thermoemission intensity J on the metal temperature T is described by Richardson's formula:
            J   =   F . T2 .
e-w/(k.T)    ,
 where J is the current density of emitted electrons [A/cm
2] - current per unit area of the emitting surface of the metal, T is the absolute temperature of the metal [°K], w is the output work of electrons [eV], k=8.62.10-5 eV/°K is Boltzman's constant. Electron emission has the character of a quantum tunneling effect (§1.1, passage "Tuneling effect").
   F is a material-dependent constant [A/(cm
2.°K2)], for tungsten it has a value of F~ 60 A/(cm2.°K2). The output work of electrons from tungsten is w=4.5 eV and it starts emitting electrons when heated to a temperature higher than 2000 °C, but effective emission only occurs at temperatures of 2300-2500 °C.
For the X-ray cathode, instead of pure tungsten, a thorium-coated tungsten filament is sometimes used, which has a lower electron work function of only 2.6 eV. A small amount of thorium, mixed into tungsten, in a wire heated to about 2500°C, drifts to the surface layer, where it causes more efficient thermoemission of electrons. The cathode from this thoriated tungsten effectively emits electrons already at a temperature of 1700-1900 °C. This lower operating temperature extends the life of the cathode by about three times.
   If there were no positive voltage on the anode, these emitted electrons would form an electron cloud around the cathode and their repulsive force would prevent further thermoemission (this is the case around the filament of a light bulb). However, at a sufficiently high positive voltage (>60kV) at the anode, thermoemission electrons are continuously diverted away from the cathode and rapidly move towards the anode, an electron cloud is not formed. However, if the anode voltage is relatively low (<40kV), part of the emitted electrons will no longer reach the anode and a larger or smaller electron cloud remains around the cathode, preventing stronger thermoemission of electrons. Stronger heating of the cathode no longer leads to higher thermoemission and to a greater electron current through the X-ray.
 Cathode in the shape of a flat emitter
Some new X-ray tubes instead of the classic spiral fiber have a heated cathode using the so-called flat emitter technology. It consists of a rectangle of heated thin sheet, masked by several holes. By setting a negative voltage between the cathode and emitter slits, a very sharply localized incident focus on the anode can be more accurately achieved.
--> Anode
Electrons emitted from the cathode are attracted to the anode *) with a high positive voltage, while they are accelerated by a strong electric field to the kinetic energy E = U.e, given by the high voltage U between the cathode and the anode (ie E = approx. 20
¸200 keV). Just before the impact on the anode, it obtains an electron with charge e and mass me a very high velocity v = Ö(2.e.U/me) (for voltage U = 60kV, the electrons will have a kinetic energy of 60keV and an impact velocity of approximately 150000 km/s, which is half the speed of light). Upon impact with the anode, the electrons brake rapidly, converting some of their kinetic energy to hard electromagnetic radiation - X-rays of two types: bracking and characteristic radiation (the origin and properties of these two types of radiation are discussed below). This X-ray leaves the anode and flies out of the tube (Fig.3.2.1 left).
*) The anode, the electrode located opposite the cathode, was formerly also called anticathode, especially in the cathode ray tubes.
   The anode is made of a heavy material (most often tungsten), which has a high electron density, so the incident electrons are sharply braked by a large repulsive force, which, according to the laws of electrodynamics, turns part of their kinetic energy into braking electromagnetic radiation - X-ray photons. However, the efficiency of this process is relatively small - only about 1% of the total kinetic energy of electrons is transformed into X-ray photons, the rest is converted into heat. The reason is that only about 1% of electrons penetrate deep enough inside the atoms of the anode material, up to the L or K shell, only where large Coulomb electric forces act, causing a sharp change in the speed of the electrons and thus effective excitation of hard braking ratiation. The other electrons transfer their kinetic energy to the electrons and atoms of the crystal lattice, which results in heat.
The X-ray tube can be considered the simplest particle accelerator (§1.5 "Elementary particles", part "
Charged particle accelerators") - it is a linear electrostatic accelerator of electrons, the source of which is a hot cathode, the (inner) target is the anode, the braking (+characteristic) X-rays comes out.
 X-ray tube volt-ampere characteristic
For the electronic operation of the x-ray tube, it is important how the electron current [mA] by the X-ray tube depends on the anode voltage [kV] - the volt-ampere characteristic - and also on the incandescent current [A] of the cathode. When the cathode is heated (e.g. with a current of approx. 5A, to a temperature of approx. 1500 °C), as the anode voltage increases, the electron current through the cathode gradually increases (still more electrons from the cloud around the cathode reach the anode) and then reaches saturation - all electrons released by thermoemission fall on anode and there are no more free electrons that could fly to the anode at a higher anode voltage.
   The cathode current characteristic of the X-ray tube is also important - the dependence of the resulting electron anode current on the cathode glow current. This characteristic is different at different anode voltages. Generally, as the glow current increases, the anode current also increases initially, but only up to a certain value. At a low anode voltage (<40kV), saturation occurs - increasing the glow current no longer leads to an increase in the anode current: the electric potential is not sufficient for all thermo-emitted electrons to fly to the anode, an electron cloud is formed around the cathode. Only at a high anode voltage (>60kV) do all the electrons released by thermoemission fall on the anode and the effect of the electron cloud and saturation does not occur.
--> 3rd electrode - grid ?
In addition to the cathode and anode, in some types of X-ray tubes, we can rarely find a third electrode - a wire grid, located between the cathode and the anode, in close proximity to the cathode. The electrical voltage applied to this grid very sensitively modulates the flow of electrons (i.e. anode current) and thus also the intensity of X-radiation. Applying a higher negative voltage to the grid can very quickly interrupt the anode current and thus the emission of X-rays (sometimes used for fast x-ray cinematography).

Braking X-rays
Braking radiation is a consequence of the laws of Maxwell's electrodynamics, according to which every uneven ("accelerated" or "decelerated") movement of an electric charge emits electromagnetic waves -
see §1.5 "Electromagnetic field. Maxwell's equations.", Larmor's formula (1.61 '), monograph "Gravity, black holes and space-time physics". Therefore, even when the electron is braked after hitting the anode, the sharper the braking (the greater the deceleration a in the mentioned formula), the more intense and harder the electromagnetic radiation is generated - see also §1.6, passage "Braking radiation".
  The effective cross section for the production of braking radiation is generally given by the highly complicated Bethe-Heitler formula (derived from quantum radiation theory, corrected by the Sauter and Elwert factors of the Coulomb shielding of the electron shell). For a not very wide range of kinetic energies Ee of incident electrons (tens to hundreds of keV) and proton numbers Z of target material (medium to heavy materials), the overall efficiency of brake radiation production h can be approximated by a simplified formula :
h = E e [keV] . Z . 10 -6   [photons / electron] .
By converting the number of electrons n
e to the current I = ne .qe/t and by substituting the value of the charge of the electron qe = 1.6.10-19 C (= 1.6.10-16 mAs) from this relation the resulting flux of photons IX [number of photons /s.] braking radiation depending on X-ray tube current I [mA] and anode voltage U [kV] :
X = U. I. (Z /1.6). 10 10   [photons / s.] ,
which will be used below in the section "
Setting X-ray parameters". Only a relatively small part (only about 1%) of the original kinetic energy of the incident particle changes to braking radiation when braked in the matter. Most of the energy, with multiple Coulomb scattering, is eventually transferred to the kinetic energy of the atoms of the anode substance - it is converted into heat.
  Total energy spectrum of X-rays
(braking + characteristic), emitted from the anode of the X-ray tube, is drawn below in Fig.3.2.5 at the top right. The graphic form of the energy spectrum I(E) of continuous braking X-rays is approximated by the so-called Kramers formula :
                           I (E) = K. I. Z . (Emax - E) ,
where I(E) is the relative intensity of energy photons E , K is a constant, Z is the proton (atomic) number of the anode material, E
max is the maximum energy of X-ray photons, given by the kinetic energy of incident electrons. It is clear that I(Emax) = 0 and the formula is valid only for E < Emax .
  It is logical that the efficiency of brake radiation production is higher for high Z - large electric Coulomb forces act around such nuclei, causing abrupt changes in the velocity vector of the incident electrons that get close to the nucleus. The efficiency of braking radiation [number of photons /electron] increases with energy Ee incident electrons. Low-energy electrons are usually scattered on the outer shells of the atoms of the anode material and emit soft radiation, which often does not even reach the X-ray energy range. The higher the energy of the incident electrons, the more likely they are to penetrate deeper into the anode atoms, close to the nucleus, where the greatest electrical forces act, significantly changing the electron velocity vector, leading to higher energy and efficiency of braking X-ray production. However, the overall energy efficiency - the ratio of the total energy of the emitted photons to the energy of the incident electrons - is lower for higher energies (due to the higher percentage of low-energy photons). And the heat losses in the target (anode) are higher.
   The braking X-rays produced by the X-ray tube have a continuous spectrum from energies close to zero to the maximum energy, given almost by the value of the anode voltage - Fig.3.2.1 in the middle (here is the spectrum after partial filtering of the softest component - see below). The energy of the braking radiation depends on the speed (acceleration) at which the electrons are braked on impact with the anode surface. The individual electrons penetrate at different depths into the atoms of the anode material, thus emitting different wavelengths or energies of photons. Those electrons, which "softly" brake with repeated multiple scattering on the outer electron shells of the anode atoms, emit a series of photons of low-energy braking (and characteristic) radiation; some of them fall into the area of soft X-rays, others into the area of UV and visible light (this resulting low-energy photons are often absorbed in the anode material and do not fly out). The deeper the electrons penetrate into the interior of the anode atoms, the closer to the nucleus, the faster the intense Coulomb forces change their velocity vector and the harder the braking X-rays are produced. The shortest wavelengths arise for electrons that have penetrated to the level of the K shell and closer to the nucleus, where they can be braked almost on-time. Depending on the impact factor of individual electrons relative to the anode atoms, which is random, all possibilities are continuously realized - such a different degree of electron braking causes a mixture of radiation of different wavelengths or photon energies - the result is a continuous spectrum of braking radiation. Low-energy X-ray photons are the most represented in this continuous spectrum, only a very small percentage at the end of the spectrum corresponds to high energies, close to the energy of incident electrons, given the high voltage between the cathode and the anode of the X-ray tube (see Fig.3.2.5 below).
The wavelength and energy of X-rays
By its nature, X-radiation are electromagnetic waves of short wavelength of about 10
-9-10-11 m, which, however, are emmited as quantum - photons - with an energy of about 5keV-200keV ( "The particle-wave duality"). Earlier (until the 1960s) it was customary to characterize X-rays with a wavelength of l and in the older literature was given the so-called Duan-Hunt relation lmin [nm] = h.c/e.U @ 1.234/U [kV] between the voltage U [in kilovolts] at X-ray tube and the minimum wavelength lmin [in nanometers] of the resulting braking X-rays *). A Kramer's formula for the spectrum was given in the form I(l) = K.Z.I.[(l/lmin) - 1]/l2 (in this form it was compiled by H.A.Kramers in 1923; at that time X-rays were described only by wavelength).
   This manner was very disadvantageous and misleading, especially in relation to the creation mechanism of this radiation in X-ray tubes, where the values of the accelerating voltage in [kV] occur. Now is abandoned long ago, the X-ray spectrum is expressed fundamentally by the photon energy E
X [keV], which in X-ray tube is derived directly from the voltage U (maximum energy EXmax @ U, mean energy <EX > » U/3] and the Duan-Hunt relation has lost its importance.
*) The Duan-Hunt formula actually just a differently written relation EX = h /l between the energy of the photon EX in [keV] and the wavelength l in [nm] for the situation, when all the kinetic energy E = U.e of the electron of charge e , accelerated by the voltage U, is converted into a photon X-rays (corresponds to the energy EXmax and the wavelength lmin ).
Characteristic X-rays
In addition to braking X-rays with a continuous spectrum, a certain smaller amount of characteristic X-rays with a line spectrum (characteristic pair of peaks K
a, Kb) is emitted, the energy of which does not depend on the anode voltage, but is given by the anode material; for the most commonly used tungsten, these are the 59.3+67.2 keV peaks (and also the L peak around 10keV), which appear as "bumps" on the continuous curve of the spectrum (Fig.3.2.1 in the middle).
   The characteristic X-rays are caused by two processes :
¨ Direct process of the impact photoeffect at the internal energy levels of the electron shell in the atoms of the anode material - fast electrons penetrate into the atoms and eject bound electrons from the K and L shells. When electrons jump from the L shell to the emptied shell K (K-series), or from the shell M to L (L-series), the difference of energies is then radiated in the form of photons of electromagnetic radiation - characteristic X-radiation (cf. also with Fig.1.1.3 in §1.1).
¨ Indirect process of photoelectric absorption of braking radiation - braking X-rays, generated by the above-mentioned mechanism during the braking of accelerated electrons, interact with other atoms inside the anode substance, among others by a photon photoeffect (described in §1.6, part " Interaction of gamma and X-rays ", Fig.1.6.3 left), emitting electrons from the inner shells, followed by an electron jump and the emission of characteristic X-rays, similar to the previous case.
   The impact electron photoeffect and the emission of photons also occur when electrons jump in the outer shells, but the energy of these photons is low and this radiation is covered by continuous braking radion at the beginning of the spectrum.

   A certain minimum (threshold) anode voltage is required for the formation of characteristic X-rays, higher than the binding energy of electrons on the K-shell of atoms of the anode material
(for tungsten it is about 70keV, for molybdenum 20keV). If the anode voltage is lower, only continuous braking radiation is generated in the X-ray, and when the threshold voltage is exceeded, the spectrum contains both continuous braking and peaks of characteristic X-rays.
  The proportion of characteristic X-rays in the total spectrum of the X-ray tube depends on the anode material and the anode voltage. For a tungsten anode, it is approximately 30% at a voltage of 100 kV and only about 3% at a voltage of 200 kV.

X-ray tube design
Unlike conventional electron tubes used in low-current electronics, X-rays tubes have a relatively robust design (they resemble screens or transmitter electron tubes in size), given by two circumstances. On the one hand, it is a very high voltage, reaching hundreds of kilovolts. The second circumstance is thermal heating: electrons incident at high speed on the anode convert only a small part of their energy into X-rays, the vast majority of their kinetic energy is converted into heat - the anode of the X-ray tube is heated strongly. To dissipate this heat, the anode must have a relatively massive construction; in addition, anode rotation or cooling is used (described below). One of the technical parameters is maximum power of the X-ray tube [kW] - peak electrical power input of the X-ray tube, which the X-ray tube can still "withstand" without overheating and thermally damaging.
  The most commonly used material for the anode of an X-ray tube is tungsten, a heavy and heat-resistant metal. To improve the thermal properties of the anode, especially the heat capacity, rhenium-alloyed tungsten (10%) is often used, or the anode is composed of several layers - alloyed tungsten, molybdenum, graphite. For X-ray tubes for X-rays around 20keV for mammography, the anode is made of molybdenum.
  X-rays tubes can be divided into two main groups, which govern their design (+ the third group of special constructions listed below) :
¨ X-rays tubes for industrial irradiation and radiotherapeutic use ,
which do not require focusing of electrons to an almost point focus and which have a fixed (non-rotating) anode. High energy and X-ray intensity are common requirements here; the anode is actively cooled by the flow of cooling medium through its interior.
¨ X-rays tubes for X-ray diagnostics
with focusing of the electron beam into the focus and mostly with a rotating anode
(to prevent local overheating of the focus). Below we will deal mainly with these X-rays tubes for radiodiagnostics. The anode target material is mostly tungsten, for low X-ray energies (around 20-40keV) molybdenum is used as the anode target material; the X - ray tube is additionally equipped with a beryllium exit window - see below "X-ray mammography".
¨ Special types of X-ray tubes  
Microfocus X-ray tube have an extremely small impact focus of electrons on the anode, of the order of micrometers. This is achieved by placing a special set of electrodes (electron optics - "objective") between the hot cathode and the anode, focusing electrons from the cathode into a very narrow beam incident almost point on the target-anode. They provide very high sharpness and resolution of the image, but only limited power (intensity, fluence) of X-rays. They are used for X-ray microscopy and CT defectoscopy (see below §3.3, section "
Radiation defectoscopy").
  For special purposes (especially spectrometric and micro -X-rays), the X- ray tubes with a frontal transmission anode (Target Transmission X-ray Tube ) are constructed, where the beam of accelerated electrons impinges on the thin front-located anode, the resulting X-rays passing through the material of the thin anode to the outside of the tube where it is used. It can also be designed as the above-mentioned microfocus.
  In addition to the usual tightly closed (sealed) evacuated X-ray tubes, so-called open X-ray lamps are sometimes constructed. They have a metal casing that the user can open, replace the cathode filament and anode material (tungsten, copper, molybdenum, etc.) as needed, and close the tube again and evacuate.

Fig.3.2.2. Special microfocus X-ray tube with transmission anode for X-ray microscopy

Historical development of X-ray tubes
X-ray tubes originally evolved from discharge lamps, which are gas-filled glass tubes with electrodes to which a voltage of the order of hundreds of volts is applied. The next stage was Crookes cathode ray tubes - discharge lamps with very dilute gas, on the electrodes of which a high voltage of the unit of up to tens of kilovolts is applied. The classic radiant discharge practically no longer occurs here, but the ionization of the atoms of the diluted gas releases electrons, accelerated by a high voltage towards the anode - the cathode radiation originated. In addition to the fluorescence of the flask or inserted objects, there is also a secondary penetrating photon radiation - X-rays
(discovered by Roentgen and independently by other researchers), braking and characteristic. Cathode ray tubes also played an important role in atomic physics, with their help J.J.Thomson discovered electrons, which allowed them to penetrate the structure of atoms.
  The first "cold cathode" X-rays tubes were actually Crookes cathode ray tubes with specially modified electrodes. An important milestone was the creation of a vacuum X-ray tube with a hot cathode, constructed by W.D.Coolidge in 1913 (shown in Fig.3.2.1 on the left). Later, with increasing performance, the anode rotation as well as other technical improvements and special designs were added to X-ray imaging diagnostics (see below). Experiments with X-ray laser sources are currently being performed- whether the excitation of characteristic X-rays in a high-temperature plasma generated by a laser beam or braking X-rays on the impact of accelerated electrons on a target. On the sidelines, we can note that the most complicated special "X-ray tube" (sources of X-rays) can be considered wigglers and undulators of electron synchrotrons (see §1.5, section "
Charged particle accelerators").

Electron focusing, focal point
In order to achieve good sharpness and resolution of the projection shadow transmission image in X-ray diagnostics, it is necessary that the X-ray beam comes from an almost point source. In X-ray tubes for X-ray diagnostics, the red-hot filament - a tungsten spiral - is embedded in a recess or focusing slit of the cathode, which has a negative polarity, so that its repellent effect clusters electrons into a narrow strip *). After acceleration by high voltage, the electrons then fall into a relatively sharply localized place of the anode - the impact focus, which has a rectangular shape due to the elongated shape of the filament. Real, optical focus the resulting X-ray is a geometric projection of this radiating surface on the anode, i.e. the impact focus, into a plane perpendicular to the beam of radiation used for imaging. The originally rectangular impact focus is reduced in the longitudinal direction due to the inclined, tilted surface of the anode; its projection in the direction of the display has an almost square shape, usually 0.5-2 mm in size.
*) These X-rays tubes usually have two cathode fibers - shorter and longer. By switching the heating current, one or the other fiber can be heated and thus the size of the impact focus on the anode can be changed.
  Some new X-rays tubes, instead of the classic incandescent fiber, have an incandescent cathode solved by the so-called flat emitter technology. It consists of a rectangle of hot thin sheet metal, masked by several holes. By adjusting the negative voltage between the cathode gap and the emitter, a very sharply localized impact focus can be achieved more precisely.
Asymmetry of the X-ray beam from the focus, heel effect
In the first approximation, the X-ray is emitted from the impact focus isotropically, with the same intensity in all directions. However, some of the incident electrons penetrate below the surface of the anode and the X-rays generated there are partially absorbed and attenuated as they pass through the anode material. This leads to a change in the shape of the radiation pattern from the target at the anode, to a certain angular asymmetry of X-ray beam emanating from the chamfered anode: for an angle of about 30° in the direction of the cathode, the radiation intensity is about 5% higher than in the center (0°), in the opposite direction (to the anode disk) about 15% lower. This shaping of the radiation characteristics are sometimes referred to as anode heel effect, some "heel shape", "skew". This phenomenon may manifest itself in some minor inhomogeneity of the X-ray image, especially during exposures of large imaging fields, or during X-ray mammography. This inhomogeneity is smooth and gradual, so it does not interfere with visual evaluation; however, digital evaluation is sometimes computer-corrected.
 Anode cooling and rotation
As mentioned above, the vast majority
(almost 99%) of the kinetic energy of the electrons hitting the anode is converted into heat. This released heat must be effectively dissipated to prevent overheating of the anode. At low powers, passive infrared radiation from the heated anode to the surroundings is sufficient. X-ray tubes for high performance (without focus, e.g. in industrial use) have an actively cooled anode - inside the anode there is a cavity through which the cooling liquid flows.
  In diagnostic X-ray tubes, electrons fall into a small, sharply localized spot on the anode - the impact focus - about 1 mm in size. At higher powers, this impact focus on the anode can heat up strongly locally. It is necessary to ensure that the temperature of the focus is lower than the melting point of the anode material
(usually tungsten). Local overheating of the focus, where the electrons fall, can be prevented by rotating the anode *): the cathode is eccentrically placed in the X-ray tube, the anode in the shape of a conical disk (about 5-10 cm in diameter) rotates around the longitudinal axis, so that the electron beam always falls to a different place the circumference of the anode, making the heating and heat dissipation more uniform (Fig.3.2.3 left). Although X-rays emanate from the same place - the focus, which is against the stationary cathode, this place is due to the rotation of the anode constantly formed by another physical part of the anode disk; the heat is thus better dissipated in the anode material.
Anode rotation
Because the anode is located inside a high vacuum tube, its rotation cannot be ensured by a mechanical transmission from the outside (via the shaft). No bearing is so tight that no air enters the tube over time - the vacuum would be broken. Rotation of the anode is driven electromagnetically: inside the anode neck of X-ray tube is mounted on the bearings a metal cylinder connected by a shaft on the anode - serves as a rotor. From the outside the X-ray tube are disposed coils supplied with alternating current - those forming the stator, giving the rotating magnetic field which, by electromagnetic induction (eddy currents are induced in the rotor), rotates by a roller and an anode inside the tube (Fig.3.2.3 left). From an electromechanical point of view, such an X-ray tube is actually a small asynchronous electric motor. The rotation speed of the anode is usually 50Hz (3000 rpm), also 10-12,000 rpm is used for high power X-ray tubes. A certain problem is the wear of the bearings on which the anode rotor is anchored. These bearings are highly mechanically and thermally stressed, they are inside the vacuum space out of the possibility of maintenance and lubrication (only "dry" lubrication with silver or lead metal powder is used) - their wear is usually the main limiting factor of X-ray tube life.
  In some X-ray tubes, hydrodynamic lubrication of bearings with a thin layer of suitable molten metal is sometimes used (a kind of "aqua-planing" of the shaft in liquid metal, with minimal friction). One suitable metal is gallium, which has a low melting point of about 130 °C and a sufficiently high boiling point of 2204 °C, so that even at relatively high temperatures of several hundred °C, the vacuum does not contaminate its vapor. Furthermore, such a lubricating contact surface in the bearing efficiently dissipates heat from the anode. Before the actual operation, after switching on the device, the bearing is first heated and only after the melting of the lubricating metal does the rotation of the anode begin, which is then maintained continuously even outside the exposure, until the device is switched off. The bearing is heated and the required temperature is maintained by the effect of eddy currents induced in the rotor (these are the same eddy currents which, by their interaction with the rotating magnetic field of the stator, drive the rotation of the anode). Special so-called eutectic alloys are also used to lubricate the anode bearing metals which are liquid even at normal temperatures (eg gallium, indium and tin in an alloy of suitable ratio, with a melting point of -10 °C).
From a mechanical point of view, the rapidly rotating massive anode behaves like a flywheel, preserving its vector of rotational momentum. If we try to tilt the X-ray tube with a rotating anode (change the direction of its axis), due to the gyroscopic effect, the rotating anode puts up resistance and its bearings are stressed by considerable forces. This is especially the case with CT tomography devices, where the X-ray tube orbits around the examined object relatively quickly. Therefore, X-ray tude with double-sided anchoring of the anode axis are sometimes used here. The shaft of the rotating anode, passing through the whole X-ray tube, is mounted in bearings at both ends. The cathode portion of the X-ray tube then has two protrusions: one on the side for mounting and feeding the eccentrically located cathode, the other in the middle for mounting the second anode bearing. When the rotary anode shaft is anchored on both sides, the gyroscopic forces are distributed and the bearings are significantly less stressed.

Fig.3.2.3. Design of X-ray tubes used in radiodiagnostics.
Left: Classic X-ray lamp with rotating anode. Right: X-ray tube rotating as a whole (STRATON type), with the front anode in direct contact with the oil cooling bath and with the magnetic deflection of the electrons from the cathode.

Although the rotation of the anode prevents local overheating of the impact focus on the anode, during longer operation the anode heats up strongly as a whole and this heat is only slowly transferred by infrared radiation through a vacuum out of the X-ray lamp to the cooling medium. It is therefore necessary to observe certain time delays between individual exposures in order for the anode to cool down. Another disadvantage of the rotating anode is the wear of the bearing inside the vacuum flask, which cannot be lubricated or otherwise maintained from the outside. In addition, when the bearing wears, unwanted fumes are released into the vacuum space of the X-ray tube.
 X-ray tubes rotating as a whole
For higher performance, a new construction arrangement of the X-ray tube rotating as a whole, with direct cooling of the anode, was therefore developed. A beam of accelerated electrons from an axially positioned cathode, deflected by the magnetic field of the deflection coils (located outside the tube) *) impinges peripherally on the opposite front anode, which is in direct contact with the cooling oil bath from which the X-ray tube is immersed - Fig.3.2.3 on the right. The resulting heat from the impact focus is thus immediately dissipated away. The X-ray tube rotates as a whole around its longitudinal axis connecting the cathode to the center of the anode, the X-rays emanating in the lateral direction (similar to a conventional Coolidge-type tube). The heating and anode voltage is conducted to the X-ray tube by means of collecting rings, on which electric brushes slide
(slip-ring technology, similar to that of electric motors for direct current). The main advantage of this design is the substantially better cooling of the anode, which is in direct contact with the cooling medium, while there are no mechanically moving parts inside the vacuum space. The bearings on which the entire X-ray tube is mounted are easily accessible and can be effectively lubricated. This leads to the possibility of achieving higher performance and significantly extending the life of the X-ray tube.
*) The current through the deflection coils must be precisely set depending on the accelerating anode voltage: the higher the voltage [kV] at X-ray tube is set, the higher the current must flow through the deflection coils so that the electron beam is properly bent and hits the desired location at the anode edge. By electronic control of the current in the deflection coils, it is thus possible to set the desired position of the impact focus of the electrons on the anode. By controlling the deflection current, it is possible to define several foci that can operate simultaneously in multiplex operation.
   Furthermore, X-ray tubes of this design can be significantly smaller and lighter at the same or higher power than conventional X-ray tubes with a rotating anode. This is very advantageous in new technologies of high - speed multi-slice CT devices, where the rotational mechanics are strongly stressed by centrifugal, gravitational and gyroscopic forces. The first type of X-ray tube rotating as a whole (rotating envelope tube) is Straton
(developed by Siemens in 2004).
 High performance x-ray tubes with rotating anode
Another newer type of X-ray tube is Vectron
(from the same manufacturer), with high performance - max. anode current 1300mA at voltage up to 90kV and 800mA at 150kV. Unlike the Straton, it reverts to Coolidge's earlier classic design of a fixed X-ray tube with a rotating anode, with bearings lubricated by a molten metal eutectic alloy. The anode has a high heat capacity and is cooled by infrared radiation through the vacuum into the cooling medium. Instead of oil, water is used as a cooling medium (high voltage sources - generators - are also cooled with water), which has a higher specific heat capacity than oil. The cathode is formed by flat emitter technology. It has a very small focus on the anode of 0.4×0.5 mm, independent of the kV setting, even at high X-ray power. The electron beam is deflected electronically very quickly (4000 times per second) and creates two foci in the multiplex mode (this creates two overlapping projections in the z-axis).
   Also other manufacturers supply high-performance high-resolution X-ray tubes - GE (Performinx HDw), Philips (iIMRC), Toshiba (Megacool Vi).

Electric power supply of the X-ray tube
The X-ray tube, as an electronic source of radiation, requires an appropriate power supply, supplying electrical energy generating X-rays and providing other functions necessary for the correct operation of the device. The X-ray tube has three basic power supplies :
The heating current source
for the X-ray tube cathode. It is a heating transformer, they supply a low voltage of usually 6-12V and a current in the range of approx. 0.5-10 A on their secondary winding, with the possibility of continuous regulation (see below "Setting the X-radiation parameters").

Fig.3.2.4. Electric power supply of X-ray tube
Above: High anode voltage source. Middle: AC voltage for anode rotation. Bottom: Cathode heating voltage.
Note: This arrangement of X-ray tube under high voltage, with a hot cathode, rotating anode and X-ray emission I practically show on the "work table" as an experimental demonstration.

¨ High voltage source
- anode voltage for accelerating electrons in X-ray tube. This is a voltage in the range of mostly about 20kV-150kV; it can also be lower for special X-ray machines for spectrometric use, and up to 400 kV in industrial applications. The basis of this source
(also called a generator) is a high-voltage transformer, which transforms the mains voltage (220V/380V) upwards - either directly from the mains voltage to the required value, or more recently via an electronic oscillating circuit. The high-voltage transformer has a high conversion ratio (given by the ratio of the number of turns on the primary and secondary windings) of the order of 1000 or more.
  Newer devices use high-frequency sources high voltage. The mains voltage is first rectified and smoothed. This DC voltage is supplied to the high-frequency oscillator (inverter), which uses thyristors to generate an AC voltage of about 10kHz with sharp edges. This is then in the high voltage transformer converted to high AC voltage, which is further rectified and smoothed (see below). The advantage of this solution is that the high-frequency transformer can have significantly smaller dimensions and weight at the same power than a conventional transformer with a frequency of 50Hz. This high-frequency oscillator and transformer solution, enabling highly efficient transformation of electrical energy - power, current, voltage - is called an inverter in electronics.
  The value of the anode voltage can be regulated either continuously or stepwise in suitable steps. This is achieved using an autotransformer which is preceded before a high voltage transformer. The autotransformer regulates the mains voltage in the range of approx. 20-220V, which is then multiplied by the high-voltage unit by a constant ratio (approx. 1: 1000). In the case of electronic high-frequency sources, the high-voltage regulation is carried out by means of frequency control.
  The AC high voltage is rectified by vacuum or semiconductor diodes. The simplest rectification is one-way ("single- pulse") by means of one diode connected in series
(or without rectification, the rectification is performed by the X-ray tube herself *), when the emission of radiation occurs only in the positive half-period of alternating current.
*) In fact, the X-ray tube itself is basically a diode that can take care of the rectification itself. For older and simpler devices, therefore, the X-ray tube was supplied with alternating voltage, while the emission of X-rays occurs only in half-periods when there is a positive voltage at the anode. The disadvantage of this solution is the increased proportion of the soft component of radiation (arising at the beginning and end of the half-period, when the instantaneous voltage is significantly lower), and at higher powers also the possibility of reverse current ("backfire") - in the opposite half-period, the secondary electrons emitted from the heated focus of the anode, can accelerate towards the cathode, which they bombard with high kinetic energy and can damage it.
  Two-way rectification is more perfect ("two-pulse") using 4 diodes in a bridge Graetz connection, or 2 diodes in a double secondary winding, where there is always a positive (pulsating) voltage at the anode and the X-ray tube operates in both half-periods of AC voltage. In the past, sometimes a three-phase mains supply of a three-phase high-voltage transformer is used and 6 diodes in a bridge connection are used for rectification - only a slightly pulsating DC voltage is generated (which never drops to zero, but only by about 15% of the peak voltage), the pulsations have a frequency of 300Hz; it is sometimes referred to as six-pulse smoothing. A three-phase high-voltage transformer can have two triple secondary coils, phase-shifted by 60°; after rectification by means of 12 diodes in a bridge circuit, only a slightly pulsating DC voltage is obtained (with a frequency of 600Hz it fluctuates only by about 4% - "12-pulse smoothing"), close to the right smoothed DC voltage.
  All these cumbersome solutions of heavy current electrical engineering are already abandoned here, in principle high-frequency (HF) high-voltage sources are used. With these high-frequency high-voltage sources, capacitor smoothing can be performed very well after rectification, thus obtaining a minimally pulsating DC voltage (with a high frequency of only small pulses).
   The value of high voltage on X-ray tube is expressed in thousands of volts - kilovolts [kV]. If a pulsating voltage is applied between the anode and the cathode of the X-ray tube, the maximum energy of the emitted X-rays is given by the maximum positive value of the anode voltage, which is expressed as [kVp] - "number of kilovolts in the peak".
¨ Power supply for anode rotation ,
which is an alternating voltage (in the simplest case mains 220/380V), applied to the stator coils, creating a rotating magnetic field for X-ray anode rotation. At a mains voltage frequency of 50Hz, the fundamental frequency of rotation of the anode is 3000rpm; by segmenting the stator coils, lower speeds of 1500, 1000, 750, 600 rpm can be achieved. To achieve higher speeds, the X-ray stator must be powered by an electronic oscillator, providing a frequency higher than 50Hz. A higher starting voltage is first applied to the stator, which is reduced to about 1/3 after spinning to maintain synchronous speed. The speed of the anode (rotor) can be electronically monitored on the basis of the current flowing at a given voltage through the stator coils when the rotating magnetic field is excited (after reaching the synchronous speed this current decreases significantly). Only after the anode has been rotated to the required speed can the high anode voltage start and the exposure begin. In order to prevent the X-ray anode from rotating unnecessarily for a long time, the opposite phase of alternating voltage is connected to the respective stator coils for a while after the exposure, whereby the rotating magnetic field reverses its direction and the rotation of the anode is electromagnetically braked
(for some types, a separate winding is installed in the stator for braking, a DC supply for magnetic braking of the rotor is also used). Exceptions are anodes with molten metal (gallium) lubricated bearings, which rotate continuously, even between exposures, as described above.
  Current X-ray devices also use other electronic power circuits to supply electric motors for sliding and rotating movements, X-ray tubes cooling, as well as control and regulation electronics, including detection and evaluation circuits.
X-ray tube cover
During actual operation in X-ray machines, the X-ray tube is encapsulated in a special metal cover (made of aluminum alloys) of cylindrical shape *). The cover is shielded from the inside by lead sheet approx. 3 mm before unwanted penetration of X-rays into the surroundings. At the ends of the housing there are bushings through which voltage is applied between the anode and the cathode by means of well-insulated high-voltage cables, and a low heating voltage is then applied at the cathode end. In the middle part *) of the cover there is an exit window (of course unshielded, to which the X-ray tube must be turned with its impact focus) made of a light material, mostly acrylic glass, through which the X-ray beam comes out for the respective use. In power X-rays tubes, the space between the X-ray lamp and the walls of the package is filled with a cooling medium - transformer oil. The oil environment also increases the electrical strength of circuits - prevents high voltage electric shocks. To eliminate the mechanical stress on the cover due to the thermal expansion of the cooling oil, a rubber expansion membrane is built in a suitable place on the wall of the X-ray tube cover
(prevents, for example, the cover from bursting when the X-ray tube overheats).
*) For Straton X-ray tubes, the housing has a double cone shape and the exit window is near the end where the anode is. The appropriate location of the impact focus on the anode opposite the output window in the housing is ensured by deflection coils (incl. the angle of their rotation), located from the outside in the narrowed part of the housing - Fig.3.2.2 on the right.
  Smaller, low-power X-ray instruments sometimes use a compact design: a high-voltage transformer (for the anode) with a winding of heating current for the cathode is built into the common housing together with the X-ray tube. Only 220V mains supply is then conduct to such a compact system.
Collimation and localization system
The following is a collimation system, consisting of a tubus with adjustable orifices defining the geometric shape of the X-ray beam. The apertures are adjusted so that the X-ray beam covers only the displayed area and other parts of the body are not unnecessarily irradiated. For visual localization and adjustment of the displayed field a light navigation system is installed in the X-ray tube collimation system - the light from the filament lamp or LED is guided by optical projection through the collimation system so as to achieve a conformity between the visible light field and the X-ray field. Before the examination, it is possible to adjust the position of the displayed field on the film cassette or display panel, as well as on the surface of the patient's body.
Mechanical design of X-ray devices
The cover with X-ray tube and collimation system, together with the opposite film cassette or imaging panel, are mounted on special stands of several types and constructions, according to the required X-ray imaging methodology - Fig.3.2.5. For skiagraphic imaging, the X-ray is most often mounted on top of a vertical stand (column tripod mounted on the floor or ceiling mount) with the possibility of easy mechanical movement. The film cassette or display panel is mounted at the bottom of the stand, again with the possibility of sliding. Between them is a sliding bed with the patient for examination while lying down. For standing (or sitting) imaging, the X-ray tube with the collimation system is rotated horizontally, the opposite cassette or flat-panel is on a separate vertical stand (so-called vertigraph). The horizontal movement (travel) of the X-ray tube can be realized by means of rails mounted on the ceiling or floor of the examination room. Displacements of individual parts of the X-ray system can be manual or motorized, using electronically controlled electric motors. For new systems, the so-called autotracking is implemented - automatic synchronous coupling of display panel and X-ray tube displacements (equipped with a collimation system). Some systems for sciagraphy and sciascopy have the ability to turn or tilt the X-ray stand, imaging panel, and the bed for various angles, from horizontal to vertical - such a system is called an X-ray tiltable wall and has a wide range of uses.
  For flexible sciascopic imaging, the X-ray tube and the opposite imaging detection system are often mounted on a special stand in the shape of the letter "C" - the so-called C-arm - Fig.3.2.3b., or the so-called U-arm - Fig.3.2.3c. These arms can be rotated using electric motors to different angles around the patient, allowing flexible display in different projections. These systems (which are sometimes mobile) are used in a wide range of applications, such as digital subtraction angiography (see below DSA - Fig.3.2.3), X-ray navigation of interventional procedures, afterloading in radiotherapy (see §3.6, section "Brachyradiotherapy" - Fig.3.6.7) and others.
  A separate category of X-ray device design solutions is transmission X-ray tomography CT, where the X- ray tube and the opposite electronic detection system are mounted on a portal rotary stand - gantry - described in more detail below "
X-ray tomography - CT", Fig.3.2.4. This includes also special constructions of X-ray imaging devices, installed directly on IGRT radiotherapeutic irradiators (see §3.6, section "Isocentric radiotherapy", Fig.3.6.1c) or tomotherapy (§3.6, part "Modulation of irradiation beams", Fig.3.6.4a).
  Further technical details of the construction of X-ray devices and their accessories are already outside the scope of this physically focused treatise...

X-ray tube aging and damage
Like almost any device or electronic component, even X-ray tubes can be subject to adverse changes due to "aging", load and operational damage. Over time, material degradation, vacuum, and wear of rotary bearings occur. Several types of adverse changes in X-ray tube properties can occur :
"Burning out" of the cathode filament
The red-hot cathode, providing electrons by thermoemission, is mostly made of tungsten wire, which is highly heat-resistant (melting temperature is 3400 degrees). However, when annealing at high temperatures, tungsten evaporates slowly, often unevenly, so that the spiral can become quite thin in some places. This can limit the life of the cathode spiral.
--> Violation of the X-ray tube vacuum
Small leaks can occur between the glass and metal parts of the x-ray bulb, allowing air to gradually enter the vacuum environment of the x-ray tube over longer periods of time. From the metal parts, especially from the heated ones, and from the stressed rotary bearings of the anode, unwanted fumes also evaporate and release into the environment of the X-ray tube.
--> Local overheating of the impact focus on the anode
At high powers, the impact focus can be strongly locally heated to temperatures briefly exceeding the melting temperature of the anode material. This results in small micro-damages, various pits and cracks. The production of X-rays is then no longer perfectly homogeneous...
--> Wear of the anode rotary bearings
The high temperature and rotation frequency lead to a large mechanical load on these bearings. The wear of the bearings leads to greater noise during the operation of the X-ray tube and deterioration of the stability of the focus. Wear can be prevented to a large extent by lubrication of the bearings, in the case of high-performance X-ray tubes, lubrication using a molten metal eutectic alloy is used.

Setting of X-ray parameters
To optimize X-ray diagnostics, it is necessary to set suitable X-radiation parameters. In the electrical circuit of the X-ray tube we can regulate and set two basic electrical parameters as required
(the third is only time, the fourth is realized by mechanical arrangement) :
¨ Anode voltage U [kV] ,
which is a high voltage supplied between cathode and anode, determines the maximum and mean energy of the photons of the resulting X-rays, its "hardness". The maximum energy of X-rays in [keV] is numerically practically equal to the anode voltage U in [kV], the mean energy is slightly higher than 1/3 of the max. energy. With increasing anode voltage, the whole spectrum of X-rays shifts towards higher energies
(shorter wavelengths) and increasing the relative proportion of higher energies (harder short-wave components).
   In practice, the anode voltage varies in a wide range from about 20kV to 200kV (depending on the type of displayed structures), in the industrial use of X-rays then higher.
The energy - hardness - of X-rays emitted is often referred to in X-ray jargon as the "quality" of X-radiation.

Anode current I [mA]
flowing through the X-ray tube, determines the intensity (fluence) of X-radiation emitted by the X-ray tube I
X. It is most easily regulated by changing the heating of the cathode - the glowing stream - and thus the temperature of the cathode fiber. The glow current can be regulated simply by means of a rheostat in the glow circuit (in the glow transformer circuit), more recently with special electronic circuits equipped with transistors and thyristors. At higher heating of the cathode fiber, more electrons are emitted, a larger stream of electrons flows through the X-ray tube and a higher intensity of X-rays is emitted (but see "Volt-ampere characteristics of X-ray tube"). The average X-ray tube current is in the range units of mA to about 200mA, the peak current can be significantly higher (in pulse mode). E.g. an X-ray tube with a tungsten anode (Z=74) supplied with an anode voltage U = 120 kV at a X-ray tube current of 1 mA emits X-rays with an intensity IX of approximately 6.1013 photons /s. (follows from the approximate relationship for braking radiation production efficiency given above in "X-ray tube", passage "Braking X-rays").
¨ Exposure
is the total amount of X-rays photons, which determines the quality of X-rays images and also the radiation exposure of the patient. It is given by the product of radiation intensity I
X (photon fluence /s.) and exposure time T - it is therefore proportional to the product of anode current by X-ray tube [mA] and exposure time [s]: "milliampere-seconds" mA.s = Q, which is the total charge Q electrons or the electrical amount that passes through the X-ray lamp during exposure (the coefficient of proportionality h is given above in the section "Braking X-rays"). At Q= 1 mAs, approx. 1013 of photons is radiated from the anode, of which only a small part is used for X-ray imaging - of most photons flying in different directions, only a relatively narrow conical beam is selected by collimation, low energy photons are further removed by filtration (see below).
Note: The energy of X-rays does not depend on the magnitude of current, time or electrical amount.
  For the acquisition of common skiagraphic images of soft tissues, the exposure to X-rays of about 2
- 6 mAs is used, for the skeleton about 20 - 80 mAs, for CT even 200 mAs. In modern X-ray devices capable of operating in a pulsed mode with high instantaneous power, a high current value [mA] at a short exposure time [s] is preferred to achieve the desired exposure [mAs] - thus reducing the risk of blurring the image with patient movement.
   Based on empirical experience, the recommended value of anode voltage [kV] and exposure [mAs] is determined for each type of X-ray examination, providing a quality image of the required structures at a relatively low radiation exposure. Some devices also use automatic exposure, which electronically switches off the anode voltage in the generator - and thus the exposure - after reaching a certain preset "amount" of X-rays. For the purposes of automatic exposure, the flux of transmitted X-rays is monitored by means of ionization chambers placed behind the film cassette or behind the flat panel. For digital imaging detectors, presetting the total number of pulses stored in the digital image can also be used to interrupt the exposure. Similarly, exposure optimization works for CT instruments, where according to the signal level from the detectors in the pre-planning radiographic display of SPR (topogram), the optimal current values [mA] can be automatically adjusted during self-diagnostic scanning - ATCM (Automatic Tube Current Modulation), see below "CT".
The already generated X-rays are subsequently treated by filtration :

Filtration, collimation 
The soft X-rays of longer wavelengths and the low energy of photons at the beginning of the continuous spectrum of X-radiation are of no significance for diagnosis, they are usually absorbed in the skin and shallow layers of tissue - causing only unwanted radiation exposure. Therefore, it is removed by filtration - an aluminum or copper plate about 1.5-4 mm thick is inserted into the radiation path, which absorbs the soft component of X-rays to a large extent, while transmitting the harder component - see the shape of the spectra in Figure 3.2.5. The initial partial filtration of the generated X-rays is already created by passing through the material of the anode, the glass flask of the X-ray tube, the cooling oil, the material of the cover and the outlet window
(thus the X-ray tube glass flask itself acts inherently as a partial filter, equivalent to about 0.5 mm Al; similarly the cooling oil and the X-ray tube cover window).
   In some special cases when we need sharper and more selective filtering of certain areas of the energy, so used filtration K-edge (K-edge filter). It is based on significantly increased ("resonant") absorption of photon radiation at energy equal to or slightly higher than the binding energy of electrons on the K-shell of atoms of the material used (see §1.6, section "Interaction of gamma and X-rays", Fig.1.64). By combining a standard filter (Al, Cu) and a filter made of a suitable heavier material using the K-edge effect, we obtain a bandpass filter, selecting a certain section of energies from a continuous spectrum of X-rays. This is especially true for mammography, where a molybdenum or rhodium filter cuts off photons of higher energies than about 20-23keV to achieve better contrast (see "X-ray mammography" below). Also for DEXA methods - analysis of absorption using two X-ray energies (see below "CT with 2 X-ray tubes- DSCT: Dual Source and Dual Energy CT").
The geometric delimitation of the X-ray beam is performed by collimation (mentioned above) .

Fig. 3.2.5. Basic scheme of X-ray examination.
Left: Arrangement of X-ray tube, filter, DAP-meter, primary and secondary diaphragm, film or detector. Right: Energy spectra of X-rays from X-ray tube, after filtration and after passing through the patient.

Scattered X-rays
When X-rays interact with matter, Compton scattering of photons on free or weakly bound electrons occurs, among other things. These scattered photons fly out of the tissue with lower energy and in different directions. The proportion of scattered radiation is greater the larger the patient (and is also higher for harder X-rays from tube - higher voltage [kV]). Scattered radiation degrades the quality of the X-ray image - it reduces its contrast. The possibility of suppressing scattered radiation is mentioned in the following paragraph. This Compton-scattered X-rays further cause some small radiation exposure even outside the X-ray beam itself from the X-ray tube. This radiation exposure is not high, scattering albedo of the human body for X-rays is less than 1%
(albedo was discussed in §1.6, section "Interaction of radiation passing through matter") . At a distance of 1 m, the radiation dose is about 0.1 mGy /1 mAs, in 2 meters only about 0.005 mGy /1 mAs. Nevertheless, during X-ray examinations, radiation workers should not stay in the examination room during the exposure, unless necessary, but in a shielded control room - the exception is, for example, interventional radiological procedures.
Filters and apertures for X-ray imaging
In the practical diagnostic application of X-rays, it is important to use filters and collimating screens - Fig.3.2.5. Suitable primary apertures ensure the geometric delimitation of the X-ray beam, reaching only the necessary examined area (a sharp image with high spatial resolution is ensured by the fact that the radiation comes from an almost point focus on the anode by X-ray tube, as described above). Immediately behind the X-ray tube is placed a primary filter, most often made of aluminum sheet, which absorbs low-energy photons (from the beginning of the continuous X-ray spectrum), which are not usable for imaging (they would penetrate only into the subcutaneous tissue) but would increase the patient's radiation exposure
(mentioned above). This is followed by a tubus with adjustable diaphragms for the geometric delimitation of the X-ray beam (size of the imaging field). A thin plane-parallel ionization chamber is usually mounted on the outlet window of the tube for monitoring the exposure to X-rays, the so-called DAP meter (see below "Radiation load during X-ray examination"), allowing to determine the radiation dose of the patient during X-ray examination - Fig.3.2.5 top left.
   A secondary aperture is then placed between the patient and the film (or screen or imaging detection system, flat panel). It is a lattice formed by parallel or diverging absorption lamellae (lead strips) which, through their gaps, transmit only primary X-rays passed in the direction of the original beam, while secondary Compton-scattered photons (moving in other directions) are absorbed in the partitions. It is therefore a collimator, which attenuates the radiation only minimally in the primary direction, while the attenuation of the obliquely transmitted scattered radiation is considerable. The quality of the secondary screen is determined by the grid density (number of lamellae per centimeter) and the grid ratio (the ratio between the distance of the absorbent strips and their height). Suppression of secondary scattered radiation significantly improves the contrast of the X-ray image. On the other hand, the secondary aperture also absorbs some of the useful X-rays
(eg with a Bucky aperture, the attenuation is about 1.8 times), so it is necessary to increase the exposure. Three types of secondary screens (grids) are used :
- Converging focused Bucky-Potter screen (approx. 10 slats /cm). The Bucky screen has relatively thick partitions (approx. 1 mm), which would be projected into an X-ray image and have a disturbing effect. This disturbing raster is eliminated by moving the aperture during exposure, blurring its image and disappearing into the overall background.
- Parallel fine Lysholm screen (40-60 slats /cm).
- Ultrafine Smith screen (density >100 lamellae /cm). Due to too high absorption, it is not used in practice.
Terminological note:
In radiodiagnostics, it is often customary to use the not very precise generic name "aperture". More precisely, however, these are collimators and filters.

Visualization and recording of X-ray images
X-rays, carrying density information after passing through the displayed tissue, is invisible to us, so it is necessary to "make it visible" or register it using suitable material and electronic methods. There are basically three ways to display this X-ray :
Visual observation of the image on a luminescent screen
A luminescent screen is a plate or foil on which is coated with a layer of suitable material (such as zinc sulphide, .......), in which fluorescence is produced upon interaction with ionizing X-rays. Thus, we can see a projection density image on the luminescent screen via X-rays. The advantage of this simplest method was the possibility of continuous dynamic observation - sciascopy. However, the main disadvantage is the low sensitivity and high radiation exposure of both the patient and the radiologist...
- Imaging on photographic film
Ionizing X-rays cause a photochemical reaction in suitable photographic materials, mostly containing silver bromide. This photographic emulsion, coating on the surface of a plastic film, forms a photographic film in which the incident X-rays form a latent photographic image - is described in more detail in §2.2 "
Photographic detection of ionizing radiation". After development, we receive a negative density photographic image in X-rays.
- Scanning with an electronic imaging detector,
which registers incident X-rays, converts it into electrical impulses and, after procesing by complex electronic circuits, creates digital density images in computer memory - this most advanced method is described in more detail below "
Electronic imaging X-ray detectors"...
Amplifiers and digital sensors of X-ray image; indirect and direct digitization
Direct X-ray photographic film for display belongs to the past, and is gradually replaced by more sophisticated technologies. To increase the sensitivity of the scanning X-ray suitable image enhancement methods are used in the image, more recently methods of electronic image capture. This makes it possible to significantly reduce the required intensity of X-rays and thus the radiation dose for the patient, as well as to reduce the undesired exposure for radiation workers.
- Amplifying foils
 During photographic skiagraphy, amplifying luminescent foils are attached to the film, the task of which is to convert X-rays into light, which is exposed by the photographic film. It consists of a layer of phosphor dispersed in an emulsion of gelatin or nitrocellulose. Use calcium tungstenate (emits blue light), lanthanoxid bromide (blue light), gadolinium-carryover (green light), barium chloride (blue light)... The film has to be sensitized to the color of the light from the phosphor. It is already abandoned.
- Xeroradiography ,
where a positively charged semiconductor (selenium) plate is used instead of film. The incident photons of X-rays there evoke a photoeffect, in which the photoelectrons locally compensate for the original positive charge - a latent electrostatic image is created ; in the copier, powder particles of dye are then attracted to differently charged places of the plate, which are finally printed on paper where a visible image is created. It is also abandoned.
- Memory foils
replace the film in the X-ray cassette and retain a latent electron image after radiation exposure. The sensitive layer usually contains europium atoms (
BaFCl: Eu2+, instead of Cl it can be iodine or bromine). The impact of X-rays photons excites in the sensitive layer of the film, electrons are released from the europium atoms, which are trapped in the halide metastable levels of so-called "electron traps" - a latent electron image is formed . This latent image is made visible after exposure by photostimulation using a laser infrared beam: a "trapped" electron is released into the conduction band, after which the electron is captured on the excited surface in the luminescent center, followed by deexcitation to baseline, accompanied by photon emission of light. This light is registered by a sensitive photomultiplier, the generated electrical pulses are sampled and converted by an analog-to-digital converter into digital image information. The device that reads and digitizes the latent image is called a digitizer - it is an indirect digitization of an X-ray image. It will be temporarily used for some time...
- Image intensifiers
During sciascopy, the image on the fluorescent screen ("shield") is amplified by a special image tube - image intensifier. The light image created by the impact of X-rays photons on the fluorescent screen of the inlet window of the tube causes a photoelectric effect on the enclosed photocathode by its emitted photons of visible light. The emitted photoelectrons are accelerated by a voltage between the photocathode and the anode (approx. 10-20 kV) and directed by electron optics to a second luminescent screen, where they create a reduced inverted image, but its brightness is more than 1000 times greater than the original image. This image is then optically captured by a video camera and displayed on a TV screen or computer monitor. It is still used for older devices, new devices are already supplied with digital flat-panels.
- Electronic imaging detectors - flat panels 
All of the above image amplifiers or television sensors are only a temporary technical solution. New systems are equipped with an electronic digital image sensor on a compact so-called flat panel, consisting of a scintillator (e.g. cesium iodide CsI: Tl or based on gadolinium Gd
2O2S:Tb - GOS; see §2.4 of scintillators, the "Scintillators and their properties") and semiconductor optoelectronic amorphous silicon (a -Si). The detection panel consists of a large number of elements - cells, pixels, assembled into an image matrix of about 2000 x2000 elements, and more. The most perfect imaging detectors are semiconductor pixel detectors (SPD), see §2.5 "Semiconductor detectors". The pulses from the individual elements of the detector are stored in multiplex mode with the help of an analog-to-digital converter (ADC) directly into the computer's memory - they create a digital X-ray image. Only this technology of so-called direct digitization belongs to the future... - it was described in more detail below in the section "Electronic X-ray imaging detectors".
Note: The terminology of "direct" and "indirect" digitization is temporary and will be abandoned soon: all X-ray images will be automatically primarily digital
(using of technology now called "direct digitization").

Electronic imaging detectors of X-radiation
The former X-ray imaging using photographic film or a luminescent screen is now generally being replaced by electronic imaging detectors - Fig.3.2.6. The advantage is significantly higher detection sensitivity and wide possibilities of electronic and computer image processing (digitization). All these imaging detectors are based on modern technologies called quantum optoelectronics (photonics), which use an internal or external photo effect to convert photons into electrical signals.

¨ Image intensifier
Electronic imaging with image intensifier was widely used in the 1960s - 1980s - Fig.3.2.6 left. The image intensifier is a special vacuum tube with two windows - input and output. On the inside of the inlet window is a layer of scintillator (mostly cesium iodide) and below it a thin metal layer of photocathode. The incident X-rays cause flashes of light in the input scintillation layer, which by photoefect eject electrons from the photocathode. The electrons thus generated are then attracted by annular accelerating and focusing electrodes, to which a high positive voltage is connected (gradually increasing up to about 30 kV at the anode at the output scintillator). This electro-optical system, acting as a conjunctive "electric lens", throws electrons onto an output scintillator (mostly ZnS: Ag), where the accelerated electrons produce intense flashes. Thus created reduced, inverted but very clear ("amplified", intense) image is then captured by an optical TV camcorder and (analog) displayed on the TV screen. Later, digital CCD camera imaging with computer image recording was used. Image intensifiers or TV sensors were just temporary technical solution, the future belongs to fully digital X-ray image sensors - flat panels :

Fig.3.2.6 Electronic imaging detectors in X-ray diagnostics.
Left: X-ray acquisition using an image intensifier. Middle: Flat-panel with indirect (scintillation) and direct (semiconductor) conversion of X-rays into electrical signals. Right: Ring-arranged CT detectors with fast ceramic scintillators and photodiodes.

¨ Flat - panels
Modern and more advanced electronic X-ray imaging detectors are so-called flat panels
( flat - they have the planar shape of a thin plate), which provide signals for direct digital X-ray image. The detection panel consists of a large number of elements - cells, pixels, assembled into an image matrix of about 2000 x 2000 image elements, and more. The level of the electrical signal from each pixel is proportional to the intensity, resp. the number of X-rays photons incident on a given location of the flat-panel. From electronic multiplex registration circuits (multiple read-out) the image signal via the ADC is fed to the image matrix of the computer, in the individual elements ( pixels) of which information about the intensity of X-rays from the corresponding location of the irradiated object is stored. In terms of the method of detection and converting X-rays into electrical signals, flat panels of two types are constructed :
- Scintillation detection (indirect conversion) ,
when photons of X-rays impinge first on the layer of scintillator material (most commonly used cesium iodide CsI: Tl) *) in which they evoke flashes of visible light
(scintillators and their use for the detection and spectrometry of ionizing radiation is discussed in detail in §2.4 "Scintillation detection and gamma-ray spectrometry"). This scintillation light then enters the semiconductor photodiodes (mostly silicon - a-Si - amorphous silicon is used as the semiconductor material), in which an electric charge is released by the internal photo effect (electrons and "holes") and the light is thus converted into an electrical signal - Fig.3.2.6b. The term "indirect conversion" here means, that X-rays are first converted in the scintillation layer into visible light, which is then converted into an electrical signal in the photodiodes. This flat panel design is currently the most widely used.
*) Scintillation crystals of some flat panels have a "fibrous" design. They are composed of densely spaced thin vertical needles. This design reduces the lateral scattering of light scintillations, leading to a sharper image.
ISS technology
Some new types of flat panels with scintillation conversion have a somewhat curious design with "opposite orientation" with respect to Fig.3.2.6b: the displayed X-rays are coming on the side of the layer of photodiodes and reading TFT transistors, passes through this layer and then impinges on the CsI scintillation crystal. Scintillation interactions then usually occur in a layer near the photodiodes, thereby reducing lateral scattering by scintillation and improving resolution. This solution, referred to as ISS (Irradiated Side Sampling) technology , is made possible by the electronic miniaturization of the photodiode and reading circuit layer, which is thin and has virtually no effect on the transmitted X-rays.
- Semiconductor detection (direct conversion) ,
where photons of X-rays fall directly into semiconductor detectors (suitable material is CdTe,
CdZnTe, or Si for lower energy). Absorbed X-rays create electron-hole pairs, electrons in a strong electric field drift to the anodes, where they generate short electrical pulses (approx. 10-9 s). The amplitude of these pulses is proportional to the energy of the absorbed X-ray photons. Thus, photons X release electric charges through their interaction and are directly converted into an electrical signal (the general principle of semiconductor detectors is given in §2.5 "Semiconductor detectors") - Fig.3.2.6c. They are also constructed in smaller dimensions (in the order of centimeters - MEDIPIX ) with a very high density of miniature image elements (high image resolution) and are used in special laboratory methods such as X-ray microscopy (mentioned above), or animal X-ray imaging. Flat panels with direct semiconductor detection are still only rarely used in clinical X-ray diagnostics, but they certainly have a future. Compared to scintillation conversion, they have significant advantages and provide new possibilities in X-ray imaging :
Spectrometric Photon-counting X-ray imaging
During X-ray diagnostics, the examined tissues and organs are irradiated with X-rays with a continuous spectrum of a wide range of energies - almost from zero to the maximum energy, given by the value of high voltage in the X-ray tube. In scintillation conversion, all signals from the photodiodes are registered integrally, regardless of the energy of the X-photons from which they originate. Including electronic noise of detection system, which may be dominant at low X-ray flux. However, direct semiconductor detection allows independent registration of individual X-ray photons: "photon couting". And then electronic analysis of the amplitudes of these pulses - performing spectrometry of the detected X-rays.
Terminological note: Instead of the name "photon-counting imaging", which was introduced in the field of radiodiagnostics, the name "photon-spectrometry imaging" would be more apt from a physical point of view, as energy (spectrometric) analysis of detected X-rays photons is important here (rather than simple "photon counting").
  By suitable setting of the lower discriminant level of the amplitude analyzer
(to a value corresponding to approx. 20 keV) it is possible to completely cut off the electronic noise of the detector. We obtain contrast images with reduced noise, given only by statistical fluctuations of the detected X-rays. This allows you to reduce the radiation dose while maintaining good image quality.
  By using two or more different amplitude- energy windows of the analyzer, images can be obtained for different energies of transmitted X-rays. We can then use the material-specific difference in the absorption of X-rays of different energies for material-tissue differentiation of the displayed structures, which otherwise show the same density attenuation in a conventional X-ray image. It can be, for example, a material differentiation of the composition of kidney stones. This possibility of material analysis in X-ray images is discussed in more detail below in the section "
Dual- and multi-energy X-ray imaging. Material composition analysis.".
Adverse effects at spectrometric photon-counting detection

Under optimal conditions, photon-counting detectors can provide better energy separation with less overlap than with dual energy X-ray technology. But there are some adverse physico-electronic influences that reduce this energy separation. X-ray photons absorbed around the boundaries of the detection pixels can be divided - shared - by neighboring cells in the detector. In the material of the detector, excitation and deexcitation of electrons on the K-shell also occurs with the emission of fluorescent photons, that can escape and be detected in neighboring cells. At high intensities (flux) of X-rays, the dead time and the pile-up effect of semiconductors are also adversely affected, whereby some signal pulses overlap and are not registered separately, but as a single pulse. This signal-sharing between cells, dead time, and pile-up effect of semiconductor detectors can impair spectral separation, spatial resolution, and detection efficiency.
Terminological note:
 Please do not confuse indirect and direct X-ray conversion in flat panels with indirect and direct digitization of X-ray images! It has nothing to do with it, here it is always a direct digitization, only with a different physical-technological design.
  In both cases of the flat panels, the electrical signal from the photodiodes or semiconductor detectors is sensed by a special matrix of TFT (Thin Film Transistors) transistors implanted in thin-film integrated circuit technologies on a glass support. Scanning, so-called read-out, takes place in multiplex mode in the X and Y directions - it provides coordinate pulses about the position of the X-ray photons detection location in the flat panel. These coordinate pulses are converted to digital form by an analog-to-digital converter (ADC) and stored in the corresponding memory addresses in the image matrix of the computer - a digital X-ray image is created. The transfer of image data to the acquisition computer from some modern flat panels is solved wirelessly using modems (WiFi).
  Electronic imaging flat-panels are also used in verification and dosimetric systems of so-called image- guided radiotherapy with modulated beams - in isocentric irradiators with a linear accelerator and in cybernetic gamma knifes (§3.6, part "Isocentric radiotherapy" and "Stereotactic radiotherapy SBRT"); they are sometimes abbreviated EPID (Electronic Portal Image Device).
¨ X-ray detectors for CT
The current CT X-ray detectors work on a similar principle as flat-panels with indirect conversion. They consist of a large number of semicircular arranged elements, each of which is a small scintillation crystal (scintillators based on rare earth-doped silicon oxides, such as lutetium and yttrium - LYSO, or gadolinium oxisulphide, with a very short flash duration are now used), optically coupled with photodiode; see below "X-ray tomography - CT", section "X-ray detectors for CT ".
  Instead of scintillation detectors, semiconductor detectors are gradually being introduced, which convert X-ray photons directly into electrical signals. This method, called photon-counting CT, provides more quality images with lower noise and the possibility of spectrometric analysis .......
Properties of electronic X-ray detectors
An important characteristic of electronic imaging detectors is their spatial resolution, which is the smallest distance of two "point" objects at which they still appear as two separate structures; or equivalent to the half-width of the point object image profile (FWHM). At shorter distances, both objects appear as one, they are not distinguished. The resolution is given mainly by the size of the individual pixels of the detector (this limits the smallest possible point that can be displayed), it is also affected by the scattering of X-rays and light in the detector and the processes of converting X-rays into electrical signals. As in photography, resolution is often measured at the maximum number of lines per millimeter [lp/mm], which can still be distinguished. Modern flat-panels theoretically reach up to 10 lp/mm, which corresponds to a resolution of 0.1 mm; in practice, however, the real resolution is around 2-5 lp/mm. The quality of X-ray images in terms of real resolution is sometimes quantified in detail using the so-called modulation transfer function MTF, which indicates using Fourier harmonic analysis, which details of the object can be displayed with a given contrast.
  Another important parameter of electronic X-ray detectors is the sensitivity of the sensor. It is reported numerically using the detection quantum efficiency DQE (Detection Quantum Efficiency), which is the percentage of X-ray photons incident on the detector that is actually recorded by the detector and used to create the image (the rest is uselessly absorbed by the input window or detector material, without a scintillation or electrical response)
.   Electronic imaging detectors enable more detailed analysis of fine structures using enlarged image sections - so-called "zoom" or "magnification". In the case of image intensifiers, this is an analog zoom (achieved by changing the voltage), in which a smaller part of the input area is "stretched" over the entire image at the output. In the flat panel it is a digital zoom - additional software zoom of the selected part of the image, without changing the resolution. If we want to maintain a high image quality on the magnified image (maintain the same signal-to-noise ratio as in the basic image), it is necessary to increase the number of incident photons, which will increase the dose. Digital X-ray systems, including CT, are equipped with software providing a number of other options for computer image editing - post-processing.

Dual- and multi-energy X-ray imaging. Material composition analysis.
Different tissues and organs differ in their chemical composition, which may or may not be reflected in their different densities. If two adjacent structures in the body have the same or close absorption coefficient
(linear attenuation coefficient) for the X-rays used, they will be indistinguishable from each other on X-ray images - they will appear identical, even if their material (chemical, elemental) composition is significantly different. Differentiation or classification of different tissue types by standard X-ray imaging is therefore very difficult and often impossible. Standard X-ray diagnostics, using X-rays from a continuous-spectrum from X-ray tube, is basically only anatomical-morphological, density-based, not functional-physiological.
  A certain possibility of at least partial resolution of the material composition of the displayed structures is measurement - imaging - at different X-rays energies - X-rays spectrometry. By using two or more different energy windows, images can be obtained for different energies of the transmitted X-rays. We can then use the material-specific difference in the absorption of X-rays of different energies for material-tissue differentiation of the displayed structures, which otherwise show the same density attenuation in a conventional X-ray image. This could, to some extent, overcome the traditional limitations of X-ray diagnostics to a mere density morphology ..?..
  Different types of substances (and tissues) differ not only by specific values of linear attenuation coefficients
m for X radiation of a certain energy, but also by somewhat different dependence m(EX) of absorption for different energies EX of radiation X. This is due to different electron density configuration at different atomic and the molecular composition of the analyte. For X-rays of energies used in X-ray diagnostics (approx. 20-150 keV) there is an interaction with tissue by photoeffect and Compton scattering - §1.6 "Ionizing radiation", section "Interaction of gamma and X-rays", Fig.1.6.3. For the effective cross-section of the photo effect, the approximate dependence s(EX) ~ Z5/EX 3 applies, where EX is the energy of X-rays photons and Z is the proton - atomic number of the substance (Compton scattering absorption is only slightly dependent on EX energy). For the passage of X-rays through a substance of density r, the linear absorption coefficient will be m(EX,r) ~ r. Zeff 4 /EX 3, where Zeff is the average - "effective atomic number" of this irradiated substance.
  Mathematical analysis of the exponential laws of absorption I = I
o.e-m(Ex).d for individual energies EX and tissue types with absorption coefficients m(EX) (by logarithm the relevant exponential equations are converted to linear) can determine the proportion of absorption in different tissues with different effective atomic number. This can in principle be used to additionally distinguish different types and compositions of tissues based on differences in the density images of the same site, obtained with different energies EX of radiation. This method of differential density analysis DEXA (Dual Energy X-ray Absorptiometry) is similar to that used for density images in "Bone Densitometry" (see Fig.3.2.11 below).
  Not only does this provide more detailed anatomy images, but in some cases it allows different types of tissue to be distinguished (eg bones, blood vessels, adipose tissue), different types of kidney stones, deposition of sodium urate crystals in joints (gout), or quantification of contrast medium distribution in myocardium.
Virtual X-ray images

A notable benefit of two- and multi-energy X-rays imaging is the ability to subtract - suppress, remove - structures of a particular material from images: create virtual images, in which are remove certain significant structures that can be disruptive and overlap some other important details. Differences in atomic numbers make it possible to distinguish calcium and iodine using the dual-energy imaging technique. It is, for example, the identification and removal of bone images, which improves the direct visualization of blood vessels with iodine contrast agent. In a similar way, images of calcified plaques
(which over-cover the lumen of the vessel) within the vessels can be identified and removed from CT images in angiography with contrast agent, allowing a clearer visualization of the patency of the vessels and areas of stenoses. Or create virtual images with iodine removal to more clearly assess the underlying soft tissues, or better visualize smaller kidney stones (which could be overwhelmed by the presence of iodine contrast agent).

Angiographic image of a section of the aorta.

Up: Calcified plaques at several sites of the stenosis over-cover the lumen of the vessel.

Down: Calcium plaques can be identified and subtracted in the 2-energy analysis, providing a clearer visualization of vascular patency in the area of stenoses
(indicated by arrows)

(Angiographic images were taken by MUDr. .... from the Radiodiagnostic Institute of the University Hospital Ostrava)
! replace with our image!

Realization of dual- and multi-energy X-ray imaging
Dual or multi-energy X-ray imaging can be technically performed in basically four alternative methods :
At the resource level :
-> Use of two X-ray tubes with different anode voltages. This is described below in the section "CT with 2 X-rays - DSCT: Dual Source and Dual Energy CT", Fig.3.2.9.
-> Use of one X-ray tube, whose anode voltage is electronically switched between low value (approx. 50 kV) and high value (approx. 120 kV). It is desirable that this switching be fast enough so that there are no motion changes between the energies (when used in CT, it is recommended in milliseconds).
At the detector level :
-> A dual-layer detector with one X-ray tube with the use of two layers of detectors that selectively detect the low- and high-energy component of X-rays (mentioned above in the section "X-rays detectors for CT", section "Dual-layer CT detectors").
-> Spectrometric Photon-couting X-ray imaging with one X-ray tube and semiconductor detectors, enabling flexible energy spectrometry of detected X-rays (described in the section "Electronic X-ray imaging detectors", paragraphs "Semiconductor detection (direct conversion)" and "X-ray detectors for CT"). This method is the physically most perfect, providing a multi-energy imaging.
  We deal with the physical and technical aspects of these methods in the sections "Electronic X-ray imaging detectors" - "Spectrometric Photon-counting X-ray imaging", "X-ray detectors for CT" and "CT with 2 X-rays - DSCT: Dual Source and Dual Energy CT ".

X-ray planar imaging - skiagraphy, sciascopy
X-ray projection
The human body is a complex 3D system of a large number of differently arranged tissues, organs, bones, body cavities, etc. During X-ray transmission imaging, these individual structures can be mually "ovesshadowing" and overlap each other - this can prevent their good display and recognition of possible anomalies. This interference and the overlap of the displayed structures substantially depends on the angle of transmission beam. As a rule, it is possible to find the projection angle for which the lesion is best shown, without disturbing from surrounding structures. Based on the long-term experience of radiologists, certain projections are prescribed for planar X-ray examinations of each organ or area, that provide the best imaging - eg anterior-posterior AP projections, anterior-posterior PA, left LL (l
atero-lateral, also SIN-sinister) or right RL (right-lateral, also DX-dextrum) side projections, oblique projections left LAO (left anterior oblique), LPO (left posterior oblique), or right RPO, RAO and other projections and special positions. The problem of overlapping structures is largely eliminated in CT X-ray tomography - see "X-ray tomography - CT" below, which provides images from different angles and projections.
  In terms of X-ray imaging and processing, planar X-ray diagnostics are divided into two groups :
In simple X-rays imaging, called skiagraphy, incident X-rays passed through the examined tissue on photographic film containing silver halides (silver bromide), in which photochemical reaction leads to the release of silver from the bond in the compound - formed latent image, which is in development in the developer visualized with the density of grains of colloid silver; the remaining silver bromide is dissolved in the stabilizer. The blackening density of the film is proportional to the amount of X-rays passed. The resulting X-ray photographic image represents a negative imaging of tissue density: sites with low density (soft tissues) have lower absorption and therefore high blackening, sites with high density (eg bones) absorb X-rays more and are therefore shown light (low blackening) on the film.
   For X-ray imaging, special films are used, the emulsion of which is thicker and contains an increased content of silver halides compared to conventional photographic materials *). Films are produced in various sizes - the smallest fields of approx. 2x2cm are used for dental X-ray diagnostics, the largest formats of approx. 43x43cm for lung imaging, or 96x20cm on the spine. At X-ray imaging, films are stored in a special light-tight cassette, provided with metal marks and letters at the edge, which are projected onto the film during exposure, are visible after development and ensure the geometric orientation and identification of the image. In the dark chamber, they are then removed from the cassettes, special concentrated developers are used to develop them, providing high contrast and saturation of the blackening of the film; the process of developing, setting and drying is carried out in developing automats.
*) The photochemical sensitivity of films is relatively low for X-rays. To increase the sensitivity (and thus reduce the required amount of X-rays, reduce the radiation exposure of the patient), amplifying luminescent foils were attached to the film, now all this is abandoned.
   Overall, however, the use of films and the "wet process" is in decline, it is mostly already abandoned. The future belongs to electronic digital X-ray imaging (see below). In connection with this, in the case of modern digital devices, the difference between skiagraphy and sciascopy is largely blurred - in a computer system it is possible to choose whether the recording of a digital image will be static or dynamic.
¨ Skiascopy
Skiascopy or fluoroscopy is a continuous visual observation of an image of the transmitted X-rays, originally on the fluorescent screen ("shield"). Direct sciascopy was used very often in the past, but due to the high radiation exposure of the examining radiologist (and also the patient), it has already been abandoned. Indirect sciascopy is performed on devices equipped with an image intensifier and electronic image capture, more recently direct electronic digital image capture by flat panel (see below). This indirect sciascopy is now used to investigate of dynamic processes (coronary arteriography, transhepatic cholangiography, ...) and in interventional procedures where visual inspection is required - X-ray navigation of precise work performed inside the body *) - insertion of various probes and catheters, implantation of pacemakers, coronary angioplastics, insertion of vascular or uterorenal stents, ... etc. - see below "Subtraction angiography", Fig.3.2.7. In radiotherapy, it is the introduction of radiophores by afterloading during brachytherapy
(see §3.6, section "Brachyradiotherapy").
*) To reduce the radiation exposure, pulse mode is now used in sciascopic imaging : the X-ray does not glow continuously, but periodically turns on only for short moments (approx. 0.1 sec. with a repetition rate of approx. 4 frames/sec.), during which produces an image. To improve the visual quality of the images thus generated sequential frames use is sometimes so called recursive filter, consisting of the weighted summation of several consecutive images.
   X-ray tube and an opposing imaging detection system are often mounted on a special "C" shaped stand - called C-arm - Fig.3.2.7b. This arm can be rotated to different angles using electromotor, which allows quality display in various projections. Similar flexible possibilities of movements of the X-ray tube and the imaging detector around the patient are provided by the so-called U-arm (independent movement of the X-ray tube and the detector on the stand) - Fig.3.2.7c. Alternatively, so-called tiltinable walls can be used; these possibilities were mentioned above, the section "Mechanical design of X-ray devices".

Contrast agents. Subtraction radiography.
One of the main difficulties in soft tissue X-ray imaging is the small differences in the absorption of X-radiation by individual tissues *), leading to low image contrast and difficulty in distinguishing some structures.
*) The tissues of the human body are mainly formed by atoms of light elements (hydrogen, carbon, oxygen, nitrogen, sodium, ...) and have a similar density of just over 1 g/cm3. Therefore, the absorption coefficients for X-rays in individual tissues do not differ much (with the exception of more massive bones and lighter aerated lungs).
   In certain cases, the natural absorption differences between the tissues can be increased and thus the resulting contrast of the X-ray image can be improved by applying suitable contrast agents. Contrast agents artificially increase the contrast of tissue imaging by causing greater differences in the X-ray absorption of the examined tissue relative to the surrounding environment. Usually we try to increase the absorption of X-rays by using substances containing atoms of heavy elements such as barium
(cavities, eg stomach) or iodine (vessels, organs). X-rays are strongly absorbed by these substances, which negatively highlighting the cavities that are filled with them (stomach, digestive tract, blood vessels). If such a substance is introduced into the examined area - the gastrointestinal tract, blood vessels, bile or urinary tract, the structure filled in this way shows a significantly increased absorption of X-rays and is clearly and contrastily displayed on the X-ray image, including possible defects and anomalies. After application, contrast agents can enter into the organ under investigation either directly (direct application to the gastrointestinal tract or blood vessels), or indirectly via metabolism (imaging of structures in the liver or kidneys).
   Contrast agents are classified according to various criteria. According to water solubility: insoluble (barium suspension, iodine substances oil and suspension) and soluble (hydrosoluble). According to their ionization (dissociation) in solution: ionic (dissociate in solution into anion bearing contrast iodine and cation) and nonionic. According to pharmacokinetics, metabolism in the body and route of excretion: nephrotropic (excreted by the kidneys) used for angiography, urography, contrast CT; hepatotropic (excreted by the liver and bile) for cholangiography. In terms of the achieved change in X-ray absorption, we divide contrast agents into two basic types :
Positive contrast agents , which increase the absorption of X-rays. The most commonly used contrast agents are based on barium and iodine .
Barium sulphate
4) is a water-insoluble compound whose suspension ("barium slurry") is used in the examination of the digestive tract.
Currently, the most commonly used contrast agents are based on iodine, which has two advantageous properties :
127I shows high absorption for X-rays of the energies used in X-ray diagnostics - it provides a good positive contrast of the structures into which it enters.
2. Iodine can form compounds with a number of organic substances that behave in the body in the necessary well-defined way. In such an organic substance, the absorption properties for X-rays are given by the bound iodine atoms, while the other organic part of the molecule determines the pharmacokinetics and distribution of the substance in the organism - where the substance "gets" or where it is taken up and excreted.
   Iodine contrast agents are used in the form of organic compounds in which iodine is tightly bound, mostly in the benzene nuclei of cyclic (aromatic) hydrocarbons. It is mainly triiodaminobenzoic acid, where 3 iodine atoms are attached to the benzene nucleus
(1,3,5-triiodo-2-aminobenzoic acid derivatives, they are usually ionic and non-ionic). Furthermore, pyridine derivatives with one or two attached iodine atoms in the molecule are used (they usually have the character of ionic substances).
   Water-soluble (hydro-soluble) contrast agents, especially ionic ones, can cause some undesirable side effects in the body, allergic reactions can be dangerous.

Negative contrast agents reducing X-ray absorption. These are mainly gases (air, carbon dioxide) that are applied to the cavities (eg the spinal canal). Today, they are practically no longer used .
   In some cases, especially in the digestive tract, the so-called double contrast is used : first a positive contrast agent (barium suspension) is applied and then a negative contrast agent - air (from effervescent powder), which expands the positive contrast agent to the walls of the examined volume.
X-ray subtraction radiography. DSA
A special method of increasing the contrast is the so-called subtraction radiography, consisting in the subtraction of two images of the same area, differing in the presence and absence, or distribution, of the contrast agent. The goal of subtraction is to highlight anatomical structures that would be little clear, indistinct, and difficult to recognize on conventional X-rays images.
   In the early days of the method (50s and 60s), film (photographic) subtraction was used, in which an X-ray image with a contrast agent was combined and overlaid with a negatively displayed image without a contrast agent. This combination (masking) created the resulting subtraction image, in which only structures filled with contrast material are visible. Further technical development, leading through analogue television subtraction, has resulted in the method of digital subtraction, which is the most perfect and now used exclusively. This method is used mainly for the selective imaging of the arterial and venous vascular bed - the contrast agent is injected at the appropriate time using a specially inserted catheter. It is called digital subtraction angiography (DSA) for arterial bed or phlebography for venous imaging.

Fig.3.2.7. a) Principle scheme of digital subtraction radiography operation. b) X-ray tube with electronic image sensor mounted on a C-arm. c) X-ray device in U-arm arrangement.

A simplified diagram of the principle of digital subtraction radiography is drawn in Fig.3.2.7a. The X-ray beam from the X-ray tube illuminates the patient's body and the transmitted radiation is detected by a digital image sensor ( flat-panel), consisting of a scintillator and a sensitive CCD image sensor. The most perfect imaging detectors are SPD semiconductor pixel detectors (see §2.5 "Semiconductor detectors"), mounted in so-called flat panels described above in the section "Electronic X-ray imaging detectors", Fig.3.2.6 in the middle. The X-ray tube and the detector are placed opposite each other on the so-called C-arm (Fig. 3.2.7b). First, a native X-ray image of the examined area without a contrast agent (formerly called a mask) is scanned into the computer's memory, and then an X- ray image after application of a contrast agent. Numerical digital subtraction of the native image from the contrast-enhanced image subtracts and cancels out all structures that have not changed (eg skeleton) and remains only what makes the two images different: contrast-filled cavities and vessels. The resulting subtraction image is created, in which only the structure filled with contrast medium is selectively displayed, while all other anatomical structures are more or less cancels.
  Proper subtraction can be adversely affected or impaired by tissue movements during the examination (in the time interval between the two images), such as breathing movements, heart pulsation, patient movement. To eliminate these adverse effects, a number of images are recorded at short intervals, from which images suitable for subtraction are selected. In addition, to monitor the kinetics of cardiac activity, the sequence of scanned images is synchronized with the ECG signal and images corresponding to end-diastole and end-systole are subtracted; it is thus possible, among other things, to obtain an image of the ejection fraction and to reveal possible disorders of the heart wall motility.
Tomographic 3D digital rotational angiography 
If angiography is performed on the C-arm, then by rotating the system (X-ray tube - detector) around the patient, a con-beam CT - imaging can be acquired and reconstructed - tomographic 3D digital rotational angiography, with the possibility of flexible spatial assessment of vascular patency at different sites.

The historical name "angio-line" 
Frequently used name "angioline" comes from the time when angiography performed on X-ray films (60th-70th years). At that time, a series - line - of cassettes with X-ray films in a row had to be prepared for the skiagraphic device, which were exchanged and exposed in quick succession after the injection of the contrast medium. They were then developed and inspected for where the contrast agent had gradually reached - or not - due to the occlusion of the vessel.
X-ray navigated interventional procedures
Subtraction angiography was originally developed as a diagnostic method. With the help of modern angiographic equipment and advanced methods of vascular medicine, in addition to diagnostics, it is possible to immediately perform the necessary interventional performance under detailed control of X-ray imaging immediately after finding out the pathological conditions in the vascular bed. These are, for example, coronary angioplasty (PTCA) - dilation of the narrowed coronary artery of the myocardium using a special catheter equipped with a balloon at the end, with possible by installation of a so-called stent, which remains stretched inside the coronary vessel and prevents it from shrinking again.

X-ray tomography - CT
Classical X-ray image is planar - it is a two-dimensional projection of tissue density to a certain plane. However, real tissue is a three-dimensional object, so a planar image that is a two-dimensional projection of tissue, can capture only part of reality. We cannot find out anything from the planar image about the arrangement of the tissue in the "deep third dimension", perpendicular to the displayed plane. Planar images have serious pitfalls in this respect - the possibility of overlapping and superposition of structures stored at different depths. Although we help here by displaying in several different projections, but the risk of a false finding or non-detection of an anomaly in the depths of the organism, covered by another structure, can never be ruled out. In the planar display occurs shine trough the X-rays from different depths, the superposition and accumulating information on the distribution density from all depht layers of tissues and organs in a common image. The resulting response in the image is the sum of the contributions from the individual layers of tissue - not only from the sites of the examined lesion, but also from the layers located above the lesion and below the lesion. In this way, the details of the structure of the examined organ at a certain depth can be overlaped by pictorial information from more distant and closer layers. The individual tissues and organs are shown in summary on the planar image, they overlap. We are not always able to unambiguously determine which organs and structures the X-rays have gone through and been weakened by. Superposition of radiation from different depths of the imaged object further leads to a reduction in the contrast of the imaging of structures and lesions.
   To overcome these disadvantages of planar X-ray diagnostics and to obtain a complex imaging of structures at different depths, transmission X-ray tomography *) has been developed to provide a three-dimensional imaging of tissue density in an organism. One of the main advantages of tomographic imaging is significantly higher contrast imaging of lesions that do not overlap on radiation from surrounding layers on transverse sections.
*) Greek tomos = section - tomographic imaging consists of certain sections (slices), primarily transverse, a larger number of which creates a three-dimensional image. The examined area is divided into a large number of thin layers (transverse sections), which are each scanned separately at many different angles, and from the local attenuation of X-rays, the density image of the layer is mathematically reconstructed in a computer. We can then view the examined area on the computer screen in individual thin layers - as if we were "cutting" the patient transversely, looking inwards at each incision and then folding it again (without damage).
The forerunner of the current CT computed tomography was motion tomography : the X-ray tube and the examination table with the patient moved in opposite directions to each other in such a way that for a layer at a certain depth both movements were compensated and a sharp image was obtained, while in the other layers the image was motion blurred and al the less clear. However, the quality and contrast of such an image were not great (completely incomparable with CT), the method is long abandoned.
   Tomographic X-ray imaging is achieved by shine trough the examined area at a number of different angles (in the range 0-180-360°): the X-ray tube and the X-ray detector located opposite it circulate around the patient's body *), while a narrow beam of X-ray shine trough the examined tissue and its passed intensity is detected and converted into an electrical signal (Fig.3.2.8a); the attenuation of the beam due to tissue absorption is evaluated. From the larger number of individual integral values obtained by passed X-radiation under a series of angles 0-360°, the absorption map is then reconstructed by back projection, creating a cross-sectional density image of the examined area in a plane perpendicular to the X-ray and opposite detector rotation axis
- see "Density image formation" below. In this image, structures stored at different depths in the organism are sensitively and with high resolution are displayed - it is a tomographic image. In such an image, we have endless possibilities of viewing the scanned object from all angles and in various layers (sections).
*) The X-ray tube and the opposite detection system are mounted on a special annular stand called gantry (gantry = portal, through-supporting construction), enabling the X-ray tube - detector system to rotate around the examination bed by means of an electric motor.
   By gradual longitudinal linear displacement of the patient with respect to the X-ray tube - detector system, we can create a series of cross-sectional images (individual layers), which placed next to each other create a three-dimensional tomographic image of the examined area. Due to the computational complexity of the reconstruction procedure, this can only be done with the help of a computer - therefore this method is called computerized tomography CT (Computerized Tomography) or computed tomography. The exact name "X-ray Transmission Computerized Tomography" did not take hold due to its length.

Fig.3.2.8. X-ray computerized tomography CT.
a) Basic principal scheme of CT. b) Principle of spiral multi-slice CT. c) Example of 64-slice CT instrument.

In addition to spatial tomography imaging, the main advantage of CT compared to conventional X-ray imaging is significantly higher contrast - is able to recognize and display even slight differences in the linear coefficients of attenuation of X-ray, that penetrates trough the examined tissue. This is primarily due to the principle of transverse section imaging using a narrow beam without being affected by adjacent layers and electronic X-ray detection, which is able to capture finer differences and a wider range of dynamics than conventional X-ray film. Methods of computer reconstruction and image filtering, as well as the possibility of flexible setting of optimal image modulation (brightness, contrast) also contribute to the excellent density resolution. Computer software for CT also has a number of tools for structural image editing, creation of three-dimensional images of certain organs, reconstruction of sections in other planes than the initial transverse in which the patient was scanned.
   The result of CT are real "anatomical sections" of the patient's body, on which organs and tissues can be seen separately, in contrast to planar X-ray imaging, where they are shown in summary and overlap.
Before starting the actual diagnostic acquisition of CT, a trial, "exploratory" or "planning" planar radiographic acquisition CT, abbreviated SPR (Scan Projection Radiograph), is usually performed; it is also known as Scanogram or Topogram. It is scanned by a stationary (non-rotating) system of X-ray tube and detectors, mostly in AP or PA projection, in which the bed and patient move over the gantry. This creates a planar image similar to classical skiagraphy. This image is then used to determine the beginning and end of the displayed area of the body. Furthermore, the SPR image can be used for automatic exposure - obtaining absorption (attenuation) data for automatic regulation of the anode current [mA] by X-ray tube ATCM (Automatic Tube Current Modulation) in order to optimize the relationship between quality image and radiation exposure of the patient. This anatomical dose modulation technology significantly reduces the radiation dose in real time while maintaining the quality of the images.

Development of tomographic imaging method. 5 generations of CT devices.
General efforts to reconstruct a three-dimensional image based on a two-dimensional image (or a set of one-dimensional projections) date back to 1917, when J.Radon derived an integral transformation (now called the Radon transformation) between a set of line integrals and a set of transverse section points. In 1963, A.Cormack applied these results and extended them to the case of X-rays passing through partial absorption by a three-dimensional object. And in 1972,
G.N.Hounsfield completed the development of the first CT instrument. 
   In the following years, the great advantages of CT were proven and these devices became very widespread. During the technical development, there were also significant changes in the design of individual electronic and mechanical parts of CT devices. In view of this technical development, CT devices are usually divided into five generations :
¨ 1st generation: X-rays from the X-ray tube were collimated into a thin beam (cylindrical "pencil" shape) and after irradiation and passing trough the patient it is detected by the opposite detector (as shown in Fig. 3.2.4a), rotating together with the X-ray tube.
2nd generation: X-rays from the X-ray tube are collimated into a fan shape and after passing through the patient it is detected by a larger number of detectors, placed in a row on a circular section opposite the X-ray tube, rotating together with the X - ray tube.
¨ 3rd generation: X-rays from the X-ray tube are collimated into the shape of a wider fan similar to the 2nd generation, but the transmitted radiation is detected by a large number of detectors placed on a circular arc in several rows (Fig.3.2.8b) - more slices - multi-slice CT. The continuation of the 3rd generation devices are the spiral high-speed multidetector MDCT systems described below .
¨ 4th generation: the detectors are arranged stationary in a complete circle (several rings lying next to each other) around the patient, while only the X-ray tube rotates.
¨ 5th generation: cardio-tomograph with electron beam - EBT - Electron Beam CT , described below, Fig.3.2.10.
   In the end, the 4th generation devices did not become very widespread, because at a higher price they do not bring significant benefits for clinical practice compared to modern design solutions for generation 3. devices (high-speed multidetector systems MDCT, see below). And the 5th generation, electron beam CT, due to its enormous complexity and cost, did not penetrate into clinical practice at all, it remained only as a technical interest...
   Along with the technical improvement of X-ray CT, the tomographic principle was also used in other imaging modalities. In addition to optical CT, scintigraphic tomographic methods were developed - SPECT single photon emission tomography and PET positron emission tomography (Chapter 4 "Scintigraphy", §4.3 "Tomographic cameras". And also the most complex tomographic imaging method of nuclear magnetic resonance NMRI (§3.4, part "Nuclear magnetic resonance").
   Note: For special technical purposes of imaging the structure of small objects, is used the so-called X-ray micro-tomography (mCT) mentioned below (§3.3, section "Radiation defectoscopy").

Formation of the density image
If, according to Fig.3.2.4 on the left, the X-ray beam emitted by the X-ray tube and falling on the examined area has an initial intensity (photon fluence in 1 s.) I
o , then its intensity I after tissue passage will be I = Io . e - Sm(i, j). Dx , where m(i, j) is the linear attenuation factor of X-radiation penetrating the tissue site at coordinates (i, j) and Dx is the magnitude (length in the beam direction) of the tissue element. The values of the coefficients m(i, j) depend on the local density and the proton number of the individual sites (i, j) of the tissue. By logarithm, this relationship can be adjusted to the form: ln (I/Io) = Sm(i, j). Dx, which states that the logarithm of the ratio of the intensities of X-rays entering and leaving - transiting the examined tissue, is equal to the sum of the products of the linear attenuation coefficients m and Dx paths that the X-rays photons at each point of the tissue overcome.
   By measuring at different positions (angles) of the X-ray tube and the detector, a number of values of the attenuation ratio ln (I/I
o) are obtained. The computer then basically solves a system of linear equations of the above shape, which obtains the values of the linear X-ray attenuation coefficients of tissue elements in individual sites (i, j) of tissue - a picture of tissue density in a transverse section is formed.
   In practice, the above straightforward procedure does not progressing. The resulting transverse CT image is obtained by computer reconstruction from one-dimensional profiles of the intensity distribution of the transiting X-ray beam when rotating the X-ray tube and opposite detectors around the examined object. For this reconstruction, the method of filtered back projection is usually used, sometimes even a more perfect (but computationally demanding) method of iterative reconstruction *).
*) Iterative reconstruction is a mathematical algorithm for finding the most accurate image by successive approximations and refinement of the initial "rough" image. It is based on an image obtained by back projection and consists of several repetitions four consecutive steps - "reconstruction loops": 1. In the first step, they are calculated from the image created by the back projection, in the opposite way, again as if the original data. 2. These are then compared with the "raw" data actually captured during detection. 3. New data for rear projection will be adjusted according to the differences found. 4. The image is recalculated and the whole operation is repeated. After several repetitions of these iterations, we obtain a clear and high-quality image. The number of iterations can be set and optimized. Correction algorithms and noise filtering methods are included in the procedure. This method provides higher quality images with reduction of noise and artifacts created during reconstruction. Furthermore, in some examinations, we can significantly reduce the patient's exposure and radiation dose, while maintaining sufficiently high-quality images.
   These reconstruction methods, which are analogous to SPECT, are briefly described in §4.3 " Tomographic scintigraphy ", passage "Computer reconstruction of SPECT " *). The central and controling software algorithm of the reconstruczion is sometimes called the kernel (core, grain).
*) I apologize to professional radiologists that it is not included in this chapter. Our professional materials are primarily focused on nuclear physics and radionuclide scintigraphy (I didn't want to duplicate it...). This angle of view may be perhaps inspiring also for colleagues in the field of X-ray diagnostics ..?..
   The density of the examined tissue is usually compared with the density of water and in the CT image is numerically presented in the so-called Hounsfield units HU = 1000. (mtissue - mwater) / mwater , introduced by a leading pioneer in the CT, G.N.Hounstfield (along with A.L.Cormack). The use of a factor of 1000 (instead of the usual 100%, where we would get decimal values) reflects a high density resolution CT. The value of HU = -1000 corresponds to zero density (vacuum, air), for water it is HU = 0, bones have a density of the order of HU = 100 ¸ 1000, sometimes even higher. Aerated lungs have HU approx. -800, fat HU = -40 ¸ -120, soft tissue density is HU = 20 ¸ 80. Such a large range of densities is not able to display linearly in brightness; also the human eye is only able to distinguish a few tens of shades of gray. For optimal image presentation, we therefore help by appropriate modulation of image brightness and contrast. If we are interested in differences in tissues with a similar density (this is usually the case in soft tissues), we use this modulation to select only a narrow part of the whole range of densities - the so-called window, whose range of densities is displayed in the entire brightness range of the screen, with eventually increased contrast setting. We get well-drawn images of the required structures and by moving the windows we can gradually obtain detailed information about tissues with different densities.

X-rays detectors for CT
The task of these detectors is to capture photons of X-rays passing through the examined tissue and convert them into electrical signals for further electronic processing for the purpose of computer reconstruction of density sections. In general, the principles set out in Chapter 2 "Detection and spectrometry of ionizing radiation" apply to this area, with the proviso that mostly detection is applied, rarely the spectrometry. The basic requirement here is a high sensitivity of X-ray photon detection and a high detection rate, ie a short dead time.
   In principle, three types of detectors can be used to detect X-rays in CT :

X-ray detectors for CT
(these are only basic schematic drawings, there is no drawn the curvature of the arc arrangement, or lamellas against scattered radiation).
a) X-ray scintillation detection, the most used method. b) Dual-layer detector registering low and high X-ray energy separately. c) System of semiconductor pixel X-rays detectors - photon-couting CT, enabling spectrometric analysis of X-rays; here are the individual detectors. d) Spectrometric photon-couting CT in the technology of tape semiconductor detectors with a common cathode and pixelated anodes.

Multidetector, multi-slice and spiral CT; Cone-Beam CT
Gradual scanning of CT images by a system of one X-ray tube and one detector, as described so far here according to Fig. 3.2.4a for easier explanation of the method, was used in the first generations of CT equipment in the 70s and 80s. Its disadvantage was considerable slowness (one cut lasted several minutes). Newer generations of CT devices already use a larger number of detectors (approx. 1000). An X-ray tube is circled, the beam of primary radiation of which is obscured by a collimator into the shape of a fan (with an angle of approx. 40°), and opposite it a corresponding circular section with a system of 300-1000 detectors (Fig.3.2.8b). The scanning time of one slice is reduced to less than 1 second.
  The first types of CT devices had a rotating part - X-ray tube and detectors - connected to the static power supply and evaluation part by a cable, which did not allow continuous rotation (after one revolution, the X-ray tube with the detector had to return to its initial position, or the direction of rotation had to be changed so that the cable did not twist). Newer types (since the 1980s) use "slip-ring" technology with electric brush scanning to power and transmit the signal, enabling fast and continuous rotation (with an unlimited number of revolutions in the same direction).
   The original generation of CT instruments scanned only one cross section of the examined area during one rotation. To increase the speed of CT examination of larger areas, it is always used in newer generations of devices several detectors, resp. several detector rings, placed side by side in the axial (longitudinal) direction - MDCT (Multi Detector CT). This allows (with a suitable shaping of the X-ray beam from the X-ray tube) the simultaneous scanning of several transverse sections side by side, the examination of several thin layers simultaneously. We are talking about so-called multi-slice CT devices - 4, 6, 8, 16, 64 and more - slice. The technical design of CT devices is constantly improving. The number of detectors and the speed of rotation of the gantry rotor increases (now approx. 0.3 s./revolution) - these are high-speed multidetector systems MDCT. Individual transverse sections can be scanned in two ways :
- Sequential scanning , where only the X-ray tube + detector system rotates, but the bed does not move with the patient. The individual layers are scanned gradually - independently by individual rings of detectors.
- In the case of so-called spiral CT (helical) , in addition, during the rotation of the X-ray tube there is a slow automatic movement of the bed with the patient (the X-ray tube path effectively appears as a spiral) - Fig.3.2.8b, c, followed by three-dimensional reconstruction; here, in principle, it is possible to achieve whole-body CT imaging. The horizontal distance by which the lounger moves between two adjacent X-ray tube cycles - the "rise" of the spiral - is called pitch factor [mm].
ECG-synchroized CT angiography

A significant technical advance in the field of cardiac imaging is non-invasive electrocardiographically synchronized angiography by multidetector computed tomography - MSCTA. The main use of this method is twofold :
MSCT native calcium score - detection and quantification of calcificates in coronary arteries (calcium content in coronary artery plaques). Usually, risk groups are assessed according to Agaston's calcium score (up to 5 groups).
MSCT coronarography - contrast imaging of epicardial coronary arteries to diagnose the extent and severity of coronary involvement in ischemic heart disease.
   By combining both methods, we can display the soft and calcified part of the plaque, determine the nature, extent and severity of the disease, with possible indications for invasive examination with determination of the method of revascularization.
Cone-Beam CT
For some purposes, a CT image with a widely collimated cone-beam CT (CBCT ) is used, which shines through the patient and impinges on the opposite flat-panel imaging (its principle is described in §3.2, passage "Electronic X-ray imaging"). The X-ray tube and the opposite flat-panel rotate around the examined object on a common gantry. Such a CBCT system is installed on radiotherapy irradiators (linear accelerators) using the image-controlled radiotherapy technique - IGRT (§3.6, section "Modulation of irradiation beams"), or is used in small dental CT devices (listed below).
CT with 2 X-rays tubes - DSCT: Dual Source and Dual Energy CT
Another technical improvement of CT consists in the construction of devices that have 2 X-ray tubes - two X-ray tube/detector systems (placed perpendicular to each other), which can scan simultaneously, Fig.3.2.9. The device is referred to as Dual Source CT (DSCT). It can work in two basic modes, providing two advantages :

1. Both X-ray tubes operate at the same voltage
Þ "dual system" - increasing the speed and shortening the acquisition time with reducing the time resolution to about 80ms. This is especially important for CT of the heart (with a higher heart rate).

Fig.3.2.9. CT device with two X-ray tubes - Dual Source CT and Dual Energy CT.

¨ 2. Booth X-ray tubes operate at different anode voltage (eg. 140kV and 80kV **)
Þ possibility of scanning with dual-energy (DECT - Dual Energy CT): each of the two X-ray tubes creates X-rays of different energies. We get twoo different density images of the same place. This allows not only to better quantify the density distribution, but also to determine the tissue composition using the method of differential density analysis DEXA (Dual Energy X-ray Absorptiometry) *) - similar analyzes of density images as in "Bone densitometry" (see Fig.3.2.11 below). It provides not only detailed images of the anatomy, but it will allow you to distinguish different types of tissue (distinguish eg bones, blood vessels, adipose tissue), different types of kidney stones, deposition of sodium urate crystals in joints (bottoms), or quantify the distribution of contrast agent in myocardial infarction (and to assess functional impairment in morphological coronary artery disease).
*) Different types of substances (and tissues) differ not only by specific values of linear attenuation coefficients m for X radiation of a certain energy, but also by a somewhat different dependence m(EX) of absorption for different energies EX of X-radiation. This is due to the different electron density configuration for the different molecular composition of the analyte. Mathematical analysis of the exponential laws of absorption I = Io.e-m(Ex).d for individual energies EX and tissue types with absorption coefficients m(EX) (by logarithm the relevant exponential equations are converted to linear) it is possible to determine the proportion of absorption in different tissues. This can in principle be used to additionally distinguish different types and compositions of tissues based on differences in the density images of the same site, obtained with different X-ray energies.
**) X-ray spectra for 80 and 140 keV are continuous and partially overlap. In addition to the different anode voltages, two different effective energies of X-rays are also achieved by special sharp filtration using the K-edge effect (mentioned above in the section "Filtration, collimation").
Multiplex DECT
An alternative to dual energy tomography DECT with two X-ray tubes is the use of a single X-ray tube
- detector system, in which performs the multiplex switching voltage in X-ray tube during spiral scanning.
2-layer CT detector
Other alternative options DECT tomography with one X-ray tube is the use of two layers of detectors, that selectively detect the low- and high-energy X-ray component (mentioned above in the section "
X-ray detectors for CT", section "Two-layer CT detectors").
Spectrometric Photon-couting CT
Furthermore, it is likely that the entire DECT technology with two X-ray tubes will be replaced in the future by a photon-couting CT system with one X-ray tube and semiconductor detectors, enabling flexible energy spectrometry of detected X-rays (described above in the section "
Electronic X-ray imaging detectors", paragraphs "Semiconductor detection (direct conversion)" and "X-ray detectors for CT").

Quality control and imaging properties of CT devices
Testing of imaging properties of CT devices is performed using special phantoms of cylindrical shapes
- see "Phantoms and phantom measurements", section "Tomographic phantoms for CT".

Electron Beam CT (EBT)
In addition to the described CT design, now "classic" with a rotating X-ray tube, a completely different, physically interesting solution has been developed that does not contain an X-ray tube at all. X-rays are created by the impact of fast electrons, fired by an "electron gun", on a metal target ring - anode, inside which the object under investigation is located (Fig.3.2.10). The electron beam from the electron gun is directed to the desired location of the target ring by magnetic deflection by means of deflection coils, powered by a suitable electrical signal. By supplying the deflection coils with alternating electric current of a suitable periodic course, the electron beam rotates at an angular frequency
w and, during this circular motion, gradually strikes individual points on the circumference of the target ring. In each affected area, braking X-rays are generated , the beam of which shines through the examined object (patient's body) at a corresponding angle. Thus, a rotating electron beam generates a rotating source of X-rays around the circumference of the target ring, as if an X-ray tube were rotating there. The braking X-ray passes through an annular collimator with radially oriented septa, which shapes it into a fan-shaped bundle.
   This X-ray, passed through the examined object (patient's tissue), is detected electronically (as with conventional CT) by means of an annular array of detectors, overlapping the collimator from the inside. With a suitable geometric arrangement of the collimator septum, they shield the X-radiation that would come directly from the target ring to the back of the individual detectors. Newer types of EBT devices have several side-by-side target rings and several ring arrays of detectors.

Fig.3.2.10. Basic principle scheme of X-ray tomography using electron beams

This design solution has two advantages :
¨ It does not contain any mechanically moving parts - the rotation of the beam is electromagnetic.
¨ Allows very fast tomography - the electromagnetically deflected beam can rotate much faster than is achievable with mechanical rotation. This is advantageous for monitoring fast processes such as gated CT - in Fig. 3.2.5, ECG trigger pulses are fed to the acquisition computer together with pulses from X-ray detectors.
  However, the disadvantage here is the considerable complexity and cost (price) of the device, due to which this type of device has not yet become very widespread in practice. It is probably not possible to expect a greater expansion of these systems in the future either, as rapid technical progress in the construction of conventional CT - high-speed multidetector MDCT systems (or with two X-rays) solves most of the advantages of EBT, are cheaper and more advantageous for common practice. The result is that electron beam CT, due to its enormous complexity and cost, did not penetrate into clinical practice at all, it remained only as a technical interest...

X-ray bone densitometry
Radiographic examination of the skeleton is one of the most common and most important X-ray diagnostics, especially the finding of fractures and other bone defects. A special method in this area is bone densitometry - a method for determining the density of bone tissue based on the degree of X-ray absorption, determined by X-ray absorption photometry (Radiographic Absorptiometry - RA).
    The simplest method is to transmission a narrow beam of X radiation with a single energy (SPA - Single Photon Absorptiometry). The disadvantage of this method is that it is not possible to determine from the total absorption of X-rays, which part is caused by bone and which part by soft tissue.
    A more advanced densitometric method is X-ray absorption photometry using two energies of the X-ray beam (DEXA - Dual Energy X-ray Absorptiometry), such as a pair of effective energies of 50keV + 100keV, or 35keV + 75keV. Here are used the different ratios of X-ray absorption in soft tissue and in bone at low energy and at high radiation energy - different values of linear attenuation coefficients
m. By mathematical analysis of the exponential laws of absorption I = Io.e-m . x for individual energies and tissue types (by logarithmization the relevant exponential equations are converted to linear) the proportion of absorption in soft tissue and bone itself is determined, from which (after appropriate callibration) can be determined bone density.
  It is calibrated with a suitable bone phantom (or hydroxyapatite), the instruments have their internal calibration phantoms. The bone mineral content is quantified using the Bone Mineral Content (BMC) parameters in [g/cm] and the Bone Mineral Density (BMD) area density in [g/cm2]. These parameters are compared with sets of reference (normal) values and relative indices (ratios) called scores are determined: the T-score compares the measured BMD values with the average BMD value of young healthy adults of the same sex; the Z-score compares BMD with mean normal values for a given age and sex. Bone Homogeneity Index (BHI) is also sometimes monitored.

Fig.3.2.11. Schematic diagram of a DEXA imaging digital X-ray densitometer. On the right is an example of a DEXA device for whole-body imaging.

Modern X-ray densitometric devices use irradiation of the examined area with a diverging ("pyramidal") X-ray beam with subsequent detection of the transmitted radiation by a digital image sensor into the computer's memory - Fig.3.2.11. Here, appropriate absorption calculations are performed between the low (L) and high (H) X-ray energy images to obtain a final skeletal density image that provides both bone mineral content and density and morphological skeletal structure information. The most advanced devices of this type make it possible to perform whole-body image diagnostics of bone tissue, determine the content of muscle mass, adipose tissue, water and minerals in individual parts of the body.
Detectors for X-ray densitometry 
For the detection of transmitted X-rays, either NaI(Tl) or CaWO4 scintillation detectors are used, or semiconductor detectors mostly based on CdZnTe (cadmium-zinc-telluride - CZT), which have a high detection efficiency. In modern imaging densitometers, the detectors are arranged in a 2-dimensional mosaic configuration with high spatial resolution - they form a digital X-ray image sensor, such as a flat panel measuring 20 × 20cm and a matrix of 512 × 512 elements that scans the entire displayed area during a single exposure.
    Bone densitometry plays a key role in the diagnosis of osteoporosis - pathological reductions in mineral and organic bone mass, leading to a weakening of bone strength. Osteoporosis is one of the most common disorders of bone metabolism and is one of the most common causes of fractures in the elderly, especially in postmenopausal women. Early diagnosis of incipient osteoporosis (osteopenia) is important for the use of effective treatment to slow or stop osteoporosis before irreversible disorders in the bone structure.
Note: Non-radiation methods of bone densitometry are also used. Ultrasound densitometry determines bone density based on the attenuation of the sound signal and the speed of its propagation in tissue.

X-ray mammography
Another important specialized method of X-ray imaging is mammography - imaging of possible inhomogeneities and areas of increased tissue density in a woman's breast, which could indicate a tumor process. In order to achieve the best possible image contrast and resolution of the smallest possible lesions, it is necessary to compress the breast using a compression plate *) and shine trough the tissue thus formed into a layer about 7 cm thick with soft X-rays with an energy of about 20 keV. Low-energy X-ray photons interact with tissue atoms primarily through the photoeffect, which provides a higher absorption contrast between tissues with small differences in density. Due to this low energy, a special X-ray tube with a molybdenum anode and a beryllium output window, focus size 0.1-0.3 mm, is used in the mammograph. A molybdenum or rhodium filter is used to filtration the X-ray beam, which cuts off photons higher energies than about 20keV
(K-edge Mo) or 23keV (K-edge Rh) - the so-called K-edge effect mentioned above in the passage "Filtration, collimation" is used.
*) Note: Compression of breast tissue also leads to one minor advantage in terms of radiation protection: compression temporarily restricts blood flow and partially causes hypoxia of tissue cells. This somewhat reduces the radiobiological effect of X-rays, as hypoxic cells are less sensitive to radiation ("oxygen effect"- §5.2 "Biological effects of ionizing radiation", part "LQ model").
    Imaging was performed with a cassette with X-ray film equipped with an image intensifier, or more recently using electronic image capture - a semiconductor flat panel with direct image digitization. A secondary Bucky diaphragm is placed between the imaged tissue and the film or imaging detector to reduce the proportion of scattered radiation, reducing the contrast of the image. The resulting X-ray image of the breast is called a mammogram or mastogram. Under suitable circumstances, it is possible to detect a tumor as small as about 4 millimeters. Mammography is suitable not only for the examination of women with symptoms or suspicion of breast cancer, but also for screening - searching for early stages of breast cancer.

Fig.3.2.12. X-ray mammography.
  The breast is inserted between the X-ray tube and the imaging flat panel and compressed with a plastic compression plate.
  On the X-ray mammographic image, we can observe normal density without defects
(top right) , or lesions of increased density that may indicate a tumor (bottom right) .

The X-ray mammography apparatus can be supplemented by a so-called mammographic stereotaxy device, which captures two images of a given lesion in oblique projections at two given angles (usually ± 15°). Evaluation of the change in the position of the lesion on these two stereo images allows precise targeting of the displayed structures suspected of the tumor process - their location and marking with a suitable marker (such as mandren, wire or dye), with the possibility of sampling by biopsy for histological examination.
Alternative mammography methods
In addition to the most commonly used X-ray mammography, there are some other examination methods based on different principles :

Ultrasonic mammography showing possible lesions based on their different density and elasticity (ultrasonography is briefly discussed in §4.6, passage "Ultrasound sonography").
¨ NMRI mammography - imaging by nuclear magnetic resonance.
¨ Radioisotope scintimammography showing increased accumulation of a suitable radiopharmaceutical in the tumor tissue (see Chapter 4 "Scintigraphy"); it can be performed as planar, SPECT or PET scintigraphy. A specific method of PEM positron emission tomography is described in §4.3, passage "Positron emission mammography (PEM)".
¨ Electroimpedance mammography sensing the electrical conductivity (impedance) of mammary gland tissues. A weak electric current is introduced into the tissue by means of electrodes placed on the skin in the vicinity of the examined area, and also by means of electrodes the distribution of electric potentials on the surface is sensed. From these data it is possible to reconstruct the spatial distribution of local tissue impedance - electroimpedance image. A different electrical conductivity is observed in the tumor tissue from the surrounding tissue.

Dental X-ray diagnostics
A separate category of smaller specialized X-ray devices are dental X-rays devices used in dentistry. There are three types of devices :
¨ Intraoral X-ray is a very simple device: a small X-ray tube of a compact design with a narrow tube is placed on the movable arm (a high voltage source of approx. 50 - 70 kV is usually encapsulated in X-ray tube cover). A small rectangle of X-ray film is placed on the back of the teeth and exposed to X-rays from the front. After developing the film, the relevant tooth (or several teeth) and its placement in the gums, including posiible defects, are displayed.
Panoramic X-ray OPG (orthopantomogram), also called DPR (Dental Panoramatic Radiograph ). The X-ray tube and the X-ray film or imaging flat panel located opposite, it rotate (orbiting) during the exposure and describe a circular or elliptical trajectory around the patient's head (jaw) so that the focal layer, given by compensating of mutual movement of X-ray tube and film (or imaging detector), is passed through the center of the jaw line along its entire length. This creates a developed panoramic image of the entire jaw. In the simplest case, this is achieved with a single motor, but a better panoramic view is achieved with devices with 2 or 3 motors, which perform rotation in multiple axes with better adaptation to the shape of the jaw. The best focusing of all jaw areas is achieved with multifocals OPG devices, where a larger number of images in different focal layers are captured, with subsequent evaluation.
¨ Dental CT is a reduced version of the classic cone-beam CT (mentioned above) with imaging detectors of about 3 x 4 - 15 x 15 cm. Sometimes panoramic X-rays are combined with dental CT into one system. Dental CT is used in maxillofacial surgery, dental implants and in the diagnosis of pathological lesions in the area.

Radiation exposure of patients during X-ray examination
In X-ray examinations, the external source of radiation is an X-ray tube, after passing through the patient's body, X-rays fall on a detector (previously a film, now a flat panel or a ring of detectors in CT), the image is created by different absorption of radiation in the tissues. The transmitted part of the X-radiation creates an image, the absorbed part of the radiation causes the radiation load in the body. The patient is irradiated only for the duration of scanning (acquisition) in the required projections. The greater the number and size of scanned projections, the intensity of the X-radiation used and the exposure time, the higher the patient's radiation exposure.
    Using the current modern X-ray technique, the radiation exposure of patients is generally relatively low. The high sensitivity of imaging detectors and the digital computer image processing enables optimization, in which high-quality X-ray images can be obtained with low radiation exposure.
    The absorbed radiation dose D [mGy] during X-ray examination of a certain area of the body is basically given by the product of the intensity of X-rays (this is given by the X-ray tube current [mA]), exposure time [s] and corresponding coefficients :
            D = G. mAs .
Coefficient G it includes a number of factors, such as the efficiency of X-ray production by X-ray tube, its energy given by the voltage [kV] in X-ray tube, filtration, distance, tissue absorption coefficients. It is measured using phantoms, most often water-filled "aquariums" (for planar X-rays), or cylinders with a diameter of 16 cm (head) or 32 cm (chest) for CT, equipped with ionization chambers, thermoluminescence or semiconductor detectors. The probability of biological stochastic effects is proportional to this absorbed radiation dose [mGy] and the size of the irradiated area [cm
    In planar X-ray diagnostics, this is quantified using the the quantity of surface dose DAP (Dose Area Product) [
2], which is the product of the input dose of X-rays and the size (area S) of the irradiated field: DAP = D. S. The effective dose Def [mSv] for the patient, expressing the effects of radiation on the organism as a whole, is then calculated as the product: Def = EDAP . DAP, where the coefficient EDAP (regionally normalized effective dose [mSv mGy-1 cm-2]) includes averaged tissue (organ) weighting factors wT for structures in the irradiated area.
    Specially calibrated radiometers, so-called DAP-meters, are used to measure the radiation dose of patients during planar X-ray imaging *), measuring product of absorbed dose and irradiated area (Dose Area Product). The DAP meter is a thin transmission plane-parallel ionization chamber mounted on the output collimator (aperture) of the X-ray tube, the area of which covers the entire (maximum) field of X-rays -
Fig .3.2.5 at the top left. The ionization current generated by the passage of X-rays from the X-ray tube toward the patient, after appropriate calibration, indicates the areal dose of DAP that the defined imaged field of the patient's body receives.
*) Radiometers of this type are sometimes also called KAP-meters , measuring the product of kerma in the air and the irradiated area (Kerma Area Product). For X-rays used in X-ray diagnostics, the kerma and dose values are practically identical. As discussed in §5.1 "
Effects of radiation on matter. Basic quantities of dosimetry .", passage "Exposure, kerma, terma", the relationship between dose and kerma D = K.(1-g) applies, where g is the fraction of energy of released charged particles that are lost during radiation processes in the material. For diagnostic X-rays, the value of g is only fractions of a percent.
    The product of the kerma and the surface is the integral of the kerma in the air over the beam surface in a plane perpendicular to the central axis of the beam. The value of the product of the kerma and the area is almost independent of the distance from the focus of the X-ray tube (if we neglect the attenuation of the radiation in the air, the backscattered radiation and possibly the X radiation arising outside the focus).
    For CT imaging, the radiation exposure is quantified using a quantity linear dose DLP (Dose Length Product) [], which is the product of the absorbed dose D and the length L irradiated area: D = DLP . L.
    The effective dose D
ef [mSv] for an X-rayed patient, expressing the stochastic effects of radiation on the organism as a whole, is then calculated as the product:
ef = EDAP . DAP for planar display, or Def = EDLP . DLP for CT imaging,
where coefficients E
DAP or EDLP - regionally normalized effective dose [mSv mGy-1 cm-1 ] - include averaged tissue (organ) weighting factors wT for structures in the irradiated area (§5.1 "Basic quantities of dosimetry", passage "Radiobiological efficiency of radiation") .
    How these measured exposure parameters are used to determine radiation doses and their optimization in X-ray examinations is briefly discussed in §5.7 "Radiation exposure in radiation diagnosis and therapy".

Other imaging diagnostic methods. Hybrid imaging systems.
X-ray diagnostics is the oldest and so far the most important imaging method in medicine. With technical progress, especially in the field of electronics and computer technology, some other alternative imaging methods of medical diagnostics have developed. They are:
Ultrasound sonography , Nuclear magnetic resonance , Thermography ; recently, electroimpedance imaging of tissue has begun to be applied. The whole chapter 4 then describes in detail another important imaging method - Scintigraphy.
    These methods are physically compared in §4.6 "
Relationship between scintigraphy and other imaging methods", where their diagnostic benefits and their complementarity in the algorithm of complex diagnostics are discussed. In recent years, hybrid imaging systems have been increasingly emerging, combining CT X-ray imaging with scintigraphic SPECT or PET imaging, or with nuclear magnetic resonance NMRI. And also hybrid imaging + irradiation technologies in radiotherapy (§3.6 "Radiotherapy", part "Modulation of irradiation beams", passage "Hybrid integration of imaging and irradiation technologies").

The basic aspects of evaluation and acquisition of diagnostic information from images of the mentioned radiological modalities are given in §4.7 "Visual evaluation and mathematical analysis of diagnostic images".

------------------------ Small physical-technical interest ---------------------- -

X-ray telescopes
The X-ray diagnostics discussed above is a transmission method based on the passage of X-rays through an object. X-radiation are a means of analyzing the structure of an object under investigation. However, there are situations where we primarily need to search for and display the sources of X-rays themselves, their position and distribution in space. And "at a distance", whether small (analytical and examination methods of materials) or large (detection of X-rays in universe) - to perform X-ray telescopy.
     Telescopes working in the field of X-radiation must have completely different optics than normal visible light telescopes. In the optical telescope, spherical or parabolic lenses and mirrors are used, on which light rays fall at a large angle (almost perpendicular) and refract or reflect so that they converge and intersect near the focus where the image is formed. The lenses are not usable for X-rays at all. Under certain circumstances, reflection on a mirror is applicable, but the beam must strike the reflecting surface of the mirror at a very small angle, almost tangentially. Due to the high energy of the photons and the short wavelength of the X-radiation, the X-rays must fall very obliquely, practically "sliding" on the surface, because when incident at larger angles (or even perpendicular), the X-rays photons would penetrate below the mirror surface and interact individually with electrons and atoms of its material (photoeffect or Compton scattering, see §1.6 "
Interaction of gamma and X-rays") - part would pass, most would be absorbed in it; no display would be created. At a very oblique impact, from the point of view of the photon, the number of free electrons in the metal surface per unit length will increase geometrically (the electron density will increase effectively), so that the X-ray photon will interact collectively with a large number of free electrons, similar to the reflection of a light wave from a metal surface: according to the laws of electrodynamics, the electromagnetic wave-photon is reflected from the metal surface at the same angle as the angle of incidence. Or, from the point of view of the reflecting surface, the wavelength of the incoming radiation (its projection) is effectively extended, which will therefore behave similarly to light. It's a bit like throwing a stone very obliquely, almost parallel to the water surface and he bounces, or it jumps above the surface several times. However, in practice this mechanism only works for soft X-rays, up to about 30keV.

Fig.3.2.13. Principle of a mirror X-ray telescope

X-ray optics are therefore based on an almost tangential impact, where the X-ray beam impinges almost parallel to the surface, only then is it reflected. Such reflective surfaces must be very precise and smooth, their "roughness" must not exceed a thousandth of a mm. The X-ray telescope consists of one or more very precisely shaped coaxial metal surfaces, inclined at a very small angle with respect to the optical axis of the system (Fig.3.2.13). This reflecting surface may have a conical shape, but the combination of paraboloid and hyperboloid provides optimal optical properties. The reflecting surfaces are arranged almost parallel to the incident rays, which are therefore reflected at a very small angle - first from the paraboloid and then from the hyperboloid - to the focal plane, where they form an image of the X-ray source from which the X-rays came. Here the radiation is sensed by detectors. The more advanced types of X-ray telescopes are multi-mirror, consisting of a number of very precisely shaped, carefully adjusted and embended in each other coaxial parabolic and hyperbolic mirrors. The central part of the system is shielded by an X-ray absorbing material.
    X-ray telephoto of this kind were constructed in the 1960s and 1990s mainly by R.Giaccomi and his collaborators. They constantly improved them and installed them on space satellites: UHURU in 1970, HEAO-2, Chandra in 1999. With these instruments was revealed many X-ray sources in space - X-ray binaries, supernova remnants, neutron stars, galaxies active nuclei, the clouds of ionized gas in galaxy clusters
(on the X-radiation from space and its origin see §1.6 , the "Cosmic ray" passage "Cosmic X and gamma rays"). Current X-ray telescopes achieve very good angular resolution (<0.5 arcseconds), spectral (energy) resolution (1eV), as well as high luminosity and sensitivity to X-rays also higher energies. They are the basic tools of the so-called X-ray astronomy.
    In the area of even shorter wavelengths, ie higher photon energies, X-ray telescopes are followed by the issue of gamma-telescopes, briefly discussed at the end of §4.2, section "
High-energy gamma cameras").

3.3. Radiation measurement of mechanical properties of materials
The properties of the interaction of different types of ionizing radiation with matter provide a number of possibilities for non-contact non-destructive measurement of some mechanical and structural properties of various objects and materials. Most of these methods work in the experimental setup schematically shown in Fig.3.1.1a,b in §3.1. The analyzed object or sample is irradiated with a suitable type of radiation, whereas the detector measuring the changings of the primary radiation caused by the analyzed object.

Radiation measurement of thickness and density
If we shine through a material with a given value of the linear attenuation coefficient
m, the absorption and attenuation of radiation is exponentially dependent on the thickness of the material - see the section "Absorption of radiation in substances" in §1.6 "Ionizing radiation". By measuring the radiation absorption, the thickness of the material and its changes can be determined without contact (in the basic arrangement according to Fig.2.8.1 on the left). For thin light materials such as paper or plastic foils, beta radiation is suitable, the source can be radionuclides 90Sr + 90Y (harder radiation b Eb= 0.546 + 2.27MeV, mass half-thickness of absorption d1/2 = 90mg/cm2, suitable for thicker layers), 85Kr (Eb = 670keV, d1/2 = 23mg/cm2 ), 147Pm (Eb = 224keV, d1/2 = 5mg/cm2). To measure denser materials (such as metals), g- radiation is used, emitted by radionuclides 241Am (60keV), 137Cs (662keV), 60Co (1173 + 1332keV). With these methods, it is possible to measure (scan) the relevant objects, or even continuously monitor the thickness of the produced foil or rolled metallurgical material on the conveyor belt.
    The attenuation of radiation as it passes through the substance is also significantly dependent on the density of the monitored material. With a known (constant) material thickness, we can monitor the density of the material and its changes based on the attenuation of the transmitted radiation beam. Densitometers of this type are used, for example, in monitoring the transport of substances by pipeline (in the chemical or food industry) or belt conveyors (coal treatment plants, dosing of components in metallurgy, etc.).
Note: If the measured sample is accessible from one side only, it is sometimes used to measure the thickness and density of the scattering method: the object is irradiated with a beam of radiation and the intensity of Compton backscattered radiation (Fig.2.8.1 in the middle) is monitored, which depends on the thickness and density of the material. This is the case with pipe walls, boilers and closed vessels, or in boreholes (logging measurements). However, the accuracy and sensitivity of these scattering methods is lower than for transmission methods.

Radiation level meters
Radiation level meters determine the height of a liquid column based on the attenuation of the radiation beam
g by the liquid, depending on whether the radiation passes through liquid or air. In addition to liquids, bulk materials can also be monitored in this way. This non-contact method is important where other methods cannot be easily used - for example in overpressure and underpressure vessels, or for liquids aggressive or heated to high temperatures. The most commonly used g- emitters are radionuclides 137Cs (662keV) and 60Co (1173 + 1332keV), or 241Am (60keV).
    The geometric arrangement of the radiator and detector is most often horizontal, where the radiator and the detector are placed on the sides of the tank opposite each other and detect the level
(in the simplest case, one radiator and one detector are used, whose response after amplification switches the relay contacts when the level reaches the measured point). In a vertical arrangement, the source and detector are below and above the tank, so that the attenuation of the radiation as it passes through the liquid column is exponentially dependent on the level height in the tank; the level height can be monitored continuously.

Neutron measurement of humidity
This method is based on the elastic scattering of fast neutrons on hydrogen nuclei (contained in water), which of all elements most effectively scatter and decelerate neutrons. The humidity meter consists of a source of fast neutrons (usually
241Am in a mixture with beryllium, activity of about one hundred MBq to units of GBq) and a detector of slow neutrons (see §2.6). Either a transmission arrangement can be used, where the attenuation of the flux of neutrons from the source due to their scattering on the hydrogen nuclei is measured, or reflective, where the increase in the flux of slowed neutrons due to their scattering in the surrounding material containing hydrogen nuclei is measured.
    The neutron method of measuring the moisture content of materials is used in many industries, eg in the chemical industry, construction, agriculture, mining. It is most often used to measure the moisture content of bulk materials such as soil, sand, mortar mixtures, ores, coal and coke, grain, etc.
Note: The response of neutron flux is given by the total volume humidity. If it is necessary to determine the mass humidity, sometimes combined neutron + gamma probes containing a neutron source and a g- radiation source (eg 137Cs) and a neutron detector and a g detector are used (it is often combined into one compact probe). From the response to radiation g it is possible to determine the density of a material, from the neutron response its moisture; the conversion of bulk moisture to mass moisture can then be realized electronically in the evaluation unit of the device.

Radiation defectoscopy
Another method, based on differences in the absorption of penetrating radiation in substances, is radiation defectoscopy. During casting, cooling, welding, machining and operation of products and parts of machines and equipment, inhomogeneities, cavities, cracks and similar internal defects can occur, which impair the mechanical properties of the part and can lead to machine failure. Defectoscopy in general is a method of non-destructive examination of structure ("defects") in the macroscopic consistency of a material.
    Radiation defectoscopy allows by non-destructive analysis of inhomogeneities to detect possible cracks and other anomalies in construction materials and finished products. The basic scheme of defectoscopic measurement is similar to the above-described
X-ray diagnostics in medicine. The analyzed object is irradiated with a collimated beam of penetrating radiation X or g, while the transmitted radiation is displayed on a photographic film - radiography, or is displayed on a fluorescent screen or electronic detector to a computer - radioscopy. The absorption of ionizing radiation depends on the thickness and density of the material (see the exponential relationship in the section "Absorption of radiation in substances" in §1.6), so that the weakened areas are reflected in greater blackening of the film.
    Possible inhomogeneity or crack will appear on the film after development as a local defect in an otherwise homogeneous blackening of the emulsion. Films and developers are used, ensuring the highest possible steepness of the blackening curve, so that even small inhomogeneities of transmitted radiation are displayed in sufficient contrast. The blackening of the film is usually evaluated visually using special transmission lamps, or they can be evaluated photometrically. Now the films are gradually being abadonded - the transmitted radiation is detected by a sensitive electronic detector. The detector, a digital semiconductor flat imaging detector (
flat panel), detects the intensity of gamma or X-ray radiation that passes through the material and the occurrence of a defect (cracks, cavities, etc.) is reflected in a change in the intensity of the measured radiation at a given location.
  For defectoscopy of steel objects, either X-rays with an energy of about 60-400keV from technical X-rays tube, or
g- rays from suitable radionuclides are used - iridium 192 Ir, selenium 75 Se, cesium 137 Cs, cobalt 60 Co, occasionally also hard braking radiation g with energy up to 10MeV produced by a linear or circular electron accelerator (braking of electrons on a target). Hard gamma radiation or braking radiation must be used especially when shine trough of metals (steel) with a thickness greater than 100 mm, where ordinary X-rays are no longer sufficient. For radiography of thinner layers, on the other hand, softer photon radiation is more suitable, which provides a higher contrast of the displayed small defects (from the radionuclides, the aforementioned Ir-192 or Se-75 is suitable).
    Defectoscopy is used especially where high demands are placed on the quality of materials and components. These are, for example, gas pipelines, turbine blades, reactor pressure vessels, bridge structures, etc.
X-ray microscopy, micro-CT
So -called micro-X-ray tubes are used for structural analysis of small objects (such as electronic components or small castings). The special X-ray tube has a very small impact focus (only a few micrometers), so the X-ray beam emanates almost from a point source and provides high sharpness and image resolution. The measured sample is placed very close to the X-ray tube and the film or imaging detector at a greater distance - there is a projection magnification of the image. For this purpose, X-ray lamps with a thin front so-called transmission anode are sometimes used (see §3.2, section "X -rays", passage "Special types of X-ray tubes"), which allows you to bring the displayed object as close as possible to the focus on the anode and thus achieve high magnification at a small distance between the illuminated object and the imaging detector. For X-ray microscopy (XRM) mainly soft X-rays of approx. 20 - 60keV are used. Low-energy photons of X-rays interact with the atoms of the investigated substance mainly by the photoeffect, which provides a higher absorption contrast between areas with small differences in density. In large special laboratories, very soft X-rays (approx. 2 - 10keV - around the K- or L-edge of the absorption spectrum of the examined material, where the absorption and imaging contrast is greatest) from the synchrotron undulator are used for X-ray microscopy (see §1.5, section "Charged particle accelerators", section "Accelerators as synchrotron radiation generators"), with the possible use a crystal monochromator and a Fresnel zone plate, acting as a connecting lens of X-ray optics. Either scan mode or display using special pixel detectors is used. These are very demanding laboratory methods!

X-ray microscopy with a special microfocus X-ray tube with a transmission anode.

CT X-ray tomography or micro-tomography (mCT) is also used for detailed 3D analysis of small objects, the principle of which is analogous to the above-described medical X-ray tomography (§3.2, section "X-ray tomography - CT"). The main difference is that the X-ray tube and the detection system do not rotate during the measurement, but the displayed object rotates between the static X-ray tube and the imaging detector. The passed X-rays, measured by an imaging detector for a number of different angles of the rotating sample, are reconstructed into cross-sectional images, the set of which forms a 3-D image of the analyzed object.

Safety inspection of materials
The X-ray inspection of baggage, used in recent years at airports, is also based on the principle of radiation measurement of mechanical properties (density) of materials. Small X-ray machines - an X-ray tube and an opposite electronic imaging detector - are installed at the baggage counter for air traffic, between which checked baggage passes. The resulting absorption image is immediately projected on the display, sometimes with a "pseudo" color display (artificial assignment of colors to grayscale), in order to recognize mainly metal objects (such as weapons).

X-ray diffraction analysis of the crystal lattice structure
When X-rays fall on a substance with a crystal structure, diffraction of part of the X-rays occurs, during which this radiation ir reflected from the regular structure of the crystal lattice - X-rays elastically scattered on the electrons of the measured crystal. Subsequently, interference of this difracted X-radiation may occur. In the diffraction interference picture is then encoded information about the internal structure of the crystal.
    At the incidence of monochromatic X - rays with a wavelength of l » 0.1 nm (comparable to the distance between the ions forming the crystal lattice), the rays may be amplified in one direction, weakened or disturbed in others. X-rays are amplified and form an interference maximum if the so-called Bragg condition is met, so that the path angle of two rays is an integer multiple of the wavelength of the radiation: n. l = 2.d. sin J, where J is the angle formed by the incident beam with the crystal plane, d is the distance between two adjacent crystal planes (lattice constant), l is the wavelength of the X-ray radiation used and n = 1,2,3, ... is integer. In this situation, the intensity of the scattered waves adds up. For a given crystal having a lattice constant d is thus interfering peaks at the diffraction is reached only at certain values of l and J. Usually, 1st order angular spectra (n = 1) are scanned, where the maximum is most pronounced, only to distinguish some details, higher order spectra are analyzed (low intensity - significant prolongation of exposure time).

Fig.3.3.2. Principle arrangement of X-ray diffraction analysis of crystal lattices

Apparatus for measuring the diffraction called diffractometer consist of a goniometer, in whose center is stored analyte and on whose one arm is a source of X-radiation and the other arm a detector measuring the intensity IX of X-radiation. By turning the goniometer, the angles J are measured, for which the maximum intensity IX of the "reflected" X-ray, i.e. the interference maximum, is detected in the reflection mode. The measurement in the transmission mode is rarely used, when the diffraction of X-radiation passed through the sample is measured. Possibly, display of the diffraction pattern on photographic film or electronic imaging detector. The X-ray continuous spectrum monochromator is included either in the primary beam or in the secondary diffraction radiation path. For a detailed analysis of the structure of single crystals, the angles of the X-ray tube J1 and the detector J2 are measured independently and another goniometer is included, enabling the sample to rotate even in the perpendicular direction (Bragg-Brentan diffractometer). X- ray tube with microfocus is used in X-ray microdiffraction (its construction was described above in the section "Special types of X-ray tubes" and the use in the previous paragraph "X-ray microscopy; micro-CT "); in the most demanding applications also synchrotron X-rays (§1.5, section "Synchrotron radiation generators").
    X-ray diffractometry is used in many areas of materials research, especially for the analysis of the crystal structure of substances - both single crystals (single crystal X-ray structural analysis) and polycrystalline and powder materials (powder X-ray structural analysis). Also in the non-destructive analysis of archaeological and artistic objects.
Note: It can also be used to decompose continuous (polychromatic) X-rays and obtain a monochromatic component.

Positron annihilation spectrometry
Positron annihilation spectrometry is used to analyze local electron densities and configurations in substances. It is based on spectrometric measurements of the positron lifetime in substance
(PLS - Positron Lifetime Spectroscopy). If we irradiate the analyzed material with positrons, fast positrons slow down in the substance in the path of a few micrometers (in a time of about 10 picoseconds) and under normal circumstances they can (via an unstable positronium) annihilate with electrons. However, they can be retained in places of structural irregularities in the crystal lattice (pores with reduced electron density) and annihilate with a delay with electrons from the surroundings. In order to determine the lifetime of positrons, we first need to detect the moment of formation (emission) of the positron. This is possible when the radioactive source of positrons synchronously emits gamma radiation from the excited level of the daughter nucleus.
    The investigated material is thus locally irradiated with a mixed
b+- g emitter (most often 22 Na), while the lifetime of positrons is determined by measuring the delayed coincidences between the detection of the g photon from the radiating radionuclide (at 22Na it is g 1274 keV) and detection the g 511 keV annihilation photon.
    In the case of the most common use of a b+- g emitter Na22 the detection radiometric apparatus consists of two gamma-detectors (scintillation or semiconductor) :
1. A detector set to a 1274 keV photopeak of gamma daughter nuclide deexcitation radiation
22Ne. This is the "start" impulse of the time coincidence analysis, indicating the moment of positron formation.
Detector set to 511keV of annihilation radiation generated by e
- e+ positron annihilation. This is the "final" impulse of time coincidence analysis, determining the moment of positron extinction .
    In solids maters without structural defects the lifetime of positrons is about 0.25ns, in positrons annihilating in defects it is extended to about 0.75ns. With this method it is possible to observe defects in material structure of about 0.1 to 1 nm - dislocations, vacancies, clusters of vacancies, clusters, or precipitates. It is used to monitor the technology of preparation of various materials (plastics, metals, conductors, insulators, semiconductors) and also to monitor the influence of the environment and technologies on materials (fatigue and "aging" of materials, thermal and radiation effects, etc.).

3.4. Radiation analytical methods of materials
The methods of atomic and nuclear physics, as well as the properties of different types of radiation, provide important tools for the analysis of the material and elemental composition of various objects. Most of these methods work in the experimental arrangement ideologically shown already in the introductory Fig.3.1.1a,b in §3.1. The analyzed object or sample is irradiated with a suitable type of primary radiation, the interaction of which with atoms or nuclei creates secondary radiation, which "brings out" some information about the composition of the material. This radiation is detected and analyzes. From the large number of different atomic, nuclear and radiation analytical methods, we will present only a certain selection of the typical and more frequently used ones, with an emphasis on the physical nature, without excessive technical details.

X-ray fluorescence analysis
This method of non-destructive determination of the chemical (elemental) composition of substances is based on the measurement of the secondary fluorescence characteristic X-rays induced by the primary irradiation of the examined sample. The measured sample is most often irradiated with photon radiation - either X-rays from an X-ray lamp or gamma radiation from a suitable radionuclide - Fig.3.4.1
(irradiation with charged particles is mentioned at the end of this passage). The interaction of this photon radiation with the atoms of the examined sample cause, among other things, a photoeffect (see the passage "Interaction of gamma radiation" in §1.6 "Ionizing radiation") mostly on the K shell (if the radiation energy is higher than the binding energy of the electron on this shell), after which when the electrons jump from the higher shell (L) to the released place, a characteristic X-radiation (K series) is emitted, whose energy is unambiguously determined by the proton number Z of the atom. If a photoeffect occurs on the shell L, then the characteristic X-radiation of the series L is emitted by the electron jump from the shell M.
    The energies (spectral lines K
a,b) of fluorescent X-rays are characteristic for each element, the amount of emitted photons of characteristic radiation is directly proportional to the number of atoms of a given species, thus a measure of the concentration of a given element. By spectrometric analysis the energy (wavelength) of the resulting fluorescent radiation can be used to determine which elements are present in the sample under investigation, and the amount (concentration) of these elements in the sample can be determined according to the intensity of the individual fluorescence peaks.
    The method of exciting characteristic X-rays by primary gamma rays is sometimes referred to as XRF (gamma-induced X-ray Fluorescent Emission).

Typical arrangement of radiation source, analyte and detector in X-ray fluorescence analysis.
At the top right is the detailed structure of the peaks K
a,b of the characteristic X-ray, measured by a semiconductor Ge(Li) detector.

The energy of the primary excitation radiation g or X is most suitable only slightly higher than the binding electrons on the shell K (or L) in the atoms of the analyzed elements; then the highest effective cross section is for the photo effect. Therefore, different irradiation sources are used for lighter, medium and heavy elements. Thus, in addition to the X-ray lamp, radionuclides emitting soft X-rays such as iron 55Fe (X Mn L-series 5.9-6.5keV), curium 244Cm (X Pu L-series 12-23keV ) are used for the irradiation of the examined samples for the analysis of light elements, for medium-heavy elements americium 241Am (g 60keV) , for analysis of heavy elements such as gold, tungsten, lead, uranium, etc., then cobalt 57Co (g 122 + 136keV), cesium 137Cs ( g 662keV), cer 144Ce (g 140keV).
    Scintillation detectors are used to detect characteristic X-rays for simpler and indicative measurements (such as geological survey and ore prospecting, metal content control in metallurgy, etc.), but a high-resolution semiconductor detector must be used for more accurate and complex laboratory analysis, and multichannel analyzer. For quantitative analysis, a correction for interfering Compton scattered radiation
g must be made and, of course, careful calibration of the device.
   Characteristic X-rays have four very close energy lines (related to the fine structure of electron levels K and L), which are referred to as K
a1, Ka2, Kb1, Kb2 - Fig.3.4.1 top right. E.g. for lead these energies are 72.8, 74.97, 84.8, 87.3 keV, for gold 66.99, 68.81, 77.9, 80.1 keV, for iron the energy of X-rays is only 6.4 keV, for aluminum 1.5keV (for these low energies it is practically no longer possible distinguish lines Ka and Kb). Thus, for light elements, the X-ray energy is very low and difficult to detect. X-ray fluorescent analysis is therefore particularly suitable for determining content of heavier elements.
   To excite characteristic X-rays, primary irradiation with charged particles that ionize the atoms of the substance is sometimes used, followed by deexcitation and emission of X-rays. The method of particle-induced X-ray emission is called PIXE (Particle-Induced X-ray Emission). It usually irradiates by protons with an energy of about 2-4 MeV, the surface layer of the sample is analyzed to a depth of about 5
   X-ray fluorescence analysis has the great advantage of being fast, accurate and reproducible, does not require any chemical processing of samples, which can be used in all states of matter. The examined material is not damaged in any way and no artificial radioactivity is generated. It is possible to examine entire objects, without the need for sampling - this is a non-destructive method. It is therefore suitable, among other things, for the analysis of the composition of art objects, which can help their temporal or authorial classification, finding out the origin, as well as verifying their authenticity.

Activation analysis
This nuclear-analytical method is based on the irradiation of a test sample with such radiation (type and energy) that enters the nuclei of the investigated atoms and causes nuclear reactions there. During these reactions, radiation (especially gamma) is emitted, but mainly unstable isotopes are formed from originally stable nuclei - radionuclides, which subsequently decay by radioactive transformation
a or b with subsequent emission of photons g. Either the secondary radiation emitted during the nuclear reaction itself is measured, but above all the g- spectrum is measured, emitted by radionuclides generated by nuclear reactions due to primary irradiation. The analysis of this spectrum determines the elemental composition of the sample (qualitative and, if necessary, quantitative).
    Neutron activation analysis NAA 
(also called induced or instrumental neutron activation analysis INAA , see below) is a highly sensitive method of analysis of chemical composition of substances, based on neutron capture (reaction n, g) in the nuclei of the test substance, thus forming radioactive nuclei (see §1.3 "Nuclear reactions"): NAZ + n ® N+1B*Z; B* ® B + gP; N+1BZ ® N+1C*Z+1 + e-(b) + n; C* ® C + gD. During this reactions, two types of gamma radiation are emitted: immediately after neutron capture, it is gP radiation, followed by radioactive decay of activated nuclei, gD radiation is emitted - lower part Fig.3.4.2. Irradiation of the examined sample with neutrons thus results in the formation of radionuclides - "activation" of the sample; after by spectrometric analysis of which energies and radiation intensities (especially g) emitted from the activated sample, can be determined the relevant radionuclide and "traced" the corresponding (inactive) starting nuclide contained in the sample, the activation of which this radionuclide was created (Fig.3.4.2). Using a suitable calibration, its content (concentration) in the examined material can also be determined.

Fig.3.4.2. Typical procedure for neutron activation analysis.

Neutron irradiation of the analyzed samples is performed either in irradiation chambers in a nuclear reactor as shown in the figure (nuclear reactor is a powerful source of neutrons, see §1.3, section "Fission of atomic nuclei"), or using neutrons from special accelerators, so-called neutron generators (§1.5, part "Charged particle accelerators", passage "Neutron generators"). In the laboratory conditions and in the terrain is also used radionuclide neutron sources, a blend of the alpha-emitter with a light element (e.g. a -radionuklide 241Am in a mixture with beryllium, reactions a,n occur), or a heavy transuranic radionuclide (most often californium 252), during the spontaneous fission of which neutrons are released (§1.3, "Transurans"). For neutron activation analysis, mainly slow neutrons with energies of about 0.001-0.55 eV are used , which have a high effective cross-section of capture by many nuclei. From neutron sources, which usually provide fast neutrons with energies of the order of MeV, the neutron beam is first led to the moderator and the samples are irradiated only by slowed neutrons.
    For complex NAA, the detection of gamma radiation from neutron-irradiated samples is usually performed by semiconductor
g-spectrometers with high energy resolution (§2.5 "Semiconductor detectors") in order to identify the exact energies of gamma radiation and to distinguish peaks often in close proximity. For some simpler applications, where it is enough to measure the representation of one or a few elements, scintillation detectors can be used (which do not have such good energy resolution, but have higher detection efficiency - §2.4 "Scintillation detection and gamma-ray spectrometry"). If the measurement of the activated sample is extended by the simultaneous - coincidence - detection of two or more quanta of emitted gamma radiation by means of two spectrometric detectors, the method is referred to as coincidence activation analysis CINAA (Coincident INAA). The method is suitable when the activation results in radionuclides with a cascade deexcitation emitting a pair of photon quanta (such as 60Co). Coincidence measurement then sharply reduces the background of interfering impulses. Detection can optionally be combined with position-sensitive detectors (such as semiconductor pixel detectors) that register soft g- radiation or charged particles, especially electrons b-, which are emitted by activated nuclei in coincidence with photons g. In this way, the spatial distribution of the analyzed element in the sample can be displayed.
    In terms of adjustment of measured samples, two methods of activation analysis are used :
¨ Instrumental INAA activation analysis, where the irradiated sample is measured directly, without chemical treatment, on a g- spectrometer. This is the simplest and most common way to implement NAA. In the respective device, the neutron source and the g- spectrometer are sometimes integrated in one compact device, which can be used not only in the laboratory, but also in the field. Such measurements can also be performed in a non-destructive way: We irradiate the analyzed object or its part with a neutron source, we measure the induced radiation g, after which we can return the object to its original use (unless it is activated too strongly by long-term radionuclides).
¨ Radiochemical RNAA activation analysis, in which the sample is first subjected to chemical separation after irradiation - either to remove interfering radionuclides (which could overwhelm the analyzed radionuclides or to interfere with them), or to increase the concentration of required radioisotopes. This method is used less often for considerable labor and laboratory complexity.
    In terms of the time relationship between irradiation and measurement, neutron activation analysis is divided into two categories :
l Subsequent - delayed gamma-neutron activation analysis DGNAA (Dellayed Gamma-ray Neutron Activation Analysis), where gamma radiation measurements from the sample are performed after the end of the neutron irradiation (as in Fig.3.4.2 in the middle). The "subsequent" (delayed) radiation gD is measured here, arising from b -radioactivity of activated nuclei N+1BZ ® N+1C*Z+1 + e-(b) + n by deexcitation of excited levels of the daughter nucleus: C* ® C + g. This is the most commonly used method, suitable where neutron activation produces radionuclides with a longer half-life (minutes and longer).
Immediate ( prompt *) gamma-neutron activation analysis PGNAA (Prompt Gamma-ray Neutron Activation Analysis), when the measurement of emitted g-radiation is performed during neutron irradiation (Fig.3.4.2 on the right). Radiation g of two types (origins) here is measured from the irradiated sample with a gamma spectrometer : 1. Immediate photons gP , usually emitted very quickly after neutron capture from excited levels of activated nuclei B* ® B + g. 2. Radiation gD arising subsequently from b -radioactivity of activated nuclei N+1BZ ® N+1C*Z+1 + e- (b) + n by deexcitation of excited levels of the daughter nucleus: C* ® C + g. This method is suitable when neutron activation produces short-term radionuclides (which would usually decay during the time between irradiation and sample measurement), or stable nuclides, or radionuclides with pure b- decay or a small proportion of g- radiation (then applied here the immediate photons gP generated after neutron radiation capture). The prompt NAA automatically falls into the category of instrumental activation analysis mentioned above.
*) In a sense, the method of neutron stimulated gamma emission can also be included in this category, when the sample is irradiated with fast neutrons, the inelastic scattering of which leads to excitation of nuclei in the analyzed sample. During subsequent deexcitation, g- radiation of characteristic energies for individual nuclides is emitted. The presence and concentration of the respective elements and their isotopes can be determined by spectroscopic detection of this g- radiation, performed during neutron irradiation. This method was also experimentally tested for the purpose of in vivo gamma imaging in medicine (see §4.3, passage "Neutron stimulated emission computed tomography NSECT").
    Neutron activation analysis can achieve extremely high sensitivity (it allows to detect even 10-12 g of element in 1 g of sample), so it is suitable for detecting trace amounts substances, eg trace element content in plant and animal tissues, water pollution, purity of semiconductor materials, etc.
Note: For special purposes of biological research, in vivo neutron activation analysis is sometimes used : the relevant part of the organism is irradiated with neutrons (from reactor or neutron generator), followed by a gamma imaging of the distribution of induced beta radioactivity (accompanied by gamma), mapping the distribution of the test substance in tissues and organs.
    In addition to neutron activation, proton and gamma-activation analysis are also rarely used, in which protons accelerated on an accelerator, such as a cyclotron, are used to activate the nuclei of the sample
(causes reactions of proton capture [p, g], or reactions of type [p, n], [p, d], [p, a]), or high-energy gamma radiation (causes photonuclear reactions [g, n], [g, p], at higher energies more particles can be ejected from the nucleus [g, 2n], [g, d], [g, 2p], [g, a] ), arising as braking radiation by electrons accelerated in betatron, microtron or linear accelerator.

Mössbauer spectroscopy
Mössbauer spectroscopy is a non-destructive analytical method based on the so-called Mössbauer effect of resonant nuclear absorption of
g radiation - see §1.6, section "Interaction of gamma radiation". The sample is irradiated with monochromatic radiation g and the detector measures the intensity of transmitted or "reflected" (resonantly scattered) radiation as a function of subtle changes in radiation energy g, which varies in a narrow range due to Doppler effect by precisely controlled mechanical movement of the source relative to the sample by a linear motor. Radiation g it must have an energy exactly corresponding to the excited level of the core of the sample under examination. The Doppler effect compensates for the energy loss of the reflected nuclei, resonant absorption of photons g, accompanied by a maximum of absorption and subsequently emission of a photon of the same energy.
    This method is applicable to substances containing elements which form as daughter nuclei of suitable radioisotopes and have excited levels emitting radiation
g ; the samples are irradiated with radiation g from such a radioisotope. The fine position of the absorption maxima depends on the properties of the chemical bond in which the atoms containing the analyzed nuclei participate, on the properties of the crystal lattice, as well as on the internal magnetic and electric fields in the crystals. By analyzing the fine structure of the Mössbauer spectrum (which is the dependence of the absorption of g on the feed rate of the source relative to the sample), some internal chemical and physical properties of the investigated material can be determined.
   This method is suitable for
57Fe, 57Co, 129In, 119Sn, 121Sb *). It is mainly used on materials containing iron 57Fe. It allows the analysis of the distribution of iron in the material in various crystallographic positions, its degree of oxidation, analysis of ferromagnetic materials, alloys, minerals, etc. For analytical purposes, the samples are made into a thin film or powder (weighing several grams).
*) The number of suitable elements (nuclei having a suitable radionuclide emitting
g radiation from a suitable excited level of a stable daughter nucleus) suitable for this analysis is very limited , so the significance of Mössbauer spectroscopy is not comparable to such methods as activation analysis, X-ray fluorescence analysis or defectoscopy...
   In the Mössbauer spectrometry of iron, the radionuclide 57Co is used as the radiation source g, which decays to an excited 57Fe nucleus with a half-life of 270 days by electron capture. This nucleus emits 692keV (0.14%), 136keV (11%), 122keV (87%) and 14.4keV (9%) g radiation when deexcited. It is the radiation g of the partial transition from the excited level with an energy of 14.4 keV, that is suitable for excitation of resonant nuclear absorption due to its low energy. Due to the high Debye temperature Fe (360 °K), the Mössbauer effect occurs even at normal laboratory temperatures, wheres the Doppler frequency shift required to compensate for the reflection of the 57Fe core is achieved by mechanical displacement of the source at speeds of only the order of 1 mm/s.
    Note: High sensitivity Mössbauer effect of resonant nuclear absorption
g -radiation 14.4keV of 57Fe was used in 1960 by R.V.Pound and G.A.Rebka to measure the gravitational frequency shift in the Earth's gravitational field, which was an important test of Einstein's general theory of relativity as a physics of gravity and spacetime - see §2.4 "Physical laws in curved spacetime" in the book "Gravity, black holes and space - time physics".

Mass spectrometers and separators
Mass spectrometers and separators, used in physical chemistry and radiochemistry, work in a similar arrangement as the magnetic spectrometer of charged particles according to Fig.2.6.1 on the left - see §2.6, section "
Magnetic spectrometers". The analyte is ionized in the ionization chamber, the formed cations of charge e are accelerated by an electric field and ions with a constant velocity v are selected in a velocity filter (consisting of, for example, a crossed electric and magnetic field). These then fly through the input slit into the magnetic field of intensity (induction) B, in which they describe a circle with radius R = (v/e.B) .m, proportional to the mass m. The ions of different weights describe different paths and thus fall on different places of the base - the device thus separates ions of different weights (given the weight of the core). By changing the magnetic field, ions of corresponding masses are gradually focused into the detector - a mass spectrum is created. In the mass separator, a suitable target is installed instead of the detector, on which the incident ions of the selected mass are absorbed.
    Magnetic mass spectrometry is a complex method for the most accurate analysis of the representation of elements and their individual isotopes in the analyzed substances. Magnetic mass separation makes it possible to isolate completely pure samples of a precisely defined isotopic composition, but only in very small quantities.

Gas ionization analyzers
As ionizing radiation (
a or b) passes through a gaseous medium, absorption and ionization depend on the density and composition of the gas. The flow ionization chamber with a built-in emitter a or b can thus serve as an analyzer for checking the composition of the gases.
Fire detectors
The ionization fire detector consists of two electrodes with an air gap. Radiation
a from the applied layer of radionuclide (most often 241Am, approx. 30 kBq) generates an ionization current in gas between the electrodes. In the presence of smoke between the electrodes, the absorption of the gas environment changes and thus the ionization current changes, which is registered by the electronic circuits of the fire alarm system.
Electron capture radiation detector - ECD

To detect compounds with high electron affinity (such as the Freons, chlorinated pesticides and other halogenated compounds ) may be a radiation electron capture detector (ECD - Electron Detector Capure). It consists of a cylindrical ionization chamber filled with an inert gas (eg argon), one electrode (cathode) of which is provided with a layer of a low-energy radiator b, usually 63Ni (activity approx. 300MBq). The emitted radiation b creates an ionization of the gas atoms, a certain ionization current flows through the chamber. When a gas containing high electron affinity atoms enters the chamber, these atoms absorb the electrons in the ionized gas and the ionization current through the chamber decreases, which is electronically registered. Such chambers are often used as a terminal detector in gas chromatography columns.

Nuclear magnetic resonance - analytical and imaging method
Nuclear magnetic resonance (NMR) is a very complex physical-electronic method, based on the behavior of magnetic moments of atomic nuclei under the action of an alternating radio frequency signal in a strong permanent magnetic field. This originally analytical method was later improved and developed even as an important imaging method.
Note: We have included nuclear magnetic resonance among nuclear and radiation methods, even though it does not contain any ionizing radiation. However, it is a method based on the findings of nuclear physics - the properties of atomic nuclei. A physical phenomenon called nuclear magnetic resonance - NMR, was investigated in the 1940s (F. Bloch, E.M.Purcell) and was initially used in chemistry as sample NMR spectrometry . In the 1970s and 1980s, NMR imaging methods also began to develop (pioneers were P.Lauterbuer, P.Manfield, A.Maudsley, R.Damadian, 1977) - see below.
    We will try to briefly outline the principles and methodology of NMR. However, due to the considerable principal and technical complexity of NMR (only scintigraphy can partially compete with it), we must observe the maximum brevity...
Physical principle of NMR
Phenomenon of nuclear magnetic resonance it can generally occur during the interactions of atomic nuclei with an external electromagnetic field. Each nucleon (proton and neutron) has its own "mechanical" angular momentum - spin (nucleons belong to fermions with spin 1/2, see §1.5 "
Elementary particles"). According to the laws of electrodynamics, this rotational angular momentum of nucleons creates - induces - its own elementary magnetic moment mp = 1.41.10-27 J / T, equal to 2.79 times the so-called Bohr nuclear magneton *) - it is discussed in more detail in §1.1, passage "Quantum momentum, spin, magnetic moment", paragraph " Magnetic moment ". Due to the spins of their nucleons, atomic nuclei therefore generate a very weak magnetic field - they have a certain magnetic moment m. However, only atomic nuclei with an odd nucleon number have spin and magnetic moment, because the spins and magnetic moments of paired protons and neutrons cancel each other out - they are zero. The magnetic moment of the nucleus is formed by an unpaired nucleon - a proton or neutron. Magnetic resonance imaging can therefore be observed only in nuclei with odd nucleon numbers - especially hydrogen 1H, then in 13C, 15N, 19F, 23Na, 31P, etc.
*) Nuclear magneton
mN is a physical constant expressing the proton's own dipole magnetic moment induced by its spin: mN = e.h /2mp , where e is the elementary electric charge (proton, electron), h is the reduced Planck's constant, mp is the rest mass of the proton. In the system of SI units, its value is approximately mN = 5.05.10-27 J /T. It is analogous to the Bohr electron magneton me = e.h / 2me , which, however, is 1836 times larger, in the ratio of the mass of the proton and the electron. It is interesting that even a neutron, although electrically uncharged, has a non-zero magnetic moment mn = -0.966.10-27 J /T somewhat smaller and of the opposite sign than a proton. It turns out that the magnetic moment of nucleons has its origin in their quark structure (§1.5., part "Quark structure of hadrons" and §1.1, passage "Magnetic moment").
Magnetic moments of nuclei in a magnetic field 
Under normal circumstances, due to the thermal motion of atoms, the directions of spins and magnetic moments of individual nuclei are chaotically "scattered", their orientation is random and disordered (Fig.3.4.4a), elementary magnetic fields cancel each other out on average, on a macroscopic scale the substance shows no magnetic properties
(we do not mean ferromagnetic substances, where it is the effect of electrons in atomic shells) . However, if we place the analyzed substance in a strong magnetic field (of intensity or induction B of the order of several Tesla), the magnetic moments of the nuclei are oriented in the direction of the vector B of this external magnetic field (at least partially).- the magnetic moment of the nuclei is parallel to the magnetic field lines (Fig.3.4.4b). The stronger the magnetic field, the more perfect this arrangement *). Outwardly, this results in non-zero magnetization vector M in the direction of the external magnetic field induction B. The magnitude of the magnetization vector is proportional to the strength of the external magnetic field B and the percentage of concordantly oriented mag. moments of nuclei in matter. A sufficiently strong magnetic field B is now mostly realized by means of a superconducting electromagnet, the winding of which must be permanently cooled by liquid helium (physical principles of superconducting magnets are briefly discussed in §1.5, section "Electromagnets in accelerators", passage "Superconducting electromagnets").
*) However, the extent of this arrangement is actually very small ! In commonly used magnetic fields 1-3T, for every 1 million hydrogen nuclei, only about 7-20 nuclei are on average in a state of uniform orientation. The vast majority of nuclei are as a result of thermal motion, it is oriented in different directions, including the opposite one (this is expressed by Boltzmann's law of distribution.) In this sense, it is necessary to take Fig.3.4.4b only as a symbolic scheme, which shows only those few nuclei that acquire concordant orientations.
  Since conventional material, e.g. water or tissue, contains about 1022 hydrogen nuclei per 1 gram, even a small excess of oriented nuclei provides a measurable magnitude of the magnetization vector and the radio frequency response signal.
Larmor frequency, radiofrequency excitation and relaxation 
In the magnetic field B, the nuclei (with a non-zero magnetic moment
m) behave as magnetic dipoles, which are acted upon by a pair of forces m.B . This will cause the core to rotate the axis of its magnetic moment around the direction B - it will perform a precessional movement (similar to the precessional movement of a gyroscope or children's "spinning top" around the vertical direction in the gravity field) by the so-called Larmor frequency
wL = g . B  , or    fL = g .B /2p ,
where g is the gyromagnetic ratio of the nucleus, which is the ratio of the magnetic moment of the nucleus and its "mechanical" moment of inertia [ rad · s -1 · T -1] . The precession movement occurs when the external magnetic field changes or the angle of the magnetic moment in this field changes and lasts as long as the mag. the moment does not stabilize in the rest position.
    If we send a short alternating electromagnetic signal into such a magnetically polarized substance by means of another coil (high-frequency - HF, or radio-frequency - RF)
(whose frequency resonates with the above-mentioned Larmor precession fL of a given type of nucleus in a magnetic field), the direction of the magnetic moment of the nucleus temporarily deviates from the direction determined by the vector B of the external magnetic field (Fig.3.4.4c) *). The deflection of the magnetization vector is caused by the magnetic component of the excitation RF pulse. The angle of this deflection is proportional to the amplitude (energy) of the RF pulse and its duration. The most commonly used RF pulses are 90° or 180°.
*) Fulfillment of the resonance condition: The nuclei are able to efficiently receive energy from an alternating electromagnetic field only if the Larmor frequency of the nucleus precession and the frequency of the electromagnetic pulse are the same. The preceding nuclei thus resonate with an electromagnetic pulse at a given Larmor frequency - hence the name "magnetic resonance".
    After the unwinding of the excitation, signal occurs relaxation
(at a constant rotation Larmor frequency) at which they emit electromagnetic waves with decreasing intensity until the magnetic moment of the spiral return back again in the direction B. These electromag. waves will induce alternating voltage in the receiving coils - "echo" radiofrequency signal **). This relaxation signal (sometimes referred acronym FID - Free Induction Decay) , has a sinusoidal course with exponentially decreasing amplitude (see below Relaxation times). It is a useful signal that carries information about the inner structure of the analyte. Frequency of this signal is equal to the above-mentioned Larmor precession and for a given force B of the external magnetic field is determined by the gyromagnetic ratio g of the nucleus, ie the type of nucleus, the amplitude of the relaxation signal is proportional to the concentration of nuclei of the given species- thus nuclear magnetic resonance can be used to analyze of the composition of substances : what elements and in a what concentration are contained in the sample. E.g. for hydrogen nuclei (protons) the gyromagnetic constant has the value g = 2.675.10-8 s-1 T-1 and in the magnetic field of induction 1Tesla Larmor's NM the resonant frequency is 42.574MHz, at 1.5T it is 63.58MHz - the area of radio waves (short waves) . For heavier nuclei is proportionally lower .
**) Phasing of a large number of nuclei : The NMR receiving coils are, of course, not able to detect the relaxation radiation of one or a few nuclei. To obtain a measurable signal, deexcitation of a large number of nuclei (> about 1012 ) is required, namely synchronously and at the same phase ! If phasing disruption occurs, the MNR signal drops sharply or disappears (cf. below "Relaxation times - T2").
    Gnoseological note :
Quantum behavior:
For the sake of clarity, we have not yet explicitly included the quantum behavior of the magnetic moment, we considered its continuous behavior. The orientation of the magnetic moment vector of nuclei in a magnetic field actually acquires discrete quantum states - parallel (0°), perpendicular (90°) and antiparallel (opposite, 180°) with the direction of the vector B
  magnetic induction of an external magnetic field. The basic, energetically lowest state is parallel, while the perpendicular or antiparallel configuration has a higher energy- excited state. From the fundamental to the excited state of the magnetic moment, the nuclei pass by absorbing a quantum of electromagnetic energy, which must be exactly equal to the difference in energy between the two states. The respective frequency corresponds to the resonant Larmor frequency. During deexcitation, an electromagnetic signal of the same frequency is then emitted . The precession rotation of the magnetic moment of nuclei in a magnetic field is again just our model idea of how to clearly explain the behavior of nuclei in a magnetic field ...

3.4.4. Nuclear magnetic resonance - simplified schematic representation.
The magnetic moments of the nuclei in the analyte normally have chaotically scattered directions.
By the action of a strong magnetic field B, the mag. moments of nuclei partially orient in the direction of the vector B .
By sending a RF electromagnetic field, these oriented nuclei deviate from the B direction, eg by 90°. After switching off this RF field, a relaxation occurs, during which the deflected nuclei when its return at precession rotation will emit an electromagneic NMR signal with exponentially decaying amplitude.
Simplified schematic diagram of NMR imaging equipment. For simplicity, only one radio frequency (RF) coil is drawn, which electronically switches alternately to transmit and receive modes; usually there are separate transmitting and receiving RF coils.
(ADC = analog-to-digital converter, DAC = digital-to-analog converter) .

Radio frequency coils
RF coils are a kind of "antennas" of the NMR system, that transmit excitation RF signals towards the analyte, or receive response RF signals from the relaxing nuclei in the analyte. In principle, the same coil can be used as the transmitting and receiving coil, which is electronically switched to the transmitting and then to the receiving mode
(as symbolically drawn in the diagram in Fig.3.4.4d). However, better detection of the response NMR signal can be achieved by using a separate receiving RF coil. Due to the relatively high Larmor frequency (tens of MHz), RF coils have a very simple design: they are formed by a loop of wire of circular or rectangular shape, which is placed close to the analyzed material (sample or area of interest in the organism). Sometimes they are suitably shaped (bent into a saddle or cylindrical shape) to achieve better homogeneity of the RF signal in the analyzed area.
  A short but very strong radio frequency alternating current, of high amplitude, is introduced into the transmitting coil in various time sequences, instantaneous power up to tens of kW. In the receiving coil, a response signal is then induced from the relaxing nuclei, on the contrary, with a very low amplitude (of the order of millivolts), which for further electronic processing must be significantly amplified in a narrowband high-frequency amplifier. For NMRI imaging (see below), special RF receiving coils of various sizes and shapes are used to tightly encircle the analyzed area - for imaging the brain, joints, spine, etc.
NMR spectroscopy and analysis
NMR spectroscopy
is performed in such a way, that the frequency of the excitation RF signal gradually increases, this signal intermittently supplies the coils in the transmitting mode, there is always a switch to the receiving mode and the intensity of the RF signal is measured - echo - transmitted by a sample placed in the magnetic field B
o during the back relaxation of the magnetic moments of the nuclei. The frequency at which the resonant maximum occurs, the Larmor frequency, determines the type of nucleus (the highest is for hydrogen - 42.6 MHz for B = 1Tesla), the intensity of the resonant maximum determines the concentration of the relevant atoms in the sample. All nuclei of one isotope, inserted into the same magnetic field, should resonate at the same frequency by themselves. However, if the atoms of these nuclei are part of chemical compounds, the distribution of electrons in their environment differs and these electrons cause electromagnetic shielding of the nuclei. The effective magnetic field acting on the nucleus is then no longer Bo, but B = Bo . (1- s), where the shielding factor s , describing the shielding intensity, slightly depends on the chemical composition of the analyte. This change in the effective magnetic field causes a so-called chemical frequency shift in the NMR signal spectrum .
    Another effect affecting the fine structure of the NMR spectrum is the mutual interaction of the nuclei of neighboring atoms mediated by valence electrons. As a result of these interactions, the splitting of the resonant maxima of the studied nuclei is observed into 2-4 lines at a distance of about 20 Hz - there is a multiplicity of signal .
    Detailed analysis of frequencies, intensities and multiplicities in the NMR spectrum can therefore provide information on the chemical composition and structure of organic and inorganic substances. Modern NMR spectrometers are computer controlled, and the induced NMR signal is analyzed using a Fourier transform .
Relaxation times
After switching off the high-frequency excitation field, the deflected nuclei relax in the magnetic field - they return in a spiral path to the original equilibrium state in the direction B
o (which we refer to here as the "z" axis), which is observed in the receiving coil as a free reverberation of the induced RF signal with an exponential decrease in amplitude. The rate of this relaxation (or fading time) is influenced by the interaction of nuclear spins with surrounding atoms and the mutual interaction between nuclear spins. The NMR signal also encodes information about the surrounding atoms and molecules - information about the chemical composition and structure of the substance. The decay time of the resonant signal is characterized by two relaxation times T1 and T2 .
    Relaxation time T
1 , sometimes called spin-lattice (the name comes from the original use of NMR for the analysis of solids with a crystal lattice) , represents the basic time constant of relaxation of magnetic moments of nuclei from the deflected position to the equilibrium position in the direction of the permanent magnetic field. It captures the speed at which the deflected core releases energy to electromagnetic waves and the environment during relaxation, while the longitudinal magnetization in the z-axis direction returns to the original value of Mo according to the exponential law: MZ = Mo.(1 - e -t / T1) . It is defined as the time, during which the longitudinal magnetization at relaxation reaches (1-e)- times the original value Mo, whereby the signal drops to 63% (if the excitation of the magnetic moment of the core by 90° was performed).
    The relaxation time T
2 , sometimes called spin-spin, expresses the time constant with which, due to the mutual interaction of spins and magnetic moments of adjacent nuclei, leading to the dephasing of the precessional motion of magnetic moments, the magnetization decreases in the transverse direction x-y: M XY = Mo XY e - t / T 2 . T2 is defined as the time during which the transverse magnetization MXY decreases e-times.
    Note: The receiving coil in the MRI actually detects a shorter relaxation time marked T2* after the excitation pulse has ended. In addition to the relaxation time T2, it is caused by a steeper decrease in the transverse component of the material magnetization due to small changes in the inhomogeneity of the magnetic field, leading to desynchronization. In MRI imaging, this phenomenon is usually negative, it can be corrected or eliminated in the so-called "spin-echo sequence" - see below.
    The relaxation times T1 and T2 are the result of the interaction of resonant nuclei with their surroundings and characterize the chemical properties and structure of the investigated material. In medical use, they are often significantly different for healthy and tumor tissue.
  In the most commonly used external magnetic field of 1.5 T, the relaxation times T
1 and T2 of water and some human tissues (in the physiological state) have the following approximate values :

Tissue type:   water   oxygenated blood non-oxygenated blood   fat     muscles   proteins gray matter brain white matter brain   liver     kidneys  
T 1 [ms] 4300 1350 1350 250 880 250 920 780 490 650
T 2 [ms] 2200 200 50 70 50 <= 1 100 90 40 70

Relaxation times are characteristic of different substances and tissues - they depend on the concentration of nuclei, temperature, size of molecules, chemical bonds. It can be seen from the table that, for example, hydrogen nuclei tightly bound in fat or protein molecules relax much faster than protons weakly bound in water molecules.
NMR imaging - MRI
The NMR method originally served as an analytical method for the composition and structure of samples. Advances in electronics and computer technology in the 1970s and 1980s made it possible to use the NMR signal to create an image of proton density in an object under investigation. This created the NMR imaging method (NMRI - Nuclear Magnetic Resonance Imaging; the word "nuclear" is often omitted and the abbreviation MRI is used) - Fig.3.4.4d.
    In order to be able to detect NMR signals separately and locally from individual places of the examined object (organism or tissue) and use it to create an image , it is necessary to ensure spatial-geometric coding of coordinates in the examined object. This can be achieved by superimposing an additional gradient magnetic field in the direction of the X, Y, Z axis on the main constant homogeneous field B
o. These gradient magnetic fields in the direction of each X, Y, Z axis are generated by a respective pair of gradient coils. By changing the gradient magnetic field, we achieve that the magnetic resonance will always occur in a different place of the examined object. By this gradient magnetic coding of spatial coordinates we can then perform NMR imaging.
Gradient coils 
are "additional" electromagnets located in suitable places inside the main strong electromagnet. They are wound with copper wire or metal tape, dimensioned for relatively high currents of tens or hundreds of amperes. Gradient coils are supplied in pulse sequences with a relatively strong current (approx. 500A) from electronically controlled sources, which allow fast and accurate setting of the strength and direction of the excited magnetic field - an additional gradient field. They produce gradients in the range of about 20-100 mT /m. In order for MRI imaging not to take an enormously long time, the rate of gradient changes needs to be relatively high - it reaches about 100-200 Tm-1 .s-1; it requires a certain voltage (approx. 50-300V) to overcome the inductance of the gradient coils - the power supplies of the gradient coils are relatively robust (power). Strong current surges in the gradient coils when interacting with the magnetic field cause mechanical vibrations, which causes considerable noise during MRI. Longitudinal gradient coils (in the Z direction ) have turns wound in the same direction as the main coil, X (gradient in the left-right direction) and Y (gradient in the up-down direction) are formed by saddle-shaped coils with vertically wound turns.
    Note first the longitudinal gradient field Bz(z) in the Z direction. His superposition with the main mag. field Bo causes the actual "local" value of the magnetic field B = Bo + Bz(z) to depend on the z coordinate : B = B(z). If we send a high-frequency pulse of a certain frequency f to a sample placed in this slightly inhomogeneous gradient magnetic field, the magnetic resonance signal will be transmitted by atomic nuclei only from a thin layer of the sample with coordinate z , for which the resonance condition f = g .B(z) /2p is satisfied. By varying the frequency f of high-frequency excitation pulses, or the intensity of the longitudinal gradient field Bz, is changes the position of the layer, in which the magnetic resonance response signal is generated. In this way, information about the dependence of the spatial distribution of the density of the nuclei in the direction of the Z axis is captured - the electronic-geometric coding of this coordinate is achieved - the layer z .
  The representation of the spatial distribution of the density of nuclei in a given layer z in the transverse directions X and Y is then obtained by the action of another, transverse, gradient magnetic field in the direction of the X and Y axis, whereby the investigated layer decomposes into elementary volumes - "pixels", in which is determined intensity of the relaxation NMR signal, and also its decay times. By changing these gradient fields, data are obtained for individual sites in the z layer and their computer reconstruction yields a cross-section image of the proton density in the examined layer z (Fig.3.4.4d right). By electronic analysis of relaxation times of the NMR signal is also generated cross-sectional images in the relaxation times T
1 and T2 (referred to as T1 or T2 - weighted images). The set of cross-sectional images for different values of the z-coordinate then creates a 3-dimensional tomographic image of the investigated area in proton density and relaxation times in individual "voxels". Using computer graphics, it is then possible to create images of any sections of the examined area, which are brightly modulated in a wide range of shades of gray (from white to black), to distinguish the structure of tissues and organs.
    The basic subject of NMRI imaging is hydrogen nuclei - imaging of proton density and relaxation times. This is why NMRI is sometimes referred to as "hydrogen topographic imaging". The intensity of such an NMR image mainly reflects the amount of water at each locationin the examined tissue and the nature of the binding and distribution of hydrogen molecules in the cells and extracellular space, as well as the distribution of fat and proteins. Based on these structural differences, different tissues can be distinguished from each other in MRI images - such as water, muscle, fat, gray matter and white matter in the brain.
    In general, two basic information is captured locally in NMRI images :
1. Density distribution of nuclei producing nuclear magnetic resonance - most often the proton density PD of hydrogen in the tissue. PD images essentially capture the anatomical structure of tissues and organs, and are largely similar to CT X-rays, which map the electron density of tissues.
2. Distribution of relaxation times T
1 and T2 related to the chemical composition and structural state of the tissue in individual places. Such images are called T1 and T2 - weighted .
    About to what extent to which the proton density PD and the times T
1 and T2 will be represented in the resulting MRI image, - how and with what this image will be modulated - "weighted" - is determined by pulse sequences: time sequence of transmitted RF pulses and "echo" response signals (will be discussed in more detail below) .

Fig.3.4.5 MRI images of the brain (transaxial section, without pathology) in proton density, relaxation times T 1 and T 2 and in a special FLAIR sequence to suppress the water signal.
(MRI brain images were taken by Jaroslav Havelka, MD, head of the MRI RDG department at the University Hospital Ostrava )

Proton densities and especially relaxation times are different not only for different types of tissues (see table above), but also differ depending on the physiological or pathological condition of the same tissue. This makes NMRI imaging an important diagnostic method in medicine, including in the field of tumor diagnostics.
Note: As with X-ray diagnostics, NMRI also uses contrast agents to increase the contrast of images of certain structures (eg cavities or blood vessels), but not on a density but on a magnetic basis - ferromagnetic compounds, mostly based on gadolinium . Pulse sequence in NMRI
In medical MR imaging, it is desirable to create images with sufficient high contrast between different tissue types so that the MRI radiologist can best answer the clinical diagnostic question. Optimal image contrast between different tissues with different densities and rexation times can be achieved by suitable excitation of magnetic moments of nuclei and subsequent measurement of their response MR signal: by setting parameters of pulse sequence - time sequence of transmitted electromagnetic excitation pulses RF
and subsequent measurements of the "echo" of the electromagnetic signal emitted by the relaxing nuclei. The first parameter here is the intensity (energy) of the transmitted radio frequency excitation pulse (RF), which determines the predominant angle of deflection (tilt) of the magnetization vector of the analyzed nuclei - 90° or 180°. The higher the excitation intensity radiated into the analyzed target tissue, the higher the percentage of reversal of the magnetic moment of the nuclei and the stronger the response signal and more time is required for relaxation. Another parameter is the time interval TR , in which we repeatedly apply individual radiofrequency excitation pulses. The shorter this interval, the less time there is for T1 relaxation. The third parameter is the time TE (echo time) between the excitation pulse and the detection of the response resonant signal. The longer this time, the less nuclei with a shorter relaxation time T2 will contribute to the measured resonant signal. The completely approximate values of the pulse sequence times for obtaining the basic types of MRI images at B = 1.5 T are :
PD: TR = 1000 ms, TE = 5-30 ms; T 1 -weighted: TR = 10 ms, TE = 5-30 ms; T 2 -weighted: TR = 1000-2000 ms, TE = 80-100 ms.
    In connection with these regularities, several significant sequences of transmission of excitation radiofrequency pulses and subsequent detection of response relaxation signals have been developed (sometimes called "MRI techniques" in MR jargon ) :   
-> Saturation - recovery sequence in which 90° RF pulses are transmitted at regular intervals. Upon arrival of each RF pulse, the magnetization vector rotates 90° and relaxation begins with different times T1 in different tissues. When another RF pulse arrives, the z-component of the magnetization will be different in different tissues. With a suitable repetition period TR of excitation RF pulses, we can set the optimal contrast of the desired tissues at times T1 . This simplest MRI technique is now rarely used, it has been replaced by the inversion-recovery sequence below, providing higher contrast.
-> Spin - echo sequence consisting of a 90° RF pulse followed by a 180° RF pulse. After the magnetization vector has been flipped into the xy plane due to a 90 ° RF pulse, T 2 (resp. T2 *) relaxes, during which phasing occurs. However, the subsequent 180° RF pulse has a "refocusing" effect - it flips the individual spins in the xy plane by 180 ° and the spins are phased again. The result is an echo signal in the receiving coil, the amplitude of which depends on the relaxation times T 1 and T 2 of the tissue (unfavorable T2 * does not apply here, because the effect of magnetic field inhomogeneity on phasing is eliminated by 180 ° pulse phasing) . The character and contrast of the display can be adjusted using the times TR and TE. With short TR and short TE we get T 1-weighted image, long TR and short TE provide a proton density image, long TR and long TE provide a T 2 -weighted image. Due to this variability of imaging options, spin-echo is the most commonly used MRI technique.
-> Inversion - recovery sequence, consisting of a sequence of 180 ° and the following 90 ° RF pulse. The initial 180 ° pulse inverts the magnetization vector to the opposite, after which T 1 relaxation takes place . With a time interval TI - inversion time , a 90 ° RF pulse then follows, which flips the magnetization vector into the xy plane. A RF signal dependent on T 1 is detected in the receiving coilrelaxation time of the displayed tissue. The contrast of the image can be adjusted appropriately using the TI time. A significantly more contrasting image can be achieved than with the saturation recovery technique.
By a special setting of the time T1 = T
1 .ln2, the suppression of the image of the tissue having this relaxation time T 1 is achieved . By setting the short inversion time TI (approx. 140ms with a 1.5T magnet) - the so-called short time inversion recovery STIR - the suppression of the fat signal is achieved in the image . Conversely, by extending the time TI (to about 2600ms) - fluid attenuation inversion recovery FLAIR - we can achieve suppression of the water signal. Other fine details and anomalies in the structure of the examined tissues can then be better assessed on such "cleaned" images.
-> Gradient - echo sequence begins with a 90 ° RF pulse (which tilts the magnetization vector to the xy plane), after which a magnetic field gradient is applied. The nuclei in adjacent atoms will thus show a precession with a slightly different Larmor frequency, which will cause spin phasing. The application of the second mag. gradient with the opposite sign, which rephases the spins and at this point the echo is measured. Used to obtain a T 2 -weighted image.
-> .......... sequence ............ ? add more sequences? ........... ? complete the picture of the graphic sequence diagram? ...
    Computer analysis of MRI images obtained with appropriate sequences
(mentioned above) can create special image modulations - such as water or fat signal suppression images . Other special sequences are used for functional MRI (mentioned below) :
-> Susceptibility weighted imaging ( SWI ) shows tissues with slightly different magnetic susceptibility. It uses an extended gradient-echo sequence for display in T 2 * . Its main variant is Blood oxygenation level dependent (BOLD) , see fMRI below .
-> Diffusion weighted imaging (DWI) shows the diffusion of water inside tissue elements, manifested by Brownian motion of molecules. Using a spin-echo sequence with the application of 2 gradients, a subtle effect is registered, in which Brownian-moving water molecules show a different phasing-phasing relationship when reversing the mag. gradient; this leads to a slightly weaker T 2 signal.
MRI Magnetic Resonance Spectrometry MRI
Magnetic resonance imaging (MRS) can be supplemented by the magnetic resonance spectrometry (MRS) described above, which enriches this examination with additional physiological information . Chemical analysis is performed here by analyzing the chemical shift of the Larmor frequency imaging structures in-vivo, eg choline or lipid levels. Chemical shifts are very fine, so this method is demanding not only in terms of signal analysis, but also requires high intensity (recommended at least 3 T) and homogeneity of the magnetic field.
Functional magnetic resonance imaging - fMRI 
Magnetic resonance imaging may be a suitable method for non-invasive imaging of the function of various tissues and organs (along with "molecular" imaging in nuclear medicine - .....). So far, fMRI has found application mainly in functional brain imaging, mapping neuronal activity . Neurons
(which do not have internal energy stores) they need to get sugar and oxygen quickly for their increased activity. The hemodynamic response to this need causes an increase in blood perfusion at a given site, but mainly a greater release of oxygen from the blood than inactive neurons. This leads to a change in the relative levels of oxygenated oxyhemoglobin and non-oxygenated deoxyhemoglobin in the blood at sites of neuronal activity.
    In this respect, two basic methods of indirect mapping of neuronal activity are used :
- Local increase of perfusion at the site of increased neuronal activity - perfusion fMRI .
- Change in the ratio of oxygenated and non-oxygenated blood at the site of neuronal activity. The method is called BOLD fMRI (Blood Oxygen Level Dependent). Changes in the relative levels of oxy- and deoxy-hemoglobin can be detected based on their slightly different magnetic susceptibility. Basic hemoglobin without bound oxygen (deoxyhemoglobin) has slightly paramagnetic properties, but when oxygen is bound to it (oxyhemoglobin), it behaves slightly diamagnetically. If more deoxyhemoglobin accumulates at a certain site in the brain tissue, a slightly stronger MRI signal is obtained from it than from the sites where deoxyhemoglobin predominates.
  MRI functional imaging of the brain is performed after neurological activation, either motor
(eg movement of fingers) , visual, linguistic or cognitive.
The physical-electronic implementation of NMRI 
NMR imaging isthe most complex imaging method. The operation of the device for NMR imaging is electronically very complicated and demanding, so it must be controlled by a powerful computer with sophisticated software - Fig.3.4.4d. In the multiplex mode, the process of transmitting a sequence of radio frequency pulses, modulation of gradient magnetic fields, sensing and analysis of relaxation signals of magnetic resonance, reconstruction and creation of the resulting images, as well as a number of other transformation and correction procedures are synchronously controlled. Since these are harmonic (sinusoidal) waveforms, scanning and reconstruction are performed using Fourier analysis - in the frequency so-called K-space. It is a set of matrices defined in the memory of the MRI evaluation computer, into the individual elements of which the frequencies, amplitudes and coordinates of MRI signals are recorded. From these "raw" data, the resulting MRI images are created using Fourier transform and other analytical methods.
    Note: Electron paramagnetic resonance (EPR) is based on a similar principle as NMR . The magnetic moments of the electron shells of atoms are used here .........

3.5. Radioisotope tracking methods
Radioisotope tracking or indicator methods are used to monitor the hidden movement and distribution of matter within physical, chemical or biological systems, or in various technological devices. A suitable "labeled" substance with bound radionuclide is introduced into the system - the so-called radioindicator, whose movement and behavior in the system is then monitored on the basis of detection of ionizing radiation emitted during radioactive transformations of nuclei in the radioindicator. The movement and distribution of the radio indicator can be monitored in two basic ways :

Radioisotope tracking methods are used in many fields of science and technology, industry, agriculture and especially medicine. Here we will briefly mention a few technical and general biological applications, we will focus in more detail below on applications in nuclear medicine.
    Radioisotope tracking methods were first tested in 1913 by the chemist G.Hevesy, who found that radioisotopes have the same chemical behavior as stable isotopes of the same element. However, unlike stable isotopes, radionuclides can be "visible" through the penetrating radiation generated by the transformation of nuclei.

Radioisotope scintigraphy and nuclear medicine
Nuclear medicine
is a field dealing with diagnostics and therapy using radioactive substances in open form, applied to the internal environment of the organism.

  Radioactive isotopes react chemically in the same way as stable isotopes of the same element - therefore they behave in the organism's metabolism in the same way as non-radioactive isotopes of a given element. However, due to the fact that radioactive isotopes are "visible" through penetrating radiation, which arises during radioactive transformations of their nuclei, it is possible to monitor the movement and metabolism of elements and compounds containing radionuclides - radioindicators - in the body and thus investigate the functions of individual organs. Depending on the organ whose function is to be examined, the specific substance (radiopharmaceutical) shall be labeled with an appropriate radioisotope. After application to the body, the movement and metabolism of this substance is monitored - mainly imaging with a gamma camera, possibly supplemented by measurement of samples (blood or urine).
Radioisotope diagnostics in vivo - scintigraphy
In radionuclide diagnostics in vivo in nuclear medicine, the patient is administered (usually intravenously, sometimes orally or by inhalation) a small amount of a suitable
g - radioactive substance - the so-called radioindicator or radiopharmaceuticals. The radioindicator used is specific to individual organs and types of examinations. The applied radioactive substance enters the metabolism of the organism and is distributed there according to its chemical composition - physiologically or pathologically it accumulates in certain organs and their parts and is subsequently excreted or regrouped. Gamma radiation emanates from the deposition sites of the radioindicator, which, due to its penetration, passes through the tissue out of the organism. Using sensitive detectors, we measure this radiation g and thus determine the distribution of the radioindicator in individual organs and structures inside the body.
    The most perfect devices of this kind are gamma cameras (scintillation cameras) - using them we display in radiation
g the distribution of the radioindicator in the organism. This method, called scintigraphy, thus makes it possible to obtain not only anatomical information, but mainly to tell about organ functions and metabolism. By mathematical evaluation of scintigraphic images, we can obtain curves of the time course of the radioindicator distribution and calculate dynamic parameters characterizing the function of the relevant organs.

Schematic arrangement of the entire process of scintigraphic examination - from the application of a radioindicator to the patient, through the process of scintigraphic imaging with a gamma camera, evaluation, mathematical analysis and quantification, to the interpretation and determination of diagnosis.

The tomographic gamma camera SPECT (Single Photon Emission Copied Computerized Tomography) slowly rotates around the patient's body, scans scintigraphic images from various angles and then uses computer reconstruction to create cross-sectional images (sections perpendicular to the camera's axis of rotation), from which computer graphics can be used to construct spatial (3-dimensional) images of the distribution of the radio-indicator in the organs inside the body.
    The PET gamma camera (Positron Computerized Tomography) detects photons of gamma annihilation radiation (511 keV energy) flying in opposite directions during the annihilation of positrons, emitted by
b+ radioindicator administered to the patient. These photons of annihilation radiation are coincidentally detected by an annular scintillation detectors, and by computer reconstruction of the line projections of the coincidence sites, images of cross-sections are generated and, if necessary, 3D images similar to SPECT.
    Nuclear medicine provides specific methods for the examination of virtually all organs and thus cooperates with a wide range of clinical disciplines. The most widespread use is mainly in  cardiology, nephrology, neurology, oncology, thyrology, gastroenterology.
    Nuclear medicine methods are among the least burdensome non-invasive diagnostic examination methods. Due to the high sensitivity of the detectors, only a very small amount of radiopharmaceutical is applied to the patient, which is needed to obtain quality image information. The radiation exposure in methods in nuclear medicine is comparable (and often smaller) as in X-ray examinations *).
*) During X-ray examination, the source of ionizing radiation is a device (X-ray tube) and the radiation dose depends, among other things, on the number of images performed, or on the extent of the area scanned during CT. In scintigraphy, the source of radiation is not a diagnostic device, but the patient himself, resp. its investigating body. Thus, we can take any number of scintigraphic images without changing the radiation exposure of the patient.
  Radionuclide scintigraphy is described in detail in Chapter 4 "
Radioisotope scintigraphy".

Radiation-guided surgery - sentinel nodes
An important radioisotope tracking method of nuclear medicine is local radiation measurement with a closely collimated miniature gamma-ray detection probe in radiation-guided surgery in the detection of so-called sentinel nodes. In the surgical treatment of cancer, it is important to remove not only the primary tumor, but also, if possible, other tissues into which the tumor cells could be infiltrated. These tumor cells spread from the primary site mainly through the lymphatic pathways, so that the lymph nodes around the tumor site are the first to be affected. If we apply a suitable radioindicator of colloidal state to the peripheral part of the tumor lesion (most often
99mTc nanocolloid, particle size approx. 50-600 nm, activity approx. 40-150 MBq), it will propagate through the lymphatic pathways and capture and accumulate in those nodes, that are lymphatically associated with the tumor site. The first such node in the lymphatic "watershed" of a tumor foci is called the sentinel node. The accumulation of the radioindicator in the nodes can be displayed scintigraphically. However, the most important thing is to monitor the radioindicator during the actual surgical procedure, when using a collimated detection probe, the surgeon can find a sentinel node containing the radioindicator directly in the operating field.
*) Along with the radioindicator, a blue dye is applied at the same time, which also penetrates the nodes, so that the surgeon can recognize the sentinel node also by its blue coloration.
    After application of the radioindicator, scintigraphic imaging is performed with the displayed nodes marking, then the patient goes to his own surgery, during which a detection gamma probe is used both for perioperative sentinel node detection and for radioactivity detection in an already operated node. This is followed by histological examination of the sentinel node to classify the type of tumor, which will help optimize the further course of therapy. ......"...."......

In vitro diagnostics. Radioimmunoassay - radiosaturation analysis
In nuclear medicine, in vitro radioisotope diagnostic methods are also used, where (non-radioactive) samples taken from patients are analyzed using radioisotope techniques - radiochemical and at the same time biochemical. Most often it is a radioimmunoassay (RIA) or radiosaturation analysis (RSA), which is used to highly sensitively determine the concentration of complex biological substances in the blood serum - hormones, tumor markers and other biologically important substances
. It is based on an immunochemical reaction antigen with a specific antibody (Ab - antibody). A competitive immunoreaction is used in which the radiolabeled Ag* antigen "competes" for binding sites on the antibody (which is present in a limited amount in the reaction mixture) with the unlabeled antigen. An appropriate antibody labeled with the appropriate radionuclide Ag* (usually I-125-radioiodine) is added to the sample analyzed, which reacts with given hormone to form an insoluble Ag*-Ab complex. After removal of the unbound fraction (rinsing with water), a compound remains in the sample, the activity of which will depend on the concentration of hormone in the analyzed sample - the amount of labeled antigen-antibody Ag* -Ab complex is inversely proportional to the concentration of antigen to be determined. The more test substance present in the primary sample, the smaller the amount of labeled Ag* -Ab antigen-antibody complex formed and the lower the activity in the final sample. It is measured in a well scintillation detector (see §2.7, section "Automatic measurement of a series of samples") . These methods are of great importance for endocrinology.
  RIA or RSA methods reached their greatest development in the 1970s and 1980s, when they were widely performed in radiochemical RIA laboratories at the departments of nuclear medicine. Then they gradually moved from nuclear medicine workplaces to clinical biochemistry laboratories. Since the 1990s, they have been gradually extruded and replaced by fluorescent and chemiluminescent optical methods, without the use of radionuclides and ionizing radiation...

Radioisotope therapy
In addition to diagnostics, nuclear medicine also includes therapy with open radionuclides, eg treatment of hyperthyroidism and thyroid cancer, blood diseases, palliative and curative therapy for various types of tumors, joint diseases - see below for more details in §3.6 "Radiotherapy", part "
Radioisotope therapy".

Nuclear medicine - interdisciplinary field
Nuclear medicine, due to the physical nature of its methods and instrumentation used, is an interdisciplinary field. Besides doctors
(specialist and certified in the field of nuclear medicine), nurses and laboratory technicians, are working in team work as well as experts from other professions - physicist, electronics, radiochemicist, pharmacist. Along with medical and physical-technical aspects, considerable attention is also paid to radiation protection of workers and patients in the workplaces of nuclear medicine when working with radioisotopes (see Chapter 5 "Biological effects of ionizing radiation. Radiation protection").

A detailed description of the principles, methods and clinical use of nuclear medicine is in Chapter 4
Radionuclide scintigraphy " .

- photographic imaging of the distribution of the beta-radioindicator in the examined preparations in close contact of the photographic emulsion with sample is described in §2.2 "Photographic detection of ionizing radiation", passage "

3.6. Radiotherapy
Radiotherapy is a physico-medical field using the biological effects of ionizing radiation for therapeutic purposes. The vast majority of it is a therapy of tumor diseases, cancer - radiation oncology, to a lesser extent, some degenerative and inflammatory disorders are treated with radiation. Recently, so-called radiosurgery has sometimes been used, especially for vascular and neurological malformations
(see the "Stereotactic radiotherapy" section). Before we focus on our own radiotherapy, we will mention some biological aspects of cancer, diagnostics and non-radiation therapeutic methods (chemotherapy, biological therapy) - to put the issue in a broader context.

Origin of tumors and their classification Diagnosis of cancer
Chemotherapy and biological treatment Radiotherapy - physical and radiobiological factors
Basic methods of radiotherapy, gamma irradiator, MLC Radiotherapy planning, DVH
Modulation of irradiation beams, IMRT, IGRT Stereotactic radiotherapy, gamma-knife, CyberKnife
Hadron radiotherapy - proton, C-14 Brachyradiotherapy
Biologically targeted radoisotope therapy with open beta and alpha radionuclides

Tumors - their nature and origin
Higher organisms, including us humans, consist of billions of cells of various types and functions in a variety of tissues and organs. These cells have a limited lifespan, they disappear after a certain period of time
(mostly by apoptosis - §5.2, passage "Mechanisms of cell death"), while they are mostly replaced by new cells, created by the division of existing cells. The process of cell death and division is under normal circumstances rigorously controlled to ensure tissue homeostasis. During their lifetime, cells are often exposed to various harmful influences from both the outside and the inside. The resulting small damages are mostly repaired by intracellular repair mechanisms, and the cells repaired in this way can then perform their original functions again (see §5.2, passage "Repair processes"). In case of severe damage, the cell usually dies. However, it can sometimes happen that "moderate" damage is not fatal for the cell, nor can it be repaired flawlessly; such a cell can continue to divide, but with disturbed genetic information in the DNA - with a mutation. This, under certain circumstances, can lead to a violation of the control of the process of cell division and ultimately result in the pathological uncontrolled multiplication of cells - the emergence of cancer.
    Tumors, especially malignant, are among the most common and most serious diseases, threatening the health and lives of patients. During the formation of a tumor, pathological tissue mass (neoplasm) is formed, usually irreversible, in which uncontrolled proliferation of tumor cells takes place, at the expense of healthy tissue; there is no feedback in the body to stop this growth. Tumor cells can grow into the surrounding tissue and migrate through lymphatic or blood vessels to other parts of the body (establish metastases). With its uncontrolled division of the mass of tumor cells, it suppresses the surrounding healthy tissue, disrupts it and can thus violate the function of important organs. The cause of such a condition it is not exactly known
*), it lies deep inside the cell structure, probably in mutational changes in DNA. Prevention and causal treatment of tumor diseases - cancer - is therefore difficult.
*) Only some risk factors were observed, which contribute to the formation of tumors or increase the probability of their occurrence. They are various chemical substances, so-called carcinogens, such as some cyclic hydrocarbons and cigarette smoke, the composition of food. Or biological effects - some viruses (so-called oncoviruses ), whose RNA can (via so-called reverse transcriptase) enter the DNA of eukaryotic cells and alter their genetic information, they can cause tumor transformation of cells (or they may not lead directly to tumor formation, but prevent an immune response that would be able to recognize tumor cells and destroy them). Then there are genetic factors , hereditary predisposition (hereditary genomic imprinting eg in Nyemegen syndrome, Ataxia teleagiectasia, Bloom's syndrome, Fanconi anemia, Xeroderma pigmentosum, Li-Fraumeni syndrome of the mutated TP53 gene encoding p53); it is caused by a specific mutation in the tumor suppressor genes (TP53 gene mutation in Li-Fraumeni syndrome, NF1,2 in neurofibromatosis, BRCA1,2 in hereditary breast and ovarian cancer, APC in colorectal polyposis, WT1 and Wilms' kidney tumor, RB1 in retinoblastoma, MLM gene mutation in malignant melanoma). Of the physical influences, it is ultraviolet radiation acting on the skin and especially harder ionizing radiation, as discussed in detail in §5.2 "Biological effects of ionizing radiation".

Is cancer a new civilizational disease (which did not occur in the past) ?
This opinion is often found among people, but it is wrong. Cancerous diseases undoubtedly began to appear immediately after the emergence of multicellular more complex organisms with specialized cells for different functions. Even then, some cells with unrepaired genetic damage in their DNA "forgot" what function they should perform in the organism and could begin to multiply uncontrollably.
  Most cancers affect soft tissues, so their remains cannot be found in paleontological and archaeological findings. Only bone tumors leave distinct fossil traces. It was possible to find a sample of turtle bone approximately 240 million years old, destroyed by bone cancer. In human ancestors, tumor destruction of bone was found in an approximately 2 million year old Australopithecus bone. In historical times, cancer was already described in ancient Egypt. However, until the 19th century cancer was considered relatively rare. People then suffered and died from many other diseases. Today we are able to cure most of them
(some have been practically eradicated). But tumor diseases, especially at a more advanced stage, are still difficult to treat even now.
  The increased incidence of cancer in recent times is due to more perfect diagnostics, but probably unfortunately also to the worse lifestyle of most people in a consumer society.

Carcinogenesis - the formation of tumors
Under normal circumstances, a multicellular organism is a system of individual tissues and organs, consisting of a large number of cells, performing their function in a harmonious community for the benefit of the whole organism. This cooperation and "social behavior" of cells is ensured by very intricate and complex regulatory processes, including signals from monitoring the external environment, transmission of signals to the internal environment of the cell, cellular response, evaluation of signals and their coordination with other signals. Cells that do not fit into this regulatory mechanism (due to damage or loss of their function) are eliminated by the mechanisms of "programmed" cell death, apoptosis. Also, cell proliferation is precisely controlled to meet the needs of the tissue and the organism - dynamic balance, tissue homeostasis. Disruption of regulatory mechanisms can cause various pathological conditions and diseases of the body. One of them is a violation of the regulation of cell division: in some cell a genetic change (mutation) occurs that allows it to survive, divide and produce daughter cells that do not "listen" to the regulatory mechanisms of tissue homeostasis. This can create a gradually expanding population of mutated cells - a clone of tumor cells, multiplying at the expense of healthy tissue.
    Tumor formation (carcinogenesis) is a complex multi-step process in which several mutations gradually accumulate, which do not harm the altered cells, but on the contrary favor them and allow their rapid division regardless of the needs of the organism. The following factors are important for the origin and development of cancer :
-> Cell cycle deregulation
To maintain tissue homeostasis (a balanced number of functional cells of a given tissue) it is necessary to control the rate at which cells form, develop and die - cell cycle regulation. Due to some mutations, autonomic growth (mitogenic) factors in the cell, their autocrine production, or loss of sensitivity to signals that stop the cell cycle may occur. In particular, p21 and p27 proteins are involved in cell cycle regulation and division, which may bind to and inhibit the activity of certain kinases (serving as regulators of DNA replication and cell division); the content of these regulatory proteins in tumor cells tends to be reduced. On the contrary, the activity of some kinases
(§5.2, section "Cells - basic units of living organisms", passage "Proteins, enzymes, kinases") is increased in tumor cells. It is mainly a tyrosine kinase - an enzyme which transfers phosphate to the hydroxyl group of the cyclic amino acid tyrosine, thereby affecting the function and activity of the respective protein. Increased epidermal growth factor receptor (EGFR) tyrosine kinase activity leads to increased intracellular signaling, disrupting cell cycle regulation. This increased EGFR activity may be due to increased expression of the EGFR ligand (which is the epidermal growth factor EGF), or by a mutation in the EGFR tyrosine kinase domain that results in sustained ligand-independent activation of the mutated receptor. Another receptor which promotes cell growth and division, the epidermal growth factor receptor HER2 (Human Epidermal Receptor), also called erbB2, in the increased presence of which there is an excessive division of cells, which can lead to the formation of a tumor. Also, deregulated signal transduction through the P3K / Akt / mTOR phosphatidylinositol-3-kinase signaling pathway provides cells with stimuli for unrestricted growth and survival, which can lead to tumor growth.
    Disruption of cell cycle regulation can lead to the proliferation of such altered cells independently of the environment, independent of the needs of the tissue and the organism - it is usually the first stage of carcinogenesis.
-> Inhibition of apoptosis
Another important mechanism for maintaining tissue homeostasis is the regulation of the rate at which "excess" cells in the tissue population die. The usual way in which cells undergo controlled death is apoptosis
(described in more detail in §5.2, section "Effect of radiation on cells", passage "Mechanisms of cell death", where the internal and external signaling pathways of apoptosis are discussed) - "programmed" cell death, actively controlled by the cell . The apoptotic program is potentially present in all cells, it is triggered by internal signals (DNA damage, hypoxia, ...) or external "death signals" that the cell receives from regulatory mechanisms in the tissue. Properly functioning apoptosis, triggered by external regulatory mechanisms from the tissue, provides effective protection against excessive cell proliferation. Apoptosis triggered by internal mechanisms then acts as a protection against the survival and proliferation of mutated cells with damaged DNA. Inhibition (blockage, damage) of apoptosis - apoptic resistance of cells - allows developing tumor cells to survive and multiply, despite the organism's interest. Apoptosis can be disrupted, for example, by altering the TP53 gene encoding the p53 protein, increasing the concentration of the anti-apoptotic Bcl-2 gene in mitochondria (protecting mitochondrial membranes - preventing cytochrome c penetration and caspase chain triggering proteolytic degradation in the cytoplasm) and other unexplored factors.
Note: This is mainly a suppressed apoptosis induced by external regulatory mechanisms of the tissue. However, with strong irradiation of tumor cells, which causes severe irreversible DNA damage, internally activated apoptosis occurs .
-> Immortilization of cells
Most common somatic cells can
only reach a certain limit in the number of their divisions, the so-called Hayflick limit (about 40-60 cycles); then the cells lose their ability to divide. This is due to mitotic shortening of DNA telomeres (about cell cycle, telomere shortening, cell senescence, apoptosis, etc. see also §5.2 "Biological effects of ionizing radiation"). This limit in the number of divisions would automatically stop the growth of the tumor population. Increased occurrence of an active enzyme called telomerase (in co-production with tankyrase), or mechanisms of homologous recombination of telomere sequences (discussed in §5.2, part "DNA, chromosomes, telomeres"), however, they are able to ensure complete replication of DNA ends (prevent telomere shortening) - gaining unlimited replication potential, so-called immortilization -"immortality"of cells. Unregulated and unrestricted division of clonogenic tumor cells, which break free from the mechanisms of tissue homeostasis, is a typical feature of tumor diseases.
-> Inhibition of immunogenity
The immune system is fundamentally capable to recognize tumor cells and destroy them. Some tumor cells, however, they lose their immunogenicity *) or the immune system is impaired - these cells are outside the control of the immune mechanisms and may cause their uncontrolled proliferation.
*) Many tumor cells have the CD47 protein on their surface, which protects them from white blood cells (physiologically, this protein occurs on the surface of blood cells to protect them from its own white blood cells) and therefore cannot be destroyed by the immune system.
-> (neo)Angiogenesis
In the initial stages (tumor size up to about 0.5-1 mm), tumor cells are supplied with oxygen and nutrients by diffusion from the surrounding intercellular environment of the tissue in which the tumor grows. As the number of tumor cells increases, this supply is no longer sufficient - there is hypoxia of the tumor tissue, accompanied by the expression of special interleukins, especially HIF1 (hypoxia-iducible transcription factor), inducing the production of vascular endothelial growth factor VEGF (as well as VEGF mRNA expression). The regulatory mechanisms in the tissue can respond to this by angiogenesis - the formation of new blood vessels, ensuring the blood supply to the tumor tissue and its rich supply of oxygen and nutrients, as well as flushing out metabolic waste. Tumor growth is dependent on angiogenesis - it necessarily requires a sufficient supply of nutrients and oxygen, which are provided by newly formed blood vessels. Each increase in tumor volume is associated with the growth of new capillaries.
    Tumor neo-angiogenesis or neovascularization is an important milestone in the progression of cancer, allowing tumors to grow to macroscopic size, threatening tissue and whole organism. Blood flow to the tumor tissue also allows the spread of tumor cells through the bloodstream - the formation of metastases :
-> Formation of metastases
In most cancers, the most dangerous feature of tumors is the ability to create secondary foci, so-called metastases (Greek: meta=change, stasis=place -> change of place, relocation). Metastases are the cause of more than 90% of cancer deaths. Tumor cells can grow into the surrounding tissue, separate from the original tumor and migrate to other parts of the body through the lymphatic or blood paths, settle there and create secondary foci - establish metastases. The tumor cell first penetrates into the surrounding tissue and then into the blood or lymphatic stream. If it survives in this watershed, it can be carried by it and migrate to other tissues and organs. There it can leave this riverbed and penetrate the tissue in a new place, establish itself there and begin to divide. Metastasis is often a chain: from the secondary metastasis, tumor cells are released and migrate and establish other, tertiary, metastases, etc. The result can be the spread of cancer throughout the whole body - generalization.
   The issue of carcenogenesis is very complex, with a number of unexplored factors. In addition to mutations of known and coding genes in DNA, so-called "genetic litter" or "unnecessary DNA" sequences can also may manifest, which via the respective RNA can function as regulatory "triggers" or "switches" of intracellular processes (see also §5.2, part "DNA, chromosomes, telomeres") .
    So to sum it up briefly, cancer occurs when some (originally healthy and functional) cells in the body undergo several mutations in their DNA and begin to divide uncontrollably, and at the same time the body's immune mechanisms unable to stop this growth. The occurrence of cancer is a stochastic phenomenon, it is not possible to specifically predict if and when it will occur in a given individual; the probability of its occurrence can only be estimated statistically.
Dependence on the age and size of the organism ?
The mechanisms of carcinogenesis and the stochastic occurrence of cancer lead to the expectation that the more cells an organism has in its body and the longer it lives, the greater the risk of developing cancer. The time factor of the growth of cancer risk with the age of a specific individual is indeed almost always applied - the longer someone lives, the more time their cells have to mutate; and at the same time the immune processes also weaken during aging. In dependence on the size of the organism, however, the situation is more complicated. Some large and long-lived animals hardly suffer from cancer, even though they have many more cells and live as long or even longer than humans. They are, for example, bowhead whales, elephants, but also cows and horses. On the other hand, some small animals such as mice and rats, with a relatively short life of about 4 years, are highly susceptible to tumors. One of the factors behind this is faster metabolism (
-> formation of a larger amount of mutagenic metabolic products) and faster cell division (-> more frequent mitotic reproduction of potential errors in DNA). In large and long-lived animals, which usually also have a slower metabolism, also genetic factors were found, consisting of an increase in the TP53 gene, which codes for the p53 protein that triggers apoptosis, or in general a faster development of tumor-suppressing genes. There is a vague hope, that some of these molecular-genetic anti-tumor strategies will be artificially applied in clinical oncology in the future ..?..

Types of tumors
Tumors and tumors disease
(lat. tumor = swelling, edema), often collectively referred to as carcinoma or cancer, are characterized by a large number of species and great variability. They are divided according to several criteria :
l According to health severity :
Benign tumors (lat. Benignus = harmless, friendly, generous ) are usually localized and isolated from the surrounding tissue by encapsulation, do not grow into other tissues and do not form distant metastases. They do not have to create major damage or difficulties for the organism (they can often remain in the tissue - but beware of the risk of malignant degeneration!), if necessary, they can usually be successfully surgically removed.
Malignant tumors (lat. malignus = evil, evil-bearing ), grow destructively and infiltratively - tumor cells grow into the intercellular spaces of the surrounding tissues (which suppress and disrupt), the cells are released and spreads through the blood or lymphatic route to other tissues and organs, where they often form distant secondary "daughter deposits", so-called metastases (Greek meta = change, stasis = place ® change of place, relocation). Even after removal of the primary tumor site, metastases can continue to grow and form more metastases. Due to metastatic spread (dissemination) can lead to the uncontrolled spread of the disease, often to the whole organism (generalization).
l According to the organ from which the tumor primarily originates : - eg breast, lung, bronchogenic, prostate, etc.
l By name of cells and tissues. The name of the tumor is formed from the Latin name of the tissue, organ or original cells from which the tumor arose and the ending "-oma" is added to this name - e.g. melanoma or melanoblastoma (skin tumor of melanocyte cell ), glioma or glioblastoma (primary brain tumor, also astrocytoma ), lymphoma (tumor growth of lymphoreticular tissue), etc. A malignant tumor of hematopoietic tissue, manifested by an increase in white blood cells (which are immature and do not perform their normal function), is called leukemia .
l According to the location of metastatic involvement - eg metastasis of breast cancer to the liver, skeleton, etc.
l By tissue and cell nature :
Epithelial tumors called (in the narrower sense) carcinomas are the most common type of cancer. According to the cell layer from which they originate, they are further divided into squamous cell and basal cell carcinomas (these names come from skin tumors).
According to the microscopic appearance of cancer cells (for histologic observation) and shape of the tumor growth is sometimes used designation papillary (warty - tumor forming warty or fimbriate formations), tubular (forming a tubular structure), medullary (marrow), ductal (outlet pipe), lobular (lobe ) and the like.
Benign tumors arising from the glandular epithelium are called adenomas .
- Mesenchymal tumors , called sarcomas, come from connective tissues. They occur less often.
    The resulting terminology of specific types of tumors is often formed by combining the names of individual categories - eg prostate adenocarcinoma, osteosarcoma in the skeleton, etc.

Some particular species of tumors
Tumor diseases are very diverse, about 100 types and subtypes of tumor diagnoses have been described. Here we briefly mention several more well-known and more frequently occurring types of cancer :
-> Skin tumors primarily affect the skin, but they can grow into deeper layers of tissue, some types even establish metastases. Malignant melanoma is a cancer that arises from the neoplastic proliferation of melanocytes. A risk factor is UV radiation (including excessive tanning in the sun). Melanoma initially spreads radially in the skin, later it can metastasize to lymph nodes or hematogenously to internal organs. In this case, it is very dangerous, it is one of the most malignant tumors! Basaliom - basal cell skin cancer - develops in the basal cell layer of the skin. It is the most common cancer of the skin. UV radiation also contributes to its formation. It is mostly benign and rarely metastasizes. It is removed by excision or laser. Squamous cell carcinoma (spinalioma) arises from flat cells of "squamous" shape, which are found in the epithelial tissue of the middle layer of the skin. It can later metastasize (mainly by the lymphatic route), so it should be removed by excision even with a safety margin. The name squamous cell is also used for squamous tumors of the lining of hollow body organs, respiratory tract, uterus.
-> Lung cancer - bronchogenic - is a common disease especially in men, the main risk factor is smoking. Malignant proliferation of epithelial tissue cells usually occurs. The two main groups are small cell and non-small cell lung cancers. Non-small cell carcinoma is more common. It is divided into adenocarcinoma (arising in the marginal lung tissue), squamous cell carcinoma (originating from the epithelium of the large airways) and large cell carcinoma. Small cell carcinoma is made up of small cells shaped like oat grains, it soon metastasizes through blood and lymphatic vessels, and is more difficult to treat. In addition to primary lung tumors, metastases of malignant tumors from other tissues (such as breast cancer, .....) can also occur in the lungs; Circulating tumor cells can easily become trapped in the capillary blood vessels of the lungs.
-> Breast cancer is the most common cancer in women. The most common (about 80%) is ductal carcinoma of the breast, which arises from the cells lining the ducts of the mammary gland. Furthermore, lobular ca of the breast (10-15%) arising in the lobules of the mammary gland. According to histology, breast cancers are further divided according to the presence of tumor markers/receptors on the surface of the cells, such as HER2+. This is important for biological treatment. In addition to the lymph nodes (most often in the armpit), breast tumors often metastasize to the skeleton, liver, and lungs. For early diagnosis of breast cancer, regular mammography is recommended after the age of 40 (described above in §3.2, passage "X-ray mammography").
-> Gynecological tumors in women - uterus and cervix, ovaries. One of the risk factors is infection with oncogenic HPV viruses. Cervical tumors arising from the epithelial layer are most often squamous cell carcinomas (80%) and adenocarcinomas (20%). Among uterine tumors, the most common epithelial tumors are endometrial adenocarcinomas (83%). Ovarian and fallopian tube tumors are mostly epithelial carcinomas.
-> Prostate cancer is one of the most common oncological diseases in men. Adenocarcinoma of the prostate most often develops in the peripheral part of the gland. A specific feature of prostate tumor cells is increased expression of the prostate-specific antigen PSA. An important diagnostic indication of prostate cancer is therefore an increased level of PSA in blood samples (it is also somewhat elevated in benign hyperplasia and inflammation of the prostate). Advanced prostate cancer can often metastasize to lymph nodes in the pelvis and to the skeleton. Some prostate tumors are hormone-dependent, hormonal therapy (castration, estrogens) can be effective for several years; however, after a longer period of time, they often become hormonally refractory. Ca prostate tumor cells have one more specific property that is beneficial for diagnosis and therapy: in addition to secreting PSA into the blood, they have an increased occurrence of the transmembrane receptor Prostate Specific Membrane Antigen PSMA on their cell wall. Appropriate ligands such as PSMA-617 were developed for this receptor and radioactively labeled both with diagnostic radionuclides 68Ga or 18F for scintigraphic imaging, and with therapeutic radionuclides 177Lu or 225Ac for effective biologically targeted radionuclide therapy. This precise theranostic approach makes it possible to cure even metastatic hormone-refractory prostate cancer - see below Prostate cancer - Radioisotope therapy can win over prostate cancer!".
-> Kidney and bladder tumors. The most common kidney tumor is adenocarcinoma of the renal parenchyma, from the epithelium of the renal tubules (sometimes called Grawitz's tumor). The most common bladder tumors are papillary tumors arising from the urothelium, which tend to have the shape of growths protruding into the bladder cavity.
-> Liver tumors. Primary hepatocellular carcinoma, arising from malignant degeneration of liver cells of hepatocytes, occurs relatively rarely. However, the liver is a frequent site of metastases of various malignant tumors (cancer of the large intestine, breast, lung, etc.), because blood flows through it from the digestive tract, which i.a. they "clean" from metabolites. Circulating tumor cells can therefore be trapped in the bloodstream of the liver.
-> Colorectal cancer - colon tumors are mostly adenocarcinomas arising from the glandular cells of the intestinal mucosa. A risk factor is polyps - growths from the wall of the intestine, from which cancer can develop over a long period of time. In the advanced stage, it often metastasizes mainly to the liver.
-> Pancreatic cancer arises mainly from the exocrine part of the pancreas, most often it is an adenocarcinoma from the epithelium of pancreatic ducts through which pancreatic enzymes pass. It is initially asymptomatic, usually diagnosed at a late stage, curative therapy is not yet successful. From the point of view of endocrinology and biochemistry, pancreatic cancer is included among the so-called neuroendocrine tumors. Their specific feature is the presence of somatostatin receptors, which can be used for biologically targeted radionuclide therapy using somatostatin analogues labeled with 177Lu ("Neuroendocrine tumors" is mentioned below).
-> Thyroid cancer can be differentiated (papillary and follicular), medullary, undifferentiated (anaplastic). Cells of differentiated thyroid cancer, including possibly metastases, retain the ability to accumulate iodine, which is the basis of its scintigraphic diagnosis and mostly successful therapy with radioiodine 131I (it is described in more detail in §4.9.1 "Thyreological radioisotope diagnostics" and below in the section "Biologically targeted radionuclide therapy", passage "Carcinoma therapy thyroid gland with radioiodine 131I").
-> Tumors in the skeleton can be of two origins. Primary bone tumors originating in cells found in the bones, called bone sarcomas, such as osteosarcoma, are rare. Bone metastases are far more common, arising from the malignant growth of tumor cells originating from tumors of other tissues - most often of the breast, prostate, uterus and others. Most skeletal metastases are hematogenous. The vascularized inner environment of the bone is a "nurture ground" for the capture and multiplication of tumor cells, which enter the bone tissue and begin to bind with stromal cells. Osteolytic deposits in bones caused by multiple myeloma are a special case; this is not metastasis in the usual sense.
-> Brain and CNS tumors. Primary brain tumors often arise from cells of the supporting brain tissue (neuroglia) - gliomas. The most common are astrocytomas, from less malignant pilocytic ones to very aggressive and dangerous glioblastomas. Furthermore, oligodendrogliomas, which are also less aggressive. In addition to primary brain tumors, the brain tissue is even more often affected by metastases of tumors from other tissues and organs (breast cancer, lung cancer, malignant melanoma, ...), whose hematogenously migrating cells have an increased probability of being caught, thanks to the high blood flow to the brain. Intracranial tumors, including metastases, up to a size of about 3 cm can be relatively successfully treated with stereotactic radiotherapy (it is described below under "Stereotactic radiotherapy"), primarily with the Lexell gamma-knife.
    Hemato-oncological tumors - malignancies of the blood and blood-forming organs - form a separate specific group of often diffusely localized lesions :
-> Leukemia is caused by tumor growth of certain types of white blood cells - leukocytes, which are immature and do not fulfill their normal function. It is accompanied by the suppression of normal hematopoiesis. Myeloid leukemia affects the formation of monocytes or granulocytes, lymphatic leukemia affects the formation of lymphocytes. Depending on the speed of the course, leukemia is classified as acute or chronic. The category of myeloproliferative diseases also includes polycythemia, which is an increase in erythrocytes in the blood. This increases blood viscosity, which can impair blood circulation.
-> Lymphomas form a group of tumors of the lymphatic system, which have varying degrees of malignancy. They are most often manifested by enlargement of the lymph nodes ......... According to the type of tumor cells, they are traditionally divided into Hodgkin and non-Hodgkin. Monoclonal antibodies such as rituximab can be used for the therapy of lymphomas, for biologically targeted radionuclide therapy the monoclonal antibodies anti-CD37 labeled 90Y ibritutomab tiuxetan or anti-CD20 177Lu-tetulomab (mentioned below "Immunotherapy of lymphomas").
-> Multiple myeloma is a malignant disease in which clonal plasma cells - myeloma plasma cells - multiply in the bone marrow, which is the seat of hematopoietic cells from which all types of blood cells arise. Genetic changes occur in myeloma cells that allow them to grow out of control. Through the bloodstream (along with normal functional blood cells), they spread to other places in the bone marrow and sometimes can infiltrate other tissues as well - multiple lesions often arise. Tumor myeloma cells produce pathological cytokines, monoclonal immunoglobulin (also called paraprotein), some of which disrupt the balance between osteoblastic and osteoclastic activity in the skeleton. Myeloma lesions often attack the bone tissue from the inside, creating a number of ostreolytic foci in the affected bones (they are not bone metastases in the usual sense). Pathological fractures may occur in these deposits of decalcification. In addition, hematopoietic disorders (mainly red blood cells – anemia) and often kidney damage occur.
An interesting fact is that monoclonal antibodies, which are important for biologically targeted antitumor therapy and diagnostics, are obtained from myeloma cells using the so-called hybridoma technology (described below in the passage "
Monoclonal antibodies").

Anatomical extent - progression - of tumor disease (staging)
To assess the possibilities of optimal therapy, the progression of cancer is crucial - how the tumor growth in the body has spread, how far it has penetrated. The anatomical extent (progression, staging) of cancer is often assessed according to three "TNM" criteria : T-tumor, N-nodes, M-metastases; the higher the number, the greater the range and propagation :
- extent of the primary tumor: T0 (no signs of primary tumor), T1 (tumor up to 2 cm in size), T2 (2-4 cm), T3 (tumor larger than 4 cm), T4 (larger tumor growing into other structures).
N - presence and extent of infiltration in regional lymph nodes: N0 (no infiltration in nodes), N1 (metastasis in one node, <3cm), N2 (bilateral metastases <6cm), N3 (metastases > 6cm). Alternatively, N2 indicates 2-4 and N3 more than 4 metastases in regional nodes. There are different T and N numbering conventions according to the type and location of the tumor.
M- presence of distant metastases: M0- without metastases, M1 - occurrence of distant metastases.
    Tumor TNM classification is not used for hematological and lymphatic malignancies (leukemia, lymphomas), which are not localized but diffuse.
    A simpler classification of the progression is into four stages of cancer, also called FIGO classification
(Federation International of Gynecology and Obstetrics) :
Stage I.
- a smaller tumor with local growth, without any dissemination (corresponds to T1, N0, M0).
Stage II.
- larger tumor with local growth, without dissemination or with minimal regional infiltration (corresponds to T2, N0-1, M0).
Stage III . - large local tumor with regional infiltration (T3-4, N2, M0).
Stage IV. - tumor involvement with infiltration into other tissues or with distant metastases (corresponds roughly to T2-4, N2-4, M1).
    The choice of treatment method and its success depend on all these aspects. In general, the success of therapy is greatest in isolated well-differentiated tumors in the early stages without metastatic infiltration (eg T1-2, N0-1, M0). In the late stages with extensive metastatic infiltration (generalization), treatment is difficult and usually not very successful ...
Degree of tumor cell differentiation - grading
Tumor tissue cells were formed by mutation and malignant transformation
(see above "Carcinogenesis - tumor formation") of originally normal cells of a certain healthy tissue or organ in which the tumor originated. Thus, they carry many of the properties of these original cells, but some of their other properties differ. The degree of tumor differentiation - the extent to which tumor cells differ from the cells of the normal tissue from which they originated - is referred to as grading (lat. Gradus = state, degree of a particular process ). If the tumor cells retain some of the properties of the original tissue, from by degeneration arose, it is a differentiated tumor. However, tumor cells often lose the properties of the original tissue - an undifferentiated (anaplastic) tumor is formed, which is largely autonomous, without binding to the regulatory mechanisms of the original tissue. Tumor grading is sometimes quantified using scores : G1 (well differentiated), G2 (moderately differentiated), G3 (poorly differentiated), G4 (undifferentiated - anaplastic). The prostate ca system uses a multi-level Gleason grading score system (up to 10 degrees).
    The risk of cancer also depends on the size of the tumor cells. The small cell type of tumor usually has rapid infiltrative growth with frequent metastatic dissemination.
Cellular heterogeneity of tumor tissue
In the initial stages, after its formation, the tumor is formed by a substantially homogeneous population of mutated cells that have escaped the regulatory mechanisms of tissue homeostasis and initiated uncontrolled division. In later stages, however, histocytological analyzes have shown that tumor tissue often contains two or more clones of cells with different biological properties - it is already heterogeneous. This is due to the increased fragility of DNA and the genetic instability of tumor cells, which may undergo further mutations upon repeated division. Tumor heterogeneity complicates treatment, because different parts of the tumor may have different radiosensitivity. A similar effect is also caused by possible hypoxia of some parts of the tumor
(see "6R" below, section "Oxygen effect").
    The specific type and nature of the tumor can be most reliably determined by histological analysis of a sample of tumor tissue under a microscope. An experienced pathologist usually recognizes the origin and type of tumor cells *) and also whether the tissue is benign or malignant. Histological examination should always precede therapy.
*) However, tumor cells that look similar under a microscope and are histologically classified in the same category, may have genetically different causes malignant behavior. Regulatory mechanisms not encoding RNA derived from as yet unexplored "genetic litter", "unnecessary" DNA (it is also mentioned in §5.2, section "
DNA, chromosomes, telomeres") may also be involved . This can significantly complicate the chain of diagnostics ® therapy of cancer.
    A more detailed classification of tumors and their clinical properties is beyond the scope of this physically focused treatise.

Diagnostics of tumor disseases
The success of any therapy depends to a large extent on careful diagnosis - both on the primary examination before treatment and monitoring the response during therapy and subsequent long-term follow-up. This is increasingly true of cancer. Malignant tumors are characterized by some specific characteristics :
- They are structures with a higher density than the surrounding tissue; also for ultrasound they usually show increased echogenicity than the surrounding tissue.
- They consist of metabolically active cells - they usually have increased metabolism.
- They usually have increased vascularization, increased blood flow and increased energy consumption. However, they may also contain hypoxic districts.
- Tumor cells may contain some general cellular antigens on their surface, or they may carry specific antigens.
- In addition, tumor cells may contain some special receptors in their cell membrane.
    All of these features can be used for diagnostic imaging and for targeted therapy .
× Primary diagnostics of tumors
This involves the finding of the primary tumor, its location and extent before surgery, as well as the discovery of possible metastases to determine the progress of further treatment. In addition to visible or tactile superficial and shallow lesions, primary tumor diagnosis is performed mainly using physical imaging methods -
X - ray diagnostics (planar, now mainly CT - §3.2), ultrasound sonography, radionuclide gammagraphy (planar, SPECT, now mainly PET - chapter 4), NMRI nuclear magnetic resonance. Tumors located on the walls of body cavities and tubes (stomach, intestines, uterus) can be recognized by optical endoscopic methods. It should be noted, that none of the imaging methods alone will determine the malignant nature of the disease! Imaging methods must therefore be combined with biochemical and especially histological methods (see below).
    Because some tumors can produce specific substances (either by the tumor cells themselves or when they interact with the body's immune system), biochemical analytical methods of blood or tissue samples are also important . These are mainly various types of tumor markers - complex organic molecules (mostly protein composition), whose increased expression in the body is the result of the tumor process - such as determining the concentration of PSA at the prostate, or markers CA19-9, CEA, AFP. Recently, the (immuno) histochemical determination of Ki-67 antigen
(or MKI67 weighing 360 kDa; the name comes from a study in Kiel, Kiel - clone 67 in bowl 96) , which is associated with cell proliferation - with ribosomal DNA transcription. Furthermore, determination of the apoptotic gene p53 (or its mutation) or the anti-apoptotic gene Bcl-2. Or cytological examination by flow cytometry.
"Molecular" gammagraphic imaging  
Most imaging methods provide only morphological-anatomical information on the presence, size and shape of tissue, differing in density from the environment. In the CT and NMRI images, we show the tumor mass, which with its density or proton density and relaxation times T1, T2 differs from the surrounding tissue, but we do not capture whether there are viable and proliferating tumor cells in the displayed anomalous tissue. Although gammagraphy (scintigraphy) does not excel in spatial resolution, it captures the functional metabolic properties of lesions - blood circulation, metabolism, drainage and other functions of tissues and organs - at the "molecular-biochemical" level
(see §4.9.6 "Oncological radionuclide diagnostics"). In particular PET display distribution of 18F-fluoro-deoxyglucose (FDG), 18F-3-fluoro-3-deoxy-thymidine (FLT), 18F-fluorocholine and labeled monoclonal antibodies, provides contrast images of viable and proliferating tumor lesions.


Example of PET/CT scintigraphy with 18FDG in a patient with lymphoma.
    PET images show multiple foci of increased glucose metabolism in the lymph nodes of the neck, left axilla, mediastinum, retroperitoneum, pelvis and groin, indicating viable tumor neoplasia.
    Physiologically, 18-FDG deposited in the brain, in the hollow of the kidney, bladder, light diffusion in the liver and spleen, the variable is the accumulation in the myocardium
    The transversal section shows separately a significant deposit in the area of the supraclavicular (marked by arrows) for the assessment of metabolic accumulation using SUV values [g/ml] :  
    The SUVmax of the bearing in the nadklicek is 11.5, with the reference SUVmax liver 2.9 and the mediastinum 1.4.

(PET / CT images were taken by
Martin Havel, MD, Ph.D.,
head of the PET / CT department
of KNM FN Ostrava )

This method is also suitable for monitoring the response of tumor tissue to radiotherapy, as it displays metabolically active tumor tissue, in contrast to inactivated cells; it is thus possible to monitor the "success" of the therapy. Among other things, it is able to recognize tumor recurrence (with proliferating cells) from other structures, necrotic or connective tissue. The 18F-FMISO and 18F-FETNIM radioindicators show cellular hypoxia, which is important for tumor angiogenesis and for planning radiotherapy (radiosensitivity, oxygen effect - see below "Physical and radiobiological aspects of radiotherapy").

                      Normal whole-body scintigram bone                      Multiple metastases (breast ca) to the skeleton

    To display bone metastasis, at an early stage of infiltration, proves best bone scintigraphy (whole-body scintigraphy in PA and AP projection, with possibly targeted SPECT images of suspicious sites) after osteotropic radiopharmaceuticals, which are phosphate complexes, whose accumulation reflects increased osteoblastic activity in response to tumor bone destruction. Planar whole-body scintigraphy of the skeleton is useful to supplement with a combined SPECT/CT image with image fusion, to specify the anatomical location of the lesions. In thyroid cancer, the diagnosis is based on scintigraphy after application of radioiodine
131 or 123I.
    Labeled peptides that bind to peptide receptors on the surface of some types of tumor cells are mainly used in neuroendocrine tumors containing somatostatin receptors
(the cyclic peptide somatostatin is a hormonal substance that has a inhibitory effect on the production of certain hormones, especially growth; Greek soma = body, statizo = stand, stop ). An artificial somatostatin analogue, octreotide, labeled with indium - 111In-pentetreotide (OctreoScan), or 68Ga -DOTATOC for PET scintigraphy is used to visualize the respective tumors and to predict the effect of somatostatin analog therapy. The logical sequence of diagnostics using these radioindicators is their labeling with therapeutic radionuclides with application for radioisotope biologically targeted radiotherapy, see below "Radioisotope therapy", section "Radionuclide therapy of tumors and metastases".
    In addition, some non-specific indicators of tumors are used, the increased accumulation of which in tumor tissue is based on their ability to penetrate pathologically altered permeability of walls and capillaries and bind within viable cells. Used
99mTc-MIBI and tetrofosmin mainly in lymphomas and mammary tumor (mammoscintigraphy). In oncological diagnostics, gallium scintigraphy (mostly planar whole-body imaging, supplemented with possibly SPECT images) with 67Ga-citrate. Chemically, Ga ions are analogs of Fe ions, bind to the transport protein transferrin, and accumulate in proliferating tumor tissues, particularly lymphomas. Gallium-67 scintigraphy is now abandoned and replaced by PET scintigraphy with 18FDG. Indirectly, tumor processes can sometimes be inferred from other scintigraphic examinations, such as dynamic scintigraphy of the kidneys and liver.
Multifactorial image analysis - radiomics
Using the methods of "machine learning" and artificial intelligence, procedures of so-called radiomics were developed - sophisticated image analysis in conjunction with statistical processing of data from a large number of patients, which can recognize even hidden informations, directly invisible visually. These analyzes can reveal some similarities and coincidences, helping to refine the likely diagnosis of tumor types, or predict response to therapy and course of disease. It is discussed in Chapter 4, §4.7, paragraph "Multifactorial statistical analysis of images, radiomics".
Histological examination
If anomalous tissue (neoplasm) that could be of tumor origin is found by imaging or other examination methods, histological examination must be performed: a small sample of tissue is taken by biopsy and then viwed under a microscope. Histological examination is also performed on "suspicious" tissues removed during surgery. According to the shape, size and other characteristics of the cells, it is usually possible to distinguish whether it is a benign or malignant tissue and what kind of possible tumor cells they are (as mentioned above). All this macroscopic and microscopic diagnostic information determines the optimal way to treat cancer.
× Diagnostics for cancer therapy planning
If the primary diagnosis of cancer is confirmed, the stage of preparation of therapy begins - a decision on the basic strategy and methodology of therapy: whether it will be a surgical solution, chemotherapy, radiotherapy, eventual its combinations (see "Cancer therapy" below). If radiotherapy is planned, X-ray CT, NMRI and gammagraphy (especially PET) images can be used to determine the exact location and extent of the tumor site and to plot the regions of interest (ROI) of the irradiated volume - GTV, CTV and finally PTV in the irradiation plan (see below the section "
Planning of radiotherapy"). The base is now used CT images (respectively NMRI), but these mainly reflect morphological page, but do not capture the behavior of biological tissue. It is therefore useful to also gamagraphic views, especially PET images of the distribution of 18FDG or 18FLT, respectively 18F-choline. By analyzing these PET images (eg by determining SUV levels - see §4.2, section "Scintigraphic image quality and detectability of lesions"), we can determine the "Biological Target Volume" (BTV) of tumor tissue formed by viable proliferating cells. By transferring of these images to a radiotherapy planning system and computer image fusion (CT, NMRI) + PET we can then refine the volume of the target lesion, especially CTV for IMRT radiotherapy. According to the distribution of tumor cell viability, we can further modulate the dose within the tumor site and increase (escalation, boost) doses to risk areas within the tumor.
l Predicting the response to cancer therapy
Predicting the biological response to planned treatment is a very difficult task in medicine in general. In cancer therapy, certain basic information can already be obtained from the results of primary imaging and histological diagnostics. However, there are ways to indirectly assess the specific behavior of tumor tissue for the planned type of treatment using gamma imaging methods, especially PET.
l Monitoring the distribution of cytostatics and monoclonal antibodies
The success of chemotherapy depends, among other things, on whether the cytostatics or monoclonal antibodies used accumulate (uptake) sufficiently in the tumor foci. Nuclear medicine can be used to predict the chemotherapeutic effect. Methods for radionuclide labeling of some chemotherapeutics have been developed: after "trial" diagnostic application of a small amount of such labeled radioindicator, we can display its distribution and assess how selectively it is taken up in tumor tissue (as well as in healthy tissues that could thus be undesirably affected by by radiation) - in this way, the respective chemotherapeutic agent will be taken up during the actual therapeutic application. A similar "trial" diagnostic application of a smaller amount of
g- radioindicator can be used in radioisotope therapy (see "Radioisotope Therapy" below).
l Imaging Dendritic Cell Migration
One of the basic preconditions for the success of immunotherapy with dendritic cells activated by antigens of a particular tumor tissue (see "Cancer Therapy" below, "
Immunotherapy" section) is their migration to peripheral lymph nodes and then to the tumor locus. If we label these activated dendritic cells before their re-application to the body using a suitable radioindicator (111In-oxin is tested), while maintaining their viability, we can scintigraphically map their migration to the lymph nodes after their application.
    The predictive role of imaging cell apoptosis is discussed in the following paragraph.
× Monitoring the biological response to cancer therapy
In addition to the primary diagnosis, it is desirable to monitor how successful the therapy is and what its side effects are on healthy tissues, organs and the whole organism. In terms of time relation to therapy, monitoring of biological response can be divided into prediction of biological effect before therapy (mentioned above) or at the beginning of therapy, monitoring of early response during therapy and monitoring of late response and overall long-term development disease after treatment. To assess the late tumor response, the basis for monitoring the tumor site in CT or NMRI images - a comparison of the size (volume) of the displayed tumor lesion before and after therapy to assess the reduction of tumor mass. However, CT and NMRI images capture only the morphological situation, not the biological development of tumor tissue and the metabolic activity of cells - we do not recognize in them what part of the depicted lesion of different density is formed by viable tumor cells and what part by necrotized or connective tissue. It is therefore useful to use "molecular" gammagraphy using SPECT and PET methods to reliably monitor the tumor response. These are mainly the already mentioned images of the
18 FDG or 18FLT distribution, performed before and after therapy (or during therapy), on which possibly we compare SUV values. Molecular gamma imaging allows the visualization of important factors influencing the response of tumors to therapy. Another "line" - radiobiological modeling - is the assessment of the radiotherapeutic effect using the quantities TCP, NTCP, UTCP, discussed below in the section "Prediction of the radiotherapeutic effect - the probability of cure of a TCP tumor and damage to normal NTCP tissue".
    Immediately after the end of radiotherapy, no macroscopic change can be detected in the irradiated tumor - the changes take place first at the molecular level. Only a few weeks to months apart, these nitrocellular processes result in the extinction of most of the cells in the tumor population; only this can be accompanied by observable morphological changes.
l Early tumor response - imaging of cell apoptosis
However, functional molecular imaging in gammagraphy provides other unique possibilities. Radioindicators have been developed to monitor one of the main radiobiological mechanisms of cancer therapy (both radiotherapy and chemotherapy): cell apoptosis. In cell apoptosis
(see §5.2, section "Effect of radiation on cells", section "Cell apoptosis") in the early phase, among other things, irreversible membrane depolarization occurs, the uncovering of phospholipids on the cell surface, increased permeability of the plasma membrane, then at a later phase the integrity of the cell wall is violated and finally to cells disintegration and their phagocytosis. It is in the early phase of apoptosis that the special radiopharmaceuticals shows an affinity for apoptotic cells *): they either bind to phospholipids on the surface, or penetrate the cell membrane and accumulate in the cytoplasm of apoptotic cells. The result is a selective accumulation of radioindicator in apoptotic cells and tissues, while they hardly enter in tissues formed by normal viable cells or necrotic tissues.
By gammagraphic imaging of the distribution of these radioindicators
(it is appropriate to use dynamic gammagraphy - time factor of accumulation) we obtain positive images of those places, where apoptosis occurs most intensively - whether due to irradiation, cytotoxic substances or ischemia. By molecular imaging of the distribution of cell apoptosis, we can monitor the very early response of cells and tissues to therapy (radiotherapy or chemotherapy), already at the beginning and during therapy. It basically allows the prediction of the tumor response: we apply "experimentally" one or two fractions and on gammagraphic images we can assess whether apoptosis is taking place in the target tissue sufficiently intensively, or whether there are regions of apoptotic resistance in the heterogeneous tumor. Early imaging of apoptosis can play a significant role in biological individual ("personalized") therapy of a particular patient.
*) Three types of radioindicators of apoptosis are in the stage of laboratory development and preclinical studies (see also §4.8 "Radionuclides and radiopharmaceuticals for scintigraphy") :
- Protein 99mTc-Annexin V (for SPECT imaging) - binds to phospholipids on the surface of apoptotic cells;
- 18F-ML-10 [2- (5-Fluoro pentyl) -2-methyl malonic acid] - a small molecule that penetrates the wall of apoptotic cells and accumulates in their cytoplasm. Approx. 400MBq is applied.
- Peptide 18F-CP18 [triazole-containing pentapeptide] - maps Caspase-3 activity, accumulates in apoptotic cells.

Combination of diagnostics and therapy - theranostics
New diagnostic imaging methods, especially molecular imaging in nuclear medicine, allow to integrate individual (personalized) diagnostics and targeted therapy (or prevention) of serious diseases into a common field, for which it was newly using the name theranostics
(created by composing names: therapy + diagnostics => Theranostics). Scintigraphy makes it possible to determine the concentrations of biologically active substances directly at the sites of their targeted action, which enables optimal and individual dosing, with the possibility of predicting effects and monitoring the results of therapy - it is discussed in more detail in §4.9, section "Theranostics".

Therapy of tumor diseases
In terms of goal and effectiveness, we generally distinguish between two types of treatment: Curative therapy (Latin cura = treatment) with the aim of complete cure of cancer, especially in the localized stage. In more severe and advanced cases, then palliative therapy (Latin pallium = mantle), alleviating and slowing down the course of the disease and its difficulties. In terms of time sequence, we also recognize two procedures: Induction therapy - initial treatment to achieve remission of the disease. After this primotherapy (possibly also simultaneously), adjuvant therapy is often applied - auxiliary, supportive or securing treatment (Lat. adiuvo = support, help), especially to reduce the risk of recurrence due to possible micro-seeding around the original tumor.
  The treatment of tumors is currently based on three main methods: surgery, chemotherapy and radiotherapy, with these three main therapeutic approaches are often combined - multimodal treatment. In the surgical treatment of cancer, physical removal is performed - resection or ablation
(Lat. Ablatio = removal, distant) of the tumor tissue. It is desirable to remove not only the primary tumor with the "safety margin", but also, if possible, other tissues into which the tumor cells could be infiltrated: these are mainly the surrounding lymph nodes located in the lymphatic "river basin" of the tumor location (see also above §3.5, passage "Radiation-guided surgery - sentinel nodes"). In addition to classical surgical techniques, radiofrequency ablation and stereotactic ablative radiosurgery SRS (sterotactic radiosurgery) are also used - see the "Stereotactic radiotherapy" section.
  For non-surgical treatment of cancer, we would ideally need some "magic missiles" that would penetrate the body in a non-invasive manner, target and destroy only the tumor cells, while maintaining undamaged healthy surrounding tissue. Because tumor tissue is made up of cells that are not very different from the healthy cells of the surrounding tissues, we do not have such an ideal and selective "shot": treatment focused against tumor cells will always more or less damage even some healthy cells, tissues and organs. However, there are certain physical and biological factors that at least partially promote the targeted destruction of tumor cells and minimize damage to healthy tissues.
  In terms of the place of action in the body, we can divide the therapy of cancer into two methodological approaches :
l Local tumor control , in which we try to stop tumor growth and destroy cells in a particular tumor site of known location and extent. It is performed mainly by the method of radiotherapy - targeted delivery of a high radiation dose to the tumor site. This approach is effective for well-defined tumors of small or medium size, without distant metastases.
l Systemic therapy performed by application of suitable drugs into the organism, which enter the tumor foci and stop the proliferation or kill the tumor cells there. These include chemotherapy with cytostatics and biological therapy, and in part targeted radionuclide therapy. This whole-body systemic action has its advantages and disadvantages. The advantage is the action on multiple tumor foci and hidden metastases, the presence and location of which we sometimes do not even know. The disadvantage is the side effects on healthy cells and tissues, into which the chemotherapeutic agents also enter, through the blood and lymphatic way. Systemic therapy is chosen in cases of more extensive tumor disease with metastatic infiltration. And also as an adjuvant therapy to reduce the risk of recurrence and metastasis.
  Below, first will be briefly described methods of systemic and targeted chemotherapy and biological treatment, then more detailed methods of targeted radiotherapy using physical and radiobiological aspects.
Chemotherapy and biological treatment
Under chemotherapy generally means treating diseases by administering chemicals - drugs that are the product of chemical synthesis or isolated from natural materials (especially plants) - and which cause desirable (bio)chemical reactions in the organism. Chemotherapy of cancer is most often performed using cytostatics - substances that stop or inhibit the growth and division of cells
(Greek: kytos = cavity, cell; statikos = stopping). Their preferential antitumor effect is due to the fact that they act primarily on rapidly dividing cells. However, they also affect healthy physiologically dividing cells in the body, which leads to undesirable side effects. A number of cytostatics are known, which act by different mechanisms and at different stages of the cell cycle. There are two basic mechanisms of action of cytostatics :
1. Action on DNA , which disrupts cellular function, prevents replication, it can be evaluated by the cell cycle control nodes as irreparable damage ® activation of the internal signaling pathway of apoptosis (in this respect the mechanism is similar to ionizing radiation).
2. Effects on other cellular structures, especially microtubules, which violates the very act of cell division - it acts as a "mitotic poison". The cells are thus inactivated, they cannot divide further, they undergo apoptosis either directly or following a "mitotic disorder" (see §5.2, passage "Mechanisms of cell death").
    In more detail, four diffrent mechanisms of action of cytostatics are distiguished, according to which these substances are divided :
l Alkylation cytostatics - react with bases in the DNA, e.g. guanine, by their alkylation - by transferring the carbon radical group C2nH2n+1 (alkyl). This leads to DNA cleavage or the formation of a two-stranded junction. This damage to DNA inactivates cells, prevents them from dividing (DNA strands cannot untwist and separate), and ultimately leads to apoptosis. A cytostatic effect has long been observed with nitrogen mustard analogues. From this group, chlorambucil, cyclophosphamide or ifosfamide are used (own cytostatic effect has only their metabolite oxycyclophosphamide formed in the nucleus), as well as fludarabine and bendamustine .
Platinum cytostatics also belong to this group. The longest used cytostatic of this species is cisplatin (cis- [PtCl
2 (NH3)2]), an inorganic molecule that binds to guanine bases in a DNA molecule; more recent are the organic compounds carboplatin and oxaliplatin, where platinum atoms are attached to cyclic ("aromatic") hydrocarbons.
l Microtubule inhibitors - react with microtubules in cells, prevent the formation of a mitotic spindle - mitotic poisons. These are two types of chemicals that have opposite mechanisms of microtubular action, but result in similar cytotoxic effects :
- yew terpenides - taxanes. The alkaloids paclitaxel (contained in yew Taxus brevifolia ) and docetaxel (from yew Taxus braccata) are used, which stabilize microtubule polymers (inhibition of microtubule depolymerization) and prevent chromosome separation during anaphase.
- vinca alkaloids - vincristine, vinblastine, vinorelbine, which bind to tubulin and prevent its polymerization into microtubules.
l Antimetabolites blocking the synthesis of purine and pyrimidine DNA bases required for cellular replication. Substances with a structure similar to purines and pyrimidines - fluoropyrimidines, especially methotrexate or 5-fluorouracil (5-FU) preparations are used. More recently, the 5-FU precursor, capecitabine, is preferably used, from which the active substance 5-fluorouracil is formed by enzymatic transformations only in the body, preferably in tumor tissue (the enzyme thimidine phosphorephilase, which participates in the final phase of the conversion of inactive capecitabine to active 5 -FU, is contained in tumor cells usually in a significantly higher concentration than in cells of healthy tissue - selective effect).
l Topoisomerase inhibitors in S-phase of the cell cycle prevent the untwist of the DNA double helix during the replication process (in which the enzyme topoisomerase is involved). This leads to the induction of DNA breaks that can cause cell death. The original substance of this kind was camptothecin, an alkaloid isolated from the Chinese tree Camptotheca acuminata. However, its synthetic derivatives topotecan and irinotecan have more suitable pharmacological properties.
l Antitumor antibiotics are originally substances that inhibit the growth and multiplication of microorganisms, therefore used in the treatment of infectious diseases, as well as antifungal agents and the like. They are secondary metabolites of microorganisms (bacteria, mold), many of which are now prepared artificially (by synthetic or semi-synthetic methods). As most of them contain in their chemical structure several groups of cyclic hydrocarbons ("benzene nuclei") characteristic of anthracene, they are also called anthracycline antibiotics. In addition to the antibiotic effect, in some of these substances have also been found an immunosuppressive effect and a cytostatic, antitumor, antiproliferative effect. The mechanism of the cytostatic effect is probably binding to DNA - their molecules have the ability to be incorporated between DNA base pairs; DNA breaks occur, intercalation bonds are formed with a transcription disorder (tight connection of both strands of DNA prevents its copying before cell division and transcription into RNA), DNA breaks down (the effect is similar to that of radiation or alkylating cytostatics). The enzyme topoisomerase is also blocked, which is involved in changes in the spatial arrangement of DNA during replication prior to cell division; when it is blocked, the individual parts of the DNA do not come together, that break down. An example is doxorubicin, bleomycin, epirubicin, idarubicin, mitomycin C.
    This group also includes, in part, rapamycin, originally isolated from the bacterium Streptomyces hygroscopicus discovered in soil on Easter Island Rapa Nui; it is also called sirolimus
(Lat. siro = pit dug in the soil, limus = mud, sludge). It ranks among the macrolide antibiotics that block protein synthesis in microorganisms by binding in ribosomes. Due to its immunosuppressive effects, it is used in transplants as protection against adverse immune reactions that can lead to transplant rejection. By inhibiting the protein kinase (mTOR *), it blocks a number of intracellular processes, leading to a decrease in cell proliferative activity. It prevents cells from moving from the G1 phase of the cell cycle to the S phase, causing cell cycle arrest. It increases the sensitivity of tumor cells to radiotherapy and the effectiveness of chemotherapy. Rapamycin thus also belongs to the group of kinase inhibitors listed below, where its new derivatives temsirolimus and everolimus are used.
*) Rapamycin, mTOR
mTOR (mammalian Target Of Rapamycin ) - this very misleading name comes from the fact that the relevant protein kinase was first discovered when rapamycin was applied to the ca of breast. An alternative name is "mechanic target of rapamycin". It was later shown that mTOR also works in other types of tumor cells. The PI3K/Akt/mTOR signaling cascade is significantly involved in the process of carcinogenesis, and its inhibition may be an important factor in cancer therapy.
l Antioxidants are known primarily as cancer prevention. However, it has been shown that some antioxidants (such as reveratrol, genistein, baikalein) damage DNA and kill dividing cells. This could be used in anticancer therapy. Their advantage is that, despite their genotoxicity, they do not have mutagenic effects. So far, it is in the stage of biological research.
l Bisphosphonates act as inhibitors of osteoclastic bone resorption. They can therefore be used for the secondary treatment of bone tumors, especially metastases, where they act primarily against bone erosion; it is not a cytostatic. Effective nitrogen bisphosphonate is mainly zoledronic acid. In combination with eg docetaxel, there is an additive synergistic antitumor effect - potentiation of the cytostatic effect.
    Individual cytostatics are sometimes combined, eg FOLFOX (oxaliplatin + 5-fluorouracil + folic acid), XELOX (capecitabine + oxaliplatin), FOLFIRI (5-fluorouracil + irinotecan) and others. A common disadvantage of classical cytostatics is their systemic non-specific effect - they act not only on tumor cells, but also on healthy physiologically dividing cells. This leads to a number of often serious side undesirable (toxic) effects. Therefore, a new variant of targeted "transport" of a suitable cytostatic preparation directly to cancer sites is being tested, using microcapsules up to 5 mm in size. Such tiny capsules, formed from a suitable organic substance and carrying a cytostatic inside, pass through the vascular system and the fine blood capillaries after application. They can be monitored sonographically or by magnetic resonance imaging; the moment they reach the tumor, they can be disrupted by an ultrasound wave, releasing the cytostatic at the required place - in the tumor.

Chemical structure of some cytostatics used in chemotherapy of cancer

Targeted biological therapy
In recent years, knowledge of molecular biology and genetics has developed rapidly, revealing, among other things, complex mechanisms of cellular communication and specific molecules that are important for malignant cell transformation. These could become the target of specific therapeutic interventions: to identify and target certain structures in tumor cells in order to prevent further proliferation of tumor tissue. Targeted biological treatment is based on these mechanisms, the strategy of which is directed against selected types of molecules and their signaling pathways involved in the malignant behavior of cells of the respective tumor types. Together with more effective therapy, these procedures make it possible to reduce undesirable side effects. Currently, the main interest is focused on the so-called growth factors (stimulating cell growth and division) and their receptors, especially EGFR, HER2 and VEGF (see below).
    A new class of drugs is being developed that selectively block the activity of these oncogenic proteins, with minimal damage to normal cells. These are mainly two groups of substances with different mechanisms of action :
l Monoclonal antibodies
are special proteins from the group of immunoglobulins
(or fragments thereof), which are obtained from a cloned population of one species of activated B-lymphocyte from the plasma of an immunized organism. The monoclonal antibody therefore has precisely defined properties and specifically binds to the respective receptors. Some monoclonal antibodies seek to approach an ideal therapeutics - a "magic arrow" that would hit only target pathological cells and have no detrimental effects on other healthy cells. However, this cannot yet be achieved 100% ..!..
 Structure of monoclonal antibodies
The molecular weight of monoclonal antibodies is around 150 kDa. The structure of immunoglobulin protein molecules is often schematically represented by the shape of the letter "
Y " (in the figure on the left - a). The branched part - arms - consists of two heterodimers, it is formed by four polypeptide parts, arranged in two mirror-identical pairs of "heavy" and "light" chains. They are internally linked by a disulfide bond (SS). Light and heavy chains contain constant and variable regions. The variable regions, located at the ends, contain short amino acid sequences (sometimes called "hepervariable"), that determines the binding antigenic specificity of an antibody. Therefore, these arms are referred to as Fab (Fragment antigen binding). The "foot" in the antibody scheme consists of two heavier chains, referred to as Fc (crystallizing fragment ) . This constant region of Fc is responsible for the effector functions of the antibody (interaction with T-lymphocytes, macrophages) - activation of systems leading to the destruction of target cells.
 Preparation of monoclonal antibodies - hybridoma technology
Because we can't cultivate directly the desired clones of activated B-lymphocytes efficiently enough, the preparation of monoclonal antibodies is complex biochemical technology.
Very suitable "auxiliary carriers" for preparing of monoclonal antibodies have proven to be myeloma cells (otherwise known as tumor cells of myeloma, a hematooncological disease of the bone marrow caused by uncontrollable proliferation of myeloma cells), which, due to their unlimited replication capabilities and longevity, are very suitable for cultivation and fusion in vitro. Thus, the formation of these "helper" myeloma cells is first induced in experimental laboratory animals. These are then harvested and cultivated in vitro. Meanwhile, in another laboratory animal, an injection of a particular antigen elicits an immune response with B-cell activation and subsequent production of antibodies. These B-lymphocytes are taken from the lymphatic system (usually a spleen sample) of the animal used and then fused in vitro with a colony of myeloma cells. From this fusion, hybrid cells (called hybridomas) are formed, which retain the properties of both myeloma cells and the desired B-cell clone, divide rapidly, and produce antibodies of B-cell (used in the fusion). Using special separation methods, only hybridomas producing only one desired antibody clone - a monoclonal antibody - are selected from them.
    This resourceful biotechnology was first developed in 1975 by G.F..Kohler and C. Milstein in the Molecular Biology Laboratories of the University of Cambridge and the Institute of Immunology in Basel (for this method they received the Nobel Prize in 1984).

Monoclonal antibodies .
a, b) Schematic diagram and illustration of the basic structure. c) Mouse monoclonal antibody. d, e, f) Chimeric, humanized and human antibody.

The laboratory animals used in the process of preparing monoclonal antibodies are almost always mice, so that the mouse monoclonal antibody (c) is primarily generated. For its human use can sometimes lead to undesirable immune reactions - immunogenicity due to the development of human-anti-mouse antibodies HAMA (Human Anti-Mouse Antibodies). On the one hand, this prevents (makes it impossible) the binding of the monoclonal antibody to the target antigen and, on the other hand, can also lead to immune anaphylactic reactions. Therefore, there is an effort to replace parts of the molecules (which do not encode antigen binding regions) with human immunoglobulin sections - to humanize antibodies using sophisticated biochemical-genetic methods ("genetic engineering"). From the original hybridone line producing murine antibodies of a given targeting, the RNA was prepared and further in a reaction catalyzed by an enzyme reverse transcriptase, a complementary cRNA -. By polymerase reactions were multiplies segment encoding the antigen binding site. By this genetic sequence are replaced the corresponding region in the human immunoglobulin gene. Thus formed resulting new gene is inserted into a suitable recipient mammalian cell, which then synthesizes a monoclonal antibody with a predominant human immunoglobulin content. The humanized monoclonal antibody has constant regions from a human immunoglobulin, and only the variable region, encoding antigenic specificity, is derived from a murine antibody. The antibodies thus transformed have the desired antigenic specificity and show only minimal immunogenicity. In this way, chimeric (cross) antibodies (content of about 60% human antibody), humanized (content of more than 90% human antibody) or 100% human antibody are generated - in Fig. d), e), f). The more "animal" the antibody, the greater the risk of immunogenicity can be expected. Humanization of monoclonal antibodies leads to a reduction in unwanted immunogenicity, but on the other hand also to a possible reduction in their effectiveness ...
 Antibody fragments 
The larger the protein molecule, the slower and more difficult it is to penetrate the target tissues. Therefore, instead of "whole" antibodies, suitable antibody fragments are prepared, containing only regions with preserved antigenic specificity. Most commonly, these are Fab´ fragments containing domains necessary for antigen binding, but not part Fc this interacting as effector.
 Nano-antibodies - Nanobodies
The most consistent "miniaturization" of monoclonal antibodies are the so-called nanobodies, consisting only of the variable domain of the antibody with the F
ab H heavy chain, without the light chain domain and, of course, without the Fc part. The name "nanobody" comes from its namometric size - length approx. 4 nm and width approx. 2 nm, molecular weight is only around 15 kDa. However, the ability of nano-antibodies to bind the relevant antigen remains the same as for conventional antibodies. In addition, nanoantibodies exhibit better stability, hydrophilicity and water solubility, deep tissue penetration, and rapid clearance from the bloodstream, contributing to their very good resulting binding affinity under various conditions.
    Nanobodies have excellent properties especially in the diagnostic imaging of tumor tissues. Images with a high signal-to-noise ratio can be obtained already in the early period after application, due to the rapid accumulation of the nanobody and rapid depletion in the blood.
    However, the therapeutic efficacy of nanobodies alone is limited by the absence of an effector Fc fragment. However, if the nanobodies are chelately conjugated with a beta or gamma therapeutic radionuclide, the therapeutic potential is preserved -
biologically targeted radionuclide therapy.
 Nomenclature of monoclonal antibodies 
The nomenclature of monoclonal antibodies is elaborated so that the basic type and the most important properties (targeting) of a specific preparation (antibody) can be identified from the name. The name consists of 4 parts: prefix - designation of the target structure - biological type (origin) - suffix (it is always - mab: monoclonal anti body) :

Monoclonal antibody name :
prefix - - target structure - - biological origin - - suffix
- ci (r) - vascular system
- tu (m) - tumor
- li (m) - immune system
- m (o) - mouse
- xi - chimeric
- zu - humanized
- u - human
- mab
( monoclonal
anti body)

Example are rituximab, a chimeric monoclonal antibody (-xi-), whose variable portion mediating contact with the tumor cell CD20 antigen (-tu-), is of murine origin and the remainder of the antibody is of human origin.
  In the literature, other more detailed divisions are sometimes given - in the target structure for specific types of tumors (
-co (l) - colon tumor, -go (v) - ovarian tumor, -ma (r) - breast tumor, -me (l ) - melanoma, -neu (r) - nervous system, -pr (o) - prostate tumor, -vi (r) - viruses .....), in biological origin other possibilities (-a- rat, -e - hamster, -i- primates, .....); however, we do not encounter these names in practice... In some preparations, in addition to the monoclonal antibody itself, the name is also given biochemically conjugated substances (eg ibritumomab tiuxetan - ......).
 Biological effects of monoclonal antibodies
In order to produce the desired effect
(therapeutic or diagnostic), the antibody must first reach the target tissues and cells. Therapeutic monoclonal antibodies are administered intravenously by slow infusion of the solution. Monoclonal antibodies have relatively large molecules with masses around 150 kDa, min. 100 times larger than conventional cytostatics. Therefore, they have slower distribution kinetics and more difficult to penetrate tumor tissue, with slow diffusion through the interstitial space. Their distribution in the tumor is usually inhomogeneous, especially in larger tumors.
    Monoclonal antibodies have the ability to react with the particular antigen against which they are targeted. If the target structure is a receptor ligand, this ligand is neutralized. If the target structure is a receptor on the surface of the cell membrane, the signaling pathway associated with it is blocked. Many monoclonal antibodies thus have inhibitory effects on certain ligands and signaling pathways
(or sometimes stimulatory ones).
    A frequent goal of treatment is the elimination - kill of a certain type of cells - the depletion process. The condition for binding to target cells is the presence of appropriate receptors. The antigen-binding complementarity of an antibody is given by the variable regions at the ends of the Fab chains, while the fixed Fc region mediates subsequent effector functions on target cells. After successful binding of the antibody, three different mechanisms of the resulting cytocidal effect can occur :
- Activation of complements - membrane glycoproteins C1-C9, which by their proteolytic effects attack cytoplasmic cell membranes and cause their penetration. The cell dies and the released chemicals cause an inflammatory reaction with the accumulation of leukocytes (cf. §5.2, passage "Mechanisms of cell death").
- Induction of phagocytosis - fixed Fc region of bound antibody (lower part of the letter " Y " in the scheme) specifically binds to the Fc receptor of some types of leukocytes, especially macrophages, which thus recognize and subsequently phagocytose tumor cells.
- Induction of apoptosis after antibody binding to the cell surface, with destruction of mitochondria and proteolytic caspase chain (detailed explanation in §5.2., passage "Apoptosis"; in the picture on the top right "External signaling pathway of apoptosis").
Immunogenicity of Monoclonal Antibodies 
Monoclonal antibodies, as immune-active agents, when administered to the body can form "anti-antibodies against antibodies", which can neutralize their effect and, in addition, to causing adverse anaphylactic effects - has already been discussed above. Immunogecality most often occurs in murine antibodies (HAMA, in about 10%), rarely in chimeric antibodies, very rarely in humanized. Before using murine antibodies, it is therefore desirable to perform a laboratory biochemical test for HAMA antibodies, its possible positivity should then be a contraindication to the use of these products.
 Use of monoclonal antibodies
In addition to oncology (see below), monoclonal antibodies are also used against autoimmune diseases, organ transplant rejection, inflammatory diseases. Also as antibacterial and antiviral.
    Monoclonal antibodies are mainly used in oncology for stand-alone targeted biological therapy of cancer (see below). In addition, cytotoxic substances or radioisotopes can bind to them, which only "guide" these antibodies, approach or bind them to the target cells - antibody conjugates with suitable effector components are formed.
 Radiolabeled antibodies - radioimmunoconjugates
An important new method of biologically targeted therapy of cancer is the combination of targeted binding of monoclonal antibodies with the biological effects of ionizing radiation from radionuclides. Radioimmunoconjugates have a beta or alpha radionuclide bound in an antibody molecule, for a biologically targeted radionuclide therapy - see below "
Radioisotope therapy with open emitters". The monoclonal antibody does not serve as a primary therapeutic here, but provides the "delivery" and get closer of the radioactive substance to the tumor cells. The advantage of these radioimmunoconjugates is that in order to achieve a biological effect, it is not necessary to bind them to each tumor cell, but emited radiation by the effect of "crossfire" has a radius of action of several tens of cell diameters (see below Fig.3.6.8 in "Radioisotope therapy"). They are therefore effective even in the neighboring cells which have insufficient expression of a tumor antigen.
    Some monoclonal antibodies labeled with a gamma or positron radionuclides, are used in nuclear medicine as radioindicators in scintigraphic diagnostics- §4.8 "
Radionuclides and radiopharmaceuticals for scintigraphy", passage "Immunoscintigraphy". It is mainly in tumor diagnosis, but also, for example, in the diagnosis of inflammatory foci using antigranulocyte monoclonal antibodies.
  Monoclonal Antibodies in Oncology
A number of monoclonal antibodies are used in oncology therapy, some of which are briefly listed :

- a chimeric monoclonal antibody that competitively binds to the extracellular domain of epidermal growth factor EGFR (HER1) and inhibits the binding of other possible tumor growth activators; panitumumab has similar effects .  
Trastuzumab (herceptin) acts as a monoclonal antibody against HER2 (Human Epidermal Receptor), binding to the extracellular domain of HER2 and thereby blocking epidermal growth factor access to its receptor; this prevents activation of the signaling pathway of cell processes and tumor growth (which is HER2 - positive). Pertuzumab, which binds to the dimerization domain of HER2 and thus prevents its dimerization with other HER receptors, has similar effects. The combination of trastuzumab + pertuzumab is being tested to increase the effect of HER2-signaling blockade (possibly + docetaxel), which shows synergistic activity. The combination of herceptin with aromatase inhibitors (such as anastrozole or ietrozole) has also been tested in hormone-dependent tumors. And in general, herceptin in combination with the chemotherapeutic agents capecitabine or 5-fluorouracil and cisplatin is indicated in various HER2-positive metastatic tumors.
Bevacizumab (Avastin) is a humanized monoclonal antibody to Vascular Endothelial Growth Factor (VEGF), which captures circulating VEGF in plasma, thereby inhibiting tumor neoangiogenesis. Ranibizumab (Lucentis), which is used in ophthalmology in the treatment of vascular-related macular degeneration of the retina, has similar effects.
Rituximab - a chimeric monoclonal antibody of the IgG1 type, specifically directed against the CD20 antigen, is a frequently used preparation in malignant B-lymphomas.
Ipilimumab is a monoclonal antibody that activates CTLA-4 targeting, where cytotoxic T cells can recognize and destroy tumor cells (it turns off the inhibitory mechanism and allows T cells to function). It is used to treat melanoma, non-small cell lung cancer, ca bladder and ca prostate.
Nivolumab is a human IgG4 anti-PD-1
monoclonal antibody, also acting as a control node inhibitor, that blocks activated T cells from attacking tumor cells. It is also used in malignant melanoma (in combination with ipilumomab), lung ca and kidney ca.
Atezolizumab acts as an inhibitor of PDL1 programmed cell death ligand. It is mainly used in non-small cell lung ca.
    Another possible mechanism involved in the anti-tumor effect of some monoclonal antibodies is the "labeling" of a cell on the surface of which the relevant receptor is present; the cells thus labeled are then attacked and destroyed by the body's immune processes. The efficacy of monoclonal antibodies in tumor therapy depends on the presence and function of appropriate receptors on tumor cell membranes. If these receptors are scarce, or are dysfunctional or mutated, the antibody is ineffective ...
    Monoclonal antibodies can also be "carriers" to which a suitable chemotherapeutic or radionuclide binds. A long-used radioactive preparation of this species is Ibritumomab Tiuxetan labeled with
90Y (Zevalin) for non-Hodgkin's lymphomas. More recently, lilotomab labeled 177Lu, targeted against CD37, is used here. However, the 177Lu or 225Ac labeled anti-PSMA ligand (PSMA-617) in prostate cancer appears to be the most promising, which is able to cure even advanced metastatic hormone-independent prostate cancer.
For more details see below "
Radioisotope Therapy", passae "Radioimmunotherapy".

Dual-targeted (bispecific) monoclonal antibodies
Various factors are involved in cancer (as well as some other serious diseases - inflammatory, autoimmune, degenerative) and there are a number of complex processes taking place with multiple signaling pathways between cells. Although monoclonal antibodies (MAb) can effectively affect specific signaling pathways, it is often difficult to achieve satisfactory curative results with one given preparation. Therefore, different types of chemotherapeutical are sometimes combined. However, this can bring additional problems with unwanted interference, side effects, the possibility of unwanted immune responses.
In the field of monoclonal antibodies, however, an alternative "more elegant" solution is also possible in principle: Develop such antibody molecules that have two binding sites aimed at two different antigens of the given process. These so-called bispecific antibodies (BsAb) with double targeting ("two in one") can then show better therapeutic effects than classical unidirectionally targeted monoclonal antibodies. Two mediators are targeted by a single molecule.
    Various specific ligands binding to different receptors can be incorporated into the Fab arms of antibody molecules in several ways. Either by cross-linking two antibodies with two different pairs of variable Fab fragments with different antigen binding. Or, instead of one of the chains, e.g. "b", another "b'
" with a different targeting is chemically attached to the Fab arm - a Fab' fragment is created. Or, alternatively, heavy and light chain variable regions cloned from two different antibodies are created in the IgG antibody. There are other more complicated options...
    We briefly list some bispecific antibodies that have already found clinical application :
-> Blinatumomab in malignant B-lymphoid cells in non-Hodkinson's lymphoma, specific for CD19 and CD3 on effector T-lymphocytes.
-> Mosunetuzumab is a monoclonal antibody used to treat follicular lymphoma. It binds bispecifically to CD20 contained in B-lymphocytes and to CD3 found in T-cells. T-cells are thereby stimulated to destroy lymphoma tumor cells.
-> Emicizumab is a bispecific antibody for the treatment of hemophilia A. It binds to both activated coagulation factor IX and factor X, whose activation it mediates.
-> Amivantamab is a bispecific monoclonal antibody used to treat non-small cell lung cancer (a special type with an insertion mutation of exon 20 of the epidermal growth factor receptor EGFR). It targets the epidermal growth factor (EGF) receptor and the mesenchymal-epithelial transition (MET) receptor.
-> Cadonilimab binds simultaneously to PD-1 and CTLA-4 antigens, which are expressed at much higher levels in tumor tissues compared to normal tissues. These two targets mediate increased cellular cytotoxicity and cellular phagocytosis. Bispecific cadonilimab (which replaces the combination of nivolumab and ipilimumab) shows promising efficacy in cervical, gastric, hepatocellular carcinoma.
-> Faricimab (RG7716) is used in ophthalmology for the therapy of vascular-related macular degeneration. It targets both the vascular endothelial growth factor VEGF-A and the angioprotein inhibitor ANG-2. It stabilizes the blood vessels in the retina and shows better effects than the currently used aptamer pegaptamib (Macugen) or the monoclonal antibody ranibizumab (Lucentis).

l Mimetic antibodies, Affibody
In addition to "real" monoclonal antibodies, so-called mimetic antibodies are also used - peptides or small proteins (with a molecular weight of about 3-20 kDa), which like antibodies can bind to antigens, but which are structurally not similar to the relevant antibodies
(the name comes from the Greek mimesis = mimicry, imitate). The main representative of these substances are the so-called affibodies consisting of three helices with 58 amino acids with a molecular weight of about 6 kDa. Their use for diagnostic imaging and targeted therapy is in development.
l Kinase inhibitors (thyrosine kinase inhibitors) are substances that block the signaling pathways of certain kinases (one or more), thereby inhibiting cell division and stimulating apoptosis. This can lead to slower tumor growth and attenuation of tumor angiogenesis. Kinases are enzymes (§5.2, section "Cells - basic units of living organisms", section "Proteins, enzymes, kinases"), which transfer the phosphate group from the adenosine triphosphate ATP to the acceptor, which has an OH group - the phosphorester of the acceptor molecule is formed. The tyrosine kinase transfers the phosphate to the hydroxyl group of the cyclic amino acid tyrosine bound in the protein, thereby activating the protein. Tyrosine kinase inhibitors (tinibs) are small molecules that bind to an appropriate site in ATP (adenosine triphosphate) to prevent phosphorylation of substances that are part of the intracellular signaling pathways by which a chemical signal captured by a receptor on the cell surface is transmitted to target structures in the cytoplasm or in the core. One of the important targets of biologic therapy is the epidermal growth factor receptor EGFR signaling pathway (Epidermal Growth Factor Receptor), also known as a human epidermal receptor, Her-1 (human epidermal receptor 1), a transmembrane glycoprotein (molecular weight about 170000). Another kinase that affects the regulation of cell growth, including angiogenesis, is the serine/threonine kinase mTOR (mammalian target of rapamycin). Thus, mTOR inhibitors may interact in particular by attenuating angiogenesis.
    Inhibition of kinases involved in oncogenic signaling pathways may suppress the proliferation of a given tumor cell clone. A certain advantage of these substances is that (unlike large protein molecules) they can penetrate cells by passive transport
, so that their activity is not linked to the presence of the respective receptors on the membranes of the tumor cells.
    Gefitinib is a quinazoline derivative that inhibits EGFR growth receptor tyrosine kinase activity (especially in the EGFR activating mutation), with erlotinib having a similar effect. Lapatinib binds to the intracellular portion of the HER2 growth receptor and inhibits its tyrosine kinase activity; it also acts as a dual inhibitor - in addition to HER2, it also acts on the intracellular activity of the HER1 receptor (ie EGFR). Imatinib primarily blocks BCR-ABL tyrosine kinase in some leukemia species; newer and more effective inhibitors of this type are nilotinib and dasatinib. The multikinase inhibitors sunitinib and sorafenib suppress the kinase activity of platelet-derived growth factor receptors, VEGFR and others (KIT, FLT3, ...) - they act as inhibitors of angiogenesis. For this purpose, the new mTOR kinase inhibitors temsirolimus and everolimus (derivatives of rapamycin - sirolimus mentioned above in the "anthracycline antibiotics" category), which block the P3K/Akt/mTOR phosphatidylinositol-3-kinase signaling pathway, are also being tested. The kinase inhibitor vemurafenib, which specifically inhibits the V600-mutated form of the B-raf protein, is also being tested; shows promising results in the treatment of malignant melanoma.
In some cases, it is useful to combine monoclonal antibody therapy with an appropriate tyrosine kinase inhibitor - for example, in a HER2-positive tumor, trastuzumab and subsequently lapatinib.
l Aptamers (Lat. aptus = capable, Greek meros = part) are short fragments of RNA or DNA (oligonucleotides, peptides; molecular weights 3-18 kDa) - specially prepared and sequenced ligands with high binding affinity to specific target molecules. They take on different three-dimensional structures, they are able to bind to different biomolecules (antibodies, growth factors, hormones, enzymes, amino acids). They can act as targeted inhibitors and also as "carriers" suitable therapeutic substances - "escort" aptamers. Their use, so far experimental, is an alternative to monoclonal antibodies. Aptamers can be produced artificially by biochemical methods in a wide range: RNA serves as a "library" of nucleotides, from which ligands of desired properties are prepared by repeated combinations with tumor antigens and selections (SELEX method - Systematic Evolution of Ligands by EXponential enrichment). With a bit of exaggeration, we can say that here we are artificially "imitating" the natural evolutionary process of natural selection (however, the selection here is not made by nature, but by a researcher - a biochemist). The selected aptamers created in this way can then be sequenced and produced artificially by biochemical methods; in this they have an advantage over monoclonal antibodies (which are prepared by immunizing of laboratory animals). In a sense, aptamers can be thought of as man-made, synthetic chemical antibodies.
    An example is pegaptanib, which binds to the vascular growth factor VEGF and thus prevents it from stimulating angiogenesis. Pegaptamib is currently used in ophthalmology
(under the name Macugen , bound to a polyethylene glycol polymer) to suppress unwanted excessive angiogenesis in macular degenerationof retina; however, it is now gradually being replaced by the monoclonal antibody ranibizumab (Lucentis), or most recently by the bispecific monoclonal antibody faricimab.
    The designation of aptamers by radionuclides could be promising - either
g- radionuclides for scintigraphic diagnostics, or b or a radionuclides for biologically targeted radionuclide therapy. For example, an anti-tenascin-C aptamer labeled with 99mTc and 111In (for glioblastoma), or an Anti-MUC1 aptamer labeled with 99mTc or 186Re (for breast cancer) is tested.
    A significant problem with aptamers is the way to "deliver" aptamers to specific pathological tissues and cells in the body - so that they are not damaged or broken down or destroyed *). Aptamers have a short lifetime (a few minutes to hours) and are quickly degraded from the bloodstream in the kidneys. Experiments are carried out with binding of aptamers to the surface of nanoparticles. After the nanoparticles come into contact with the cell, the aptamers react with specific receptors on the cell membrane and influence the processes in the cell.
*) This disadvantage does not apply to pegaptanib, which is not used systemically, but is applied directly to the eye (intraocular injection into the vitreous) and is therefore locally bioavailable well in the vascular system of the retina.
    Biological treatment of cancer includes several other special procedures :
Gene therapy of tumors is still in the stage of laboratory development, but its more significant application can be expected in the near future. Two different pathways of gene tumor therapy are being developed :
- A straightforward procedure within tumor cells seeks to "correct" a genetic variation that has led to malignant cell transformation by a targeted change. The difficulty of this approach lies not only in the laboratory biochemical complexity of introducing specific genetic information using a suitable RNA vector and reverse transcriptase, but also in the fact that several different genetic changes (mutations) are involved in the malignant transformation of cells. This is probably not only the changes in known DNA coding sequences, but also in the as yet unexplored "genetic junk", "unnecessary" DNA.
- In an alternative approach, the target of gene therapy is not directly tumor cells, but other cells and tissues that, under the influence of targeted genetic intervention, are modified so that they can begin to produce certain active substances that effectively block tumor growth.
l Telomerase inhibition is so far an experimental method directed against one of the above-mentioned factors of carcinogenesis: overcoming the Hayflick limit of cell division - their immortilization - due to active telomerase acting in tumor cells. Antitelomer vaccination is performed by applying telomerase to the body in order to elicit an immune response that would kill the telomerase in the tumor cells and thus prevent their unrestricted division. Unfortunately, telomerase inhibition also affects other cells, such as hematopoiesis, where telomerase performs its physiological function. Experiments with combined telomerase inhibition together with inhibition of tankyrase to potentiate the effect are also being experimented (telomerase and tankyrase work "synergistically in tandem cooperation"). Therapy based on telomerase inhibition can only be successful where telomerase is active. Recently, however, it has been found that telomerase is not the only factor in the immortilization of tumor cells, but mechanisms of homologous recombination of telomere sequences work similarly (it is discussed in §5.2, part "DNA, chromosomes, telomeres"). This "bad news" somewhat reduces the promising therapeutic potential of telomerase ...
l Hormone therapy of tumors is based on the fact that some types of tumor cells contain receptors for hormones - they are hormone dependent, their origin and development is dependent on the level of certain hormones. In breast cancer cells there are estrogen receptors, in prostate tumors for androgens.The growth of tumor cells can inhibit hormones with the opposite effect (hormone antagonists) or prevent the synthesis of hormones (castration in the prostate, ovariectomy in the breast ca), both methods are often combined, with the possibility of blocking receptors in hormone-dependent forms of tumors (especially breast cancer), which leads to the cessation of tumor cell proliferation. Selective estrogen receptor modulators such as tamoxifen are used for this purpose. Aromatase inhibitors (an enzyme involved in the synthesis of estrogens from testosterone, estradiol is formed) such as anastrazole, ietresol, formestane, also block the production of estrogens.
l Immunotherapy generally represents a targeted intervention into the body's immune system for a therapeutic purpose - to restore, strengthen, or modify the functions of the immune system. Unfortunately, in advanced cancer, the immune system usually does not respond to the tumor cells of one's own body (immunosuppression). One of the goals of tumor immunotherapy is to label and "make visible" tumor cells for the body's immune system, which can then "take care" of their destruction. Genetic changes in tumor cells result, among other things, in the emergence of new antigens, different from non-tumor cells. These tumor antigens may become a desirable target for immune responses, but only if we "serve" them properly to the immune system.
As the vector for the purpose of immune antitumor vaccination are particularly useful group of special cells from white blood cells called dendritic cells of the immune system, with many numerous protrusions on the surface (Greek dendron = tree, the cells have tree-shaped branched protrusions - dendrites). These cells initiate immune responses, differentiate foreign and the body's own substances (they can recognize various antigens, including tumor cells), are capable of phagocytosis of foreign particles. Subsequently, they mature, exposing on their surface parts of absorbed proteins (in our case tumor antigens) and thus activating T-lymphocytes, which completes the "destructive" immune response involving effector monocytes transforming into macrophages. Dendritic cells can thus become an effective tool for autologous cellular immunotherapy, which stimulates the body's own immune system to "engage" in the fight against cancer.

The procedure of anti-tumor immune vaccination consists of several stages :
- Sampling of tumor tissue, isolation and cultivation of tumor line cells;
- Collection of peripheral blood, separation of leukocytes and monocytes by leucopheresis;
- Growing in vitro cultures of native (immature) dendritic cells;
- Activation of dendritic cells by uptake of cells of a given tumor line (their antigens);
- From the activated (mature) dendritic cells, which expose tumor antigens on their surface, a final vaccine is prepared, which is applied back to the organism;
- Activated dendritic cells migrate to the lymph nodes, where they activate effector T cells;
- Activated cytotoxic T-lymphocytes recognize and kill tumor cells (N).

Dendritic cells can be obtained by culturing for several days from monocytes extracted from the patient's peripheral blood *). In addition, a sample of tumor tissue is taken. If then (in vitro, using disrupted or apoptotic tumor cells) dendritic cells absorb the tumor antigens from the tumor tissue and then thus stimulated cells are applied back into the body, they have the ability to stimulate the immune system for "fight" against to the original tumor cells. Subsequently, activated effector T-cells travel into the tumor site and selectively attack the tumor cells. The method is still in the stage of experimental clinical studies.
*) To obtain a larger number of leukocytes and monocytes, a special sampling separation method called leucopheresis or leukapheresis (Greek leukosis = white , these are white blood cells - leukocytes; afairesis = take ) is used.: the blood circulates through a centrifugal separation unit, where the leukocytes are separated, while other components of blood (especially plasma) return to the patient's circulation.
    Although tumor cells produce antigens that, in principle, the immune system can use for their identification and subsequent targeted destruction by cytotoxic T-lymphocytes (CTL), here exters an inhibitory mechanism of antigen associated with the CTLA-4 proteinn, the binding of which to the CTL receptor shuts down the cytotoxic response. This mechanism, on the one hand, prevents excessive adverse immune reactions (autoimmunity), but on the other hand allows tumor cells to survive. The anti-CTLA-4 monoclonal antibody ipilimumab (MDX-010), which binds to CTLA-4, blocks its inhibitory function and allows CTL to continue to destroy tumor cells, has been shown to enhance anti-tumor CTL immunity.
   Another newly tested method based on the immune system is the blocker protein CD47 - anti-CD47. The CD47 protein is physiologically present on the surface of blood cells and its task is to protect them from its own white blood cells. However, many tumor cells also have the CD47 protein on their surface, which protects them from white blood cells and therefore cannot be destroyed by the immune system. By applying an anti-CD47 blocker, the immune system is stimulated to kill the tumor cells. Although there is also a loss of blood cells, which need to be supplemented ...
    The above-mentioned monoclonal antibodies also belong to the category of immunotherapy. Molecular biological chemotherapy (immunotherapy, monoclonal antibodies) is often suitable to combine with classical cytostatics - to enhance the therapeutic effect. For example, the combination docetaxel with trastuzamab is used, 5-fluorouracil + oxaliplatin (FOLFOX), or 5-fluorouracil + irinotecan (FOLFIRI), with bevacizumab or cetuximab, capecitabine with trastuzumab or lapatinib, and several others. Recently, the application of cytostatics and monoclonal antibodies labeled with therapeutic beta-radionuclides, which represent combined molecular chemo-radiotherapy, has been tested. Some such methods and preparations are mentioned below in the part "
Radioisotope therapy", section "Radionuclide therapy of tumors and metastases" and "Radioimmunotherapy".
Abscopic effect

In some cases a synergistic effect of immunotherapy and radiotherapy is observed. Rarely occurs so-called abscopic effect
(i.e. off target - Lat. ab = outside, away; scopium = target, angle of wiev) when after local radiotherapy certain tumor lesions, recede systemically even other lesions that have not been irradiated. Radiotherapy may induce an immunoeffect against further metastases of the same tumor (for more details, see §5.2, passage "Bystander-Abscopal effect ").

Alternative methods
In addition to chemotherapy and radiotherapy, some alternative methods of cancer therapy are sometimes used or tried. We will mention two based on temperature :

- local heating of the target tissue to a temperature higher than 43 °C, causing inhibition of DNA and protein production, together with a reduction in tumor vascularization. Tumor tissues usually respond more sensitively to heat than normal healthy tissues. In healthy (normally perfused) tissue, vasodilation occurs when heated, which removes heat more efficiently through the blood and reduces heating. The vessels formed by neoangiogenesis in the tumor are chaotic, functionally imperfect and possibly compressed with tumor mass. Tumor vessels are not able to effectively regulate blood flow, so when the tumor is heated, vasodilation does not occur and the tumor heats up more than the surrounding healthy tissues. In particular, large and hypoxic tumors, which are less sensitive to radiotherapy, are therefore suitable for the treatment of hyperthermia.
  At temperatures above 43 °C, denaturation of proteins
(including cell membrane proteins and microtubules, causing changes in membrane potentials and ion concentrations) begins to occur, leading to cell death, predominantly by necrosis. In addition, heat stress express the HSP heat stress proteins (Heat shock proteins), the most common is Hsp70, which in addition to its anti-stress effect (bind to a hydrophobic amino acid sequence partially damaged proteins, which allow the correction to the correct spatial arrangement; further promote the degradation of damaged proteins), their "chaperone" activities allow binding to the antigen, and transport to cell membranes, where these antigens are presented (via the transmembrane glycoprotein MHC 1) and thus stimulate the immune system - the formation of cytotoxic T-lymphocytes specific for a given type of tumor. Furthermore, enzymatic cell repair mechanisms (such as excision repair, homologous recombination, non-homologous end-joining - see §5.2, passage "Repair processes") are heat-sensitive - thermolabile. This can be used for the synergetic effect of combining hyperthermia with radiotherapy or chemotherapy - thermoradiotherapy (hyperthermic radiotherapy) or thermochemotherapy.
   Non-invasive heating of the tumor inside the body can be achieved by electromagnetic waves or ultrasound. Recently, a promissing method of high-intensity focused ultrasound HIFU (High-Intensity Focused Ultrasound) has apeared. Focusing of ultrasonic waves is achieved using a specially shaped (concavely curved) transducer. High-intensity ultrasound focuses on the tumor, within which energy is converted into heat. The tissue temperature rises to 65 °C, during which thermal ablation occurs - the temperature kills the tumor cells, but when properly targeted, does not damage the surrounding healthy tissues. A rapid and short-term increase in local temperature (within 2-3 seconds) destroys the target tissue by coagulation necrosis - we literally "cook" the tumor, while the surrounding structures are not damaged. HIFU therapy is suitable to perform during MRI navigation.
   Furthermore, the hyperthermic method could activate chemotherapeutic drugs directly in tumors. The chemotherapeutic is "wrapped" in heat-sensitive microscopic particles ( liposomes - particles coated with a fat layer) and applied to the bloodstream. At a normal body temperature of about 37 °C, the particles pass through the blood vessels undisturbed and have no toxic effects on the body. When they enter the tumor in this way, they can be locally heated by the focused HIFU waves to a temperature higher than 42 °C, at which point the liposome envelope becomes porous and the drug is released directly into the tumor. The therapeutic effect of the drug is thus targeted in the tumor, with the minimization of undesirable side effects in other parts of the body.
Cryotherapy (Greek cryos = cold )
(also called cryosurgery) consists in the application of very low temperatures in order to destroy the target tissue (cryodestruction). It is used in various medical fields (dermatology, ophthalmology, gynecology, surgery) even for the treatment of non-malignant diseases. In oncological indications, it is the destruction of a tumor by freezing by introducing freezing probe - a cryocauter, cooled mostly by liquid nitrogen. The rapid freezing of the tissue causes its damage by the formation of ice crystals inside and outside the cells, with subsequent necrosis of the frozen cells. To achieve the desired effect, the target tissue must be cooled to below -20 °C with a high freezing rate
approx. 30 °C/sec. Slow freezing would dehydrate the cells, which could survive after thawing. On the contrary, the subsequent thawing should be much slower so that the cells are exposed for a long time to mechanical damage by recrystallizing ice and also to the toxic action of the intracellular fluid, in which the concentration of salts and ions has risen sharply. The cryotherapy method is used for tumors accessible by direct application of the cryocauter, most often for skin lesions.
Note: Issues of chemotherapy and other non-radiation methods we have outlined here only briefly and marginally, due to the complexity of the interpretation of the principles and current possibilities of cancer therapy; further details of complex biochemical reactions (often not yet fully explored) in chemotherapy and biological treatment lie beyond the scope of our physics-oriented treatise on radiotherapy ...

of tumor diseases (cancer) is based on the effects of ionizing radiation on living tissue
(the mechanisms of these effects are described in detail in §5.2 "Biological effects of ionizing radiation").
Can radiation heal cells ?
The name "Radiotherapy" may give the impression that "radiation can cure the damaged or mutated cells?". Non! - ionizing radiation can only kill cells
! The therapeutic effect here consists in the targeted elimination of tumor cells in the organism.
   Sufficiently high doses of ionizing radiation are able to inactivate and kill cells, in this case tumor cells. In tumor tissue, it is necessary to destroy mainly clonogenic stem cells, the unrestricted division of which causes cancer. Radiation damage to the tumor's vascular supply can also play a significant role in stopping tumor growth. Radiotherapy can thus be an effective local (or local-regional) method of cancer therapy
(and possibly also some other focal diseases and disorders).
    The goal of classical radiotherapy is the reproductive sterilization of clonogenic tumor cells by radiation-induced apoptosis. In stereotactic radiotherapy, in addition, an ablative approach is even applied - immediate destruction of cells by necrosis, caused by a high single dose of radiation.
 Radiotherapy - local treatment
Unlike the systemic chemotherapy and biological treatment discussed above
("Chemotherapy and biological treatment"), external radiotherapy is a local method that is able to successfully eliminate specific tumor lesions in known locations. If the cancer is disseminated - metastatic, the final result of external radiotherapy is problematic: after the successful treatment of one lesion, progression may occur in other places...
   Radiation eradication of tumor cells can be part of effective curative therapy to completely cure cancer, especially at the localized stage. In more severe and advanced cases, then palliative therapy, alleviating and slowing down the course of the disease and its difficulties. After surgical removal of the tumor lesion, the so-called adjuvant radiotherapy is often applied - auxiliary, supportive or securing treatment after surgery, to reduce the risk of recurrence due to possible micro-infiltration in the vicinity of the original tumor. In certain cases, preoperative so-called neoadjuvant radiotherapy is used before surgery - to reduce the extent of the tumor
("downstaging") and thus improve its operability, as well as to reduce the viability of tumor cells and thus reduce the risk of local or metastatic infiltration (during tumor surgey, tumor cells may be release into the environment, lymphatic system and bloodstream). The surgery is then performed about 6 weeks after radiotherapy, when the acute radiation symptoms have disappeared and late changes have not yet occurred (see below "Side effects of radiotherapy - radiotoxicity").
    Perioperative (intraoperative) radiotherapy is rarely used - direct irradiation of the tumor site or its remnant, uncovered during surgery. This method is complicated by the fact that it is necessary to either install an irradiation device in the operating room, or to be able to quickly transport the patient to the irradiation room and then back to the operating room. From this point of view, mobile devices with a miniaturized X-ray tube or an electron gun and a target, which is applied to a target lesion (eg a cavity after resection of the tumor itself) and irradiated with low-energy X-rays (with an energy of several tens of keV) with a high dose and ionization density, which can in principle also be used in laparoscopic operations.
    The majority of this §3.6 will be devoted to radiotherapy of tumors.
Combination chemo-radiotherapy
To improve the results of cancer treatment, in some cases it is useful to combine the two above-mentioned therapeutic modalities: concomitant - accompanying, complementary, additional therapies, a special case of multimodal therapy. The benefits of simultaneous application of chemotherapy with radiotherapy can be basically of three types :

  Additive effect - the effect of radiotherapy and chemotherapy for the destruction of tumor cells adds up (without direct interdependence). The additive effect of chemotherapy occurs on both the irradiated cells of the target volume and chemotherapy can also cause the elimination and attenuation of micrometastases outside the irradiated volume.
  Radiosensitizing effect - chemotherapy enhances the biological effect of ionizing radiation on cells (increased DNA fragility due to chemically bound cytostatics, inhibition of DNA repair mechanisms, or by appropriate temporal influencing of the cell cycle to a phase more sensitive to radiation, eg G2) - potentiation radiotherapy. One such cytostatic agent that increases the sensitivity of tumor cells to radiotherapy is sirolimus (rapamycin- mentioned above as an anthracycline antibiotic); the combination of sirolimus + radiotherapy is better tolerated in terms of side effects than the combination of radiotherapy with most other chemotherapeutics.
Tumors with increased expression of EGFR (= HER1) and HER2 growth receptors generally have increased radioresistance, as the intrinsic signaling pathways of these receptors are involved in activating DNA repair processes upon ionizing radiation damage. By applying targeted biological treatment against growth factor receptors - cetuximab, trastuzumab, gefitinib, lapatinib, etc., the growth factor signaling pathways are interrupted, DNA repair capacity is reduced and therefore an increase in tumor radiosensitivity can be expected.
l Anti-repopulatory effect - cytostatics, and in particular some targeted biologic therapies, such as monoclonal antibodies against growth factors, reduce the repopulation of tumor cells during time-prolonged fractionated radiotherapy, leading to a more efficient killing of a larger fraction of tumor cells by radiation.
  In chemosensitive tumors, chemotherapy can also cause a reduction in the volume, a kind of "shrinkage" of the tumor (similar effect to neoadjuvant chemotherapy). Such a reduced tumor is then easier to treat with radiotherapy, both by reducing in the number of tumor cells alone and possibly by improving blood flow and oxygenation, which leads to increased radiosensitivity due to the oxygen effect. The resulting effects of chemo-radiotherapy are analyzed in more detail below in the section "
Prediction of the radiotherapeutic effect".

Physical and radiobiological factors in radiotherapy
The optimal therapeutic effect of radiation is achieved by co-production of two types of factors :

Physical factors - selective introduction of a sufficiently high dose of radiation into a pathological lesion by a suitable irradiation technique, using physical properties of radiation. We will discuss these physical aspects in detail in most of the text of this chapter; here we first briefly analyze the radiobiological aspects :
Biological factors - the type of tumor and the properties of the surrounding healthy tissue. For radiotherapy is very important radiosensitivity of specific tumor type, as well as the difference in radiation sensitivity between tumor and healthy tissue *). Lymphomas, leukemia, seminoma are highly radiosensitive. Carcinomas, such as prostate adenocarcinoma, are moderately radiosensitive. Gliomas, sarcomas, melanoma, squamous cell carcinoma of the skin are radioresistant.
*) It is the risk of damage to the surrounding healthy tissues and organs, that is the main limiting factor in delivering a sufficiently high dose to the tumor site. By critical organ or tissue we mean a structure in the organism whose radiation damage would have serious health consequences, or in the case of vital organs even death. Therefore, a certain so-called tolerance dose must not be exceeded during radiotherapy in these critical organs to prevent their irreversible damage.
Therapeutic ratio
For the possibility of achieving a good curative effect of radiotherapy, the most important thing is often not the actual radiosensitivity of the tumor tissue, but rather the ratio of the radiosensitivity of the tumor and surrounding healthy tissue - the so-called therapeutic ratio TR (Therapeutic Ratio). It can be quantified in different ways *): by comparing the dose-response curves of cell survival N/N
0 (Fig.5.2.3c) for tumor and surrounding healthy tissue (from the ratio of gradients or areas under these curves), or with the help of biologically effective dose BED, or by comparing the probability quantities TCP and NTCP (see below "Prediction of radiotherapeutic effect - TCP, NTCP"). To quantify the therapeutic ratio of TR, we obtain different indices TRN/No , TRBED , TRTCP , whose numerical values are different and must be considered separately. The therapeutic ratio can be improved by fractionation of radiation, combination with chemotherapy, improvement of oxidation of the tumor site (overcoming hypoxia, use of densely ionizing radiation with high LET) - is discussed below.
*) Therapeutic options were previously evaluated using the so-called Paterson graph, which shows the dependence of the relative number of killed tumor cells on the radiation dose. The same graph also shows the dose dependence of the risk of irreversible damage to the surrounding healthy tissue. In the favorable case, the curve for tumor tissue is on the left, the curve of the probability of complications of healthy tissue is shifted on the right. The width of the gap between these two sigmoidal curves is sometimes called the therapeutic width. More complex evaluations are now performed using special TCP and NTCP graphs - see below "Prediction of radiotherapeutic effect - TCP, NTCP".
Time fractionation of irradiation
Tumor tissue that is in a state of intense (pathological) cell division is usually more sensitive to radiation than healthy tissue
(it is discussed in more detail in §5.2 "Biological effects of radiation"). Fractional irradiation is usually used, where the total dose is divided into a number of smaller daily doses *), applied over a number of days (approximately 3-5 weeks, see below "Fractionation in practice").
*) Single irradiation of small lesions with a high dose (or substantial reduction of the number of fractions to 2-5) allows the Sterotactic radiotherapy thanks to the possibility of very precise targeting of the canceroletal (or radioablative) dose into tumor, with less load on the surrounding critical tissues and organs.
    There are basically two reasons for the time fractionation of the radiation dose :
1. Healthy tissue cells usually have a higher ability to repair radiation damage than tumor cells. When dividing the dose into a number of smaller fractions, applied individually after completion of the repair processes in the cells, the resulting cumulative biological effect on tumor tissue is generally higher than on healthy tissue, which has a greater regenerative capacity.
Radiobiological aspects of radiation fractionation are discussed below. The radiotherapeutic effect on the tumor tissue itself is sometimes expressed by a probabilistic quantityTCP, defined below.
2. In each tissue, including tumor tissue, there are cells at different stages of the cell cycle, in which they have different radiobiological sensitivities. Therefore, a single dose of radiation may not be optimal for all cells, which may not be in the most sensitive phase. If the total dose is divided into several fractions with a suitable time interval, then after each such dose, the part of the cells which has just reached the most sensitive phase at that time can be most efficiently destroyed.
Dependence of radiation-biological effect on dose and its time schedule - LQ model
The dependence of deterministic radiation effect on dose and its time schedule is analyzed in detail in §5.2 "
Biological effects of ionizing radiation", part "Dose-biological effect relationship", where the so-called linear-quadratic model (LQ) is introduced - see the section "LQ model", Fig.5.2.3c. There is also derived a basic equation of dependence between dose D and the surviving fraction of cells N/No in (semi)logarithmic scale :
             -ln (N / No ) = a .D + {2. [(1-e - l .T ). (1-1 / l .T)] / l .T } . b .D 2 - ln2.T / T2r ,
a and b are the factors indicating the probabilities of damage a- and b -processes, T is the irradiation time, l is the rate of cell repair, T2r is the doubling time of the number of cells by repopulations. The coefficient in angle brackets {...} is the so-called Lea-Catcheside factor, which captures the effect of cell repair during irradiation. The linear member a.D is dominant for early-reacting tissues (with higher cell proliferation), the quadratic member b.D2 is more pronounced in late-reacting tissues. Basic linear-quadratic dependence N/No on the dose (D) is shown on a reduced scale for illustration below, in Fig.3.6.0 a. The general equations of the LQ model have rather theoretical significance; for practical applications in radiotherapy, simpler special relationships for specific irradiation conditions and techniques are derived from them (see below "Irradiation fractionation"). These general radiation-biological mechanisms are approached in practice by some other individual biological influences, which are sometimes difficult to include in one LQ model.
Individual biological factors - " 6 R "
Biological effect of ionizing radiation in relation to the radiation dose, its time schedule and possibly volume distribution is influenced by several factors and biological processes during irradiation (whose names can be formulated so as to beginning with the letter "R") :
× Radiosensitivity
of the irradiated tissue is given by the sensitivity of individual cells to radiation damage; it generally varies considerably for different types of tissues. In the linear-quadratic model, radiosensitivity is implicitly contained in the coefficients
a and b. However, each tissue is in fact a heterogeneous cell population, containing cells with different radiosensitivity - with different coefficients a, b: the resulting survival curve [ln (N/N0)] (D) is then a superposition of several different LQ curves.
is the ability of cells to repair their important structures, especially DNA, damaged by ionizing radiation or other influences
(cell repair processes are described in more detail in §5.2 , section "Repair processes"). Repair processes have a time dimension: they take a certain amount of time (given by the coefficient l - the rate of cell repair), and the repair must be carried out before further damage prevents successful repair. The repair processes, which take place continuously during irradiation, thus lead to a "dose rate effect". In the LQ model, the repair is included in the additional coefficient RG º {...} in the b- member.
× Repopulation
Upon exposure to ionizing radiation, some cells die, but other cells normally divide and may eventually replace destroyed cells. This cell repopulation and tissue regeneration is provided by clonogenic stem cells. Repopulation is quantified by the rate of recovery of a certain number of cells, or the time T
2r of doubling of the number of cells. In the LQ model, the repopulation is captured in the additive term RP º ln2.T/T2r . Exponential tumor growth is assumed here, which is approximately met only in the initial stages of growth of miniature tumors, with the growth of the tumor the growth rate slows down.
× Redistribution
Different cell types at different stages of the cell cycle are differently sensitive to ionizing radiation. During the actual exposure, the so-called redistribution of cells can occur - a change in the representation of different types of cells in the tissue *). During irradiation, more clonogenic stem cells and cells in the G1 and G2 phases of the cell cycle decrease, while effector daughter cells and M and S phase cells in general will be relatively greater representation. Stem cells are more radiation-sensitive than mother cells (see also §5.2 ); the goal of radiotherapy is to kill tumor stem cells. The redistribution effect leads to changes in intratumorous radiosensitivity during irradiation, as well as to specific side effects on healthy tissue *) .
*) Complicated temporal dynamics of cell redistribution
The processes of cell redistribution during irradiation have a complex time dynamics and occur both in the tumor target tissue, as well as in the surrounding healthy tissues affected by radiation. In the first part of the exposure, clonogenic stem cells are declining faster, because they are (due to the faster cell cycle) more sensitive. This is followed by a gradual loss of daughter effector cells, which reduces the function of the irradiated tissue. A regulatory mechanism is in place to preserve tissue functionality, leading to a partial loss of division asymmetry stem cells, which begin to divide symmetrically into two effector cells each; thereby (at the cost of loss of stem cells) the functionality of the tissue is temporarily preserved. If exposure continues, as the number of clonogenic cells falls below a certain critical level (threatened by stem cell disappearance and subsequent tissue death), another regulatory mechanism occurs to trigger accelerated stem cell repopulation to maintain their population necessary for tissue regeneration. This reduces the production of effector daughter cells, which are no longer sufficient to cover the functional need for tissue - there is a clinical manifestation of deterministic radiation effect in healthy tissue, acute radiotoxicity (cf. the passage "Acute radiation sickness" in
§5.2; on the side effects of radiotherapy - early and late radiotoxicity, is briefly discussed below - "Strategic goal of radiotherapy"). If the exposure continues (fractional irradiation) and the tolerance dose of the tissue is not exceeded, the proliferated stem cells can produce effector cells in sufficient numbers to ensure the basic (albeit reduced) functionality of the tissue; some "emergency steady state" may occur with the reduction or disappearance of previous acute problems in healthy critical tissue. In the case of a tumor lesion, on the other hand, it is desirable to deliver a sufficiently high dose to overcome the repopulation of clonogenic tumor cells - to reduce them to a zero level, leading to the death of tumor tissue.
Reoxygenation - oxygen effect
The atoms of oxygen, contained in water and other molecules in the tissue, play a dominant role in the radiobiological effect - oxygen radicals and peroxides generated by radiation effectively damage DNA in cells. During tumor growth, as the tumor mass increases, there is often a lack of oxygen in the cells, so-called hypoxia. Hypoxia occurs especially when the tumor grows faster than the capillary vascular network of tumor neoangiogenesis. In hypoxic tumor cells, a much slower metabolism takes place (they often remain in the G
0 phase of the cell cycle) and during irradiation there is a lower formation of oxygen radicals - these cells therefore have reduced radiosensitivity, they are radioresistant. The surviving fraction of these radioresistant hypoxic cells may be a potential risk of cancer recurrence.
    Irradiation can cause some reoxygenation (reduction of hypoxia) of tumor foci: reducing the number of tumor cells reduces total oxygen consumption and reducing the tumor can also reduce intratumorous pressure, loosen capillaries, and improve blood flow and supply to the remaining cells. The effect of reoxygenation is positive for radiotherapy - it increases the radiosensitivity of the tumor tissue *) and thus improves the therapeutic effect when using doses with limited tolerances of healthy tissues. The oxygen effect is significant especially when using sparsely ionizing radiation (photon radiation
g or X is most often used), where the indirect radical mechanism of the radiation effect predominates. In densely ionizing radiation, where there is an increased proportion of direct intervention mechanism (and also increased recombination of radicals), the effect of oxygen (oxygenation) on radiobiological effects is less significant (see "Hadron radiotherapy").
*) Influence of oxygen content on radiosensitivity - so-called oxygen effect - is sometimes expressed by the factor OER ( oxygen enhancement ratio ), which indicates the relative increase in the biological efficiency of radiation in the presence of oxygen (normooxidation) compared to its absence. The ratio of OER between completely anoxic and normoxic tissue reaches a value of about 2.5.
    The processes of redistribution and reoxygenation vary widely between different tumor tissues and are difficult to predict. It is therefore difficult to introduce them into the LQ model, they are usually considered separately. All these individual effects can result in changes in the radiation sensitivity of cells and tissues during irradiation
, leading to further deviations from the dependencies of the ideal LQ model. Recently, the influence of the so-called bystander effect has also been discussed (see §5.2 "Biological effects of ionizing radiation", passage "Bystander-Abscopal effect"), which could perhaps somewhat correct the effect of mild inhomogeneities in tumor tissue irradiation - increase the effect in underexposed parts target tissues.
× Volume factor - Radiation volume
The sixth factor, which is sometimes important for the resulting radiobiological effect in the tumor tissue and especially in the critical organs, is the volume distribution of the radiation dose - the size of the irradiated volume (Radiated volume). At the cellular level, the biological effect is determined primarily by the size of the dose; the same is true for local tissue effects. Therefore, in organs with a serial arrangement of functional parts (spinal cord, esophagus, intestine, optic nerve), the resulting radiobiological impact is dependent on the maximum local dose: at high local dose, the serial organ can be radiation "disrupted" with irreversible impairment of its function. In contrast, volume organs with parallelby arranging functional parts (lungs, liver, etc.) they tolerate high local irradiation well, even above 80Gy (causing failure of only negligible functional parts), but even relatively weaker irradiation (approx. 20Gy) of their entire volume can significantly impair their function - the resulting organ effect depends on the average dose per whole organ.

For various radiation sensitivities of tissues and organs and their division into serial and parallel, see
§5.2, section "Local tissue and organ radiation effects".
Irradiation fractionation according to the LQ model
The parameters
a, b, l, T2r in the equation of the dependence of the surviving fraction of cells on dose D and irradiation time T according to the linear-quadratic model, as well as other biological factors of redistribution and reoxygenation, are different for individual tissue types, especially for healthy and tumor tissues. This dependence can be used to optimize the resulting radiation response of tumor tissue with respect to healthy tissue using a suitable time schedule - fractionation - of radiation dose. The total radiation dose D is distributed into individual fractions d i (i = 1,2, ..., n) with irradiation times t i . For a detailed analysis, the general Lea-Catcheside dose-time integral (derived in §5.2, "LQ model") can be used in the equations of the LQ model. However, in the case of evenly distributed fractions d i º d (D = n.d), the duration of which t i = t is short in comparison with the total duration T of radiotherapy treatment, the simplified equation can be used for total therapy, substituting the following values for dose and time variables :
In the linear a - member we substitute the total dose D = n.d, which is proportional to the number of damages by the double a -process (a-processes in individual fractions do not interact with each other, they are composed linearly).
l In the quadratic b- member it is important, that the square of the dose D2 in the LQ model during fractionation is not D2 = n2.d2 (as would result directly from the squaring of the relation D = n.d), but the number of fractions n appears in 1.power: b .D2 = b .n.d2. The exact derivation of this fact lies in the solution of the general Lea-Catchesid integral. However, it also follows from the physico-biological mechanism : according to the theory of dual radiation action (see §5.2, section "Intervention and radical theory of radiation effect"), quadratic dose dependence refers to a single absorbed dose, in this case to dose d of one fraction - b .d2; the total effect is formed by the sum of n independent fractions, ie n.b.d2. The individual fractions here do not interact with each other.
l Time T in the Lea-Catchesid coefficient at quadratic b-member we replace by the exposure time of the irradiation fraction t (during which the repair occurs).
l For the time T in the additive repopulation member, we take the total time T of the given radiotherapy (assuming that the continuous repopulation of cells occurs during the whole therapy at approximately constant rate).
    The resulting equation of the LQ model for (regularly) fractionated irradiation with the total dose D = n.d during the total time T divided into n fractions with sub-doses d and exposure times t , will then be :
            -ln (N / No) = a .D + n . { 2. [1- (1-e -l .t / l .t)] / l .t }. b .d - ln2.T / T2r   .
The repair mechanism may be more pronounced in LDR brachytherapy, especially in the late stages of permanent brachytherapy (see "
Brachyradiotherapy" below). In EBRT teletherapy, the exposure time of the individual fractions t is short (about tens of seconds to several minutes) due to the rate of cell repair: t << 1 / l , so the Lea-Catcheside factor {...} can be set equal to 1 (there are no interactions between the individual fractions) and the resulting effect will be :
           -ln (N / No) = a .D + b . n .d 2 - ln2.T / T2r = ( a + b .d) .D - ln2.T / T2r   .
When a single short-term irradiation (t = T << T
2r), does not appear the additive repopulating member (ln2.T/T2r ) ® 0.
    For a basic analysis of fractionated radiotherapy can be neglected temporal effects reparation and repopulation - we come from the basic equation LQ model :
             E º -ln (N / No) = a .D + b.D 2   ,
where the logarithm of the surviving fraction of cells N/N
o is denoted for brevity by E (a kind of "irradiation efficiency"). With regularly fractionated irradiation, D = n.d, so simple algebraic adjustments gradually give (justification of first power of n1 at b .D2 was given above) :
            E = a .n.d + b .n.d 2 = n.d (a + b .d) = D. a . [1+ (b / a) .d] = D. a . [1 + d / (a / b)]   .
We see that the radiobiological effect increases with increasing dose on fraction d and also depends on the value of the
a/b ratio for the irradiated tissue. At high doses per fraction, the radiation effect is significantly higher; at a given dose D, the effect is highest when applied once, in one fraction (n = 1, d = D). If we apply a larger number of fractions n with a lower dose d to the fraction, we must increase the total dose D to achieve the same biological effect .
Biologically effective dose BED

The logarithmic irradiation efficiency E is therefore proportional to the total dose D with the coefficients
a and [1 + d /(a/b)]; it is this second coefficient that expresses the relationship of the biological effect to the fractionation of the dose and the a/b ratio of a given tissue. To express the dependence of the biological effect of radiation on dose fractionation, a derived biophysical dose quantity biologically effective dose of BED (biological dose equivalent) is introduced :
             BED º E / a = D. [1 + d / (a / b)]   .
It can be said that BED = (physical dose)
x (proportionality coefficient); this proportionality factor [1 + d / (a/b )] (relative efficiency) shows how the biological effect of irradiation depends on the fractionation and the ratio a/b for a specific irradiated tissue. Since limd ®0 BED = D, the BED is a fictitious dose that would lead to the same biological effect if the total dose D were supplied in an infinitely large number of infinitesimal fractions (or for an infinitely long time with an infinitesimally low dose rate - if however, we neglect the time factor of cell repair and repopulation).
    Specific BED values are expressed in dosage units [Gy] provided with an index given by the numerical value of the
a/b ratio for a particular tissue - BEDa/b . E.g. a dose of 60Gy, applied in 30 fractions of 2Gy, will form in the early reacting tissue (fast-growing tumor tissue, skin) with a/b = 10, biologically effective dose of BED10 = 60. (1 + 2/10) = 72 Gy10 , while in late-reacting tissue (lungs, liver, kidneys) with a/b = 3, BED3 = 60. (1 +2/3) = 100 Gy3 .
    The concept of BED is important above all as a useful tool for comparing the effects of different fractionation regimes: the total dose D
1 applied in n1 fractions d1 gives the same (equivalent) biological effect as the dose D2 in n2 fractions of size d2 if it leads to the same BED value: D1.[1 + d1/(a/b)] = D2. [1 + d2/(a/b)]. Also, the above-mentioned therapeutic ratio of TR radiation sensitivity of tumor tissue and surrounding healthy tissue can be quantified as the BED ratio for tumor tissue "TU" and normal healthy tissue "NT": TRBED = BEDTU / BEDNT .
LQL model
In hypofractionation regimens with high doses per fraction (d > 5, 10 or more Gy/fraction), which allow the advanced conformational techniques of stereotactic radiotherapy and HDR brachytherapy described below, a certain discrepancy was shown between expected and observed effects: the classical LQ model at these higher doses per fraction somewhat overestimates the biological effect of radiotherapy, predicting higher damage to normal NTCP tissue. As if the curves of the surviving fraction of cells -ln (N/N
o) (on a log-linear scale) at higher doses actually showed an increased proportion of the linear component than the quadratic. To capture these clinical findings, use is sometimes empirical model modification LQ called LQ-L model (linear-quadratic-linear), which for higher doses / fraction (greater than 2. a/b , in practice > cca. 6Gy) adds additional linear component increasing the surviving fraction of cells. Other modifications of the LQ model lead to similar results - generalized gLQ model, USC (universal survival curve), KN (Kavahagh-Newman) model, PLQ (Pade Linear Quadratic), LQC model (linear-quadratic-cubic), see §5.2, part "Deviations and modifications of the LQ model".
Fractionation in practice
The same radiation dose applied in a shorter time (at a higher dose rate) has greater biological efficiency - cf. also Fig.5.2.3 in §5.2, part "
LQ model". From a radiobiological point of view, the most effective would be a single irradiation *) of a given deposit with the required radiation dose of several tens of Gy. However, the problem here would be the high acute radiotoxicity to the surrounding healthy tissues, which always receive a certain (albeit smaller) dose of radiation together with the target foci. Therefore, it is necessary to divide the curative radiation dose into a larger number of smaller parts - fractions. By suitable fractionation it is possible to achieve that in the time interval between fractions there is a partial reparation and regenerating healthy tissue, which is then able to tolerate the load of the next dose. However, this also increases the tolerance of the tumor cells (although usually less than in the cells of healthy tissue), so that it is necessary to increase the total dose to the tumor site.
*) However , let us consider still point 2 in the paragraph "Time fractionation of irradiation"..?..
    The most common fractionation, normofractionation, consists in the application of about 2Gy 1
x per day (5 days a week), for a period of 5-8 weeks (5w-8w), total dose about 60-80 Gy. From a radiobiological point of view, the optimal fractionation scheme depends on the type of tumor, whether it is slow or fast growing. In fast-growing tumors, the so-called hyperfractionation is used, in which more smaller doses (approx. 1.2 Gy) are applied at shorter time intervals, eg 2-3 x per day, to limit the rapid repopulation of clonogenic tumor cells. The slower dividing cells of healthy tissues (due to the time interval between fractions) can be regenerated, the risk of late radiotoxicity is reduced. The whole irradiation process is often shortened and accelerated here - so-called accelerated radiotherapy. Are used classically 2 fraction/day, as well as the CHART mode (Continuous Hyperfractinated Accelerated Radiotherapy), hyperfractionated irradiation 3x day, and continuous even over the weekend. The opposite procedure, so-called hypofractionation, when only 2x or 1x weekly is irradiated, it is used in palliative therapy, HDR brachytherapy and sometimes also in radioresistant and slow-growing tumors.
    Single (one-time) irradiation with a high dose of tens of Gy radiation is used in the so-called stereotactic radiotherapy, described below
("Stereotactic radiotherapy. Gamma-knife."); here the radiobiological efficacy is no longer precisely described by the LQ model (in addition to apoptosis, immediate cell death by necrosis also applies), the so-called LQL model or other high-dose modifications described in §5.2, section "Deviations and modifications of the LQ model" are introduced.
    In addition to the regular doses, which are part of the used fractionation regime, certain additional or additives doses are sometimes applied, so-called boost (additional increase) - "saturation of". Reasons for boost application may be radiobiological (improvement of local tumor control with respect to individual tumor conditions and surrounding tissues) or technical (when for medical reasons or for irradiator failure the whole irradiation series does not go according to schedule -
with an additional dose it is necessary to appropriately compensate the time dependences of the reactions of the tumor and healthy tissues). To determine the doses in the boost, it is appropriate to use radiobiological modeling based on the LQ model; however, it is often based on empirical experience. A special technique is the so-called concomitant boost CB (concomitant - concurrent), in the case of hyperfractionated radiotherapy, 2 doses are administered sequentially daily: one for the total target volume of PTV, the other only for the inner part of GTV, containing the macroscopic volume of the tumor itself (see "Planning radiotherapy" below). This increases the dose in the that sub-volume of PTV, in which there is a higher risk of recurrence. Using the IMRT technique of modulating beam intensity using an MLC collimator (see "Modulation of irradiation beams" below), the maximum tumor dose (GTV), somewhat lower dose in the CTV region with potential for micro-seed, and minimized dose in surrounding critical tissues can be achieved relatively accurately. The two batches of concomitant boost can then be combined and applied simultaneously within one daily fraction. This advanced technique, called SIB (Simultaneous Integrated Boost) gradually replaces sequential concomitant boost. 
Prediction of radiotherapeutic effect - probability cure of tumor TCP and damagie of normal tissue NTCP 
Successful radiotherapy - the cure of a cancer - involves killing as many clonogenic cells in the tumor site as possible, that would be able to regenerate the tumor (recurrence) if they survived. To quantification this basic goal of radiotherapy and to predict the success of treatment, the quantity TCP (Tumor Cure/Control Probability ) was introduced - the probability of cure the tumor. The radiobiological effect and the behavior of cell populations have a stochastic (probabilistic) character, according to Poisson statistics. The probability that after irradiation it will not occur to redistribute clonogenic cells and tumor growth, is given by the exponential relation TCP = e
- N , where N is the number of surviving clonogenic cells in the lesion after irradiation *). Substituting for N from the LQ model, we get a double exponential relation for the dependence of TCP on the dose D .:
             TCP(D)  =  e-No.e-(a.D + b.D2) =  e-No.e-a . BED  ,
where N
o is the original number of clonogenic cells in the tumor before irradiation (No is on the order of 1010 -1012 cells). The graph of this function is S-shaped - for low doses up to about 20-30Gy, TCP is close to zero (almost no therapeutic effect), then increases approximately linearly, and for doses above 80-100Gy, the "saturation state" of TCP ® 1 (100% effect) - red TCP curve in Fig.3.6.0b; however, specific values are different for individual types of tumor tissue, they depend on radiosensitivity (on values a, b).
*) This remarkably simple relationship results from the more complex laws of mathematical statistics. Overall cell survival is a stochastic random variable of a binomial character, governed by the Poisson statistical distribution. If we hawe a mean the number of clonogenic cells N, then the probability P(n) is a random phenomenon that survives n cells is given by P(n) = (N n /n!). e -N . For n = 0 we obtain P(0) = (N0 /0!).e -N = e- N . It is the decrease in the number of remaining clonogenic cells to n = 0 that can be considered as a guarantee of definitive liquidation of the tumor; its probability is therefore TCP = e- N.
    To express adverse biological effects on normal healthy tissue during radiotherapy, an analogous quantity of NTCP (Normal Tissue Complication Probability - the probability of complications from damage to normal tissue NT, especially critical organs (see below). NTCP comes from the same Poisson stochastic laws of killing and survival of cells by radiation as TCP, but applied to the surrounding healthy tissue, irradiated with a certain portion of the tumor dose D . When substituting from the LQ model, in addition to the relevant parameters N
o, a, b, l, T2r for the given NT tissue, it is necessary to include the volume factor, with regard to serial or parallel tissue type :
           NTCP(D,V)  =  e-No.V-k.e-(a.D + b.D2) =  e-No.V-k.e-a . BEDNT  ,
where V [% /100] is the relative proportion of the volume of irradiated normal tissue, the parameter k describes the volume effect
(k = 0-1; parallel organs with a large volume effect have a higher value of k than serial organs with a small volume effect), BEDNT is the biologically effective dose for normal NT tissue. In the parameter No (which no longer has the immediate significance of the initial number of cells, as is the case with TCP) the requirement for a minimum number of surviving cells (or their percentage; these are stem clonogenic cells) in NT is implicitly included so that their deficit does not lead to exceeding the functional reserve competent critical organ and its necessary function has been maintained. The dose dependence of NTCP(D) has a similar sigmoidal shape as TCP(D), but is shifted horizontally to higher doses of D (normal tissue receives only a small part of the tumor dose D, respectively only a certain part of the NT volume is irradiated) - green NTCP curve in Fig.3.6.0b. In radiobiological modeling in radiotherapy, the effort is to maximize TCP (®1) and minimize NTCP (®0), although it is sometimes quite difficult ...
Functional modeling of TCP and NTCP
Instead of primary above-derived 2-exponential functions, the sigmoidal course of dose-response curves TCP(D) and NTCP(D) is modeled in practice using "secondary" so-called probit-function (from eng. probability ) of Gaussian shape
               F(D,D50,m,V) = (1/Ö2p)-An [(D/D50.V-k) - 1]/me-x2/2dx ,
where D50 is the dose value with a 50% probability of the studied effect (tumor elimination or complication in normal tissue) and m is the slope parameter of the TCP(D) or NTCP(D) curve in the linear section (maximum value of derivation according to dose D). D50 and m play the role of form-factors of the sigmoidal shape of the curves. V is the relative proportion of the volume of irradiated tissue, k describes the volume effect. For TCP the parameter k = 0 (so V 0 = 1 - volume does not apply), in the case of NTCP the value of the parameter k> 0 models the normalization of dose D per volume, with respect to the serial or parallel type of critical tissue (volume effect mentioned above, "sixth R"). In case of uneven irradiation of critical NT tissue (as in practice) a suitable correction is made for NTCP determination - instead of dose D the so-called equivalent uniform dose of EUD is introduced, recalculating the irradiation effect of individual sub-volumes V i (total number N, ie. i=1
SN Vi = Vtot = 1) with partial doses d i for uniform irradiation of the whole critical organ: EUD = (i = 1SN Vi .d ik )1/k (weighted sum of partial volume contributions of the given NT body with volume factor k). Modeling of the radiotherapeutic effect using TCP and NTCP was first introduced by J.T.Lyman, G.J.Kutcher and C.Burman in the 1980s ( LKB model ).
 Methodological note: Unified concept of TCP and NTCP
From the point of view of radiotherapy, TCP and NTCP are independent quantities, related to different tissues with different parameters of radiosensitivity and with conflicting requirements of radiobiological effect. In the professional literature, therefore, they are mostly introduced as separate models. However, the basic ideological aspects have TCP and NTCP in common: the same radiobiological mechanism of cell survival and killing (quantified in the LQ model) and the probability character with the Poisson distribution. Here we have tried to outline a unified theory, deriving both TCP and NTCP from the same initial "baseline principles" as the Poisson statistical distribution and LQ model of cell survival dose dependence. The basic approach is then exactly the same, only with NTCP the percentage of irradiated NT volume and the volume factor of a given NT tissue are introduced. The TCP and NTCP models are thus unified into one concept. An open problem in this approach, however, remains the expression of the minimum number (percentage) of clonogenic cells in NT that must survive to maintain long-term functionality of critical tissues and organs - with respect to functional reserve of relevant NT, their parallel or serial character .

Fig.3.6.0. Some radiobiological aspects in radiotherapy - graphical representation (model examples).
Basic LQ model of biological dependence on dose. b) Graphical representation of the dependence of TCP, NTCP and UTCP on dose D. c) Quantification of the success of radiotherapy with conventional irradiation (top) and conformal radiotherapy IGRT (bottom).

TCP and NTCP are sometimes combined to assess overall radiotherapy optimization. Introducing the so-called probability of uncomplicated treatment UTCP (Uncomplicated Tumor Cure Probability) :
                   UTCP = TCP. (1 - NTCP) ;
In the case of irradiation of several critical tissues NT
1 , NT2 , ..., NTn , adjacent to the target volume, the probabilities of NTCPi complications in the i-th critical organ appear in the product of coefficients (1-NTCPi ) :
                UTCP = TCP . i=1P n (1 - NTCPi ) .  
The curve of UTCP(D) dependence on the radiation dose has a bell shape (blue curve in Fig.3.6.0 in the middle) - it is zero at small (insufficient) doses, it increases to the maximum at the optimal dose D
opt and then decreases again to zero for too high doses that damage healthy critical tissues. The dose of Dopt , corresponding to the maximum of UTPC, expresses the optimal dose in terms of the relationship between the achieved probability of TCP tumor eradication and the acceptable level of probability of NTCP damage to healthy NT tissue. The position and height of this UTCP maximum significantly depends on the precision of the irradiation methodology: using conformal IGRT radiotherapy or stereotactic radiotherapy (described below), the UTCP maximum shifts to higher doses, due to reduced irradiation volume of critical tissues ® reduction of NTCP, better tolerance of surrounding tissues (Fig.3.6.0c).
    To optimize radiotherapy, all of the above derived dose functions are obtained by conversion from DVH dose-volume histograms in 3D radiotherapy planning (see "
Radiotherapy Planning" below), using special computer software; the areas under the DVH curves represent the relative "partial volumes" of irradiated NT tissues. Also, the above-mentioned therapeutic ratio of TR is sometimes expressed by the ratio of these values: TRTCP = TCP / NTCP. The evaluation of all these parameters is not always unambiguous, opinions on the "weight" of tumor eradication and side effects on healthy tissue sometimes differ. However, due to the danger of cancer, it should be remembered that (with the exception of fatal damage to important critical organs) the most serious complication is tumor recurrence !
Time factor - the influence of cell repair and repopulation

In principle, the influence of time factors - the influence of cell repair and repopulation during irradiation on the resulting biological effect can be "built into" the derived biophysical dose quantities BED and TCP used in radiotherapy. By replacing the simplified equation of the LQ model -ln (N / N
o ) = a .D + b .D2 by general equation with Lea-Cathesid cell repair factor and additive repopulation term, we get for BED and TCP more general expressions :
             BED = D.[1 + {2.[1-(1-e-l.t/l.t)]/l.t}.d/(a/b)] - T.ln2/(a.T2r) , ® TCP = e-No.e-a . BED .
Similarly for NTCP. However, too many parameters - and thus degrees of freedom - complicate radiobiological modeling and often make it ambiguous. In practice, we usually suffice with simple laws, supplemented by empirical experience (cf. the above-mentioned section "Fractionation in practice") ...
  Cell repair causes that if we apply two radiation fractions of dose d , the radiobiological effect is lower than with one dose irradiation 2d. E.g. [N/N
o](2x2Gy) < [N/No](4Gy).
  During irradiation, cell repair takes place in healthy tissue and in the tumor, but at different rates. For lower doses, more tumor cells are usually killed than normal (late-reacting) tissue cells. At high single doses, the curves of normal and tumor tissue may "cross", and the effect on healthy tissue may be greater. Fractionated therapy has a higher effect on tumor tissue and a lower effect on healthy tissue, while this desired difference increasing with the number of fractions.
  In fractionated radiotherapy, cell repopulation of tumor cells between fractions may also occur , as the total treatment time is relatively long. For fractionated therapy tumor tissue with coefficients
a, b and doubling half of T2r repopulation, with the total dose D divided into fractions d during the total duration of treatment T , for the biologically effective dose of BED the LQ model (without repair, but with repopulation) is based on the relation: BED = D.[1+d/(a/b)] - T.ln2/(a.T2r). Thus, the repopulation time factor reduces the biological effect of irradiation, especially for fast-growing tumor cells (here, the reduction in BED is estimated to be 0.5 Gy/day), which needs to be compensated by increasing the total dose.
  The model of continuous exponential repopulation with a fixed doubling half-life T
2r during the whole irradiation is only approximate, in fact the above-mentioned effect redistribution of cells (fourth "R") with complex temporal dynamics, including progressive repopulation, occurs in the irradiated tissue. Its exact inclusion is difficult, the additive term would have the shape of an integral, model eg (ln2/T2r).0nT(1-k.e-v.t) dt, with a time-varying rate coefficient v(t) of the repopulation rate change. For practice, however, a simpler approach is sufficient. The time Tacc from the start of therapy, when the accelerated repopulation of clonogenic stem cells begins (sometimes referred to as the accelerated repopulation delay time Tdelay), is important. Reduction of the radiobiological effect after this time, it can be simply written using the relation E s -ln(N/No) = D.(a+b.d) - K.(T-Tacc), where the empirical factor K (according to the previous theoretical approach K = ln2 /(a.T2r)) expresses the degree of reduction of the biological effect - at the same time it indicates the daily dose [Gy] needed to destroy the newly formed cells on this day, ie the dose needed to compensate for the repopulation of clonogenic cells. Using BED (º E/a) this relation can be written in the form BED = D.[a+d/(a/b)] - K.(T-Tacc)/a. In head and neck tumors, the time Tacc is about 28 days, while with prolongation of the total duration of T therapy losing the biological effectiveness of the radiation by about 0.8 Gy per day. This relationship is sometimes used for shorter times T <Tacc , where the equivalent dose, on the contrary, increases relatively.
    In any case, the approximate nature of this approach must be borne in mind. The K parameter is probably not constant during therapy, it is small at the beginning and increases with time - progressive repopulation. Time T
acc the onset of accelerated repopulation depends on the applied radiation dose and its timing (fractionation). And both of these variables are, of course, different for different types of irradiated cell population.
Combination chemo-radiotherapy
Irradiation is sometimes appropriately combined with chemotherapy
( concomitant - accompanying, complementary therapy, mentioned above), which can either have an additive effect (independent cytotoxic cell killing) or, in certain circumstances, increase the effect of radiotherapy (so-called potentiation) - to increase the radiation sensitivity of irradiated tissues either by inhibiting DNA repair mechanisms, or by appropriately influencing the cell cycle to a phase more sensitive to radiation (eg G2). Optionally, neoadjuvant chemotherapy is applied prior to irradiation to reduce the extent of the tumor, so that the tumor foci can be irradiated more selectively while sufficiently protecting the surrounding critical tissues and organs.
  The biological effects of combined chemoradiotherapy can in principle also be quantified using a suitably modified LQ model. For this purpose, three basic mechanisms of the biological effect of simultaneous chemo-radiotherapy need to be analyzed separately :
¨ The additive effect is given by the sum of the radiobiological effect of radiotherapy and the cytotoxic effect of chemotherapy, which are independent of each other. The effect of concomitant chemotherapy on the resulting biological effect can then be expressed in the LQ model by adding a new independent term Ech which expresses the efficacy of cytotoxic cell killing in chemotherapy; the value of Ech represents the logarithm of the surviving fraction of cells after separate chemotherapy. The overall additive effect of chemoradiotherapy is then: E º -ln (N/No) = a .D + b .D2 + Ech .
Radiosensitization (potentiation) effect of chemotherapy, which enhances the biological effect of ionizing radiation on cells. It is included in the LQ model by adding a new coefficient - sensitization factor s , which expresses the increase in radiation sensitivity of cells due to chemotherapy. This coefficient effectively multiplies the radiation dose in a- and b -member, so the final effect of radiochemotherapy with sensitization is: E º -ln (N/No) = a .D .s + b .D2.s2 .
Inhibition of repopulation by cytostatics and in particular by some targeted biologic therapies, such as monoclonal antibodies against growth factors, reducing the repopulation of tumor cells during fractionated radiotherapy, leading to the elimination of a larger fraction of tumor cells by radiation. In the LQ model with time factor -ln (N/No) = a.D + {2.[(1-e-l.T).(1-1/l.T)]/l.T}.b.D2 - ln2.T/ T2r this slowing of cell repopulation can be captured by modifying the additive member RP º ln2.T/T2r - by multiplying the doubling half-life of repopulation T2r by a coefficient greater than 1. At the same irradiation with radiation dose D during time T thereby reduces surviving fraction of cells N/No the tumor. Inhibition of repopulation due to appropriate chemotherapy prolongs both the half-life of repopulation T2r and the time Tacc from the beginning of therapy, when accelerated repopulation of clonogenic tumor cells occurs. Thus, in connection with the analysis of the previous paragraph, there is less loss of biological efficacy with increasing the duration of fractionated therapy, thus effectively increasing the effect of therapy.
    In chemosensitive tumors, chemotherapy can also cause a reduction in the volume, a "shrinkage", of the tumor (similar effect to neoadjuvant chemotherapy ). Such a reduced tumor is then easier to treat with radiotherapy, both by reducing the number of tumor cells itself, and possibly by improved oxygenation and thus increased radiosensitivity due to the oxygen effect. However, these effects can no longer be objectively included in the LQ model.

Strategic goal and methods of radiotherapy
The "strategic goal" of radiotherapy is therefore the selective elimination of the tumor foci with the least possible damage to the surrounding healthy tissues - so that their functionality is not endangered. Irradiation of the surrounding tissues can never be completely avoided, but the so-called tolerance dose in critical tissues and organs must be observed *). It is necessary to introduce a sufficiently high dose of radiation into the target area, lethal for tumor cells - tumorous cancero-lethal dose (approx. 50-150 Gy) in such a way, that the surrounding healthy tissues are not enormously damaged. The task of radiotherapy in clinical practice is to find the optimal compromise between these two conflicting requirements. In this chapter, from a physical point of view, we will briefly describe how the basic strategic goal of radiotherapy is achieved by various methods of irradiation.
*) By critical organ or tissue we mean such a structure in the organism, the radiation damage of which would have serious health consequences, or in the case of vital organs even death. Therefore, during radiotherapy, a certain so-called tolerance dose must not be exceeded in these critical organs in order to prevent their irreversible damage. The various radiation sensitivities of tissues and organs and their division into serial and parallel are discussed in §5.2 , passage "Local tissue and organ radiation effects".
    Theoretically, all tumors could be locally curable by radiotherapy, but the obstacle is the limited tolerance of healthy tissues and organs, the irradiation of which cannot be avoided - radiotoxicity to healthy tissues.
Side effects of radiotherapy - radiotoxicity, secondary malignancy
As mentioned above in several places, a common limiting factor in achieving a sufficiently high cancerolethal dose in tumor foci is radiotoxicity (also called radiation morbidity) to surrounding healthy tissues and organs, which are always partially irradiated with tumor tissue. In terms of time, there are basically three types of unwanted consequences of radiotherapy :
- Early acute radiotoxicity
manifests itself within a few days to weeks from the start of irradiation. It is caused by the loss of stem cells in rapidly proliferating tissues with a short cell cycle, where continuous and rapid production of daughter effector cells is required. The loss of these stem cells upon irradiation soon leads to depletion of effector cells, which results in impaired function of the affected tissue or organ. The complex time dynamics of division of rapidly proliferating cells during prolonged irradiation (asymmetric division, accelerated repopulation, ...) was discussed above in the LQ model, passage "Redistribution". Mucous membranes (esophagus or intestines), bone marrow, epidermis are affected by early radiotoxicity. Clinical manifestations of early radiotoxicity, if not very severe, are usually temporary and due to the gradual replacement of missing cells (by dividing stem cells) they disappear within a few weeks. Early radiation toxicity can later turn into late toxicity (often occurring with increased early toxicity) - referred to as consequential late radiotoxicity.
Very early radiotoxic effects
When irradiated with high doses with radiotherapy, very early symptoms such as fatigue, nausea, and dry mouth may appear relatively quickly, within a few hours. These symptoms of very early radiotoxicity are not caused by the mechanisms of radiation killing of cells - a larger number of cells are although damaged, but this damage will manifest itself later, only during the mitosis of these cells. Very early radiotoxicity is caused by irritation of regulatory centers by direct action of released ions, radicals and other products of radiolysis.

- Late radiotoxicity
manifests itself with a delay of many weeks to months (sometimes several years) after irradiation. It is the result of tissue damage with a slow recovery of effector cells, where therefore the proliferation rate of stem cells is low. Damage to these stem cells (which do not divide continuously, but only with the loss of effector cells and the need for their replenishment) occurs already during irradiation, but it manifests itself only when there is a need for their division, which is unsuccessful (mitotic death). This occurs with a longer time interval and manifests itself in the depletion of daughter effector cells in the affected tissue. Late radiotoxicity is manifested in connective tissues (subcutaneous, submucosal), bone, muscle (myocardium), eye lens, kidney or lung. The clinical consequences of late radiotoxicity are usually permanent.
    Radiation-induced heart damage - cardiotoxicity, manifested by atherosclerosis of the heart arteries, fibrosis, or damage to the heart valves, occurs during radiotherapy in the chest area, e.g. during irradiation of breast tumors. Pulmonary toxicity at higher doses may cause fibrosis of lung tissue. Deterioration of renal function - nephrotoxicity - can occur after radiotherapy in the area of the abdomen and urogenital tumors.
  By optimizing the irradiation regime, especially by precisely directing the irradiation beams and by appropriate fractionation of radiation doses, undesired damage to healthy tissues can be largely minimized or reduced to an acceptable level.
- Very late stochastic effects
The very late effects of radiotherapy manifest themselves several years to decades after irradiation. Due to the long time interval since the irradiation, it is very difficult or impossible to determine the causal connection between the irradiation and the adverse effect. Their clinical manifestation and course cannot be distinguished from spontaneously arising diseases in persons who were not treated with radiation.
In addition to the desirable deterministic effects on tumor tissue, on which radiotherapy is based, as well as adverse radiotoxic effects on healthy tissue, secondary post -radiation malignancies may occur over time due to the stochastic effects of that portion of the radiation, that was absorbed outside the primary target tumor
(including scattered radiation) and irradiated other tissues and organs. It is difficult to distinguish this radiation-induced carcinogenesis from spontaneous cases (after all, already the first emergence of cancer indicates an increased predisposition of the patient to these diseases). The risk or incidence of these secondary radiation-induced malignancies is estimated at about 3% / 60Gy. It is sometimes debated whether exposure from frequent verification imaging in IGRT-guided radiotherapy methods may also contribute to secondary malignancies (see below)..?..
 Basic methods of radiation therapy
According to the basic method or path by which radiation is "transported" to the desired target site (affected tissue or organ), the radiotherapy methodically divided into three modalities :

Teletherapy - irradiation "at the distance" using beam of radiation from an external irradiator, referred to also as EBRT - external beam radiotherapy (this "distance" of irradiation is only relative, the irradiator is usually a max. of 80 cm away from the body).
Note: However, I do not use the name teletherapy much here, because of its misleading resemblance to the charlatan methods of "teletherapy" = "distance treatment"...
Brachytherapy - irradiation "at close range" - insertion of closed radionuclide emitters, radiophores, into the tumor tissue or in its immediate vicinity.
Radioisotope therapy - application of open radionuclides in a suitable chemical form directly to the organism (most often intravenously, sometimes orally). Radionuclides then enter the target tumor tissues via a metabolic pathway and destroy the tumor cell "from the inside" - Biologically Targeted Radioisotope Therapy BTRT.
    These three basic radiotherapeutic methods will be discussed below, mainly from a physical point of view.
Note:  In conventional radiotherapy, it is usually required that the target lesion (volume) be irradiated as homogeneously as possible with a sufficiently high dose of radiation. However, this requirement does not apply to brachytherapy and stereotactic radiotherapy, where the dose distribution is strongly inhomogeneous, with a steep drop from the target site (will be described below).

External irradiation with gamma, X and electron radiation (teleradiotherapy)
The most common method of radiotherapy is irradiation with a collimated beam of penetrating radiation from an external irradiator. X- radiation of higher energies (approx. 100 keV) were used especially in the past *), now they are used, for example, for irradiation of skin lesions. Radiotherapy with high-energy heavy particles will be discussed below in a separate section "
Hadron Radiotherapy". Here we will deal mainly with irradiation with high-energy gamma radiation.
*) The main disadvantage of X-ray therapy was the inability to achieve a sufficient dose of radiation in the deepertumor lesion without imposing enormous radiation on healthy shallow tissues, especially skin. This shortcoming was largely addressed by the use of penetrating photon radiation with significantly higher energies of several MeVs, where the skin and superficial tissues ceased to be a limiting factor, as the maximum dose shifted in depth. First it was cobalt and cesium irradiators (they were introduced in the 50's), later hard radiation generated by betatrons and now by linear accelerators.
Terminological note:
Kilovoltage - Megavoltage
Sources of radiation are divided into two categories in radiological "jargon" (in connection with historical development) according to the radiation energies produced :
Kilovoltage - providing energy of quanta up to 1000 keV, is mostly produced by X-ray tubes;
Megavoltage - providing energy of quanta over 1 MeV, is produced by accelerators (or radionuclides with correspondingly high energy of radiation
g; the exception is cesium 137Cs with energy Eg 662 keV, which is still classified as "megavoltage").
These slang terms are not physically appropriate and can be misleading. We do not use them in our treatise.
    The intensity of the radiation decreases with the square of the distance from the source. At greater distances, we obtain a more favorable ratio between the amount of radiation that falls on the surface of the body and the amount of radiation penetrating deep. During deep irradiation, it is therefore irradiated from a distance of min. 60 cm from the surface of the body.
Gamma - irradiators
Radiotherapy is currently performed mainly by penetrating gamma radiation, produced either by radioisotope irradiators
137Cs (g 662 keV) and 60Co (g 1173 + 1322 keV) (for radionuclides see §1.4 "Radionuclides"), or arising as braking radiation (bremsstrahlung) *) under the impact of high-energy electrons accelerated in a betatron or linear accelerator (for energies Ee approx. 4-40 MeV) to a suitable brake target made of heavy metal (Fig.3.6.1b) - here the radiant energies are in units up to tens of MeV (see §1.5 "Elementary particles", part "Charged particle accelerators"). The target is mostly made of tungsten, on a plate about 2-3 cm thick, it is thinned to about 3 mm at the point of impact of the electron beam - it works in "transmission" mode, braking radiation coming from the target in the direction of the original electron beam is used. The robust design of the target ensures heat dissipation (most of the kinetic energy of the electrons is converted into heat).
*) Terminological note: radiation X or g ?
In §1.2 "Radioactivity", part "
Radioactivity gamma", we introduced a terminological agreement that photon radiation emitted from atomic nuclei is called radiation g (even in the case, when it has a low energy of a few keV), while the radiation generated by the jumps of electrons in the atomic shell and the braking radiation of electrons is called X-rays (even in the case, when it has a higher energy of tens and hundreds of keV). However, for braking radiation generated in accelerators at energies of several MeVs, such terminology (a kind of "megavolt X-radiation") would be misleading, although sometimes used. This radiation lies deep in the g- region of the classification of the electromagnetic spectrum, it has even significantly higher energy than the usual g- radiation from radionuclides. Therefore, this high-energy braking radiation we will be called a gamma radiation.
    At present, the betatrons have been completely pushed out high-frequency linear electron accelerators - LINAC, which are smaller, more flexible and provide high radiation intensity - Fig.3.6.1b,c
(physical principles and construction of accelerators are described in more detail in §1.5 "Elementary particles and accelerators", part "Charged particle accelerators"). For larger accelerators for energies around 20 MeV, for geometric reasons LINAC is placed perpendicular to the gantry and the electron beam is electromagnetically deflected in the transverse direction of irradiation (Fig.3.6.1b). The deflection electromagnet also serves as an energy filter of electrons, which deflects only the electrons of the required energy (momentum) in the desired direction - the other electrons end up on the walls of the tube (Fig. 3.6.1b shows, for simplicity, the bending of the electron beam at an angle of 90°, but usually 270° is used, allowing a better focusing of the electron beam). Smaller accelerators up to 6 MeV can also have a compact "rectilinear" design without deflection of the electron beam (is seen below in Fig.3.6.4a, c).
Irradiators for radiotherapy, equipped with linear electron accelerators with energy mostly 6 or 18MeV, are currently supplied mainly by two main manufacturers: American Varian (Palo Alto, California, originally a manufacturer of klystrons and accelerators) and Swedish Elekta (Stockholm, also produces Lexell's gamma-knife ). Newly, there are the manufactures of Tomotherapy and Accuray (which merged), producing special tomotherapeutic and stereotactic robotic systems, equipped with compact linear accelerators mostly 6 MeV - see below "Tomotherapy; Stereotactic radiotherapy".

Left: Continuous spectrum of braking gamma radiation generated by the impact of electrons of energy Ee 6MeV and 18MeV from a linear accelerator on a target. In the middle: Line spectrum of g- radiation of radionuclides 137Cs and 60Co. Right: Percentage depth dose dependence for different photon energies (in the water phantom).

Braking photon radiation, caused by the impact of electrons of energy Ee on a target, has a continuous energy spectrum with a predominance of lower energies (up to 1/3 Ee), which decreases continuously from its flat peak (around 1/8 Ee) and then ends at maximum energy just below the value of electron energy Ee. It should be noted that the mean energy of this radiation is significantly lower than the original energy of the electron beam from the accelerator. E.g. when using electrons accelerated to energy Ee= 6MeV, the maximum in the spectrum of braking radiation is around 500 keV, the mean energy is about 1.5 MeV, while the proportion of photons with a maximum energy approaching 6MeV is already very small (units %); the usual statement that "we irradiate with 6MeV energy" is therefore somewhat misleading. It is worth noting that the depth distribution of the dose in water (and tissue) is almost identical for g -radiation of 60Co and braking radiation from the accelerator Ee = 4MeV; and is only slightly different for an accelerator with Ee = 6MeV. These lower energy accelerators are therefore basically interchangeable with a cobalt radiator (the cobalt source has slightly larger dimensions of the radiator itself and therefore a larger half-shadow in the beam).
Effective cross section for the production of braking radiation is generally given by the rather complicated Bethe-Heitler formula (derived from quantum radiation theory, corrected by the Sauter and Elwert factors of the Coulomb shielding of the electron shell). For a not very wide range of energies of incident electrones Ee and proton numbers Z of the target material (medium to heavy materials), the overall efficiency of braking radiation production h can be approximated by a simplified formula :
h   =   Ee [kev] . Z . 10-6 [photons/electron]   .
Only a relatively small part (only approx. 1%) the original kinetic energy of the incident particle changes to braking radiation during braking in the matter. Most of the energy is eventually transferred to the kinetic energy of the atoms of matter by multiple Coulomb scattering - it is converted into heat.
    It is logical that the efficiency of braking radiation production is higher for high Z - large electric Coulomb forces act around such nuclei, causing abrupt changes in the velocity vector of the incident electrons that get close to the nucleus. The efficiency of braking radiation [number of photons /electron] increases with energy E
e incident electrons. However, the overall energy efficiency - the ratio of the total energy of the emitted photons to the energy of the incident electrons - is lower for higher energies (due to the higher percentage of low-energy photons). And the heat losses in the target are higher.
Contamination of the photon beam by electrons and neutrons
The resulting beam of high-energy braking radiation
g is always somewhat contaminated by electrons released during the interaction of photons with the material of the target, homogenization filter, screens and collimators. There is Compton scattering, photoeffect, electron-positron pair formation. In all these processes, fast electrons are emitted from the material. At higher energies, above about 10MeV, photonuclear reactions also occur and release neutrons (see below). Secondary particles, which contaminate the photon beam, reduce the depth effect and increase the radiation dose even outside the direction of the primary radiation, they contribute to the increase of the radiation dose outside the target volume.
Homogeneity of the irradiation beam

The beam of braking
g- radiation, diverging conically from the interaction point in the target, has a significantly higher intensity in the central direction than in the peripheral parts (radiation diagram of braking radiation has a "lobe" shape in the direction of high energy electrons - §1.6, part "Interaction of charged particles - directly ionizing radiation"), left in Fig. :

Left: The directional radiation diagram of the braking radiation from the target leads to an inhomogeneous intensity distribution across the beam. Middle: Achieve a homogeneous distribution with a suitably shaped homogenization filter. Right: Examples of homogenization filters for different energies.

To achieve a homogeneous distribution of radiation troughout the entire required beam width, a homogenization filter is inserted into the radiation path *) - a rotationally symmetrical metal absorber disk-shaped in the middle of a strongly thickened (into a cone), which by the higher absorption in the central part, equalized radiation intensity across the beam cross section - flattening filter. The higher the energy of accelerated electrons Ee, the more strongly the braking radiation is collimated in the axial direction and the thicker the central part of the homogenization filter is needed. For lower energies, homogenization filters are usually made of aluminum, for higher energies, heavier metals (iron, tungsten) and suitable alloys are also used. Homogenization filters are usually replaceable and each of them must be precisely shaped depending on the energy used and also on the required field size and the distance in which a homogeneous radiation intensity is to be achieved. In addition to the homogenization of the irradiation beam, another positive factor is the filtering out of electrons with which the high-energy photon beam is often contaminated. Also, the spectrum of braking radiation across the beam is not exactly the same - in the central part, the proportion of harder radiation is slightly higher than in the peripheral parts; however, the homogenization filter further emphasizes this difference.
*) A separate homogenization filter is not used for cybernetic gamma knifes (CyberKnife), where it is irradiated with a relatively narrow central part of the braking beam (see Fig.3.6.4c below), the homogeneity of which can be ensured by suitable shaping of the target material. After all, with stereotactic radiotherapy it is not necessary to achieve homogeneous irradiation of the target tissue. Recently, homogenization filters have generally been abandoned not only for stereotactic, but also for conventional irradiators, as new sophisticated computer planning systems can accurately plan the dose distribution for any course of the dose profile of the photon beam (this profile is measured dosimetrically and inserted into the planning system).
However, it is necessary to take into account a slightly higher contamination of the photon beam with an electrons.
Collimation and monitoring of the irradiation beam
The braking beam is further collimated by a system of fixed forming orifices
(primary orifice just behind the target and secondary orifice behind the homogenisation filter). A part of the irradiation head is also a radiation monitoring system, which by means of ionization chambers indicates dose rates in the irradiation field *). Finally, the beam is collimated ("modulated") to the desired final shape by a system of movable apertures - the most perfect collimation system is the so-called MLC collimator (see below, Fig.3.6.3). A light localization system is installed in the head for visual aiming and adjustment of the irradiated field - the light from the filament lamp or LED is guided by optical projection through the collimation system of the radiator so that the agreement of the visible light field and the radiation field is achieved. In modern isocentric irradiators, a detector is built into the gantry opposite the irradiator (flat-panel imaging, its principle is described in §3.2, section "Electronic X-ray imaging"), allowing to display the beam after passing through the patient - to create so-called portal images - X-ray images of patient structures using high energy ("megavolt") photon radiation. This portal display system is abbreviated EPID (Electronic Portal Image Device). This system also allows for "in vivo" verification dosimetry to be performed operatively on the beam passed through the patient.
*) Radiation doses and monitoring units
In addition to the standard units of radiation dose Gray [Gy], in connection with phantom measurement (monitoring) of radiation beams, so-called monitoring units
MU are often used in practical radiotherapy (Monitor Unit - 1 MU = --> 0.01 Gy (1 "centigray"). The monitoring chamber indicates the dose of 100 MU, when a radiation dose of 1 Gy in a water phantom (at a field size 10x10cm) is delivered in the isocenter of irradiator.
Depth "build-up" effect of hard photon radiation
  In photon radiation, the radiation dose is caused by secondary electrons, arising from photoeffect, Compton scattering and at higher energies also the formation of electron-positron pairs (see §1.6 "Ionizing radiation", section "
Interaction of gamma and X-rays"). When high-energy radiation g is used, Compton scattering predominates and the secondary electrons have a predominantly primary beam direction as well as high energy; they cause more and more ionization. Thus, as high-energy radiation passes through the tissue, the number of secondary electrons initially increases and ionization increases. At a certain depth, the equilibrium of charged particles is established and then the ionization begins to decrease, as the photon beam is gradually attenuated by absorption in the tissue.
    For hard photon radiation, therefore, the maximum radiation dose is no longer on the surface (as is the case with soft radiation), but shifts somewhat in depth (the so-called build-up effect - onset of dose with depth), depending on the radiation energy. The depths of the maximum dose in the tissue for different energies of photon radiation are approximately: 1MeV ... 4mm; 5MeV ... 1cm; 10MeV ... 2.5cm; 25MeV ... 5cm. Although this effect alone can not be used for depth selective irradiation from one direction, but it has a significant effect on the skin-sparing effect and the superficial tissues when isocentric radiotherapy.
    At greater depths, the equilibrium state of ionization already occurs and the dose D (dose rate) decreases with the depth d according to the standard exponential dependence D
~ e - m .d with a linear absorption coefficient m(r, Eg) given by the tissue density r and the radiation energy Eg - the higher the energy, the slower the decrease (it is derived in §1.6 "Ionizing radiation", section "Radiation absorption in substances", Fig.1.6.5).
    High-energy hard radiation
g *) therefore has the advantage of less absorption (even in the bones) and thus a better "geometric" ability to get the required dose of radiation selectively to a deeper target location, with relatively lower absorption and radiation exposure of other tissues, especially skin.
*) From above, however, the optimal energy of photon radiation is limited to about 20MeV, because at higher energies frequent photonuclear reactions occur (see §1.6, section "Interaction of gamma and X-rays"), due to which the beam is contaminated with neutrons. These neutrons scatters in the tissue and cause radiation exposure even outside the direction of the original beam, ie outside the target volume. In general, it should be noted that at energies higher than 10MeV occur g- activation of the irradiator materials, which are exposed to the radiation beam - target, homogenization filters, collimators, bed and other components are weakly radioactive even after the end of the exposure! Short-term radionuclides (15O, 11C, 13N, in trace amounts further 24Na, 29P, 34Cl, 35S, 38Ca, 38,42,43K) are also formed in the irradiated volume of the patient, but in such a small amount that their contribution to the radiation dose is completely negligible (<10-5 %).
Electron irradiation

For irradiation of surface and shallow lesions, the primary electron beam from the accelerator (energy of the MeV unit, approx. 4-12MeV) is sometimes used. In the arrangement according to Fig. 3.6.1b, the electrons from the accelerator do not fall on the target, which is pushed aside
(of course, the homogenization filter is also displaced), but they are led through a collimating tube directly into the patient's body. The primary narrow electron beam (approx. 3 mm in diameter) is guided, instead of on a target, on a scattering foil to scatter the electrons over the entire irradiation field. In some systems, the electron beam is swept to the desired width by electromagnetic deflection coils (similar to an electron beam in a classic screen). Electron irradiation is suitable for surface lesions at a small depth below the surface (up to about 5 cm), which can be irradiated only from one direct direction (field) and where at a depth below the irradiated deposit there are tissues or organs that should not be irradiated with a higher dose of radiation. Compared to gamma radiation, the electron beam has a sharp decrease in dose towards the depth of the tissue: the maximum range of electrons in the tissue in centimeters is approximately 1/2 of the energy used in MeV, the mean range is about 1/3 of this energy.
    For high electron energies, an analogous mechanism of the "depth build-up effect" is manifested
(cf. Fig. 3.6.5a below), which was mentioned above for hard photon radiation: to a certain depth, the absorbed dose increases somewhat, then - after the equilibrium of charged particles has been established - it begins to decrease rapidly as the electron beam is inhibited and attenuated by interaction with tissue. If high-energy electrons need to irradiate the surface layers of the skin, the build-up effect is undesirable and tissue-equivalent boluses (also mentioned below) are used to suppress it, which leads to an increase in the surface dose and a reduction in the depth dose.

" Make the invisible visible " - display of radiation beams
Ionizing radiation used in radiotherapy is invisible to our eyes, we can register them only using special methods of detection and spectrometry
(Chapter 2 "Detection and spectrometry of ionizing radiation"). For better clarity, however, it would be appropriate to somehow directly "make visible" this radiation, respectively its interaction with the substance. One of the methods was described in §2.2 - 3-D gel dosimeters; however, it is a relatively complicated and demanding method, it is used very rarely... There are two other ways to directly and easily "make visible" the passage of ionizing radiation through a substance: Cherenkov radiation in an optically transparent medium (also in water) and scintillation radiation (preferably in a liquid scintillator). We used these methods experimentally for electron and photon beams at our workplace and for proton beams at PTC.
.Cherenkov radiation
During the passage of fast electrons - whether primary or secondary - through the medium, a weak visible so-called Cherenkov radiation is emitted
(§1.6, passage "Cherenkov radiation"). The following figure shows an example of the "visibility" of an irradiating electron and photon beam in water using this Cherenkov radiation :

Cherenkov radiation generated in an aqueous phantom during irradiation with electron and photon radiation beams.
Left: A cylindrical phantom (diameter 20 cm and height 18 cm) filled with water was irradiated with a wide (magnetically scattered) beam of 9MeV energy electrons from a linear accelerator.
During passing through the upper part of the phantom, fast electrons generated Cherenkov radiation to a depth of about 4.5 cm, when the energy of the electrons fell below the threshold level of 260 keV.
When irradiating the same phantom with a beam of photon radiation (max. energy 6MeV, beam with a diameter of 4cm) secondary electrons along the g beam form Cherenkov radiation - with a deep decrease in intensity as the primary photon beam weakens as it passes through water (just below the surface, a slight increase in intensity is initially seen - build-up effect to a depth of about 1 cm, discussed below) "Secondary radiation generated by X and g interactions").
In the upper and lower part, optical reflections of light from the cover and from the bottom of the phantom are visible. Due to the relatively weaker intensity of the images, the images contain a higher amount of disturbing noise.
Acknowledgments: Irradiation of the water phantom on TrueBeam and CyberKnife devices was performed in cooperation with colleagues: Ing.L.Knybel, Ing.L.Molenda and Ing.B.Otáhal.

Display of radiation beams in a liquid scintillator
Another option for displaying the passage of beams of ionizing radiation through a substance in a suitable phantom is the use of a liquid scintillator
(liquid scintillators and their use for internal measurement of beta-radioactive samples are discussed in §2.6, section "Detection of beta radiation by liquid scintillators").
  At our department, we used a liquid scintillator
(in a very unconventional way) to map and visualize the radiation beams - electron, photon, proton - used in radiotherapy. We filled a glass measuring cylinder with a diameter of 6 cm and a height of 44 cm with 1 liter of liquid scintillator (we used dioxane scintillator with a density of 0.95 g/ml) and placed it under the irradiation head of the respective irradiator - electron Varian, photon CyberKnife, proton IBA. From the side, we observed and photographed the scintillation radiation generated in the scintillator along the passage of the irradiation beam :

Scintillation radiation generated in a cylinder (diameter 6 cm and height 44 cm) filled with a liquid scintillator during irradiation with electron, photon and proton radiation beams.
a), b): A cylindrical phantom filled with a liquid scintillator was irradiated with a wide electron beam of 6 MeV and 18 MeV from a linear accelerator.
c), d):
When irradiating the same phantom with a photon beam - max. energy 6MeV, beam 1.5 cm and 3.5 cm in diameter - secondary electrons along the g beam generate scintillation radiation - with a deep decrease in intensity, as the primary photon beam weakens when passing through a liquid.
Acknowledgments: Irradiation of the scintillation phantom on TrueBeam and CyberKnife devices was performed in cooperation with colleagues: Ing.L.Knybel, Ing.L.Molenda and Ing.B.Otáhal.
e), f), g): When irradiated with narrow ("pencil beam") proton beams of energy 100, 170 and 226 MeV, the protons penetrate to different depths depending on the energy, with a prounced Bragg maximum.
Acknowledgments: Irradiation with proton beams from the IBA cyclotron was performed in cooperation with colleagues: Ing.P.Máca,Ing.M.Andrlík,Mgr.L.Zámeèník, Ph.D. , Ing.M.Navrátil, Ph.D. (and consultations with colleagues Ing.V.Vondráèek and MUDr. J.Kubeš, Ph.D.) from the PTC proton center.

Analysis and discussion of the images :
When irradiated with a wide electron beam of energy 6MeV (a), a bright blue glowing trace to a depth of about 26 mm can be seen in the scintillator, where the electrons are already braked.
Electrons of energy 18MeV (b) continue to a depth of about 78mm, while scintillation radiate. However, the interaction of these high-energy electrons with the scintillator atoms also produces intense photon braking radiation, which is penetrating and continues to depth.
The angular distribution of the emitted photons of braking radiation depends on the energy of the primary charged particles. At low energies, the braking radiation is emitted practically isotropically in all directions from the point of interaction. As the energy of the electrons exciting the braking radiation increases, the mean angle of the emitted quanta becomes smaller and smaller - at high energies of the incident charged particles, the braking radiation is preferentially emitted in a narrow cone "forward" in the direction of impact of the primary particles. The directional radiation pattern of high-energy braking radiation has the shape of a sharp "lobe" in the direction of the primary beam.
Thus, in addition to a clear scintillation pattern of electrons in the upper part, we also observe a weaker narrow beam of braking radiation, continuing to the bottom of the phantom.
¨ When irradiated with a photon beam from CyberKnife (narrow and wide - c, d ) with a continuous spectrum with a maximum energy of 6MeV, we see a significant scintillation trace from secondary electrons across the entire cylinder - the photon beam penetrates deep to the bottom (and would reach even deeper), with a slight depth drop in intensity as the photons are gradually absorbed as they pass through the liquid.
It is interesting to compare these images with the above representation of the same irradiation beams using Cherenkov radiation in an aqueous phantom.
When irradiated with proton beams (e, f, g), we see a significant scintillation trail, which amplifies and ends with a bright Bragg maximum; the radiation no longer continues to a greater depth. The depth of the Bragg maximum increases with proton energy.
The blue "halo" around the proton beam is caused by secondary electrons ejected from the matter as the protons pass. With lower proton energy, this "halo" is wider - electrons are less collimated in the direction of the primary beam; this is especially evident at the end of the trajectory around the Bragg peak, where the protons are already considerably slowed down.

Irradiation field
From a geometric point of view, radiation for radiotherapy can be divided into one or more areas of certain shapes and intensities and from different directions - the so-called irradiation fields. For surface lesions, one irradiation field of softer photon radiation (or electrons) is usually sufficient; for lesions deposited in depth, a larger number of suitably shaped irradiation fields are used
(converging or opposite fields, "crossfire" of four fields and many other combinations). Various absorption filters, screens, wedges (Fig. 3.6.1b ') or special collimators (see below) are often used to form the shape of the radiation beam (and thus also the isodose curves), suitable filters are used to influence the energy spectrum of the radiation. To compensate for the irregular shape of the surface, or to adjust the dose on the surface and in depth, the so-called compensatory bolus (Greek bolos = lump, wad, piece ) is sometimes used - a suitably shaped tissue-equivalent material of a certain thickness, which is applied to a suitable place on skin, or inserted off-surface into the radiation beam.
Note: A more detailed description of irradiation techniques of this kind lies outside the scope of our physical treatise. From a physical point of view, they are not very interesting and, moreover, they are gradually being pushed out more and more by more advanced and accurate IMRT and IGRT techniques - see below.
   The most perfect deep irradiation technique is isocentric irradiation with high-energy radiation with suitable shaping and modulation of the irradiation beam - IMRT, with possible imaging navigation IGRT - and stereotactic irradiation with narrow sharply collimated beams of radiation (with radiation navigation); the most complicated is hadron radiotherapy. These methods are successively described in detail below.

Isocentric radiotherapy
The main strategic goal of radiotherapy - effective selective irradiation of the tumor loci with the least possible damage to surrounding tissues - is achieved by irradiating the tumor site with a collimated beam from multiple directions *) so that the intersection of beams, ie focus or isocenter, where doses add up, it was localized to the tumor site - Fig.3.6.1a. The surrounding healthy tissues then receive a reasonably lower dose, divided into a larger region. Simply put, healthy tissue (its individual sites) is irradiated only once, while the tumor is irradiated each time.
*) For this purpose, the radiator is mounted on a special round stand, the so-called gantry (gantry - portal, continuous supporting structure ), enabling controlled rotation of the radiation source around the patient by means of electric motors.

Fig.3.6.1. Movement isocentric radiotherapy with a collimated beam of gamma radiation.
a) Basic idea scheme of radiotherapy with a rotating irradiator. b) Arrangement of the irradiator with a linear accelerator. c) Example of a modern IGRT irradiator.

Collimated fields and radiation beams for radiotherapy
From a general physical point of view, the properties of ionizing radiation were described in §1.6 "Ionizing radiation"
(fields and beams were then mentioned in the section "Fields and beam, radiation intensity"). The primary radiation from the accelerator (electron radiation or proton radiation for hadron therapy) usually emits in a precisely defined direction, in a narrow beam (which is then further modified, filtered and shaped). However, the radiation g (and possibly X), arising in radionuclides (cesium or cobalt), or excited as secondary braking radiation after the impact of the primary electron beam from the accelerator on the target (Fig.3.6.1b), is emitted in practically all directions (high-energy braking radiation has only a higher intensity in the central direction, which is corrected by a homogenization filter). In order to create an irradiation beam for targeted (tele)radiotherapy, it is necessary to shield the vast majority of this diffuse radiation and transmit only the radiation in the required direction - to perform the collimation of the radiation. The simplest collimation is roughly shown in Fig.3.6.1a - tube-shaped collimator. More complex collimation systems are used to accurately shape the irradiation beam, the most advanced of which are the electronically formable MLC collimators described below ("Irradiation beam modulation").
    In the middle part of the (homogenized) beam of radiation defined by the collimator, there is an approximately homogeneous intensity distribution. At the edges, the intensity does not suddenly decrease to zero, as would follow from an idealized geometric configuration, but decreases continuously. Absolutely sharp collimation cannot be achieved in practice for two reasons :
- Geometric blurring due to the non-zero size of the primary source (manifests itself especially in radioisotope sources) .
- In the case of penetrating high-energy radiation g, partial translucency occurs trough the edges of the collimator .
    In the marginal parts of the collimated beam, a kind of "half shadow" is created. Next to this geometric penumbra, the scattering of the radiation beam in the tissue also applies
(this scattering is significant especially in the electron beam). At higher energies, the photon beam in the tissue is sharper, there is less scattering penumbra. These two effects - geometric and scattering - create in the dose distribution in the tissue the resulting dose half-shadow in the marginal parts of the radiation beam, which must be taken into account when planning radiotherapy, can significantly affect the isodose curves.
Technical note: For the sake of simplicity, the irradiation beams are shown in Figures 3.6.1 and 3.6.2d by lines (arrows) of constant width. In reality, however, the radiation beams have a diverging geometry - with distance from the radiation source are expanding .

Radiotherapy planning
The combination of physical and biological factors in most cases enables sufficiently effective and selective irradiation of the pathological lesion. In clinical radiotherapy, the patient's own irradiation is always preceded by a very important and demanding process of radiotherapy planning, the result of which is the so-called irradiation plan, containing all the specific details of the irradiation process for the patient. A properly designed radiation plan is a basic prerequisite for successful radiotherapy.
    The main basis for creating an irradiation plan are detailed diagnostic images of the irradiated area. At present, it is mainly the tomographic X-ray images (
CT), or on nuclear magnetic resonance (MRI) and scintigraphy imaging, in particular positron emission tomography (PET). These images serve both for the precise localization of the tumor site together with the determination of its size and shape, as well as a detailed anatomical-density map of the distribution of tissue densities and the location of organs.
Radiotherapy simulator
In exact radiotherapy planning, a so-called simulator is used - a device that mimics the entire irradiation process and allows its optimization. The classic simulator is a diagnostic X-ray device with an image intensifier, the X-ray tube of which is mounted on a rotating isocentric arm and is equipped with a system of adjustable apertures, enabling the imitation of a beam of radiation as it will then be used on its own therapeutic irradiator. The simulator enables localization of the target volume and topometry of tumor deposits, aiming of the beam and modeling of field geometry and irradiation parameters, drawing of orientation and reference points and markers on the patient's body.
  Instead of the classic simulator, the so-called virtual simulator - X-ray imaging device CT is now often used for advanced irradiation technologies (IMRT, IGRT), equipped with a aiming system and special software for dose planning. The planning software first converts the Hounsfield units of the CT image to the electron density of the individual tissues. Furthermore, the target volumes and critical organs are marked on the pictures. Then the images are overlaid with the characteristics of the radiation beams (energy, dose distribution - isodose curves). The marked structures are then displayed in the transformed BEV (
Beam's Eye View) mode - from the point of view of the radiation beam. In conventional planning these overlapping images are then sought to find the most favorable irradiation conditions for delivering the desired dose to the target volume. The number of irradiation fields and their shape, dose rate, angles and other parameters for optimizing the irradiation plan are set. The so-called inverse planning will be mentioned below in connection with the IMRT and IGRT techniques.
    Images from CT examinations are thus directly included in the planning of therapy - 3D-planning, which is followed by the so-called 3D conformal radiotherapy (3D CRT), or even more advanced therapy with modulated IMRT-IGRT beams. The transfer of data from the CT, via the planning computer to the computer controlling the irradiator, provides the possibility of shaping the irradiation fields based on accurate spatial knowledge of the internal anatomy around the target tissue of the patient; the radiation beams are thus adapted to the target volume and protection of the surrounding critical tissues.

    From these data and the required radiation dose in the target tissue (this dose depends on the tumor type - its radiosensitivity), as well as the maximum tolerance dose in the surrounding critical organs, the intensity, energy and geometric parameters of the radiation beam are calculated, including precise radiation positions and angles. Batch fractionation is further determined . The whole process of planning and subsequent radiotherapy is now largely automated using computer software that works in several basic stages (Fig.3.6.2) :

Fig.3.6.2. Some basic stages of computer radiotherapy planning.
a) Diagnostic X-ray (CT) image of the irradiated area. b) Drawing of areas of interest of the target volume and critical tissues, selection of the irradiation procedure, number and shapes of beams (fields) and radiation intensities. c) Optimization of the irradiation plan using dose-volume histograms of DVH. d) Control of the function and movements of the irradiator by the resulting irradiation prescription.

l Analysis of diagnostic data, choice of treatment strategy - curative or palliative therapy, combination with surgery and chemotherapy, localization of the target tumor volume.
Processing of initial X-ray images from CT (Fig.3.6.2a). Tissue density (expressed on CT in Hounstfield units) is converted to electron density. This takes into account the inhomogeneity of the tissues (different electron densities of soft tissues, water, air, bones) during the passage and interaction of the irradiation beam. The electron density of the substance is directly proportional to the linear energy transfer LET (the amount of energy loss per unit path) and thus the local ionization in the tissue and the absorption of radiation - radiation dose distribution. To refine the irradiation plan, it is also appropriate to take into account gammagraphic images of PET (eg by computer fusion of CT+PET images), which map the viability of tumor tissue - it was discussed above in the section "Diagnosis of cancer".
l Drawing the regions of interest ROI into the picture - especially the target volume of the tumor lesion, then the risk critical tissues and organs (Fig.3.6.2b). These areas of interest are drawn in individual transverse sections, the program combines them (using interpolation) into a three-dimensional volume. Perpendicular frontal and sagittal sections can also be used when drawing ROI.
Target volumes irradiation
Target irradiation volume (Target Volume TV) means the corresponding region of tissue localization and size (volume), to which must target canceroletal desired dose. For successful curative radiotherapy, it is necessary to apply a lethal dose of radiation not only to the actual volume of the macroscopically detected tumor deposit, but also to some neighboring areas - the so-called safety margins, reducing the risk of insufficient irradiation of structures that could be affected by cancer and subsequently cause recurrence of the disease. In connection with this, we have three or four consecutive target volumes in radiotherapy : 

Basic target volume GTV
(Gross Tumor Volume) represents the intrinsic volume of a macroscopically detected tumor lesion, imaged using an appropriate image (mostly CT). Some other neighboring areas - lem - margin - are then added to this initial, basic or gross volume .
¨ Clinical target volume of CTV
To safely ensure local control of the irradiated lesion, it is necessary to apply a lethal dose of radiation not only to the actual volume of the macroscopically detected GTV tumor site, but also to those neighboring areas where, according clinical experience, microscopic spread of tumor cells could already be hidden. Therefore, we increase the irradiated target volume of GTV by a clinical safety margin - the so-called clinical target volume of CTV is created (Clinical Target Volume). It is analogous to the safety margin in the surgical removal of visible tumors.
¨ Internal target volume of ITV
Due to internal physiological changes in the position of the target volume within the organism (eg due to respiration
*), variable filling of the bladder and intestines, peristalsis, swallowing, heart pulsation), further expansion to the internal target volume ITV (Internal Target Volume) is needed, which includes the entire pathway of internal movement of the target tissue.
*) Tumor tracking
Advanced methods of stereotactic radiotherapy use the so-called tracking of tumor - monitoring the movement of the tumor due to respiration, its inclusion and correction, which allows to reduce ITV and thus minimize the exposure of the surrounding healthy tissue, possible escalate the dose to the tumor site itself. These "tracking" methods of respiratory gating or respiratory synchronization are discussed below in the "
Stereotactic Radiotherapy" section.
Resulting planned target volume of PTV
Due to the expected small deviations in the reproducibility of the position of the patient and the irradiator during fractional irradiation, it is sometimes necessary to further expand the target volume by the so-called position edge. This creates the resulting Planning Target Volume (PTV), which is drawn in the irradiation plan. PTV includes the CTV and an widening rim for ITV organ and tissue movement, as well as for expected irregularities in irradiation settings. The resulting PTV is thus a unification (outer envelope) of all partial target volumes: PTV = GTV
È CTV È ITV È [position edge].
l Entering the required radiation dose in the target tissue and the maximum allowable dose in critical tissues. This is based mainly on empirical experience, which results in the coefficients a, b of radiosensitivity of a given type of tumor tissue and tolerance doses for healthy critical tissues (discussed above in the section "Physical and radiobiological factors"), passage "Prediction of therapeutic effect").
l Selecting the basic irradiation method - the number and geometric configuration of the radiation fields, energy and intensity of the beam - its eventual modulation, the number of fractions.
l Calculation of the distribution of local dose in the thus mapped tissue - the so-called isodos curves are construct (illustratively is seen on Fig.3.6.2d). In practice, the radiation beam is never homogeneous, as is the absorption of radiation in tissue, so that the spatial distribution of radiation intensity and the radiation dose is usually a complex form (highest dose is usually in the central part of the beam, decreases towards the edges). The spatial distribution of the radiation dose is often mapped using the so-called isodose curves - imaginary lines representing the connection of points with the same dose. Usually, isodose curves are plotted for certain percentages from the site with the maximum dose, e.g., isodoses of 80%, 50%, 20%, and the like (reminiscent of contours on the map).
l Optimization of the irradiation plan. For this purpose, are often constructed volume histograms of dose, so called Dose Volume Histograms DVH (Fig.3.2.6c). These histograms provide a plot of the 3-D dose distribution using a illustrative 2-D curve display. Each marked area of interest has its own DVH curve. On the horizontal axis is the dose (in Gy or in % of the max. dose), on the vertical axis is the volume (in % of the volume of the marked structure). DVHs show the dose exposure of the target volume (PTV) and individual identified critical organs (NT).
DVH dose-volume histograms 
Dose-volume histograms indicate, how much of the volume of the target or critical tissue will receives a particular dose. From the point of view of the basic radiotherapy strategy, it is desirable that the largest possible volume of the target (tumor) PTV tissue receives the highest possible percentage of the required dose (ideally 100%). At the same time, that the smallest possible volume of critical healthy tissue NT received the lowest possible part of the dose. This dosage exposure target volume and critical organs indicated schematically shown in a dose-volume histogram DVH. The optimization of radiotherapy here consists in optimizing the areas under the DVH curves - the largest possible area under the target PTV volume curve and the smallest possible area under the histograms of NT critical organs (the proportions of NT/PTV areas under the DVH curves represent the relative "partial volumes" of irradiated NT tissues).
  In a more detailed analysis, we can further improve radiotherapy optimization by converting dose and dose distribution values from standard DVH to quantities BED, or to TCP + NTCP and derived UTCP (these quantities were defined and discussed above in "Physical and radiobiological factors in radiotherapy", passage "Prediction of radiotherapeutic effect - TCP, NTCP").
Data transfer to the irradiator coordinate system .
Creation of an irradiation prescription , according to which the functions and movements of the irradiator are controlled during the actual irradiation (symbolically in Fig.3.6.2d) - mainly the angular positions of the irradiator, exposure times, radiation beam geometry using MLC collimator modulation.
    Current planning computer systems are also able to implement so-called inverse planning (see below), in which the planning system calculates the parameters and movements of the irradiator so as to achieve the primarily required dose distribution in the target volume (lesion) and do not exceed tolerance doses in surrounding tissues.
Dosimetry and verification in radiotherapy
In order to ensure the necessary accuracy of radiotherapy, verification methods are also needed, ensuring the delivery of correct therapeutic radiation doses to target lesions and tolerance doses to at-risk critical tissues and organs, taking into account changes in their position, anatomical shape and size at different fractions of irradiation.
    For dosimetric verification, ionization chambers or diode detectors are most often used. These detection elements can be individual (with mechanical shift), in a linear or two-dimensional matrix arrangement, or in a cylindrical structure for measuring isocentric irradiation. They are placed in the irradiation beam either directly ("in the air") to map the intensity of the beam, or they are inserted into suitable water or plastic phantoms, modeling typical anatomical structures for irradiation. Dose monitoring is also performed "in vivo", directly when irradiating a patient to whose body dosimeters are applied to the appropriate sites. An elegant method of verification and at the same time in vivo dosimetry is the use of images from the portal flat-panel of the irradiator
(flat-panel principle is described in §3.2, passage "Electronic display X-rays"), their calibration and quantification - the so-called portal dosimetry EPID (Electronic Portal Dosimetry image).
    Rarely, 3-D gel dosimetry systems are also used, allowing to determine the spatial distribution of the dose in the irradiated volume (the gel is filled into a phantom modeling the irradiated structure). This method is quite demanding both in the stage of phantom creation and in terms of evaluation (a more detailed description is in §2.1, section "
3-D gel dosimeters"), it is used only for research and development work.
    For the quantitative assessment of the agreement of irradiation plans and dose distributions, the so-called gamma-analysis is sometines used, the output parameter of which is
g -index (0 <g£ 1); the closer the gamma value is 1, the better the agreement.
    In modern irradiation systems, the irradiation and verification technology is integrated into one irradiation device. High demands are placed on the accuracy and reproducibility of the geometric position of the patient relative to the radiation beam - so that the irradiated target volume is precisely set in the coordinate system of the irradiator. Various markers drawn or placed on the surface of the body and laser sights are used for this. Opposite the irradiator, detectors (display flat-panels) are built into the gantry, enabling the creation of so-called portal images during irradiation; even more perfect is the X-ray system con-beam CT. These IGRT -guided radiotherapy methods are described in the following section "
Modulation of irradiation beams". In classical stereotactic radiotherapy of intracranial lesions, a stereotactic frame is used for precise targeting of the target lesion, in cybernetic irradiators special stereoscopic X-ray imaging and aiming systems - see the section "Stereotactic radiotherapy. Gamma-knife" below.
    The proper interplay of this complex "technological chain" requires the cooperation of an experienced radiological physicist.

Uncertainties in radiotherapy
Every physical or technical, diagnostic and therapeutic method is burdened with greater or lesser inaccuracies, errors, uncertainties. Naturally, even during the complex chain of radiotherapy, we encounter a number of uncertainties. The "input" primary uncertainties are :
¨ Location and extent of the disease , the uncertainty of which is given by the accuracy and sensitivity of diagnostic methods, event. it can be affected by artefacts of imaging methods, inaccuracies in the patient's settings, his general movements and the movements of organs inside the body.
Radiobiological factors - radiosensitivity of tumor and healthy tissues (parameters a, b in the LQ model, the rate of cell repair and repopulation) is known only approximately and on a flat-rate basis, and it can vary considerably for individual patients. This leads to uncertainties in the basic prescription of the radiation dose and its fractionation.
During the process of planning and implementation of radiotherapy, this is followed by other uncertainties :

Inaccuracies in plotting ROI - defining target volumes and critical structures in planning images. This process is highly dependent on the experience of the planing radiotherapist.
Uncertainties in irradiation technology - accuracy and stability of energy and intensity of the primary irradiation beam, uncertainties in the monitoring system, accuracy of the collimation system, geometric setting of distances and isocenters, accuracy of transfer of irradiation plan parameters to the irradiator control system.
¨ Inaccuracies and disturbances during the patient's own irradiation - variability in the positioning of patients under the irradiator, fixation and movement of the patient, movement of tissues and organs inside the patient during irradiation.
    New knowledge in the field of radiobiology, together with advances in diagnostic methods and technical improvements in irradiation technologies, especially their integration with imaging modalities, make it possible to gradually reduce or eliminate these numerous uncertainties. This increases the radiobiological, dosimetric, geometric and overall accuracy of radiotherapy.

Modulation of irradiation beams
Radiotherapy with modulated beam intensity - IMRT
  In order to perform sufficiently intense and homogeneous irradiation of the tumor and to protect of surrounding tissues, it is necessary to shape the beam to achieve maximum irradiation of the target volume of the given geometry (size and shape) and the dose in the surrounding environment is reasonably lower, so that the surrounding tissues and organs are obscured (shielded) - protected against radiation. For this purpose, suitably shaped filters (masking blocks of various shapes) and apertures or collimators, defining the field size, are inserted into the radiation beams. This was previously done manually for each irradiation field and was very laborious
(the workplace was equipped with a mechanical workshop, where the covering blocks were cast, cut and machined). With technical development, therefore, more universal mechanically movable apertures have been created. By dividing these apertures into independently moving segments, a very flexible multi-lamellar collimator MLC ( Multi Leaf Collimator) was constructed, mounted on the output of the photon beam of braking radiation from the accelerator - Fig.3.6.3.
    The invention of the MLC multi-leaf collimator, which led to the introduction of IMRT modulated beam intensity irradiation, marked a major revolution in radiotherapy: it enabled high-precision specially targeted irradiation of tumor lesions of various shapes and sizes, with maximum protection of surrounding healthy tissues and critical organs. And in a relatively easy and reproducible way, without the laborious making of non-standard tools.
    MLC collimators have a larger number of lamellae (approx. 60-120 sheets) 5-10 cm thick, which can be moved independently by means of small electric motors. This makes it possible to create an opening of any shape for the radiation beam, or several openings dividing the bundle into several parts. The edges of the lamellae are suitably shaped to mimic a diverging irradiation beam to reduce "half-shadow". The entire collimator can be further rotated. The small electric motors driving the slats are computer controlled - the MLC collimator is electronically formable.
    Irradiation is performed from several directions, while during irradiation with the help of electric motors the position of individual lamellas of the collimator changes - the intensity is modulated across the radiation beam and thus the dose is regulated in individual parts of the irradiated volume. The beam of radiation is as if divided into individual rays with different intensities. By combining several fields modulated in this way from different directions, a more optimal dose distribution, selective irradiation of the target tissue with better protection of the surrounding tissues and critical organs (which is obscured by appropriate MLC shaping) is achieved. This makes it possible to irradiate even irregular tumors, with maximum protection of healthy tissues in the vicinity of the tumor. The method is called IMRT (Intensity Modulated Radio Therapy) - radiotherapy with controlled (modulated) beam intensity. This is ensured by the construction of special MLC collimators, which modify - shape, modulate- radiation beam at the output of the radiator (linear accelerator). Dose intensity modulation is achieved by superposition of overlapping radiation fields during rotation of the irradiator with different positions of the MLC lamellae. The edges of the lamellae projectively "copy" the shape of the irradiated target volume, transmit an intense beam into the tumor bed and shield the surrounding tissues and critical organs.

Fig.3.6.3. Electronically adjustable collimators for precision radiotherapy with modulated IMRT beam.
a) The multi-lamellar collimator MLC with the help of motor-shifted shielding lamellae allows to flexibly shape (modulate) the radiation beam from the accelerator for radiotherapy with modulated beam IMRT. b) Micro-MLC (mMLC) - miniaturized MLC as an extension to a standard irradiation head with MLC, for therapy with narrow sharply collimated beams. c) Binary (bipolar) MLC for tomotherapy. d) An iris-collimator with an electronically (motorized) controlled hole size - aperture - can replace a whole set of fixed collimators with circular holes of different diameters in a cybernetic gamma knife.

From the point of view of time control, the modulation of the IMRT irradiation beam can take place in two modes :
l Intermittent mode (step-and-shoot), where the collimator lamellae are in motion only during pauses between irradiations. The MLC collimator forms the desired aperture, through which irradiation is performed. Then the irradiation is stopped, the lamellae are moved to the next position (or the collimator is turn slighly), the angle changes to gantry and another dose of irradiation takes place. It is actually an improved technique for a large number of static fields.
l Continuous mode (dynamic, sliding windows) - the collimator lamellas move smoothly, relocate and modulate the beam into the desired shape during irradiation. The movement of the collimator lamels in synchronization with the continuous rotation of the collimator and the entire irradiator on the gantry, is electronically controlled by the appropriate software. According to the irradiation plan, when the irradiator is rotated on the gantry (gradually by up to 360°), the dose rate changes and thus irradiation with the modulated beam takes place. This process is sometimes referred to as Intensity Modulated Arc Therapy (IMAT) - intensity-modulated angle radiotherapy.
    Another improvement of this system is called AMCBT ( Arc-Modulated Cone Beam Therapy) - angularly modulated therapy with conical beams, or VMAT (Volumetric Modulated Arc Therapy) - volume modulated arc therapy, or RapidArc. It contains an irradiation beam controlled by a modulated MLC collimator and a controlled rotation of the irradiator around the patient. In some systems, the primary intensity of the beam is also continuously modulated by regulating the flow of electrons in a linear accelerator. Controlled moving of the bed with patient, turning of the collimator and angular shift of the irradiator gantry are also possible.
    The VMAT technique therefore allows the dynamic change of some parameters during irradiation: - The gantry can move at a variable speed; - The position of the individual lamellae of the MLC collimator is variable during rotation; - The collimator can rotate. The dose and geometry of the beam is thus continuously modulatable during the rotation of the radiator on the gantry.
  Thanks to another, angular-intensity degree of freedom (a number of finely adjustable irradiation beam angles with individually set beam shape and intensity are available), the selectivity of the radiation dose delivered to the target tissue is further improved and the irradiation time is shortened. By modulating the dose rate during irradiation, healthy tissues and critical organs are better protected, the whole-body dose is reduced.
    In addition to the standard MLC collimator, two modifications are used :
(mMLC) - a miniaturized multi-lamellar collimator for irradiation with narrow sharply collimated beams in so-called stereotactic radiotherapy (see below). It is mostly used as an attachment mounted on a standard irradiation head with MLC for radiotherapy with modulated IMRT beam (Fig.3.6.3b).
Binary or bipolar MLC (Binary MLC) - slot-shaped with a plurality (64) of linearly arranged lamellae that open and close, thereby modulating the irradiation beam in the plane of the transverse section (Fig.3.6.3c). It is used in so-called tomotherapy (see below, Fig.3.6.4a). The slats of binary MLCs are driven electromagnetically or pneumatically, which achieves a very fast response of opening and closing of slats (tenths of a second).
  The method of intensity modulated radiotherapy, analogous to IMRT, has recently been introduced even in proton radiotherapy - the so-called IMPT (Intensity Modulated Proton Radiotherapy), see "Hadron Radiotherapy" below.

Image-guided radiotherapy - IGRT
The high potential for accuracy, flexibility and conformity of IMRT technology (as well as gamma-knifes, see below) can only be used effectively in co-production with a very precise method of verifying the targeting of the irradiation beam to the target volume. Such verification of the target volume can be performed by displaying the area of the irradiated lesion and surrounding structures before each irradiation or fraction, followed by computer comparison with initial planning images, evaluation and transfer of results to the irradiator coordinate system. In this case, it is on-line navigation according to the picture. According to the current images obtained immediately before each individual irradiation, the position of the patient and the targeting of the tumor site can be adjusted as required. This achieves a precise setting that is updated accordingly each day of treatment. Only after the patient's position has been verified in this way, the own irradiation is started.
    Such additions and improvements irradiation technology is referred to as IGRT (Image-Guided Radiation Therapy) - radiotherapy controlled (navigated) by image *), controlling the patient's position - target volume and surrounding structures - during the treatment process using radiological imaging methods. It combines the IMRT irradiation technique with the imaging verification technique (Fig.3.6.1b, c). It allows the display of the target volume and surrounding structures using a display device connected to the irradiator. A simpler method is the above-mentioned portal imaging (lower flat-panel in Fig.3.6.1b, c), which displays bone structures well, but often does not provide sufficient contrast for imaging soft tissues, including the lesion itself.
*) IGRT is sometimes considered in the narrower sense only as a verification method ; however, see the section "Hybrid integration of imaging and irradiation technologies" below.
    For quality imaging, the irradiator can be further equipped with an additional X-ray imaging system (sometimes called In Room CT - CT in the irradiation room, Synergy, or OBI - On-Board Imager System - an imaging system mounted directly on the irradiator), which is used to accurately control the position of the patient and target tissue before irradiation; it can be performed before each irradiation fraction. The OBI imaging system is mounted on the irradiator gantry (linear accelerator) perpendicular to the irradiator's central axis. The X-ray imaging system rotates together with the gantry and is positioned to have the same isocenter as the high-energy beam of the irradiator. Prior to irradiation, an X-ray planar or CT image is performed on the irradiator with a widely collimated cone-beam CT, which shines trough the patient and impinges on the opposite flat-panel imaging (its principle is described in §3.2, passage "Electronic X-ray imaging"). The X-ray tube and the detector rotate around the patient on a common irradiator gantry. The resulting current images are compared with planning reference images- initial CT or planar images from the simulator and, if necessary, appropriate position correction or beam shape modification by collimator MLC; in case of larged differences also changes in the irradiation plan, its re-optimization. This allows the elimination of errors in patient positioning between the individual irradiation factions, or changes in the position and size of the target tissue and the surrounding anatomical conditions during radiotherapy. X-ray display provides up-to-date images of structures and organs before the irradiation procedure and, based on them, the accuracy of radiotherapy is optimized by controlled IMRT. The new systems enable verification not only between individual fractions, but also inter-fraction CT imaging during the rotation of the irradiator with IMRT.
    Very precise localization of the tumor and surrounding structures helps to improve the therapeutic ratio - to irradiate the tumor site with a sufficiently high dose and to a large extent eliminate the harmful radiobiological effect on the surrounding healthy tissues and organs. In principle, ultrasound imaging can also be used for IGRT, but the recognition of structures in these images is more difficult and is practically not used in teletherapy. Ultrasound navigation is used in some methods of brachytherapy, as described below and shown in Fig.3.6.7 on the right. A perspective method of IGRT is navigation using nuclear magnetic resonance MRI - a combination of LINAC+MRI (described below in the passage "Hybrid integration of imaging and radiation technologies"), which provides a better visualization of the structures of especially soft tissues.
    The IGRT verification method is most often used in connection with irradiators with a modulated IMRT beam according to Fig.3.6.1b, c *), so it is IMRT + IGRT. On the IGRT image navigation are further based the high-precision tomotherapy and stereotactic gamma-knife, Leksell, and especially cybernetic radiotherapy methods, described below. IGRT can be supplemented by a system of correction for respiratory movement, the so-called Respiratory Motion Technology or Real-Time Position Management ( RPM ), which allows monitoring of the change in the position of the target volume depending on the patient's respiratory cycle - this is used for selective irradiation of the target volume only in selected part of the respiratory cycle, so-called respiratory gating, or when synchronizing the movements of the irradiator with the brathing cycle (respiratory synchronization). All these procedures are now integrated into the DART (Dynamic Adaptive Radiotherapy) system, which makes it possible to evaluate the results obtained during IGRT and, based on them, to operatively adapt the parameters of the irradiation procedure so that the dose distribution is optimal.
*) Occasionally was used the integration of linear irradiator with a CT imaging in the CT-on-Rails design. A CT scanner, mounted on rails, is installed in the irradiation room together with the irradiator itself. In the opposite end of the room, it can be used for basic CT imaging. On the rails, the CT scanner can then be moved to the radiotherapy position, where the deviations of the target volume and other structures are checked in comparison with the planning images, with subsequent correction of the patient's position. This system has again been used in some new carbon-12 hadron radiotherapy systems (see "Hadron radiotherapy" below), where it is problematic to mount an On-Board imaging system on very robust and complicated gantry.
    Adaptive radiotherapy guided by the image IGRT - "daily" CT or MRI - is carried out by two on-line procedures of reoptimization of the radiation plan :
--> Position adaptation (ATP) in which the plan adaptation is performed according to the new current position of the patient under the irradiator. The position of the isocenter is updated and the original contours (ROI) can be applied to the modified plan.
--> Shape adaptation (ATS) based on new anatomy of target tissue structures and surrounding organs. Reoptimization of the radiation plan is performed by changing the shape and size of the ROI (either automatically or adjusted by the radiation oncologist).
 Biologically guided radiotherapy - BGRT
Anatomical and functional-biological multimodality imaging methods are increasingly included in the irradiation process, together with modeling of molecular-cell radiation response of tumor and healthy tissue (molecular imaging, monitoring of radiotherapeutic effect including early detection of apoptosis, discussed above in section "Diagnosis of cancer"). The sum of these approaches makes it possible to gradually achieve "biologically guided" radiotherapy BGRT (Biologically Guided Radiation Therapy), adapted to individual conditions specific to the patient and tissue - radiotherapy controlled (guided) by molecular imaging.
  This area mainly includes the biologically targeted radioisotope therapy by open emitters (discussed below), where the complex theranostic approach discussed in §4.9, section "Combination of diagnostics and therapy - theranostics" is being developed.

Hybrid integration of imaging and irradiation technologies
The accuracy of the localization of anatomical structures in CT and NMRI imaging, as well as the targeting of the isocentre in the irradiators, is already very high, approx. 1 millimeter. However, the use of this accuracy to actually target the dose to the desired site may be hindered by variability in patient position and organ mobility within, as well as changes in their size and anatomical proportions (see "Conformal radiotherapy" below). It is therefore desirable to continuously on-line monitor the position of the patient and internal anatomical structures using an imaging system directly on the irradiator. This creates images of organs and anatomical structures in the irradiator coordinate system, to which the modulated irradiation beam can respond in feedback by modifying and correcting irradiation conditions to achieve the exact desired dose distribution in the target volume and surrounding tissues. Image-guided radiotherapy (IGRT ) and tomotherapy therefore require the merging of the imaging and irradiation device into a single hybrid system - Fig.3.6.1c and Fig.3.6.4a, c. Before each irradiation, it is then possible to take "daily images" of the target tissue and surroundings, according to which it is possible to perform a possible correction of the position, operative update of the irradiation prescription, or correction of the irradiation plan.
    In IGRT systems, hybrid combination [LINAC + CT] is already standard. Hybrid combinations of irradiator with the nuclear magnetic resonance NMRI imaging system are under development (the NMRI principle has been described above - "Nuclear magnetic resonance"). From this combination is expected to better visualiza the structures (target tumor as well as surrounding tissues and critical organs) - especially soft tissues - on NMRI, which would allow to make a more perfect adjustment and targeting of the radiation dose from the irradiator. It is therefore a two-mode MR-IGRT technology. The system has already been tested [60Co + NMRI] - two or three cobalt irradiators (equipped with MLC collimators) combined with simultaneous magnetic resonance imaging. For the highly desirable combination [LINAC + NMRI], a significant technical problem so far is the mutual negative influence of both modalities - influencing the operation of a linear accelerator by a strong magnetic field of a NMRI superconducting electromagnet and interfering with NMRI display by strong electromagnetic signals generated during accelerator operation.
The radiation dose in the tissue - both desired and unwanted - is caused by fast secondary electrons, generated in the tissue by the interaction of primary photon (X, gamma) radiation. In a strong magnetic field, these fast electrons will have their paths deviated from their original direction - the Lorentz force acts on them in the direction perpendicular to the motion and perpendicular to the transverse magnetic field. This can alter the resulting radiation effect of the irradiation beam. In dense and homogeneous tissues, this effect is relatively small (electrons brake quickly in the tissue on a short path), but at the interfaces between substances with different densities, they can manifest themselves. Significant changes in the dose distribution can occur especially at the air-tissue interfaces - the so-called electron return effect occurs here. Fast electrons that have already left the tissue at the interface are no longer slowed down in the air, they spirally bend in the magnetic field and can return back to the tissue; this increases the dose at the interface.
    Several structural solutions of the LINAC/MRI combination with different shaping of the electromagnets, perpendicular or longitudinal orientation of the magnetic field and radiation beam, different positions or rotation of the patient were experimentally built. These prototypes were mostly quite complicated, suboptimal for clinical practice and did not achieve full online connection of the irradiator with MRI navigation. In the end, the most feasible solution turned out to be the simplest in principle: to place another ring gantry on the outside of the superconducting solenoid, with the help of which the LINAC (with a target) would rotate, which would inwardly irradiate the patient lying inside the MRI tunnel :

A simplified diagram of the hybrid combination of a linear accelerator (LINAC) with magnetic resonance imaging (MRI) - IGRT with nuclear magnetic resonance navigation.

(Note: There is only one LINAC; two are drawn in the picture only to illustrate the different position when the gantry rotates)

With special methods of active magnetic and electromagnetic shielding, mutual interference can be minimized and both modalities can work independently and simultaneously. Other effects (such as changes in dose distribution due to the magnetic field) can be corrected in software and included in the radiation plan. Thus, it was possible to realize a functional hybrid LINAC/MRI combination. Two types of devices of this kind were developed and began to be used clinically :
- MRIdian from the manufacturer ViewRay Technologies (MRI 0.35 T + Linac 6 MeV), completed in 2014 (previously this manufacturer supplied the aforementioned 60Co+MRI, later replaced cobalt with a linear accelerator).
- MR-Linac Unity, which was developed in the years 2012-18 in cooperation between the manufacturers Electa (manufactured Linac 7 MeV) and Philips (MRI 1.5 T).
    Navigation of adaptive radiotherapy by magnetic resonance MRI is undoubtedly a very interesting method from both a physical and a clinical point of view. It is legitimately assumed (at least hypothetically...) that it can significantly contribute to the precision and better success of radiotherapy, especially in situations of complex anatomy of soft tissues, where the irradiated tumor tissue is closely adjacent to at-risk critical tissues and organs. Moreover, these anatomical structures often move relative to each other during the radiotherapy process, changing shape and size. These soft tissues - target and surrounding healthy - tend to be difficult to visualize using conventional cone-beam CT navigation on the irradiator. However, they are shown with excellent contrast on MRI images and can be focused and contoured even without the use of fiducial markers. It is mainly for tumors of the pancreas, prostate, rectum, ... In general, however, MRI navigation of radiotherapy is expensive, laborious and time-consuming. And the spectrum of suitable diagnoses is quite narrow. It is necessary to assess in which cases this time and effort is justified and necessary. So far there is only limited evidence of its actual better efficacy and superiority..?..
    The scintigraphic method of PET positron emission tomography is very suitable for primary tumor diagnosis (see Chapter 4 "Scintigraphy", part "PET cameras"), especially with the use of radiopharmaceutical 18FDG. The metabolic cellular activity of the tissues is displayed. However, this method is also suitable for monitoring the response tumor tissue for radiotherapy, as it displays metabolically active tumor tissue, as opposed to inactivated cells. Among other things, it is able to recognize tumor recurrence from other processes (eg from the consequences of previous tumor treatment). This monitoring of the success of radiotherapy can be performed off-line, but a hybrid combination of [LINAC + PET] radiotherapy irradiator with PET imaging is also possible.
    Another interesting hybrid combination that may be implemented in the future is the combination [hadron 12C-irradiator + PET], where the dose distribution from accelerated carbon core beams is monitored by annular positron emission tomography (PET) camera detectors displaying annihilation photons generated in areas of Bragg maxima from positrons b+ -radioactive 11C - see below "Hadron radiotherapy", passage "Radiotherapy with heavier ions", fig.3.6.6. And in the distant future a possible hybrid combination [antiproton irradiator + PET], where a PET camera mounted on antiproton irradiator gantry could monitor the dose distribution in the tissue by detecting annihilation radiation from positrons arising secondarily from antiproton interactions in the tissue (see "Antiproton radiotherapy" below).

Conforming, adaptive radiotherapy. Inverse planning.
  All these gradually evolving methods lead to better irradiation selectivity - a higher dose in the target tissue and a reduction in the dose to the surrounding healthy tissues. They allow better dose distribution in the target volume - so-called conformal radiotherapy (conform = adapt), also referred to as three-dimensional conformal radiotherapy (3DCRT). In this technique, a three-dimensionally defined target volume is selectively and homogeneously irradiated with the desired high radiation dose, which drops sharply outside the target volume, so that the surrounding healthy tissues are irradiated with a substantially lower dose. The size and shape of the irradiated area is adapted to the irregular volume of the tumor lesion. The dose distribution can be adapted for tumor foci of various shapes, including the situation where the tumor foci are closely adjacent, or partially surrounds critical organs and tissues. IMRT uses a number of irradiation fields at different angles, which adapt to the shape of the lesion and "copy" its contour. The modulation of the irradiation beam makes it possible even to partially cover it certain parts of the target volume, that interfere with a critical organ, on which the tumor may push (or partially surround it). The radiation dose in the target tissue is then compensated by stronger irradiation from other fields. As a result, sufficient and almost homogeneous irradiation of the tumor site can be achieved with significant protection of adjacent critical organs (isodose curves can be concavely "curved" around the critical organ). This result is achieved in IMRT by inhomogeneous transport of partial radiation doses to the lesion, adapted to the irregular shape of the tumor and the anatomical situation in the environment.
    Conformal radiotherapy techniques make it possible to selectively increase (escalate) applied radiation doses in target tissue by reducing the dose in critical organs. The sum of these IMRT + IGRT methods is also sometimes called adaptive radiotherapy (ART) - irradiation is adapted to each patient individually, it changes with specific anatomical conditions, even over time in the same patient *). Operative continuous IGRT includes, in addition to the three dimensions of spatial imaging, even a time factor - sometimes referred to as 4D-radiotherapy.
*) The irradiated patient is not immobile and unchangeable object! There are a number of events that can change the patient's internal anatomical proportions somewhat. Respiratory movements, intestinal peristalsis, changes in tissue volumes due to the dynamics of the disease and due to the therapy itself take place. This can lead to differences in the position of target volumes of up to units of centimeters. This can have a significant effect on the accuracy of selective radiotherapy; without correction for these facts, the target lesion could be partially "missed" during irradiation and healthy tissue could be irradiated instead..!..
Correction for respiratory movements

is especially important when irradiating tumors in the lungs and chest area. There are basically two modifications of the method for eliminating the disturbing effect of respiratory movements on irradiation :
- Respiratory gating , when the irradiation beam is switched off and on so that the irradiation takes place only in the selected defined phase of the respiratory cycle (eg in the expirium period).
- Tumor tracking , where the scanned breathing movements are electronically transmitted via a computer to the irradiator control system, which "shifts" the beam in the rhythm of the breath so that it is still directed to the target lesion - respiratory synchronizitaion.
    Said complex method of radiotherapy planning, where the initial requirement is the distribution of radiation dose and computerized planning system determines the optimal shapes, intensity and irradiation time of each modulated radiation fields, sometimes referred to as "inverse planning" :
Inverse Planning 
The name "inverse" comes from the fact that some stages are "reversed" here, compared to earlier conventional planning procedures (conventional planning is sometimes referred to as "forward"). First, the target volumes and structures of critical tissues and organs on individual CT sections are accurately marked. After entering the required dose into the target tissue and the maximum permissible dose for the surrounding healthy tissues and critical organs, a 3D-model of the dose is created. The planning system then designs the number and shape of the irradiation fields, the dose rates, the times, the angles of the gantry irradiator; this was previously done manually in conventional planning. Each partial radiation field (from a given angle) is virtually decomposed into individual surface elements - pixels, the distribution of which is controlled by the positions of the lamellae of the multi - leaf collimator MLC; this distribution is computer-optimized so that the spatial distribution of the dose corresponds to the required values. An important output part of the computer irradiation plan is therefore the data on the position of the lamellae and the angle of rotation of the MLC collimator. All this data is transferred to the irradiation computer, which according to them electronically controls all movements of the gantry, collimator leafs, accelerator power and other irradiation parameters. The adjective "inverse" is likely to become unnecessary in the future, as no other type of planning will exist ...
  Besides to the high purchase price (and higher operating costs), a certain disadvantage of all these advanced radiotherapy methods is greater time-consuming irradiation process and a somewhat higher whole-body dose of radiation (even outside the directly irradiated area), arising as a result of more frequent diagnostic and monitoring irradiation.
    Paradoxical note:
 Accurate irradiation of a defined tumor site with a rapid decrease in the radiation dose to the environment is certainly very desirable and leads to a better protection of healthy surrounding tissues. On the other hand, it can sometimes paradoxically have a certain disadvantage: in the case of micro-seeding of tumor cells into the vicinity of the defined lesion, recurrence of the disease may occur more easily after the end of radiotherapy than with earlier methods, where the surroundings of the target volume was relatively strongly irradiated. It is therefore necessary to pay increased attention to the definition of a sufficient area of the envelope around the lesion itself and its incorporation into the target volume.

A special modern variant of IGRT-radiotherapy navigated by CT images is the so-called tomoradiotherapy. The prefix "tomo" expresses the fact that the irradiation takes place gradually in a series of narrow transverse sections perpendicular to the longitudinal axis of the patient, defined by a beam from the orbiting accelerator (Fig.3.6.4a). Diagnostic imaging and therapeutic irradiation technology is integrated into one system. CT imaging provides up-to-date images of the target tissue and surrounding structures before each irradiation procedure ("daily CT") and, based on them, optimizes the positions and accuracy of radoiotherapy with controlled modulation of the beam intensity.
    An interesting variant of the tomotherapeutic apparatus was realized, using the same linear accelerator as a source of radiation for imaging and for therapeutic irradiation :
¨ CT imaging is realized as a transmission g -CT, where instead of the X-ray tube there is a linear accelerator with a target, producing in the "low dose" mode (with reduced energy and especially with many times lower beam intensity) photon radiation ("megavolt X-radiation") fan-shaped collimated ("Cone Beam"), which shines trough the patient and is registered in the opposite direction by a set of detectors (arranged in a circular section, most often it is a multipixel single-row xenon ionization chamber, more recently scintillation detectors with ceramic materials) similar to classical CT. As with CT diagnostics, along with the rotation of the accelerator and detectors, the bed is moved with the patient (helical or spiral scanning), followed by the reconstruction of the density images.
¨ The same linear accelerator, after switching to high-dose power mode (without using a homogenization filter), orbits around the patient and irradiates the target tissue localized in the previous diagnostic CT step. This irradiation takes place with a modulated beam using an MLC collimator *): for different angles, the intensity of the g-beam can be greater or less (or radiation completely switched off), or the beam suitably shaped so that the radiation dose avoids critical tissues. As in the previous step, together with the orbiting of the accelerator and detectors, the bed and the patient are moved in a controlled manner - helical or spiral tomotherapy is performed. Modulation of the dose intensity is achieved by superposition during the rotation of the irradiator with different positions of the lamellae of the binary MLC, modulation in the longitudinal direction then by means of superposition during the overlap of the individual sections.
*) Since tomotherapeutic irradiation takes place only in a narrow beam rotating in a plane perpendicular to the translation axis, it is sufficient to modulate the irradiation beam intensively only in one direction (plane). Therefore, a special somewhat simpler multi-lamellar collimator MLC is used here, sometimes called binary (bipolar) MLC, which, however, has a faster opening and closing response of the lamelae (Fig.3.6.3c). It typically consists of 64 slats with a pneumatic drive mechanism.
     During high-performance irradiation, CT detectors must be switched off or removed, as high radiation flux would overwhelm them and could damage them (in further development, the detectors are expected to be switched on even during irradiation and to continuously modulate feedback intensity). This elegant, accurate, and highly integrated system is sometimes referred to as "HI-ART" ("Highly Integrated Adaptive Radiation Therapy").
  The first prototype of a tomotherapeutic irradiator (Corvus system) were constructed in 1993 by M.Carol (Nomos Corp.), T.R.Mackie, P.Reckwerdt et al. (Univ. of Wisconsin). For further development and production of these devices, the company Tomotherapy Inc. was founded in 2002 based in Madison, Wisconsin, USA, which supplies these systems commercially. It later merged with Accuray, a company that produces CyberKnife.
Tomotherapy with 60Co
In principle, a radionuclide emitter 60-cobalt (g 1173 + 1322 keV) can be used as a source of photon radiation for tomotherapy, as a replacement for LINAC (as discussed above in the section "External irradiation with gamma, X and electron radiation - teleradiotherapy", dose distribution for g- radiation 60Co is very similar to LINAC 4 or 6MeV). The classic cobalt irradiator, for this purpose, is equipped with a binary multi-leaf collimator and an opposite imaging detector (flat-panel). It can operate in the same configuration as in Fig.3.6.4a, or using two or three 60Co sources - one for flat-panel imaging ("daily CT"), the other for self-therapy. However, such systems are used only very rarely, because radionuclide sources in teleradiotherapy are generally abandoned (with the exception of Leksell's gamma knife).

Fig.3.6.4. Some special gamma irradiation techniques (top principle, bottom device). a) Tomotherapy. b) Lexell's gamma knife. c) Cybernetic gamma-knife.

Stereotactic radiotherapy - SBRT. Gamma - knife.
Stereotactic Body Radio Therapy (SBRT ) is a very accurate high-dose irradiation of a small target volume, usually a large number of targeted thin beams of intense ionizing radiation, with a sharp decrease in radiation dose outside the target volume
(sometimes referred to as the "zone effect"). Each individual beam is relatively weak and does not cause significant radiobiological effects on its tissue path. However, if these rays are directed to a common focus - target tissues, their summation results in a high effective dose capable of damaging and inactivating tumor cells. Outside the focus, the radiation dose decreases sharply, so that already at a distance of a few millimeters from the focus, the dose already corresponds practically to the dose from one beam. Using the so-called stereotactic targeting, the target volume is precisely spatially defined by transferring the diagnostic image to a 3-dimensional coordinate system (without direct visual inspection). Based on the coordinates that locate given places, it is possible to achieve highly selective irradiation of even a small target deposit with a high dose of radiation, with relatively low damage to surrounding tissues. Due to its high accuracy, the method is sometimes referred to as stereotactic radiosurgery SRS (Sterotactic RadioSurgery) *) - allows a single-time ablation irradiation with a high dose, which eliminates the lesion (tumor or malformation). This method is suitable where classical surgery is difficult or unsolvable (eg fine structures in the brain). This targeted irradiation with a "gamma-knife" can then replace the classic surgical intervention - without surgical burden and surgical complications (bleeding, infection). High accuracy (1-2 mm) enables effective treatment of even small tumors near important centers or in areas with a complex anatomical structure. Irradiation is usually performed once or with a small number of fractions (2-3).
*) After all, this method is used not only for cancer therapy, but also for "radiosurgery" removal of vascular malformations or neuropathological (eg epileptic) foci in the brain - disposable focal intracranial irradiation. Stereotactic radiosurgery is a non-invasive alternative to "bloody" surgery.
Terminological note: The term stereotaxy was created by combining the words: stereo = spatial and taxe= intervention in the right place (Lat. tactio = touch). It is also used for accurate surgical procedures.
    In classical radiotherapy, standard single doses of about 2 Gy are applied in 20-40 fractions, the radiobiological mechanism is cell apoptosis with the aim to reproductive sterilization of clonogenic tumor cells; the resulting effect is described by the LQ model. In stereotactic radiotherapy, a high single dose (in the order of tens of Gy) is applied to a small target lesion either one-time or in a few fractions (1-5 fractions). With a one-time dose of tens of Gy, in addition to apoptosis, immediate cell death in interphase - necrosis is already partially manifested (radiobiological effect is no longer precisely described by the LQ model, its high-dose modifications are sometimes used - LQL model, gLQ model, USC (universal survival curve), KN (Kavahagh-Newman) model, PLQ (Padé Linear Quadratic), see §5.2, passage "Deviations from the LQ model and its modifications". Tumor cells are affected by such a large radiation dose (with high dose rate), that nitrocellular repair will not take place and cell repopulation not in progres, all cells are "killed" - the tumor sterilization effect becomes ablative. Therefore, in addition to the name stereotactic radiotherapy SBRT, the term stereotactic ablative radiotherapy SABR, SABRT (Stereotactic Ablative Body RadioTherapy) is also used; sometimes, with a bit of exaggeration, the association with the English word "saber" is mentioned - it is an effective and elegant weapon against tumors... It is interesting to note, that at the high-dose SABRT is more pronounced the abscopic effect (otherwise rare) of the anti-tumor immune response (§5.2, passage "Bystander-Abscopal effect").
    Stereotactic radiotherapy makes it possible to precisely target a high radiation dose to the tumor focus, while maximally saving healthy tissues. This can achieve high local control - effective destruction of the tumor lesion - even near important critical organs and complex anatomical structures, with a lower risk of side effects and complications
(lower radiotoxicity, less risk of secondary radiation-induced malignancies). Stereotactic irradiators - Lexell's gamma knife and CyberKnife - irradiate with about 10-30 times higher spatial accuracy than conventional linear accelerators.
Note: Similar goals are achieved by somewhat different means - beams of heavy particles - the Hadron radiotherapy, below.
Leksell Gamma-Knife

Now already classic device for high-precision isocentric radiotherapy is Leksell Gamma-Knife LGK
(first prototype developed in 1967 by neurosurgeon L.Leksell and radiologist B.Larsson with coworkers at the Karolinska Institute in Stocholm). Radiotherapy takes place by precisely targeted irradiation of a pathological site in the brain with gamma radiation from a large number of solid radioactive sources 60Co (g 1.173 + 1.332 MeV), whose narrow collimated rays from different directions intersect in a common focus, into which a pathological district of brain tissue is positioned by stereotactic localization. Large total radiation doses from all intersecting rays act in the focus, outside this focus the dose decreases sharply and is already 100 times smaller in the vicinity of a few millimeters from the focus; corresponds to the dose from a single beam. The emitters are arranged on a hemispherical surface and are equipped with collimators, that directs (transmits trough the channels) the beams of radiation g to the center (Fig.3.6.4b above). In the basic type of device there are 201 small encapsulated cobalt sources with activities of about 1GBq, evenly distributed on a hemisphere with a diameter of 400mm, which gives a dose rate of about 3 Gy/min. in the isocenter.
    Definitive precise collimation is performed by secondary collimators in special collimation helmets (there are several types of them in the accessories of the instruments, or their segments can be moved by motor). Prior to the actual radiotherapy, a coordinating stereotactic aiming frame is attached to the patient's head with four screws (Fig.3.6.4b below), enabling on the X-ray or MRI imags to mark the position of the pathological lesion together with contrast markers on the frame, and assign the displayed structures to the three-dimensional coordinate system of the irradiator. More preferred here is magnetic resonance imaging, which provides a more contrast imaging of the soft tissues of brain structures. The tomographic image of the brain, which also shows the marks of the stereotactic frame ( fiducial markers), is transferred to the planning system and serves to precisely set the target volume to the focus of the gamma knife. Rays from some 60Co sources can be discarded as needed (by a shielding "plug" in the helmet), if they pass through critical structures that should not be exposed to radiation (such as the optic nerve, ocular lens, brainstem). Irradiation time depends on the size and type of lesion, it is on the order of tens of minutes. Tumors are irradiated with a single dose of about 20-25 Gy, up to 100 Gy (necrotizing ablation dose) is used in malformation radiosurgery. If the target volume is larger or irregular in shape, the bed is moved with the patient so that the focus moves in the lesion and there is a gradual irradiation of the entire target volume - multi-isocentric irradiation. Brain tumors for LKG therapy should not be larger than 3 cm and their number should not be greater than 5. However, in some workplaces 20 lesions are irradiated, mostly with palliative intent. The therapy is basically one-time, but in case of recurrence or appearance of new metastases, the treatment is repeated, even 3 times.
    In new types of gamma-knife have increased the irradiation space (using 192 cobalt emitters with cylindrical geometry, without collimator helmets - these are replaced by a motor-controlled conical collimator with 8 independently moving segments with 576 holes), which allows irradiation even of other target volumes in the head and neck area (up to vertebrae C1, C2).
The gamma knife is used to treat mainly brain tumors and metastases, meningiomas, auditory nerve neurinoma, ocular uveal melanoma, vascular and neurological malformations, and pituitary adenoma.
    Leksell's g- knife has three disadvantages :
¨ Its construction is basically single-purpose - adapted for the therapy of brain lesions (the innovated model also allows irradiation of lesions in the neck area) .
¨ Radioactive emitters 60Co have a half-life of 5.27 years, they gradually weaken and need to be replaced. This is a very complicated and expensive matter...
¨ An unpleasant and uncomfortable aiming stereotactic frame attached to the head for the patient (this fixation should not be used in young children, whose skull is not yet quite strong).
    Nevertheless, Leksell's gamma-knife is intensively used in larger workplaces specializing in diseases of the central nervous system. It is the most accurate device for tumors in the head area.
Universal and cybernetic gamma-knife
With the technical improvement of the "classic" IGRT isocentric radiotherapy using g- radiation, precise irradiation with narrow beams with millimeter accuracy is also possible here. This gradually achieves the properties of a gamma knife for universal use, for various irradiated localizations, not just the brain *). In addition to precisely working IGRT irradiators with an MLC collimator, resp. mMLC (micro-multileaf collimator - Fig.3.6.3b) - in the classical, VMAT, or tomotherapeutic arrangement, "cybernetic (robotic)" stereotactic irradiators with a sharply collimated beam were also developed.
*) For accurate stereotactic irradiation, however, the brain is the most suitable object, as it is enclosed in the skull, which can be well fixed and thus ensure sufficient accuracy (<1mm) of targeting the beams to the target bearing. In other locations, the problem is the mobility of anatomical structures due to respiratory movements, peristalsis, filling and emptying of cavities, muscle mobility, etc. Some of these movements are corrected with advanced irradiation technologies (eg. respiratory gating in respiratory movements).
Cybernetic gamma-knife , CyberKnife
Devices of this kind are precisely functioning cybernetic image-guided irradiators - a complex computer-controlled system, consisting of several basic components (Fig.3.6.4c) :
Radiation source g - compact linear accelerator (LINAC) of electrons with an energy of about 6MeV, equipped with a target converting energy of electrons for braking g- radiation. A homogenization filter is not used here.
¨ Narrow collimators for setting different diameters of the irradiation beam. Either a set of mechanically interchangeable collimators with different aperture sizes is used, or the collimator can be equipped with a variable iris diaphragm, whose electrically moving segments allow automatic on-line setting of various apertures - diameters of the irradiation beam during irradiation (Fig.3.6.3d). Some types of devices are also equipped with a multi-lamellar collimator MLC.
¨ Cybernetic arm on which the irradiator is mounted: the movements of the irradiator are ensured by a special stand - "cybernetic (robotic) arm" with servomotors controlled by a computer (Fig.3.6.4c), with a large range of possibilities of movement of the irradiator. With these servomotors, the irradiator flexibly moves around the patient with all degrees of freedom - it can to shift, turn angularly, rotate around the bed - and purposefully irradiates the tumor lesion with a large number of narrow beams, at various angles, with appropriate doses of radiation. In some new systems, not only the irradiation head moves "robotically", but also the bed with the patient, which takes over part of the movements ("degrees of freedom") of the irradiator (Fig.3.6.4c below).
¨ Stereotactic X-ray imaging system, equipped with two orthogonally placed X-ray tubes (one of which can be seen in Fig.3.6.4c) and imaging flat-panels (located either on stands under the lounger or recessed under the floor), scans the irradiated area, while stereoscopic X- ray images of significant structures in the patient's body can be used as a stereotactic base (reference system). Thus, a fixed external stereotactic frame (as in classical Lexell's stereotactic radiotherapy) is not necessary, as a "stereotactic frame" serve certain significant structures in the body of the irradiated patient :
Some parts of the skeleton - vertebrae (spinal, lumbar, thoracic, cervical), skull structure;
- Directly irradiated tumor - if the difference in density between the lesion and the background is shown distincly enough on X-images;
For reliable navigation of stereotaxy (especially in the area of soft tissues), special so-called fiducial markers ("reliable" - Lat. Fiduacia = faith, trust, reliance, coverage) are sometimes implanted in the vicinity of the tumor - easily recognizable reference locatization orientation markers in the number of 3-6, implanted around the target tissue. Usually gold grains of about 2x5mm size are used.
Note: Gold as a material for fiducial marks has two advantages: 1. It is an inert metal, well tolerated by tissues. 2. Due to the high density, gold grains appear in high contrast on navigation X-images.

On-line tumor tracking

These locating "reference points" or structures (own anatomical or implanted) are marked on the irradiation plan and the X-ray navigation system of the irradiator then constantly monitors them and accordingly them controls movements of the irradiator or robotic bed with the patient. Before each sub-dose from a certain direction, a stereotactic image is taken, which is compared on a computer with the initial images that were used to create the irradiation plan. If the position of the target tissue deviates (due to movement of the patient or movement of target structures within the body) from the planned position, the computer system calculates the appropriate beam alignment correction and the cybernetic arm is adjusted to the new correct position to emit the next dose. Continuous scanning and comparison of current images with default images allows you to operatively correct the position of the irradiator, so that even when the position is changed (e.g.patient movement), the irradiation beam is still aimed precisely at the target lesion.
    The integration of the irradiator with the X-ray imaging device into one system thus ensures optimal on-line image-controlled angular-dose modulation of the irradiation beams (IGRT-SBRT). The above-described technique of image-guided (navigated) IGRT radiotherapy is brought to complete perfection here!
    The main device of this kind is the CyberKnife (Fig.3.6.4c below).
Correction for respiratory movements
For accurate irradiation of target volumes in the lungs, chest and partly also the abdomen, it is useful to equip the system with a device for synchronization and correction of respiratory movements. It is usually performed using an optical laser system with sensors or mirrors attached to the patient's chest, which electronically monitors breathing movements. There are two basic ways to eliminate the effect of respiratory movements on the geometric accuracy of irradiation :
- Respiratory gating - is a simpler way in which the irradiation beam is switched off and on, so that the irradiation takes place only in the selected defined phase of the respiratory cycle (eg in expiration period).
- Respiratory synchronization - opto-electronically monitored breathing movements are transmitted to a computer, which first creates a "breathing curve". With the help of this curve, the sensed signals of respiratory movements are then electronically transmitted to the servomotors of the irradiator arm, which "sways" in the rhythm of the breath, so that the irradiation beam is still directed to the target foci - respiratory tracking .
Note: A certain litle problem of the whole process of "respiratory tumor tracking" may be the positional relationship between the monitored markers (optical reflectors or metal fiducials) and the target tumor during the whole respiratory cycle. A planning CT scan for SBRT in the chest and abdomen is commonly taken with breath holding and represents the target volume, fiducial markers, and surrounding anatomical structures only when exhaling or inhaling; does not provide information on possible changes in position between the tumor (or its deformation) and markers in other phases of respiration. To solve this problem, it is desirable to take two CT images - during exhalation and during inhalation, on which changes in the distance between individual markers and defined parts of the tumor on both images are then evaluated.
    In general, tumor tracking allows for more accurate targeting radiotherapy using individual and reduced margin in ITV, which can be used to better protect healthy tissue or to escalate the dose in the tumor itself.
Large number of beams, high accuracy and selectivity
Even with cyber irradiators, their integration with the CT imaging device into one system ("In-Room CT") is sometimes (occasionally) used for accurate imaging and targeting of the target lesion and surrounding critical tissues immediately prior to exposure (but this is not necessary if there is a planning CT nearby).
    Flexibility of the irradiator's movements allows you to irradiate the target volume with a large number of thin beams from various directions, in an angular range of almost 360° (except for the direction from below under the lounger). This achieves higher accuracy and selectivity of the radiation dose delivered to the target tissue (with a high dose gradient outside the target volume), with the possibility of respecting shape and anatomical anomalies - well avoiding critical tissues. In general, all these precise stereotactic techniques are suitable for radiotherapy of small tumor foci, up to about 3-5 cm, in areas with a complex anatomical structure. Image-guided irradiation with photon beams from a linear accelerator moving on a cybernetic arm makes it possible to irradiate the target area from many angles and minimize the radiation load of the surrounding tissues.
    It is worth noting that, unlike the other teletherapeutic methods mentioned above, the cybernetic gamma knife is not an isocentric technique: the irradiator does not have a rotating gantry and its beam can be directed at any angle. However, the isocentric mode can be achieved, if necessary, by suitably controlled movements of the radiator by means of the servomotors of the cybernetic arm. It can be said that the cybernetic irradiator can operate in 6D positioning mode: classic movement in the x, y axes and movement in three further rotations.
    For precise irradiation with on-line tumor tracking, we can talk about 4D radiotherapy - 3 spatial dimensions and 1 temporal dimensions. The inclusion of radiobiological processes then represents a new 5th dimension - in a way it is 5D radiotherapy.
    In principle, at proton beam, or a heavier ion beam, can also be used for stereotactic radiotherapy, in addition to using the effect of the Bragg maximum depth dose (see below). However, the great complexity of proton beam targeting in a robust gantry does not yet allow accurate proton stereotaxy by the online-guided image.
The first prototypes of cybernetic irradiators have been developed since the late 1980s, mainly in the laboratories of Stanford University (J.R.Adler et al., inspired by the first type of Leksell gamma knife and an effort to improve it), using an industrial robot Fanuc (Japanese Fanuc developed within the electromechanical company Fujitsu). On this basis, the company Accuray (based in Sunyvale, California) was founded in 1991, which significantly improved this robotic radiotherapy stereotactic system and has been supplying it under the name CyberKnife since 2001 (Fig.3.6.4c below). Another type of stereotactic irradiation system is Novalis (manufactured by Brainlab), which also has continuous X-ray scanning, but uses a special Micro-Multi Leaf collimator (mMLC, mentioned above, Fig.3.6.3b) to collimate the photon beam, which can flexibly shape the irradiation beam using computer control.

Hadron radiotherapy
Hard electromagnetic radiation - gamma or X - is the most common type of radiation used in the treatment of tumorous diseases. A number of precise techniques have been developed to selectively direct this radiation to tumor foci
(discussed above). However, a certain disadvantage here is the not very advantageous depth profile of the radiation dose :
    In conventional irradiation of deeper lesions with photon beams, most energy is transferred to the tissues located on the surface and at shallow depths in the body *), before they hit the tumor itself. With increasing depth of penetration into the tissue there is a slow exponential decrease - black curve in Fig.3.6.5a
(it would be similar in the case of irradiation with electron beams - a red curve, the intensity of which decreases rapidly with depth; it is not suitable for deep irradiation). Thus, in a deeper-placed tumor, the photon beam transmits the largest dose of radiation to the tissues in front of the tumor, only then (partially attenuated) radiation passes through the tumor and continues through even the healthy tissues behind the tumor. So healthy tissues and organs receive a relatively large radiation dose before and after the tumor... This leads to the risk of damaging important tissues and organs in the areas of radiation application. In anatomically more complex places, it is often difficult to decide which lowest radiation dose to use in order to ensure a therapeutic effect, without permanent damage to important tissues and organs.
*) "Depth effect" of high-energy g-radiation ("build-up effect" mentioned above in the introduction to the section "External irradiation with radiation g and X") is relatively small and is no longer dealt with here.
  Thus, in each individual photon beam from a given direction, the sites in front of the target tissue are irradiated even slightly more than the tumor itself, and the sites behind the target area are also exposed to only a slightly smaller radiation load. When irradiating from multiple directions, the total radiation dose at the target site will ultimately prevail, but the dose gradient and selectivity may not always be sufficient, especially when irradiating tumours in close proximity to important tissues and organs. However, there are physical mechanisms *) that allow this selectivity of the irradiation to be increased by achieving a more favorable dependence in the depth distribution of the dose: it is irradiation with heavy charged particles, often referred to as "hadron therapy".
*) Radiobiological factors also apply here. The biological effect of radiation is related to the ionization density given by the loss of radiation energy per unit path, the so-called linear energy transfer LET (§5.1 "Effects of radiation on matter. Basic quantities of dosimetry ."). Electron and photon radiation has a low LET, it is sparsely ionizing radiation. In contrast, fast protons, heavier ions, pions, neutrons, as well as products of nuclear reactions in the tissue, have a high LET - they show "dense" ionization and strong radiobiological effects, even for hypoxic tumors. The oxygen effect is significant especially when using sparsely ionizing radiation (photon radiation g or X is most often used ), where the indirect radicals mechanism of the radiation effect predominates. In densely ionizing radiation, where there is an increased proportion of the direct intervention mechanism (and also increased radicals recombination), the effect of oxygen (oxygenation) on the radiobiological effects is less significant. There is more frequent damage to the affected cells - the cells are inactivated, stop dividing and die by apoptosis. In addition, this higher radiation efficiency is accompanied by the possibility of better depth targeting this "heavy radiation" to the desired location.
  By hadron radiotherapy we mean irradiation with heavier particles - protons, heavier nuclei (ions),
p- mesons or neutrons (possibly antiprotons in the future), which collectively belong to the category of hadrons - particles showing strong interaction (see §1.5, passage "Systematics of elementary particles" and "Elementary particles and their properties"). However, when protons and heavier nuclei irradiation, does not use strong interactions, but electromagnetic interactions, by which these heavy charged particles intensely ionize the irradiated tissue (see the section "Common aspects of hadron radiotherapy" below). First, let's describe proton radiotherapy.

Fig.3.6.5. Hadron radiotherapy with proton beams.
a) Bragg curves of the depth dependence of the effective dose of radiation in the tissue when irradiated with gamma radiation, high-energy electrons and accelerated protons. b) Selective irradiation of the tumor site with a beam of protons of such energy, that the Bragg maximum lies in the depth of the tumor localization. c) Principle schematic representation of a proton radiotherapy workplace.

Proton radiotherapy
If we irradiate tissue with a beam of accelerated protons (with an energy of about 100-200MeV and a speed of about 1/2 the speed of light), the dose-dependence curve, so-called Bragg curve
(see §1.6 "Ionizing radiation", section "Interaction of charged particles"), has a completely different shape than for gamma radiation, as seen in Fig.3.6.5a (blue curve). During their flight, fast protons interact with matter in three ways :
¨ Coulomb interactions with electrons in atoms
The main mechanism, by which fast-flying charged protons lose their energy, is the inelastic electric interaction with the atomic shells of matter - ejection of electrons of atoms (Fig.1.6.1 top center). These secondary electrons are then a major factor in the radiobiological effect in the tissue. Due to the fact that protons are almost 2000 times heavier than electrons, interactions with individual electrons practically do not affect the movement of protons - the path of protons remains straight and the loss of energy of protons in matter is practically continuous.
    The secondary electrons from the proton beam have significantly lower energy than from the photon beams
(where the energy of the secondary electrons can approach the energy of the primary photons, ie several MeVs). Protons with an energy of the order of 100 MeV are relatively "slow" (speed max. c/2) - and only for this max. velocities are able to accelerate the secondary electrons. A simple kinematic consideration shows that the maximum energy of the secondary electrons here can then be about 50 keV. In reality, however, the energy of most electrons is much lower (protons in their rapid passage through the atomic shell suffice Coulombic to transfer only a small amount of energy to electrons) - usually only tens of eV (see spectrum in Bethe-Bloch formula in §1.6 "Ionizing radiation", passage "Charged particle interactions") .
¨ Coulomb interactions with atomic nuclei
For protons flying very close to the atomic nucleus
(with a small impact parameter) there is a repulsive Coulomb force which, due to the large mass of the nucleus, elastically deflects the proton from its original linear path. According to the law of conservation of momentum, the reflected core moves to the opposite side. In light materials (eg a hydrogen nucleus - a proton), a reflected nucleus can gain considerable energy. These effects may contribute to the partial lateral scattering of the proton beam.
¨ Nuclear reactions
Upon direct "intervention" of the nucleus
(with almost zero impact parameter), the proton enters the nucleus, where it can trigger a nuclear reaction (§1.3 "Nuclear reactions and nuclear energy", passage "Types of nuclear reactions"). The nucleus can emit a secondary proton, a deuteron, an alpha particle or a heavier ion, one or more neutrons, gamma photons. From the resulting secondary radiation from nuclear reactions, its penetrating component can be negatively manifested in proton therapy - gamma photons and "stray" neutrons, which fly to greater distances and can cause unwanted radiation exposure of surrounding tissues outside the target volume. An interesting use of nuclear reactions for imaging is mentioned below in the section "Nuclear reactions in hadron therapy and the possibility of gamma monitoring"; for a possible increase in the effectiveness and selectivity of proton therapy in the section "Proton-boron therapy ".
¨ Braking radiation protons in light materials is practically negligible, in contrast to electrons (§1.6, section "Interaction of charged particles").
    To illustrate the interpretation of the interaction of proton radiation with tissue, in comparison with other types of radiation, we will duplicate here the important figure 1.6.1 from §1.6
(we will be mainly interested in the part of "Protons 200MeV" at the top center and the corresponding curves on the right) :

Fig.1.6.1. Interaction of fast charged particles with matter.
Top left: Schematic representation of ionization mechanisms in the passage of beta
- and alpha particles .
Top middle:
Three basic mechanisms of proton radiation interaction with matter and braking of protons.
Bottom: Interaction of positron beta
+ radiation with a substance, ending in annihilation of a positron with an electron.
Right: Bragg curves of depth dependence of absorption and specific ionization along the path of gamma photons, accelerated electrons and protons.

When a charged particle passes through a substance, the linear transfer of (ionizing) energy is directly proportional to the electron density of the substance (which increases with density r and the proton number Z of the substance) and indirectly proportional to the square of the velocity of the charged particle, here the proton. Fast protons entering the tissue therefore initially ionize relatively little. As protons slow down and their velocity decreases, the ionizing effects increase - as the proton moves more slowly, the effective time of the electrical Coulomb action on the electrons in the atoms increases, so it is enough to transfer more energy and pull out more electrons.
    The dose distribution depending on the depth thus has a characteristic shape: as fast protons pass through the tissues, the initially absorbed dose is relatively low and almost constant, until it approach the end of the proton's reach in the tissue. Towards the end of the range, the dose increases sharply, reaches a maximum and then follows a very rapid decrease to zero. Fast protons forwards most of their energy in a narrow depth region of the so-called Bragg peak, just before their maximum range and stopping; here the densest ionization and the largest radiation dose occur. Approximately 70% of the proton's energy enters the region of the Bragg maximum and is absorbed there. Tissues lying in front of this maximum are irradiated with a significantly smaller dose
(only about 30% of energy is transferred here), tissues lying beyond this maximum they even get almost no radiation dose, because the protons do not reach there at all; after braking, the proton is neutralized by electron capture (hydrogen is formed) and further ionization no longer continues. Thank to this specific depth dependence of the dose with the Bragg maximum, it is possible to apply higher doses to the target volume, compared with standard photon radiotherapy and at the same time protect from radiation the surrounding healthy tissues, especially those that lie deeper behind target volume.
Note: Absence of Cherenkov radiation
Unfortunately, the proton beam of the used energy of approx. 200MeV can not be displayed using Cherenkov radiation in water in the same way as an electron or photon beam
(picture above in the passage "Cherenkov radiation"). The basic reason is that these protons are relatively "slow". A threshold energy of approx. 460 MeV is required for the emission of Cherenkov proton radiation. But the secondary electrons ejected from the atoms along the proton beam in the tissue usually have a very small energy of tens of eV or a small part of the keV unit (discussed above, see also the spectrum of the Bethe-Bloch formula in §1.6 "Ionizing Radiation", passage "Interactions charged particles"), much lower than the threshold energy of 260keV for the formation of Cherenkov electron radiation in water... However, it can be very well displayed in a liquid scintillator (picture in passage "Visualized invisible").
    The depth that occurs Bragg peak of the substance, is given by the energy of the proton; proton energy 200MeV makes this depth in the tissue of about 25 cm. Changing energy of proton beam can adjust the depth in which there is a maximum radiation dose. This can sensitively modulate the dose distribution within the body :
Proton beam modulation
The proton beam from the accelerator is relatively narrow and has a certain fixed energy, so the Bragg maximum is relatively sharp, so that the protons would transmit a sufficient radiation dose only at a narrowly defined location at a certain depth. The width of the Bragg peak for monoenergetic protons is only about 2 cm, which is often much less than the size of the tumor, which is also usually irregular in shape. Therefore, in order to sufficiently irradiate the entire tumor volume, it is necessary to shape and expand the proton beam, both in the transverse direction and in depth. There are basically two ways to proceed :
Passive modulation
By using suitable "deceleration" filters (wedge or stepped thickness) we scatter energetically the proton beam, so that we achieve the extension of the Bragg peak to the required dimensions. Depending on the thickness of the filter, the energy of the protons in certain parts of the beam is reduced so as to achieve the required irradiation of the tumor throughout its depth. Deceleration filters are mechanically made often in the form of modulating disks, which rotate in a controlled manner in a beam of proton radiation.
For the transverse shaping of beam, forming apertures and compensators are used, either fixed or formed individually for the patient according to the shape of the tumor and the irradiation plan.
A certain undesirable side effect of filters and screens is the formation of secondary neutrons, which are released during the interaction of high-energy protons with the atomic nuclei of the materials used. These parasitic neutrons contaminate the proton beam.
Active scanning
We irradiate the target area from each direction with a suitable meandering "scan", changing the energy of the particle beam and moving the maximum dose to different depths; gradually, the entire target volume is irradiated. However, most of the accelerators used - cyclotrons - do not have the ability to continuously change the energy of the beam, they have a fixed energy. The energy of protons is changed (reduced) externally by means of deceleration filters, mostly graphite degraders at the output. Synchrotrons have variable energy, but they are rarely used due to their greater complexity and cost.
New systems have developed magnetic deflection and narrow proton beam scanning ("pencil beam"), which is very flexible: no laborious individual apertures and compensators are required and no secondary neutrons are produced. Only this technique belong the future of proton therapy...
  A combination of hadron irradiation with conventional photon irradiation is also used to irradiate larger volumes of tissues.
Secondary particles in proton therapy
In addition to Coulomb interactions with electron shells (in which are ejected electrons causing radiobiological effects ), a small part of protons undergo nuclear reactions in the material - other secondary particles are formed, protons, photons, neutrons, deuterons,
a-particles (Fig.1.6.1 top middle). Secondary neutrons and photons can "travel" outside the target tissues and irradiate more distant tissues and organs (with a possible risk of secondary malignancies). Measurements have shown that the total fraction of energy escaping through secondary radiation is about 1-2% of the primary energy of protons. An interesting use of secondary radiation is mentioned below in the sections "Nuclear reactions in hadron therapy and the possibility of gamma monitoring" and "Proton-boron therapy".
Construction of the proton irradiator

The source of the proton beam - and thus the most important part of the hadron therapeutic system (Fig.3.6.5c) - is the accelerator. It is most often a cyclotron or synchrotron
(for accelerators, see §1.5, section "Charged particle accelerators"), the use of powerful linear proton accelerators can be expected in the future. Although according to the classification in §1.5 it is a "small accelerator" *), the accelerator laboratory occupies relatively large spaces - one large room (hall) with its own vacuum accelerator tube surrounded by strong electromagnets, shielding, as well as several smaller rooms with air conditioning, power and control electronics.
*) Cyclotrons for proton energies of 250 MeV tend to have a diameter of about 4-5m, synchrotrons about 6-8m. Significantly smaller compact accelerators are also being developed, cyclotrons with superconducting electromagnets, which could be mounted directly into the irradiator gantry.
    Combinations [cyclotron
-- > linear accelerator], sometimes called "cyclinac", are also tested. Behind the smaller cyclotron, which provides protons or heavier ions of fixed energy (approx. 30 MeV), there is a linear accelerator (linac) with a high gradient, which further increases the energy of particles to a value to reach the Bragg maximum at the required depth of the tumor lesion. This technology would allow easy, fast and flexible electronic beam energy regulation - active 4D scanning for moving organ therapy.
  Experiments with laser acceleration of protons are promising - §1.5, passage "
Laser accelerators LWFA". The intensity is, however, still very low, perhaps they can improve and apply in the more distant future..?..
    Accelerated protons are bring out from the accelerator by means of electromagnets and trough a vacuum conveying tube is fed into the irradiation room. On one accelerator may be connected to several irradiation rooms, where the individual sub-beams are led by transport tubes equipped with deflection electromagnets - Fig.3.6.5c
(only the main tube is drawn here, branching transport tubes to the irradiation facilities are not drawn due to space). In the end irradiation head *), "nozzle", in the irradiation room, the proton beam is shaped (as mentioned above in the passage "Proton beam modulation") by means of other precisely controlled electromagnets and enters the irradiated tissue. Proton beam can be focused by a strong magnetic field to a narrow "pencil" beam.
*) The irradiation head is often mounted on a special stand, the so-called gantry, enabling controlled rotation around the patient's body for isocentric radiotherapy (cf. Fig.3.6.1 above). The rotation of the proton radiation beam can be performed using a combination of mechanical movement of the gantry and controlled magnetic fields of electromagnets - it is a very robust and complex device (whose purchase price is close to the price of the primary cyclotron!) .
    As in conventional photon radiation therapy is often used here fractionated irradiation from multiple directions, intensity modulated beam (IMPT - Intensity Modulated Proton Therapy) in analogy to the above modulation for IMRT photon beams
("Modulated radiotherapy intensity"). An additional advantage of hadron radiotherapy is the ability to depth-adjust the area of maximum energy transferred to the tumor site for each beam. At a given energy, all heavy charged particles (protons or heavier accelerated nuclei) reach roughly the same place (depth) in the tissue, where they stop and transfer the maximum of their energy. Thanks to this increased selectivity, it is possible increase the focal dose (and thus increase the likelihood of effective destruction of tumor cells) without more serious damage to surrounding tissues.
  In special cases, proton one-time therapy is also performed for benign malformations or therapy for eye tumors.
    One of the other favorable physical properties of heavy particle beams is their minimal lateral scattering. A proton, whose mass is 1836-times larger than an electron, is only minimally deflected when interacting with the electron shells of atoms, it flies in one direction "forward" *). This feature also contributes to better targeting of the radiation dose to the desired location.
*) A larger part is directed in this direction of the primary beam secondary electrons released during the interaction of heavy fast particles with matter.
Advantages of proton and ionic radiotherapy
In summary, proton radiotherapy has three basic advantages :

A well-defined pathway and radiation dose during the movement of protons in the tissue, which can be regulated by proton energy.
The area of maximum dose distribution is narrowly localized and can be precisely adjusted by the energy of the particles. In the path before the Bragg maximum (at a smaller depth) the radiation dose and ionization density are relatively low Þ relatively small radiobiological effect. In the area of the Bragg maximum (locaized inside the tumor) the dose is high, the radiation densely ionizes and has a high radiobiological effect.
× At greater depths than the proton range, beyond the Bragg maximum (behind the tumor), the radiation dose is practically zero - healthy tissues behind the tumor are not affected by protons.
    This leads to a better opportunity to precisely target a high radiation dose to the tumor focus, while maximally conserving the surrounding healthy tissues. This can achieve high local control - effective destruction of the tumor site - especially in deeper tumors near important critical organs and complex anatomical structures, with lower risk of side effects and complications
(lower radiotoxicity), lower scattered radiation, lower risk of secondary radiation-induced malignancies.
In summary:

Radiotherapy with heavy charged particles allows maximizing the radiation dose in the target tumor volume and minimizing radiotoxicity in the surrounding healthy tissues.
    These advantages also have radiotherapy with heavier nuclei (ions) and, in theory, pions. In some cases, this is approached by the possibility of monitoring the radiation dose along the beam, eg by PET imaging of secondary radioactive nuclei emerging along the beam of high-energy carbon nuclei
(see the section "Nuclear reactions in hadron therapy and gamma monitoring options" below). Other physically interesting aspects of hadron radiotherapy are discussed below.
Note: Similar goals are achieved by somewhat other means - cybernetic gamma knife - the above stereotactic radiotherapy SBRT .

Bragg curves dependence of the dose depth distribution in the tissue (water phantom) for different kinetic energies of protons ( left ) and 12C nuclei ( middle ). Right: Example of the depth dependence of the radiobiological effect (surviving fraction of cells) on tissue irradiation (with radiosensitivity a~0.35) with a 150 MeV proton beam.

Radiotherapy with heavier nuclei (ions)
Accelerated protons (energies up to 250 MeV) are currently the most commonly used for hadron therapy. However, heavier accelerated particles - alpha particles or lithium, beryllium, boron, carbon, etc. nuclei
(whose wider use perhaps can be expected in the future), also have a similar and somewhat greater effect *); is referred to as hadron radiotherapy with heavy ions ("heavy ions" are considered to be atoms heavier than hydrogen, deprived of all or part of their electrons). The arrangement is basically similar to Fig.3.6.5c, the technology is even more demanding than with proton radiotherapy (synchrotron for 12C has a diameter of 20-25 meters!).
*) The Bragg curve of the depth distribution of the radiation dose has a slightly different shape for heavier ions than for protons. Bragg's peak has a slightly sharper increase before the maximum (lower irradiation of the tissue before the tumor). However, beyond the Bragg maximum, the dose does not drop as sharply to zero as for protons: the curve here has a kind of "tail" (representing about 10% of the dose in the input plateau), stretching about 2 cm to greater depth - the interaction of heavy ions with tissue atoms lead to fragmentation and sharply reflected lighter ions (mostly protons) are formed, which at high energy have a longer range than primary heavier ions.
    Of the heavier nuclei, accelerated carbon
nuclei 12C (carbon ions) are particularly suitable for radiotherapy. They are relatively easy to obtain (by ionizing carbon dioxide gas with electrons) in an ion source to accelerate, and show a high radiation contrast in the region of the Bragg maximum. In addition, nuclear reactions of 12C nuclei in tissue produce 11C nuclei that exhibit b+ -radioactivity, allowing scintigraphic monitoring of dose distribution in irradiated tissue by PET (see the section "Nuclear reactions in hadron therapy and gammagraphic monitoring options" below). Accelerated oxygen nuclei 16O have similar properties, including the formation of b+ -radioactive nuclei 15O, also suitable for PET monitoring of dose distribution.
Radiotherapy using mesons
p -
p- - negative pions, have a particularly significant radiation maximum at the end of their range in the substance (for their origin and properties see chapter 1.5, section "Properties and interactions of elementary particles", passage "Mesons p and K "). In addition to the usual mechanism of the Bragg peak (longer effective time of interaction of a slower moved charged particle with the atomic envelope of the substance), this contributes to this effect by the fact that at the end of its path the p- mesons are trapped in the nuclei of atoms (in tissue eg in cores of carbon 12C, oxygen 16O, nitrogen 14N). During this capture of the p- -meson by the nucleus, its reaction with the proton (p- + p+ ® n0 + 140MeV) releases energya bout 140 MeV, that is higher than the binding energy, so that the excited nucleus usually cleaves into a- particles, deuterons, neutrons and protons (for heavier nuclei, 6Li or 12C also occur between the fragments). E.g. for carbon there is a reaction p- + 12C ® 2a + 3n + p, whereby particles a carry away kinetic energy of about 30MeV and neutrons about 70MeV (the remaining 40MeV is used to overcome the binding energy of the nucleus). By braking these fragments, considerable ionization energy is transferred at a given site, i.e. a dose of radiation that effectively kills the tumor cells.
p- are obtained by bombarding target nuclei (eg carbon or beryllium) with protons accelerated to high energies, greater than about 500 MeV, in a large accelerator (eg synchrocyclotron). A certain problem is the very short lifetime of these particles p-, about 10-8 seconds, so they cannot be distributed to more distant irradiation facilities. The range of pions with energies of 50-100 MeV in the tissue is about 10-25 cm. When pions decay at the end of their path, in addition to the useful transfer of large ionization energy, there is also some undesirable scattering of the radiation dose; also, fast neutrons flying away from the point of interaction of the pions cause a certain radiation dose outside the target volume. Due to its high technical complexity and cost, this method is still only in the stage of experimental testing in a few of the largest accelerator centers...
Antiproton radiotherapy

Other unusual particles that could potentially be beneficial for targeted radiotherapy, are antiprotons p' - negative protons p
- (their properties have been described in §1.5, part "The properties and interactions of elementary particles"). Accelerated antiprotons after entry into tissue ionize similar to normal protons - initially low ionization density slowly increases and just before braking is a significant increase of ionization in the Bragg peak. However, after braking, in addition, the antiproton annihilated with a proton or neutron in the atomic nucleus of the irradiated substance (tissue) to form p-mesons, typically: p' + p®2p+ +2p- +po. Secondary mesons p- they can behave as described in the previous paragraph; in general, positive and negative pions decay rapidly into muons and neutrinos, the neutral pion into two quantum gamma (these particles usually escape from the site of interaction). Additional quanta can be emitted from "affected" nuclei (unless they are hydrogen nuclei) by the mechanism of nuclear reactions. Antiproton annihilation thus releases additional energy up to several hundred MeV at the site of the Bragg peak, which significantly increases the radiation effect at the site of the Bragg maximum - about 3 times compared to protons. An accompanying phenomenon during interactions are also positrons, whose annihilation gamma-photons of energy 511keV can be detected using a PET camera and thus monitor the actual distribution of the radiation dose in the tissue.(similar to that mentioned below for radiotherapy with 12C carbon nuclei ) - Fig.3.6.6. A certain disadvantage of antiproton therapy is the slightly higher radiation dose outside the target volume (including the whole body dose), caused by penetrating pions, neutrons and g, flying in all directions from the site of antiproton interaction.
    Antiprotons p- can be prepared by bombarding target nuclei with accelerated protons in reactions p + p
® 2p + p + p' and p + n ®2p + n + p'. The kinetic energy of the protons must be at least 5.6 GeV, but to achieve a higher yield over 20 GeV, which can only be achieved on large accelerators. The resulting antiprotons fly with considerably high kinetic energies of several GeV, so for radiotherapeutic purposes, it is necessary to slow them down the energies of about 100-200MeV in decelerators. This resulting energy determines the range of antiprotons in the tissue and thus the depth of the Bragg maximum of the radiation dose. The method is in the stage of laboratory testing in the largest nuclear laboratories (CERN, FERMILAB); due to the extraordinary complexity and cost, the introduction of this interesting method into clinical practice cannot be expected in the foreseeable future ...
    Conclusion: Due to the mentioned problematic aspects (predominant disadvantages), extraordinary complexity and cost, neither pion nor antiproton radiotherapy in the foreseeable future probably will not used...
Neutron therapy
In principle, neutron radiation beams can in principle also be used for radiotherapy. Either they are fast neutrons, which collide with nuclei in the tissue, especially hydrogen nuclei, to form accelerated protons that have strong ionizing effects. Fast neutrons have high LET and radiobiological efficiency, but depth dose distribution in the tissue is not more advantageous than with gamma radiation. In addition, the neutron beam is difficult to collimate and modulate, and exhibits considerable scattering "to the sides" of the original direction in the tissue.
  An interesting unconventional method for increasing the selectivity of neutron irradiation of tumor lesion is called neutron capture therapy (NCT - Neutron Capture Therapy) by means of slow neutrons. In this therapeutic procedure, suitable atoms whose nuclei have a high effective neutron capture cross section are bound to the tumor site by means of a suitable compound, which is preferentially taken up and accumulated in the tumor tissue (see §1.3., 1.6) - boron enriched in isotope
10B is used. Special boron compounds (BSH-mercaptododecarborate, or BPA-dihydroxyboralfenylalanine) have been developed for brain tumors, which penetrate only marginally into healthy brain tissue, but are selectively taken up in tumor tissue cells that have a disrupted blood-brain barrier. Bor-Deoxy-Glucose can be used for metabolically active tumors elsewhere in the body.
  The tumor deposit prepared in this way is then irradiated with a beam of low-energy neutrons (with energies of about 1eV-10keV), which slow down (moderate) to thermal energy as the tissue passes and are then trapped in the boron nuclei, wherein by the reactions (n,
a): 1n + 10B ® 11B* ® 7Li + 4He disintegrate the boron core occurs and the emission of helium (ie alpha particles) and lithium nuclei. The resulting alpha particles and lithium nuclei, carrying away considerable energy released in the reaction, have a very small range in the tissue, it stops about 10 µm from the reaction site, so that the ionization energy is transferred practically directly inside the respective tumor cells, which can be effectively destroyed, without radiation damage to the surrounding tissues. The described method is currently experimentally tested in brain tumors of glioblastomas (and also brain metastases of cutaneous melanoblastoma).
    The source of neutrons for radiotherapy can be either a nuclear reactor
(§1.3, part "Nuclear reactors"), but in laboratory conditions neutron generators are more advantageous, special small accelerators of charged particle, mostly deuterons, with a tritium target (§1.5, part "Charged particlesaccelerators", passage "Neutron Generators"), or radioisotope source consisting of a mixture a-radionuclide with light element (such as mixtures of americium with beryllium, reacting a, n), or heavy transuran radionuclide (typically californium-252) during the spontaneous fission neutrons are released ( §1.3, "Transurans"). For capture therapy, neutrons at first slowed in a moderator.
  Neutron capture therapy dot not exceed the framework some experimental studies
(in practice not worked too) and is now used only sporadically... Regarding the use of boron, on the contrary, the "Proton-boron therapy" below, is more promising here.

Nuclear reactions in hadron radiotherapy and the possibility of "in-beam" gammagraphic monitoring
When irradiated with high-energy hadrons (protons, pions, fast neutrons, antiprotons), most of these particles interact with the atomic shells of the irradiated substance (tissue); this gives their ionizing and radiobiological effect. However, a small part of these particles also undergo a number of nuclear reactions with the nuclei of atoms in the tissue, during which various secondary particles and fragments of nuclei are formed. In these nuclear reactions, secondary (or tertiary) radiation is emitted, which can in principle be detected and used to monitor the dose distribution in the tissue. This is a hadron activation analysis using the resulting gamma radiation
(only this high-energy photon radiation is penetrating, can fly out of the irradiated tissue and be detected). The principle of activation analysis was given in §3.4, part "Neutron activation analysis" (where it was mainly neutron activation, but similar principles apply to irradiation with protons and heavier ions). Gamma radiation in nuclear reactions here is basically created by two mechanisms :
Deexcitation of the excited levels of target nuclei formed after the reaction. This radiation has a line spectrum with a number of energies (in a wide range of about 100keV-10MeV) and can be analyzed using gamma spectrometers with scintillation or semiconductor detectors (HPGe) - see section "" below. Gammagraphic imaging for such high energies is difficult with standard scintigraphic collimators, special slit collimators are tested, experimentally also high energy cameras with electronic collimation, using Compton scattering kinematics (§4.2, part "Alternative physical principles of scintillation cameras", passage "Compton camera").
× Positron emissions from positron radionuclides formed during nuclear reactions. These positrons in the tissue annihilate with electrons to form pairs of annihilation radiation photons gamma energy 511 keV. The distribution of positron radionuclides can be imaged using a PET gamma camera.
  In terms of time, it is a secondary radiation of two types :
¨ Prompt gamma radiation, emitted immediately during the reaction or immediately after the reaction due to deexcitation of the excited nuclear levels in the irradiated substance. This radiation must be measured "on-line" directly during irradiation.
Subsequent - delayed - gamma radiation, emitted by radionuclides generated after the reaction, which have a certain shorter or longer half-life. Here we can also measure and display "off-line" with a certain time interval, with respect to the half-life of the analyzed radionuclide.
In-beam PET monitoring 
From the point of view of gammagraphic monitoring of the passage of a radiation beam through a substance (tissues), such reactions that lead to the formation of positrons are of particular interest: either e
+ are emitted directly, or nuclei showing positron b+ -radioactivity are formed. Annihilation of positrons with electrons is accompanied by the emission of pairs of opposite quanta of annihilation radiation g of energy 511 keV. These photons can then be detected by the positron emission tomography method of PET (see "PET cameras" in Chapter 4 "Radionuclide scintigraphy") - Fig.3.6.6a. The most common positron radionuclides generated by the high-energy proton irradiation of tissue, are: 11C (T1/2 = 20.3 min.), 15O (T1/2 = 122 s.) and to a lesser extent 10C (T1/2 = 19.3 s.). As there is some correlation *) between radiation dose and induced positron radioactivity (or direct positron emission and annihilation radiation), this allows scintigraphic monitoring of dose distribution in irradiated tissue using PET - "the invisible making visible" online, in situ, or off-line with the use of emerging positron radionuclides with a not too short half-life (lower parts of Fig.3.6.6 b,c).
*) This correlation can be negative or positive, as shown by the curves in Fig.3.6.6b,c. It depends on the mechanism of reactions and the kinetic balance of the irradiating particles and the induced radioactive nuclei. In proton therapy, this three above mentioned radionuclides are formed by the ejection of neutrons from carbon and oxygen nuclei in the region of high proton energy, before the Bragg maximum; the correlation is negative - the greatest activity is induced in the region of low LET fast protons, in the region of Bragg's maximum there is none (Fig.3.6.6b).
  During irradiation of the
12C nucleus, two kinds of strip reactions take place: 1. A neutron is entrained from the 12C nucleus, which produces b+ -radioactive 11C, which continues in flight and stops at the site of the Bragg maximum - a positive correlation between the dose of 12C and the induced b+ -radioactivity (Fig. 3.6.6c). 2. The fast-flying 12C nucleus ejects a neutron from the carbon or oxygen nucleus in the tissue, creating 11C, 15O or 10C, which remains at the site of its origin (ie outside the Bragg maximum) - again a negative correlation. There is also a positive correlation between the dose distribution and the intensity of annihilation g- radiation in the case of a nuclear reaction of antiprotons or pions p- at the point of their cessation (Bragg maximum).
  For PET monitoring of dose distribution, a case of positive correlation is more suitable, even under conditions of sufficiently high induced
b+ -activity. With a negative dose-activity correlation, the PET images are of poor quality, only the passage of the tissue bundle outside the target volume can be monitored (Fig.3.6.6b below); is briefly discussed below.
Note: b+ -radioactive isotopes are induced even in classical irradiation with g- beams with an energy higher than about 10 MeV, but the activities are very small, insufficient for gamma imaging.

Fig.3.6.6. a) Possibility of gammagraphic "in-beam" monitoring of hadron radiotherapy using positron emission tomography. b) Negative correlation between dose distribution D and induced
b+ -radioactivity in the proton beam. c) Positive correlation between dose distribution D and induced b+ -radioactivity in irradiation with accelerated carbon nuclei.

A typical example of nuclear reactions enabling PET-monitoring with a positive dose correlation is radiotherapy with accelerated carbon nuclei (ions), where during nuclear reactions *) in the tissue from the part of the 12C nuclei, the b+ -radioactive 11C nuclei are formed, which continue to fly and it stops, as well as the basic nuclei 12C, at the place of the Bragg maximum (Fig.3.6.6a, c). By gammagraphic PET-imaging of the positron radioactivity thus induced, we obtain an image of the distribution of the sites in which the 12C + 11C nuclei stopped and delivered the maximum radiation dose. With a PET camera installed on the irradiator, we can monitor the distribution of the dose in the target tissue and in the environment - thus control the course of radiotherapy similarly to the IGRT method (mentioned above). This method is called in-beam PET monitoring - monitoring directly in the irradiation beam using PET - Fig.3.6.6a.
*) "Peripheral" nuclear reactions, so-called strip reactions (see the section "Mechanisms of nuclear reactions" in §1.3 "Nuclear reactions"), in which a neutron is detached from the flying nucleus 12C during interaction with the nucleus in the tissue 12C ® 11C + n, are mainly used here. This creates a neutron-deficient carbon nucleus 11C (continuing in motion), which is b+ -radioactive: 11C ® 11B + e+ + n with half-life T1/2 = 20.3min.; it disintegrates only after stopping at the site of Bragg's maximum. Subsequently, the positron e+ is annihilated with the electron: e+ + e- ® 2 g, while these two oppositely scattering quantum g with energies of 511 keV can be used for PET scintigraphy. This displays the distribution of 12C + 11C core stop points, which are also the sites of the largest radiation dose in the Bragg maximum. Due to the short half-life of 11C, PET imaging must be performed immediately after hadron irradiation - either directly with a PET camera installed in an irradiation facility, or within a few minutes on a camera in another room in the workplace.
    In proton radiotherapy, radionuclides 11C and 15O are formed by the ejection of neutrons from carbon and oxygen nuclei in areas where protons have high energy, ie in the input beam to a distance of about 1-2 cm before the Bragg maximum. Activated nuclei remain "standing" in their original places in the tissue (or are reflected only at short distances). In the region of the highest dose, the activation is zero (the energy of the slowed-down protons is subthreshold, not enough for the reaction). Here, too, PET images may provide some information on dose distribution, but there is a negative correlation between the absorbed dose and the induced activity on the PET image.
    Scintigraphic imaging of PET is also very suitable for monitoring the biological response of tumor tissue to radiotherapy in general (both hadron and conventional
g radiation ), as it is able to monitor the cellular activity of the tissue - to distinguish the remaining (or recurrent) viable tumor cells.
    Positron radionuclides
(11C, 15O, 13N, ...), produced along the path of the hadron beam, can thus also be measured by means of a PET camera. Although this technique is suitable for additional adjustment and correction of the irradiation beam, it does not allow real-time on-line monitoring. It may be effective for less perfused structures (such as a scaffold), but in well-perfused tissues, rapid biological leaching and movement of induced radioactivity occurs. Due to the generally low induced activity, longer PET acquisition times are needed.
Hadron-gamma-activation analysis - prompt gamma monitoring of the irradiation beam
During the passage of the hadron beam through the irradiated tissue, there are a small percentage of interactions with the nuclei of the irradiated substance to form excited nuclei, which then deexcite by emission of prompt gamma radiation. Thus, another possibility of in-beam monitoring is the analysis of this deexcitation
g- radiation generated along the hadron beam. Scintillation gamma cameras with slit collimators, that provide 1-D projection of prompt gamma radiation along the path of proton rays, are tested. This method is suitable for monitoring in pencil beam mode, but is not applicable to irregularly shaped fields in scattering mode, where different parts of the field reach different penetration depths.
Proton-boron capture therapy
Recently, the possibilities of further increasing the effectiveness of proton therapy for the selective killing of tumor cells using nuclear reactions of protons with appropriate substances incorporated into tumor tissue have been explored. Most promising one is the proton-boron capture therapy PBCT (
Proton-Boron Capture Therapy). Therapy of this kind is performed in two steps :
1. A suitable compound containing boron atoms, isotope 11B , is first captured in the tumor site. Increased metabolism of tumor cells compared to normal cells can be used for this. If we also apply, for example, glucose with chemically attached boron atoms - Bor-Deoxy-Glucose (it is analogous to the well-known Fluordeoxyglucose 18FDG, used in PET scintigraphy), it will accumulate more in tumor cells. Other such compounds are mercaptododecarborate (BSH) or dihydroxyboralfenylalanine (BPA). For chemical coupling it is sufficient to use natural boron, which contains 80% of the isotope 11B and 20% of the isotope 10B.
2. The tumor deposit prepared in this way, in the cells of which boron is contained, is then irradiated with a proton beam. During the interaction of protons with boron nuclei, nuclear reactions p + 11B ® 3 a *) occur, in which three alpha-particles with an average energy around 3 MeV are emitted, the total released energy has the value Q = 8.7 MeV. These alpha particles are immediately braked about 10 µm from the reaction site, so that high ionization energy is transferred practically directly inside the respective tumor cells, which can be effectively eliminated by DNA birefringence, without radiation damage to surrounding tissues.
*) This reaction proceeds in 3 stages: first the capture occurs - the fusion of the proton with the boron nucleus 11B to form the excited carbon nucleus 12C*, which immediately decays into beryllium 8Be and alpha particle 4He, after which 8Be is immediately cleaved to 2 alpha particles: p + 11B ® 12C* ® 8Be + 4He (3.8MeV); 8Be ® 4He + 4He (2.4 + 2.4MeV). The overall result is the emission of three alpha particles: p + 11 B ® 3 a . The reaction has an increased effective cross section of 1.2 barn for proton energies around 700 keV, which corresponds well to the slowed protons in the region of the Bragg maximum. The proton-boron reaction also emits gamma photons with a main peak of 718 keV, which could in principle be used for "in beam" gamma monitoring of the distribution of the alpha-particle dose along the proton beam in the target tissue and around; however, due to the low concentration of boron, this weak radiation will be difficult to detect against a much stronger background of secondary radiation arising from nuclear reactions of protons with carbon and oxygen nuclei in the tissue, especially the annihilation gamma 511keV (Fig. 3.6.6).
    The result of such combined proton therapy, "enhanced" by secondary alpha-particles from nuclear reactions, is selectively higher radiobiological efficacy compared to the protom beam itself. It is a molecular - biologically targeted proton therapy. A significant improvement in the therapeutic effect can be expected, especially due to densely ionizing alpha particles, especially in hypoxic and radioresistant tumors. However, the basic condition for the success of this method is a sufficiently efficient and selective uptake of boron in the target tumor tissue.

Common aspects of hadron radiotherapy
Somewhat unusual name "hadron therapy" originated because particles that interact with a strong interaction - so-called hadrons - are used here
(see §1.5, passage "Systematics of elementary particles", "Elementary particles and their properties" and 1.6 ). However, in the case of its own therapeutic effect, the electromagnetic interaction leading to the ionization of the substanceis used here in particular; a strong interaction is seen with p, antiproton and neutron or proton capture therapy. What is important for a given application is that, in the end, the particles are heavy and electrically charged, with high radiobiological effects and specific depth distribution of radiation dose. In p-, neutron or antiproton therapy, it is a bit of an exaggeration to say, that at the end of the particle path, inside the target tumor tissue, there is a kind of miniature "nuclear explosion" whose energy effectively kills tumor cells, with minimized radiation damage to surrounding tissues. Even with proton or ion therapy, it can be said that the radiation dose in a sense "explodes" at the site of the Bragg peak, which should be located inside the tumor.
    The relationship between dose and biological effect is also basically given by the standard linear-quadratic (LQ) model mentioned above (it is analyzed in detail in §5.2, part "
LQ model"), from which, however, there are some deviations. In addition to the square, minor corrections should be included, including higher powers of the dose, originating in multiple interactions of densely ionizing radiation with the DNA structure.
    All "hadron" methods outlined above - proton therapy (+ ion, muon or antiproton) and neutron capture therapy, we present here mainly because they are very interesting in terms of nuclear and radiation physics. Only proton therapy has so far developed into the wider therapeutic practice, the others are still in the laboratory testing stage.

For irradiation of smaller volumes of target tissue, it is sometimes possible to use the so-called brachytherapy *) - a method of local radiotherapy, in which the radiation source is in close contact with the tumor site. The condition for the usability of brachytherapy is the mechanical availability of the lesion. In organs affected by cancer, the radionuclide radiation source is introduced (by puncture or implantation) either directly into the tumor bed (interstitially), or is introduced intracavitatively into body cavities (eg into the uterus), or intraluminally into tubes, or it is attaches to the surface of the tumor (so-called mulch).
*) Greek brachys = short - this is radiation from a short distance, "at close range", in contrast to teletherapy as radiation "at a distance". From this point of view, we can compare it with open radionuclide therapy, which is "completely close" - at the cellular level; see the discussion below "Radioisotope therapy with open emitters - the closest possible “brachytherapy".
    Selective irradiation of the tumor bed is achieved here by the highest radiation intensity in the immediate vicinity of the emitter, while decreasing sharply at greater distances - in vacuum it would be approximately with squared distance, in tissue it is even faster due to the exponential absorption of radiation. It is therefore possible to concentrate a very high dose of radiation on the tumor site, usually without the risk of more serious damage to the surrounding healthy tissues.
    If we have a radionuclide gamma emitter - radiophore - of radius r
0 with activity A [GBq], located in the tissue with the absorption coefficient m for the emitted energy of gamma radiation, then the resulting intensity I of gamma radiation (and thus the dose rate D' ) in the surrounding tissue at a distance r (> r0) will be given by the product of the geometric inverse quadratic dependence r -2 and absorption exponential dependency :
                   D'(r) ~ I (r)  =  G . A / r 2 . e -m . r  ,  r> r0  . 

Left: Dependence of gamma radiation intensity on the distance from a point or spherical radionuclide source of activity A in the material environment.
Scintillation radiation of a
192Ir brachytherapy radiophore (400 GBq activity) immersed in a flask with a liquid scintillator.

For r < r0 - inside the source - the course of radiation intensity is different, depending on the construction, material and distribution of radioactivity within the radiophore (however, this is not reflected in brachytherapy). Initially, for close distances outside the source, a quadratic decrease predominates, at greater distances the exponential decrease is more pronounced (attenuation of radiation in the tissue).
    In terms of time, brachytherapy is divided into temporary and permanent brachytherapy :
Temporary brachytherapy
Closed radionuclide emitters of a longer half-life - radiophores - are introduced into the target tissue for a specified period of time T. Of this application time is then proportional the radiation dose D - simply expreses: D
~ G. A.T, where G is the radiation dose constant for a given type of radiation resp. radionuclide, A is the activity of the radiophore (which due to the long half-life of the radionuclide can be considered constant during a relatively short exposure).
  As a radiation source for brachyradiotherapy, radium
226 Ra (a- decay, T1/2 1602 years) in the past has long been used, whose decay products (such as 214 Pb, 214 Bi, 214Po , ...) they are gamma emitters. Radium had some disadvantages, eg radon is formed during its decay (for closed radiophorers however, it does not penetrate out), as well as low intensity of g- radiation leading to long exposure times (approx. 2 days) - it was LDR brachytherapy (Low Dose Rate, < 2 Gy/hour). Therefore, radium was gradually replaced by some other artificial radioisotopes: cobalt 60 Co , cesium 137 Cs and especially iridium 192 Ir. With sufficiently high radiophor activities (approx. 400GBq, mostly 192Ir), the exposure time is reduced to tens of minutes - HDR brachytherapy (High Dose Rate, >10 Gy/hour).
  Rarely is used the so-called pulsed brachytherapy PDR (Pulsed Dose Rate), in which the radiation dose is delivered fractionally during one brachytherapy application: the radiophore moves in a sequence of repetitive steps - in "pulses". Neutron radiophores are also used experimentally (especially with a
252Cf californium), where neutron radiation has a high LET - higher ionization density and a stronger radiobilogical effect (even on hypoxic tumors).

Fig.3.6.7. Two basic techniques of brachyradiotherapy.
Left: During temporary brachytherapy, from the holes of the head of the shielded box, the radiophores by afterloading are led through hoses to the target area and then returned to the container after exposure. Right: During permanent interstitial brachytherapy, small radiophores are permanently implanted into the target tissue using applicators.

Radiators, or radiophores, for brachytherapy are closed encapsulated radioisotopes, the envelopes of which are in the shape of needles, tubes, wires or rollers. A significant improvement of the brachytherapy technique is the so-called afterloading (afterloading - addtional load, introduction) : an inactive hose - applicator is first inserted into the target area or body cavity, which is precisely adjusted (with the help of a radiophore modeling mark, placed at the end of the application tube). For precise settings, as well as for batch distribution planning, X-ray navigation is performed (most often C-arm, at least 2 perpendicular projections to obtain a spatial-volume image, or CT), ultrasonography or NMRI. Own radiator - a radiophore - in the shape of a small cylinder (mounted on a guide wire) is then introduced into this hose for a specified period of time, which is returned to the shielded box after the end of the exposure. The former manual afterloading has now been replaced by automatic afterloading, in which radiators are inserted and returned using electronically controlled motors. The shielded container may contain one or more radiophores *), which can be introduced either gradually or simultaneously into different transport paths - Fig.3.6.7 on the left.
*) On the idea diagram in the left part of Fig.3.6.7, all possibilities are shown for generality. However, current HDR afterloading devices usually contain only one more powerful 192Ir source. This is for economic and technical reasons: radiators are expensive and decays (need to be replaced); the mechanism for simultaneous introduction of multiple emitters is technically complicated. This intensive radiophore is gradually introduced into the various transport pathways, if necessary, to achieve the desired dose distribution in the target volume, into which the individual applicators are introduced; this insertion is fast enough. For precise uniform implementation applicators in the desired geometry by irradiation plan is sometimes used special grid template, similar to the one schematically shown in the right part of Fig.3.6.7.
    Compared to the manual application of radiophores, automatic afterloading has two main advantages :
- It significantly reduces (or even completely eliminates) the radiation exposure of workers when performing brachyradiotherapy.
- By moving the sources in the applicator, we control the time for which the source is in certain positions, thus achieving the required distribution of the radiation dose in the target volume. With this controlled modulation of the dose distribution, the automatic afterloading significantly precise the therapeutic effect.

    Miniature electronic X-ray sources could be promising emitters for brachytherapy
 (working on the development of laser micro-X-rays) with adjustable dose rate and selectable effective X-ray energy. They would also have the advantage of better radiation safety and the absence of radioactive waste.
    The radiobiological effect of brachytherapy can be described by a general linear-quadratic (LQ) model (given above in the section "Dependence of radiation biological effect on dose and its timing"), but including specific factors of spatial and temporal dose distribution in different types of brachytherapy :

HDR brachytherapy , in which individual applications last several minutes (up to tens of minutes), in terms of time distribution of the dose is similar to fractionated external radiotherapy EBRT. Reparation during the fraction is not applied, the dose rate factor is negligible. However, cellular repopulation of tumor cells between fractions may occur, with a total treatment duration of relatively long. For HDR therapy of tumor tissue with coefficients a, b and doubling half-life T2r of repopulation, with total dose D divided into fractions d during total treatment time T, for biologically effective dose of BED the LQ model (without repair but with repopulation) is based on: BED = D. [1 + D / (a/b)] - T. ln2 / (a .T2r).
l LDR brachytherapy, in which there is a continuous exposure with a low dose rate with a relatively short total treatment time, the repopulation of tumor cells is significantly limited (a constant dose rate to the target volume takes place). For LDR exposure of tumor tissue with coefficients a, b and repair rate constant l , with dose rate R and time T , for the biologically effective dose of BED the LQ model (with repair but without repopulation) comes out a simplified relation: BED = R.T [1+ 2R / (l . (a/b)]. (1-1/l.T).
The case of permanent LDR brachytherapy is discussed below :

Permanent brachytherapy
In this method, radiators with a shorter half-life (days to tens of days) are introduced into the target tissue permanently and have a long-term effect until their disintegration and fading. Dose rate R decreases exponentially with a half-life T
1/2 of the radionuclide used: R(t) ~ G.A(t) = G.Ao.e-(ln2/T1/2).t, the total radiation dose is given time integral D ~ G. 0nAA(t) dt = G.0nAAo.e-(ln2/T1/2).t dt ~ G.Ao.T1/2. The total radiation dose is thus given by the initial applied activity Ao of the implanted emitters.
    The radiobiological effect is generally expressed by the equation of the LQ model (given above in the section "Dependence of the radiation-biological effect on the dose and its time schedule"), containing the time factors of cell repair and repopulation. In the initial stages of permanent brachytherapy, there is practically no cellular repopulation due to the constantly supplied sufficiently high dose rate. From the thus reduced equation of the LQ model (with repair, but without repopulation) it follows for the biologically effective dose of BED in permanent brachytherapy of tumor tissue with coefficients a, b and repair rate constant l, using a radiophore with half-life T1/2 and initial dose rate Ro , simplified relation: BED = Ro.(T1/2/ln2).{ 1+Ro/[(l+ln2/T1/2).(a/b)] }. The initial dose rate here is proportional to the initial applied activity: Ro ~ G. Ao .
    Due to the exponential decrease of the dose rate with time, the radiation effect is highest in the initial phase. After expiry about 3-4 half-life the dose rate decreases to such an extend, so that the cell repair mechanism are sufficient to eliminate the relevant changes - deterministic radiation effects disappear, continued
(and increasingly weakening) the radiation dose is no longer therapeutically effective *), is "unnecessary" ("Wasted dose" - "barren" ) - cf. with a discussion of the effect of dose rate in §5.2., part "LQ model".
*) The proportion of ineffective (unnecessary, wasted) dose may be somewhat lower (and the therapeutic effect thus reasonably higher) than would result from the LQ model, due to two circumstances :
If tumor regression occurs during irradiation, the tumor tissue may "shrink" and the remaining tumor cells may become closer to the radiophores, increasing their dose rate.
2. If there is hypersensitivity to low doses (see §5.2, section "
LQ model", section "Deviations from the LQ model", Fig.5.2.4c), deterministic radiation effects can continue secretly even at low doses.
    Permanent interstitial brachytherapy consists in the implantation of a large number (several dozen) of small radiophores, using suitable needle applicators of manual afterloading (under X-ray or ultrasound navigation), or using a template
 the shape of a regular grid, directly into the target tissue - Fig.3.6.7 on the right. The most commonly used grains *) containing radioiodine 125 I (g 35keV, X 27keV), each with an activity of approx. 10-20MBq, or 103Pd (X 21keV), or 131Cs (X 33keV). Gold 198Au has a short half-life of 2.7 days, but too high an energy of g 412keV, so it is no longer used. Irradiation is relatively long-term (half-life 125I is 60 days, 103Pd has T1/2 = 17 days, 131Cs is T1/2 9.7 days), in terms of dose rate, it is LDR (or even VLDR - Very Low Dose Rate) brachytherapy, while due to the low energy of photon radiation, the radiation exposure of the environment is minimal, virtually all radiation is absorbed in the tissue. The method is suitable for slow-growing tumors, it is mainly used for prostate cancer. In addition to 125I, beta emitters such as 106Ru (®106Rh, half-life 174 days, max. energy 3.5 MeV) are also rarely used to irradiate small deposits in more complex structures, eg in ophthalmology.
*) Several grains of the radiophore can be joined together using a special fiber (so-called strand). As they are gradually ejected from the needle applicator, the grains are arranged linearly at the same distance, which ensures a more homogeneous irradiation of the target area and also prevents individual shifts (migration) of individual grains in the target tissue or even their escape from the target tissue.
    Computer planning systems (although not as complicated as IGRT) are currently used also for brachytherapy, which determine the dose distribution in the target tissue based on images of the source position (simulated by markers in the afterloading applicator) and determine the times and movements of the source of the afterloading control system (or activities and positions of permanently applied radiophores), such that the required dose distribution is achieved in the irradiated target volume.
  The so-called bystander effect (see §5.2 "Biological effects of ionizing radiation", passage "Bystander-Abscopal effect") could also partially contribute to a more effective irradiation of tumor tissue, which could perhaps somewhat correct the effect of mild inhomogeneities in tumor tissue irradiation - increase the effect in underexposed portions of the target tissue.

Tumor therapy: ionizing radiation - or chemistry ?
Currently, the therapeutic use of ionizing radiation is very important and beneficial here. However, this method only affects the consequences, but does not solve the causes of the disease.
 Radiation also often behaves like an "elephant in porcelain" in the body !
 It can be hoped that the future of cancer treatment will lie more in advanced chemical, biochemical and immunological methods - at the cellular and molecular level .

Radioisotope therapy with open emitters - the tightest possible "brachytherapy"
If we apply a radioactive substance to the body, it enters the metabolic process in a way that is determined by the chemical form of the substance - its biodistribution, pharmacokinetics. If we manage to label with a suitable radionuclide a substance that is selectively taken up and accumulated in the tumor tissue, we can get a very effective way of radiation elimination of the tumor: the ionizing radiation emitted by the radionuclide attacks the cells "from the inside". For radioisotope therapy, radionuclides emitting beta or alpha radiation with a short range in the tissue (max. millimeters) are suitable, which have strong local radiobiological effects and at the same time do not reach the surrounding healthy tissues. Therapeutic radiopharmaceuticals are labeled (conjugated) with these radionuclides. The beta or alpha radionuclide delivers a biologically effective dose of radiation selectively to accumulating tumor cells and the surrounding microenvironment. From the point of view of the above-mentioned division of radiotherapy methods
(according to the method of "targeting" the radiation dose to the affected tissue), radiotherapy using open radionuclides can be described as the closest possible brachytherapy (permanent) - directly at the cellular level.
    However, due to its physical, chemical and biological specifics, this targeted therapy with open radionuclide emitters forms a separate radiotherapy category and is mostly performed in nuclear medicine workplaces. It is called biologically targeted radionuclide therapy (BTRT) or molecular radiotherapy, sometimes endoradiotherapy. Unfortunately, we do not yet have such a suitable substance for most cancer processes; it will be shown below, when such a radioactive substance it at least partially "we have" and how it can be used for effective radionuclide therapy. The application of open radionuclides is also used for non-tumor therapy, especially in hyperthyroidism and in radionuclide synovectomy
(see below "Thyroid therapy" and "Radionuclide synovectomy").
    The methodical approach here is completely different from external therapy and brachytherapy. In classical radiotherapy with external beams, we need to know the exact localization and extent of the tumor foci, which we then irradiate precisely directed radiation beams
(Fig.3.6.8 left), or introduce into the lesion's proximity a brachytherapeutic radiophore. In radioisotope therapy, we need to know the biological (biochemical, pharmacokinetic) properties of the tumor tissue, according to which we apply a suitable radiopharmaceutical to the metabolic environment of the organism (Fig.3.6.8 on the left), which gets inside the tumor deposit (it is picked up there biologically by the mechanism ligand --> receptor, or based on the specific metabolic activity of the cells) and by its radiation destroys the tumor cells "from the inside" (Fig .3.6.8 on the right).
    The necessary selective accumulation of the radiopharmaceutical in the target tissue is achieved in two main ways :
--> By binding labeled ligands to receptors on the cell surface. Monoclonal antibodies such as ibritumomab tiuxetan (anti-CD20, labeled 90Y - Zevalin), tositumomab (anti-CD20, labeled 131I - Bexxar), tetraxetan-tetulomab (anti-CD37, labeled 177Lu - Lilotomab) are used for lymphoma therapy - below "Radioimmunotherapy of lymphomas". For neuroendocrine tumors, peptide somatostatin ligands (labeled 90Y, more recently 177Lu) are used - below "Neuroendocrine tumors". It appears to be the most important are ligands PSMA (labeled with 177Lu or 225Ac) in metastatic prostate cancer (below "Prostate cancer"). The labeled substance binds selectively in places with a high density of the respective receptors and is internalized after binding to the receptor.
--> Incorporation of the radiopharmaceutical into cells based on their specific metabolic activity. This is the case with the uptake of radioiodine 131I by cells of the thyroid gland, including metastases of the differentiated thyroid carcinoma.
    These mechanisms of selective incorporation into cells lead to the accumulation of the radionuclide in the tumor and the emission of biologically effective ionizing radiation.
--> Simple methods of direct local application of radiopharmaceuticals are marginally used, especially in joint cavities ("Radionuclide synovectomy"), but they do not have the character of biologically targeted therapy.
  We basically don't need to know the location of the tumor site, the radiopharmaceutical can be taken up even in sites we don't know about yet (eg micrometastases) - the radiopharmaceutical "finds" tumor foci itself. Thus, this targeted therapy to kill tumor cells uses radionuclides, which are bound to a suitable "transporter", whose task is to selectively transported to target tissues a sufficient amount of a radionuclide, so that the dose of emitted radiation killed tumor cells, while surrounding healthy tissues and organs should be damaged as little as possible - not be irreparably compromising functionality (in accordance with the above stated general "strategic goal" radiotherapy "Strategic goal and methods of radiotherapy").

Fig.3.6.8. Biologically targeted radionuclide therapy ("molecular radiotherapy").
Left: Comparison of the methodological approach of external radiotherapy and radioisotope therapy. Middle: Time dependence of the surviving fraction of cells in the tumor during radioisotope therapy. Right: The "crossfire" effect of hard radiation
b on tumor cells.

Biologically targeted radionuclide - radiopharmaceutical - therapy provides several advantages over other treatment approaches in oncology. Unlike classical radiotherapy ("Radiotherapy"), radiation is not applied to the body from the outside - the problem of accurately "hitting" known lesions and the impossibility of therapy for hidden lesions, radiation interference with healthy tissues - but is delivered through a systemic biochemical pathway, similar to chemotherapy or biological treatments ("Chemotherapy", "Biolog Therapy"). They can also be specifically picked up in unknown micrometastases or even in individual isolated cells (for example in the bloodstream). In contrast to chemical and biological therapy, carrier biochemical molecules do not act directly, but only transport radionuclide atoms, whose high cytotoxic effects after uptake are no longer dependent on the variability of signaling pathways (with event mutations).
    A significant advantage is the possibility to visualize the biodistribution of radionuclides - using gammagraphic methods to visualize the desired accumulation of the radiopharmaceutical in target tissues as well as the unwanted uptake in healthy tissues and organs. This possibility is perfectly utilized in the concept of theranostics
(described in more detail in §4.9, passage "Combination of diagnostics and therapy - theranostics"), combining diagnostic imaging with biologically targeted therapy. This analysis of diagnostics and therapy in an individual patient makes it possible to determine whether the respective therapy will be effective, even before it is started. And on the basis of dosimetric analysis of radiopharmaceutical distribution in target tissues and healthy critical tissues, determine the optimal dose - activity. This achieves a precise individualized approach to radionuclide therapy.
    Compared to other methods of systemic therapy, in indicated cases the biologically targeted radionuclide therapy shows good efficacy with minimal toxicity. Unfortunately, it is still not well known in the wider oncology community, so it is not used either at all or only as a modality of last choice when all other methods have failed. And that may already be a bit late, in patients "decimated" by previous failed treatment
(although even here there are cases of achieving almost complete remission in infaustuous patients in the terminal stage...).

Physical and biological factors
Similar to external radiotherapy, radionuclide therapy is achieved by co-production of two basic factors :
--> Physical factors
- type of radionuclide, type of emitted radiation (
a, b, g) and its energy, half-life.
    For therapy with open radionuclides, only radiation with low penetration (short range) is suitable, especially beta radiation (the range of which in the tissue is usually less than 5 mm), or Auger electrons (with a very short range of the order of nanometers), or alpha radiation (also a short range of tens of micrometers). The short range of this radiation in the tissue ensures that virtually all the energy is deposited in the target volume and the effect of the radiation is thus localized to the organ or area of tissue in which the radioactive substance has been taken up *), with minimal damage to surrounding healthy tissues.
*) However, radiation exposure to other tissues and organs may occur due to partial undesired uptake of the used radiopharmaceutical in these tissues and during metabolic processing and clearance of the radiopharmaceutical (blood, urinary tract or GIT).

Tab.3.6.1. Some radionuclides suitable for biologically targeted therapy (their properties and uses will be described below).
The range (missile) of radiation in the tissue depends on the type and energy of the respective quanta. For radiation
b, the maximum range is given by the maximum energy in the continuous spectrum; however, only a small percentage of electrons have this max. energy b. More important here is the mean range, which represents about 1/3 of the maximum range - it is given by the mean energy in the spectrum b (§1.2, part "Radioactivity beta"). For radiation a, if it is "monochromatic", there is practically no difference between the maximum and medium range (the difference is only when two or more alpha lines with significantly different energies are emitted).
The radionuclides are listed in the table according to the range of the respective radiobiologically most effective quantum, alpha or beta.

Therapeutic beta or alpha radionuclides should meet several criteria :
l The nuclide should have a high proportion of corpuscular beta or alpha emission and a low gamma component.
l The half-life should correspond to the biological kinetics of the target ligand. The effective half-life of the radiopharmaceutical, which is the result of the physical half-life of the nuclide and the half-life of the biological elimination of the ligand, then determines the duration of therapy (or its faction).
l If the daughter nuclide is also radioactive - the radionuclide decays by the conversion series ("In vivo generators" in nuclear medicine"), the decay chain should not contain intermediates with a long half-life (which could be released from the target lesion and cause radiation exposure to healthy tissues - more details is discussed below in the section "Alpha and beta radionuclides for therapy").
l Radionuclide atoms should be able to form stable compounds or conjugates with the necessary biomolecules - radioligands (Fig.3.6.10 a).
    The goal of radionuclide therapy is the inactivation and destruction of tumor cells by apoptosis induced by ionizing DNA double-strand breaks. A LET of about 100-200 keV/micrometer is required for a double-strand break caused by a single particle, which allows for alpha particles capable of producing dense ionization tracks. For the more sparsely ionizing beta radiation, is need a sufficiently high intensity (fluence) of quanta, causing rapid fiber breaks at various places in the DNA
(faster than the cellular repair mechanisms are sufficient to repair them).
--> Biological and radiobiological factors
- radiosensitivity of pathological cells in comparison with cells of healthy tissues and organs, pharmacokinetics of therapeutic radiopharmaceuticals (their uptake in target tissues and other tissues). The basic requirement is high accumulation in target tissues and low accumulation in healthy tissues.

    The biokinetics of therapeutic radiopharmaceuticals can be influenced pharmacologically to some extent
(eg by discontinuation of TSH or applications of thyrogen in the thyroid gland, or applications of rituximab in lymphomas).
 Internalization and externalization of radiopharmaceuticals in cells
The degree of accumulation of the radiopharmaceutical in the target tissue is certainly the basic parameter determining the success of the therapy. However, the effectiveness of the therapy also depends on the mechanisms of transport of radiopharmaceuticals into tumor cells - their internalization inside - and penetration outside these cells - externalization. Internalization of macromolecules is mediated by cell receptors or endocytosis of the cell membrane. When a ligand activates a receptor on the cell membrane that is associated with a G-protein on the inside of the membrane, the receptors are phosphorylated together with the ligands and internalized into the cell. First, it forms a so-called endosome, which then fuses with a lysosome and the ligand is broken down. The degradation products, including the radioactive atom, then remain in the lysosomes for a certain period of time and can act with ionizing radiation. For metabolic accumulation
(such as radioiodine in thyroid cells), the internalization time is determined by the rate of formation of the resulting metabolites and their excretion from the cells into the bloodstream.
    After a certain time, the atoms of the therapeutic radionuclide or its daughter isotope are excreted from the cells - externalization and enter the bloodstream. The further fate of the externalized radiopharmaceutical then depends on its (bio)chemical composition. It is generally excreted mainly by the kidneys, to a lesser extent by the GIT.
    Rapid internalization and delayed externalization are desirable for effective radionuclide therapy.
    Internalization and externalization has a specific meaning for radiopharmaceuticals labeled with alpha-radionuclides with a chain decay series
- it is discussed below in the passage "In vivo radionuclide generators".
 Physico-radiobiological effects
For large and heterogeneous tumor lesions, it is appropriate to use a radionuclide with high energy
b -radiation (such as 90Y with a maximum energy of 2.3 MeV and a range in the tissue of about 5 mm, which represents about 100-200 cell diameters) for the so-called "crossfire effect": this radiation can destroy even those tumor cells that are not in direct contact with the bound radiopharmaceutical (cells that do not have the appropriate receptors, or that the radiopharmaceutical does not penetrate them inside the tumor, perhaps due to hypoxia). These cells get into the "crossfire" of hard radiation from a radionuclide bound to surrounding cells (shown in Fig.3.6.8 on the right).
    However, for the eradication of smaller tumor foci, or tumors infiltrating normal tissues in a diffuse form, this effect could cause increased radiation exposure to surrounding healthy cells. Here, on the other hand, radionuclides with a lower energy of radiation beta, such as
177Lu with a maximum reach in the tissue of about 2 mm, or alpha-radionuclides are suitable. Shorter penetration of 177Lu radiation captured in tumor cells, into surrounding tissues may lead to a more favorable tumor/healthy tissue effect ratio.
    For cancer diseases involving larger and smaller tumors, is tested the so-called tandem therapy
90Y - 177Lu, or 177Lu - 225Ac, event. "cocktail" co-administered radiopharmaceuticals labeled with high- and low-energy beta radionuclides, or alpha radionuclide (cf. "Alpha and Beta Radionuclides for Therapy" below).
    The effect of radioisotope therapy may be further contributed by the radiation-induced biological bystander effect
(described in §5.2, section"Effect of radiation on cells", passage "Bystander-Abscopal effect"), thanks to which not only the cells directly affected by the radiation are damaged, but also some surrounding cells that have not been directly irradiated. Thus, some cells that did not take up the radiopharmaceutical, can be also eliminated. In late stages of radionuclide therapy, when radiation doses are already low, biological efficacy of irradiation may increase the effect of hyper-radiosensitivity to low doses of radiation (§5.2, part "LQ model" passage "Deviation from the LQ model", Fig.5.2.4c). This reduces the proportion of "unnecessary wasted dose" in the later stages of therapy.

Fig.3.6.9 Schematic comparison of the effects of beta and alpha radiation on DNA.
Beta electrons with a low ionization density cause mostly simple breaks in DNA that the cell can repair. A number of simple fractures are required to induce apoptosis.
Alpha particles with a high ionization density cause double breaks in DNA, which usually result in cell death by apoptosis.

Alpha and beta radionuclides for therapy
So far, beta-radionuclides are mainly used for radioisotope therapy (middle and lower part Tab.3.6.1) - discussed belowe. Recently, however, also alpha radionuclides have been increasingly used here (upper part of Table 3.6.1) , whose radiation has a high LET - high ionization density - to create double DNA breaks, which leads to high radiobiological efficiency of cell killing. In the case of alpha radiopharmaceuticals, while doing so, high radiation energy is released in a very small volume, which leads to a lower radiation exposure of the surrounding tissues. Comparing the a- and b- emitting radionuclides in terms of use for biologically targeted therapy, we can emphasize the following differences :
The weight of a -particles is about 7000 times greater than b -particles (electrons).
Energy of the a particles is »10-30 times greater than the b particles : alpha typically 4-8 MeV, beta approx 0.2-2.25 MeV.
The electric charge of a -particles is 2 times larger than b -particles (alpha: +2, beta: -1 of elementary charge |e|) .
Ionization density (linear energy transfer LET) of a-particles is about 100 times larger than b -particles. For alpha particles with energies of 4-8 MeV, the LET in the tissue is about 100 keV/micrometer, at the end of the path in the Bragg maximum it can increase locally up to 300 keV/micrometer. For beta particles with typical energies of hundreds of keV, LET is only about 0.2 keV/micrometer.
¨ Effective range of a particles in tissue is substantially shorter than b -particles. In alpha the range is about 2-5 cell diameters, in beta hundred cell diameters.
    This comparison shows that alpha radionuclides have locally higher radiobiological efficacy than beta, but due to the short range, in
a -radiation almost not occurs the "crossfire effect". However, low penetration (short range in the tissue, approx. 50-90 mm) and high ionization density (LET tens to hundreds of keV/mm) allow effective destruction of tumor cells with minimal collateral damage to the surrounding healthy tissue.
    From the point of view of radionuclide therapy, alpha and beta radionuclides are thus largely complementary to each other :
--> Beta particles with a typical energy of approx. 0.5-2.3 MeV and a low LET of ~0.2 keV/micrometer have a longer path, micrometers to centimeters, representing approx. 5-150 cell diameters - they are more suitable for larger tumors, diffuse or residual disease. Each beta electron is able to fly through and hit many cells in which it causes ionization damage (cf. the "crossfire" effect). But in the case of micro-lesions or isolated cells, the efficiency is low, most of the beta particles fly away outside the lesion, without effect.
--> While alpha particles (with energy ~4-9MeV and high LET ~100-300 keV/micrometer), due to their short path length in the tissue of about 40-100 micrometers, which is only ~1-3 cell diameters, can be effective even for small lesions and millimeter-sized micrometastases, or even for isolated solitary tumor cells. If alpha particles get into them, or preferably when an alpha-emitting radionuclide is internalized inside a cell, they can cause irreparable DNA breaks with high ionization density.
Radiobiological effects of reflected nuclei 
During the emission of alpha particles - heavy helium nuclei with high kinetic energy - occurs due to action and reaction, a back reflection daughter nuclei with kinetic energy of about 100keV
(§1.2 passage "Backward reflection of nuclei"). The nuclei reflected in this way brake very quickly in the tissue on a path of about 500 nm, along which they cause dense ionization of the substance with a high LET (of the order of hundreds of keV/mm). If this occurs inside the cell, it can cause double DNA breaks or damage to the mitochondria, which can result in apoptosis. The reflected nuclei thus contribute somewhat radiobiological effect, which, however, is located here in the very close vicinity of the alpha-decay site, only 0.5 micrometers.

In vivo radionuclide generators
Most alpha-radionuclides used in nuclear medicine are converted by the whole decay series
(§1.4, passage "Decay series") and after their application to the organism they behave as "in vivo radionuclide generators" ("In vivo generators" in nuclear medicine") - (Fig.3.6.10b,c). The advantage here is the emission of several alpha-particles (typically 4 alpha/decay, see eg 227 Th - where is 5 alpha) with high energies of about 4-8 MeV, which leads to high radiobiological efficiency.
    A certain problem with in vivo generators is the redistribution of daughter radionuclides - release - dissociation - of daughter atoms from chemical bonding in radiopharmaceutical molecules due nuclear back reflection during alpha-particle emission (§1.2, passage "Nuclear reflection") and differences in chemical properties of daughter atom (oxidation number). Already at the first a-decay, the chelation of the radionuclide with the ligand is disrupted by the back reflection of the nucleus (and thus the daughter atom) and the chemical transformation of the original atom. Subsequent a -emitting daughter atoms are then free. When this occurs on the cell surface, the lymphatic and bloodstream can transmit the radioactivity thus released to non-tumor tissues (Fig .3.6.10 b). If the daughter atom escapes from the target tissue during its transformation, the effectiveness of the therapy is reduced by a few subsequent energy alpha-particles from the decay series. These released radioactive atoms can then migrate and be taken up in other tissues (eg bone marrow) and cause unwanted radiation exposure there. This side effect will be alleviated, if the radiolabeled ligand penetrates inside the cells quickly enough and is internalized therein (Fig.3.6.10 c); then nascent charged daughter atoms, which are highly reactive, can remain bound in the cytoplasm inside the cells long enough to be able to repeatedly decay there and transfer all their radiation energy. This cells internalization of the daughter radionuclides in the generator in vivo can also reduce or prevent undesired redistribution of radioactivity outside the tumor tissue.

Fig.3.6.10. Radiolabeled biomolecules as radiopharmaceuticals for diagnostics and therapy in nuclear medicine.
a) Chemical binding of a radionuclide to a biomolecule (eg a monoclonal antibody) using a bifunctional chelating molecule (eg DOTA). b) Release - dissociation - of the daughter atom from the chelating bond as a result of the backscattering of the nucleus and transformation of the oxidation number of the daughter atom during radioactive transformation (eg alpha). When a radionuclide is bound to the cell surface, the released daughter radionuclide may migrate by metabolism to healthy tissues (and cause unwanted radiation exposure there) . d) Prevention of this leakage by internalization of the radiopharmaceutical inside the cell.

For these reasons, it is also desirable that the decay chain of the radionuclide used does not contain daughter intermediates with a longer half-life, that could be released from the target lesion and cause radiation exposure to healthy tissues. E.g. actinium 225 Ac (225Ac(10d.; a) ® 221Fr(4,8m.; a) ® 217At(32ms.; a) ® 213Bi(46m.; b-) ® 213Po(4ms.; a) ® 209Pb(3,3h.; b-) ® 209Bi(stab.)) is more advantageous in this respect than the otherwise promising thorium 227 Th (227Th(18,7d.; a) ® 223Ra(11,4d.; a) ® 219Rn(4s.; a) ® 215Po(1,8ms.; a) ® 211Pb(36,1min.; b-) ® 211Bi(2,2min.; a) ® 207Tl(4,8min.; b-) ® 207Pb(stab.)), where 223U is problematic with a half-life only slightly shorter than the basic 227Th.
    In general, however, all radiopharmaceuticals, including daughter radionuclides in an in vivo generator, are taken up not only in tumor cells, but more or less also in other tissues, where they cause undesired radiation exposure...

"Tandem" therapy with alpha + beta radionuclides
During radionuclide therapy, alpha and beta radiation can complement each other with their biological effects - work "synergistically in tandem cooperation" :
-> Densely ionizing alpha radiation can also be effective against micrometastases or even solitary tumor cells .
-> Beta radiation, due to the "crossfire" effect, can be more effective in larger and heterogeneous tumor lesions .
    For example, tandem therapy with
225Ac-PSMA and 177Lu-PSMA is being tried in prostate cancer.
Radionuclide therapy: beta- yes, beta+ no !
b+ have essentially the same radiobiological effects as electrons b- of the same energy. However, positrons are annihilated with electrons after braking in the tissue, which generates a penetrating annihilation g- radiation with an energy of 511 keV, which places a strong radiation load on surrounding and more distant tissues and organs. Positron b+ -radionuclides are therefore unsuitable and unusable for therapy !
Radionuclide therapy with Auger electrons
In addition to radiation
a, b, g, some radionuclides emit also conversion and Auger electrons (+ Coster-Kronig electrons). They are formed by the internal conversion of (virtual) photons of g -radiation and characteristic X-rays inside the atomic shell (§1.2, passage "Internal conversion of gamma and X-rays"). In contrast to the continuous spectrum of electrons b- radiation from the nucleus, the conversion and Auger electrons have a discrete spectrum consisting of several monoenergetic lines. Auger electrons have significantly lower energies than b- rays, usually keV units or less. Auger electron emitting radionuclides have the following specific properties in terms of biologically targeted therapy :
× It is emitted several (approx. 5-20) electrons per decay.
Very short range of Auger electron - in water or tissue it is of the order of nanometers .
× Their LET is locally »> 20 times higher than beta (energy »100keV).
In the immediate vicinity of »5 nm from the radionuclide, a locally high micro-dose arises from Auger electrons (according to a fictitious theoretical conversion » 104-107 Gy).
    These properties result in locally high radiobiological (cytotoxic) efficiency killing of cells in the range of a few nanometers, which, however, can only be applied if the Auger electron-emitting radionuclide decays directly inside the DNA molecule of the cell, in particular between the DNA strands. So far, two ways to bind the appropriate radionuclide to DNA molecules in vivo are being tested : 1. Labeling of pyrimidine derivatives that are incorporated into DNA in vivo. 2 . Labeling of DNA dyes that bind to DNA via hydrogen bonds.
    Intensive sources of Auger electrons are mainly radionuclides decaying by electron capture, such as
125I, 77Br, 123I, 124I, 111In, 67Ga, 201Tl, ...; they are also emitted by some "pure", metastable g- radionuclides (including 99mTc). In the case of some therapeutic beta- - radionuclides, they are produced by the internal conversion of the accompanying gamma-radiation. One such radionuclide with a high proportion of Auger electrons is terbium 161Tb.
 Effect of gamma radiation
Radiation gamma in radionuclide therapy have virtually no therapeutic effect, and its presence may even cause undesired irradiation of other organs and tissues than the target lesion. However, in the case of mixed beta-gamma emitters, gamma radiation can be advantageously used for scintigraphic imaging of the biodistribution of the radiopharmaceutical in the organism (
Chapter 4 "Radioisotope scintigraphy" ) and for dosimetric monitoring of the course of therapy (see below "Planning, monitoring and dosimetry of radionuclide therapy"). Such mixed therapeutic radionuclides with a useful component of radiation g are mainly 131I [g 364keV (81%)], then 153Sm [g 70keV (5%) and 103 keV (28%)], 186Re [g 137keV (9%) ], 177Lu [g 113keV (3%) and 208keV (6%)], 166Ho [g 48-58keV (9%) and 81keV (6%)], or 225Ac [g line 100keV(1.6%), 218keV(5%), 440keV(26%)], and several others. Detailed gamma spectra of these and other radionuclides are given in §1.4, section "The most important radionuclides".
    For pure higher energy radionuclides
b, such as 90Y, braking radiation may be used for detection (gammagraphy), but the accuracy of localization and determination of organ activity is significantly worse (for yttrium-90, however, there is a possibility to use very weak annihilation radiation for PET scintigraphy with better resolution than braking radiation - is analyzed in §1.4, passage "Ytrium 90 Y") .
    Note: It is worth noting that the basic requirement of "as much beta as possible, as little gamma as possible" for radionuclides for therapy is exactly the opposite to radionuclides for diagnostics (scintigraphy), where the main component must be gamma radiation, while beta radiation, which increases the radiation exposure, should be contained as little or not at all (as is ideal for
99mTc) - see Chapter 4. "Radioisotope scintigraphy".

Application of a therapeutic radiopharmaceutical
 Examination before application
Before the actual application, it is necessary to make sure whether the tumor cells or its metastases will accumulate the used radiopharmaceutical. Preliminary basic information provides clinical results - what type of tumor it is
(with histological data), what receptors the tumor cells should express. Actual confirmation of radiopharmaceutical accumulation (or expression of the necessary receptors) by tumor cells is performed using scintigraphic examination when applying the diagnostic activity of the relevant radiopharmaceutical. The accumulation of the radiopharmaceutical in the tumor tissue must be higher than the accumulation in the surrounding healthy tissues and other organs, mostly in the liver (the so-called Krenning score of 1-4 is sometimes used to assess selective uptake, it should be higher than 2 for successful therapy).
    Furthermore, it is necessary to take into account contraindications (such as pregnancy or lactation), risk circumstances such as impairment of kidney function (creatinine clearance <40 ml./min.), liver, hematopoiesis.
Therapeutic radiopharmaceuticals are usually administered intravenously, sometimes orally
(for thyroid cancer therapy), then into cavities and joint capsules (accumulation scintigraphy is not performed here, as these mechanisms are not present here). Carriers ("transporters") in radiopharmaceuticals have a wide chemical range - from simple inorganic compounds (such as chlorides or iodides), through more complex organic substances (.., peptides,...), to very complex labeled monoclonal antibodies. Intravenous application of therapeutic radiopharmaceuticals is usually performed by slow infusion over a period of approximately 10 - 30 minutes.
    The optimal dosage of activity [MBq] of a therapeutic radionuclide is based on a compromise between the maximum required radiation effect in the target tissue and the radiotoxicity of the preparation, which is often inadvertently taken up and radiatively burdenes also other tissues and organs. Based on thorough laboratory and clinical trials, the recommended applied activity [MBq] is determined for each therapeutic radiopharmaceutical, usually calculated on the patient's weight
(or on the body surface area - an empirical formula from the patient's height and weight). Individual refinement can be performed on the basis of an MIRD analysis (Determination of radiation dose from internal contamination. MIRD method.).
   Figure 5.5.1 from §5.5 we shown here again for clarity :

Fig.5.5.1. Radiation doses from the distribution of radioactivity inside the organism. Left: Source and target organs in the body. Middle: Time dependence of activity in source organs. Right: Time dependence for determination of doses by MIRD method.
< E> is the mean energy [J] deposited in the tissue of weight m per one decay of the used radionuclide
(beta electrons or alpha particles are considered). If this mean energy is given in nuclear units [eV], there a conversion factor of 10-19 is also used.
The figure roughly simulates the situation after the penetration of radioiodine 131I into the organism. Radioiodine is rapidly taken up in the thyroid gland, then metabolized and excreted by the kidneys into the bladder, from where it periodically leaves the body during urination.

Radionuclide therapy is performed either once or in several cycles, usually 4, with an interval of 4-8 weeks. In addition to the radiopharmaceutical itself, other auxiliary substances such as antihistamines against allergies and amino acids to protect renal toxicity are often administered.
    For application to the joints
(radionuclide synovectomy - see below), the values of the recommended applied activity are derived from the size of the joint.

Radiation dose and its distribution in radionuclide therapy
In radionuclide therapy, the distribution of the dose, as well as its time variability, in tumor and healthy tissue is usually more complex than with external beam radiotherapy. This is due to time-varying biological processes of uptake (accumulation) and leaching (clearance-loss) of radioactive substances in different types of tissues; the physical half-life of the radionuclide used also contributes to this. The basic relations for the radiation dose from the homogeneous distribution of radioactivity in the substance were derived in §5.1, passage "
Radiation dose from radioactivity"; here we will modify them appropriately.
    The radiopharmaceutical with A
inj activity is usually rapidly (within a few hours) partially absorbed after administration in the target tissues; the rest leaves the body mainly through the urinary tract. In the simplest case, in the target deposit of mass m the activity Ao = a .Ainj , given the accumulation capacity a of the given tissue, accumulated evenly (accumulation is often expressed in %). This activity Ao causes with its emitted particles in the given deposit a dose rate Ro [Gy/s] = Ao . <E>. 6.10-12 /m, where <E> [MeV] is the mean energy of the short-range particles (mostly beta, resp. alpha), which is absorbed in the investigated deposit (coefficient 6.10-9 is the energy conversion factor between MeV ® Joule units, including also the weight conversion g ® kg). Then, this accumulated activity will decrease approximately exponentially with time t : A(t) = Ao .e - k .t with an effective rate k = ln2/(T1/2phys + ln2/T1/2biol), given by the physical half-life T1/2phys of the used radionuclide and the biological half-life clearance of radiopharmaceutical T1/2biol from tissue. The dose rate in the lesion will decrease at the same rate. The total cumulative dose D received in the target tissue after the time T has elapsed is then D(T) = 0nTR(t) dt = (Ro/k).[1- e-k.T]. This radiation dose, together with its time dependence, can then be substitute to a linear-quadratic model with a time factor of reparation l and repopulation T2r , as derived above ("LQ model") :
- ln(N/N
o) = a.D + {2.[(1-e-l.T).(1-1/l.T)]/l.T}.b.D2 - ln2.T/T2r .
In the general case, a complex equation arises for the surviving fraction of N/N
o cells, which, however, assuming an irradiation time long compared to the effective half-life of the radioactivity in the target volume (and neglecting cell proliferation) is simplified to: -ln(N/No) = a.D.{1 + Ro/[(l+k).a/b]}]. From a comparison with the corresponding formula for external fractionated radiotherapy, it can be seen that the Ro/l ratio here has a similar role as the fractionation dose d.
    In radionuclide therapy, an important parameter is the resulting radiation dose - whole body, therapeutic dose in target lesions, unwanted dose in healthy (critical) tissues and organs. If we measure the time course of activity A(t) in a certain tissue (organ, lesion) of mass m, we can determine the total cumulative radiation dose using the relation
            D = 0
nAA(t) dt . <E>. 6.10-19/m   , or   D = AS . <E> .6.10-19/m   ,
where A
S is the total so-called cumulative activity (introduced in §5.5, "Internal contamination") in the examined volume during the entire time since application (t=0-A). We will deal with this further down in the passage "Planning, monitoring and dosimetry of radionuclide therapy", Fig.3.6.11.
    During the exposure, that is continuous with decreasing dose rate, in addition to radiation destruction of cells, proliferation (repopulation) of tumor cells in the lesion may also occur, especially in the later stages of therapy. As long as the dose rate is higher than the critical value of ln2/(
a.T2r), the number of cells in the tumor deposit will decrease, later when radiation declines, tumor cell proliferation may predominate - compare with the above-mentioned "wasted" dose in permanent brachytherapy. It is therefore desirable to apply such a high activity of radiopharmaceuticals, that a high dose rate from the accumulate radioactivity in the tumor tissue will rapidly kill, if possible, all clonogenic cells even before cell repopulation predominates. In this respect, however, radiotoxicity is a frequent obstacle for those healthy tissues and organs, in which the radiopharmaceutical is also inadvertently taken up...
    Dose escalation - the applied activity - is permissible only with careful dosimetric control and radiation monitoring of radiopharmaceutical uptake in target deposits and unwanted uptake in healthy tissues and critical organs - "
Determination of radiation dose from internal contamination. MIRD method.". From a clinical point of view, it is necessary to take into account the patient's general state of health and possible risk factors (such as disorders of blood formation or kidney function). Although the MIRD method in all its complexity is not yet applicable in routine therapy, at least whole body dosimetry (by measurement of the dose rate at a distance of about 1-2 m from the patient's body, see below) and laboratory hematological analysis of collected blood samples to monitor for adverse hematological toxicity, should be performed (in more serious cases of hematological toxicity, it is necessary to proceed with subsequent autologous blood cell transplantation). The role of dosimetry in the optimization of radionuclide therapy is discussed below in the paragraph "Planning, monitoring and dosimetry of radionuclide therapy".
"Waste" radioactivity 
During the entire course of radioisotope therapy, the radioactive substance used leaves the organism, mainly through the urinary tract. The highest content of radioactivity in urine is at the beginning of therapy, when a significant amount of free unbound radioactive substance is excreted. Later, during the destruction of the tumor tissue, the originally bound radiopharmaceutical is released, which has already "fulfilled its role". The relatively high content of radioactivity in the urine (to a lesser extent also in other excrements) in patients during radionuclide therapy should be taken into account from the point of view of radiation protection - handling of radioactive waste (see §5.6 "
Radiation protection in workplaces with ionizing radiation").
Note: Please do not confuse the excreted "radioactivity of the waste" with the above-mentioned radiobiologically "unnecessary" or "wasted" dose !
Planning, monitoring and dosimetry of radionuclide therapy
The methods of dosimetric planning and verification described above for external beam radiotherapy (section "Planning for radiotherapy") are not applicable to radionuclide therapy. The situation here is more like chemotherapy or biological treatment (described above in the section "Chemotherapy and biological treatment"). However, we do not have to do radionuclide therapy completely "blindly" (empirically, at a flat rate), as is the case with pharmacological treatment. Methods of the detection and imaging of emitted ionizing radiation offer certain possibilities for monitoring, dosimetric measurements and individual dosing in targeted therapy with open radionuclide emitters. Determination of radiation doses can be performed by measuring biokinetics - the rate of accumulation and subsequent excretion of the used radiopharmaceutical in defined areas of interest of tumor lesions and healthy tissues and critical organs. It is possible to use mainly quantitative scintigraphy on gamma cameras (planar, SPECT or PET, in combination with CT), in a simpler case, whole-body measurement of radiopharmaceutical retention, as well as measurement of the activity of blood samples.
    For external dosimetric monitoring of radionuclide therapy it is advantageous, if the radionuclide used has, in addition to the main component
b or a, also a not neglibigle component of the radiation g - this is the case, for example, with radioiodine 131 I, samarium 153 Sm, lutetium 177 Lu or actinium 225 Ac. This radiation g freely (with some absorption) passes through the tissues and by its external detection by scintigraphy can determine the location and in principle the activity in tissues and organs. For pure higher energy radionuclides b, such as 90 Y, braking radiation may be used for detection (gammagraphy), but the accuracy of localization and determination of organ activity is significantly worse (for yttrium-90, however, there is a possibility to use very weak annihilation radiation for PET scintigraphy with somewhat better resolution than braking radiation - is analyzed in §1.4, passage "Ytrium 90 Y").
    In principle, three procedures can be used for dosimetric monitoring in radionuclide therapy :
-> Whole body dosimetry
during which it is determined what total - whole body - radiation dose the patient will receive as a result of the application of a certain activity of the used radiopharmaceutical. And possibly what are its time dynamics. It can be performed for both diagnostic and therapeutic applications if necessary. It is most simply performed using a gamma-radiometer (GM or scintillation detector, dose rate meter) located approximately 1-2 meters from the center of the patient's body. A more recent method is using a gamma camera in whole-body scanning mode, where we also get an image of the distribution of the radiopharmaceutical in the target sites and unwanted uptake in healthy tissues. The conversion factor for converting the measured number of pulses "cps" to the activity "MBq" is determined either using a phantom measurement with the standard of the radionuclide used, or more simply using the first measured value at time t=0, which corresponds to the actually applied activity A
WB(t=0). This coefficient is then used even to convert the measured values nto activity in the patient's body during further measurements.
    Values from repeated measurements over several days are usually fitted with a biexponential function, modeling the initial faster and later slower decline of total radioactivity in the organism :
WB(t) = A0 . ( e-l1 . t + e-l2 . t )    ,
where A
0 = AWB(t=0) is the initial applied activity (Fig.3.6.11 left). This creates a curve of instantaneous whole-body activity AWB(t) for time t, whose integration from 0 to A yields the value of cumulative total activity ASWB = 0nAAWB(t) dt in the patient's body. The absorbed whole-body dose of DWB is then directly determined as the product of this cumulative ASWB activity in the body and the mean energy <E> emitted per 1 decay of the used radionuclide, converted to 1 kg (with relevant coefficients): DWB[Gy] = ASWB.<E>.6.10-12/MWB, where MWB is the patient's weight. Alternatively, using MIRD analysis, it is determined as a product of this cumulative activity in the target volume and the S-factor: DWB[Gy] = ASWB.SWB. This S-factor expresses the mean absorbed dose per unit of cumulative activity. For the target volume of the whole body for the most frequently used radionuclide 131I, an approximate empirical formula was determined for the S-factor: SWB[Gy/MBq.h] = 1.3.10-4 MWB-0.92, where MWB is the patient's weight in kg.
    Determination of whole-body radiation dose during radionuclide therapy can be useful mainly in three directions :
- Verification of whether the correct value of the indicated therapeutic activity of the given radiopharmaceutical was successfully applied. However, it may be more useful to use whole-body dosimetry "in the opposite direction": Based on the known value of the applied activity AWB(t=0), perform a quantitative calibration of scintigraphy - determination of the conversion factor [cps/MBq] between the activity and the measured number of pulses [cps] in the image.
- From the point of view of radiation protection, the whole-body dose is the basic quantity determining the patient's radiation load and the level of risk of stochastic and possibly deterministic effects.
- For some types of therapy (thyroid glands with 131I), a maximum whole-body dose (mostly up to 2 Gy) has been established, which must be observed.

Fig. 3.6.11. Dosimetric measurements for radionuclide therapy.
Left: Whole-body A
WB activity and resulting whole-body DWB dose it caused. Middle: Accumulated radioactivity in the target deposit (tumor - marked ROI) ATum and the radiation dose DTum induced by it. Right: Specific radioactivity ABlood in the blood - in 1ml samples taken.
Illustrative values of time t[days] on the horizontal axis and activity A[MBq] on the vertical axis roughly correspond to the situation during the therapy of metastatic ca thyroid gland with radioiodine 131I .

--> Dosimetry of target tissues (tumor lesions) and healthy (critical) organs
It is used to determine the absorbed therapeutic radiation dose D
Tum in the target tissue (tumor lesion "Tum" or other pathological tissue), or what unwanted dose healthy tissues and critical organs are burdened with. This is the most important thing we need to know for the successful course of radionuclide therapy. Dosimetry of lesions is suitable for predicting response to treatment and for determining the desired therapeutic activity, taking into account limitations arising from radiotoxicity. However, the methodological procedure here is quite demanding and has its limitations. For smaller and weakly accumulating lesions, it is usually not possible to determine the relevant results of the radiation dose...
    The measurement is performed with a gamma camera using quantitative scintigraphy
(§4.1 "Essence and types of scintigraphy", passage "Quantitative scintigraphy"). The camera must be calibrated using phantom measurements to convert the displayed number of "cps" impulses to "MBq" activity in the lesion. In a simplified way, this calibration can be performed with the help of whole-body dosimetry - the first measured whole-body image at time t=0, which corresponds to the actually applied AWB(t=0) activity. A scintigraphic imaging of the area in which the examined lesion is located is performed, the region of interest (ROI - Fig.3.6.11 in the middle) of the displayed lesion and the area of the tissue background is drawn on the image; this background is subtracted. Based on the calibration coefficient, the impulses from the ROI are then converted to activity. It is measured repeatedly, on the day of application and then on subsequent days, according to the half-life and pharmacokinetics of the radiopharmaceutical used.
    Radioactivity values from repeated scintigraphic measurements over several days are usually fitted with a biexponential function, modeling the initial accumulation of the radiopharmaceutical in the lesion and the later slower excretion - decrease - of radioactivity in the lesion
(2-compartment model of accumulation and excretion of the radiopharmaceutical) :
Tum(t) = [lTumAcc /(lBloodExc - lTumExc)].AApl . (e-lTumAcc . t - e-lBloodExc . t )   ,
where A
Tum(t) is the activity in the lesion at time t, AApl is the applied activity, lTumAcc is the rate coefficient of uptake (accumulation) of the radiopharmaceutical in the tumor lesion from the blood, lBloodExc is the rate coefficient of the degradation of the radiopharmaceutical from the blood, and lTumExc is the rate coefficient of the excretion of the radiopharmaceutical from the tumor lesion. From the bloodstream, the radiopharmaceutical is partially absorbed in the tumor lesion, from where it is continuously washed back into the bloodstream, and at the same time it is partially absorbed again from the blood in the tumor site; however, most of the radiopharmaceutical is constantly excreted from the circulation, mainly through the urinary tract. The result of this dynamic balance between uptake and excretion of the radiopharmaceutical from the tumor site is the time dependence of the activity in the lesion of a characteristic shape according to the curve in Fig. 3.6.11 in the middle: fast increase - rounded peak - slower decrease. Relevant speed parameters l are obtained from the biexponential interpolation-fit of the data obtained by measuring ATum activity at individual times t - Fig.3.6.11 middle.
    This results in a curve of instantaneous activity A
Tum(t) in the lesion for time t, the integration of which from 0 to A yields the value of the total cumulative activity in the given lesion ASTum = 0nAATum(t) dt. The resulting absorbed dose of DTum in the lesion is determined as the product of this cumulative activity of ASTum in the target volume and the mean energy <E> emitted per 1 decay of the used radionuclide (the emission of short-range beta electrons or alpha particles is assumed), converted per unit mass of the lesion (with the appropriate coefficients) : DTum[Gy] = ASTum.<E>.6.10-12/mTum, where mTum is the mass of the examined lesion - tumor (usually determined from the sonographically determined volume multiplied by the density of the lesion, usually 1g/1ml).
  Dosimetry of other, healthy tissues and organs can be performed in an analogous mann