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4.
Radionuclide scintigraphy
4.1.
The essence and methods of scintigraphy.
4.2. Scintillation cameras
4.3. Tomographic scintigraphy
4.4. Gated scintigraphy
4.5. Physical parameters of scintigraphy - image
quality and phantom measurements
4.6. Relationship between scintigraphy and other
imaging methods
4.7. Mathematical analysis and computer
evaluation in nuclear medicine
4.8.
Radionuclides
and radiopharmaceuticals for scintigraphy
4.9.
Clinical
scintigraphic diagnostics in nuclear medicine
4.1. The essence and methods of scintigraphy
Radionuclides in nuclear medicine
Nuclear medicine
is a field dealing with diagnostics and therapy using open
radioactive substances - radiopharmaceuticals -
applied to the internal environment of the organism; these in
vivo methods will be addressed in this chapter. In
an in vitro test , the
radiopharmaceutical is not administered to the patient's body,
but is used in the radiochemical analysis of blood samples taken;
the patient does not come into contact with a radioactive
substance, only a sample of plasma or other body fluid is used ( in vitro radioisotope methods are briefly
outlined in §3.5 " Radioisotope
tracking methods ", passage
"Radioimmunoassay - radiosaturation analysis "; now they are mostly not part of nuclear
medicine, but laboratory biochemistry).
Nuclear medicine methods are based on two basic
properties of radionuclides :
1.
Emissions of
penetrating ionizing radiation during radioactive
transformations of nuclei(detailed physical
explanation in §1.2" Radioactivity " );
2. Identical chemical behavior isotopes Þ
radioactive isotopes react chemically as stable isotopes of the
same element (§3.5 "Radioisotope tracer method").
Radioactive atoms and their
molecules - compounds "labeled" with radioactive
elements - are distributed in the body as if they were
non-radioactive, but penetrating radiation is continuously
emitted during the radioactive transformation of the respective
nuclei. This radiation allows them to be "made visible"
- to monitor, indicate, "trace" *) - and measure their
amount by detection devices during diagnosis ,
or by their radiobiological effects this radiation can be used
for therapeutic purposes .
*) From the generic name of
indicators or tracers methods that are
used only with the help of radionuclides and not only in medicine
but also in laboratory and industrial applications (§3.5 "Radioisotope tracking
methods ") .
Scintigraphic diagnostics and radionuclide
therapy in nuclear medicine
The central method of nuclear medicine is radioisotope
diagnostics in vivo : we apply a suitable (bio) chemical
substance with a bound radionuclide - the so-called radioindicator
or radiopharmaceutical - to the organism. This
substance enters the metabolism and is distributed
in the body according to its chemical composition - pharmacokinetics
- of the given radioindicator. Physiologically or pathologically,
it accumulates in certain tissues and organs, regroups and is
subsequently excreted. Radiopharmaceuticals chemical composition
determines its integration into kinetic or certain metabolic
processes - targeted input ( targeting)
into relevant tissues, organs, cells or sub-cellular elements,
including subsequent secretion. The built-in radionuclide then
allows either external detection and imaging of the
distribution of this substance ( g radiation in scintigraphy)
or monitoring of its amount in samples taken (biological fluids, mostly blood or urine) - specific methods of these examination methods are
described in detail below in " Clinical
scintigraphic diagnostics in nuclear medicine ".
In the case of therapy
, radionuclide radiation performs biological effects
on the cells of the tissue in which the radiopharmaceutical
accumulates (eg destroys tumor cells -
§6.6 " Radiotherapy", part" Radioisotope therapy ") .
Radioindicators
in nuclear medicine are applied in a small trace amount
, about 10 -9 -10 -12 grams (pico- or nanomolar
concentrations in tissues) , so they can
not chemically affect the function of the examined
organs, nor can they cause some side or toxic effects on the
organism *) They can only cause radiation exposure
, which we try to minimize by optimizing the applied activities.
*) The only exception to
this biochemical safety are radiopharmaceuticals based on murine
monoclonal antibodies.. In a small percentage of
patients, they may experience allergic reactions
due to the presence of so-called HAMA antibodies (discussed below in the section " Radionuclides
and radiopharmaceuticals for scintigraphy ").
The best known
example is the application of radioactive sodium iodide NaI 131 , which, like any
iodine, is taken up (accumulated) in the thyroid gland. By
external detection of gamma radiation emitted during radioactive b- transformations
of 131 I
nuclei, it is then possible to measure the accumulation of this
iodine or to display its distribution in the thyroid gland - §4.9.1 " Thyrological
radioisotope diagnostics
"; if necessary radiation bmay have biological effects
on the cells used in therapy when higher activities are applied .
A number of types of radiopharmaceuticals
with affinity for the kidney, liver, bone, myocardium, some tumor
or inflammatory tissues, signaling receptors have been developed
, for the function of which the substance is an indicator
(§4.8 " Radionuclides and
radiopharmaceuticals for scintigraphy ") . The degree of local
accumulation of radiopharmaceuticals depends on the intensity of
local metabolic and functional processes in organs and tissues.
Possible malfunctions can be located, analyzed and possibly with
the help of scintigraphic display. and quantify.
Or the radionuclide is injected into the
bloodstream and the dynamics of its passage
are monitored heart, lungs and large vessels - in this
case without metabolic binding to a specific organ or tissue (§4.9.4, part " Dynamic radiocardiography " and " Radionuclide gated ventriculography ", or §4.9.8, part " Perfusion scintigraphy of the brain ") ; again with the
possibility of analysis and quantification.
"Molecular
imaging"
With the development of organic chemistry,
biochemistry and cell biology, some radiopharmaceuticals have
been developed whose labeled molecules have affinity for very
specific cell types or processes at the subcellular level. With
the help of scintigraphy and a suitable radiopharmaceutical, it
is possible to purposefully examine not only the function of a
certain organ or tissue, but also to selectively affect a certain
type of metabolic or transport pathway, such as enzyme or
receptor binding or antigen-antibody reactions. For this purpose,
special radiopharmaceuticals (both for diagnostics and for
therapy) have been developed and are still being developed, which
are characterized by their effects at the molecular level
. With a bit of exaggeration, these methods of local measurement
and imaging of the physiological response are referred to as
" in vivo biochemistry ".
Note: Of course, the name " molecular
imaging " does not mean that we are perhaps
imaging the molecules themselves (unfortunately we cannot do
that), but we are depicting a distribution of the radiolabel that
is a consequence and reflection of specific biochemical
reactions at the molecular level.
Scintigraphy
The passage and distribution of a radioindicator thus reflects a
specific physiological or pathological condition
or function of the relevant organs and tissues.
For its assessment in the simplest cases, it is sufficient to
simply measure the intensity of radiation g emanating from a
certain place (eg from the thyroid gland -
to determine its accumulation) by a collimated detection probe. For
more comprehensive diagnostics, however, we need to measure - map
- display - the entire distribution of
the radio indicator , including local details and
anomalies. A method called scintigraphy or gammagraphy
is used for this :
Scintigraphy: |
Scintigraphy or gammagraphy is a physical-electronic method of imaging the distribution of a radioindicator in an organism based on external detection of outgoing gamma radiation |
Terminological note:
The more apt name of gammagraphy
- gamma-ray imaging - is unfortunately used relatively rarely;
the less precise name of scintigraphy
came from the fact that scintillation detectors are now
technically used here. In the future, scintillation detectors are
likely to be replaced by semiconductor detectors (see below "
Alternative physical and technical principles of gamma cameras "), with the name
"scintigraphy" losing its justification. Out of
inertia, however, it will undoubtedly persist.
Scintigraphy or scintigraphic examination
is often also called a scintigraphic study in
the "gimmick" of nuclear medicine. . It dates back to
the days when scintigraphy was a new experimental research
method to study physiological processes in the
body.
In most of the
text of this chapter (§4.1-4.8) we will deal with the physical
principles of scintigraphic imaging and technical
solutions of devices for gammagraphic imaging. The
clinical use of scintigraphy in nuclear medicine is
summarized in the last §4.9 " Clinical scintigraphic
diagnostics in nuclear medicine
". And the therapeutic use of radionuclides
is discussed in §3.6 "Radiotherapy", part " Radioisotope
therapy ".
Types of scintigraphy
Before we deal with specific physical-electronic methods for the
implementation of scintigraphic imaging, we will briefly
introduce the division (classification,
categorization) of scintigraphic methods. In terms of time
, scintigraphy can be divided into two types :
In terms of spatial-geometric, we can divide scintigraphy again into two categories :
In terms of complexity and interpretation of scintigraphic examination, we can distinguish two basic categories :
Radiation
exposure during scintigraphy
The radiation exposure of a patient during scintigraphy
examination depends on the type of radiopharmaceutical
applied (chemical form and radionuclide used) and on the
applied activity (see §5.7 " Radiation exposure during radiation diagnosis and
therapy "). In the case of
pure g- radionuclides
(such as 99m Tc), the radiation exposure is relatively low, as most
of the penetrating radiation g
passes through the tissue and carries its
energy out. There is a significant difference between X-ray
diagnostics and nuclear medicine in the laws of radiation
exposure.During X-ray examination, the source of ionizing
radiation is a device 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.
Basic
principles of scintigraphic imaging
How to achieve gammagraphic imaging ?
An idea might arise to use photography for this
: Radiation g is an electromagnetic wave of the same physical nature
as light. If we want to display an object using light (reflected
or actively emitted), we use the laws of geometric optics
and use a contact lens to project an image of the object
on a sensitive photographic layer and expose it for some time - a
photochemical reaction creates a latent image becomes a visible
image of different densities of silver grains in the photographic
emulsion - see the picture on the left.
Fig.4.1.1. Comparison of photographic imaging options in visible
light and gamma radiation.
It would be very pleasant if the patient could
be "photographed" in this way in g radiation - in the picture
in the middle .. Unfortunately, this is not possible! Radiation g will not refract
like light when it strikes the lens. As shown in §1.3, radiation
g interact
with each substance, and thus with the material of the optical
lens, in three ways :
1. Photoeffect- here terminates the
incoming photon and therefore all will reach the sensitive layer Þ is not used for imaging.
2. Compton scattering
- here would be some scattered photons g could strike the sensitive
layer and elicit a photochemical reaction there, but the
scattering angle is essentially random and always different,
regardless of the angle of incidence. Compton-scattered radiation
thus produces no image, but only a more or less monotonous
graying or blackening of the film. Thus, even Compton scattering
is not applicable for photographic imaging in gamma radiation *).
*) However, this statement is not
completely absolute, it only applies to photographic images. At
the end of §4.2 it will be shown that Compton scattering of
radiation g can in principle be used for electronic
collimation in so far experimental so-called Compton
cameras .
3. Formation of e - e +
-pairs (if the primary radiation g had energy >> 1MeV) -
here the primary photon g disappears and the secondary photons of annihilation
radiation always fly in opposite directions *), but each time at
a different angle in space Þ the same as in Compton
scattering.
*) This fact is used for
electronic collimation in positron emission tomography
(PET) - see §4.3, section " Positron emission tomography PET ".
We would reach the same conclusions if we tried to use a hollow
mirror instead of a lens to display it in g radiation . Only
the simplest pinhole displayin
Fig.4.4.1 on the right it also works for gamma radiation, it is
used in Pinhole type collimators (they are described below in the section " Scintigraphic
collimators ") .
Radiation g therefore do not
apply the laws of refraction and reflection Þ for radiation g there is
no refractive or reflective optics ! In
no way can we specifically influence the direction of motion of
photons of radiation g *). Only for soft X-rays, under certain circumstances, reflective
mirror optics still work anyway , but only for very
small angles of incidence-reflection - see
the appendix " X-ray telescopes " at the
end of §3.2.
*) Physically speaking, only gravity can
influence the direction of motion of photons g (due to its
universality) . Although such gravitational lenses of
gigantic dimensions are abundant in space (see
§4.3, passage " Gravitational
lenses. Optics of black holes .
" In the book "Gravity, black holes and
space-time physics" ) , they are not feasible in laboratory conditions on
Earth; even if we could make miniature black holes
with the required properties, the quality of their images would
not be very good and, most importantly, they would immediately kill
us with their gravity and quantum radiation (§4.7 "Quantum radiation and thermodynamics of
black holes" in the same book) .
The only way
to achieve an image in g radiation is collimation - shielding g radiation from
all unwanted directions and releasing only radiation from the
required direction. This creates a collimation projection
in gamma radiation. This is how g scintigraphy methods are
"treated" by g radiation - see " Scintigraphic
collimators " below.
Exceptions are special methods using so-called electronic
collimation using coincidence detection of two or more
primary or secondary photons. These principles are used mainly inpositron
emission tomography , or for so far
experimental Compton cameras (see
section " Compton cameras " and " High
energy gamma cameras
") or Compton telescopes in
astrophysics - some "telescopes without lenses and
mirrors".
Motion
scintigraph
Historically, the first type of instrument to perform
scintigraphic imaging of the radioactivity distribution was a
motion scintigraph , sometimes called a scanner
. The first device of this kind was built in 1951 by B. Cassen
and his colleagues, their main manufacturer in the 60s and 70s
was the company Picker (Fig.4.1.2
right). It is in principle a simple device,
schematically shown in Fig.4.1.2:
Fig.4.1.2. Motion scintigraph.
Left: Principle diagram of the movement scintigraph
(bottom middle is an example of a thyroid
scintigram) . Right:
Scintigraph Picker 500i at KNM Ostrava.
A collimated scintillation detector
*) is mounted at one end and an electromagnetic pen
at the other end on a common massive arm moved
by an electric motor . The detector is uniform
meandering motion moves the measurement object W ,
the radiation g (which is detected only from the area just below the
collimator on its axis) is converted into electrical pulses,
which (after amplification and amplitude
discrimination or. Reducing excessive frequency) leads to electromagnetic coil. For each pulse, a
ferromagnetic core provided with a pen (punch) at the end is
ejected from the coil of the electromagnet, which prints
a mark on the paper over the inking tape.(comma). Each
comma represents, depending on the reduction setting, a hundred
or a thousand pulses or the like. The higher the radioactivity of
the place above which the collimated probe is located, the higher
the frequency of pulses the probe will send to the solenoid coil
and the denser the pen will type the lines of the image as it
moves over the paper. The result is a display of the invisible
distribution of the radio indicator using the visible density
of lines on the paper (Fig.4.1.2 in the middle) - a
scintigraphic image W * is created . In addition to paper, some instruments
also recorded scintigrams on photographic film ,
which made it possible to better distinguish details in the
frequency of pulses.
*) To increase the detection efficiency,
relatively large scintillation crystals with a diameter of up to
15 cm, equipped with multi-hole focused collimators, were
used . Thus, radiation g
from a relatively large spatial angle was
concentrated on the surface of the crystal from the focus at the
investigated site .
The advantage of the motion
scintigrapher was simplicity and perhaps also
the fact that it provided an image in a 1: 1 scale. However, it
had some major disadvantages . In the first
place, it is a very low measurement efficiency :
only a small part of the g photons is always detectedonly from the place above
which the detection probe is currently located - radiation from
all other places escapes uselessly. Furthermore, the probe moves
relatively slowly over the patient and takes a long time to scan
the scintigraphic image. If the distribution of the radio
indicator changes with time during the measurement, we are not
able to capture and display these changes - the motion
scintigraph does not allow dynamic scintigraphy
. For these reasons, movement scintigraphs have not
been used since about
the end of the 1980s (they have been used
for thyroid scintigraphy for the longest time, Fig. 4.2.1 in the
middle of the bottom) - they were
completely replaced by scintillation gamma cameras
.
A scintillation camera is a device that captures photons of radiation g simultaneously from the entire field of view, converts them into electrical impulses and then uses them to create a scintigraphic image of the distribution of the radio indicator in this field of view. |
The
principle of the scintillation camera
Scintillation cameras, or gamma cameras
, are so far the most perfect devices for
scintigraphic imaging of radioactivity distribution. It is a very
complex device both in its principle and in its technical
construction .
The first
scintillation camera was designed by HOAnger in 1958. In the
initial experiments, he used a single-hole collimator and exposed
the scintillation in a thin crystal of larger diameter to a
photographic plate. He achieved a striking improvement by
attaching photomultipliers (originally 7 photomultipliers) to the
crystal, which sensed flashes in the scintillation crystal and
converted them into electrical pulses that were electronically
evaluated. The company started producing the first scintillation
cameras with 19 photomultipliersNuclear Chicago in 1964,
soon to be Picker (a leading manufacturer of motion
scintigraphs); later in Europe by Intertechnique, Philips,
Gamma , in Japan by Toshiba .
The schematic
diagram of Anger's scintillation camera is shown
in Fig.4.2.1 :
Fig.4.2.1. Schematic diagram of a
scintillation camera (analog).
Note: For
clarity, only two photomultipliers F1 and F2 are shown. In fact,
there are a larger number of photomultipliers - min. 19 (for
older cameras with a smaller field of view), 32, 64 and more.
Detection
of radiation g and determination of the place of
its origin
Consider (model) investigated object W , in which
there are three localized deposits A , B
, C of increased concentration of g- radio indicator.
From each place of deposition of radioactivity, radiation g is emitted
isotropically on all sides , which, due to its penetration, emanates from the object W
out. In order for an image g
to be imaged, a collimation
projection must first be performed . To do so, the
radiation alleging g put in the path of the lead plates
pierced with plenty small parallel holes . Only
those photons g that move exactly in the direction of the axis of the
holes can pass through this collimator . Other
photons that go "obliquely" are absorbed on the lead
partitions between the holes. The collimator thus creates a planar
projection of the radio indicator distribution into the
blue marked plane in Fig.4.2.1. A thin large-area
scintillation crystal is placed here . Each photon of
radiation g that passes through the collimator causes a
scintillation flash of a large number of photons of (visible)
light in the crystal. Crystal scintillations are sensed and
converted into electrical pulses by a system of
photomultipliers, optically adhered to the crystal. For
simplicity, only two photomultipliers are drawn in Fig.4.2.1 - F1
and F2 .
Let us now observe the
"fate" of the individual photons g emitted from the
interior of the object W under investigation . In particular, any photon g 'that flies in a
direction other than exactly perpendicular to the collimator face
(i.e., parallel to the orifice axes) is absorbed
at the baffles between the collimator orifices, does not
fall on the crystal, and is not detected . The
photon g A , which flies in the
"right direction" from position A ,
passes through the collimator opening and causes at position A´
in the crystal scintillation, whose photons propagate in all
directions in the crystal. A photomultiplier F1
, which is close to the site A´ of
scintillation, will receive a relatively large number of photons
from this flash, so that the pulse at its output will have a high
amplitude, while the distant photomultiplier F2
will receive only a small portion of these photons and its pulse
will be very low. For the photon g B
from position B, scintillation occurs
approximately midway between the photomultipliers F1
and F2 , so that the amplitude of their pulses
will be approximately the same. For photon g C
(radiated from bearing C ), which strikes and
causes scintillation near the photomultiplierF2
, the photomultiplier F2 will receive much more
light than the photomultiplier F1 , and the
amplitude ratio of their pulses will be the same.
In general, most light enters the
photomultiplier, which is closest *) to the
flash point (the point of interaction of
the photon g with the crystal) - therefore a
pulse is generated at its output, the amplitude of which is larger
than the amplitude of pulses from more distant photomultipliers.
from a given flash. The localization of the flash positions is
thus performed by a kind of electronic-geometric " triangulation
", determined as the "center of gravity" of the
signals from the photomultipliers.
*) The photomultiplier receives the largest
portion of light when scintillation occurs directly below the
center of the photocathode. From scintillations at more distant
locations, fewer photons will fall on the photocathode, so the
output signal has a lower amplitude.
Thus, we see that by
comparing the amplitudes of the pulses from the
individual photomultipliers, it is possible to calculate
the position of the flash in the crystal, and thus the place
in the patient's body from which the photon g was emitted.
Pulses from individual photomultipliers (of which there are a
larger number - 16 (for older cameras with
a smaller crystal) , 32, 64 and more), are
led to an electrical circuit called a comparator
(based on a resistive matrix), where the pulse amplitudes are compared and the
resulting X and Y coordinate pulses are
generated - these already carry direct information about the position
of scintillation in the crystal and thus also about the position
of the place in the organism from which the respective
gamma photon was emitted. After amplification, these X
and Y pulses are fed to the deflection plates of
the oscilloscope screen , where they determine the
position of the flash on the screen (this
was the case with older analog gamma cameras
used in the 1960s and 1970s) .
Amplitude
analyzer
In addition to coordinate analysis, pulses from all
photomultipliers are fed to the summing circuit
- from the point of view of this circuit, the whole scintillation
camera behaves as one large scintillation detector of radiation g . These summation
pulses whose amplitude is proportional to the energy of
the absorbed radiation g, are then sent to an amplitude analyzer
*) ( selector pulses according to
amplitude) - for each flash is thus
determined not only by its position (coordinate
pulses X, Y) but also the energy of the photon g , which this flash
caused. The analyzer window is
set to transmit only pulses corresponding to the photopeak
- total absorption of radiation gin the crystal. If the
radionuclide used has more radiation energies g , the window is
usually set to the "main" (strongest) photopeak, or
measurements in multiple windows set to individual
photopeaks shall be used.
*) The principle and role of the amplitude
analyzer in radiation spectrometry is described in §2.4 " Scintillation
detectors ".
For correct radiometric
measurements on each spectrometric instrument , the basic
condition is to set the analyzer window to the photopeak of the
gamma radiation of the used radionuclide. In the case of a
scintillation camera, in addition to the detection efficiency,
the correct adjustment of the analyzer window is necessary to
suppress Compton scattered radiation and to ensure the alignment
of the photomultipliers to achieve good field of view homogeneity (see below passage " Adverse
effects with scintigraphy and their correction ," part "Compton scattering g ) .
In older types of gamma camera
settings window analyzer photopeak manually, with modern digital
cameras is implemented automatic setup and
tuning window analyzer - called. Peaking ( Picking
) or Auto Peak ( automatic tuning peak
). By comparing the frequency of pulses in the lower and upper
half of the window, this window analyzer automatically tunes
in the middle photopeak (see figure) :
Formation
of analog scintigraphic image
The pulses behind the amplitude analyzer, called Z
(have nothing to do with the third
dimension coordinate!) Are uniform
"trigger pulses" - they say: "Yes, a 'correct' g photon has now
been registered and the X and Y coordinate pulses are
valid." The Z pulses are fed to the grid of
the oscilloscope screen; here it cancels the negative bias for a
moment, causing the cloud of electrons to emerge from the
cathode, focusing and accelerating in the "electron
cannon" and flying towards the screen screen. In the
meantime, the X and Y
coordinate pulses have already appeared on the accelerating
plates , whereby the electron beam is deflected in the
appropriate direction and falls into the appropriate place ( A
* , - depending on where the photon is emitted g A
, B or C ) of the screen
screen, where it emits a flash of light . As the
flashes gradually come to the screen screen as if they were
"raining" there, these analog images are sometimes
called "expired images".
In this way, the invisible
distribution of the radio indicator in the examined
object W , via physical-electronic detection of
invisible gamma radiation, is displayed in the form of a density
of visible flashes in the corresponding places of the
screen - a scintigraphic image W * is created .
Radioactive structures (lesions) A, B, C in the examined object
are displayed as sites A
* , B * , C *
with increased number of flashes on the screen.
The described scintillation camera according to Fig. 4.2.1 provides analog scintigraphic images on the oscilloscope screen. This image is present here only for the duration of the photon scan by the g camera, after the scanning ("patient departure") this image disappears. To preserve it, it was photographed from the screen with a camera whose shutter was open while the pulses were being stored. The so-called persistent oscilloscope was also often used , on the screen of which the flashes did not disappear immediately, but remained here for an adjustable time and only then gradually faded until they disappeared.
Digital
scintigraphic images
The above-described photographic method of recording (analog)
scintigraphic images has the disadvantage that it cannot be
post-edited (perhaps intensifying dark underexposed and weakening
light overexposed areas) and, most importantly, can not
be quantified . Therefore, with the development of
desktop minicomputers in the 1960s, there was an effort to
supplement (and later replace) oscilloscopic imaging of analog
scintigraphic images by digitizing them and storing them
in computer memory . The scheme of operation of such a
gamma camera equipped with an acquisition computer is shown in
Fig.4.2.3 :
Fig.4.2.3. Right:
Creation of a digital scintigraphic image by AD-conversion of
analog X, Y coordinate pulses, their storage in the image matrix
of the computer memory and display on the monitor screen.
The scintillation camera itself and the
relevant electronic circuits for amplification, comparison,
summation and amplitude analysis of pulses are identical as in
Fig.4.2.2. Only the oscilloscope screen in the right part is
replaced by a special circuit - the so-called analog-to-digital
converter ADC ( A
nalog-to- D digital C onverter)
and computer memory . The
actual conversion process is triggered by the trigger
pulse Z , which indicates that a valid photon of
radiation g has been detected. The amplitudes of the X and Y
coordinate pulses are then converted by the ADC converter into
digital (numerical) information - a bit combination - and sent to
the corresponding cell address in the computer. A certain sequence
of cells is set aside in the computer's memory to write
these digitized pulses; these cells are software-arranged into a
so-called image matrix - it is usually 64x64,
128x128, 256x256 cells (exceptionally also 512x512 cells, for
cameras with a rectangular field, even the image matrix is ??not
square). Each cell in the image matrix topographically
corresponding to a specific location in the displayed
object W . The field of view of the gamma camera is thus divided
by a grid into small squares - pixels (picture element), which
correspond to individual addresses in a defined part of the
acquisition computer's memory.
Before the start of the acquisition,
the contents of all cells are reset. If a digitized pulse arrives
at a cell from the ADC converter, its content is
increased by 1 . Thus, photons of radiation g , converted into
electrical pulses and digitized, gradually populate the cells in
the image matrix of the computer memory, according to the
location of the radiation, with ever-increasing values ??of their
content - a digital scintigraphic image formed
by the numerical content of the image matrix cells
in the computer memory. The numerical content of each of these
memory cells ( pixels ) is directly proportional
to the radioactivity corresponding place in the
organism, resp. its columnar projections from the entire depth of
the displayed area. The image matrix from the computer's memory
is then electronically displayed
("mapped") on the computer monitor screen.
FRAME mode, LIST mode
Process described above cumulative explicit recording the
scintigraphic image into memory is called a frame mode ( " image method ") . For special purposes (for phase dynamic studies and
iterative tomographic methods - §4.3, part
" Computer
reconstruction of SPECT
", " Reconstruction of
PET images ", " TOF - time localization of the annihilation
site " ) is sometimes used so called list mode
(" list method ") , where only a list of X and Y
coordinate values of successive pulses (together
with time stamps) is sequentially loaded
into memory and the own images are created additionally only
after the acquisition is completed.
Digital
scintillation cameras
With the development of electronics, especially the construction
of fast and miniaturized ADC-converters and microprocessors, the
digitization of the scintigraphic signal is no longer limited to
the conversion of analog X, Y coordinate pulses according to
Fig.4.2.3. With current so-called digital cameras
, each photomultiplier already has its
analog-to-digital ADC converter at its output . The
calculation of the coordinates of scintillation in the crystal is
not performed in an analog comparator, but in a digital
microprocessor, which already directly "install-fits"
the respective addresses in the computer's image matrix with the
relevant numerical information. In addition, the gain of the
preamplifier of each photomultiplier via a DAC converter is
controlled directly from the computer, which allows more accurate
and operative camera calibration - adjustment ( tuning
) and setting of appropriate corrections for homogeneity and
linearity.
Construction arrangement of scintillation
cameras
Gamma camera
detector
A large-area scintillation crystal
of a gamma camera with mounted photomultipliers (their number is usually 19 to about 100) and appropriate electronics is built into a special
robust housing (a kind of "pot"),
providing light tightness and radiation
shielding against external ionizing radiation. The metal
housing also shields the photomultipliers against an external
magnetic field. At the bottom of the camera housing is a
mechanism for attaching a collimator, which must
be tightly attached to the crystal. The collimators are
exchangeable, during manual exchange they are usually fastened
with screws, for automatic exchange the collimators are fixed
with special motorized holders. For SPECT cameras, there are also
touch sensors for mechanical protection of the patient
and the detector when the camera moves towards the patient.
Stand and gantry for
mounting detectors
The entire camera detector is then mounted on a special stand
equipped with electric motors for mechanical movement of
the camera - shift in the vertical, or. and horizontal
direction and rotation of the detector. For SPECT tomographic
cameras, the stand is made in an annular arrangement as a
so-called gantry , enabling the use of an
electric motorrotation of the camera around the
examined object. There are usually two detectors mounted on the
gantry, which can be angled around the axis of the lounger - a
"double-headed" camera. Additional electric motors
ensure radial movement of the detectors towards
the center and away from the center, so that it is always
possible to set the smallest possible distance between the body
surface and the collimator face.
Examination lounder
Under the camera detector, there is a bed for
the examined patient - perpendicular to the stand, or enters the
gantry. Manually or motorized, it allows horizontal movement
in a sufficiently large range (up to 2m) to be able to drive with the whole patient under the
camera or through the gantry and take pictures of different parts
of the body. To a lesser extent(approx. 60
cm) a vertical shift is also realized. The
lounger should be sufficiently robust (load
capacity min. 180 kg) and stable, ensuring
mechanical positioning with the possibility of locking. The
support plate of the lounger in SPECT cameras is made of a
material with low absorption of gamma and X-rays (when shooting from the front and back through the
lounger) . With the couch down and the
camera detector turned vertically, scintigraphic examinations of
patients can also be performed sitting or standing.
To realize the whole body
scintigraphy ( whole body imaging) is using the
electro-chair with the patient slowly moved in
the longitudinal direction, so that the individual parts of the
patient's body gradually enter the field of view and are detected
by the camera detectors; the acquisition computer consists of a
whole-body scintigraphic image - "sliding" whole-body
scintigraphy.
Auto-Body-contouring
To achieve the best possible resolution, the gamma camera
(collimator face) should be placed as close as possible
to the patient's body surface (trigonometric
analysis is performed below in §4.5, section " Spatial resolution ") . Auto-contouring
or body-contouring is a useful opto-electronic
tool for ensuring optimal quality of scintigraphic imaging in
whole-body and SPECT examinations : when moving the bed and
rotating the camera, using electronic position sensors, the
camera detectors on the gantry are automatically moved by
electric motors so that they "copy" the patient's body
and the collimator face is still as close as possible to the
patient's body surface (automatic
"body contouring") .
Auto-contouring is
realized by means of two rows of infrared LED diodes
and two rows of opposite photodiodes , placed in
two strips mounted on opposite edges of the camera detectors.
Electronic circuits regulate the radial position of the gamma
cameras so that the infrared rays from the outer row are
interrupted, but not from the inner row (closer to the front of
the collimator). The distance of the detector is thus constantly
kept in the range between the two rows of LEDs <->
photodiodes, approx ..-... mm.
Fig.4.2.4. Construction arrangement of a scintillation camera.
Left: Uncovered scintillation camera detector -
collimator, crystal, system of photomultipliers and electronic
circuits.
Right: Example of an assembled planar camera
with one detector (top) and a SPECT tomographic camera with two
gantry detectors (bottom).
In the left part of Fig.4.2.4 is a disassembled
detector of a smaller older camera (PhoGamma
Nuclear Chicago, with 19 photomultipliers) ,
removed from the shielding package. Below we see a collimator,
above it is a thin circular scintillation crystal, to which
photomultipliers are optically attached via light guide blocks.
In the upper part of the detector there is the appropriate
electronics, especially the preamplifier for each
photomultiplier, adjustment circuits, for digital cameras also
analog-to-digital converters and microprocessors for determining
coordinate pulses. Newer scintillation cameras have a larger
rectangular crystal, equipped with a larger number of
photomultipliers.
In the right part of Fig.4.2.4 there
is an example of two installed cameras. Above is a smaller planar
camera with one detector on a simple stand(PhoGamma
HP from 1973, with Clincom evaluation device; on the left next to
the camera stand there is a stand with interchangeable
collimators) , at the bottom there is a
larger SPECT tomographic camera (from 2002)
with two detectors ("heads")
mounted on gantry *) and motorized movement of a lounger for
whole-body scintigraphy.
*) Other structural arrangements of
scintillation camera detectors are also used sporadically
(Anger-type camera detectors themselves are designed almost
identically for different types and manufacturers; other
alternative technical solutions are mentioned below). Instead of
the classic circular gantry, the detectors can be placed on
special arms , the movements of which are
electronically controlled by servomotors.. The
advantage here is greater flexibility of different detector
positions (including the possibility of simultaneous independent
sensing of two patients by each detector separately). In addition
to "universal" cameras, special single-purpose
cameras with a fixed detector configuration are
sometimes used , such as 3 or 4 detectors connected in a triangle
or square, designed for scintigraphy of the heart (myocardium) or
brain.
These constructions generally
have not proven...
The electronic
circuits of the scintillation camera have been described
above (in the section " Principle
of the scintillation camera ") only in a general and simplified way, rather from a
physical point of view. Scintillation cameras are equipped with a
number of other electronic circuits for adjustments and for
corrections of physical-electronic influences. Circuits for coordinate
pulse correction, for example, are important X, Y -
shape and size of the image, especially the correction of the dependence
of the image size on the energy of the detected gamma
radiation - so that the scale of the image is not dependent on
this energy.
Scintigraphic
collimators
The primary " optical member " of a
scintillation camera through which radiation g is the first to
pass is the collimator *). In terms of gamma
imaging, the collimator has an analogous function as an optical lens
when photography. Its task is to make the most perfect projection
of the distribution of radioactivity in the examined
object using g- radiation into the plane of a large-area scintillation
crystal. Therefore, the final quality of the
scintigraphic image largely depends on the properties of the
collimator .
In general, the collimator is a
special screen made of a shielding material
(mostly lead, sometimes tungsten), defining the direction
of the photonsincident on the scintillation crystal as
well as the field of view of the camera. Most often it is a plate
with a large number of densely and evenly spaced holes
- channels - of a certain shape, size and
direction. Without attenuation, only photons flying in
the direction of the axis of the collimator's orifices
pass through the collimator (and impinge on the crystal) , or
only with a small deviation, ie almost perpendicular
to the collimator front and the crystal surface. Other photons in
other directions are absorbed in lead partitions
( septs ) between the holes, they do not fall on the
crystal and are not detected.
*) From the general point of view of
radiation physics and radiation detection, collimators were
discussed in §2.1 " Methodology
of ionizing radiation detection",
paragraph" Shielding, collimation and filtration of detected
radiation "and in §3.1" Nuclear and radiation methods
", section" Collimation of ionizing
radiation "). In
scintigraphy, collimators have an imaging role .
electronic collimation - see below Positron
PET emission tomography
.
Collimators
for scintillation cameras are usually interchangeable
- there are several types of collimators with clearly defined
properties, which govern their use. radiation energy gfor which they are
optimized, according to the resolution and sensitivity (detection
efficiency). The imaging properties of collimators are discussed
in more detail in §4.5 " Physical
parameters of scintigraphy
".
Here we give a brief overview of the
basic types of collimators - Fig.4.2.6. First we will discuss
collimators parallel holes - channels -
perpendicular to the scintillator crystal cameras, which are by
far the most common type - this image of an object formed in the
detector has the same magnitude of 1: 1 as the display
object, independently of the distance from the source collimator (this distance however, the spatial resolution of
the display depends significantly (see below) .
Fig.4.2.6. Left: Basic types of
scintillation camera collimators (gamma camera crystal is in the
up position, just above the collimator). Right:
Example of a robust high energy collimator HE and a subtle low
energy collimator LE HR (and UHR) .
Collimators
for different energies
The most basic criterion according to which collimators are
divided is the radiation energy g
for the scintigraphic imaging, of which the collimators are
optimized. According to this gamma radiation energy, the
collimators have different thicknesses of the partitions
(septums) between the openings *), sufficient to absorb the
radiation of a given energy.
*) The thickness of
the partitions
when optimizing the design of a collimator for
gamma photon energy required is based on the requirement that
only gamma rays pass through the holes, while in the partitions
(septa) therebetween is effectively absorbed by
. If gamma radiation penetrated the partitions to a greater
extent, it would degrade the imaging properties of the
collimator, especiallydisplay contrast (it is discussed in §4.5, passage " Irradiating the collimator septa ", Fig.4.5.3) . A large
thickness of the baffles would be required for complete
absorption of gamma photons, leading to very low detection
efficiency. Irradiation of 5% is taken as a
sufficient criterion for achieving a reasonable level of septa
irradiation, without significant deterioration of the image
contrast . According to the trigonometric analysis in Fig.4.5.3b in
the passage " Irradiation of the collimator septa ", this leads to the
condition for the transmission factor e -m .s.L
/ (2d + s) <0.05, where d is the diameter of the
holes, L their length, sthe thickness of the baffles and m is the linear absorption
coefficient of the collimator (lead) material for the required
gamma energy. This gives a limitation for the thickness
of the collimator baffles s > (6.d / m) / [L - (3 / m)]. The
optimal is the smallest possible thickness of
the partitions allowed by irradiation - so that the
septa shades the smallest possible area of the detector and the efficiency
(luminosity) of the collimator is the best
possible .
The absorption coefficient of the
collimator (lead) material strongly depends on the gamma
energy , on which the required thickness of the
baffles depends . For low energies around 150keV, where for lead
ism » 21.4
cm -1 , eg
for a collimator with holes with a diameter of 2 mm and a length
of 25 mm, the required thickness of the partitions is based on » 0.3 mm (thin lead foil) . For higher energies around 400keV, where m is » 2.5 cm -1 , significantly thicker
partitions s » 4.5 mm are needed .
According to gamma energy we have 4
basic types of collimators (Fig.4.2.6 left) :
Recently, it has been constructed :
Appropriate selection of the collimator
according to the energy of the emitted gamma radiation has a
fundamental effect on the quality of the scintigraphic image. For
low energies, such as 140keV 99m Tc, we use Low Energy collimators , which
provide the best resolution. If we used a robust HE
collimator for high energies here, we would get an image with
lower resolution and lower detection efficiency, on which, in
addition, the lead septa between the holes of the collimator *)
would be disturbingly visible. We can also use the Pinhole
collimator (see below "Collimators
with special geometry") , which
provides a quality image, but with lower detection efficiency.
For higher energies, such as 364 keV 131 I, the collimators are Low Energycompletely
unusable, significant irradiation between the septa completely
degrades the image into a shapeless "daub" (it is discussed in §4.5, passage " Irradiation of the collimator septa ") . It is imperative that
we use a High Energy collimator here (the holes and partitions of the collimator can also be
seen in the picture) or Pinhole .
Pinhole is the only type of collimator that is widely independent
of energy.
*) This disturbing structure of the holes and septa of the HE
collimator can be suppressed by a stronger smoothing of the image
(approx. 4 x S9), at the cost of a lower resolution - pictures on
the right.
Scintigraphic images of a thyroid phantom filled with 99m Tc ( top
) and 131
I ( bottom ), imaged using a Pinhole
collimator , Low Energy HR and High Energy .
The disturbing display of the holes and septa of the HE
collimator can be suppressed by a stronger smoothing (filtering)
of the image - pictures on the right.
Collimators according to
resolution and sensitivity
Another criterion for the division of collimators is their
required resolution and sensitivity
(efficiency - "luminosity"). However, this only applies
to low energy LE
collimators ; with robust collimators for high and medium
energies, we cannot achieve good resolution or high sensitivity
due to the thick partitions between the holes (and thus the low hole density) .
Depending on the resolution and sensitivity, therefore,
low-energy collimators are further :
The number of collimator
holes
depends on the type of collimator and its size (area) of the
camera's field of view. With the current planar / SPECT cameras,
the field of view is around 55 x 45
cm . The total number of holes for the basic types of
collimators is then approximately:
HE - 8000 holes; ME - 15,000
holes; LEAP - 80,000 holes; LE HR (UHR) - 140,000
holes.
The holes are usually hexagonal in shape.
Spatial resolution of a
gamma camera
The spatial resolution of a camera is determined by two
components: the internal resolution of the detector and
the resolution of the collimator (for
a more detailed analysis, see §4.5, section " Spatial resolution ") . The resolution of the collimator
is determined by the diameter of the holes and
their length . HR collimators with narrow and
long holes (the length of the holes is
given by the thickness of the collimator) have
better resolution than thinner HS collimators with larger and
shorter holes. The spatial resolution of the gamma camera
significantly depends on the distance displayed
structures from the collimator front. From each hole of the
parallel collimator we can draw an imaginary cone
defining the area from which gamma radiation can pass through
this hole to the camera detector (radiation
from places outside this cone is absorbed by the lead septa of
the collimator) . With the distance from
the collimator, this detection cone widens ,
which significantly worsens the geometric
spatial resolution of the image projected by the collimator on
the scintillation crystal of the gamma camera (trigonometric analysis is performed below in §4.5,
section " Spatial resolution ", here for clarity. 4.5.2 :) .
Fig.4.5.2. Deterioration of the positional resolution of
the gamma camera with increasing distance h from the
collimator front. The image of the point source becomes more and
more "blurred" with increasing distance, the PSF
expands and the spatial resolution of the FWHM deteriorates -
Fig. d). Deterioration of the spatial resolution
is accompanied by a decrease in the brightness of the image, but
the total number of pulses is the same in all images and the area
(integral) under the
PSF function is also the same for all distances.
The gamma camera (front
of the collimator) should therefore be
placed as close as possible to the surface of
the patient's body. For collimators with a different arrangement
of holes (see below) ,
the geometric situation is more complicated, but in principle the
same rule applies to the deterioration of the
spatial resolution for greater distances from
the collimator face.
Detection efficiency
of the scintillation camera
The detection efficiency
(sensitivity) of the camera is given by the efficiency
(luminosity) of the collimator and the internal detection
efficiency of the detector (discussed in more detail in §4.5, section " Detection efficiency (sensitivity) of the gamma
camera " ). Efficiency (transmittance, luminosity) the
collimator is given by the diameter of the
holes and their length , but in the opposite
ratio to the resolution. The larger and shorter the holes, the
higher the detection efficiency. The efficiency or luminosity of
collimators is generally very low - around 1-2%.
Interestingly, with gamma cameras,
when using parallel collimators, the detection
efficiency (sensitivity) does not depend on the distance h
of the displayed source from the collimator front! The
imaging of the point source in a wide range of distances 0-30cm
from the front of the collimator in Fig.4.5.2 d) shows a
deterioration of spatial resolution and decreased image
brightness, but the total number of pulses is the same in all
images, area (integral)under the PSF function is the same for all distances. This
surprising behavior is due to the specific properties of
geometric collimation in parallel collimators.
We can clearly illustrate this according to the schematic drawing
in Fig. 4.5.2 b) as follows: As the source moves away from the
collimator front, the number of photons incident on the
individual holes decreases quadratically as 1 /
h 2 .
However, the number of holes through which radiation can pass to
the detector increases quadratically in
proportion to h 2 . These two opposing trends cancel
each other out , so the total photon flux - collimator
efficiency - does not change with the
distance between the source and the collimator.
Note: This
rule does not apply to special convergent or Pinhole
collimators , the detection efficiency changes
significantly with distance - it increases or decreases (see
§4.5, section " Imaging
properties of special collimators
") .
However, this distance sensitivity
independence of parallel collimators only applies to situations without
a substance-absorbing environment - in vacuum or in
air . In practical scintigraphy, however, there is a tissue
environment between the displayed structures with
distributed radioactivity in the organism and the gamma
camera.with which gamma radiation interacts, leading to the absorption
and attenuation of gamma radiation. This
attenuation of gamma-ray absorption, also called attenuation
, is reflected in scintigraphic images by an artificial
reduction in the number of pulses from structures deposited
at greater depths, compared to structures closer to the surface.
In such a case, the statement that the detection efficiency
(sensitivity) does not depend on the distance of the
displayed source from the collimator face no longer applies
. Here, the detection efficiency decreases significantly
with the distance - depth - of the displayed source !
Collimators with special
geometries
In addition to collimators with parallel holes - channels -
collimators with otherwise geometrically arranged holes are used
for some special purposes (Fig.4.2.6. In the middle) :
The imaging properties of collimators are
discussed in more detail in §4.5 " Physical
parameters of scintigraphy
". Here, for clarity, we will only duplicate graphs of the
dependence of the spatial resolution and detection efficiency
(sensitivity) of the gamma camera with basic collimators on the
distance :
Fig.4.5.6. Depending spatial resolution FWHM ( left
) and the detection efficiency of S ( right
) gamma camera to face away from the source of various kinds of
collimators ..
Imaging
properties of the most important types of collimators with
different geometric arrangement of the holes, we tested using
linear orthogonal grid (its
construction is described in " Phantoms and phantom
measurements in nuclear medicine" Image" Grid ")
:
For a collimator with parallel
holes (such as LE HR left) brings all linear
display a grid, wherein for a greater distance from the front
collimator deteriorates the spatial resolution (blur matrix). In a
convergent collimator (such SmartZoom with the
convergent center part) the image of the center part increases
with increasing distance.For the Fan Beam
collimator (which is convergent in the transverse direction,
parallel in the axial direction), they increase only in
the transverse direction with increasing mesh distance ,
they remain the same in the axial direction .
The most significant dependence on
the object distance exhibits the collimator Pinhole:
right at the forehead we get an image magnified
many times , with increasing distance the zoom decreases and for
distances above about 20cm the image is already reduced
.
All pictures also show a general
trend of deteriorating resolution (and thus contrast in the
image) with the distance from the front of the collimator.
Scintigraphic images
and their evaluation
The whole process of scintigraphic diagnostics is schematically
shown in Fig.4.2.5. After application of tracer
leads to its distribution in certain parts of
the body (uptake in target tissues and
organs, or the flow of radiotracer vessels and heart cavities) , this distribution with external detection based
radiation g show scintillation camera, computer generated digital scintigraphic
images which both evaluate visually ,
on the one hand, we can mathematically analyze the
investigated processes and calculate quantitative
parameters using curvesfunctions of individual bodies.
Finally, an interpretation of all these partial
data and results begins , which, together with the results of
other methods, will result in the diagnosis being made
in the final protocol .
Fig.4.2.5. Schematic
representation of the whole process of scintigraphic examination
- from the application of the radioindicator to the patient and
its uptake in target tissues and organs, through the process of
scintigraphic imaging with a gamma camera, visual evaluation of
images, mathematical analysis and quantification, to
interpretation and diagnosis.
The methodology of mathematical analysis and computer evaluation of scintigraphic studies will be discussed in more detail below in Chapter 4.7 " Mathematical Analysis and Computer Evaluation in Nuclear Medicine ".
Adverse
effects of scintigraphy and their correction
In scintigraphy, there are some unfavorable and disturbing
phenomena that can worsen the quality of the image and
thus, in the extreme case, even lead to incorrect
interpretation of scintigraphic examinations in the
sense of false negative or false
positive findings. Here are six basic adverse effects
that occur in general in every scintigraphy , ie
in planar scintigraphy and SPECT tomographic scintigraphy. Other
adverse and disturbing phenomena specific to SPECT (such as
instability of the axis of rotation or artifacts arising during
reconstruction) and PET (random false coincidences) will be
mentioned below in §4.3.
![]() |
Fig.4.2.7.
Volume and activity distortion in images of
lesions of various sizes (left) and specific
activities (right). Left: Images of sources with the same specific activity, but different sizes, appear differently clear. Right: Images of sources of the same size, but of different specific activities, appear to vary in size. (measured on PhoGammaLFOV camera, FWHM = 6mm) |
These unwanted side
effects are marked in small lesions
smaller than 2.FWHM (twice the resolution), while in
lesions larger than about 3-4.FWHM are practically
negligible (they appear only at the edges of the lesion
image). The phenomenon is particularly negative in small
negative lesions - small areas of reduced
radioindicator concentration against the background of
higher radioactivity concentration. Here, the effect of
irradiation from the surroundings into the image of the
lesion can completely erase the visibility of
such a small lesion, which disappears in statistical
fluctuations; we say that such a lesion is not
detectable (see below " Scintigraphic image quality
- detectability of lesions" Fig.4.2.9) .
Overall, due to the" blurring "of the
positive image (" hot ") lesions occurs decrease
in the observed activity (number of pulses in
the image) in the negative (" cold ") lying on
her back rise . the common result is a reduction
in the contrast of the image (see also below " the quality of scintigraphic images -
detectability of lesions
") .
Correcting
distortion of the activity
Have been developed methods for correcting this distorted
view of activity - the so-called. Partial Volume
correction ( PVC), which is desirable for
quantitative image analysis, such as SUV determination
(see " Scintigraphic image quality - lesion
detectability " below).
Theoretically, reconstructive algorithms could be used
based on the knowledge of the response function of the
point source function PSF (point spread function) - the
above-mentioned method of resolution recovery
, but it fails at higher statistical fluctuations. In
practice, simple multiplication by correction
factors is sometimes used, which indicate the
ratio of the actual volume activity of the lesion to the
apparent activity in the image. These coefficients are
strongly dependent on the size of the displayed object
and on the spatial resolution of the scintigraphic image;
their specific values ??are determined on the basis of phantom
measurements . In the literature, the values of
inverse correction coefficients, so-called recovery
coefficients RC (recovery
coefficient) depending on the diameter of the spherical
lesion for different values of FWHM resolution are
tabulated or plotted . To use this correction method
correctly, you need to know the actual sizeimaged
lesions, which is practically only available for hybrid
systems combining scintigraphy with anatomical CT imaging
(see below §4.3 "Tomographic cameras", section
" Image
fusion, hybrid tomographic systems "). For small lesions of about 1 cm with a
resolution of FWHM @ 8 mm, the correction coefficient is 1 / RC @ 5, for
smaller lesions or worse resolution (as is usually the
rule at greater depths), its value is even higher. This
leads to a large correction error ,
which is practically unusable at values
of the correction coefficient approaching ~ 10.. Of
course, the above-mentioned resolution recovery
method is also unusable .
If the size of the displayed lesion is smaller
than the spatial resolution of the camera, the
differences in the volume of this structure will be
reflected only in the number of accumulated pulses in
pixels of the image location (brightness or intensity of
the displayed lesion). This effect is sometimes used to
assess changes in heart wall thickness in the SPECT
myocardium.
Note: The effect of
volume (size) and intensity distortion is manifested not
only in scintigraphy - it occurs wherever the image shows
convolutional blurring . And that is, to
a greater or lesser extent, in practically all imaging
methods ...
Fig.4.2.8. Influence of registered number of photons on image
quality in terms of statistical fluctuations (noise) - image
quality improves with increasing number of photons.
Above: Photographic portrait exposed with
different number of photons of light .
Bottom: Gammagraphic image of a phantom ( Jasczak , filled with 99m Tc radionuclide ) accumulated by
a scintillation camera with different numbers of g- photons in the
image.
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Compton scattered radiation in scintigraphy |
If, coincidentally, a
photon is scattered in the tissue at such an angle that
this scattered photon passes through the collimator
orifice and is detected by the camera crystal (in Figure a ) , then these photons of g záøení radiation
can be detected from a false location -
the gamma photon coming from a location other than where
it was originally radiated during the radioactive
transformation (a similar
"false localization" effect may occur when
Compton scattering g radiation in the material of the scintillation
crystal itself) may occur . These
randomly coming scattered photons g ´ would be
artificial reduced the contrast of the
scintigraphic image.
Fortunately, however, these
false scattered photons g ´ have lower
energy than the "true" direct and
primarily detected photons g (part of the energy
was transferred to the electron e - during
scattering in matter ) , so they
usually do not fall into the photopeak (in Figure b ) . By carefully setting the analyzer
window to the photopeak of the given radiation g, we can
therefore largely eliminate the
Compton-scattered radiation g ´ . More effective
suppression of compton. Scattered radiation can be
achieved through a narrower window , or
its slightly asymmetrical adjustment towards higher
energy - but at the cost of reduced detection efficiency (elimination of part of the primary photons) and the risk of a slight deterioration in the
homogeneity of the camera's field of view (in a narrow and asymmetrical window, matching
the response of individual photomultipliers may not be so
perfect) .
Scatter correction
However, a small proportion of Compton scattered photons
(scattered at a small angle) still affect
energetically into the photopeak area and can be
detected. Some types of scintillation cameras use special
electronic procedures to correct or eliminate
these remaining scattered photons .
Pulses for each pixel of the image are registered in two
or three energy windows (DEW - Dual
Energy Window, TEW - Three Energy Window; instead of the
word "Energy" "Photopeak" is
sometimes used and abbreviations are written DPW or TPW) : 1 . window just in front of the photopeak (with a
high proportion of scattered radiation), 2 . main
central window of the photopeak, 3. window just behind
the photopeak. Three matching images are created for
these energy windows. By interpolating the number of
pulses registered in the auxiliary windows before and
after the photopeak, the fraction of scattered
photons corresponding to the main window of the
photopeak is determined for each pixel - a " scattering
image " is formed , which is
subtracted from the main image in the basic window.
More
complicated methods of scattering correction were also
tested with a larger number of analyzer windows, or with
different homogeneity correction matrices for different
windows; however, with significantly greater complexity,
the results were not demonstrably better than with the
basic TEW method. Algorithms for obtaining (modeling) 2D
or 3D scattering distribution are being developed for
SPECT tomographic scintigraphy, with implementation into
reconstruction procedures of type MLEM, OSSEM (below
§4.3 "Tomographic scintigraphy" ...., section
" Computer reconstruction of SPECT ", passage " Iterative
reconstruction ").
Another source of false impulses
can be the so-called pile-up the
effect of the cumulative electrical response of two
quantum g , detected almost simultaneously (see §2.4,
section " Scintillation spectra of
radionuclides "). This
is manifested at high frequencies (fluxes) of g photons .
In most cases, these pulses fall off the photopeak and
are not detected. However, if there is a pile-up effect
on two simultaneous Compton scattered photons, the
resulting pulse may fall into a photopeak with its
amplitude - it is detected and may contribute to the
degradation of the scintigraphic image contrast.
Physical
parameters of scintigraphy
Resolution, detection efficiency, homogeneity and other
parameters of the scintillation camera are defined and discussed
below in §4.5 " Physical
parameters of scintigraphy - image quality and phantom
measurements ". The methods of
their measurement and testing are discussed in the work " Phantoms
and phantom measurements in nuclear medicine ".
Errors and pitfalls of
correction methods - correction artifacts
It should be noted that no correction methods are
"self-saving", but they can have their pitfalls. Errors
of correction methods can be divided into two categories :
Undercorrection, overcorrection, and correction artifacts can lead to similar deterioration (or even the risk of misinterpretation) of scintigraphic studies as uncorrected studies. Experience shows that in order to correctly interpret the findings, it is necessary to carefully compare images without correction and images with correction by the "trained eye" of an erudite expert, who must also take into account the specific anatomical and positional circumstances of the patient.
Scintigraphic
image quality - lesion recognition
The above-mentioned adverse effects mean that the scintigraphic
image is far from accurate and perfect - despite useful
information, disturbing statistical fluctuations
(noise) overlap , the image is out of focus and
often low in contrast . This imperfect quality
leads to the fact that some more subtle structures of the
examined object are not visible on the scintigraphic image - we
say that such lesions are not detectable . In
the diagnostic practice of nuclear medicine, such a scintigraphic
image is optimal , which, in addition to
objectively measurable physical parameters, also suits human subjective
visual perception evaluating physician. So what
parameters of the examined object and its image determine the objective
image and the best possible recognition of
lesions ?
The basic laws result from the properties
of scintigraphic imaging and from the statistical analysis of the
resulting image data. In the left part obr.4.2.9 shows
scintigraphic images of simple structure (lesion) of the circular
shape of the size (diameter) d and the
specific activity A surrounded by a
homogeneous environment - background - specific activity of B
. The prominence of the lesion against the background can be
characterized as the contrast of the object C obj = (A - B) / B (or ´ 100 [%]). Scintigraphic imaging
produces an image in which the lesion is shown as a structure A *
and the background as a constant (more or
less wavy) area B * (depending
on statistical fluctuations) .
Fig.4.2.9. Analysis of contrast and statistical fluctuations of
scintigraphic imaging of lesions (phantom
measurements on a PhoGamma LFOV camera) .
If we compare the original object with its
scintigraphic image, we see two main differences :
¨ 1. Blur and contrast reduction
Due to imperfect spatial resolution, the sharp contours of the
original object A were blurred and the
difference between image maximum A * and background B * decreased
- image contrast C img = (A * max - B *) / B * is lower
than the contrast of the object C obj : C img < C obj . Assuming a circular lesion and Gaussian convolutional
blur (the response function of the point
source of the PSF camera is in the form of a Gaussian curve with
a half-width FWHM) the relationship between
the contrast of the object and the image is given by an
exponential expression:
C
img = C obj . e - (FWHM
/ d) 2 ,
where FWHM is the camera resolution and d is the size
(average) of the lesion. For large lesions (d> 4.FWHM), the
contrast of the image hardly changes (C img @ C vol ). However, in
small lesions, comparable or smaller than the FWHM camera
resolution, the contrast degradation is very significant, C img «C obj (at a typical camera resolution of 10mm,
the contrast of a 1cm lesion decreases almost 3-fold, in a 5mm
lesion more than 30-fold !) .
¨2. Statistical fluctuations - noise
Due to the quantum stochastic laws of radioactive decay, emission
and detection of quantum radiation g , all parts of the
scintigraphic image show statistical fluctuations
- noise is covered over the image of the object
. As shown in §2.11 " Statistical fluctuations and measurement
errors ", the magnitude of this
noise at each point of the image is given by the square root of
the average accumulated number of pulses n : s = ± Ö (n) . The relative
statistical fluctuations s / n = 1 / Ö (n) are lower the higher the number of pulses
accumulated in the individual cells of the image. Constant
background B is thus shown as an area whose points
fluctuate roughly between B * ±
Ö (B *), ie with sB = ± Ö (B *). Similarly, the point values in the A* image
fluctuate statistically. If these fluctuations are too high,
comparable to the average values of the difference between A *
and B *, these differences can easily be "lost" in them
and the corresponding structure will not be visible in the image.
Disturbing statistical fluctuations are thus a fundamental
limiting factor for the recognizability *) of small and
not very contrasting lesions on the scintigraphic image.
*) Without statistical fluctuations, an
artificial increase in steepness would be
possible (contrast) display of a scintigraphic image on the
screen to achieve visibility of even small and low contrast
lesions. In addition, appropriate deconvolution filtering
(using the inverse modulation transfer function MTF) could
correct the camera resolution , resolution
recovery - computer "focus" of the image - and
reconstruct all details from the displayed object (see " Filters and filtering ", section " Band focus filters
"). Unfortunately, statistical fluctuations deprive us of
most of these possibilities in practice ...
The statistical analysis of image
data shows that we can only recognize (and
statistically prove) in the image a structure (lesion) whose
contrast C img satisfies condition
C
img > 4
/ Ö (B *).
It is a condition of the statistical significance of the
difference A * - B * information in the lesion image to the surrounding
fluctuating background B *.
Signal
- noise
In analogy with the analysis of electrical signals in low-current
electronics, the terms are introduced for the quantitative
description of image properties :
Signal S
is the difference in image intensity (its
"brightness", number of accumulated pulses) between the investigated structure (lesion) and
environment. In our case it is given by the difference: S = A * max -B *.
Noise N
represent disturbing statistical fluctuations in the image. For
our case, background fluctuations are important, so the noise is
given by the square root of the average accumulated number of
pulses in the background image: N = s B = Ö (B *).
Like the signal quality of the
electronics, the quality of the image given by
the parameter:
Signal to noise ratio SNR (Signal-Noise Ratio),
SNR = S / N = S / O (B *).
The above statistical condition for the detectability of a lesion
can then be expressed as follows: A lesion can be seen in an
image only if its signal-to-noise ratio is SNR> 4
.
If we take into account the effect
of resolution and statistical fluctuations, by combining the
above relationships we can basic formulate the condition
of lesion recognition as follows :
C
obj >
4. e (FWHM / d) 2 / Ö (B *).
Only such a lesion will be visible in the image, which will have
sufficient contrast Cobj (in the accumulation of radioactivity), geometric size d
large enough compared to the resolution of the FWHM camera and
the number of accumulated pulses will be large enough so that the
relative statistical fluctuations are not too high. The image of
the lesion is better the larger, more contrasting the lesion and
the higher the density of accumulated pulses in the image. And
the smaller the size and contrast of the lesion, the higher we
need to accumulate the number (density) of pulses in the image
for its successful imaging. For the display of these small and
not very contrasting lesions, the best possible resolution of the
camera is also crucial in order to avoid an enormous degradation
of the contrast of the lesion during the imaging.
Positive
and negative lesions
One of the differences between "cold" (negative) and
"hot" (positive) lesions is that well acumulating hot
lesions can have high contrast Cobj even many hundreds of percent, while with cold lesions
the contrast can reach a maximum of 100%. Therefore, we observe
well-displayed and small (but contrastively accumulating) hot
lesions, such as inflammatory or tumor foci in classical skeletal
scintigraphy or 18 FDG PET. Smaller cold lesions are difficult to observe,
especially when they are deeper (perhaps inside the liver or
lungs).
Deep-seated lesions
Phantom measurements in the left part of Fig. 4.2.9 (similar to
the measurements above in Fig. 4.2.7) were performed without a
scattering environment (in the air) and near the front of the
camera collimator. They simulate an idealized situation of superficial
lesions .
If the lesion is deposited at greater
depths in the tissue, four other adverse factors apply,
further reducing imaging contrast and impairing lesion
detectability :
v A greater distance from the collimator front leads to
poorer resolution (higher FWHM), which reduces the C img contrast in the
image according to the exponential dependence above .
v Absorption of g radiation from the
lesion as it passes through the tissues (attenuation) reduces the
number of useful pulses detected in the lesion image.
v Radiation from other layers of
tissue can be added to radiation g from the lesion. This
primarily reduces the contrast of the object C vol in the respective
planar projection and thus the contrast in the image. This effect
is largely eliminated in SPECT and PET tomographic imaging (see
§4.3 "Tomographic scintigraphy" below).
v Part of the gamma radiation is scattered Compton in the
tissue material. Part of this scattered radiation is detected and
also reduces the contrast of the lesion image (as shown above in
Figure 4.2.8).
In the right part of Fig.4.2.9 is a
phantom display of positive and negative lesions deposited on the
surface and at different depths in the tissue (simulated by water
with dissolved 99m Tc activity ). In deep-seated lesions, their appearance
deteriorates sharply, especially in the case of negative
("cold") lesions.
How
can image quality and detectability of lesions be improved ?
The recognizability of small structures (lesions) in
scintigraphic imaging is determined in practice mainly by the
following factors :
× Geometric size of the lesion;
× Accumulation of radioindicator in the lesion compared
to the surrounding tissue ® contrast of the lesion;
× Depth of lesion placement ®
attenuation of radiation, interference with radiation from other
layers;
× Spatial resolution of the scintigraphic system -
contrast in the image;
× Detection efficiency (sensitivity) + acquisition time ®
number of accumulated pulses ® statistical fluctuations.
The size and location of the lesion
is determined by the anatomical situation of the patient, the
resolution and sensitivity of the camera are basically determined
by its construction, but we can partially influence them by a
suitable choice of collimator. There are basically four ways in
which we can improve the image quality and capture of lesions :
l Increase the primary contrast of the lesion
This can be achieved in some cases by choosing a suitable
radiolabel, which is as selectively captured in
the diagnosed lesion.
l Increase the applied activity
of the radio indicator, which will increase the detected
number of pulses and reduce the relative statistical
fluctuations. However, this encounters the problem of increased
radiation exposure of the patient and, at high
activities, also for a dead time detection device.
l Increase the image acquisition
time ,
which proportionally increases the number of stored pulses in the
pixels of the image and reduces the relative statistical noise.
However, too long an acquisition time brings problems with the movement
of a patient who cannot stand motionless under the
camera detector for so long. In dynamic scintigraphy, this
solution is usually not applicable at all, because the
acquisition time of individual images is determined by the time
dynamics of the investigated process.
l Perform a suitable computer filtering of the
image ,
which can improve its quality and help identify minor defects - it is discussed in more detail in the discussion
" Filters and filtering of scintigraphic
images ". It is mainly optimized smoothing of
statistical fluctuations ( Low-pass filters - smoothing ) and artificial improvement
of resolution - resolution recovery
(" Bandpass filters - focusing ").
In general, the larger, more
contrasting, and shallower depths below the surface of the body,
the easier to see the lesion in the tissue.
Quantification of positive lesions on
gammagraphic images - SUV
One of the most common tasks of radionuclide gammagraphy is to
display the accumulation of a suitable
radioindicator in lesions (especially tumor) - not only to
recognize the lesion in the image, but also to determine the quantitative
rate of radioindicator accumulation in
the displayed tissue. A simple relative
criterion of the significance of the displayed lesion is the contrast
of the C img = (A * max - B *) / B* image
discussed above between the activity (accumulated number of
pulses) in the A * lesion image and the surrounding B *
background. To assess the severity of tumors in different
patients, as well as in monitoring the time course of tumor
size and metabolic activity in a given patient (most often monitoring the biological response
of tumor tissue to therapy) , the
degree of accumulation of the relevant radiopharmaceutical in
images from various independent scintigraphic studies should be
evaluated and compared . For the absolute
(semi) quantitative expression of the selective uptake of
the radioindicator in the tumor, in comparison with the
average distribution in the rest of the body, the so-called standardized
accumulation value SUV ( Standardized Uptake Value
) is often used.. It expresses the ratio of the local accumulated
concentration of the radioindicator in the lesion to the average
concentration in the whole body (ie to the
applied activity normalized to the patient's weight) :
SUV = C / (A inj / M).
Here, C [kBq / cm 3 ] is the tissue concentration of radioactivity (volume
activity) in the lesion, A inj [MBq] is the applied activity, M [kg] is the
weight (weight) of the patient. The values of volume activity of
lesion C and applied activity of A inj must be corrected at the same time (especially when using short-lived radionuclides such as
99m-Tc or 18-F) . Concentration C
radioactivity in the lesion is determined from the gamma image
using the appropriate conversion and correction factors :
C = h -1 . (A * -B *). RC -1 .V tum -1 ,
where h [imp.
s -1 MBq -1 ] is the detection
efficiency (sensitivity) of the camera , RC is the so-called
recovery coefficient of correction for the " partial
volume effect " (mentioned above
in the section " Adverse effects of
scintigraphy ", writes
" Partial volume effect ") , V tum [cm 3] is the volume of the lesion.
Therefore, if we measure the value SUV
= 1 at some point in the image , the volume activity is
the same as the average activity in the whole body - it means
that the radio indicator is not captured
here . The higher the SUV value > 1
we get, the more selectively the given
radioindicator accumulates in the given place, the higher
the metabolic activity of the respective tissue.
Either the SUV max
calculated from the A * max value of the most intense pixel in the lesion image is
used, or the SUV mean (SUV 50% ) determined from the average value in pixels within
the area of ??interest (ROI) of the lesion, sometimes the SUV 70%etc. If there is
otherwise an approximately homogeneous distribution of the
radioindicator outside the examined lesion, the SUV max is approximately
equal to the contrast value C img and other SUV 50 or SUV 70 values can be determined simply as the ratio of the
number of impulses in the tumor (or its defined part - ROI) and
in the tissue background (" tumor to background ratio
"). However, it is desirable to make a correction to the partial
volume effect using the RC recovery
coefficients (as mentioned above) .
SUV analysis is performed mainly on
PET images of 18 FDGs and other radiopharmaceuticals with tumor
accumulation - see also §3.6, section "Diagnosis
of cancer ". For medium storing
tumor lesion SUV value in the range of about 2 ¸ 5, for the well
and selectively accumulating bearings can then be SUV> 10.
Note .:
Quantification of SUVs with planar and SPECT
scintigraphy performed only quite rarely . SUV
domain primarily tumor scintigraphy, PET (see
below §4.3, section " Positron emission tomography
PET "), which is mainly
used to quantitate the accumulation of 18 F-FDG. Here we discuss the issue of SUV in
general terms in connection with common properties
scintigraphic images and the information contained therein.
Disadvantages and
pitfalls of SUV quantification
The determination of SUVs can be a useful tool for assessing the
severity (metabolic activity, possible aggression) of tumors and
the effectiveness of the biological response to their therapy. It
should be kept in mind certain drawbacks
consisting in a function values obtained SUV on
a number of factors and parameters:
¨ Particular, it is the exact actual
value of activity in relation to the calibration
of the meter applied activity and sensitivity
(detection efficiency) gamma camera. It also depends on the time
between application and examination, which by
radioactive decay and pharmacokinetics significantly affects the
amount of radioindicator accumulated in individual tissues,
including the examined lesions. Correction to this time can be
difficult because different types of tissues and tumors
accumulate the radioindicator at different rates . The
only way to minimize this time factor is to keep the same
time interval between application and scintigraphy. It
is also necessary to read the activity values ??remaining in the
syringe or tubing during application.
¨ Hydration and levels of metabolic
substances (eg sugars) in the patient's blood, functional state
of the kidneys, liver and other organs.
¨ It is also a dependence on the weight and body
constitution of the patient . In patients with the
higher the fat content , which accumulates very little in
the radiopharmaceuticals used, overestimates the
measured SUV values. This can be a problem when comparing
different patients with each other, or if a given patient changes
weight between exams. Correction of SUV to
patient weight can be performed approximately by normalization to
standard reference values of patient weight 70kg
and body surface area S = 1.75m 2 , using empirical
relationship between weight M , height H , body
surface area S and adipose tissue fraction: SUV M-corr
= C / (A inj ) .43.8.M 0.425.H 0.725 .0.0072. Due to this correction, the measured value of
SUV M-corr
in the shown lesions is lower in more massive patients than the
uncorrected value of SUV, on the contrary it is lower in more
subtle patients.
¨ Marking of areas of interest (ROI) of
examined lesions on the scintigraphic image is individually
dependent and is not very reproducible. SUV values (especially SUV mean ) are very sensitive to small
differences in the size and position of marked ROIs
.
¨ Absorption of gamma radiation in the tissue
, causing attenuation - attenuation - of the
signal from deeper lesions (see section
"Adverse effects with scintigraphy and their
correction "). The correction for attenuation is not always accurate
and reliable.
¨ Computer image editing - various kinds
of filters, methods and algorithms reconstruction by tomographic
images can significantly (and non-linear) influence the accumulated number of pulses in the
evaluation of lesions and tissue background. this leads to large
arteficikálním differences in measured values SUV.
¨ The effect of partial volume ( partial volume effect - as described above in
the passage " the volume and activity bias ") causes distortion
displayed lesions in terms of activity and size. to correct for
this effect use RC coefficients, which are
difficult to determine (obtained by phantom
measurements) and their values depend on
the imaging properties of a particular camera. To use them, it is
also necessary to know the diameter of the
displayed lesion.
Due to these difficulties in
determining a specific exact SUV value, this parameter is valid
only in comparative studies of larger patient
populations, where individual deviations and inaccuracies are randomized
. When comparing changes in scintigraphic images in a particular
patient, the SUV value (which in practice
cannot be determined with an accuracy of better than 30%) needs to be "taken with a grain of salt" !.
Author's
skeptical note on SUV:
The importance of "accurate" absolute quantification of
SUVs using all sophisticated correction methods is sometimes overestimated
. To gain my own experience, I would like to recommend the
following experiment to colleagues : Try to
compare the SUV values determined by the above
complex procedure with the values ??obtained from a simple ratio
of the number of pulses from the ROI in the lesion
image and the number of pulses in the ROI of a suitable reference
healthy tissue . The relative results will be very
similar , at least in terms of assessing the severity of
the metabolic activity of the tumor and the biological response
to therapy ..? .. - I welcome your experiences ...
Relative SUV
These pitfalls of accurate SUV determination show that identical
conditions cannot be maintained in paxi during repeated
scintigraphy of the patient before and after therapy. Therefore (in connection with the above note) the relative SUV rel
is introduced as the ratio SUV
rel = SUV tumor / SUV reference tissue , where all problematic values ??of applied activity,
detection efficiency, partial volume effect, patient weight,
application time are truncated . We basically
get the value of the tumor / background ratioexpressing
the relative rate of uptake of the radiolabel in the analyzed
lesion compared to the tissue background. The SUVrel can be obtained
from the scintigraphic image very simply by comparing the number
of impulses from the lesion ROI and the ROI of a representative
tissue background (eg liver or aorta ROI is
used as reference tissue; identical ROI must be observed when
repeatedly evaluating the same patient) .
Technical failures of scintillation cameras
With such a complex electronic device as a scintillation camera,
there are many possibilities for mild and more serious technical
failures. In terms of their location, we can divide them into two
groups :
1. Disruptions of electrical power supplies and
mechanical movements of the camera
Electrical power supplies for cameras are often burdened with
long-term power, they become hot, cooling fans "seize",
....
Current gamma cameras in their
electro-mechanical part contain a number of sensors, regulation
and control circuits, which is certainly correct in terms of
successful and safe operation of the device. Sometimes, however,
it is too "recombined", so that even an insignificant
deviation can lead to blockage of mechanical movements and thus
the practical usability of the camera, with the need for service
intervention.
2. Disorders of imaging properties in the field of
view of the camera
Practically all these disorders can be clearly seen in the image
of homogeneous distribution of gamma radiation (whether it is a homogeneous area or irradiation of a
crystal without a collimator with a point source from a
sufficient distance - see " Phantoms and phantom
measurements " , section
" Testing and calibration of camera
image homogenity ":
When properly functioning, the image of the homogeneous
distribution of radioactivity should also be homogeneous
(Figure a ), with the only tolerances arising from static
fluctuations in the number of pulses accumulated due to
quantum-stochastic gamma photon emission events.
The local circular outage (usually sharply demarcated, often with a visible rim) in the field of view is the result of interrupted
detection by scintillation from a specific location of the camera
crystal. The cause can be either a failure of the respective photomultiplier
, or its preamplifier or some other circuit
through which the detected pulses pass. In Figure b is
a failure of one peripheral photomultiplier. Repairing a
preamplifier is not a bigger problem. However, replacing a
defective photomultiplier is technically very difficult. After
the electrical disconnection, the photomultiplier must be
carefully "peeled off" from the sicon grease (ensuring
optical contact with the scintillation crystal light guide),
thoroughly clean the area and seal a specially selected
photomultiplier with the same properties as other
photomultipliers. This work will take an experienced electronics
all day, including subsequent adjustments and the resulting
calibrations of the camera detector.
A series of minor inhomogeneities in
Figure b , corresponding to the positions of the
photomultipliers, may not indicate a malfunction, but are usually
caused by misalignment - " detuning
""- position photopeak from individual
photomultipliers. Upon proper adjustment -" tuning
"( tuning ) and a new correction matrix homogeneity
generally receive homogeneous field of view.
The most serious disorder
scintillation camera is cracked crystal . The
image field of view appears as a distinctive irregular (tortuous
or branched) pulse failure line, with a positive rim (Fig. d )
This fatal failure can occur in basically two ways :
×
By mechanical pressure
or impact on a very brittle crystal.
It is enough for a screwdriver, phantom holder or other object
heavier than a few grams to fall on the crystal without a
collimator. When replacing collimators, the crystal may break
when there is a foreign object on the mounted collimator, for
example a pencil ..! ..
×
Thermal stress
when the temperature of the crystal changes unevenly or rapidly.
Larger temperature gradients due to thermal expansion can cause
considerable mechanical stresses in the crystal, which can result
in cracking. The crystal is particularly temperature sensitive
when the collimator is removed. In this situation, it is not even
recommended to open windows in the room or turn on the air
conditioning.
A cracked crystal is an irreparable
defect in a scintillation camera. The entire detector must be
replaced (crystal +
photomultipliers + preamplifiers) for a new
detector, assembled in the factory. This is an expensive
affair, about 200 thousands of dolars !
New
and alternative physical and technical principles of
scintillation cameras
Practically the only type of scintillation cameras used so far in
nuclear medicine are Anger-type cameras
described above (with the exception of PET cameras described below in §4.3 on tomographic
scintigraphy) . Despite the clear success
of the use of these cameras in nuclear medicine, two
basic disadvantages of this solution were also known
from the very beginning . The first is the need to use a lead
collimator , through which only the radiation g passes in a
precisely defined direction, but the vast majority of incident
photons are captured in the partitions between the holes Þ low
detection efficiency (sensitivity) of camera. The second
disadvantage stems from the limited accuracy with which a system
of photomultipliers and electronic circuits is able to locate the
position of a scintillation flash in a large-area scintillation
crystal Þ imperfect spatial resolution .
Therefore, since the 1970s, alternative
physical-technical solutions of scintillation cameras
have been designed and experimentally tested , eliminating the
first or second disadvantage, or both at the same time. These
alternative solutions have not yet gone beyond laboratory
experiments, but with the development of technologies in
the field of microelectronics and new materials, there is a real
hope in the near future to bring some of these structures into a
practically usable form, or even to replace existing
scintillation cameras in the more distant future.
Wireframe
cameras
Wireframe cameras are based on the simple principle of a
position-sensitive multi-wire ionization chamber
, which was developed for monitoring and displaying traces of
particles formed during interactions on accelerators (see §2.3,
section " Drift ionization chambers "). The
detector itself is made up of a large number (even several
hundred) of thin wires - electrodes stretched in a gas charge in
two layers in a mutually perpendicular direction - determined by
the X, Y coordinates. When a photon enters radiation gionization occurs at the appropriate site.
The electron cloud drifts from this point to the nearest
electrodes, where an electrical signal is generated. The
intersections of the electrodes thus received signal the location
of the interaction of the detected photon. The ionization cloud
of electrons can reach several nearby electrodes; the evaluation
electronics then determine the coordinates using the weighted
averages of the signals from the various electrodes. The point of
impact and interaction of the photon can be determined with an
accuracy of about 0.1 mm. Cameras of this type are especially
suitable for imaging with low-energy radiation g .
Semiconductor multidetector gamma cameras
One of the basic factors limiting the internal resolution of an
Anger scintillation camera is the uncertainty with which a system
of photomultipliers and subsequent electronic circuits is able to
locate the position of a scintillation flash in a large-area
scintillation crystal. Therefore, the internal resolution of the
Anger camera cannot be reduced below approx. 3 mm in practice.
The concept of a multidetector
camera is that instead of one large-area scintillation
detector equipped with a number of photomultipliers, many
separate miniature detectors - pixel semiconductor detectors (see §2.5 " Semiconductor detectors ") are used , placed in a matrix next to each
other. Gamma radiation is transferred directly
here to electrical signals without the need for scintillators and
photomultipliers. The signal from each of the detector is
processed independently (in
multiplexed mode) , whereby the positional
coordinates (x, y) are determined simply by the position (i, j) of the
mini-detector in the detector array, and lead directly into the
pixel array in the computer ( the pixel it pixel
) - fig.4.2.10 :
Fig.4.2.10. Principle of multidetector semiconductor camera.
Left, center: The crystal of a
multidetector camera consists of a large number of regularly
arranged miniature semiconductor pixel detectors. Right:
Special arc configuration of semiconductor CZT detectors and
multi-pinhole collimators for SPECT myocardium.
Photon detection is performed in individual
pixels independently, so the internal spatial resolution is given
by the size (pitch) of the detector pixels (unlike the Anger camera, where the coordinates of
scintillation are determined triangulation according to the
response of different photomultipliers) .
If a sufficiently dense grid of pixel detectors is created, we
can achieve a very good internal spatial resolution (even below 1mm) ; the total
resolution then depends on the collimator used. Optimized
collimators for multipixel semiconductor cameras should
have square apertures the size of pixel
detectors (minus the thickness of the
baffles) , which would geometrically
overlap with the detection pixels with their apertures everywhere
in the field of view.
So far, this type of camera has been
produced only with a small field of view of about 5 x 5 cm, for a unique
use for scintigraphy of small objects (small
laboratory animals) , now it is beginning
to be produced in the standard size of classic cameras. This
category also includes electronic imaging detectors for X-rays,
so-called flat-panels (described
in §3.2, section " Electronic
imaging of X-rays ", flat
panels with "direct conversion", which probably belongs
to the future ...) . Gradually, planar and
SPECT cameras of standard dimensions with semiconductor detectors
are also being used.
For this semiconductor
gammagraphy (planar and SPECT
"scintigraphy") , semiconductor
CZT (Cadmium-Zinc-Tellur ) detectors. Cadmium
and zinc CZT telluride is a semiconductor
detector operating at room temperature , which converts
gamma and X-rays directly into electrical impulses with high
efficiency (physical aspects see §2.5
" Semiconductor detectors ", passage " Cadmium-Zinc-Teluride
(CZT) detectors " ) . A comparison of the average basic parameters of a
standard Anger camera (with NaI
scintillation crystal (Tl) and photomultipliers) and a semiconductor camera with CZT detectors (2.5 mm in size) is in the
following table :
Camera type | Internal spatial resolution | Detection efficiency (for 99m Tc) |
Energy resolution | Max. pulse frequency |
Anger camera with NaI (Tl) | 4 mm | 60 cps / MBq | 10% | 3 . 10 5 cps |
Semiconductor CZT camera |
2.5 mm | 85 cps / MBq | 6% | 6 . 10 5 cps |
Thus, compared to conventional Anger cameras, semiconductor
DHW cameras have better spatial resolution
and energy resolution, slightly higher detection efficiency
(sensitivity) and shorter dead time of detection.
The use of CZT
detectors for positron emission tomography of PET
is also promising , instead of BGO / LSO scintiblocks with
photomultipliers (see below " Positron
emission tography of PET
", Fig . 4.3.5) . In addition to better detection efficiency and spatial
resolution, a somewhat shorter coincidence time can be achieved (for better TOF) . So far, it is
being tested experimentally on smaller PET models. The advantage
of semiconductor detectors is also theirsindependence
from the magnetic field , which allows use in hybrid PET
/ MR systems .
In nuclear cardiology, stationary semiconductor
CZT ( Cadmium-Zinc-Tellur id) cameras with a
special "cardiofocal" detector arrangement are
beginning to be used for SPECT of the myocardium
( Fig. 4.4.10 on the right). The detectors are placed in
the camera gantry along an arc covering an angle
of approx. 90-180 °. The detectors are equipped with " multi-pinhole
" collimators directed cardiofocally into the center of the
gantry. Unlike the classic rotary SPECT (described
below " Tomographic scintigraphy SPECT ") , data storage takes
place stationary, detectors and collimators are
in a fixed position relative to the patient 's body, all SPECT
projections are obtained simultaneously . This
achieves higher detection efficiency and faster processing.
However, it is a single-purpose device for SPECT
myocardium in cardiology.
Advanced universal
stationary semiconductor SPECT cameras are being developed
- see below " SPECT Stationary Multidetector Cameras ".
Compton cameras
In the paragraph on adverse effects of scintigraphy, we
classified Compton scattering of g- rays in tissue as an adverse
event that worsens the quality of scintigraphic images.
However, with the appropriate mechanical configuration and
electronic interconnection of two or more detectors, Compton
scattering g in the detection system itself can be used for " electronic
collimation " and g field imaging without
the use of mechanical collimators
(using the Compton scattering for gamma imaging, suggested
Everett, Fleming, Todd and Nightengale in 1977) . The principle of operation of such a so-called Compton
camera is schematically shown in the following figure
4.2.11 :
Fig.4.2.11. Schematic representation of the principle of
electronic collimation using energetic-angular reconstructions of
the paths of primary ( g ) and Compton scattered ( g ') gamma-ray photons.
The camera itself consists of two (or several)
consecutive detectors providing positional and energetic
information about the detected quantum g :
In the first thin detector 1
(replacing the classical lead collimator) there is a Compton
scattering of photons of incoming radiation g (by different angles J ) , which then
continue their movement to the second more massive detector 2
, where they are fully absorbed.
In the coincidence mode, the positional coordinates of the impact
of the primary photon g (x 1 , y 1 ) and the energy E 1 are detected.transmitted to the electron at Compton
scattering in the first detector, as well as the positional
coordinates of the impact (x 2 , y 2 ) and the energy E 2 of the Compton scattered photon g 'absorbed in the second
detector. Based on the geometric comparison of the positions (x 1 , y 1 ) of the primary and
(x 2 , y 2 ) scattered gamma
photons, the angle J of the compton scattering is determined . This angle J is then related to
the energy E 1 of the Compton scattering and the energy E 2 of the scattered
radiation g', which allows (according to the relation for the
angular-energy distribution of Compton scattered radiation E g ' = E g / [ 1 + (E g / m oe c 2 ). (1 - cos J ) ], given in §1.3) to kinematically reconstruct
the path of the photon to determine the incidence angle j at which the
primary photon g flew to the first camera detector from its source.
Photopeak measurement E g = E 1 + E 2 then it makes it possible to
eliminate those unwanted photons which were scattered by Compton
before coming to the first detector, similarly to Anger's camera.
This creates an incident
cone with a vertex at (x 1 , y 1 ) and an apex angle J
, on the mantle of which lie the possible
trajectories of the incoming photon. The set of these incident
cone shells from individual detected photons can then be used for
computer reconstruction of the resulting
scintigraphic image of radioactivity distribution in the scanned
object: pixels corresponding to the intersection of individual
cones (ellipses, circles) plane(Fig. 4.2.11
on the right is an example of the reconstruction of the image of
a point source, arising as an intersection of elliptical
projections of incident cones of photons emanating from this
source) .
In the scattering detector 1
, a multidetector system of semiconductor detectors Si, CdTe or
GaAs with a thickness of about 5 mm is used, a high effective
cross section for Compton scattering is required here. The
absorption detector 2 can be an Anger crystal
system NaI (T1) or BGO or LSO with photomultipliers and
electronics evaluating the position of the flashes. Recently,
however, even in this second detector, the Anger camera has been
replaced by a semiconductor multicrystalline detector. In
addition to spatial and energy resolution, high demands are also
placed on the good operation of the Compton cameratime
resolution of coincidence (similar to PET detectors -
see §4.3).
Compared to mechanical collimators, electronic
collimation can lead to a significant improvement
in detection efficiency (sensitivity), as g photons are used
from a much larger spatial angle (electronic collimation, but of
a different kind, is of great importance in positron emission
tomography, see PET below ).
Apart from laboratory experiments, Compton's
cameras have not yet been implemented ,
they will probably remain only a physical-technical interest ....
High-energy gamma cameras
The need to imaging the distribution of high-energy g radiation arises
mainly in two areas :
×
1. Gammagraphic imaging of the distribution
of radioactive substances emitting hard gamma radiation
(their distribution in samples, tissues and organs), or depicting
the distribution of atomic nuclei excited by external
radiation that emit high-energy radiation g during
deexcitation (such as the NSECT method - see "Neutron- stimulated
emission computed tomography "
below). However, this is a relatively marginal issue ...
×
2.
Gamma-telescopic
imaging sources of gamma radiation in space -
supernovae, neutron stars, accretion disks in black holes (see eg " Astrophysical significance of
black holes " in the book
Gravity,
black holes and space-time physics ) and other turbulent astrophysical processes; on g radiation from
space, see also §1.6 " Cosmic radiation ",
section "Cosmic X and gamma radiation".
Imaging with high-energy gamma rays - hundreds of keV to
tens of MeV - is much more difficult than with soft g -radiation (60-500keV). For such energies, the
collimators have poor spatial resolution and luminosity due to
the significant irradiation of the septa between the apertures,
and the scintillation crystals of standard gamma cameras used are
too thin to achieve reasonable detection efficiency. A suitable
solution here is the above-mentioned principle of the Compton
camera , in a modification optimized for high energies.
A simpler type of Compton telescope , used on space
stations to detect g- rays from space sources, consists of a larger ionization
chamber (drift wire or projection) in which the energy of
scattered g- rays and reflected electrons and the direction of
scattered radiation or reflected electrons are measured .
Fig.4.2.12. Some principles of gamma cameras for high energy.
Left: Combined Compton-Anger high
energy gamma camera. Right: 3-Compton
gamma-telescope with many detection layers.
Fig. 4.2.12 on the left
schematically shows the principle of operation of a combined
Compton-Anger gamma camera for high energies. Deteèní camera
sensitive volume consists ionization drift-time
projection chamber (TPC - Time Projection Chamber
) with a gas filling (the ionization detector, see §2.3 " Ionization Chamber "). When a high-energy photon g flies into this
working space , Compton scattering in the gas
charge occurs , and for higher energies, the formation of
electron-positron pairs., followed by annihilation of a
positron with an electron to emit a counter - pair of gamma
photons with energies of 511keV. The path of reflected or paired
electrons is sensed based on the ionization electrons that these
high-energy particles generate along their paths. They are
detected by a matrix of several hundred miniature pixel
ionization chambers, working in proportional or Geiger
(avalanche) mode. Or semiconductor detectors can be used. This
cell matrix forms a 2-D position sensitive electron detector. The
ionization electrons from the individual paths of the fast
charged particle drift into different chambers (perpendicular
projection of the path into the nearest chambers) for different
lengths of time; by evaluating these geometric and temporal data,
the 3-D path of reflected or paired electrons and positrons in
the chamber space can be reconstructed. The working chamber is
surrounded on all sides by scintillation crystals with
photomultipliers ( Anger camera), scanning
Compton scattered and annihilation photons, with scintillation
positioning and radiation energy. By a complex coincidence
evaluation of pulses from the matrix of pixel detection chambers
and from the photomultipliers of the Anger camera, it is then
possible to geometrically reconstruct the direction
(angle) from which the detected primary high-energy photon g arrived - to
realize gamma-ray imaging .
Fig. 4.2.12 on the right shows the
principle of a gamma-telescope based on repeated Compton
scattering in layers of position-sensitive semiconductor
detectors. The system consists of several layers of flat
position-sensitive (2-D) detectors, stacked at equidistant
distances. After the entry of the primary g -photon with energy Eg 1 is scattered by Compton in one of the detectors, which
is accompanied by a position pulse and an amplitude pulse
carrying information about the energy loss D E 1 that the photon left in the detector during scattering.
The scattered photon continues to fly at an angle J 1 with energy E g
2 =
E g 1 - D E 1 , after which it can
be further scattered by Compton in another detection layer,
providing the appropriate position pulse and energy pulse D E 2 . The photon scatters
by an angle J 2 and continues with
the energy E g 3= E g 2 - D E 2 . Thus, repeated multiple scattering can occur until
the photon leaves the detection space. Coincidence evaluation of
position coordinates in individual detection layers determines
scattering angles J 1 , J 2 , J 3 , ...., evaluation of
pulse amplitudes determines energy losses D E 1 , D E 2 , D E 3 , ... These data substitute into Compton 's equations
Eg2 = Eg1 /[1 + (Eg1 /moec2).(1 - cos J1)] ; Eg1 = DE1
+ Eg2 = DE1 + {DE2
+ [DE22+ 4moec2.DE2/(1-cosJ2)]1/2/2} , .... ,
which allows kinematic and geometrical reconstruction of the
photon path - will provide the required value of the angle J of the incident
cone, under which the primary g-
photon flew into the detection system . And
further reconstruction by intersecting a set of projections of
incident cones of all registered photons, the resulting g- telescopic
image of the source from which the photons were emitted
is obtained. The advantage of this arrangement is that to
reconstruct the angle of incident g-photon does not need its
complete absorption in a heavy "calorimetric" detector
to determine the total energy. The energy of the incident photon
is determined by measuring the position of the first three
interactions and the energy delivered in the first two
interactions.
Thus, it is sufficient to obtain at
least a 3-fold Compton scattering, the analysis of which can be
used to reconstruct the incidence angle J - hence this system is
sometimes referred to as the 3-Compton telescope
. Analysis or. further dispersions refines the reconstruction. In
the individual layers of flat position-sensitive detectors it is
possible to use either ionization wire chambers, or better
germanium or silicon semiconductor drift detectors
, which have good energy and image resolution.
To imeging gamma radiation very
high energies , hundreds of MeV to hundreds of GeV,
special particle detectors of electron-positron pairs are
used in an arrangement similar to Fig.4.2.12 on the
right. The g- rays first fall on a plate of heavy material (tungsten),
where they are converted into electron-positron pairs, flying
almost in the direction of the original photon g . Their paths are
then monitored by layers of position-sensitive 2-D silicon
detectors (trackers), which determines the direction from which
gamma radiation came. Finally, they transfer their energy to a
calorimetric detector located below the last detection layer,
which detects the energy of the g- quantum.
4.3.
Tomographic scintigraphy
Every living organism is a three-dimensional
object and the distribution of a radio indicator has the same
character. A planar scintigraphic image, which is a
two-dimensional projection of reality, can therefore capture only
part of reality. We cannot find out anything about the
distribution of the radio indicator in the "deep third
dimension", perpendicular to the front of the collimator,
from the planar scintigraphic image. Planar scintigraphic images
have serious pitfalls in this respect - the possibility of overlapping
and superposition of structuresstored at different
depths. 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. The superposition of radiation
from different depths of the imaged object further leads to a reduction
in the contrast of the image of the lesions, which are
overlapped by radiation from the tissue background in the planar
image.
To overcome these disadvantages of planar
scintigraphy and to obtain a complex image of structures at
different depths, tomographic scintigraphy *)
has been developed to provide a three-dimensional image
of the radiolabel distribution. One of the main advantages of
tomographic imaging is significantly higher contrastimaging
of lesions (up to 10-fold) that do not overlap with tissue
background radiation on transverse sections.
*) Greek tomos
= section - the tomographic image consists of certain sections
, mostly transverse, a larger number of which create a
three-dimensional image.
Some basic principles, especially geometric
and reconstructive, have all tomographic methods in common. X-ray
transmission tomography CT was described in §3.2 "X-ray
diagnostics", part " Transmission
X-ray tmography (CT) ",
where the development of tomographic methods in general is also
mentioned.
Technical development of gammagraphic
tomography
Efforts to achieve in-depth tomography imaging began shortly
after the introduction of scintigraphy in the 1960s and 1970s.
Predecessor present gamagrafické tomography SPECT was in the 70 motion
tomography (obr.4.3.1 left): the examining table with
the patient and the collimator camera with inclined holes ( slant
poles ) using the electromotor synchronously rotated
in such a way that, for a layer in a "focal" depth,
both movements were compensated and a sharp image was
obtained, while in the other layers (above and below the focal
plane) the image was motion blurred
and thus it was darker and less distinct. Against the background
of these blurred and darkened areas, sharper and more clearly
displayed structures from the focal plane were better visible.
The depth position of the focal plane was set by the radius of
rotation of the lounger on the eccentric of the lounger motor.
However, the quality, contrast and depth effect of such an image
were not great (completely incomparable with SPECT). An image of
only one longitudinal layer was obtained at a time , in
order to create an image of another focal layer, it was necessary
to change the radius of the sliding rotation of the lounger and
start a new acquisition. More detailed tomographic imaging in
multiple layers was therefore time consuming. This method has
long been abandoned.
Fig.4.3.1. Early attempts to implement tomographic scintigraphy.
Left: Movement tomography with rotating slant-hole
collimator and rotating lounger. Middle:
Coincidence tomography using g-g
angular correlation. Right: SPECT
on a stationary planar gamma camera with patient rotation.
Interesting experiments
were also performed with 2-photon coincidence tomography
using g-g angular correlations between
the emission of cascade pairs of gamma photons in some
radionuclides (§1.2, part " Gamma radiation
", passage " Angular correlations of gamma
radiation ") - Fig.4.3.1 in the middle. Another gamma detector, or in
a more advanced version another gamma camera with a slit
collimator, in a coincidence connection was attached to
the basic imaging gamma camera at a certain place and at a
suitable angle q (eg 90 °) . Only pulses originating from the current
one were registered to create the imagethe arrival of a pair
of cascading gamma-quantum in both detectors. Thus, only one
longitudinal thin layer was displayed , defined by the
detection angle of the auxiliary coincidence detector (or a larger number of independent layers when using a
coincidence camera with a slit collimator) .
However, the range of suitable isotopes exhibiting cascade
deexcitation with angular correlation of the gamma-photons
emitted is very limited and the detection efficiency of the
coincidence system has been very low. This method did not go
beyond laboratory experiments; however, in a sense it can be
considered the ideological forerunner of positron emission
tomography. Namely, only a perfect angular correlation of 180
° between a pair of annihilation photon sin
electron-positron annihilation, it has found wide application in
coincidence positron emission tomography , see PET below .
The first attempts at SPECT
gamma tomography were performed in the
1960s and 1980s at a number of workplaces with planar Anger
cameras (the first tomographic image was
presented by Kuhl and Edwards in 1963) .
Since the (planar) cameras at the time did not have gantry and
could not rotate, the patient turned-
fig.4.3.1 right. In front of a vertically set stationary camera,
the patient sat on a swivel chair with a marked angular scale
(goniometer). A planar image was taken, the chair and the patient
were rotated by a certain angle, another image was accumulated,
etc. - approx. 16-64 images for a sequence of angles 0-360 °.
This was followed by computer reconstruction by back
projection into transverse sections . This method,
despite its mechanical clumsiness, has in fact already enabled
full-fledged SPECT imaging - with the limitation that at that
time not yet sufficiently complex software for the reconstruction
of transverse sections and their processing had been developed.
The assembly of gamma cameras on gantry
with camera rotation then became a truly successful and
routinely used SPECT tomography method.around the patient (as
described below). Recently, stationary SPECT multidetector cameras have been developed without rotation, which is likely
to displace the clumsy rotating SPECT.
Author's note:
In the 1970s and 1980s, we also performed early attempts at
tomographic imaging according to Fig. 4.3.1 at our Department
of Nuclear Medicine KNsP in Ostrava-Poruba . Our
first Pho Gamma HP gamma camera (Nuclear Chicago) with
the CLINCOM evaluation device from 1974 was equipped with a
rotating Slant Hole collimator and a rotating lounger
for motion tomography (Fig. 4.3.1 on the left).
Experiments with gg coincidence tomography we performed on a Pho
Gamma HP camera with the help of a perpendicularly oriented
collimated scintillation probe, connected in coincidence to the
flow of scintigraphic pulses (we did not have an additional gamma
camera with a slit collimator). We also tested the improvised
SPECT with a patient in a swivel chair on a Pho Gamma planar
camera, with the then latest computer evaluation device GAMMA-11
and our own developed software. Unlike
X-ray transmission tomography CT, where the image is created by
the passage - transmission - of X radiation through the
body, scintigraphic tomography creates the image by detecting
radiation emitted - emitted - from a
radio indicator inside the body. Radionuclide emission
computed tomography ( ECT ) is of two types :
1. SPECT single-photon
emission tomography using g- radionuclides registers only one emitted
gamma radiation photon from each radioactive transformation .
2. Two-photon positron emission tomography of PET
using positron (beta + ) radionuclides, resp. the resulting annihilation
radiation, where two photons emitted during
annihilation are always detected at the same time (coincidence) .
In both cases, the resulting
tomographic images are obtained by computer
reconstruction after scanning the pulses from the photon
detection. We will discuss both tomographic methods in this order
1. , 2 .
Tomographic
scintigraphy SPECT
The most common method of
tomographic scintigraphy is the so-called single-photon
emission computed tomography SPECT ( Single
Photon Emission Computerized
Tomography). Its principle is shown in Fig.4.3.2
:
Fig.4.3.2. The principle of capturing
scintigraphic images of the examined object W (here the
brain) at different angles by a rotating
SPECT camera and their computer reconstruction into the resulting
image W* of a cross section of this object.
The basic
principle SPECT
Tomographic camera SPECT its construction from the conventional
planar cameras differ only in that the special rack
in which the detector camera is fixed, so-called. gantry
circular shape (Gantry = portal,
through the support structure ) ,
allows the motor-driven rotation of the detector
around of the examined object *) - the
photograph of the SPECT camera is above in §4.2 in Fig.4.2.4 at the bottom right . To speed up
the acquisition, two detection systems ("double-headed"
SPECT camera) are most often installed on
the common gantry , there were also cameras with 3 or 4 detectors
(it did not prove in practice...) .
*) The technical construction of
a SPECT camera without a gantry is rare . The
camera detectors are mounted on special arms equipped with servomotors
, allowing the detectors to move in
space in different directions and at different angles -
with all "degrees of freedom". By suitable electronic
control of the servomotors, it is then possible to achieve a
circular movement of the detectors around the examined object
(around the bed with the patient) during SPECT.
The new stationary SPECT multi-detector cameras have a completely different principle .
Acquisition of SPECT
SPECT own tomographic scintigraphy
then consists in the fact that the camera gradually revolves
*) around the examined object and at a number of
different angles captures (planar) scintigraphic images
of the examined object - the number of these projections
is usually 32, 64 to 128 images at angles 0 ° - 360 ° - Fig.4.3.2
left. [ Note: In some cases, a smaller
range of angles is used - some projections, the quality
of which would be degraded by increased absorption (attenuation)
of g- radiation
and would not contribute to the resulting reconstructed images
(they could rather cause deterioration), are not captured. Such
is the situation with myocardial SPECT , which is sensed by a range of angles from 90 to 270
°, while the angles between 0-90 ° and 270-360 °,
corresponding to the rear and right side projections, the images
due to a significant attenuation of radiation g are not taken.]
*) SPECT stationary multi-detector cameras without rotation are mentioned below .
Orbiting
detector camera around the object to be examined is usually a stepper
( step-by-step ) - camera brook a certain angle, and
stops after a preset period the acquisition respective
projection, then again rotated through a given angle step and
charging proceeds next image. Continuous
rotation of the detector with continuous acquisition is rarely
used. From the geometric point of view may flow detectors circular
( circular ) - which is used more often, or elliptical
( non-circular ), if event. using the " auto-countouring
" system " (mentioned
above in the section" Design of scintillation cameras ") , in order to better
"copy " the body surface and keep the shortest possible
distance of the camera collimator face from the displayed
structures (to achieve the best possible
spatial resolution) .
Collimators for SPECT
Orbiting cameras during the acquisition of SPECT are usually
equipped with standard collimators with parallel holes
, the same as used in planar scintigraphy: for 99m Tc it is mostly HR collimator , for 131I HE
collimator , for 123I MediumEnergy collimator . Sometimes they
are used special collimators with a different
hole geometry: the FanBeam collimator for SPECT
of the brain, or a convergent collimator for
SPECT of the myocardium. All of these types of collimators have
been described in more detail above in the " Scintigraphic Collimators ". For stationary multidetector SPECT cameras are used Pinhole type
collimators (or special mechanically
movable collimators which, however, it is only a temporary
solution ...) .
Reconstruction of SPECT images
From this series of planar scintigraphic images - projections -
taken at different angles (these are planar
projections of the distribution of the radio indicator to
different angles) are then computer
reconstructs the image of the distribution of
radioactivity in the imaginary cross section ,
guided by the examined object in a plane perpendicular to the
axis of rotation of the camera - Fig.4.3.2 on the right. SPECT
computer reconstruction methods are described below in the
section " SPECT computer reconstruction ".
Such a reconstruction can be
performed for each row of the image matrix of angularly scanned
images, so that a whole series of "stacked side by
side" cross-sectional images is created in the computer's
memory - a kind of three-dimensional "cylinder" (for cameras with circular field of view) or "cube" (resp. square
- for cameras with a quadrangular field of view) , representing a three-dimensional image
distribution of a radio indicator in the examined object. The
cells of this three-dimensional image already have a volumetric
character and are called " voxels " (volume- volume -pixels) .
With this three-dimensional image in
computer memory, we can use computer graphics
methods to guide and display sections in any direction
on the monitor - not only primary transverse sections, but also
longitudinal and oblique sections, we can make various geometric
reorientations and other adjustments to show the desired
structure as clearly as possible. . Using computer graphics
methods, three-dimensional 3D images can be
created using suitable shading and perspective angular display
with a number of computer effects, which are artificial and may
not directly reflect reality, but are very illustrative and
effective, also for didactic purposes.
Example of 3D-imaging in myocardial scintigraphy.
SPECT
stationary multidetector cameras
The basic disadvantages of standard rotational
scanning of scintigraphic projections at SPECT are clumsiness
, slowness and low detection efficiency
. At each angle, only a small portion of the gamma photons (going in the direction of the current position of the
detectors) are registered , the other
photons emitted from the patient's body are lost. The detectors
slowly move to more and more angles; with a very long acquisition
time, the number of accumulated pulses in the images is
relatively low. We will obtain the required SPECT images only ext
post , after the measurement and computer
reconstruction. For the rotation methodit is not possible
to perform dynamic SPECT scintigraphy ( except for myocardial perfusion, where this is allowed
by ECG-gating) , or to operatively modify
the acquisition procedure.
Fig.4.3.3. The stationary SPECT camera
detects projections from all angles simultaneously using a large
number of circular multipix detectors. There is no any rotation.
The resulting data can then be continuously reconstructed into
cross-sectional images.
To eliminate this disadvantage of
rotary SPECT, it is offered to use a larger
number of smaller imaging stationary detectors
- gamma cameras, placed in a ring around the
examined object, without rotation - Fig.4.3.3
left. All projections are then shot simultaneously from
all angles . The camera is compact , it
does not contain any moving parts, the use of photons for imaging
is much more efficient - even with lower applied activity, SPECT
examination takes significantly less time. Interfering mechanical
phenomena, such as displacements of the center of rotation, do
not apply here. Cameras with multiPinhole
collimators, multiDivergent , or multiParalel
collimators are being developed.
Here, SPECT tomographic images can
be reconstructed and displayed continuously
during the acquisition, Fig. 4.3.3 on the right (similar to planar
scintigraphy and PET ) , which also enables dynamic
SPECT scintigraphy. This type of camera is currently
being used in nuclear cardiology for SPECT myocardial perfusion (§4.9.4, section " Scintigraphy
of myocardial perfusion
", see also section " Semiconductor multidetector cameras " above) . Stationary
compact SPECT cameras with multipixel semiconductor
detectors (described above " Semiconductor multidetector
gamma cameras")
undoubtedly belongs to the future , surely
sooner or later it will push out cumbersome
SPECT cameras with rotating detectors ..! ..
Computer
reconstruction of SPECT
Sinogram,
Linogram
If we successively draw images stacked at different angles J , the individual
points of the object in them will describe circles
with different radii (according to the
distance from the center of rotation) . X
and Y - coordinates of these circular orbiting points will
describe sinusoidal (or cosine) curves
R.sin J with amplitude
given by distance R from the center of rotation. Their
brightness is modulated by the accumulated number of pulses in
individual pixels of angular projections. The set of all these
graphically represented coordinates of the circulation curves for
all points in a given section creates an important 2D image
called a sinogram.. It is created simply by
gradually taking selected lines (corresponding
to a given transverse section) from individual projection images
and storing them "on top of each other" in a new image
- a sinogram. This is done for all projections (angles J ). Sinogram has
two roles in tomographic scintigraphy :
1. It is used as a data format
( sinogram-file ), into which the acquisition of primary
data from individual projection angles takes place at predefined
times per image. Some new gamma cameras (especially PET) can
operatively store data even in LIST mode (without
predefined time per frame), from which the sinogram can be
additionally created, with the possibility of computer editing.
Each row of the acquisition matrix - each transverse section -
has its own sinogram. Cross-sectional images (using the inverse Radon transform) can then be reconstructed from the sinogram data .
2. The sinogram display allows you to check
the correct course of the tomographic examination. Under
normal circumstances, the sinogram of each displayed active site
must have a smooth uninterrupted course ("wavy line").
Possibly. the patient's movement (which may lead to deterioration or artifacts) during the acquisition of SPECT is clearly seen on the
sinogram as a discontinuity smooth course.
Sinograms are also used to test mechanical disturbances during
the rotation of detectors, such as displacements of the center of
rotation (see below " Adverse
effects of SPECT and their correction ", section " Mechanical
instability of the axis of rotation ") . Based on sinograms, unwanted movements can be corrected
by software .
![]() |
Examples of sinograms
(top) and linograms (bottom) in SPECT
scintigraphy. In the middle part there are transverse
sections . a) A point source at a distance R from the center of rotation on the sinogram describes a sinusoidal curve x = R.sinJ. The weak auxiliary source at the center remains motionless on the axis of rotation (x = 0). b) The displacement of the source in the radial direction during the acquisition is reflected in the discontinuity on the sinogram. c) Sinogram of 5 active lesions in Jasczak phantom. d) Example of sinogram and linogram of the SPECT brain for imaging dopamine. receptors (" Scintigraphy of receptor systems in the brain ") |
In tomographic imaging methods (SPECT, PET, or
MRI), a so-called linogram is sometimes
constructed - an image in which the summed rows from all primary
images are "stacked" as columns linearly
side by side. In SPECT, a linogram is created by summing all
rows in each of the projection images at individual
angles and saving the resulting summation row as a column
in a new image - the linogram. This is done for all projections.
Unlike sinograms, of which there are a large number (for each
transverse section), the linogram is one for the
entire tomographic examination. Computer methods for
reconstruction of cross-sectional images (inverse
Radon transformation) have been developed .by integration along the lines in
the linogram. In some cases, the linogram may also be used to
assess the smooth running of the SPECT examination with respect
to movement artifacts or transient electronic disturbances in the
acquisition process.
Note: Universal tomographic
regularities
The regularities and relationships outlined in this part and the
relationships between circular planar projections at angles J and transverse
tomographic sections (sinograms, Radon transformation,
reconstruction methods, formulas in Fig. 4.34) apply not only to
SPECT scanning by physical camera rotation, but also with
stationary SPECT, with PET, CT, or NMRI. They are basically universal
and are used in various display modalities.
SPECT reconstruction methods
The amount of data accumulated in
individual projections at different angles implicitly
contains information about the spatial depth distribution. In
order to be able to explicitly display the depth
spatial distribution of the radio indicator in cross sections, it
is necessary to perform a computer reconstruction of the
accumulated "raw" data. They are two methods of
computer reconstruction accumulated planar images from different
angles to the desired transversal sections :
1.
Back-projection method
The analytical back-projection method by computer simulates
an inverse acquisition process SPECT study: As if the
camera detector emitted rays of radiation - of an intensity
modulated by the image (accumulated information in the cell) -
from each position (angle where it was rotating and accumulated
the relevant image) and from each of its cells back towards the
object under investigation, where these rays "draw" a
cross-sectional image in an imaginary image matrix located at the
center of rotation. The information contained in one given pixel
of the image stacked from a certain angle is transferred to all
the pixels of the emerging cross-sectional matrix located in a
line perpendicular to the detector. Different "intense"
rays from different angles then "irradiate" and
"fit" individual elements (cells, pixels) with
differently large numbers as they pass through the created image
matrix, which add up when passing through other
rays (from other angles). In places where most rays of higher
intensity pass, "hot" places with a high accumulation
of impulses are created - they correspond to places with a higher
concentration of radioindicator in the examined object.
Fig.4.3.4. The process of acquisition of SPECT and reconstruction
of the transverse section by the method of filtered back
projection.
The back projection therefore uses the back
projection of data from individual planar images into the
originally empty matrix, always at the angle at which the planar
image was created. The resulting matrix - the reconstructed image
- is created by direct addition of these projection data. Simple
rear projection has the disadvantage of a higher disturbingly
structured background with the formation of
"star-shaped" artifacts (see below). In practice,
therefore, filtered back projection FBP ( Filtered
Back Projectoion ) is used, which is a variant of the inverse
Radon transformation . The relevant mathematical formulas
are shown in Fig. 4.3.4, where the whole process of acquisition,
filtration and reconstruction of SPECT is shown from a mathematical point of view (3rd dimension is omitted) .
Examined object (patient), whose cross sectionhas
a distribution of the radio indicator A (x, y), it is captured by
the camera in a series of projections at
different angles J , thus creating images of projections p (u). These
images are then Fourier transformed into the frequency domain and
the resulting spectra p ( n ) are multiplied by a filter composed of a RAMP filter
and a user filter (see " Filters
and filtering ") . The resulting filtered spectra p F ( n ) are then converted back to the spatial region by
inverse Fourier transform (filtered images of projections p F (u) are formed ),
after which with the back projection (at the
same angles J ) the resulting cross-sectional image
A´f (x,
y) is formed.
The filtered back projection method
is the most used because it is relatively fast (fast Fourier transform algorithms are used, the values
of trigonometric functions are calculated in advance for discrete
values of angles, so that common arithmetic operations are then
used) . However, in terms of the
relationship between the actual distribution of radioactivity A
(x, y) and the reconstructed cross-sectional image A´ (x, y), it
is not exactly a mutually unambiguous representation - the image
is constructed not from local values in pixels, but by superposition
of projection rays . These projection beams are artificial
and leave traces in the resulting image that do not
correspond to the actual distribution of radioactivity in the
object under investigation. This is most pronounced in the
vicinity of deposits of increased deposition of radioactivity,
where converging projection beams form a " star-shaped
" artificial structure - the so-called star
effect . Although this star artifact is effectively
suppressed by a RAMP filter , various
"noodles" or "filaments" are always visible
in the images reconstructed by rear projection (below in the
figure on the left). These disadvantages of retrospective
projection are largely eliminated by iterative
reconstruction method. The RAMP filter, which also acts as a
"focusing" (emphasizes details and edges in the image),
is used in the reconstruction in combination with a user
"smoothing" filter to reduce statistical fluctuations,
noise, in the image. By a suitable choice of this filter and its
form-factors, it is possible to achieve optimal contrast,
detailing and noise reduction (it is
discussed in more detail in the work " Filters and
filtration ") .
2. The
method of iterative reconstruction
Iterative method *) of reconstruction looking through successive
steps - approximations such a cross-section image that
would best suit the individual scanned projections at different
angles J . It is an algebraic reconstruction technique
(ART).
*) The Latin word " iteratio
" means " repetition "; these are
recurring cycles of successive approximations.
Iterative reconstruction takes place
in the following stages :
The iterations are repeated until a certain
(preset) convergence criterion is met , such as
the required accuracy or a preset number of iterations.
It could be expected that as the number of iterations
increases, the overall image quality will increase. However,
experience shows that this is true for about 4-8 iterations
. A higher number of iterations then only increases the
statistical fluctuations, i.e. the signal-to-noise ratio
worsens .
Compared to the
back-projection method, the iterative method has the basic
advantage **) that no star artifacts are formed
(RAMP type filters are not used here). Also in areas with low
radioactivity (near background), the cross-sectional images are
"cleaner" and more contrasting - they do not contain
"filaments" as remnants of backscatter beams. Another
advantage of iterative reconstruction is the possibility of
introducing some corrections directly into the
reconstruction algorithm - correction for collimator properties,
dependence of resolution on distance from collimator ...
**) When processing SPECT images
in routine clinical practice, however, "no wonders" -
the difference from the back-projection method is often not even
noticeable, as the image quality is primarily due to insufficient
statistics, camera resolution, scattering and other disturbances
(mentioned below) with which no reconstruction method "will
do nothing"...
The iterative method of reconstruction is
much more demanding on the number of arithmetic operations, so it
could be routinely used only with the development of sufficiently
fast computers (using coprocessors) with a high memory capacity.
Improved variants of
iterative reconstruction
In order to streamline and speed up iterative reconstruction,
some newer variants and modifications of the basic iterative
procedure have been developed : EM (Expectation
Maximalization) - finding the best estimate of the image by
statistical methods ........
ML (Maximum Likekihood) - estimating maximum
likelihood principle ......
MLEM ( Maximum Likelihood
Expectation Maximization) - iteration
procedure with a preset number of iterations: before the start of
the reconstruction, the number of iterations is preselected for
which we assume the optimal image quality.
OSEM (Ordered-Subset Expectation Maximalization) - the
set of all projections is first regularly divided into several
smaller groups ( subsets = subsets ) and the iteration
step is applied to individual subsets separately. The
sub-iteration of each subset serves as an input estimate for the
iteration of the next subset. One complete iteration step is an
iteration cycle across all subsets. The product of the number of
subsets and the number of iterations in each of them determines
the effective number of iterations . From a
computational point of view, the OSEM method is approximately as
many times faster as the number of subsets we
use.
SART(Simultaneous Algebraic Reconstruction
Technique) - works simultaneously on multiple sections of a 3-D
image ............
OSSART - combination of OSEM and SART methods
............ ... ..... add .........
Hybrid reconstruction of
SPECT-CT ?
Some new SPECT / CT
hybrid systems attempt to improve the quality of SPECT
images through special iterative reconstruction, integrating
SPECT and CT data during reconstruction using local ("zone")
CT density maps of soft tissues, lung or adipose tissue, and bone
tissue. These zone CT density maps define tissue boundaries and modulate
their coefficients ("remodel") the primary
scintigraphic data of the radiotracer distribution. This achieves
a sharper boundary of bone tissues and lesions - provided that
these tissues take up the radiopharmaceutical
(eg 99m-Tc phosphonates). This modulation may also more
significantly show the differentiation between cortical and
spongy bone in the vertebrae and flat bones, or between the
cortex and cavity in the long bones. Simply put, modulation by CT
coefficients gives the SPECT images of the radio indicator
distribution a higher contrast .
A slight improvement in the quality of the
images is visible, but the model dependence is debatable
here (confrontation with classical reconstruction is
desiderable!).
Advantages
of SPECT
Compared to planar scintigraphy, SPECT tomographic scintigraphy
has three advantages:
¨ More
precise determination of the anatomical position of structures
and their shape in a three-dimensional image, when viewed from
different angles.
¨ Better
separate display of successive lesions at different depths.
¨ By
suppressing the superposition of radiation from overlapping
layers, a significantly better image contrast is achieved, which
enables more sensitive recognition of small lesions even at
greater depths.
Adverse effects of SPECT and their
correction
As mentioned above, the main advantages of
tomographic scintigraphy are the provision of a complete complete
"from all sides" image (3-dimensional image) and a
significantly higher contrast imaging of the
lesions against the tissue background. However, we will also
mention some disadvantages and pitfalls
of SPECT scintigraphy.
As with planar scintigraphy, SPECT
scintigraphy has some adverse and disruptive effects that may
degrade imaging quality. Here are five basic adverse effects, the
first two of which are also known from planar scintigraphy (however, with the SPECT method they manifest themselves
more markedly and in a slightly different way), the last three are specific to SPECT (rotary method) .
Note: The same applies to the pitfalls and possible errors of correction methods as described above in the section " Errors and pitfalls of correction methods - correction artifacts " in general scintigraphy.
Use of
SPECT scintigraphy
Spect tomographic scintigraphy represents (as
opposed to planar scintigraphy) a
significant addition and improvement to the
geometric-anatomical information on the distribution of
the radioindicator in tissues and organs. It is mainly used in
scintigraphy of myocardial perfusion (§4.9.4 " Scintigraphy of myocardial perfusion ") and brain
(§4.9.8 " Perfusion scintigraphy of the
brain "), as well as
reporter scintigraphy of the brain ( " Scintigraphy of brain receptor
systems
" ). Even in other
scintigraphic methods, such as skeletal scintigraphy or
localization of tumor diagnostics, tomographic SPECT imaging is
beneficial.
Another important
possibility, the specification of anatomical localization
, is the fusion of SPECT + CT images in
two-modal combinations of SPECT / CT (see
below the section " Fusion of PET and SPECT images
with CT and NMRI images ") .
Scintigraphic images provide important information about the
functional status of tissues and organs, but are usually unable
to provide sufficient anatomical information about the exact
location of pathological abnormalities (lesions) imaged
scintigraphically. Radioactivity does not enter the surrounding
anatomical structures (eg skeletal) , which do not capture the radioindicator and are not
visible in the scintigraphic image. It is therefore optimal to
perform a better and clearer comparison of the character, size
and location of the displayed structuressimultaneous
imaging of SPECT + CT or PET + CT images, where X-rays
of CT provide precise anatomical localization of the examined
structures.
Positron
emission tomography PET
Positron emission
tomography ( PET ) is a method of
scintigraphic imaging of the distribution of positron ( b + )
radionuclides, based on the detection of annihilation
photons formed by the interaction of emitted positrons
with electrons in the tissue of a patient to whom a positron
radionuclide was applied. During the radioactive conversion of a
positron radionuclide, a positron ("positive electron") is
emitted from the nucleus . In the material environment, the
positron gradually loses energy by collisions with the electrons
of atomic shells and zigzags change the direction of motion.
After braking (thermalization) of the positron e +during a relatively short path, the interaction with the
electron e - their mutual annihilation -
transformation of the electron-positron pair into two
gamma photons with energies 511keV, which fly apart from
the place of annihilation simultaneously in opposite directions,
at an angle of 180 ° *) - see §1.6,
passage " Interaction of charged particles -
directly ionizing radiation ",
Fig.1.6.1 below .
*) This applies exactly in the center of gravity
reference system of the positron and the electron. The energy of
photons 2 ´ 511keV is a consequence of the law of
conservation of energy (resting energy of electron and
positron is m 0e.c 2 = 511keV), the
opposite direction of 180 ° is a consequence of the law
of conservation of momentum . In the case of collisions
of positrons and electrons of higher energies, the angle of
inclination of annihilation photons would differ from 180 °. In
the material environment, however, the positron and the electron
have relatively low velocities at the moment of annihilation, so
that the emitted quantums actually fly in almost opposite
directions, with a maximum deviation of + - approx. 2.5 °. The
effect of this angular deviation is discussed below in the
section " Spatial
resolution of PET ".
Furthermore, own annihilation usually
precedes the formation of metastable bound electron-positron
system pozitronia . In the case of the so-called
orthopositronium , 3 photons g
can also be emittedwith continuous spectrum. This can only be
observed with positron radionuclides in a sparse gaseous medium;
in a relatively dense tissue environment, this phenomenon is very
rare (for details see §1.5, section "
Elementary particles and their properties ", passage "Positronium") .
The path of the emitted positron in
the substance (tissue) is "zigzag" and depends on its
energy. The mean range or range of the
positron determines the average distance of annihilation from the
point of positron emission, ie from the beta +
radionuclide position . Positrons from the point of emission can
fly isotropically in all directions, so that the points of
annihilation can be anywhere inside a sphere with a radius given
by the range of the positrons. The medium range of positrons is
thus limited the maximum physically achievable
resolution of PET (discussed below in the
paragraph " Spatial
resolution of PET ") .
Positron emission tomography uses coincidence
detection of a pair of photons of gamma annihilation radiation
(511 keV energy), which arise during the annihilation of a
positron b + with an electron and fly out of their place
of origin in opposite directions - at an angle
of 180 °. This coincident - simultaneously -
detection of a pair of annihilation photons is used for electronic
collimation of g radiation and subsequent reconstruction of
tomographic images .
Note:
For scintigraphic detection of annihilation radiation, a classic
scintillation camera with a special "heavy" collimator
with sufficiently strong septa between the holes can be used in
principle. In this mode, however, only one of the pair of photons
is always scanned - it is a single-photon planar
or tomographic scintigraphy (SPECT). However, the detection
efficiency is very low here (only one photon + low transmittance
of collimators + low absorption in a thin NaJ (T1) crystal) and
the images have poor spatial resolution (usually worse than 10
mm) due to coarse collimators. This scintigraphy is no longer
used.
Some alternatives, such as the
multidetector and Compton cameras mentioned above, are still in
the laboratory experiment stage and can only be used for
scintigraphic imaging of small objects.
Development of PET
The basic primary particles used in PET, positrons, predicted by
P. Dirac in 1928, were first discovered by C. Anderson in 1932 in
cosmic rays (§1.5, part " Elementary
particles and their properties ")
. Coincidence detection of pairs of annihilation quantum
radiation from positron radionuclides for gamma imaging was first
tested by W. Sweet and G. Brownel in their two-detector motion
scanner in the late 1950s, other PET experimental devices were
designed at Univ. of Pennsylvania, the first ring detectors
designed by R. Robertson and Z.H.Cho. A significant impetus for
the development of PET was the synthesis of 18-FDG
(fluorine-labeled 18-glucose) in 1970 and the discovery of its
accumulation in tumor tissues. In the early 1990s, PET
gammagraphy began to be used clinically in large laboratory
centers, and after 2000 it spread more and more rapidly to
clinical workplaces of nuclear medicine. After 2005, most PET
cameras are produced in a hybrid combination
with X-ray CT imaging - PET / CT (or with magnetic resonance PET
/ MRI) - cf. §4.6, part " Hybrid
tomographic systems ".
All complex oncology centers are gradually being
equipped with PET / CT devices .
Coincidence
detection ® electronic collimation of g- radiation
The photons g generated during e + e -
annihilation have three significant geometric properties:
¨ They fly out of the annihilation site simultaneously
and in the opposite direction - at an angle of 180 °;
¨ They move along straight paths
;
¨ They move at a speed of light of
300000km / s, so they can be detected at laboratory scales
practically simultaneously . These
properties enable the so-called concident detection of
pairs of annihilation photons: we place the measured positron
emitter between two detectors (small enough in
size), the outputs of which are connected to an electronic coincidence
circuit . Only pulses corresponding to the simultaneous
detection of photons in both detectors pass through this
circuit to another electronic apparatus . Due to the above
geometric properties, only photons from annihilations that
occurred on a straight line can be detected in
this way.sensitive points of both opposing detectors. If
annihilation occurs outside this space of the connector, then
even in the case of detection of one of the photons by one
detector, the other of the annihilation photons is not captured
by the opposite detector - the pulse does not appear at the
output of the coincidence circuit. Thus, when a pulse appears at
the output of the coincidence circuit, it is known that e
+ e - annihilation has occurred
at one of the points on the junction of the two detectors .
If we surround the investigated
object with a positron radionuclide by a larger number of
oppositely placed detectors in a coincidence circuit ,
we achieve targeted directional detection of
annihilation g- photons - their electronic collimation,
without the need for physical shielding with a lead hole
collimator.
Fig.4.3.5. Principle of scanning and reconstruction of positron
emission tomography.
Left: Coincident acquisition of annihilation
photons g . Middle: Image reconstruction. Right:
Scintiblock with BGO / LSO pixel crystal and 4 photomultipliers (manufactured by Hamamatsu) .
PET scanning principle
The PET scanning principle is schematically shown in Fig.4.3.5.
The PET scintillation camera detector has an annular
arrangement of segments of a large number of small
scintillation crystals in optical contact with photomultipliers
*), which detect flashes caused by the interaction of radiation g . Due to the
relatively high energy of annihilation radiation g 511keV, BGO
or LSO material with higher density and higher
detection efficiency in the area of ??higher energies g is used in
scintillation crystals instead of the usual NaI (Tl) - see §2.4. "Scintillation detection and
spectrometry", section " Scintillators and their properties " . The diameter of the
detector ring is usually 60-80 cm.
*) Individual scintillation crystals with dimensions around 4x4mm
are fixed in scintiblocks (Fig.4.3.5 on the
right) together with photomultipliers, it is described below.
The
investigated object W , in which the b + -radioactive substance is distributed , is
located inside the detection ring of the PET camera (Fig. 4.3.5
on the left). If a radioactive b + -transformation of the radioindicator nucleus occurs at
a certain point , the radiated positron e + after practically 1-3 mm (depending on its kinetic
energy *) movement in the tissue by ionization braking
practically stops and when interacting with the electron e - annihilates
: e + + e - ® 2 g , both quantums of
annihilation radiation g 1 and g 2 with energy 511keV are scattered in opposite directions
(ie at an angle of 180 °), pass through the tissue and are coincidentally
registered by an annular scintillation detector in two
places (angles j 1 and j 2 , in the picture marked: j 1 ® x 1 , j 2 ® y 1 ). The sensing ring
of the detectors located around the object to be examined thus
detects those photons which have fallen at the same time on the
opposite points of the ring. The connection of these places, the
so-called coincidence line or the response
line passes through the point where the + e -
annihilation occurred . The set of these coincidence lines from
individual pairs of detected annihilation photons (x i , y i ) then serves to reconstruct
the image of the distribution of the positron
radionuclide in the investigated object - in Fig.4.3.5 on the
right.
*) This range of positron
radiation in the tissue determines the basic limit
below which it cannot be reached with the resolution of
PET imaging. For the most commonly used 18 F, the range of positrons in the tissue is about 0.9
mm, which is substantially less than the actual resolution of the
PET apparatus. It is discussed in more detail below in the
section "Adverse effects of PET".
Differences
between PET versus planar scintigraphy and SPECT
The main difference from conventional planar or SPECT
scintigraphy is that PET detectors are not equipped with lead
collimators with many holes, as collimation is performed
electronically , leading to significantly higher
detection efficiency of PET compared to SPECT (where most of the
radiation is absorbed in the collimator septa). Another
difference is that the imaging detector of the SPECT camera must rotate
around the examined object (patient) in order to store partial
projections at different angles (this is
the case with existing SPECT cameras; with newly developed stationary SPECT cameras , acquisition takes place from all projection angles at
the same time) . For PET, the detectors do
not rotate around the patient, they are stationary
- ring detectors store data from all projection angles simultaneously
. The resulting image can then be reconstructed continuously
during the acquisition.
Coincidence PET with
double-headed rotating cameras for SPECT
In the mid-1990s, some scintillation camera manufacturers (first Adac
, then Picker , Elscint , GE and
others) developed special electronic circuits that allowed to
perform positron emission tomography on conventional two- (or 3)
- main cameras used for SPECT . Both heads,
placed opposite each other and without collimators,
rotated around the object under investigation as in the
acquisition of SPECT, but the pulses were fed to a special coincidence
device which recorded and evaluated pulses corresponding
to the simultaneous detection of 511 kV annihilation photons by
both opposing detectors. The software of the evaluation device
then performed the reconstruction of the transverse sections in
the same way as for PET.
This solution initially seemed very promising, as it
would allow to perform PET even in workplaces that do not have
expensive single-purpose equipment, a universal double-headed
SPECT camera supplemented with suitable electronics would
suffice. using a thicker scintillation crystal. However,
experience from practical use has shown that this is a sub-optimal
solutionwhich, with its detection efficiency and resolution,
cannot completely compete with single-purpose PET cameras with a
ring detector. Therefore, the manufacturers of scintillation
cameras soon withdrew from this solution and
offer separately classic double-headed cameras for SPECT and
separately PET cameras with a ring arrangement of detectors.
Three
types of coincidences in PET
In the coincidence detection of annihilation photons, there can
be basically three cases where two photons g are detected
simultaneously :
¨ True
coincidences
- direct detection of pairs of photons always coming from one
e + e - annihilation. The
annihilation site is located exactly on the line between the
opposite detectors, which during the reconstruction creates an
image of the radio indicator distribution. For not very high
frequencies the number of true coincidences increases practically
linearly with activity in the field of view, at
higher frequencies it grows more slowly due to dead time and at
very high frequencies it even decreases due to overload by random
coincidences (paralyzable dead time effect)
.
¨ Scattering
coincidences
- one or both simultaneously detected photons succumbed to Compton
scattering , which deviated their
angle. The annihilation site does not lie on the junction of the
detectors that registered this pair of photons. The percentage of
scattering coincidences increases with the (electron) density of
the material environment and their number again increases
essentially linearly with activity in the field of view (similar to true coincidences) .
If only one of the annihilation photons hits one of the detectors
and the other escapes outside the opposite detector after Compton
scattering, no coincidence is recorded.
¨ Random
coincidences
- this is the detection of photonsg originating from different
annihilations , which accidentally hit
the opposite detectors simultaneously (within
the time coincidence). The location of neither annihilation lies
on the junction of the detectors that registered it. The number
of random coincidences is proportional to the square of
the activity of the positron emitter in the field of
view.
Only true coincidences
produce a correct gammagraphic picture of the positron
radionuclide distribution. Scattering and random coincidences are
parasitic (the relevant coincidence lines are false,
they do not reflect the actual distribution of the positron
radioindicator - this is the so-called combinatorial
background ) and degrade image quality - reduce
contrast and increase noise.
Newer types of PET cameras consist of several
coaxial rings of detectors arranged side by side, which
allows the simultaneous scanning of several transaxial sections;
the field of view in the axial direction is approx. 15 cm for
current devices. In this arrangement, two types of scanning are
used :
In the so-called 2D
method, shielding baffles are inserted between
the individual detection rings , so that the coincidence lines
are scanned separately from each cross section - only in the
plane of the rings, perpendicular to the system axis.
In the so-called 3D
method, the septa are extended from the detectors and coincidence
sensing also takes place "obliquely" from the
directions between the planes of the individual rings - the
coincidences from the detectors in the different rings are also
evaluated. Thus, significantly more photons can be captured, ie
achievedhigher sensitivity . However, there is
also an increased probability of accidental coccidences (see
below), so this method can only be fully utilized with cameras
with faster detectors based on LSO scintillators.
For imaging larger parts of the body or
for full-body imaging , PET cameras are equipped
with an examination bed with a motorized controlled
movement . The computer system then combines the scanned
data from several patient positions during reconstruction into
one large whole-body tomographic image.
Scintillation Detectors for PET
Scintillation
Crystals
As mentioned above, at a relatively high energy of 511 keV
annihilation gamma radiation, conventional NaI (T1) scintillation
crystals have low detection efficiency. Higher density
scintillation materials are more suitable for PET to achieve high
detection efficiency with not too large a crystal thickness - to
achieve high detection efficiency and good spatial resolution of
scintillation localization by a photomultiplier system in the
camera's ring detector. It is also highly desirable to have a short
scintillation duration (scintillation afterglow) so that
a narrow coincidence window can be used - high time
resolution to reduce random coincidences (and the possibility of using the TOF method - see
below)
. To detect gamma radiation, a number of scintillation materials
with different properties have been synthesized (see §2.4, section " Scintillators and their properties ") . In principle, several
types of scintillators are applicable for PET :
Scintillator: | NaI (Tl) | BaF 2 | LaBr 3 (Ce) | YAlO 3 (Ce) (YAP) |
LuPO 4 (Ce) (LPO) |
Gd 2 SiO 5 (Ce) (GdSO) |
Bi
4 Ge 3 O 12 (BGO) |
Lu 2 YSiO 5 (Ce) (LYSO) |
Lu
2 SiO 5 (Ce) (LSO) |
Lu Fine Silicate (LFS) |
LuAlO
3
(Ce) (LuAP) |
Density [ g / cm 3 ] | 3.67 | 4.89 | 5.1 | 5.55 | 6.2 | 6.71 | 7.13 | 7.1-7.4 | 7.41 | 7.35 | 8.34 |
l max [ nm ] | 415 | 220/310 | 360 | 350 | 360 | 440 | 480 | 420 | 420 | 425 | 380 |
scint. afterglow [ns] | 230 | 0.8 | 16 | 30 | 24 | 60 | 300 | 41 | 40 | 33 | 11/28 |
h [photon/ MeV] | 4.10 4 | 1.8.10 3 | 6.3.10 4 | 1,6.10 4 | 1,3.10 4 | 8.10 3 | 6.10 3 | 3.10 4 | 3.10 4 | 3.2.10 4 | 9,6.10 3 |
In practice, heavier BGO
scintillators are used , more recently LSO (possibly LYSO modified) , which
has the advantage of a significantly shorter scintillation
afterglow. LYSO scintillators have similar
properties to LSO; the yttrium component causes
technologically easier growth of single crystals. A higher
percentage of yttrium (LYSO also occurs on
the composition of Lu 0.6 Y 1.4 SiO 5 : Ce), but reduces the
mineral density and the detection efficiency in comparison
with the LSO.
Based on LSO, the LFS
( Lutetium Fine Silicate ) scintillator was further
developed , which has a finer crystal structure and in
addition to basic lutetium, silicon, and oxygen (LSO) with doping
Ce, it also contains carefully tested small impurities of some
other elements such as Ca, Gd, Sc, Y, La, Eu, or Tb. This results
in slightly better energy resolution and shorter scintillation
afterglow.
LaBr 3 scintillator : Ce5% (is
hygroscopic) with very fast scintillation
is tested for TOF (see below) .
Internal radioactivity
of LSO scintillators
A minor disadvantage of lutetium -based
scintillation detectors (such as LSO and LYSO) is the higher
radiation background due to the internal
radioactivity contained in the scintillator. In addition
to the basic stable isotope 175
Lu, 2.6% of the long- lived radioisotope 176
Lu with a half-life of 3.8.10 10 years is also contained in the luterium - see §1.4,
passage " Lutetium ". During its radioactive decay, beta and gamma
radiation is emitted, which is internally detected with high
efficiency and causes an internal radiation background
*) in each detector *) about 40 pulses / sec / l gram LSO (more detailed analysis was performed in §2.4, part
" Scintillators and
their properties ",
passage" Internal radioactivity of LSO scintillators "). Beta radiation is fully absorbed in each individual
crystal independently, so it does not manifest itself in
coincidence measurements (random coincidences are negligible
here). However, gamma-ray beams, especially 300-kV photons, can
fly out of individual LSO crystals and hit other detectors, where
they can be detected immediately - creating a coincidence event
contributing to the background in the PET image. The background
thus formed is negligibly small in relation to the fluxes of the
measured annihilation radiation of the order of 10 6 photons / s in
clinical scintigraphy. However, certain problems may arise in experimental
studies of PET with low activities of the order of kBq units
at long measurement times ( animal PET ).
*) It is interesting that a
typical PET camera, consisting of about 190 blocks of LSO
crystals with a volume of about 50 cm 3 , contains a total internal radioactivity of 176 Lu of about 2.4
MBq! Each 50 cm 3 scintillation detector produces about 12,500 pulses /
s. radiation background, which significantly burdens the
electronic reading circuits. However, when coincidently measuring
higher activities (approximately 100 MBq in patients), this is
practically not applied in the resulting images .
However, this "parasitic" radiation can be used for
continuous calibration and tuning of PET detectors,
without the use of phantom sources.
Photodetectors for PET
Two types of photodetectors are used in PET to capture and
electronically register light flashes from scintillation
detectors : × Photomultipliers
are the most commonly used and proven electronic light signal
sensors from scintillation detectors - they are described in
detail in §2.4, section " Photomultipliers ". They have high and linear gain, low
signal-to-noise ratio, short output signal pulse (short dead
time). Their partial disadvantages are the complexity of the
design, the need for high voltage, larger dimensions (they cannot
be miniaturized too much), higher cost, relatively low quantum
efficiency and sensitivity to the magnetic field.
× Semiconductor detectors
are a modern alternative to photomultipliers. Their main
advantages are: small compact dimensions (miniaturization), high
quantum efficiency, low voltage, lower cost, insensitivity to
magnetic fields. Two types of semiconductor photodetectors are
used for scintillation sensing :
- Photodiodes are formed by p- and n-type
semiconductors in close contact in the pn junction. They are
connected in inverse polarity to voltage (in the reverse
direction). The impact of the photon of light excites
electron-hole pairs in the semiconductor material, whereby a
current pulse passes through the diode. At higher voltages,
secondary electron-hole pairs also form and signal amplification
occurs. If the electric voltage is set just around the breakdown
voltage of the pn junction, there will be an avalanche-like
increase in electron hole pairs when the photon strikes - it is
avalanche photodiode , operating in the so-called Geiger
mode .
- Silicon "photomultipliers" SiPM
are multipixel avalanche photodiodes - multipixel photon
counters, each element of which works independently in Geiger
mode. The output signal is proportional to the number of pixels
that have been hit by photons, and thus the number of photons
detected in the flash, they have spectrometric properties (they
are described in more detail in §2.4, section " Photomultipliers ",
section " SPM
Semiconductor Photomultipliers
").
Detector blocks for
PET
Scintillation crystals with photomultipliers (or semiconductor
photodetectors) are assembled into compact scintiblocks
in a PET camera, distributed around the circumference of the
circular gantry. Each such scintiblock is formed
by a square 2D array of crystals (BGO or LSO), connected to
photomultipliers by means of a light guide - Fig.4.3.5 on the
right. The array of crystals is usually formed from a single
single crystal using sections filled with light-reflecting
material. The usual configuration consists of a crystal measuring
5 ´ 5
cm and 3-5 cm thick, cut into an array of 8 ´ 8 partial crystals, to
which 4 photomultipliers with a diameter of approx. 2 cm are
attached via a light guide (in Fig. in the
right part of the picture) . When the
photon g of radiation hits one of the crystals, the resulting
scintillation light is shared by all four photomultipliers.
Information accuratethe position of the flash
(x, y coordinates) in the crystal field is obtained by electronic
analysis of the ratio of pulse amplitudes at the
output of individual photomultipliers, similar to a classical
planar gamma camera (described above in
§4.2 " Scintillation camera ",
Fig.4.2.1; each PET scintiblock can be considered a simple "
Anger mini-camera ") .
One ring ( ring ) of the detector is made up of about 48
scintiblocks closely adjacent in a circle diameter of about 60-70
cm (gantry), a camera comprising such parallel rings 3-5.
As mentioned above, instead of
photomultipliers, multicrystal scintillation can electronically
scan arrays of semiconductor photodiodes, or better SiPM
photomultipliers. The near future probably
belongs to compact scintiblocks LSO-SiPM,
LYSO-SiPM or LaBr 3 -SiPM. For more distant development in PET (instead of BGO / LSO scintiblocks with classical
photomultipliers or SiPm) there are promising multipixel fully semiconductor
detectors (eg based on CZT) . In addition to better detection efficiency and spatial
resolution, a slightly shorter coincidence time (for better TOF) can be achieved .
So far, it is being tested experimentally on smaller PET models.
The advantage of semiconductor detectors is also theirsindependence
from the magnetic field , which allows use also in
hybrid PET / MR systems .
Reconstruction of PET images
During the acquisition, a large number of coordinates
of coincidence lines (in the order of millions of coincidence
detections) are scanned; data are stored in the form of so-called
sinograms . By computer reconstruction of
these linear projections of coincidence sites, images of
cross-sections are created and from a set of transverse
sections, computer reorientation can be used to create sections
at any angle, or 3D images similar to SPECT. The reconstruction
uses either (filtered) back projection methods ,
which can, however, produce stellar artifacts around positive
lesions, or more computationally demanding iterative
reconstructions., providing higher quality images
without these artifacts. Another advantage of iterative
reconstruction is the ability to incorporate various properties
of specific devices and methods (such as homogeneity,
attenuation, noise, resolution) directly into the reconstruction
procedure. Conventional reconstruction methods, analogous to
SPECT, have been described above in the section " Computer reconstruction of SPECT ".
................? add special modifications
of reconstruction procedures? ........
PET imaging properties
Compared to classical single-photon planar and SPECT
scintigraphy, two-photon condensation tomography PET has two
basic advantages: significantly higher
detection efficiency (sensitivity) and slightly better
spatial resolution :
Detection efficiency ( sensitivity
) of PET
The absence of classical collimators and registration of photons
of annihilation radiation simultaneously from all
directions , using electronic collimation, leads to significantly
higher detection efficiency (sensitivity) of PET gamma
cameras compared to classical Anger cameras, where the vast
majority of gamma photons are not detected. (flies
"into space" or is absorbed in the septa of
collimators) .
The detection
efficiency or sensitivity h detection
devices and spectrometry of ionizing radiation is generally a
ratio between the number of detected pulses and the number of
incoming quanta of radiation introduced relative and
absolute efficiency, often expressed in % (it was defined and physically discussed in §2.1,
section " General physical and instrumental effects in
detection and spectrometry
", section " Detection efficiency and
sensitivity ") . For
gamma cameras where the radionuclide used is the
gamma radiation source , the sensitivity - detection efficiency -
is usually quantified as the number of pulses detected by the
camera per unit time [per second] - cps ,
relative to the activity unit [kBq, MBq] of the
radionuclide used in the displayed source; for the planar / SPECT
scintigraphy is usually 99 m Tc for PET it is 18 F . Only exceptionally expressed in % (detection efficiency-sensitivity in classical gamma
cameras is discussed in §4.5, section " Sensitivity ( detection
efficiency ) of a scintillation camera ") .
Detection
efficiency - sensitivity - of PET cameras are determined by
several physical, geometric and technical factors, which can be
divided into two categories :
l Detection efficiency hd annihilation radiation 511keV in the detection elements of the PET ring. This
"physical" detection efficiency hd of annular
detectors depends on the thickness of the
detection material h , its density and
the atomic number - despite the linear attenuation factor
m for gamma 511keV is given by the coefficient (1-e
- m .
h ) (detection
efficiency of scintillation detectors is discussed in §2.4
" Scintillation detection and gamma
radiation spectrometry ", section
" Scintillators and their properties "; scintillation materials suitable for gamma
511keV were discussed above in the section " Scintillation
detectors for PET ") . It also depends on the amplitude analyzer window
setting, scintillation afterglow and dead time. At greater
distances from the center, there is also a slight decrease in the
detection efficiency due to the oblique anglethe impact
of most of the annihilation photons on the detectors. We include
all these other individual influences in the factor f
. In our case of coincident two-photon detection
of PET in two opposite detectors with detection
efficiency h d , the resulting detection efficiency will
be given by the product h d . h d , ie
it appears in the quadrate h d
2 = [f . (1 - e - m .h )] 2
.
l Geometric efficiency hg
the PET registration of 511 kV annihilation photons is given by
the spatial angle of projection at which the
annihilation photons from the activity source are detected. Each
radionuclide source emits radiation isotropically
in all directions - up to a solid angle of 4p . Around the point source,
located in the middle of the detection ring of radius R ,
we can draw an imaginary spherical surface of this radius R
- its surface 4p R 2 will pass all photons emitted from the
source (if the detectors were placed
densely on this spherical surface, the geometric detection
efficiency 100 % ) .
Circular width detection ringHowever , S , which
has an area of ??2 p R.S, passes out of this total number only a part
of the photons, given by the area ratio of 2p R.S / 4p R 2 = S / 2R. If the PET
camera has N parallel rings of width S and radius R
, the geometric efficiency will be h g
= NS / 2R (if we neglect the gaps
between the detectors and somewhat oblique angles from the center
to the peripheral rings) . Increasing the
diameter of the ring 2R reduces the overall projection solid
angle and thus the geometric efficiency. With a rotating number Ndetection
rings in a PET camera, on the other hand, increases the
projection angle and thus the geometric detection efficiency.
By multiplying
these two partial factors of detector and geometric
efficiency, we can obtain the resulting relationship for the total
detection sensitivity h of the PET
camera:
h
= [f . (1 - e - m .
H )] 2 . N . S
/ 2R,
where f is the fraction of detected photons in the
photopeak with the specific setting of the PHA analyzer window (with possible angular dependence in the peripheral
parts) , h is the thickness of the
detectors, m is the linear attenuation factor of the detector
material used for gamma radiation 511keV (usually
BGO or LSO m = ....) , R
is the radius of the detection ring, S is its width, N
is the number of detection rings.
Current standard PET cameras with
three detection rings with a diameter of about 70 cm, each
consisting of 48 scintiblocks BGO / LSO 5 x 5 cm and a thickness of
about 5 cm, achieve a detection sensitivity of about 7-10
cps / kBq 18 F (for classic Anger planar or
SPECT cameras sensitivity only approx. 0.04-0.3 cps / kBq
99mTc, depending on the collimator used - almost 30-100
times lower) .
Spatial resolution of PET
Replacement of mechanical collimators by electronic
collimation also leads to a somewhat better spatial
resolution of PET compared to conventional Anger
cameras.
The spatial
resolution (abbreviated resolution ) of a
scintigraphic image is the
smallest distance [mm] of two point radioactive sources in the
displayed object, which are still distinguishable from each other
in the scintigraphic image as two images. We can determine it as
the width of the PSF profile in the image
of a point or line source in half the maximum height of the
profile, converted to a spatial scale in the object [mm] - it is
called FWHM ( Full
Width at Half Maximum - overall width at half maximum; resolution for
classical gamma cameras is discussed in §4.5, passage " Spatial resolution ") .
![]() |
||
Fig.4.3.6. Physical and
geometric effects on positional resolution in PET
imaging. a) Range of positrons h in the tissue (an example of three ranges in different directions is marked - strongly increased) and flight deviation of annihilation photons from 180 °. b) Projection-geometric degradation of resolution due to width d of detectors. c) Irradiation of annihilation photons between detection elements causes radial "astigmatism" in images of peripheral parts. |
The spatial resolution of PET imaging is again
determined by a combination of several physical, geometric, and
technical factors :
¨ The range
of positrons in the tissue prior
to positron annihilation represents the primary physical
limit for PET resolution. Positrons are emitted from
nuclei at high speed - with a kinetic energy of hundreds of keV
to several MeV, so from the site of radionuclide deposition
positrons fly in the tissue to a certain
distance (approx. 0.5 mm-6 mm, depending on
the radionuclide) before braking (thermalizing) and meeting with
electrons (via positronium) annihilate on a pair of 511keV photons. Therefore, since
the positions in which annihilation photons are generated are
somewhat different (and always different)
from the position of the original parent nuclei, there is some blurring
of position . The magnitude of this blur depends on the
positron radionuclide used - on the maximum and mainly the mean
energy of the emitted positrons, determining their mean range h in
the tissue (Fig. 4.3.6a) . For some
positron radionuclides used, the following is:
....... table ..............
¨ Angular blurring due to incomplete braking of positrons
. Electron-positron annihilation in the tissue then occurs with a
certain residual kinetic energy (different each time) , so that in the
laboratory reference system the angle of radiation of
annihilation photons will be slightly different from 180 °, on
average by + - 0.25 °. This angular uncertainty - variability
causes a small geometric blur , which is
proportional to the radius R
of the detection ring, with a value of 0.004.
R .
¨ Size of detectors - width d of detection elements is the dominant factor of projection-geometric
degradation of resolution. Opposite detectors of finite
(non-zero) dimensions detect radiation not only from a single (central, axial) coincidence lines, but from
the whole cone of angles. This leads to a geometric uncertainty
of the coincidence line with respect to the actual position of
the annihilation, the half-width of which is d /
2 (Fig. 4.3.6b) .
¨ Accuracy of decoding the position of
scintillation within the
scintiblock of detection elements, using a significantly lower
number of photodetectors (eg 4
photomultipliers per 64 detection scintillation elements) . The inaccuracy of this decoding (optical
multiplexing) somewhat degrades the
resolution. For this paper, the value of the half-width of
approx. D / 3 was empirically determined .
¨ Irradiation of
annihilation photons between detection elements - penetrating annihilation g of 511keV radiation can
interact with several different (adjacent) crystals. This can
cause a detected signal in another neighboring crystal than the
one on which the photon primarily strikes. As a result, the
coincidence line may be incorrectly assigned to one of the
adjacent detection elements. This effect occurs when annihilation
photons hit the detection elements in an oblique
direction , which occurs from sources farther from the
center (for sources in the center r = 0 it
does not manifest itself and the projection of coincidence lines
remains narrow, given only the size of the detection element) - Fig. 4.3 .6c . These obliquely
incident photons can interact with several different crystals,
depending on the depth of penetration into the
scintillation material. Radial projection of the source is doneextended
by the trigonometric factor k . r / Ö (R
2 + R 2 ) , wherein the coefficients for
a given depth of penetration annihilation photons into
scintillation - polovrstvou absorption of photons 511 keV in the
material of the detector h 1 /2 = ln2
/ m , where m is the linear
attenuation coefficient gamma 511 keV in the material
scintillator - Fig. 4.3.6c .
This creates a kind of radial "
astigmatism "*) - asymmetric blurring of images of peripheral
sources, more distant from the center of the ring r = 0. For
detection crystals made from BGO or LSO scintillators, an
approximate value of 12.5 was measured for the
coefficient k . This effect is clearly visible in
Fig.4.3.7. on the right in images and profiles of peripheral
point sources at distances r = 34 and 30 cm, partly also for r =
20 cm (indicatively already manifests ur =
10 cm) .
*) It is a bit similar to astigmatism
in optics - imaging defect ( aberration ) lens,
reflected by objects at greater distances from the optical axis,
or optical systems asymmetrical to the optical axis.
¨ Reconstruction
algorithms for the generation of
the resulting images of the distribution of the radio indicator
using a set of coincidence lines show minor errors and
variations. There is a non-uniformity of the density of
coincidence lines with respect to the position of sources inside
the detection ring, different types of reconstruction algorithms
and filtering. During the reconstruction, a certain common additional
coefficient of degradation of resolution *) is created , for
which the range of approx. 1.2-1.5 was empirically determined; we
will use an approximate value of 1.3 here .
*) In this physical analysis we mean standard
"classical" reconstruction algorithms of filtered back
projection or iterative reconstruction of OSEM. We
do not consider special reconstruction algorithms with built-in resolution
recovery, design modifications and PSF modeling or filtering
using inverse MTF to artificially
improve resolution. We are talking about the purely physical
properties of PET images , not about the
possibility of their additional computer improvement (which,
however, can be useful in practice ...).
These effects
lead to several contributions to the response function of the PSF
point source, which are approximately Gaussian in shape. By their
quadratic summation (geometric
average) we can then obtain the resulting
relationship for the total spatial resolution of the
PET image :
FWHM
= 1.3 . Ö [ (d / 2) 2 + h 2 + (0.004 .
R) 2 + (d / 3) 2 + (12.5 .
R) 2 / (r 2 + R 2 ) ] ,
where FWHM is the resulting half-width of the
response function of the point source, ie the total
resolution , d is the size (width) of the
detection element, h is the mean range of the positrons, R
is the radius of the detector ring of the PET camera, r is
the radial distance of the source emitter from the center of the
ring (all dimensions are in millimeters) .
Current standard PET cameras with
three detection rings about 70 cm in diameter, each consisting of
48 scintiblocks 3 cm thick with detection elements about 5x5mm in
size, achieve spatial resolution in the middle of the ring in
transverse sections around 4.2-5 mm *) (classic planar / SPECT cameras they reach such a
resolution only close to the front of the collimator, in
practical scintigraphy, where the distance - depth - of the
lesion is around 10 cm, but the resolution cannot be achieved
better than 10-12 mm).
*) Small animal PET cameras with
a diameter of approx. 20 cm with detection elements with a width
of approx. 0.5-1 mm achieve an even significantly better
resolution of approx. 1-1.5 mm .
![]() |
Fig.4.3.7. PET images of point sources
18 F located at
different distances r from the
center of the detection ring. By analyzing the ROI and
profile curves with these images, the values ??of
detection efficiency h
and spatial resolution FWHM were
measured (we measured on a PET camera GE
Discovery IQ at KNM FN Ostrava) . At the last peripheral point source at a distance of r = 34 cm, part of its image was already cut off by the edge of the field of view. The measurement was performed using a simple arrangement described in the work " Phantoms and phantom measurements ", part "Tomographic phantoms ", passage" Simple improvised phantom for measuring the imaging properties of a PET camera ". |
Our measurement of spatial resolution and detection sensitivity on PET images of point sources, located at different distances from the center of the detection ring, is marked in Fig.4.3.7. The effect of radial astigmatism can be clearly seen in the images and profiles of peripheral point sources at distances r = 34 and 30 cm, partly also for r = 20 cm (indicatively, ur = 10 cm is already evident) . For similar reasons (due to the oblique angle of incidence of most of the annihilation photons on the detectors) there is also a slight decrease in the detection efficiency at greater distances from the center.
TOF
- time localization of the annihilation site
The basic (conventional) PET method described above does not give
any information about the annihilation site on the coincidence
line, all pixels on the coincidence line are assigned the same
annihilation probability, the image is formed only by
intersections of coincidence lines. Increasing speed electronics
and the introduction of detectors with high time resolution (such as LSO scintillators) gradually
allows the use of more important "information channel"
annihilation radiation for PET: it is called measurement. TOF
( Time Of Flight ) - the time of flight of
photons gfrom the annihilation site. These photons fly in
opposite directions at the speed of light c = 300000 km / s. If
annihilation occurs in the middle of the coincidence line
"0", both photons are detected exactly at the same
time. However, if an annihilation occurred off-center, at a
distance D x, the photon g 1 will have a flight
time TOF 1
= x 1 / c
to the detector , while the second photon g 2 will fly a slightly different time TOF 2 = x 2 / c - see fig. 4.3.8.
From the time difference t 2 -t 1 = TOF 2 -TOF 1 it is then possible to determine the radial
coordinate D xplaces of annihilation on the
coincidence line: D x = c. (t 2 -t 1 ) / 2.
![]() |
Fig.4.3.8 Time localization of the annihilation site x 1 , x 2 on the coincidence line by electronic analysis of the difference in flight times of annihilation photons D TOF in positron emission tomography. |
If the coincidence detection of annihilation
radiation has a sufficiently short time resolution
, the time difference between the detection of
both annihilation quanta g can be measured , which allows (at least in principle)
to determine the place on the coincidence line D x where
annihilation occurred and where both photons were emitted * ) -
fig.4.3.8. This introduces additional information about the position
of the detector response into the system . The time resolution of
existing instruments does not yet allow accurate localization of
annihilation sites, but even the approximate location of the
annihilation photon radiation site could shorten the response
line, improve the reconstruction procedure, and improve
the signal-to-noise ratio in the resulting images.
*) If we could measure the
time differences of annihilation photon arrivals with picosecond
accuracy, this information would be enough to determine the sites
of annihilation and achieve PET imaging. There would then be no
need to reconstruct the quantifications of the intersections of
the condensation rays, but the cross-sectional image could be stored
(in polar coordinates) directly . These would no
longer be coincidence lines, but consensus points
. However, for it we do not yet have fast enough electronics and
detection technology ...
TOF is still in the stage of
laboratory development. The initial enthusiasm has not yet been
fulfilled, the method is rather a promise for the futurefor
the next generation of PET cameras. For current types of PET
cameras (2010-15) with installed TOF, the TOF time resolution is
about 500-600 picoseconds, which corresponds to the possibility
of resolving the annihilation site on the coincidence line of
about 15-18 cm. The TOF parameter is so far only minimally
usable in clinical scintigraphic diagnostics . TOF
analysis will be relevant only when its resolution from the
current 15cm can be improved to about 2cm...
An improvement in TOF resolution to less
than 400ps can be expected with the introduction of special
scintillators (such as lanthanum bromide
LaBr 3 )and modern photodetectors (silicon photomultipliers
SiPM). Data are loaded and reconstructed in LIST-mode format,
iterative reconstruction procedures (3D list mode TOF MLEM, OSEM,
...) have built-in special correction algorithms, containing data
from calibration measurements. So far, however , "no
miracles" can be expected ..!
.. - TOF will remain only a physical-technical interest
for a long time ...
Adverse
effects on PET and their correction
As with planar and SPECT scintigraphy, there are some adverse and
disturbing effects on PET imaging, which worsen the quality of
the images. We will mention here several significant adverse
effects, of which the first four are also known from planar and
SPECT scintigraphy, the others are specific for two-photon PET :
The same applies to the pitfalls and possible errors of correction methods as in the section " Errors and pitfalls of correction methods - correction artifacts " in general scintigraphy.
Construction design of PET gamma cameras
The basis of each PET gamma camera (for
which there is a less suitable name " PET scaner
") is a circular ring of
detectors with a diameter of 60-80 cm, registering pairs
of photons of annihilation gamma radiation with energy 511keV in
coincidence mode from opposite directions. The ring consists of a
large number of detectors (mostly
scintillation) , the scintiblocks (see Fig.4.3.5 on the right) are
mounted in several parallel concentric rows on a gantry
, through the inner cylindrical "tunnel" of which the
bed lies with the patient. The movement of the lounger is driven
by a servomotor, ensuring precise computer-controlled movementso
that the images captured by the rings from the individual parts
of the body are folded into the resulting PET image (sometimes even whole-body) ,
including online fusion with X-ray CT images.
![]() |
Fig.4.3.9. PET / CT positron
emission tomography examination room at the Department of Nuclear Medicine, University Hospital Ostrava. In the middle is the basic PET / CT (GE Discovery) device . Aiming and navigating laser pointers for radiotherapy planning are installed on the sides and ceiling of the laboratory . In the back of the right is on the stand applicator contrast agents for CT. |
Current PET devices are two-modal -
PET/CT for the assessment of the exact anatomical
location of imaged lesions by fusion with CT X-ray images (or PET / MRI for fusion with nuclear magnetic
resonance images) . In addition to the PET
ring, a CT ring (or MRI) is also installed on the same gantry ,
through which the bed passes in a controlled manner and, at
the same time as the in-line PET image, it also
creates CT images (or MRI)
of the patient . In addition to the physiological-anatomical
correlation , CT images also provide density maps for
the correction of attenuation of gamma annihilation radiation in
PET tissues.
The device
sometimes includes opticalaiming and navigation lasers
for precise localization of lesions when defining ROI
within the irradiation plan using PET images.
Use of PET scintigraphy
The areas of clinical use of positron emission
tomography in nuclear medicine are given, similarly to emission
planar and SPECT scintigraphy, mainly by the properties of
relevant radiopharmaceuticals , here
radiopharmaceuticals labeled with positron radionuclides (these radionuclides and radiopharmaceuticals are
briefly described in §4.8 " Radionuclides and
for scintigraphy ") . The most important area of ??PET use is oncological
diagnostics - finding out the location and nature of tumors
, which accounts for more than 90% of all PET examinations (§4.9.6 " Scintigraphic
diagnostics in oncology "
and §3.6, section " Diagnosis
of cancer") . To assess the precise anatomical localization of
lesions appear used fusion of PET with X-ray CT images
- dvoumodalitní PET / CT , or. MRI images (PET
/ MRI).
![]() |
Fig.4.3.10.
Example of PET / CT scintigraphy with 18 FDG in a patient with lymphoma. (PET / CT images were
taken by |
At a general level, PET works very well in the
field of assessing the metabolic activity of tumors,
proliferation, tissue hypoxia, the density of expressed receptors
in cells.
Specifically, herein used pharmacokinetic
properties especially 18 F-deoxyglucose FDG (hereinafter 18 FLT, 18 F-choline) which is selectively uptake in tumor
cells with increased metabolism of carbohydrates - appears metabolic
cellular activity of tissues (whereas
X-ray and ultrasound displays only morphological page) . Malignant tumors are usually characterized by glucose
hypermetabolism . This method is therefore also suitable formonitoring
the response of tumor tissue to therapy by imaging
metabolically active tumor tissue as opposed to inactivated
cells; it is possible to monitor the therapeutic response - the
"success" of therapy. Among other things, it is able to
recognize tumor recurrence from other processes (eg from the consequences of previous tumor treatment) , see §3.6, section " Modulation of radiation beams " . Monitoring of the
therapeutic response by PET consists in comparing the metabolic
activity of the tumor before the start of treatment and
after the application of therapy. The change in tumor metabolism
occurs before the change in its dimensions,
assessed by morphological X-ray or sonographic imaging methods.
PET can also be used for detection inflammatory process
in the organism (§4.9.6 " Oncological
radionuclide diagnostics. Scintigraphy of inflammation ") , in cadiology for the
diagnosis of myocardial viability (§4.9.4 " Nuclear
cardiology ") .
The metabolic activity of
tumor lesions is often quantified in PET images using SUV
values - from a general point of view, it was discussed above in
§4.2 " Quality of scintigraphic imaging ", section " Quantification
of positive lesions - SUV
". There were discussed some physical and biological factors
that may skew the absolute SUV quantification
and recommended relative SUV quantification to
compare the metabolic activity of lesions in specific patients
before and after therapy. In the case of PET imaging with 18 F-FDG, the blood
glucose level also adds to this - its increased value
reduces the accumulation of FDG, which underestimates the
SUV .
PET (/ CT)
has thus become an important functional imaging method in the
primary diagnosis, staging, assessment of therapeutic response,
recurrence search or re-staging of a number of oncological
diseases . PET images can also be advantageously used
for radiotherapy planning . It is discussed in more detail in §3.6, section " Diagnosis of cancer ".
For PET scintigraphy with the most
commonly used18 F-FDG is a problem of physiologically variable
(sometimes quite high) accumulation of FDG in a number of
healthy viable tissues, especially in the brain and myocardium (see Fig . 4.3.10)
. Therefore, PET with FDG is not suitable
for the detection of brain metastases or minor perfusion defects
of the myocardium. Non-specific increased accumulation of FDG is
also observed in inflammatory processes, wound healing, after
surgery. The displayed site with increased glucose uptake may not
always be a tumor ...
Therefore, is promissing the use of
other radiopharmaceuticals and positron radionuclides (§4.8, passage " Radionuclides and
radiopharmaceuticals for PET
") such as gallium 68
Ga , zirconium 89
Zr , iodine 124 I , copper 64 Cu (for PET)
and beta - 67
Cu (for therapy) , scandium 44
Sc (for PET) and beta - 47
Sc (for therapy) , or mixed alpha-beta + terbium 149
Tb (for both PET and
alpha-therapy) . These radionuclides can be
used to label mainly monoclonal antibodies for
PET diagnostics and to perform subsequent biologically targeted
radionuclide therapy - theranostics (§4.9, passage " Combination of diagnostics and
therapy - teragnostics ")
.
An interesting
application of PET has recently appeared in the so-called hadron
radiotherapy (§3.6 "Radiotherapy", part
" Hadron
radiotherapy "), where
irradiation with high-energy charged particles in the irradiated
tissue causes, among other things, nuclear reactions
, during which positron radionuclides are formed
. When irradiated with accelerated carbon nuclei 12 C, a positron
radionuclide 11 C is also formed , the distribution of which can be
visualized by the PET method. With a PET camera installed on a
hadron radiotherapy irradiator, we can monitor the dose
distribution in the target tissue and in the
surroundings - so-called in-beam
PET monitoring - and thus control the
course of radiotherapy (see Fig. 3.6.6 in §3.6).
PET is
also used in neurology to diagnose brain
activity and perfusion. The area of ??the brain that is active
has an increased accumulation of radiopharmaceuticals, which can
be used to assess brain activity and its association with some
psycho-neurological disorders (including
Alzheimer's disease) . Furthermore, it is a
scintigraphic diagnosis of inflammatory processes
and examination of the myocardium , where the perfusion
and viability of the myocardium can be assessed on the basis
of the consumption of special positron radiopharmaceuticals (see below §4.9.4, section " Myocardial
perfusion ") .
The fusion combination
of PET scintigraphy with X-ray CT is
very important for clinical applications imaging (anatomical - §3.2, part
" Transmission X-ray tmography CT ") , which provides
visualization of morphological and anatomical structures with
high spatial and density resolution. This information obtained
from CT can be used to increase the accuracy of the location,
extent, and nature of the lesions found in PET images. X-ray CT
thus complements the functional information
obtained by PET with the help of a radiopharmaceutical with
localization anatomical information. For modern
devices, this is implemented online in a two-mode hybrid PET
/ CT system - see below §4.6,
section " Hybrid tomographic systems " . This combination also
allows quality correction for absorption (attenuation) of the
annihilation g radiation in tissue. Recently, there is also a hybrid
combination of PET / MRI.
Positron
emission mammography (PEM)
The PET method using suitable
tumor-accumulating radiopharmaceuticals (usually
18 FDG or
FLT) is naturally also used in the
diagnosis of breast cancer . However, the
specific anatomical proportions of the breasts and the properties
of mammary lesions have led to efforts to develop smaller
single-purpose - dedicated, optimized - PET
imaging devices that would have a higher resolution for
small lesions typical of breast cancer. And also a shorter
acquisition time than with whole body PET and the application of
lower radio indicator activity. Basically, two technical
solutions of these specialized devices for PEM positive
emission mammography have been developed :
- The
first with its design resembles a classic X-ray mammography with compression
, only the X-ray machine and the detector are replaced by two flat
PET imaging cameras (in Fig. 4.3.11 on the left) , between which a breast with suitable compression is
inserted.
- The
second system is an small annular PET detector of
circularly arranged scintiblocks (approx. 48), of significantly
smaller dimensions than whole-body PET (O approx.
20 cm) , usually one ring into which uncompressed
breast is inserted . The breast hang freely inside the detector
located under the bed with the hole on which the patient lies - Fig.4.3.11 on the right .
Fig.4.3.11. Design of PEM positron emission mammography
instruments.
Left: Compression PEM mammogram with
flat PET detectors. Right: Ring PEM
mammogram with loosely inserted breasts without compression.
In both cases, concurrent
detection of a pair of annihilation photons by opposite detectors
is performed, with computer reconstruction of cross-sectional
images, as in classical PET. The advantage of these optimized PEM
devices is better spatial resolution (approx. 2-3 mm),
allowing to detect even small lesions in the breast (in case of good accumulation even under 1 cm) .
Although small PEM devices are significantly cheaper
than large universal (full body) PET cameras, the positron
emission mammography method has not become more widespread (unlike the widely used X-ray mammography - §3.2,
section " X-ray mammography "). One of the reasons is a
positron radiopharmaceutical with a short half-life, which is
difficult to access outside the larger workplace of nuclear
medicine. It serves only as an additional method to
X-ray, sonographic or MRI mammography. However, in order
to visualize the more complex extent of the disease, it is still
necessary to perform PET imaging on a larger scale, including
nodes and possibly. metastasis. PEM is a specialized peripheral
method that is uniquely used mainly in the USA and Japan in some
larg complex oncology centers ...
Specialized PEM devices in the world are
supplied by only two manufacturers: CMR Naviscan ,
California, USA and IHEP - GaoNeng Medical Equipment
, Hangzhou, China .
-------------------- small physical-technical interest ------------------- |
Neutron Stimulated Emission Computed
Tomography ( NSECT)
NSECT ( Neutron Stimulated Emission Computed
Tomography ) is a new (and so far experimental) method of spectroscopic
imaging of the concentration of certain elements in an
organism using neutron interaction . Unlike
conventional emission computed tomography SPECT (or PET), gamma
radiation is not emitted by radioactive isotopes, but by stable
isotopes in which the emission of g- radiation (characteristic
energy) is stimulated by inelastic scattering of
fast neutrons by which the analyzed area is externally
irradiated. These stable isotopes may either be a natural part of
the tissue under investigation or may be introduced as molecular
indicators (similar to contrast agents or radioniculators), eg by
a metabolic pathway.
Fig.4.3.12. Principle of neutron stimulated emission computed
tomography NSECT.
The analyzed area (sample, tissue)
is irradiated with a beam of fast neutrons
(energy approx. 7-10 MeV) from a suitable collimated neutron
source - electronic neutron generator (§1.5, part " Accelerators ", passage "Accelerators as neutron
generators") or radioisotope source ( .....) . These neutrons collide
with the nuclei of the atoms of the irradiated material, and
there are basically three types of interactions (see §1.6, passage " Neutron
radiation and its interactions
") . For our purposes, the inelastic
scattering of neutrons is important , in which the
neutron transfers part of its kinetic energy to the nucleus and
this causes an increase in its internal energy - excitation
of the nucleus.. When the nucleus returns to its
original state (deexcitation of the excited nuclear levels), a gamma
radiation photon with a precisely
determined characteristic energy , given by the
type of nucleus, is emitted . These energies of secondary g- radiation from excited
nuclei range from tens of keV to about 6 MeV. By
spectrometric detection of this g- radiation it is possible
to determine which elements are represented
(according to the energy of the line g ) and in what relative
concentration (according to the intensity - the number of photons
in the respective peak). The spatial distribution of
these g-emitting
cores can
be determined by gamma detection (gamma camera) *). Or the
spectrometric detector can sense gamma radiation at
different angles . Computer maps of spatial distribution
of concentrations of specific chemical elements in the examined
tissue can be created by computer reconstruction of positions
(angles) and energies of this neutron-stimulated g- radiation.
*) The energy of g- photons emitted
from excited levels of stable nuclei during neutron excitation is
usually too high for imaging by standard gamma
cameras. They are hundreds of keV to several MeV (eg for 16 O the E g = 6MeV, for 12 C the E g is= 4.5MeV). For
such energies, gamma camera collimators have poor spatial
resolution and luminance, and the scintillation crystals used are
too thin to achieve reasonable detection efficiency; also the
spectrometric properties are not good. Therefore, spectrometric
detectors not providing spatial information, but only the energy
spectrum, are used in current experimental methods. Information
on the position of the analyzed atoms is obtained either by
rotating the detector equipped with a collimator and scanning
from different angles, or by rotational scanning of the examined
object by closely collimated neutron beams. In this second
method, the path of the neutron beam defines the geometric
position of the examined volumes and the spectrometric detector
integrally scans all photons emitted by excited nuclei along the
path of the neutron beam.(that is, the part
of the photons that enters the detector) .
Unwanted background pulses can be significantly reduced by using
a conductor spectrometer circuit, triggered by the pulse mode of
the neutron generator. This is followed by computer
reconstruction of data from individual projects. The resulting
image is basically 4-D : for each voxel of the
3-D image, information about the energy of g- radiation ® representation of
various elements
is also stored . By selecting a specific energy (energy window),
an image of the distribution of the corresponding specific
element is obtained. Note: For the
gammagraphy of this hard g- radiation, special Compton cameras
(described above in the section "New and
alternative physical and technical principles of gamma- ray imaging, the " High
Energy Gamma Camera" passage ), which are still experimental.
NSECT can in
principle show the distribution of all elements and their
isotopes, except hydrogen (whose nuclei do not have excited
levels and therefore, there is no stimulated g- emission) and
helium (which has too high an excitation energy of 25MeV)
Neutrons are penetrating particles, so structures in the depth of
the organism can be excited and displayed, possibly correcting
the absorption (attenuation) of primary neutron radiation and
registered stimulated g-radiation. The diagnostic potential of NSECT is due to
the fact that the relative proportions of different elements,
including trace elements, are different for different tissue
types. It also differs slightly between healthy and tumor tissue.
The method has so far been tested in the early diagnosis of
breast and lung tumors. Apart from laboratory experiments, NSECT
has not yet been implemented , it will
probably remain only a physical-technical interest. ...
Note:
NSECT has some analogies and common aspects with other methods of
neutron analysis of materials , especially neutron
activation analysis NAA (INAA), described in §3.4, part
" Neutron activation analysis ". For special purposes of biological research, ineutron
activation analysis in vivo : the relevant part of the
organism is irradiated with neutrons (from a reactor or neutron
generator) followed by a standard gamma plot of the
distribution of induced beta radioactivity accompanied by gamma
photons, mapping the distribution of the test substance in
tissues and organs.
4.4.
Gated phase scintigraphy
In scintigraphic analysis of periodic actions in
an organism, this periodicity can be advantageously used in a
methodological approach called gated (triggers)
scintigraphy . In addition to scintigraphic impulses, another
electrical signal is also recorded from the camera - an ECG or a
respiratory signal - which suitably controls (triggers, gates)
the course of the acquisition.
Phase scintigraphy of
fast periodic actions
Dynamic scintigraphy with time of rapidly
changing distribution of radioactivity encounters
fundamental physical and technical problems. In order to
faithfully capture the dynamics of the monitored process, it is
necessary to use the best possible time resolution
, ie a high frequency of short-term frames. The statistical
character of radioactive decay then leads to significant statistical
fluctuations of the measured pulses in the image.
Relative statistical fluctuations are given by the expression 1 /
Ö N,
where N is the number of pulses in the image cell
accumulated over one frame. At high time resolution, the storage
time of one image is very short (of the order of 10 -2s), the numbers of
accumulated small pulses and statistical fluctuations are very
large *). The starting point is not an enormous increase in
applied radioactivity (this is usually not possible for other
reasons, especially radiohygienic), because due to the dead time,
the detection device is not enough to effectively process such a
fast flow of pulses.
*) For accurate capture of cardiac
activity, it is necessary to divide the heart cycle into very
short time intervals; since the cycle lasts approximately 1
second, the storage time of one frame should be approximately
0.03 seconds. With this extremely short measuring time, the
statistical fluctuations of the recorded pulse frequencies are so
large (tens of%) that they do not allow the individual images to
be evaluated. In such pictures, it would not even be possible to
know which organ it is - only a spray of chaotically scattered
dots would be visible.
Thus, at first glance, it seems
completely impossible to perform a detailed dynamic scintigraphy
of one cardiac cycle. Fortunately, however, there are two
favorable circumstances :
1. Cardiac activity is a periodic event
(this is true at least approximately;
2. Cardiac activity is accompanied (or
triggered) by electrical currents, that can be
detected externally.
In the case where the observed
event is periodic , ie the distribution of
radioactivity is a periodic function of time, the situation is
significantly more favorable. If we denote the scintigraphic
response function f (x, y, z, t), where x, y, z are positional
coordinates, t is time, then the following will apply to the
periodic event: f (x, y, z, t) @ f (x, y, z, t + kT), k =
0,1,2, ...., T is the period (" @"means that equality is
valid only on average, except for statistical fluctuations in
decay and registration.) A dynamic scintigram of such an event is
in principle given by a scintigraphic study of only one period
(cycle). Conversely, the periodicity of the process offers the
possibility to create a dynamic study of one cycle periods) with
a very high time resolution and at the same time with
satisfactory "statistics": we measure several hundred
individual cycles in succession with a high time resolution as
they follow each other, and then synchronously compose (absorb)
the results frame by frame based on periodicity to create dynamic
study of only one cycle:
F N (x, y, z, t) = k = 1 S N f [x, y,
z, t + (k-1) .T], t Î < 0, T).
Such a synchronously composed study F N (x, y, z, t) will be called a phase
dynamic scintigraphic study - it is a study of one
"average" or "representative" cycle, composed
of N common cycles of periodic events.
This can be done directly (without
additional information) using a computer if the period T is
exactly known and constant. In practice, however, this is usually
not met, eg the heart rate fluctuates somewhat. From
statistically strongly scattered data, the computer does not
completely "know" the individual phases of the periodic
event and then has nothing to compose synchronously. Therefore,
it is also necessary to input certain synchronization
pulses to the computer from the outside (marks) which
make it possible to pinpoint the end of one cycle and the
beginning of the next cycle. In the case of cardiac activity,
such synchronizing or gating pulses can be signals from the ECG
(R-waves), by means of which the computer always
"recognizes" the end of one and the beginning of the
next cycle.
![]() |
Fig.4.4.1. Left: Schematic of
creating a gated phase dynamic scintigraphy of the
cardiac cycle based on scintigraphic data from the camera
and synchronization derived from the R-wave of the ECG. Right: Images from many different heart cycles are added synchronously to form a series of images capturing a single average cycle. |
The principle of construction of the phase dynamic study of the cardiac cycle is shown in Fig.4.4.1. The computer program divides the time interval between two R-waves into short intervals of length D t = T / (number of frames per cycle), where the number of frames per cycle is usually chosen 32, sometimes 16. Signal derived from the R-wave (electronically from its leading edge) determines the beginning of the cardiac cycle. Pulses corresponding to the beginning of the cardiac cycle in the interval 0 to D t are recorded in the first image in the computer's memory. In the time from Dt to 2 Dt, the pulses are stored in the second frame, in the time from 2.Dt to 3.Dt to the third frame, etc. In the same way, the duration of the next cardiac cycle is divided, the beginning of which is signaled to the computer by another R-wave from the ECG; pulses registered during the first interval D t are added to the first image of the previous heart cycle, pulses registered from D t to 2. D t are added to the second image of the previous cycle, etc. This process is repeated many times, forming a phase dynamic study of one average cardiac cycle.
Cycle selection and
excluding
As is well known, the heart rhythm is never completely regular,
the heart rate and the period are more or less variable
, even arrhythmias can occur. Only those cycles whose period does
not differ too much from the average period need to be taken into
account in the calculation. Irregular cardiac cycles, ie those
whose duration is different from normal regular cycles, should be
excluded from the record . These false cycles,
caused by extrasystoles or other heart rhythm disorders, would
distort the overall dynamics of the average cycle - it would no
longer be a representative cycle. The limits for selecting the
"correct" cycles are usually chosen to be ± 10% of T. It is
necessary to exclude not only such an incorrect cycle with an
anomalous period, but also cycle following it , as it may not
begin at the correct stage of end-diastole. Due to the slightly
fluctuating cycle length, the accumulated number of pulses in the
last phase images is artificially reduced. In order not to
distort the dynamics of the terminal section of the phase curves,
an appropriate correction is made based on the number of cycles
that contributed to the individual phase frames.
LIST-mode acquisition
The method of acquisition into image matrices described above is
sometimes referred to as frame-mode . In earlier
generations of acquisition computers, where a sufficiently large
operating memory was not available, acquisition in the so-called LIST-mode
was used : the coordinates (x, y) of
individual pulses were sequentially stored in the memory as they
came one after the other. Synchronous pulses from the ECG were
also recorded into this continuous data stream under appropriate
coding. Conversion to scintigraphic images (re-framing) and
construction of the own phase study was then performed additionally.
The advantage of LIST-mode was that the data could be
subsequently re-framed with different selection of correct
cycles, which is important in some heart rhythm disorders (such
as bigimenia), when the correct phase study cannot be obtained in
frame-mode. LIST-mode has been practically abandoned for dynamic
scintigraphy for many years, now it is being used for some
iterative tomographic methods (§4.3,
section " Computer reconstruction of SPECT ", " Reconstruction of PET images ", " TOF - time localization of the annihilation
site ") .
First-pass phase
scintigraphy
The above-described method of construction of phase dynamic
scintigraphy of the cardiac cycle is performed in situations
where blood carrying a radioindicator (eg 99m -Tc labeled erythrocytes) is evenly and steadily mixed
in the bloodstream - the so-called stady-state
method. We will briefly mention the construction of phase
scintigraphy of the cardiac cycle using the first-pass
method , where the radioindicator is applied to the circulation
as a compact bolus . Acquisition from the camera
into the computer's memory starts when the bolus arrives in the
heart chamber and ends before recirculation begins (when the
structures would already overlap).
![]() |
Fig.4.4.2. Construction of a phase dynamic study of the cardiac cycle in first-pass radionuclide ventriculography. The "pulsed" curve represents the time course of radioactivity in the left ventricle during the first bolus flow. |
Fig.4.4.2 shows the time course of
radioactivity in the left ventricle during such a measurement.
The curve is "pulsed" due to the periodic filling and
emptying of the chamber carrying blood carrying the radioactive
bolus. The construction of the phase dynamic study of the cardiac
cycle is performed in a similar way as with the steady-state
method, but only a few cycles can be taken into account -
starting with the bolus arrival in the left ventricle and ending
with the onset of recirculation, when the ventricular images
would overlap.
The advantage of the first-pass method is that the
radioactivity is contained only in the chamber, so on the one
hand the problem of correction on tissue and blood background is
eliminated, on the other hand it allows scintigraphic
"view" of the heart chamber even in such directions
images of the right and left ventricles and possibly and other
structures. The main disadvantage of the first-pass method
compared to the steady-state method is the small number of cycles
from which the phase study is created; therefore, the image
quality is not very good due to statistical fluctuations. In
addition, such a study may not represent a representative cardiac
cycle, as selection of the "correct" cycles is not
feasible here and acquisition is performed immediately after
injection of a radiolabel, when central hemodynamics are affected
by stress. From the quantitative parameters, therefore, only the
ejection fraction can be objectively evaluated. The first-pass
method is now rarely used to construct phase scintigraphy of the
cardiac cycle. Used only whenbolus radiocardiography
(§4.9.4, section " Dynamic bolus angiocardiography ") .
The method of phase (gated) scintigraphy is practically exclusively used in nuclear cardiology (§4.9.4 " Nuclear cardiology ") . It is both radionuclide ventriculography (§4.9.4, the " equilibrium gated ventriculography ") , whose comprehensive analysis program Ventre is described in §3.1 " Radionuclide ventriculography " books " OSTNUCLINE - Comprehensive assessment of scintigraphy " in recent years, the mainly SPECT myocardial perfusion (gated myocard SPECT - §4.9.4, part " Scintigraphy of myocardial perfusion " ).
Author's
note:
We have been engaged in research and development of methods of
scintigraphic phase dynamic studies with high time resolution at
our workplace since 1976, practically in parallel with the
development of these methods in leading laboratories in the
world. An electronic device for R-wave detection and implantation
of synchronization pulses into a computer was designed. We
developed a program for the reconstruction and mathematical
evaluation of radionuclide ventriculography on a small computer
device Clincom (operational memory only 12k!), Which was
probably the most complex procedure in this area at that time;
then served as the basis for a comprehensive program VENTR ( " Radionuclide
ventriculography " ) on the GAMMA-11 device and later
on a PC, the OSTNUCLINE system. Several dynamic phantoms were
also constructed for this research and development work (starting
with a rotating gramophone disc and ending with a flexible
dynamic phantom of heart pulsation and circulatory pumping); some
of them are described in the work " Phantoms and
Phantom Measurements in Nuclear Medicine
", part. 4. " Dynamic phantoms
".
4.5.
Physical parameters of scintigraphy - imaging quality and phantom
measurements -
The task of scintigraphy is to provide quality, ie objective
, detailed and accurate imaging of the
distribution of radioactivity in the examined object, both spatially
and temporally (dynamic scintigraphy). We have
mentioned above several limitations and adverse effects
of a physical and technical nature that limit
the possibilities and quality of scintigraphic imaging (" Adverse effects of scintigraphy "). To assess the quality of
scintigraphic imaging, its optimization and detection of possible
errors and defects, it is necessary to analyze
and test the physical properties of
scintillation cameras. As with any complex measuring instrument,
the scintillation camera's properties can be described by several
physical parameters :
Spatial resolution of a gamma camera
A scintillation camera is an imaging device , so
the most important parameter is its resolution ,
or spatial or positional resolution
(given in length units - millimeters) :
Spatial resolution |
The spatial resolution of a scintigraphic image is called the smallest distance [mm] of two point radioactive sources in the displayed object, which are still distinguishable from each other in the scintigraphic image as two different objects. |
Equivalent definition : |
By spatial resolution we mean the width of the FWHM profile in the image of a point or line source in the middle of the maximum height of this profile, converted to the spatial scale in the object [mm] . |
The spatial resolution is thus given by the half-width of the image profile of the point or line source; it is called FWHM ( Full Width at Half Maximum ) . Two point radiation sources can be distinguished from each other in the scintigraphic image only if the distance between them is at least FWHM - Fig.4.5.1 :
Fig.4.5.1 Spatial resolution of scintigraphic imaging
- analysis of images of point sources, which are displayed as
"blurred" scattering circles.
a) By a "blurred" image of a
point source we lead a profile curve PSF
( Point Spread Function ), its half-width FWHM
indicates the resolution of the display. b, c, d, e)
Images of two point sources located at different distances from
each other. If this distance is greater than FWHM, these
resources are displayed as two separate objects. When zoomed in
to a distance equal to FWHM and smaller, the two images already
merge into one - the sources can no longer be
distinguished from each other . g)Schematic
representation of the interweaving, superposition and fusion of
images of two closely spaced point sources (on PSF profile
curves) - corresponds to the situation according to Fig. e)
.
Note: It was measured with point
sources with a diameter of approx. 1mm, 50MBq 99m Tc, on a Nucline
MB9201 camera close to the front of the HR collimator, matrix 256
x 256,
zoom 2 x, with strongly enlarged image sections, approx. 4 x . Slight
deformations of the circular shape of the images of point sources
were caused by small geometric irregularities of the holes of the
HR collimator (which are visible at this
high magnification, but do not manifest themselves in practical
scintigraphy) .
In addition to FWHM, the resolution of the
camera is sometimes characterized by the parameter FWTM
( Full Width at Ten Maximum ), which is the width of the image profile of the line
source in tenths of the maximum height of the profile. Of course,
this value is higher : in situations without a
scattering environment, it is approximately FWTM @ (1.8 - 2) ´ FWHM, in the
presence of a tissue (or aqueous) scattering environment, it is
approximately FWTM @ (2 - 2.5) ´ FWHM.
Spatial resolution
FWHM <--> modulation transfer function MTF
It should be noted that the FWHM value determined with PSF or LSF
may not to capture the quality of the display
completely objectively in terms of resolution! Scattered
radiation and irradiation of the collimator septa mainly expands
the lower part of the PSF, lower than 50% (see Fig.4.3.3d, ea Fig.4.5.4e) ,
so it may not affect the value of FWHM. For a completely detailed
physical analysis of the distinguishing properties of imaging
systems, the so-called modulation transfer functions
(MTF) are used - see the work " Theory of scintigraphic imaging and
modulation transfer functions ",
which take into account all points of PSF.
Terminological note: The spatial
resolution of a gamma camera is sometimes abbreviated to resolution
. Please do not confuse with spectrometry
energy resolution (....) or time
resolution (as the detector dead time is
sometimes called "Detector dead
time ") - it's something
completely different ..! .
The spatial
resolution of the gamma camera classic single photon - planar,
SPECT - is determined by two components :
1. Geometric resolution of collimator
Rcolim ,
which depends on the hole diameter d and
the length of the holes - channels L of
the height (thickness) of the collimator, significantly also on the
distance h of the displayed object from the
collimator:
Rcolim ~ d. [1 + (h + m) / L] ,
(4.5.1)
where m is also the width of the gap between the rear face
of the collimator and the camera crystal (its
center) - Fig.4.5.2a. The resolution of the
collimator is manifested by the fact that each point (point
source) in the object is geometrically projected on the camera
detector as a small blurred circular trace with
a diameter dependent on the R collim - Fig.4.5.2b .
![]() |
|
Fig.4.5.2
Two basic components of the spatial resolution of
a gamma camera - a collimator and a crystal +
photomultiplier. a): Trigonometric analysis of the collimator resolution . From each hole of the parallel collimator we can draw an imaginary cone defining the area from which gamma radiation can pass through this hole to the camera detector (radiation from places outside this cone is absorbed by the lead septa of the collimator) . As the distance from the collimator increases, this detection cone widens , thus deteriorating geometric spatial resolution of the image projected by the collimator onto the scintillation crystal of the gamma camera. b): The resolution of the collimator is manifested by the fact that each point (point source) in the object is geometrically projected on the camera detector as a small blurred circular trace with a diameter depending on the distance from the collimator. c): Approximate representation of the internal resolution of the camera detector, related to statistical fluctuations of detected scintillations, blurring X, Y coordinate pulses and Compton scattering of gamma photons. d): 99mTc point source display at different distances h from the front of the collimator HR - degradation of FWHM resolution with distance .. |
The spatial resolution of the collimator
improves with increasing the length of the holes - channels
(thickness of the collimators) and decreasing the diameter of the
holes. Collimators with small and longer holes have better
resolution (and at the same time less
dependence on distance) . The spatial
resolution of the gamma camera significantly deteriorates
with the distance *) of the displayed structure from the
front of the collimator (Fig. 4.5.2b, and
Fig . 4.5.6 on the
left) . The gamma camera (front of the collimator) should
therefore be placed as close as possible to the
surface of the patient's body.
*) For collimators with
a different arrangement of holes (described
above in §4.2, section " Scintigraphic
collimators ")the geometric situation is more complicated, but
basically the same rule applies to the deterioration of
the spatial resolution at a greater distance
from the collimator face - Fig.4.5.6 on the left.
Over-radiating
trough the collimator septum
For the correct imaging function of the collimator, it is
necessary that gamma radiation passes only through the holes,
while the partitions between them would not transmit radiation.
However, this cannot be fulfilled 100%, a certain part of the
radiation penetrates even through the septa. Especially at higher
energies may cause gamma over-radiating gamma
rays trough collimator septa, if these partitions between holes
thinner than optimum energy utilization gamma. Separation of the
septa impairs the contrast of the scintigraphic
image. If the low-energy LEHR collimator, optimized for 140keV 99m Tc, were used for
scintigraphy with 111 In or even with 131 I., we would get an image with greatly degraded
contrast! The effect of over-radiation trough of septa and its
influence on the quality of scintigraphic imaging is analyzed in
Fig. 4.5.3 :
Fig.4.5.3 Over-radiation of gamma photons over septa of a
collimator and its influence on the quality of a scintigraphic
image.
a) Schematic representation of the aborption of
low-energy gamma radiation in lead septa between the collimator
orifices and the partial irradiation of high-energy gamma through
the collimator septa. b) The over-radiation of
photons of higher energy trough the collimator septa causes the
enlargement of the edge parts of the point source image (red part of the PSF curve) . c)
, d) , e) Images of point
sources of radionuclides emitting various energies of gamma
radiation - 99m Tc (140keV), 111 In (245keV), 131 I (364keV), recorded with a LE HR collimator for low
energies.f) Phantom Jasczak, filled with 111 In and displayed
with a gamma camera with a collimator Medium Energy (top) and a
low-energy collimator LE HR (bottom). Radiation trough the LEHR
septum completely degrades the contrast of the image here - an
almost invaluable scintigram.
Note: The typical
"star-shaped artifact" of diagonal over-radiation is
due to the geometric configuration of the holes and lead
partitions, which are projected somewhat narrower in the diagonal
direction. For different types of collimators, the shape of this
artifact may be somewhat different, depending on the details of
the geometric design of the openings and partitions.
The proportion of radiation penetrating the
barrier between the apertures - transmission factor , is
e -m .s tran , where m is the linear absorption coefficient of the collimator
material (lead) for the required gamma energy and stran is the shortest
path that gamma radiation can penetrate the barrier from one hole
to adjacent hole. For a collimator with the diameter of the holes
d , their length L and the thickness of the baffles
s , the shortest
path of irradiation is the baffle s tran = sL / (2d
+ s), so that the transmission factor is e -m .sL / (2d
+ s) . Optimization of partition thicknessesbetween
the holes is performed on the basis of the requirement for a
sufficiently low value of the transmission factor (most often
0.05); this leads to a condition for the transmission factor e -m .sL
/ (2d + s) <0.05 (described in §4.2, passage " Scintigraphic
collimators "). This gives a
limitation for the thickness of the collimator baffles s > (6.d / m) / [L - (3 / m)].
Effective length of the collimator holes
In the formula (4.5.1) for spatial resolution, we did not
consider irradiating the collimator septa . If
irradiation occurs, it causes a seemingly effective
shortening of the actual length of the collimator holes in
terms of geometric collimation . Therefore, the so-called
effective length of collimator holes L ef = L-2.m -1 , wheremis the linear coefficient of attenuation
in the collimator material (lead); it is a correction for the
transmission of photons through two opposite baffles between the
holes. The introduction of L ef instead of L in (4.5.1) describes the
deterioration of resolution by radiation. However, this is not
very important, because the PSF does not have a Gaussian shape
here, it is widened in the lower part and leads to a
deterioration of not so much the FWHM resolution value
as theimage contrast , as seen in Fig.4.3.5d,
e).
2. Internal resolution of the camera
detector R int
Internal
(proper, intrinsic) resolution is
given by the accuracy with which the system of
photomultipliers and related electronics is able to locate
the position of scintillation in the crystal. At the
quantum level, the internal resolution is influenced by
statistical fluctuations in the production of
light photons after the interaction of g- radiation in the detector
and variations in the number of electrons emitted from the
photocathode and dynodes of the photomultipliers. These
fluctuations " blur " the
amplitudes of electrical signals from photomultipliers and thus
the values of the resulting coordinate pulses X, Y - Fig.4.5.2 c.
The use of more photomultipliers with higher quantum efficiency
and good optical contact with the crystal leads to better
internal resolution.
The thickness of the crystal
also has a negative effect here , in two ways. On the one hand,
it is a geometric "blur" of the positions of
light flashes registered by photomultipliers - scintillations
occur at different depths, and the system of photomultipliers
evaluates their positions X, Y somewhat different accordingly.
Furthermore, it is the multiple Compton scattering of g- photons in the
detector, which also causes uncertainties in the X,
Y-localization of the interaction site of primary gamma-photons.
A thinner crystal and a larger number of photomultipliers allow
you to achieve better resolution. For these reasons, thin
scintillation crystals about 0.7-1.8 cm thick are used. Current
scintillation cameras, optimized for g140keV, have a crystal
thickness of mostly 9.5mm and achieve an internal resolution of
2.5-3.5mm.
The total resolution
of the R gamma camera ( external -
extrinsic )
is then given by the geometric sum of
both subcomponents: R = Ö (R 2 int + R 2 collim ). However, in practice, the total resolution is not
calculated according to this relationship, but is measured using
point or line sources placed in the field of view of the camera
with the given collimator (" Phantoms
and phantom measurements in nuclear medicine ", section " Measuring
the positional resolution of the camera ") .
In practice, it is given in the
first place for each camera its internal
resolution , which is determined from the
production technology - typical values of internal
resolution for newer cameras are about 2.5-4 mm. The internal
resolution represents the limit value of the resolution
, which can no longer be reached for a given camera and which can
only be approached using an ultra-high resolution collimator.
Furthermore, the total
resolution of the camera with individual
collimators for the given distances of the source from the
collimator (usually 10 cm), or even in the presence of a
scattering environment (scattering
environment between the radiation source and the detector - water
and tissues, which is always present in scintigraphic imaging of
radioactivity in the body, somewhat worsens
spatial resolution and more marked contrast in the image) .
These values of the total resolution
are different, for LE HR collimators (high resolution), optimized
for 140keV 99m Tc, at a distance of 10cm from the front of the
collimator, it is about 8-10 millimeters. For collimators for
higher energies, the overall resolution (in
10 cm) is worse, about 12-15 mm. As
analyzed above, the spatial resolution of the gamma camera deteriorates
significantly with the distance of the displayed
structure from the collimator front. The gamma camera (front of the collimator) should
therefore be placed as close as possible to the
patient's body surface when scanning .
Influence of scattered
radiation
It also contributes to the deterioration of image qualityCompton
scattering of gamma radiation in the material
environment - in the patient's tissue (the
physical nature is described in §1.6, section " Interaction
of gamma radiation and X ",
passage " Compton scattering
") . Scattered radiation primarily impairs
the contrast of scintigraphic imaging, to a lesser
extent the spatial resolution of FEHM. This effect is analyzed in
Fig.4.5.4 :
Fig.4.5.4 Compton-scattered gamma radiation in
scintigraphy and its influence on the quality of
scintigraphic image.
a) If the coincidence scattering of photons in
the tissue at an angle so that scattered photons passing through
the aperture of the collimator is detected crystal camera, could
then lead to the detection of such photons scattered radiation g 'of the
false place - is detected gamma photons seemingly coming
from another places from where it was originally radiated during
the radioactive transformation.
b) These false scattered photons g ´ have a lower
energy than the "true" direct and primarily
detected photons g (part of the energy was transferred to the electron e - when
scattered in the substance) . ) , so they usually do not fall
into the photopeak. By carefully setting the analyzer
window to the photopeak of the given radiation g, we can therefore
largely eliminate the Compton-scattered
radiation g ´.
c) , d), e) Images of point
source without scattering medium (c) and with scattering medium
in photopeak measurement and including Compton scattering in a
wide analyzer window (e).
f), g) An image of the Jasczak water phantom with a
narrow photopeak window and a wide window, including scattered
radiation (deterioration of the image
contrast can be seen) .
Note: Spatial resolution of
PET
For two-photon cameras of coincidence positron emission
tomography PET , the physical principle of imaging and
analysis of spatial resolution is different - it
was described above in §4.3, passage " Spatial
resolution of PET " (spatial resolution of PET cameras is generally somewhat
better than conventional Anger cameras) .
The measurement of the
spatial resolution of the camera
can be performed in two ways :
¨ Quantitative
physical measurement
- by analyzing images of point and line sources, through which we
conduct profiles - sections, thus obtaining PSF
or LSF curves. From them, we then directly
determine the value of the half-width-resolution FWHM
in [mm] (or by a more complex analysis of
the so-called modulation transfer function MTF ....
"...") . For practical
determination of the resolution, it is more appropriate to use a line
source , the image of which we can guide independently
of several LSF profiles, or add these profiles and thus achieve
smaller statistical fluctuations.
![]() |
Fig.4.5.5. Measurement of positional
resolution of a gamma camera. Left: Line source (capillary) and two point sources for measuring positional resolution. Middle, right: The measurement was performed with a charge of 99m Tc at distances of 0, 5, 10, 15 and 20 centimeters from the front of the collimator of the HR camera Nucline TH. The degradation of FWHM resolution with distance can be seen in the images and profile curves. |
¨ Visual
evaluation of phantom images
- most often they are so-called Bar-phantoms irradiated
by a planar homogeneous source (mostly 57 Co), or vessels of complex structure, filled with a 99m Tc solution (eg Jasczak
phantom ). It is a more or less qualitative
evaluation , the value of the resolution is rather estimated from
the images; however, it is usually sufficient for practice and
comparison. The implementation of both methods of resolution
resolution is described in the work " Phantoms
and phantom measurements in nuclear medicine ", section " Measurement of positional resolution of the
camera ".
Sensitivity (detection efficiency)
of a scintillation camera
Detection efficiency or sensitivity S
of devices for detection and spectrometry of ionizing radiation -
radiometers - is generally defined as the ratio
between the number of detected pulses (quanta
registered by the detector) and the number
of incoming radiation quanta; introducing the relative
and absolute efficacy, often expressed in %
(physically defined and discussed in §2.1,
section " General physical influences and instrumentation
for the detection and spectrometry ," passage " detection efficiency
and sensitivity ") .
Single photon detection efficiency
of the classic gamma camera - planar, SPECT - is determined by
two components :
1. Geometric transmittance
(luminosity) of the collimator to gamma radiation,
indicative of which part of the incoming gamma photons collimator
pass on to the detector camera (unlike photons which are absorbed in the collimator
septa) . It depends in principle on what
part of the total area of the collimator is occupied by the
through-holes and which part absorbs the lead baffles between
them. According to the trigonometric analysis in Fig.4.5.2a, the geometric
luminosity Scollim of the parallel collimator :
S collim = (d / L) 2 . [d 2 / (d + s) 2 ]. K g . [100%] ,
(4.5.2)
where d is the diameter of the holes, s thickness
of the baffles, L is the length of the holes-channels
given by the height of the collimator. The geometric factor K g depends on the shape
and arrangement of the holes and partitions (for
circular holes in a hexagonal arrangement, K g = 0.24, for hexagonal
holes in a hexagonal arrangement, K g = 0.26, for square holes in a quangangular arrangement,
K g =
0.28). . Conventional
LE HR collimators have a geometric luminosity of about 1.2%, LEAP
about 2%, HS about 2.2%. ..............
Note: In the
equations for spatial resolution (4.5.1) and detection efficiency
of the camera (4.5.2) we did not consider over-radiatin
gtrough collimator septa . If it occurs, it leads to a
deterioration of the image quality (see Fig. 4.5.3), but at the
same time to an increase in detection efficiency. However, this
is a negative phenomenon, the quantification of which in physical
parameters is irrelevant ...
Independence of the detection efficiency of the gamma camera on
the distance
The intensity of the radiation decreases with the square of the
distance from the source (this is exactly
the case for a spot emitter). Therefore,
the detection efficiency of conventional radiometers is
significantly reduced for greater distances from the measured
sources. However, for gamma cameras with parallel collimators,
the detection efficiency (sensitivity) does not depend on
the distance of the displayed source from the collimator
front! In formula (4.5.1) the parameter h of distance does
not appear . The display of the point source in a wide range of
distances 0-30cm from the front of the collimator in Fig.4.5.2 d)
shows a deterioration of spatial resolution and decreasing image
brightness, but the total number of pulses is the same in all
images, area (integral) under PSF is the same for all distances.
This surprising behavior is due
to the specific properties of geometric collimation
for parallel collimators. We can clearly illustrate this
according to the schematic drawing in Fig. 4.5.2 a), b) as
follows: As the source moves away from the collimator front, the
number of photons incident on the individual holes decreases
quadratically as 1 / h 2 . However, the number of holes through which radiation
can pass to the detector increases quadratically
in proportion to h2 . These two opposing trends cancel
each other out , so that the total photon flux -
the efficiency of the collimator - does not
change with the distance between the source and the
collimator.
Note: This rule does not
apply to special convergent or Pinhole
collimators, the detection efficiency changes
significantly with distance - it increases or decreases (see the section " Imaging
properties of special collimators
" below) .
2. The internal detection efficiency
of the crystal and photomultipliers of the S int camera ,
indicating which part of the gamma photons incident on the
detector (i.e. released by the collimator) , is actually detected in the form of
scintigraphic image pulses by the scintillation system,
photomultipliers and analyzer system . It depends on the thickness
and conversion efficiency of the scintillator,
the gamma radiation energy , the window
width setting photopeak analyzer. The photopeak
detection efficiency of a standard gamma camera detector with a
9.5 mm thick crystal for 140keV 99m Tc is about 80%, for 364keV 131 I then about 30%.
The overall - system
detection efficiency
- sensitivity of the camera S is then
given by the product of both of these components S collim .
S int ; the main determining component is the efficiency of
the collimator. However, in gamma cameras, where the source of
gamma radiation is a radionuclide , the
sensitivity - detection efficiency - is usually quantified in a special
way : as the number of pulses detected by the camera per
unit time [per second] - cps , based on unit of
activity [kBq, MBq] of the radionuclide used in the
displayed source; for the planar / SPECT scintigraphy is usually 99mTc
, then the PET 18 F . Only exceptionally is it expressed
in %.
In scintigraphic diagnostics, we are
mostly concerned with the relative assessment of
the distribution of the radio indicator in various parts of the
examined object. In the case of so-called quantitative
scintigraphy , however, we may also be interested in the
absolute activity of the radioindicator in the
investigated area. In order to determine this real activity from
a scintigraphic image, we need to know the efficiency
( sensitivity ) of radiation detection g of
the used radionuclide by a scintillation camera. We also need to
know the detection sensitivity of the gamma camera to determine
the optimal applied activity of the radio
indicator to obtain sufficiently high-quality scintigraphic
images.
In the case of scintillation
cameras, for practical use, the detection sensitivity is
related to the radioactivity of the examined object :
Detection efficiency (sensitivity) of a scintillation camera |
The
detection efficiency, or sensitivity S, of the imaging
system is quantified as the pulse frequency N [imp./s]
measured by a scintillation camera with a
point radiation source g (located at
the desired field of view) , relative to the
activity unit A [MBq] of the source:
S = N / A. It is expressed in units [imp. s -1 MBq -1 ], or [cps / MBq] or [cps / kBq]. |
It is given for a specific type of radionuclide
and collimator. Most often, the sensitivity for planar and SPECT
cameras given for radionuclide 99
mTc, for
PET cameras 18 F . For different radionuclides, the
sensitivity of a scintillation camera generally has different
values, depending on the yield of gamma photons [%] (number of
gamma quanta / 100 conversions of the radionuclide) and their
energy [keV].
The basic physical
measurement of the detection efficiency of a gamma
camera - at a given distance and location of the field of view -
is performed with a point source of the required
radionuclide. We can thus perform detailed measurements of the
dependence of the detection efficiency on the distance in
different places of the field of view (as
can be seen in the right part of Fig.4.5.6).
Another way of measuring the "averaged" sensitivity is
with a planar source , which lies entirely in
the field of view (this method is
especially suitable for cameras equipped with a parallel
collimator) - in the image of the planar
source are averaged or. local
sensitivity inhomogeneities and measurement error may be
somewhat reduced. According to the recommended NEMA procedure (to ensure good accuracy and reproducibility) we determine the sensitivity of the camera by placing a
10cm diameter dish with 99m Tc solution with exactly known activity in the middle
of the field of view - approx. 10MBq, a layer of solution up to
1cm. We accumulate a scintigraphic image with the given
collimator and in the required configuration (acquisition time min. 100sec), in
which we determine the number of pulses in the ROI of the dish
image and convert it to 1MBq and 1sec.
In addition to the actual detection
efficiency of the scintillation crystal of the camera on
the given gamma radiation, the resulting, total - system
sensitivity depends decisively on the collimator
used . For universal collimators of the LEAP type, the
sensitivity of scintillation cameras for 99m Tc is around 150-300 cps / MBq, for high-resolution
(HR) collimators only about 50-100 cps / MBq 99m Tc. As discussed
above, for parallel collimators, the registered
number of pulses is virtually independent of distance
, while for convergent and Pinhole collimators.
(§4.2, part " Scintigraphic
collimators ") the dependence of the detection sensitivity on the
distance from the collimator face is very significant - Fig.4.5.6
on the right (it is derived below in the
section " Imaging properties of special
collimators "; the independence of
detection efficiency on distances is collimators is maintained
only for sources with a homogeneous area distribution of
activity exceeding the field of view) .
For PET positron emission
tomography cameras , which use electronic
coincidence collimation instead of mechanical collimators ,
the detection sensitivity is significantly higher
, approx. 7000-10000 cps / MBq 18 F - see §4.3, passage "Detection
efficiency ( sensitivity ) of PET ".
Fig.4.5.6. Dependences of the spatial resolution FWHM ( left
) and the detection efficiency of the S ( right
) gamma camera on the distance of the source from the front of
different types of collimators. In the box on the far right, the
entire detection efficiency curve for the convergent collimator
is plotted up to a distance of 70 cm, capturing a significant
maximum in the focus and the subsequent decrease.
Note: These
curves are only approximate and are more or less illustrative.
They were created by comparing and interpolating a series of
measurements of point and line sources 99m Tc with different collimators on cameras PhoGammaHP,
Nucline MB9201 and TH, Symbia T. Specific values ??of resolution
and sensitivity may differ slightly for individual cameras of
different types and manufacturers, butdependency trends
are captured objectively .
Influence of the material
environment
The analysis of the detection efficiency
(sensitivity) of the gamma imaging was performed from above in a
situation without a substance-absorbing environment - in
vacuum or in air . However, in practical
scintigraphy, there is a tissue environment
between the imaged structures with distributed radioactivity in
the organism and the gamma camera , with which gamma radiation
interacts, which mainly leads to the absorption
and attenuation of gamma radiation. During the
passage of the tissues from the point of origin towards the
camera detector, a certain amount of radiation g is absorbed
during the interaction with the tissue substance - due to the
photoeffect and the Compton scattering in the tissue. It leads toexponential
decrease in the frequency of N detected photons g with increasing
depth h of the radiolabel
distribution in the body: N = N o .e -m .h , where m is the linear
attenuation coefficient, depending on the radiation energy g and on the tissue
density ( for g 140keV 99m Tc this absorption
coefficient is m @ 0.15 cm -1 ). This attenuation of gamma-ray absorption, also
called attenuation , is manifested in
scintigraphic images by an artificial reductionthe
number of pulses from structures deposited at greater depths,
compared to structures closer to the surface. In such a case, the
statement that the detection efficiency (sensitivity) does
not depend on the distance of the displayed source from the
collimator face no longer applies . Here, the
detection efficiency decreases significantly
with the distance - depth - of the displayed source !
An approximate
dependence of S ~ R- 2
applies between the sensitivity of the camera S and its
total spatial resolution R (= FWHM) ( for parallel collimators it follows from the comparison
of formulas (4.5.1) and (4.5.2) ). So the better
the resolution of the imaging system, ie the smaller R =
FWHM, the lower its sensitivity - and vice versa. When
trying for high resolution (using a UHR collimator), this leads
to a lower pulse density in the image and therefore to higher
statistical fluctuations (higher noise) . In other words, resolution and sensitivity
they compete - collimators with better resolution have
lower sensitivity and vice versa.
Imaging
properties of special collimators - convergent and Pinhole
Imaging properties - display scale, spatial resolution and
detection efficiency (as well as linearity) - of these special
collimators differ significantly from basic collimators with
parallel holes. Especially, collimators with parallel
holes projected radioactive imaging structure of a
crystal detector in the size unchanged - in scale 1: 1
, while special collimators provide an enlarged
or reduced display, depending on the distance
from the source collimator: just this dependence
of the magnification ratio - " optical zoom
" is the main reason for their use. Spatial
resolution it though basically worsens with distance,
but at a different "pace" than with parallel
collimators. The detection efficiency
(sensitivity) of parallel collimators is practically independent
of distance, while that of special collimators increases or
decreases significantly with distance (as can be seen in Fig . 4.5.6 on the right) . We will give
an overview of the imaging properties of these collimators
according to Fig.4.5.7:
Fig.4.5.7 Geometric arrangement of holes and imaging properties
of collimators: Parallel, Convergent and Pinhole.
Parallel collimator
Display scale: M = 1 -
projects an image without resizing on the camera detector,
display 1 : 1 .
Spatial resolution : R kolim ~ d. (1 + h / L)
( formula (4.5.1) ) -
deteriorates with increasing distance from the collimator,
approximately linearly.
Detection efficiency (sensitivity) :
With collimation = (d / L) 2
. [d 2 / (d + s) 2 ]. K g . [100%] (
formula (4.5.1) is
independent of the distance from the
collimator. The imaging properties of parallel collimators were
discussed in detail above at the beginning of this chapter "
Physical
parameters of scintigraphy " in
the sections " Gamma camera resolution " and " Gamma camera detection efficiency ".
Convergent collimator
Display scale: M = (f + L) / (fh) - provides magnification of the
image depending on the distance h of the source from the
collimator.
Spatial resolution : R kolim ~ [d . (L + h) / L] . (cos
q) -
1 . [1 - (L / 2) / (f + L)] deteriorates with
distance in a manner similar to a parallel collimator.
Detection efficiency (sensitivity)
: With collimation = (d / L) 2
. [d 2 / (d + s) 2 ] . [f 2 / (f -h) 2 ] .
K g . [100%] .
With a convergent collimator, the detection efficiency increases
with distance from the collimator. Significant maximum
- up to 30 times higher than at the front of the
collimator! - reaches in the focus (which is usually at a distance of about 40-60 cm) , from where gamma radiation passes into the detector
through all the holes of the collimator; behind
the focus, the detection sensitivity decreases - see the curve in the box in Fig.4.5.6 on the right . However, this area of long distances is not usable for
practical scintigraphy, as there is already poor spatial
resolution (and the structure of openings
and partitions can be disturbed, which is projected on the
detector) .
Convergent collimators show an optimal
combination of resolution, efficiency and display scale at a
distance of about 15-20 cm, which corresponds well to the size
and depth of the heart. They were therefore often used innuclear
cardiology in entriculography and scintigraphy of myocardial
perfusion (§4.9.4 " Nuclear
cardiology ") .
Note:
The opposite - divergent - hole configuration is divergent
(now no longer used) in collimators , which provide image
reduction . A divergent collimator basically occurs when
we turn the convergent collimator and "put it in the
opposite direction" on the camera detector.
Pinhole collimator
Display scale:
M = L / h - provides reduction
or enlargement of the image, according to the distance
h. The image is mirror- inverted .
For larger distances h> L the image is reduced
, for smaller distances h <L from the collimator the image is enlarged
.
Spatial resolution: Rcollim = d . (L
+ h) /L is very good
at small distances (with a hole diameter d
= 2mm, at distances up to 10cm, a resolution of Rcollim approx. 3mm and
a total FHHM resolution of around 4-5mm is achieved) . At greater distances, the resolution deteriorates, but
is still better than other collimators.
Detection efficiency (sensitivity)
: Scollim = (d / 16h 2 ). cos 3 q . The detection
sensitivity of a gamma camera with a Pinhole collimator is
relatively high only in close proximity to the aperture (at 3 cm it is about 500 cps / MBq) , but it decreases sharply with distance . At
greater distances it is already very low (at 10 cm it is about 50 cps / MBq, at 20 cm only 10 cps
/ MBq) , for practical scintigraphy it is
completely insufficient (perhaps
only for monitoring high therapeutic activities 131 I) .
Due to these imaging properties, the
Pinhole collimator is suitable for scintigraphy of small
organs , especially the thyroid gland - see §4.9.1 "Thyrological radioisotope
diagnostics ",where it provides an enlarged image with very good
resolution and sufficient detection efficiency.
Convergent and Pinhole collimators
show a certain inhomogeneity and nonlinearity of the
image- resolution and detection efficiencydiffersomewhatat
different points of view, even at the same distancehfrom
the collimator front . These differences depend on theradial
distance ofthe display source from the axis of the
collimator, which is expressed at a trigonometric analysis ofthe
angleq between the collimator axis and the line connecting the
source and its image. Deviations in the values of resolution and
sensitivity are given bythe cosine of the angleq. In
practical scintigraphy, while maintaining optimal configurations,
they are not these derogations very significant
...
Homogeneity (uniformity) of
the camera's field of view
The scintigraphic image is created in the gamma camera in a very
complex way, the signals go through a number of precisely tuned
electronic and electro-optical components. However, the
individual links in this chain may show some changes, which may
cause deviations and defects in the images created.
Another important parameter of the
quality of scintigraphic imaging - homogeneity
(also called uniformity ) indicates whether
individual places in the field of view are imaged with the same
efficiency (sensitivity). By inhomogeneity of
the displayed field we mean local artificial
changes accumulated number of pulses, caused by local
changes in the sensitivity and linearity of the image.
Inhomogeneity is usually caused by different sensitivity of
individual photomultipliers or their different spectrometric
settings, deviations in the adjustment of electronic circuits,
defects or inhomogeneities in the collimator or scintillation
crystal.
Field homogeneity characterizes the
camera's ability to provide an accurate (ie,
homogeneous) picture of a homogeneous
distribution of radioactivity. By irradiating the camera's field
of view with a homogeneous flux of photons of radiation g, we obtain an image
of a homogeneous source , which should also be
completely homogeneous (except for
statistical fluctuations) . Possible
inhomogeneities in this image are visible visually
, but they can also be expressedquantitatively ,
mostly in percentages :
Homogeneity of the camera's field of view (integral) |
The
homogeneity of the camera's field of view is the maximum
deviation of the actual image created in response to the
homogeneous irradiation of the camera detector, from the
ideally homogeneous image: H = 100 [%] . (N max - N min ) / Nmean , where N max is the maximum, N min minimum and Nmean the average (mean) number of pulses accumulated in the pixels of the homogeneous source image. |
The overall homogeneity of the field of view
thus defined is referred to as integral homogeneity
. Since the human eye is more sensitive to differences in
the brightness of neighboring areas in the visual
assessment of images , the so-called differential
homogeneity may also be useful for evaluating the
homogeneity of the image . The following criterion was adopted
for its quantification :
Differential homogeneity
is the ratio of the largest difference in the
number of pulses in adjacent cells (row and
column) in the homogeneous source image, divided by the average
number of pulses in the image H dif = max(N i - N i-1 ) / Nmean . To reduce the effect of statistical fluctuations, the
determined number of pulses is averaged over 5
cells.
Whole and central field of view
It follows from the design of the scintillation camera that the
quality of the scintigraphic image is usually best in the central
part of the field of view, while in the peripheral parts it may
be somewhat degraded. Therefore, homogeneity (and sometimes other camera parameters) is often determined separately for the entire
field of view (UFOV - useful field of view) and
separately for the central part of the field of
view (so-called CFOV - central field of view). 75% of the entire
field of view is usually taken as the central part. For quality
and correctly adjusted (calibrated, tuned) cameras, the integral homogeneity in the central field
should not be worse than about 3,5%, and in the whole field of
view up to 5%; differential inhomogeneity in the central field
should be in the range of 1.5 - 3%.
As with the resolution, the
homogeneity of the scintillation camera is recognized by :
¨ The internal homogeneity of
the camera detector ( intrinsic )
- is given by the homogeneity of the scintillation crystal and
its light response, light collection, sensitivity and adjustment
of individual photomultipliers. It is measured by homogeneous
irradiation of a crystal without a collimator .
¨ Overall homogeneity of the
camera ("external" - extrinsic )
- given the internal homogeneity of the camera detector and
homogeneity (or inhomogeneity, defects) of the used collimator. It is measured with a
collimator mounted using a homogeneous surface source
, most often 57 Co.
Correction of the
inhomogeneity of the field of view
The inhomogeneity of the field of
view of the camera can be reduced or eliminated in two steps :
1. Careful adjustment
- matching the individual photomultipliers to the same
detection efficiency (same photopeak position), so-called tuning
. In earlier analog cameras, this was done manually using potentiometers
in the preamplifiers of the individual photomultipliers, with
oscilloscope control. Current digital cameras have a computer
procedure, which for each photomultiplier uses ADC
<--> DAC converters to adjust the gain so that the peak of
the photopeak of the radionuclide is exactly in the middle of the
set analyzer window.
2.
Computer correction
using a suitable matrix of correction coefficients
g ij to
correct the remaining inhomogeneous response of the detector -
Fig.4.5.8. The accumulated numbers of pulses A ij in the individual
elements (i, j) of the original uncorrected image are multiplied
by the correction coefficients g ij from the correction matrix, thus creating an image *A ij corrected for
inhomogeneity :
*A
ij
= A ij . g ij .
![]() |
Fig.4.5.8. Computer correction of
gamma camera image inhomogeneity. Left: Image a ij of a homogeneous source, showing significant inhomogeneities. Middle: Matrix of correction coefficients g ij . Right: Multiplying by correction factors creates a corrected image *a ij that is already homogeneous. Note: Instead of the usual luminance modulation , an isometric display is used here , where the height of the elements (pixels) above the base is proportional to the number of pulses contained. The curves at the top are cross-sections , taken through the center of the images. |
The matrix of correction coefficients g ij is obtained from the
scanned image h ij of a homogeneous source (which we know should ideally be homogeneous -
constant) as its normalized inverse
matrix . The correction coefficients g ij for individual elements (image
cells - pixels) i, j are calculated as the
ratio of the average number [pulses /
pixel] in the whole field of view W
to the number h ij [pulses /
pixel] in a given image location (i, j) of
a homogeneous source:
g
ij
= ( i, j Î
W S h ij) / (N. h ij ) ,
where N = W S I + W S j is the total number of
elements of the visual field image W . In
places (i, j) of the field of view with reduced detection
efficiency, the correction coefficients g ij
are slightly higher than
"1", in places of higher efficiency they are slightly
lower than "1".
The image of the homogeneous source
must be obtained under the same physical conditions *) as the
images that we want to correct with the resulting matrix. This
correction matrix g ij is stored in the memory of the acquisition computer and
its accumulated number then multiplies the
accumulated number of pulses at a given location of the field of
view during scanning (acquisition).
*) What the homogeneity of the field of view
depends :
The homogeneity of the gamma
camera imaging depends (in addition to the
mechanical, detection, optical and electronic properties of the
device) also on a number of parameters and setting
conditions - on the width and symmetry of the analyzer window
setting, on the radiation energy, on the amount of scattered
radiation, on the frequency of the pulses, on the collimator
used. It can also change over time due to instabilities in the
detector and electronic circuits. If a sufficient number of
accumulated pulses in the image (information density) is not
achieved, statistical fluctuations may also appear as
inhomogeneity - the correction matrix must therefore be recorded
with the highest possible number of pulses, min. 10,000
imp./pixel. These important aspects of measuring and correcting
inhomogeneity are discussed and documented on experimental
scintigraphic images in the section " Testing and calibration of camera image
homogeneity ", section
" Dependence of inhomogeneity on physical
conditions ".
Scintigraphic
digital cameras have both of these operations 1.
and 2. covered by the procedure calibration homogeneity
and tuning of photomultiplier tubes ( tuning ).
Regular testing of the
homogeneity of the gamma camera imaging
Virtually all disturbances and anomalies in the imaging
properties of the camera are most sensitive to the
homogeneity of the field of view. To ensure high-quality
scintigraphic imaging, it is therefore necessary to perform regular
homogeneity testing - and of course also after each
electronic intervention in the circuits of photomultipliers,
amplifiers and ADCs. In case of degraded homogeneity, it is
necessary to adjust or recalibrate the photomultipliers (tuning), updating the correction
matrix, in case of gross abnormalities electronic intervention or
correction. It is useful to archive the results
of homogeneity testing for a long time , or graphically
, for the analysis of the time trend and the detection
of causes or. deterioration of homogeneity .
How homogeneity
measurements are performed using point and area sources (and
recommended testing intervals) is described in the work " Phantoms
and phantom measurements in nuclear medicine ", section
" Testing and calibration of camera
image homogeneity ".
Linearity of gamma camera imaging
Another parameter of scintigraphic image quality indicates
whether the spatial scales and proportions in the object are
displayed faithfully and without distortion -
linearly. It is therefore the ability of the camera to display
the distribution of radioactivity without positional distortion,
to display the line source as an exact line. Special phantoms are
used to assess (and possibly quantify) the linearity of the
scintigraphic image, in which a regular geometric
structure of the radioactivity distribution is realized
. It can be either a system of a larger number of regularly
quadrangularly distributed point sources, or a system of linear
(linear) sources (so-called bar-phantom
, mostly transmission). The most perfect
phantom for the analysis of linearity of scintillation camera
imaging is the cartesian linear grid - it is described in the work " Phantoms and
phantom measurements in nuclear medicine ", passage
" Analysis of linearity of gamma
camera imaging " .
The scintigraphic image of such a
regular geometric structure should also show geometric
regularity . Possibly. the nonlinearity of
the image manifests itself in this image as distortion
and irregularities in the geometric arrangement.
We can monitor them either visually or evaluate them
quantitatively :
Linearity of scintigraphic imaging (spatial) |
The
linearity of the scintigraphic image is characterized by
the maximum deviation of the scintigraphic image of the
linear distribution of radioactivity from the exact
linear form: L [mm] = max (X -X lin ) , where X are the actual coordinates in the image and X lin are theoretical coordinate values corresponding to the exact linear curve. |
Sometimes linearity is also expressed as a
percentage, ie L = 100 [%] .max (X -X lin ) / X lin . Linearity is analyzed in two "X"
and "Y" directions perpendicular to each other
. As with resolution and homogeneity, linearity is also given for
the whole and central field of view, or in addition to the total
(absolute) linearity, the differential linearity is also given.
The spatial linearity of the display is better than about 4 mm
for high-quality and correctly adjusted cameras.
Linearity <---> Homogeneity
The linearity of the image and the
homogeneity of the sensitivity of the field of view of the
scintillation camera are closely related.
Irregularities in the efficiency of registration of
scintillations from different places of the scintillation crystal
of the camera by a system of photomultipliers will be reflected
in the image as geometric nonlinearity and at
the same time as inhomogeneity in the density of
registered pulses. It can be said that under normal circumstances
the nonlinearity of the image is the main source of image
inhomogeneity. Deviations in the regular arrangement and size of
the holes and baffles in the collimator can also
contribute to the inhomogeneity of the scintigraphic imaging ,
especially in the case of mechanical damage to the collimator.
In practical testing of the
properties (quality) of a scintillation camera, the linearity of
the image is rarely determined, as it is difficult and, in
addition, a small change in linearity (which would be difficult
to demonstrate in targeted linearity measurements) is
significantly more pronounced in inhomogeneity
of the field of view. The only case of visually
observable systematic nonlinearity is
scintigraphic images with convergent and Pinhole
collimators (§4.2, section " Scintigraphic
collimators ") .
Tomographic
resolution, homogeneity and linearity
For all the above parameters of the classical scintillation
camera, we had in mind the usual planar
scintigraphic imaging. The quality of the tomographic
image is basically described by the same physical
parameters as in the planar image. However, the planar parameters
of the SPECT *) scintillation camera detector cannot always be
transferred directly to tomographic images of transverse
sections, which arise from a complex reconstruction procedure
from many planar images at different angles. Although the
parameters of the camera detector are also decisive for the
quality of tomographic images, some other physical and technical
aspects also cooperate here.
*) Of course, there is no planar image for
the PET camera, all parameters are measured in tomographic mode,
on images of transverse sections - see §4.3, section " Spatial
resolution of PET ".
The spatial resolution of the camera
in the planar image is also decisively reflected in the
tomographic image. The radial tomographic resolution
in the transverse section image is approximately (1.1-1.3) times
the total resolution of the camera, but it can be possibly
aggravated by mechanical shifts of the center of rotation during
the acquisition of SPECT examination. The resolution in the axial
direction is given directly by the resolution of the camera with
the collimator used (at a given distance of the displayed
structure from the front of the collimator).
Possible local defectin
the homogeneity of the camera, it is projected as an annular
artifact in the rotation of the camera in the
cross-sectional image . The inhomogeneity in the cross-sectional
image can then still be caused by the absorption of g radiation
(attenuation) depending on the depth of deposition of the
respective radioactivity distribution in the tissue.
Tomographic resolution and homogeneity are
measured using special aids - phantoms, mostly
cylindrical, containing tubes (line sources), various rollers and
beads of various sizes, as well as free space for homogeneous
distribution of the radio indicator. The most commonly used is Jasczak
's phantom . The phantom is filled with a radionuclide solution
(usually 99m Tc or 18F), its SPECT or PET scintigraphy is performed and the
resolution and homogeneity are evaluated on the reconstructed
images of transverse sections in the appropriate places in a
manner analogous to the planar images. Tomographic phantoms are
described in the work " Phantoms and phantom
measurements in nuclear medicine
", section " Tomographic phantoms ".
Energy resolution and dead time of
the camera detector
In addition to the above basic parameters - spatial resolution,
homogeneity and linearity of images, which have a primary effect
on scintigraphic image quality, scintillation camera also
considers detection parameters , describing its
properties in terms of scintillation detection and radiation
spectrometry . . Although these parameters are secondary and auxiliary,
they can indirectly affect the quality of the display. Adverse
changes may also indicate a malfunction of the scintillation
camera.
The energy resolution
of a scintillation camera not only allows the separation of
different lines of gamma radiation (eg when
simultaneously imaging two isotopes), but
mainly determines the ability of the camera detector to separate
Compton - scattered radiation g from direct non - scattered
radiation (discussed above in the
section " Spatial
resolution of the gamma camera
", passage " Compton - scattered radiation " , Fig . 4.5.4) .
With the correct adjustment of photomultipliers, the total energy
resolution of classic gamma cameras (Anger
type) is about 9-12% (for
new semiconductor CZT cameras it is about 5%) .
Dead time
(sometimes called time resolution)) is the time
for which the detector processes the signal from the arrival of
one detected quantum of radiation and is not able to register any
other quantum. In scintigraphy, it can be applied to sources with
high activity , ie with a high flux of photons
of radiation g - at a frequency of many tens of thousands of registered
photons per second. The dead time of the camera leads to a violation
of the linearity of the dependence between the activity
in the source and the registered frequency of pulses in the
image, which may distort the results of the analysis of the
dynamics of the investigated processes. For quantitative dynamic
studies, especially radiocardiographic studies, the need for a correction
for dead time may arise . The important thing here is overalldead
time of the whole system camera + computer. For older types of
cameras, the total dead time was about 5 m s, for newer cameras it is
already reduced to about 1 m s. Overloading the gamma camera with high radiation
intensity also leads to deterioration of imaging
properties - resolution, homogeneity, image contrast.
All registered photons of all energies contribute to the dead
time, ie not only the corresponding photo peak in the analyzer
window, but also Compton scattered radiation.
Note:
For details on the paralyzable
(cumulative) and non-paralizable nature of dead time, as
well as methods for its measurement and correction
for dead time, we can refer to Chapter 2, passage "Dead
time of detectors ".
Energy resolution and dead time are
measured by the spectrometric methods described
in Chapter 2" Detection and spectrometry of
ionizing radiation ", in
particular in §2.4"
Scintillation detectors ".
The issue of measuring the
imaging properties of gamma cameras and the practical
implementation of testing is discussed in a separate work
" Phantoms and phantom
measurements in nuclear medicine ".
4.6.
Relationship between scintigraphy and other imaging methods
Scintigraphy is just one of several other imaging
diagnostic methods used in medicine. Each of these
methods has its uses, its advantages and disadvantages. In
principle, diagnostic imaging methods can be divided into two
groups :
¨ Anatomical-morphological
,
which show mainly the size and structure of tissues and organs.
However, anatomical imaging lacks a functional aspect - it does
not allow to recognize the biological nature of the displayed
pathological structure.
¨ Functional-metabolic
,
which map blood circulation, metabolism, drainage and other organ
functions. However, functional imaging usually does not allow the
exact localization of a pathological event or lesion in the
organism - there is no "background" of other structures
that are not displayed (because they do not have the appropriate
"function"). In addition, functional imaging generally
has a lower resolution than anatomical imaging.
To make a correct diagnosis, it is
necessary to assess both anatonic-morphological and functional
and metabolic symptoms of the disease. Only the combination of
both mentioned images will make it possible to recognize the
biological character of the depicted deposit and its exact
location.
To clarify the position and role of scintigraphy in
the spectrum of other diagnostic methods, we will briefly compare
the principles and diagnostic
statements.the most important imaging methods.
X-ray
imaging
The oldest and most frequently used imaging method so far is X-ray
imaging (see §3.2 " X-ray diagnostics "),
whether it is planar or tomographic CT imaging. The penetrating
X-rays generated in X-rays pass through the examined object
(organism tissue), while part of the radiation is absorbed
depending on the tissue density , while the
remaining part passes through the tissue and is displayed
either photographically or on a luminescent screen, or more
recently by electronic detectors . This creates an X -
ray image of the examined tissue, which is a shadow
density image showing differences in tissue density
. In certain cases, the contrast of the image can be artificially
increased by applying suitable contrast agents .
Tomographic X-ray In addition, the CT image provides images of transverse
sections with high resolution
(approximately 1 mm), from which a three-dimensional image of the
examined area can be composed.
Ultrasonic sonography
Ultrasound is a mechanical (acoustic) wave of a
substance (air, liquids, solids) with a frequency higher than the
sound audible to the human ear, ie higher than 20kHz. In matter,
a wave propagates by oscillating its particles around an
equilibrium position. In gases and liquids it propagates as longitudinal
waves , in solids it can also have the character of transverse
waves . In medical diagnostics, ultrasound with a frequency of
1-15MHz is usually used. At higher frequencies, better spatial
resolution can be achieved (due to the shorter wavelength), but
more ultrasound is absorbed in the tissue.
Ultrasound sonography or ultrasonography
is based on the propagation of sound wavesof
high frequency (several MHz), ie ultrasound , in
the elastic environment of tissues and its reflections on
inhomogeneities . The speed in the propagation of
a wave in an elastic medium is given by the relation v = Ö(M / r ), where M is the elasticity
(Young's modulus) of the medium and r its density
(specific gravity). When an ultrasonic wave strikes an area with
different density or elasticity
- the acoustic interface , there are changes in
the speed of propagation, refraction and reflection of the wave (related to the well-known Huygens principle). The reflected ultrasonic waves carry information about
the presence of structures of different density and elasticity.
Ultrasound sonography creates an image of these structures in the
examined tissue by echolocation *) of reflected
ultrasound waves. The reflected signals - acoustic echoes
- correspond in their time sequence to the spatial distribution
of reflecting structures in the investigated environment.
*) Echolocation
is a way of obtaining information at a distance, where a sound
is transmitted to the monitored environment ,
which is partially reflected from a possible
object back to the place of transmission and is captured and
evaluated there. From the time delay which
elapses from the moment of sound transmission to the moment of
reflection of the reflected wave (echo ), the distance
of the reflecting object can be determined . In nature,
this principle is used by dolphins and bats for orientation and
foraging. In marine technology, so-called sonar is used
, among other things to measure the depth of the sea. On the same
principle Radar operates radar , which
uses sound instead of electromagnetic radiation - radio waves.
Transmitting piezoelectric crystal
of the probe, put into mechanical contact with the body
surface (good passage of waves into the
skin is ensured by a layer of special gel),
is periodically deformed by the action of an alternating electric
voltage applied to its opposite electrodes, and this mechanical
disturbance (vibration) emits an acoustic wave into the tissue.
In water and tissue, sound waves propagate at an average speed of
1550m / s. As ultrasound passes through the substance, it is
absorbed, scattered, bent, and partially reflected back.
Reflection occurs at the interface of tissues
with different densities and elasticities, in which ultrasound
propagates at different speeds - ie tissues with
different acoustic impedances*). The reflected
ultrasonic waves return and cause vibrations of the piezoelectric
crystal (transducer) in the receiving part of the ultrasonic
probe, which generates an alternating electrical signal of
appropriate frequency, amplitude and time delay at the crystal
electrodes, which is further electronically processed.
*) The so-called specific acoustic
impedance is important for the propagation of tissue
ultrasound , which is the product of tissue density and
ultrasound speed: Z = r .v = Ö(M. r ). It gives the specific "wave resistance" in
the propagation of ultrasound in analogy to Ohm's law of
electricity. The acoustic impedance Z defined in this way
is somewhat analogous to the reactance in electronics
(capacitive or inductive). Viscosity the environment
leading to the absorption and attenuation of ultrasound is
analogous to the active ohmic resistance . The greater
the difference in acoustic impedances, the greater the intensity
of the reflections - echogenicity .
The probe attached to the body
surface transmits short (millisecond) ultrasound
signals at fast regular intervals , the electronic receiving
probe records the reflected signal (" echo
") and the electronic apparatus evaluates the time and
position differences of the transmitted and
reflected signal and creates an image of structures
on the screen according to their density and elasticity (so-called echogenicity or acoustic
impedance) . Ultrasound images of
echogenicity are displayed on most probes primarily in the form
of a circular segment in polar coordinates (r, j ) centered at the
point of attachment of the receiving probe. The radial coordinate
r - the depth in the tissue - is derived
from the time delay D t between the transmitted
ultrasound signal and the reception of its reflection: r [mm] = 1.55 . Dt [ m s] (at the average speed of
ultrasound in the tissue approx. 1550 m / s). Angle jit is determined
for simple devices with one receiving crystal by oscillating
rotation of the probe (manual or motorized), for probes with more
directional receiving crystals it is determined electronically. The
brightness of the individual points of the sonographic
image is modulated by the intensity of the
received reflected ultrasound signal (this
intensity should be corrected for the depth absorption of the
ultrasound, see below) , ie the echogenicity
*) of the corresponding sites in the tissue. The sonographic
image thus captures the spatial distribution of
structures with different densities and elasticities in
the examined tissue.
*) Formations that reflect
ultrasound more or less than the surrounding tissue are called hypoechogenic
or hyperechogenic . Higher or lower echogenicity
is not in itself a "diagnosis", but it is an important
feature by which the character of the examined area of tissue can
be taught.
Some technically advanced systems
have a computer transformation of the image into the usual Cartesian
coordinates , providing a much clearer presentation. The
rectangular view is provided by probes with a linear arrangement
of a plurality of receiving elements.
The signal coming from a greater depth is significantly attenuated
by absorption in the tissue *) (double attenuation -
transmitted and reflected signal), so for objective display, a correction
for attenuation is performed , either time (longer signal
reception time from greater depth) or computer.
*) For absorptionof
ultrasound in the environment occurs by the fact that due to the
internal friction of the oscillating particles, part of the
mechanical energy of the waves changes into heat. The rate of
absorption of ultrasonic waves is determined by the exponential
law I (d) = I 0 .e - m .d , wherein I 0 is the original (initial) intensity, I (d) the
intensity at depth d , m
coefficient of absorption
. The absorption coefficient m
depends on the type of substance
(its viscosity) and on the frequency . In most
substances, the attenuation by absorption is directly
proportional to the square of the frequency.
The great
advantage of ultrasonography is the simplicity of
its design, non-invasiveness and
unpretentiousness for patients. The method is completely safe and
harmless (ultrasound intensity reaches a maximum
of 1 mW / m 2 ), it does not burden the body with ionizing radiation.
Therefore, in diseases related to morphological and anatomical
changes, ultrasound examination is usually included at the beginning
of the diagnostic chain . Ultrasonography is also widely
used to evaluate the course of pregnancy.
Doppler
ultrasonography
Modern sonographic instruments also allow the analysis
of the frequency of the received ultrasound signal: the
frequency of the signal reflected from a moving object
is slightly increased or decreased due to the Doppler
effect
*), depending on whether the subject is moving toward or away
from the receiver. The Doppler frequency shift of the reflected
ultrasound can be used to modulate a common
echogenic anatomical image (color modulation is used) and thus
obtain a velocity map of the movement of
structures and fluid flow in the examined object. With the help
of this so-called Doppler ultrasonography, it is
possible in cardiological diagnostics ( Doppler
echocardiography ) to detect, for example, movements of
heart walls and valves, or jets of blood from under the heart
valve during regurgitation. It is also possible to monitor the speed
of blood flow in the venous system.
*) Doppler effect:
If the wave source moves of a certain constant frequency fo towards the observer
(receiver), this observer registers a higher frequency f
than the source actually emits. Conversely, when the source moves
away from the observer, the registered frequency is lower than
the actual one. The difference between the actual f o and the observed f
frequency (Doppler frequency shift) increases in proportion to
the velocity V of the source relative to the observer: f =
[1 + (V / v)] . f o , where v is the speed of propagation of a given
wave. This rule also applies when the source of the received wave
is the reflection of the wave from a certain
moving object (including a flowing liquid). By measuring the
frequency difference of the primary transmitted wave and
the reflected wave ("echo") we can thus determine the speed
of movement of the reflecting body.
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 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 knowledge of nuclear physics
- the properties of atomic nuclei. A physical phenomenon called nuclear
magnetic resonance - NMR, was investigated in the 1940s (F. Bloch, EMPurcell) 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" momentum - spin
(nucleons belong to fermions with spin 1/2, see §1.5 " Elementary
particles "). According to the
laws of electrodynamics, this rotational momentum of nucleons
creates - induces - its own elementary magnetic
moment m p = 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 1 H, then in 13 C, 15 N, 19 F, 23 Na, 31 P, etc.
*) Nuclear magneton m Nis a
physical constant expressing the proton's own
dipole magnetic moment induced by its spin: m N = eh / 2m p
, where e is the elementary electric charge (proton,
electron), h is the reduced Planck's
constant, m p is the rest mass of the proton. In the
system of SI units, its value is approximately m N =
5,05.10 -27 J / T. It is analogous to
the Bohr electron magnet m e = eh / 2m e,
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 m n = -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 1 gram 10 22hydrogen nuclei,
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
w L =g .B , or f L = g .B /
2 p ,
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 - RF, or radio-frequency - RF) (whose frequency resonates with the above-mentioned Larmor
precession f L
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 **). The relaxation signal (sometimes
referred acronym FID - Free Induction Decay) , a sine wave with an 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 type analysis of the
composition of substances : what elements and in 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) . It is proportionally lower for heavier cores.
**) Phasing of a large number
of nuclei : The NMR receiving coils are of course
not able to detect the relaxation radiation of one or more
nuclei. To obtain a measurable signal, deexcitation of a
large number of nuclei (> about 10 12 ) is required,
namelysynchronously and at the same stage ! If
phasing occurs, the MNR signal drops sharply or disappears (cf.
below " Relaxation times - T2 )
General 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 corresponding
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 ...
Giant. 3.4.4. Nuclear magnetic resonance -
simplified schematic representation.
a) The magnetic moments of the nuclei in the analyte
normally have chaotically scattered directions.
b) By the action of a strong magnetic field B
, the mag. moments of nuclei partially orient in the direction of
the vector B .
c) By emitting an 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 will emit an
electromagnet when they return during the precessional rotation.
NMR signal with exponentially decaying amplitude.
d)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 by increasing the
frequency of the excitation RF signal, 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 - 42.6MHz 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 B o , but B = B o . (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 decay 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 T 1 and T 2 .
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.
Illustrates the speed with which the deflected core while
relaxing supplies energy electromagnetic waves, and the ambient
temperature, the longitudinal magnetization in the axial
direction from the initial value to the M by returning
exponential law: M Z = M a (1 - e -t / T 1 ) . It is
defined as the time taken for the longitudinal magnetization to
relax (1-e)-times the original valueM o , 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 phasing out of the
precessional motion of magnetic moments, the magnetization
decreases in the transverse direction xy: M XY
= M XYo e - t / T 2 . T 2 is defined as the time during
which the transverse magnetization M XY 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 T 1 and T 2 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 T 2 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 the main constant
homogeneous field B by superimposing an additional gradient
magnetic field in the axis direction X, Y, Z. 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 B z (z) in the Z
direction . His superposition with the main mag. fieldB o causes the actual
"local" value of the magnetic field B
= B o + B z (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 fulfilled. By changing the frequency f of
high-frequency excitation pulses, or the intensity of the
longitudinal gradient field B z , the position of the
z layer in which the magnetic resonance response signal is
generated changes . 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 imaging 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 the determined intensity of the dependence of the
relaxation of the NMR signal, and also the decay time. Changing
these gradient fields to obtain data for each place in the layer of
and computer reconstruction of the obtained cross-section
image of proton density of the examined layer from
(obr.3.4. 4d right). electronic analysis of relaxation times of
the NMR signal is also generated even images of cross sections in
relaxation times T 1 and
T 2 (referred to as T
1 or T 2 - weighted images). A plurality
of images of cross sections for the different values z are
then creates 3-dimensional tomographic image of
investigated areas 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 of PD
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
T 2 related to the chemical
composition and structural state of the
tissue in individual places. Such images are called T 1 and T 2 - weighted
.
The extent to which the proton
density PD will be represented in the resulting MRI image, the
times T 1 and T 2 - how and how 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 cancer
diagnostics.
Note: As
with X-ray diagnostics, NMRI 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
sufficienthigh 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 cores. 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 mag flips.
torque and the stronger the response signal and more time is
needed 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 fully indicative 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 laws, 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 T
1 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 T 1 . This simplest MRI technique is
now rarely used, it has been replaced by the inversion-recovery
sequence below, providing higher contrast.
-> Spin - echo a 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.
T 2 *) 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
(B lood 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 realization 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 .........
Thermography
Thermography is a method of imaging the temperature
distribution on the surface of analyzed objects. The
medical use of thermography is based on the fact that some pathological
events in the body are accompanied by changes in
temperature (eg the inflammatory process by raising the
temperature), which are also reflected on the surface of
the body at a location above the bearing. Thermographic
imaging can be performed in two ways :
¨ Contact thermography
using liquid crystals. Liquid crystals are
substances that behave mechanically as liquids, but optically as
solid crystalline substances (they appear optical anisotropy).
Thermography uses the properties of some liquid crystals that,
depending on the temperature, they color differently
(thermo-mechanical changes affect the interference of incident
light). With a suitable composition, liquid crystals can be
prepared, which make it possible to display different temperature
ranges in color. The liquid crystals were previously painted
directly on a black-backed skin. They are now applied on a
special flexible foil, which is applied to the skin.
¨ Infrared electronic thermography
by scanning infrared radiation from the surface
of examined bodies with a special video camera
sensitive to infrared radiation. Each body of non-zero
temperature emits infrared radiation (thermal radiation)
- electromagnetic waves of a continuous spectrum with a
wavelength greater than visible light. Its intensity is greater
the higher the surface temperature; as the temperature rises, the
average wavelength decreases. It is used in industry and
construction - for example, in the infrared image of a heated
building, the places of poorer insulation with higher heat
leakage are clearly shown.
Thermographic imaging of
the body surface, obtained with a sensitive infrared camera, may
show areas with abnormally elevated temperatures
(differences may be less than tenths of ° C), which may indicate
inflammatory or tumorous
process in the tissue beneath this precinct. Alternatively, low
temperature areas may indicate a perfusion
disorder , perhaps due to occlusion of a blood vessel
(eg, venous thrombosis).
Electroimpedance
imaging
Certain information about the properties of tissues can be
determined by sensing the local electrical conductivity
, resp. impedance of examined 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. This method is (so far rarely) used
in mammography .
Complementarity
of methods
From this brief overview of the principle and statements of
several different imaging methods, we see that each of the
methods looks at the examined tissue or organ from a different
"viewing angle" - it examines a different
aspect of morphology or function . In
other words, it can be said in general that the individual
imaging methods are complementary to each other
- in certain aspects they complement and compose
a diagnostic "mosaic", which is then interpreted
by an experienced clinician into the final diagnosis
, from which the appropriate method of therapy
is based .
Status
and role of nuclear medicine
Scintigraphy does not provide images with as high a resolution as
CT, it does not recognize density, temperature or mechanical
consistency of tissues. The main "floor" of nuclear
medicine is the non- invasive imaging and quantification
of structures and processes in the body, which are
characterized by a specific function and metabolism.which
can be "traced" by a suitable radio indicator and
displayed by external detection of gamma radiation. The degree of
local accumulation of radiopharmaceuticals depends on the
intensity of local metabolic and functional processes in organs
and tissues. Possible anomalies and malfunctions can be located
and quantified using scintigraphic imaging. Disorders of function
in many cases precede structural disorders - anatomical and
morphological. Therefore, pathological events can be detected by
nuclear medicine methods sometimes earlier than other imaging
methods (typical examples are bone
metastases of breast, prostate or lung cancer) . The perfusion of tissues or organs
and the dynamics of blood flow through
individual parts of the heart and blood vessels can also be
analyzed very well by nuclear medicine methods.
Fusion
of PET and
SPECT images with CT and NMRI images
In this §4.6 "Relationship between scintigraphy and other
imaging methods", it was continuously discussed how the
individual diagnostic methods complement each other
in creating a comprehensive picture of the healthy or
pathological condition of individual organs and the whole
organism. The main goal is to combine anatomy with
physiology , or in other words function with
morphology , in order to better clarify the location,
biological character and origin of pathological foci and
abnormalities - assign the foci shown on the
scintigram to specific anatomical structures in
the organism. In the field of imaging methods, such multi-modalities
examinations and compare CT , SPECT , PET , NMRI and sonography images .
Scintigraphic images provide
important information about the functional status of tissues and
organs, but are usually unable to provide sufficient anatomical
information about the exact location of pathological
abnormalities (lesions) imaged scintigraphically. Radioactivity
does not enter the surrounding anatomical structures (eg skeletal) , which do not
capture the radioindicator and are not visible in the
scintigraphic image. For a better and clearer comparison of the
character, size and location of the displayed structures, it is
optimal to perform a simultaneous displayPET +
CT images, or NMRI or CT + SPECT, into a single suitably
color-modulated image - so-called image fusion .
Individual images are overlaid in various color
combinations. In these images, where it is possible to
continuously modulate the percentage of
individual images (from what percentage
will one image be projected into another) ,
we can observe the correlation of physiological
and anatomical-structural information. This will make it possible
to accurately locate scintigraphically displayed
lesions in terms of space and anatomy.
These mergers often encounter the
problem that images from different modalities were captured at
different times, at different display scales, and with different
geometric configurations.patient relative to the imaging
device. Sophisticated computer graphics programs are able to
perform affine and conformal transformations of images
to correct geometric effects - scaling and relative position of
displayed structures (translation,
rotation, reorientation, enlarging or reducing,
cross-correlation) and achieve a relatively
good "matching" of images, but some deviations from the
exact overlap of the corresponding structures may persist. For
this process of geometric alignment of images ,
the somewhat misleading name of registration
or normalization of images is sometimes used .
Procedures of this kind are applied in a number of areas of
computer image processing(eg in panoramic
photography, cartographic imaging, astrophotography) .
Hybrid
tomographic systems - combination
of PET+ T, SPECT+CT, (PET+MRI)
In order to eliminate these problems, as
well as to operatively and quickly achieve comprehensive
diagnostics, an effort was made to combine some
imaging methods into one device . Imaging device
manufacturers have developed so-called hybrid systems
, combining pairs of PET + CT or SPECT + CT devices (a combination of PET + MRI is also mentioned below) . These combined systems have three basic advantages :
*) Geometric alignment of CT images
with SPECT and PET images
When fusing functional scintigraphic SPECT or
PET images with anatomical CT images, it is important that the
structures shown in both modalities overlap geometrically
- they appear in the same place of the image. To calibrate
this precise harmonization of mutual overlapping position
shown structures SPECT <-> CT or PET <-> CT hybrid
device used scinticrafické + CT view point
sources filled with mixture of radionuclides ( 99 m Tc, or 18 F) and the contrast
agent ( " Phantoms and phantom measurements ", part"Tomographic
phantoms for SPECT, PET, CT ,
"passage" Geometric alignment of images CT images of
SPECT and PET ").
This process geometric alignment of the images are sometimes (not very accurately) called spatial
normalization and registration of images.
Hybrid combination PET / CT
in in recent years, it has become a standard
feature of PET-nuclear medicine workplaces, and its benefits are undeniable
, especially in the field of tumor diagnostics. sometimes used
simpler and cheaper "low-density"("low
dose", "non-diagnostic") localization
CT, but this solution is not entirely optimal; For added
versatility, the preferred hybrid combinations with full diagnostic
multi-slice CT (which itself may be
reducing the anode current operated in the low-dose mode) . In many cases, the combination of SPECT / CT helps to
refine the diagnosis by physiological-anatomical correlation
. Hybrid
combination of PET + NMRI
Nuclear magnetic resonance imaging (NMRI) provides, compared to
CT, better soft tissue resolution, which is particularly
advantageous in oncological diagnostics. In recent years,
therefore, efforts have been made to develop a hybrid combination
of PET / NMRI imaging devices. However, the
direct combination of "classical" PET and NMRI
technologies into one system faces a major problem: the strong
magnetic field of the NMRI superconducting electromagnet affects
the movement of electrons between dynodes in photomultipliers of
ring scintillation detectors in PET (Lorentz
force acts on electrons) ; in a
strong magnetic field, the photomultipliers stop working
. However, technologies have been developed to combine PET with
NMRI :
¨ Semiconductor photodetectors . Instead
of conventional photomultipliers, APD (Avalanche
Photo Diode) semiconductor photodiodes or SiPM
semiconductor photomultipliers are used in the PET ring detector
(see §2.4, section " Photomultipliers "). The photodetectors are sensitive to magnetic
fields.
¨ Multipixel fully solid-state detectors
(e.g. based CZT) annihilation photons (instead
scintiblokù BGO / LSO with conventional photomultipliers or
SIPMA) . In addition to better detection
efficiency and the spatial resolution can thereby be achieved
coincidence somewhat shorter time (for
better TOF) . the advantage of
semiconductor detectors is also their independence on the
magnetic field , which allows the use in hybrid systems of
PET / MRI .
Note .: Initially proposed
some temporary suboptimal solutions :
- NMRI magnetic field modulation, which
would turn on only while magnetic resonance imaging was created,
while it would be turned off when capturing PET images.
- Optical fibers - light scintillations from annular PET
scintillators would lead to photomultipliers located outside the
NMRI magnetic field using optical fibers.
These bizarre makeshift suggestions did not
work, did not work and were not used in practice. The only real
possibility of a hybrid PET / MR combination is the use of magnetically
independent semiconductor photodectors, or better
directly semiconductor gamma detectors.
A hybrid PET / MRI combination is sometimes abbreviated
as mMRI - Molecular Magnetic Resonance
Imaging : MRI provides imaging of morphological and functional
details of tissue, PET shows tissue metabolism at the
molecular-cellular level. The integrated PET / MRI system allows,
in some cases, more accurate identification and determination of
the extent and characteristics of malignancies, which can help
plan effective treatment and eventually its effect. The system
can also be used in neurology and cardiology.
However, the combination of PET
/ MRI also has some disadvantages , due
to which it cannot yet function as a routine alternative to PET /
CT (so far it is more of a specialized device) :
-Relatively long magnetic resonance imaging time, which
limits the number of patients due to the work shift and the short
half-life of 18-F; it also causes blurring of images in the chest
and abdomen with breathing movements. CT is unevenly faster,
practically does not prolong PET examinations and provides sharp
images of moving organs.
- Common contraindications MRI (pacemakers,
metal implants, stents) , while CT has no
contraindications (other than PET).
- MRI images do not yet provide accurate density maps for
the necessary exact correction of PET images to attenuate g radiation in
tissue, as provided by CT images.
- Significantly higher purchase price of equipment and
high operating costs of MRI compared to CT.
The argument that MRI is a
non-radiation method with zero radiation dose is not significant
here. PET gammagraphy alone is loaded with a relatively higher
dose, and due to the composition of patients, who will usually be
treated with radiotherapy with many times higher doses, the dose
from CT imaging is irrelevant.
Therefore, installing
a hybrid PET / MRI combination as the first or
only PET device in a complex oncology center is not
entirely optimal . A more suitable variant is the
currently proven PET / CT combination; for the extension of
complex diagnostics supplemented by a high-quality nuclear
magnetic resonance device (3 T magnet) in a separate room or
workplace, using computer fusion of PET + CT images with MRI
images. In larger workplaces, even a hybrid combination of PET /
MRI can be successfully used for many indications.
A hybrid
combination of imaging diagnostic and radiation
radiotherapy technologies is discussed in §3.6
"Radiotherapy", section " Modulation of radiation beams ".
4.7.
Visual evaluation and mathematical analysis of diagnostic images
Diagnostic images X-ray
(planar or CT) , scintigraphic (planar, SPECT, PET) and magnetic
resonance MRI can carry a lot of information
about physiological or pathological anatomical-morphological
situation and structure of tissues and organs, their function and
metabolism , the presence of abnormalities and pathological
lesions. To obtain this important diagnostic information - image
evaluation - there are basically two ways to proceed :
¨ Visual evaluation
by an "experienced eye" of an erudite expert
in the field of nuclear medicine, X-ray or MRI diagnostics. This
is the basic method of evaluation (and
before the era of digital imaging it was the only way ...) . An experienced radiologist can recognize a
number of abnormalities, disorders, lesions on well-scanned
images (with the necessary
processing - brightness modulation, filtration, corrections) . Such a description can then be an important guide for detection
and proper therapy or. pathological conditions,
as well as for assessing the response and efficacy of therapy.
¨ Quantitative
processing,
which with the help of mathematical-computer analysis
provides quantitative parameters about the
densities of various tissues and districts, radionuclide uptake
and function examined organs, the course and
rates of functional-metabolic processes ("molecular"
imaging) . The quantitative results
obtained in this way complement and refine the
visual assessment (eg the degree of
metabolic activity of the visually recognized lesion) , but they may also have their own importance for the
assessment of functional processes in the organism (functional state of the kidneys, heart, liver) . New special methods of filtering and computer image
processing can also "pull" and emphasize some details
, indistinguishable from native images - and thus help visual
evaluation.
Multifactorial
statistical analysis of images
In addition to the basic structural and functional information
mentioned above, diagnostic images may also contain some additional
information that is not directly visible (and does not
result explicitly from quantitative analysis), but can in
principle be extracted by special sophisticated computer methods of
structure recognition and character analysis
in paintings. An example could be an analysis of the relationship
between tumor size and shape (surface and volume), or the degree
of internal homogeneity or heterogeneity of the lesions shown. It
is also possible to analyze the topological shape and compactness
(such as the Hausdorf analysis of the
contour dimension) and the relationship of
the displayed structure with the surrounding tissues.
The factors extracted in this way do
not in themselves provideno individually valid
diagnostic information. Only when we confront them with
statistical sets of evaluations of images of the same
species obtained in patients with a reliably described diagnosis (including genetic character) ,
treatment method, response to therapy and overall outcome, can we
(with some probability) reveal some similarities and similarities
. This can potentially help to refine the diagnosis (by pointing to the possibility shown by this
similarity) and possibly. predict
the response to the appropriate type of therapy and the
further development of the disease.
The necessary comparison
databases can be created so that for each diagnosis
considered in a standardized way performs
multifactorial analysis of images in many patients with the
appropriate (reliably verified) diagnosis and the quantified values of the
extracted feature -factors from these images are gradually stored
in special files, including the evaluation of statistical
variance. In case of event. clinical use, then a particular
patient performs display from which a multifactorial analysis of
extracted features necessary (standardized
same manner as in Comparative databases) and
their values are compared with statistical methods - confront
- with the respective "exemplary" values from the
database. From " probability intersections"values
of several factors can be inferred to a certain sub-type of
pathology and possibly predict its behavior.
These methods are still in the stage of
experimental development, creation of necessary software,
experimental compilation of factor databases from" sample
"images ( templates ).
Mathematical analysis and
computer evaluation of nuclear medicine <<-------click
As a relic of the earlier structure of the monograph "
Nuclear Physics and Physics of Ionizing Radiation ", in
which the chapter "Physical and Technical Problems in
Nuclear Medicine" was included, the computer evaluation of
scintigraphic studies consists of a separate set " Mathematical analysis and computer evaluation
in nuclear medicine ",
which is then followed by a separate detailed monograph " Complex computer evaluation of functional
scintigraphic examinations on a PC - OSTNUCLINE system". Describing specific procedures and algorithms,
mathematical analysis and evaluation of scintigraphic studies of
individual institutions. In this chapter," Radionuclide
scintigraphy "the implementation and
evaluation of clinical scintigraphic examinations described in
§4.9" Clinical
scintigraphic diagnostics in nuclear medicine . "
Furthermore, this topic is related to work
" Filters and filtration in nuclear
medicine ".
4.8.
Radionuclides and radiopharmaceuticals for scintigraphy
In order to be able to diagnostic imaging
something at all with the help of scintigraphy, should be
incorporated into the body g -radioaktive substance - radiotracer or
radiopharmaceutical , whose distribution in
various tissues and organs are then displayed. The indicator or tracking
principle is used : radioisotopes behave chemically
exactly like stable isotopes of the same element, but are
"visible" through their radiation, which allows their
monitoring in the system using ionizing radiation detectors, in
the case of scintigraphy also imaging their
distribution.
Radiopharmaceuticals
are special diagnostic or therapeutic preparations containing radionuclides
, which are a source of radiation. A radioactive atom is
incorporate in their molecules - a radionuclide means a
suitable compound that determines the pharmacokinetics according
to our diagnostic or therapeutic requirements. The
radiopharmaceutical is composed of two main parts :
× Carrier - specific biochemical
substance providing pharmacokinetic targeting or
directing ( " targeting ") to the desired
site, tissue or organ which is to be displayed (or treat). The
carrier is its own indicator of function , which
actively or passively participates in the examined
or therapeutic process in the target structure.
In the simplest case, the carrier is water
(saline) , in which the radionuclide is
dissolved and carried by the bloodstream, or air during pulmonary
ventilation examination. However, most of them are more complex biochemically
active substances - from inorganic salts, through cyclic
hydrocarbons, chelates, dispersed colloidal particles, peptides,
protein carriers, monoclonal antibodies, immunoglobulins,
radiolabeled cells (erythrocytes or
lymphocytes) - which are selectively
taken up in target tissues. The radionuclide carrier is
selected according to the required diagnostic or therapeutic
performance.
× Radionuclide bound to this carrier,
ensuring by its emission of ionizing radiation radiation
"visibility" - signaling or indication
positions of indicator molecules (carriers) in the organism
- in our case display of its distribution. In therapy,
it then causes biological effects on
tissue cells. The binding of the radionuclide should be such as
not to alter the biochemical properties of the carrier.
Radiopharmaceuticals may also contain some stabilizing or
antioxidant excipients.
Radiopharmaceuticals are open
radioactive emitters and, after application to the body,
enter into various metabolic processes depending
on their (bio) chemical structure.. Chemical
composition of the radiopharmaceutical is determined by its
inclusion in the kinetics or to certain metabolic processes,
integrated radionuclide their radiation allows external detection
of either the distribution of the substance (for
scintigraphy) or monitoring the amount in
samples (biological fluids, mostly blood or
urine) . In the case of therapy
, radionuclide radiation performs biological effects
on the cells of the tissue in which the radiopharmaceutical
accumulates (eg, it destroys tumor cells -
§3.6, section " Radioisotope therapy with open emitters ") .
The selected radiopharmaceutical
should ideally accumulate only in the desired target
areas. In practice, however, radiopharmaceuticals are to
a greater or lesser extent also absorbed in other
tissues and organs, or create a continuous tissue
background . This undesirable pharmacokinetics
should be taken into account when evaluating scintigraphic
images, as well as when assessing side effects - radiotoxicity
- in biologically targeted radionuclide therapy (§3.6, section " Radioisotope therapy ") .
Radioindicators in nuclear medicine are
applied in small trace amounts , approx. 10 -9 -10 -12 grams (pico- or nanomolar concentrations in tissues) and therefore cannot affect function on
their own. organs under investigation, nor can they cause any
side or toxic effects on the organism. Therefore, due to the
small and practically immeasurable amount by weight,
radiopharmaceuticals cannot be dosed according to their weight
[mg] - the weight of the "active substance", as is
usual for drugs. Radiopharmaceuticals are dosed by the applied
activity in [MBq].
Note: In this respect, radioindicators
used in nuclear medicine differ significantly from
contrast agents used in X-ray diagnostics. The X-ray
contrast agent is applied in a relatively larger amount(up
to tens of grams) needed to produce a sufficient contrast of
X-ray absorption. There is a relatively high concentration in the
blood and tissues, which due to the chemical composition of
contrast agents (mostly iodine compounds) can significantly
affect the function of the examined organs. Some contrast agents
can have side effects or toxic effects, they can cause allergic
reactions. In contrast, radioindicators used for diagnostics in
nuclear medicine are biochemically safe and
usually have no contraindications.
A certain exception to
this biochemical safety are radiopharmaceuticals based on murine
monoclonal antibodies . They may have allergic
reactions in a small percentage of patientscaused by the
presence of so-called HAMA antibodies ( Human
Anti-Mouse Antibodies ); an adverse immune response
to the product - production of human antibodies against murine
monoclonal antibodies - may then occur . It is therefore
desirable to perform a laboratory biochemical test for HAMA
antibodies before using these radiopharmaceuticals, and its
positivity should be a contraindication to the use of these
products.
Compared to
other pharmaceutical preparations, radiopharmaceuticals have two
other specifics :
« Time-varying
content of a substance carrying a diagnostic or
therapeutic effect - the amount of radioisotope used decreases
exponentially over time due to radioactive transformation
(The rate of this decrease
varies for individual radionuclides, depending on their
half-life) . This is associated with a short
expiration time (which cannot be
extended in any way, regardless of eg storage temperature) .
« Remote
action - emitted ionizing radiation, especially
penetrating gamma, can have biological effects - in this case undesirable
- even outside the tissue where the drug was distributed (or even on another patient in the vicinity) .
For the use of scintigraphy in nuclear
medicine , several g-
radionuclides are
available (mixed bg , pure g , for PET then b + with subsequent emission of annihilation g radiation) in the
chemical form of a number of radiopharmaceuticals
, enabling the study of various functional processes in the
organism. Methods of production and physical (nuclear)
characteristics of individual radionuclides are detailed in §1.6
" Radionuclides
", where their measured spectra are also. Here we will
mention in particular the properties of the most important
radionuclides used in nuclear medicine. For each such
radionuclide, we draw its decay scheme, describe the methods of
its radioactive transformations, types and energy of emitted
radiation. Finally, for each radionuclide, we present the gamma-ray
spectrummeasured by a scintillation spectrometer with a
multichannel analyzer - such a spectrum can then be observed in
practice on a scintillation camera; a more detailed semiconductor
spectrum is also measured .
Radionuclides and radiopharmaceuticals for
single photon gammagraphy - planar and SPECT
Radioiodine 131 I
(+ 125 I + 123 I )
The first radionuclide used in clinical nuclear medicine was radioiodine
131
I (T 1/2
= 8 days, b - with max.
energy 606keV , the main energy g is 364keV), which is of key
importance for the diagnosis and therapy of thyroid
disease (§4.9.1 " Thyrological radioisotope diagnostics ") . It is administered
orally in the form of 131I sodium iodide. For several
years, radioiodine-labeled 131 I-o-hippuran was also used for radionuclide
nephrography and possibly renal scintigraphy, later displaced by 99m Tc-
labeled radiopharmaceuticals (see below).
The radionuclide 131 I is converted (according to the decay
scheme in the figure on the left) by b - radioactivity to excited states of the daughter nuclide
xenon 131
Xe, which is already stable (non-radioactive). The dominant
"channel" of beta-conversion is to an excited level of 364.5
keV (89%), which in 81% deexcites to baseline 131Xe and in 6%
deexcites to a level of 80keV (which then deexcites to baseline).
In 2% there is a decay to the excited level of 722keV, in 7% to
the level of 637keV. One of the excited levels of 131 Xe is the metastable
131m Xe level with an energy of 164keV,
which deeexcitates to the ground state of the 131 Xe nucleus with a
half-life T 1/2
= 12
days . Only 0.38% of 131 I decays occur at this metastable level, and in
addition, its deexcitation is subject to internal conversion, so
that only about 0.021% is emitted as 164keV gamma radiation.
After about 14 days, a radioactive equilibrium is
reached , when the activity of 131 I is equal to the activity of 131
mXe.
Fig .... Decomposition scheme and gamma-spectrum of radioiodine 131 I
The spectrum of gamma radiation 131 I is dominated by
the main photopeak capturing the energy of radiation g 364keV
. Towards higher energies, two weaker peaks, 637 and 723 keV, are
visible. In the region of lower energies we also see weaker peaks
284 and 80 keV, at the very beginning of the spectrum the
characteristic X-rays of K a,
b xenon 30keV (low-energy
lines L a, b 4-5keV on a conventional scintillation detector are not
visible) . The faint 164keV photopeak from
the metastable 131m Xe is not very noticeable because it lies in the
Compton main energy scattering region of 364keV (interferes with
the backscatter peak) - it is analyzed in " 131 I ".
For in vitro radioimmunoassay
(RIA, RSA) is then used radioactive iodine 125
I (T 1/2 = 60 days, EC, 31keV X-27, g 35keV) ....
For scintigraphy is also used radioiodine 1 23
I (T 1/2 = 13.1 hours, EC, g 159keV, X 27 + 31keV), which has more advantageous
physical properties for this purpose than 131 I - more suitable energy g and the absence of b , which leads to a
lower radiation load. Radiopharmaceuticals marked 123I are seldom used
for scintigraphy of kidneys (o-jodhipuran), more often thyroid
gland (NaI), heart (MIBG), as well as for scintigraphy of
receptor systems in the brain - 123 I-ioflupane, 123 I-IBZM (§4.9.8, part " Scintigraphy of receptor systems in the
brain ") . Compared to 99m Tc, 123 I has disadvantages in higher price, difficult
distribution (short T 1/2 ) and slightly higher radiation load.
Fig .... Decomposition scheme and gamma-spectrum of radioiodine 123 I
Technetium
99m Tc
The most important radionuclide for nuclear medicine is
metastable technetium 99m Tc (T 1/2 = 6 hours), which is a pure gamma emitter
(E g =
140keV) and is obtained mostly by beta-decay of molybdenum 99 Mo (T 1/2 = 66 hours) in the so-called generator
(see §1.2. "Radioactivity", part
" Gamma radiation ") . 99m Tc is an almost ideal radionuclide for
scintigraphy, on which basically the entire development
of nuclear medicine in the 1960s and 1990s was based;
has the following advantages :
× 1. A
pure gamma emitter with a short half-life of 6 hours
allows, without the risk of significantly increased radiation
exposure, to apply to patients the very high 99m Tc activity
(in the order of hundreds of MBq) required to obtain quality
images in SPECT or dynamic scintigraphy. Note:
After deexcitation of 99m Tc, 99 Tc is formed in the ground state. It
is also radioactive: b - -transforms into a stable core of 99 Ru (see Fig.4.8.2),
but the half-life is very long here - 2,11.10 5 years. Since the
activity of a preparation containing a given number of N o radioactive nuclei is
A = N o .
l , is
the ratio of the activities of the parent and daughter
radioisotopes in the ratio of their decay constants l , or the inverse
ratio of their half-lives T 1/2 . The relationship between the activity of 99m Tc and the activity
of the formed 99 Tc is thus given by the coefficient » 4.10 -9 . With an applied
activity of the order of 100MBq 99m Tc, the activity of the resulting 99 Tc will be only about 0.4 Bq, which is practically zero
(unmeasurable, well below the level of the natural radioactive
background, eg 40 K). From the point of view of nuclear medicine, we can
therefore consider the resulting 99
Tc to be non-radioactive .
However, the opposite situation is in the field of nuclear
reactors (see §1.3 "Nuclear reactions", passage "
Atomic nuclear fission
"), where 99 Tc, produced in significant quantities as one of the
fission products of uranium, is a difficult component of nuclear
waste with long half-lives, potentially hazardous to the
environment.
×
2. Radiation
g with
an energy of 140 keV can be collimated
very well and effectively decoded
in a thin large-area scintillation crystal of a gamma camera,
which provides images with relatively good resolution and
sensitivity.
× 3.
99mTc is easily obtained from a Mo-Tc generator
(The physico-chemical principle of
radionuclide generators is described in §1.2, part " Gamma
radiation ", passage " Radionuclide
generators " and in §1.4 "
Radionuclides
", part " Production of
artificial radionuclides
", passage " Radionuclide generators ") . These generators are mostly of the elution
type. Molybdenum 99 Mo is absorbed on a support (mostly Al 2 O 3 ) in an " insoluble"
oxide form in a "chromatographic"
column . After the radioactive transformation of the 99 Mo core into a 99mTc daughter core,
the resulting technetium atom is exported from an insoluble bond;
combines with 4 oxygen atoms to form the anion 99m TcO 4 - pertechnetate . This daughter product
is soluble in water, whereby it can be separated
from the starting molybdenum by washing with water - elution
(Fig.4.8.1 left). Since the elution is performed with physiological
saline containing a NaCl salt, the pertechnetate anions are
immediately ionically bound to sodium to form sodium
pertechnetate Na 99m TcO 4 - . In
this chemical form we obtain technetium from the elution
generator.
Fig.4.8.1. Elution 99 Mo - 99m Tc generator.
Left: Principle functional diagram of the
elution generator. In the middle: Technical
design of a sterile generator with an evacuated elution vial.
Right: Decomposition scheme of molybdenum 99 Mo to technetium 99m Tc, deexcitation to
99 Tc and
slow transformations to stable ruthenium 99 Ru.
New types of sterile elution generators use an evacuated
elution vial , into which, after "injection"
under atmospheric vacuum, the saline solution is automatically
sucked through a tube leading from the storage vial through the
sorption column of the generator with 99 Mo (Fig.4.8.1 in the middle). Within about 30 seconds,
the vial is filled with 99m Tc eluate .
In §1.2, part
" Exponential law of radioactive decay ",
passage " Mixtures of radionuclides, decay
series, radioactive equilibrium
" the general equation of subsequent decay of radionuclides A ( l A ) ® B ( l B ) ® C
(stable) was derived. If we apply this equation (multiplied by
the factor lTc to get the
instantaneous activity in [Bq] from the instantaneous number of
nuclei) to our case of the Mo-Tc generator 99 Mo ( l Mo = 0.0105h -1 ) ® 99m Tc ( l Tc = 0.1155h -1 ) ® 99 Tc ("stable") and taking into account that
87% 99 Mo
decays to a metastable excited level of 99m Tc, we get for the time dependence of the immediate
activity of the required technetium 99m Tc relation: A 99m-Tc ( t) = 0.957. A Mo (t = 0). (e -0.0105.t
- e -
0.1155.t ), where A Mo (t = 0) is the activity of 99 Mo at time t = 0 of the previous elution, time t
is in hours. The activity of the 99 Mo with time T varies according to the laws and
decay Mo
(t) = A Mo
(0) .e - l Mo .T = A Mo (0) .e - 0,0105.T . Substituting this basic decomposition of molybdenum
we obtain the resulting relationship for the instantaneous
activity of the eluted 99m Tc at time T from the delivery
of the molybdenum generator and at time t since the last
elution:
A 99m Tc (T, t) = 0,957.A Mo (0) .e - 0,0105.T . ( e - 0.0105.t
- e -
0.1155.t ) ,
where A Mo
(0) is the activity of 99 Mo at time T = 0 of the generator supply, times t
and T are in hours. To determine the actually eluted 99m Tc activity, we
must also take into account the elution efficiency
, which is usually approximately 75-85%. This time dynamics of
activity 99mTc during repeated elutions of the Mo-Tc generator is
plotted in Fig.2.1.B (d), which we present here again for clarity
:
Fig.2.1.B. Time dynamics of radioactivity in a mixture of two
radionuclides.
a) In a mixture of two independent radionuclides
X , Y , each of them is
converted according to its own half-life and the total activity
of the preparation is given by the sum of both exponential
functions.
b) , c) In the decay series of two
generically related radionuclides X -> Y, the decay dynamics
depends on the ratio of the half-lives of the primary parent
radionuclide X and the daughter, further decaying radionuclide Y ; depending on this
relation l X and l Ya transient or secular equilibrium of both radionuclides
can then be established.
d) Specific radioactive dynamics of the
radionuclide molybdenum-technetium generator during repeated
elutions of the daughter 99m Tc, resulting from the conversion of the parent 99 Mo.
After elution, the 99m Tc activity in the generator drops to almost zero, then
rises and reaches a (local) maximum 23 hours after the previous
elution, after which a radioactive equilibrium occurs and the 99m Tc activity
decreases exponentially with a half-life of 67 hours of 99 Mo. After 23 hours
from the last elution, the elution yield of 99m Tc is the highest; The generator can of course be
eluted as needed in a shorter time, but with a lower yield of 99m Tc.
These elution cycles
can be repeated many times until the activity of the parent
radionuclide falls below the applicable limit; for a Mo-Tc
generator with an initial activity of approx. 10-40 GBq it's about 7-15 days. The
relatively long half-life of the parent radionuclide allows for
long-term use of the generator, and the short half-life of the
resulting daughter radionuclide ensures a low radiation exposure
to the patient.
Note: In the
past, generators of the extraction type (by passing the
methyl ethyl ketone through an aqueous solution of 99 Mo, extracting
pertechnetate 99 Tc and separating it from the aqueous phase from the
parent molybdenum) and sublimation type (using the
difference between the volatility of molybdenum oxide and the
resulting technetium oxide). Due to their excessive complexity
and operational complexity, they are no longer used, they have
been pushed out by elution generators.
Detailed decay scheme 99mTc
is in Fig.4.8.2. Default metastable level 142keV
isomerically passes first at the level of 140.5 keV, from where
it emits primary gamma rays of energy 140,5keV
. With a very small proportion of 0.02%, there is a direct
deexcitation to the ground state, in which energy of 142.7 kV is
emitted. Photons of very soft gamma radiation of 2.17 kV are
practically not observed, as they are almost 100% subject to internal
conversion . The core of technetium 99 Tc in the
ground state (after the isomeric transition
from 99m Tc) is beta-radioactive
and with a very long half-life of 200,000 years it slowly
transforms into stable ruthenium 99Ru. In a very small
percentage, there is a direct beta-conversion of 99m
Tc from a metastable level of 142keV to 99 Ru (the ground state of the 99 Tc nucleus is
"bypassed") - mainly to excited
levels of 322 and 90 keV 99 Ru. Their deexcitation
produces g- radiation with energies of 322, 232 and 90 keV, but a
very small representation. At 99m Tc radioactivity,
soft characteristic X-rays with energies of » 2-3keV (L-series)
and » 18-22keV
(K-series) are also emitted , as well as a larger number of
low-energy conversion and Auger electrons
(approx. 4 electrons / 1conversion), mostly with energies » 1.6-3 keV, smaller amounts
» 120-140
keV.
In the standard 99m Tc scintillation spectrum
(on a scintillation spectrometer or gamma camera) we observe only
one significant photopeak of 140keV energy -
Fig. 4.8.3 on the left (on a semiconductor
spectrometer we can also distinguish a weak line of 142.7keV) . Weak peaks from excited levels of 99 Ru (arising from 99m Tc by "bypass" 99
Tc) , especially 322keV, can be seen
spectrometrically only after filtering out a strong radiating
line 140keV with a layer of about 4-5mm lead - Fig.4.8.3 on the
right in the passage " Radionuclide purity ".
Fig.4.8.2. Energy levels, radioactive transformations and beta
and gamma radiation 99m Tc.
Left: Formation of 99m Tc b - transformation of 99 Mo. Right: Detailed decay scheme 99m Tc.
The energies of the individual nuclear levels are calculated in
the left part of the figure from the ground state of 99
Tc, while in the right part they are determined from the ground
state of 99 Ru.
The pertechnetate anions 99mTcO 4- bind
relatively easily to a number of biologically important
substances (after possible previous
reduction of pertechnetate, eg with tin ions) . 99m Tc is able to create chelates with functional groups of
various organic substances and thus provide a wide range of
radioactive preparations differing in their kinetics in the
organism and uptake in individual organs.
Radioactively 99mTc-labeled radiopharmaceuticals
This produces technetium- labeled radiopharmaceuticals
that, after application to the body, are selectively
taken up in certain target tissues or organs, which can
then be detected by external detection of emitting radiation.g displayed by a
scintillation camera. Technetium-labeled radiopharmaceuticals are
widely used in static and dynamic scintigraphy of the kidneys,
liver, lungs, heart, brain and other organs, as well as in tumor
diagnosis. Here are some of the most commonly used 99m
Tc-radiopharmaceuticals :
In some applications, the eluate
alone in the chemical form of sodium pertechnetate Na 99mTcO 4 - will suffice , as technetium ions behave similarly to
iodine ions. It is mainly scintigraphy of the thyroid
gland , examination of Meckel's diverticulum,
radiocardiography. In other applications, 99m Tc atoms chemically bind to complex biochemical
molecules.
For dynamic renal scintigraphy
(§4.9.2., " Dynamic
renal scintigraphy ") is most frequently used 99 m Tc- MAG3 ( m mercapto
and cetyltrimethylammonium g
lycine) for diagnosing tubular function and renal drainage and DTPA
acid ( d diethylene t riamino p
entaoctová) for capture of glomerular filtration. For static
renal scintigraphy, it is then DMSA
(dimercaptosuccinate), which accumulates in the kidney in
proportion to the function of the relevant sites and remains
fixed there in the cortical zone in the cells of the proximal
renal tubules for several hours.
Iminodiacetic acid derivatives - 99m Tc HIDA
(.......) or EHIDA (.........) are used for dynamic
liver scintigraphy
(cholescintigraphy - §4.9.3 " Dynamic liver scintigraphy " ), which they are taken up
from the bloodstream by polygonal liver cells and further pass
through intrahepatic and then excretory bile ducts into the
duodenum.
For dynamic
radiocardiography examinations
(bolus radiocardiography and equilibrium gated ventriculography -
§4.9.4 " Radionuclide
ventriculography ","
Dynamic radiocardiography
")uses 99m Tc- labeled erythrocytes (mostly
labeled in vivo bySn-pyrophosphatepremedication),
which remain in the bloodstream for the duration of the dynamic
study.
99m Tc are usedfor scintigraphy of tissue perfusion
and their viability.- isonitrile complexeswhich,
in the form of lipophilic cations, passively penetrate the cell
membrane, enter cells and bind to cytosolic proteins and viable
cells in the mitochondria. The radiopharmaceutical accumulates,
depending on blood circulation, in healthy viable cells, while in
cells damaged (eg due to ischemia) or even dead and replaced by
scar fibrous tissue, no accumulation occurs. The distribution of
the radioindicator in the individual sites of the examined tissue
is then proportional to the regional blood flow and the viability
of the tissue cells. For scintigraphy myocardial
perfusion is most commonly used 99
m Tc- MIBI
(methoxyisobutyl-isonitrile) and 99 m Tc Tetrofosmin (§4.9.4
" scintigraphy myocardial perfusion") . The isonitrile
radiopharmaceuticals are also used for non-specific cancer
diagnosis - show increased accumulation of viable cells
with a higher energy turnover (pursuing through mitochondria).
Scintigraphy brain perfusion is used 99 m Tc-exametazime HMPAO
( h exa m
ethylpro p ylene and
misses about Xime ) - §4.9.8 " perfusion
scintigraphy brain " .
For scintigraphy, lung
perfusion applied 99 m Tc-labeled macroaggregates MAAserum
albumin, the particles of which are trapped in the capillaries of
the pulmonary circulation, in proportion to the blood supply to
the individual parts of the lungs. For ventilatory
lung scintigraphy, an aerosol of a suitable
radiolabeled inert preparation (usually 99m Tc-DTPA)
is inhaled ; a better
option is to inhale an inert radioactive gas (see "..."
below). Labeled phosphate complexes
(pyrophosphates and polyphosphates) are used for skeletal
scintigraphy , which are osteotropic
and bind to hydroxyapatite crystals; allow you to view bone
reconstruction. The most commonly used is 99m Tc- MDP (methylene diphosphonate) - §4.9.7 "Skeletal scintigraphy " .
Other radionuclides
for g
scintigraphy
Other radionuclides for single photon (planar
and SPECT) scintigraphy can be briefly
named for example.:
Thallium 201 Tl (as chloride) for scintigraphy of myocardial
perfusion , which as analog potassium enters myocyte
cross the cell membrane and accumulates there proportionally
blood flow to the site of the heart muscle.
Galium 67 Ga - citrate for scintigraphy of tumors and
inflammatory foci.
Indium 111 In is also used for a similar purpose ; in form 111In labeled
antibodies for immunoscintigraphy, eg labeled somatostatin analog
111
In-pentetreotide (OCTREOSCAN) for the diagnosis of neuroendocrine
tumors - §3.6, section " Diagnosis of cancer ", section " Molecular gamma imaging
".
For radionuclide cisternography
or perimyelography, intrathecal administration of 169
Yb -DTPA or more
preferably 111 In-DTPA is
used .
Fig ... Decomposition scheme and gamma spectrum of indium 111 In
Gas xenon 133
Xe (complex dynamic scintigraphy) was previously used
for ventilation scintigraphy of the lungs , now
gas radioactive krypton 81m Kr , obtained from the generator 81 Rb (T 1/2 = 4.85 hours) (EC) ® 81m Kr (T 1/2 =
13s). A stream of air, guided through a tube through a container
containing a layer of parent radionuclide 81 Rb, carries away the released daughter 81m Kr, which the
patient inhales, and a scintillation camera uses external
detection of radiation g to show the distribution of this 81m Kr in the pulmonary alveoli - static
ventilation scintigraphy of the lungs . For more complex
pulmonary diagnostics, it is appropriate to combine perfusion and
ventilatory scintigraphy.
The principle of the 81 Rb / 81m Kr generator is in the left part of
Fig. .... The parent rubidium 81 Rb is fixed in the solid phase in a small column,
through which a stream of elution air is passed
by means of a fan (air pump with adjustable
power) . By radioactive decay of rubidium-81, the continuously
released daughter gas krypton 81m Kr is entrained by the passing air and led to the respiratory
mask , from which the patient inhales a mixture of air
and radioactive 81m Kr.Kr. One -way valves are included in the
circumference of the breathing mask , and a mixing valve for
outside air is also connected to ensure free breathing. Exhaled
air is led to the extinction vessel (volume approx. 30 liters) , from
which, due to the very short half-life of 81
m Kr, practically non-radioactive air
emerges.
During this examination of pulmonary
ventilation , inhaled air with a trace content of
radioactive 81 m Kr enters the pulmonary alveoli , while the
emitted radiation of 191keV gamma is scanned by a gamma
camera . The scintigraphic image of the site of reduced
activity shows areas of the lung with impaired
ventilation , where krypton-81m, and thus no air, does not get
(either at all or reduced) - see §4.9.5 " Lung scintigraphy
(nuclear pneumology) ".
Fig .... Generator 81 Rb / 81m Kr.
Left: Principle of generator operation. Middle:
One of the design arrangements of the Rb-Kr generator. Right:
Disintegration scheme 81 Rb and 81m Kr; in
the black field is the scintillation spectrum of gamma radiation 81m
Kr.
Radionuclides and radiopharmaceuticals for
PET
Of the more than 100 positron radionuclides, most are not
suitable for PET imaging - due to too short or long half-lives,
inappropriate radiochemical properties, low positron content and
high intensity of unwanted electron and hard gamma radiation.
Only a few b+
-radionuclides are available for the medical use of positron
emission tomography, but in the chemical form of a
number of radiopharmaceuticals, enabling the study of various
metabolic processes in the organism. For diagnostic imaging by
PET, coincidence detection of g-
photon pairs formed in the tissue during
annihilation of positrons from a radioindicator with electrons is
used (described in detail above "Positron
emission tomography "). The following positron radionuclides are mainly used :
Fluoro-deoxy-glucose 18 FDG
By far the most commonly used radiopharmaceutical for PET is
2-deoxy-2- 18 F-D-glucose, abbreviated as fluoro-deoxy-glucose ( 18 FDG). FDG metabolism
is somewhat different from normal glucose metabolism. Like
ordinary glucose, FDG has an affinity for cells with increased
metabolism (increased need for sugar - glucose), where
it gets through the appropriate transport proteins and is
subsequently phosphorylated. However, unlike true glucose, FDG is
no longer metabolized and therefore accumulates
in the cell . As a result, there is a markedly increased
accumulation of FDG in the tumor cells, so that the
tumor foci appear with high contrast to tissue
and blood background. Oncological diagnostics
therefore makes up more than 90% of all PET examinations (§3.6, section " Diagnosis
of cancer ") . 18 FDG is also used to examine the myocardium
, where myocardial viability can be assessed based on FDG
consumption.
Glucose
it gets from the extracellular space into the cells by passive
transport through transmembrane proteins - glucose transporters.
Upon entering the cell, glucose is phosphorylated by gluokinase
to glucose-6-phosphate (analogous to FDG-6-phosphate). Normal
glucose can then be converted to glycogen or metabolized to water
and carbon dioxide. However, this metabolism does not occur in
FDG, so FDG is "trapped" and tends to accumulate
in cells.. For FDG, the only way to be excreted back from the
cell is through glucose-6-phosphatase. In cells containing low
glucose-6-phosphatase, the concentration of FDG is proportional
to glucose consumption. In contrast, in tissues that contain a
lot of glucose-6-phosphatase, the accumulation of FDG is lower
than that corresponding to glucose metabolism. Dephosphorylation
by glucose-6-phosphatase is generally very slow, so that the
concentration of FDG-6-phosphate in the tissues is kept stable
for several hours.
After radioactive decay, 18 F produces
non-radioactive oxygen 18 O and FDG-6-phosphate produces ODG-6-phosphate, which
then undergoes cellular glycolysis as normal glucose.
Unmetabolized 18 FDG is removed by glomerular filtration in the kidneys
and is excreted in the urinary tract (with
an excretion half-life of about 2 hours)..
18 FDG was first used in non-tumor diagnostics to
visualize local glucose metabolism in the brain and myocardial
glucose metabolism. However, it now has a major application in
oncology as a radioindicator for imaging the increased metabolic
activity of tumor tissues. Conversely, it is not suitable for the
diagnosis of brain tumors. And it is also not suitable for the
diagnosis of prostate cancer - due to the relatively slow
metabolism of these tumor cells and the proximity of the bladder
with a significantly higher content of FDG.
18F
Sodium -fluoride (NaF)
is used for PET scintigraphy of the skeleton. PET shows areas in
the bones with osteoblastic and osteoclastic changes that may be
related to tumor remodeling, but also to benign skeletal changes.
The advantage of Na 18F is high (up to 50% of the
applied activity) and rapid absorption in
the bones, together with the rapid degradation of unbound
radioindicator in the blood. This leads to the acquisition of
contrast images in a short time (less than 1 hour after iv
application).
NaF
is an ionic compound of Na + and F - ions. After iv application, NaF is delivered to the
bones and fluoride ions diffuse through the blood capillaries
into the extracellular fluid. 18 F-ions are exchanged for OH-ions of hydroxyapatite
Ca 10 (PO 4 ) 6 (OH) 2 to form floroapatite.
Subsequently, the installation 18 F-ions into the crystalline
structure of hydroxyapatite in bone. Increased uptake of 18 F-fluoride occurs in
malignant bone lesions due to increased blood supply, increased
permeability of capillary walls, and faster bone remodeling. The
advantage of Na 18 F is that virtually all 18 F-fluoride that is
transported to the bones by the blood is trapped in them and,
conversely, binding to serum proteins is minimal. This leads to
the rapid degradation of the unbound preparation from the
circulation and the acquisition of "pure" contrasting
skeletal images in a short time.
Special
radiopharmaceuticals for "molecular imaging" *)
With the development of organic chemistry,
biochemistry and cell biology, some radiopharmaceuticals have
been developed whose labeled molecules have affinity for very
specific cell types or processes at the subcellular level. With
the help of scintigraphy and a suitable radiopharmaceutical, it
is possible to purposefully examine not only the function of a
certain organ or tissue, but also to selectively affect a certain
type of metabolic or transport pathway, such as enzyme or
receptor binding or antigen-antibody reactions. For this purpose,
special radiopharmaceuticals (both for diagnostics and for
therapy) have been developed and are still being developed, which
are characterized by their effects at the molecular level
. With a bit of exaggeration, these methods of local measurement
and imaging of the physiological response are referred to as
" in vivo biochemistry ".
*) Name "molecular imaging
"does not, of course, mean that we perhaps visualize the
molecules themselves (unfortunately we can't do that ...), but we
visualize the distribution of radioindicators that is the result
and reflection of specific biochemical reactions
at the molecular level.
Imaging of viable
and proliferating tumor cells and tissues - §3.6, section " Diagnosis of cancer " In addition to the above-mentioned and most
commonly used fluoro-deoxy-glucose
18 FDG,
there are some other tumor radiopharmaceuticals :
18 FLT ( 18F-3-fluoro-3-deoxy-thymidine) is a radiolabeled form of
a pirimidine nucleoside. It accumulates significantly in proliferating
cells - it shows the activity of the enzyme thymidine
kinase , which characterizes the intensity of cell division.
Because a substantial increase in the rate of mitosis and cell
proliferation is a hallmark of malignant tumor tissue, 18FLT functions as a tumor-specific
PET radioindicator . It usually provides more
contrasting images of proliferating tumor lesions than 18 FDGs. It is
particularly suitable for monitoring the response of
malignant tumors to therapy . Chemotherapy and
radiotherapy often cause an inflammatory reaction in the tumor
and around the tissue, which significantly increases the
accumulation of 18FDG, making it very difficult to assess the regression
or progression of the treated tumor; Therefore, 18 FDG is not a
completely ideal radioindicator for the response of malignant
tumors to treatment, 18 FLT
is more suitable .
Thymidine
is essential for replication in dividing cells. After passage of
thymidine through the cell membrane, it is phosphorylated, which
is catalyzed by the cytosolic isoenzyme thymidine kinase-1
(TK1), and subsequently incorporated into DNA (during the DNA
phase of cell cycle synthesis). However, incorporating fluorine
to the 3 'position in thymidine prevents FLT from further
incorporating it into DNA. FLT monophosphate is not incorporated
into DNA and the cell membrane is impermeable to it - it is
therefore metabolically "trapped" inside the cells. PET
display 18FLT
detects the enzymatic activity of TK1, tracing the recovery of
nucleosides from degraded DNA. The uptake and accumulation of 18 FLT thus corresponds
to the rate of cell proliferation. 18 FLT serves as an indicator of changes in tumor
cell growth . 11 C-thymidine
was used for a similar purpose .
18
F-fluorocholine ( 18 FCH) is a fluorine-18-labeled analogue of choline
, the basic building block of cell phospholipid membranes
. It is used to visualize phospholipid metabolism in tumors. It
appears to be increased uptake in tumors of the brain, prostate,
breast, lung, esophagus.
Choline is
an important component of phospholipids in cell membranes.
Choline is phosphorylated to phosphorylcholine by choline kinase
inside cells and, after several other biosynthetic processes, is
eventually incorporated into phospholipids. 18 Flourocholin behaves in the same way. Cells with high
metabolism also have increased choline uptake due to higher
requirements for phospholipid synthesis in their cell membranes. 11 C-choline is
sometimes used for positron emission tomography, but its
disadvantage is the short half-life of 11 C (20 min.).
18
F-fluciclovin is used for scintigraphic PET diagnosis of
prostate tumors, especially in recurrent disease.
Fluciclovin
is an analog of the amino acid L-leucine, referred to herein as 18 F. It accumulates in
the tumor via amino acid transporters. An increase in
transmembrane amino acid transport occurs in the prostate due to
increased metabolism of amino acids for energy and protein
synthesis. Unlike the natural amino acids fluciclovin, it is not
metabolized and accumulates in tumor cells - a positive
PET image of the tumor.
18
FET ( 18 F-O- (2-fluoroethyl) -L-tyrosine) is an analog of the
amino acid tyrosine to which an 18 F- labeled ethyl group is attached via an oxygen atom .
It shows the accumulation of amino acids in cells. It is suitable
for imaging brain tumors glyoms , their extent, to
detect recurrence after therapy and its differentiation from
necrosis.
Tyrosine is one of
the building blocks of protein. Increased uptake into cells is
due to the higher content of L-type amino acid transporters.
However, the fluorinated FET analogue, unlike tyrosine, does not
enter protein metabolism. Therefore, 18 FET accumulates in cells and maps the increased amino
acid consumption due to increased protein metabolism in tumor
cells. 18
FET (unlike 18 FDG or 11 C-MET) is not absorbed in
macrophages and allows tumor tissue to be distinguished from
inflammatory tissue.
Other labeled amino acids are
sometimes used to diagnose brain tumors:11 C-methyl-L-methionine ( 11 C-MET), 3- 123 I-iodo- and methyl-L-thyrosine ( 123 I-IMT), two other L-tyrosine analogues L- 11 C- tyrosine and 2- 18 F-fluoro-L-tyrosine.
All these substances have similar properties.
18
F-FMISO ( 18 F-fluoromisonidazole) and 18 F-FETNIM ( 18F-fluoroerythronitroimidazole) are radioindicators
showing cellular hypoxia , which is important for tumor
angiogenesis and for planning radiotherapy (radiosensitivity,
oxygen effect - see §3.6, section "Physical and radiobiological factors of
radiotherapy ").
Nitromidazoles
passively diffuse through the cell membrane into
the cytoplasm, where they are reduced by intracellular
nitroreductases. The resulting nitro-radical R-NO can be further
reduced to R-NH 2 , which reacts with macromolecules as a strong
alkylating agent DNA, RNA, with proteins However, in cells with
sufficient oxygen, the reduced nitromidazole is rapidly (re)
oxidized and removed outside, further reactions no longer proceed
and the reaction products do not accumulate there. to form stable
covalent bonds with biomolecules whose speed is inversely
proportional to oxygen concentration intrecelulární. 18
F-labeled nitromidazole derivatives are
thus "trapped" in hypoxic tissue cells. This
accumulation of 18 FMISO occurs only in cells with active nitroreductases,
so that the accumulation of the radiolabel occurs only in living
hypoxic cells, not in necrotic ones.
Another PET radiopharmaceutical for
hypoxia imaging is 64 Cu-ATSM
(acetyl-methyl-thiosemicarbazone).
18F-Florbetaben,
18F-Flutemetamol, 18F-Florbetapir are 18F-labeled polyethylene glycol stilbene derivatives
with high specific affinity for beta-amyloid plaques
. It is used for scintigraphic visualization of amyloidosis
especially in Alzheimer's disease in the brain, it can also be
used in myocardial amyloidosis.
Immunoscintigraphy
An important methodology for "molecular imaging" is immunoscintigraphy
, based on the highly specific nature of antigen-antibody
immunological reactions. The antibody, labeled with the
appropriate radionuclide (99mTc, 18F, 68Ga, 111In, 131I, 123I), selectively binds to the appropriate tumor
marker (antigen) after administration, after which we
can locate the relevant tumor by external detection of gamma
radiation using a gamma camera. The required antibodies are
either of human origin or are obtained from the serum of
immunized animals by the method of so-called lymphocyte
hybridization, which allows the preparation of a homogeneous
so-called monoclonal antibody - §3.6, passage
" Monoclonal
antibodies ". They are
used mainly in tumor diagnosis , but also, for
example, in the diagnosis of inflammatory foci
using antigranulocyte monoclonal antibodies such as 99mTc-sulesomab and 99mTc-besilesomab (§4.9.6 " Scintigraphy
of inflammatory foci ") .
Neuro-endocrine tumors
For gamma imaging of neuro-endocrine tumors
(including pancreatic), radioindicators that bind to somatostatin
receptors are suitable . For example, 68Ga-DOTATOC
is used, 11C-5-hydroxy-tryptophan
(5-HTP) is also tested . This diagnosis can also be used to
assess (predict) subsequent radionuclide therapy using 90Y- or 177Lu-DOTA
radiopharmaceuticals - teranostics (discussed in §4.9, section " Combination
of diagnostics and therapy - teranostics ") .
Recently,
bombesin-based gastrin-releasing peptide (GRP) receptors, labeled
with e.g.18F.
Bombesin *) is a peptide composed of 14 amino
acids. It is formed in the small intestine and antrum, has a
stimulating effect on the pancreatic and gastric mucosa. It is a
potent antagonist of the neurotransmitter gastrin releasing
peptide (GRP). It has been shown that it can be expressed by
several cancer cell lines, where it can endocrine stimulate the
growth of tumor cells via bombesin receptors on the membranes of
these cells. Thus, bombesin may be a tumor marker for
prostate, lung, gastric, neuroblastoma and others.
*) The somewhat bizarre name " bombesin
" originated in the discovery of this substance, which was
first isolated from the skin of a bellied frog (Latin Bombina
bombina ).
......... add ............
Prostatic tumors
Prostate- specific membrane antigen PSMA
appears to be very promising for imaging and therapy of prostate
cancer . Radiolabeled small molecules of PSMA
inhibitors bind with high affinity to prostate cancer
cells (which highly express PSMA) , allowing scintigraphic imaging of these lesions as
well as their radionuclide therapy - depending on the
radionuclide used. For scintigraphic images
planar and SPECT can be used 99mTc-MIP-1404, for PET imaging of 18F-DCFBC or 68Ga-HBED-PSMA, recently 68Ga-PSMA-11. However, 18F-PSMA
(18F-PSMA-1007)
is best suited for PET scintigraphy of prostate tumors.
131I-MIP-1095 has been tried for targeted radionuclide therapy
of the prostate, but 177Lu-J591 and more recently 177Lu- or 225Ac-PSMA-617 have proven to be the
best so far .
In addition to diagnostics, the
possibilities of therapeutic use of monoclonal antibodies
as carriers of suitable radionuclides beta or alpha with
radiotherapeutic effect, or suitable chemotherapeutic agents. The
use of monoclonal antibodies (non-radioisotope) in chemotherapy
(biological treatment) of cancer is discussed in §3.6, section
" Therapy of cancer ", in radionuclide therapy in the section " Radioimmunotherapy ".
Labeled cytostatics
Methods for radioactive labeling of various
types of cytostatics have been developed for monitoring
and prediction of chemotherapy of cancer (cytostatics
are discussed in more detail in §3.6, section " Therapy of cancer ") . The diagnostic
application of such radiopharmaceuticals makes it possible to
show where these cytostatics are taken up and to what extent they
penetrate into tumor foci. In this way, the relevant cytostatics
will be taken up in the chemotherapy itself - according to which
the effectiveness of the treatment ("theranostics")
can be inferred .
One such
radiolabeled cytostatic is 18F-paclitaxel (currently being tested in
preclinical studies). 99mTc-MIBI , which has similar cell uptake
kinetics to doxorubicin and cisplatin , can
also be used to pre-map the distribution of some cytostatics .
Apoptotic
radiopharmaceuticals
A new interesting group are
radiopharmaceuticals for imaging cellular apoptosis
. These are organic molecules (either proteins or relatively
small molecules) labeled with a suitable radionuclide (eg 99m Tc, 18 F), which have an
affinity for cells that are in the early stage of apoptosis (programmed cell death - see §5.2 " Biological effects ionizing radiation ", passage "Mechanisms of cell death "). Proteins bind to their
surface (to phospholipids exposed on the surface of apoptotic
cells), small molecules penetrate the cell membrane and
accumulate in the cytoplasm. The result is selective accumulation
of radioindicator in apoptotic cells and tissues. By gammagraphic
imaging of the distribution of these radioindicators, we obtain
positive images of those places where apoptosis occurs most
intensively - whether due to irradiation, cytotoxic agents or
ischemia.
We can monitor the molecular imaging
of the distribution of cell apoptosis very early response
cells and tissues for therapy (radiotherapy or
chemotherapy), already at the beginning and during therapy - see
§3.6, section " Diagnosis of cancer ". They can also be used to image ischemic
foci in heart or cerebral infarction. Two types of such
radiopharmaceuticals have been successfully tested in clinical
practice for imaging apoptosis (the third
is in the laboratory stage) :
- 99mTc-annexin V - is a protein that binds
to phospholipids detected on the surface of cells undergoing
apoptosis. Annexin V is obtained from the placenta and technetium
are labeled via hydrazino nicotinamide: 99
m Tc-6-hydrazinonicotin (HYNIC)
-annexin V . For similar properties is
tested 99mTc-Duramycin ;
- 18F-ML-10
[2- (5-Fluoro pentyl) -2-methyl malonic acid] - penetrates the
depolarized cell membrane and accumulates in the cytoplasm of
apoptotic cells;
- Peptide
18F-CP18
[pentapeptide containing triazole] - maps Caspase-3 activity,
accumulates in apoptotic cells.
Radionuclides
for therapy in nuclear medicine (methodological note)
Open radionuclide therapy
is also organizationally integrated into the field of nuclear
medicine (the main method of which is the scintigraphy
discussed here) . However, from the point of view of our physical
point of view, as well as from the point of view of the mechanism
of action and purpose of use, we have included this radioinuclide
therapy in §3.6 " Radiotherapy ",
part " Radioisotope therapy with open
emitters ". Thus, we also
find radionuclides used in nuclear medicine for therapeutic
purposes.
Preparation of radiopharmaceuticals
As mentioned above, radiopharmaceuticals are composed of two
basic constituents: a radionuclide emitting
ionizing radiation and a carrier to which it is
bound and which brings it to the required target in the body - to
certain cells, target tissues and organs. Three basic
radiochemical methods are used to prepare radiopharmaceuticals - radioisotope
labeling :
× Isotope exchange reaction ,
wherein a certain stable isotope in the carrier compound is
chemically replaced (exchanged, "displaced") by its
radioactive isotope added to the reaction mixture. The resulting
labeled substance has the same chemical and biological properties
as the starting substance, because its molecules are chemically
identical to the original molecules.
× Chemical synthesis ,
in which radioactive atoms are chemically incorporated into the
appropriate site in the carrier molecule, most often by means of
a coordination covalent bond, so that the resulting complex
compound has the desired properties. For this labeling,
so-called chelates (such as EDTA,
DTPA) are often used , which bind one part
to the carrier molecule and the other part binds the radionuclide
atom.
× Biochemical synthesis
uses enzymes and microorganisms. The radionuclide, added to the
culture medium, enters their metabolic processes in living
microorganisms and then incorporates them into their respective
metabolites.
In terms of organization of
preparation, we can divide radiopharmaceuticals into two groups :
¨ Completed
radiopharmaceuticals - mass production,
manufactured and radiolabeled in the manufacturer's radiochemical
laboratory, delivered to nuclear medicine (appropriate activity
and volume) and ready for direct application to patients. Thus,
radiopharmaceuticals labeled with radionuclides with a longer
half-life (> approx. 2 days) are supplied. These are, for
example, radiopharmaceuticals labeled with iodine 131 I and 123 I, then 111 In, 201 Tl, 67 Ga, 169 Yb, recently
short-term 18 F and others.
¨ Radiopharmaceuticals
prepared at the workplace - individual preparation
The required biochemical substance - carrier - is marked
with the necessary radionuclide at the nuclear medicine
workplace, in the laboratory of radiopharmaceuticals (preparation
" magistraliter "). Thus, mainly
radiopharmaceuticals labeled with short-lived radionuclides are
prepared, mainly technetium 99m Tc from a generator, sometimes even 18 F. The actual
synthesis was previously performed using basic chemicals, now the
so-called kits are used ( kit
= set of tools, building parts) - a
compact set of non-radioactive ingredients supplied by the
pharmaceutical manufacturer, to which only the solution of the
radionuclide itself is added and the corresponding labeling
reaction already takes place automatically.
In nuclear medicine workplaces, the radiopharmaceutical
laboratory deals with the preparation and filling of
radiopharmaceuticals for their application . It is usually
performed in special boxes or fume hoods
equipped with air conditioning, ensuring laminar air
flow. Current requirements for air cleanliness (often exaggerated! - see below) *)
lead to very complex and expensive air conditioning systems. New
alternative solution, providing (without
hood and air conditioning) sterility of
radiopharmaceuticals and radiation protection of workers, are compact
automatic devices for computer-controlled
filling of radiopharmaceuticals, sometimes supplemented
by the possibility of automatic application of the prepared
solution of the radiopharmaceutical to the patient. They are
mainly used in PET for 18 FDG.
Preparation and filling of radiopharmaceuticals at the workplace
of nuclear medicine.
Left: Laminar hood for elution of Mo / Tc
generator, preparation and filling of radiopharmaceuticals at the
Department of Nuclear Medicine, University Hospital Ostrava.
Right: Compact device for automatic computer-controlled
filling of radiopharmaceuticals.
*) Author's note
- exaggerated
requirements for the preparation of radiopharmaceuticals in
nuclear medicine workplaces
From the point of view of my long-term work in the field of
nuclear medicine, I would like to make a small critical
comment on current standards and regulations for the
preparation and filling of radiopharmaceuticals in nuclear
medicine workplaces. Until the 1990s, the actual laboratory
radiochemical preparation of radiopharmaceuticals was
carried out at workplaces , in which a given radionuclide was
labeled with compounds prepared in the laboratory of
radiopharmaceuticals using chemical methods (in
test tubes, beakers, penicillins) . This
was done in conventional chemical fume hoods with or. lead
shielding (sometimes with air extraction,
other times not ...), located in standard
laboratory rooms. Following the principles of good
laboratory practice, there have never been any
problems with the sterility of the resulting
radiopharmaceuticals (" nothing
has ever happened to any patient ") . Since the 1990s, kits (significantly
facilitating preparation) and ready-to-deliver
radiopharmaceuticals have been used more and more , which
come sterile and are only filled into workplaces
for use in patients.
And at this time, paradoxically, the
requirements for the sterility of the environment
began to appear and continue to tightenIn which they have already
performed significantly easier handling ..! .. This leads to
enormous investment and operating costs, which in my opinion
virtually useless (throwing
hundreds of thousands even millions!) ...
The core of misunderstanding here is the confusion
minor " magistraliter "on - the - job training (or even
just filling), to which officials try to mechanically
transfer the demanding requirements, standards and
regulations from the mass production of medicines
in pharmaceutical plants, where these strict standards are, of
course, justified. The doctor in the internal surgery also does
not work in an aseptic environment of class "A" when
withdrawing a sterile injection preparation from an ampoule or
penicillin into a syringe for iv administration.(It would be nonsense) ..! ..
Furthermore, narrow-minded standards and
regulations regarding " registration
with us" are a serious limiting factor in the introduction
of new promising radiopharmaceuticals , already proven
and registered in the world. This leads to the lagging
behind of the field of nuclear medicine in our country,
at the expense of more advanced diagnostics and therapy of
patients.
After all, we
encounter similar bureaucratic approaches in the field of radiation
protection (cf. §5.8, concluding note " Bureaucratic requirements of radiation
protection ").
Methods of
administration of radiopharmaceuticals
In terms of application form, three types of radiopharmaceuticals
are used :
- Parenteral radiopharmaceuticals
administered most often intravenously ,
sometimes subcutaneously or intralubally . They
are mostly aqueous solutions, dispersions, colloids, suspensions.
There are high demands on sterile and pyrogenicity.
- Oral radiopharmaceuticals can be in the
form of solutions or solids. The most common are solutions or
capsules of radioiodine given during thyroid therapy, or liquid
or solid bites swallowed during examination of the esophagus and
evacuation of the stomach.
- Inhaled radiopharmaceuticals are
primarily radioactive gases (such as krypton 81m Kr) or gaseous dispersions
of labeled radiopharmaceuticals (eg 99m Tc DTPA) produced
in nebulizers, inhaled together with air during
examination of pulmonary ventilation .
Quality and purity of
radiopharmaceuticals
The properties of the
radiopharmaceutical used primarily affect scintigraphic
diagnostics; unsuitable, poor quality or contaminated
radioindicator can lead to inaccurate or erroneous diagnosis,
ineffective radionuclide therapy, or it can also have side
effects for the patient. From our physical and methodological
point of view, the purity of the
radiopharmaceutical is an important property , which can be
divided into two categories :
v Radionuclide
purity
Nuclear reactions (see §1.3
" Nuclear reactions and nuclear energy ", part " Types of nuclear reactions
" and §1.4, part " Production artificial radionuclides "), which produce their own
radionuclides used for labeling radiopharmaceuticals, usually
take place in various ways and, in addition to
the desired radionuclide, can lead to the formation of other
radionuclides (the same element or another element). The
amount of these radionuclide impurities depends on the target
used, the type and energy of the irradiating particles and
subsequently also on the method of separation
and isolation of the given radionuclide.
There are three basic sources of
radionuclide impurities (+ one special for radionuclide
generators) :
1. Target material
can never be prepared in 100% "mononuclide" purity of
the desired target nuclide. Traces of other isotopes of a given
element, or even other elements, are always present. Nuclear
reactions can then produce radionuclides other than the desired
ones in the target from these impurities.
2. Different nuclear
reactions - even with
the same nuclide composition of the target, nuclear reactions can
take place through different "channels" with different
probabilities. When irradiated with neutrons, these are most
often reactions (n, g ), but they can also occur (n, p) or (n, d), etc., when
proton irradiated, then reactions (p, g ), (p, n), ( p, d) and the
like; it essentially depends on energy. Even in a completely pure
target, a mixture of different radionuclides can be formed.
3. Radiochemical
separation The mixture
of radionuclides formed by nuclear reactions during irradiation
is a technologically difficult process which may not succeed with
100% efficiency. Trace amounts of other radionuclides - radionuclide
impurities - may thus be present in the final product .
4. For generator
radionuclides , the
radionuclide impurity can enter the desired daughter radionuclide
in two ways:
¨ From
the radionuclide impurities contained in the parent radionuclide.
¨ Traces
of the parent radionuclide may also be released into the daughter
radionuclide. E.g. in the Mo-Tc generator a small amount of
parent 99
Mo can penetrate into the 99m Tc eluate , or in the Ge / Ga generator the daughter
eluate68
Ga can release even a small amount of parent 68 Ge.
Radionuclide purity
is the share of radioactivity of the required (declared)
radionuclide in the total activity of the preparation. Usually,
however, the opposite value is given - the content of
radionuclide impurities - contaminants; it is usually
expressed as a percentage. The permissible content of
radionuclide impurities is specified for each radioindicator in
the relevant standard for its preparation (eg
for the 99m
Tc eluate , radionuclide impurities must not exceed 0.1%) .
Measurement of radionuclide impurities
Accurate determination of the content of radionuclide impurities
is performed by spectrometric measurement of
radiation g using a scintillation NaI (Tl) or
semiconductor Ge (Li) detector connected to a multichannel
analyzer. It is not easy to measure the very low (trace)
radioactivity of a contaminant in the background by many orders
of magnitude higher activity of the basic radionuclide - the weak
radiation of the contaminant is completely "irradiated"
by the radiation of the basic radionuclide. We have a chance to
measure the radionuclide purity in basically two situations :
1. A
high-energy contaminant
that emits gamma radiation with a significantly higher energy
than the basic radionuclide. In this case, the screening
method with a shielding absorbent pad can
advantageously be used for the separate detection of the
contaminant
: Place the vial with the examined preparation in a lead shield
of suitable thickness (approx. 2-5 mm), which almost completely
absorbs the intense low-energy radiation g of the basic radionuclide, but
transmits a considerable part of the weak but high-energy g -contaminant
radiation.
2. Long-term
contaminant
with a half-life several times longer than that of the basic
radionuclide. If we measure such a preparation with an interval
of 10 or more half-lives of the basic radionuclide, we obtain the
activity or spectrum of the contaminant, which is no longer
irradiated by strong radiation of the basic radionuclide. This
method has the disadvantage that it is an " ex post
" measurement long after the preparation and use of the
preparation. In some cases, however, there is no other option.
A typical example of high energy contamination is the 99m Tc eluate (E g = 140keV), which
may be contaminated with maternal 99 Mo with a strong E g
= 740keV line. To shield strong 140keV
radiation, we use a small lead container with a wall thickness of
» 4-5mm,
the transmitted 740keV radiation can already be measured with a
scintillation detector, without the risk of being flooded by
powerful primary radiation 99m Tc, which is absorbed by Pb-shielding. For quantitative
determination, it is of course necessary to have the detector
pre-calibrated with a 99 Mo standard for a given shielding and geometric
configuration . The contaminant activity measured in this way is
then divided by the total activity of 99mTc and we obtain a radionuclide impurity content of 99 Mo.
In Fig. 4.8.3 in the left part there is a
standard scintillation spectrum of 99m Tc (a bottle with an activity of approx. 10kBq attached
directly to the scintillation detector). Only a distinctive
140keV photopeak is displayed. In the right part of the picture
there is a bottle with 1GBq of 99m Tc eluate placed in a lead container with a wall
thickness of 5mm. In addition to the residual (lead highly
attenuated) peak 140keV on the spectrum, we see a peak of
characteristic X-rays of lead around 80keV and of higher energies
"popped" the two weaker peaks:
- 322keV comes from deexcitation of
excited levels of ruthenium of 99 Ru, to which a slight proportion of the beta
-radioactivity breaks down the metastable level of 142keV99m Tc (see conversion
diagram in the previous Fig.4.8.2 on the right).
- 740keV comes from trace contamination
of the eluate by mother molybdenum 99 Mo (cf. previous Fig.4.8.2 on the
left). It is from the intensity of this peak that the radionuclide
purity of the 99m Tc eluate is determined spectrometrically .
The same measurements on a semiconductor detector. are at the
bottom of the image.
Fig.4.8.3. Spectrometric measurement of the radionuclide purity
of the 99m
Tceluate ( top - scintillation spectrum, bottom
- semiconductor spectrum).
Left: Basic gamma radiation spectrum 99m Tc. Right:
The spectrum of gamma radiation measured through the shielding
layer of a 5mm lead container.
However, for a simplified measurement of the
radionuclide purity of the 99m Tc eluate , a conventional activity meter
with an ionization chamber, equipped with a suitable shielding
insert (and calibration) , is used in most workplaces .
However, due to the low sensitivity of these meters, we can
determine up to radionuclide impurities of 99 Mo exceeding hundreds of kBq (for Tc-eluates with an
activity of tens of GBq, however, the sensitivity is sufficient
to check compliance with the standard). It is sufficient to
measure the radionuclide purity of the 99m Tc eluate only for the 1st elution
from the given generator, where the risk of potential
contamination by 99 Mo is the highest; if the result is satisfactory, but
it will almost certainly apply even more to further elutions.
v Radiochemical
purity
Even chemical reactions by which
radiopharmaceuticals are prepared by labeling with the
required radionuclide do not proceed in 100% yield. Therefore, in
the resulting preparations, in addition to the active
radioactive substance itself, there is always a small
amount of unbound activity and possibly other
compounds of radioactive substances that do not carry a
diagnostic or therapeutic effect and may interfere
or cause undesired radiation exposure in non-target tissues. Radiochemical
purity is the share of the declared chemical compound of
a given radionuclide in the total activity of the preparation.
Chromatographic methods (mostly on paper or on a thin layer) are most often used to control the radiochemical purity
of prepared radiopharmaceuticals - §2.7, section "RadioChromatography ", sometimes also electrophoresis
- see the same section " RadioElectrophoresis ".
Only a certain part of the molecules
of a biological substance is marked by a radionuclide - we are
talking about a radioactive substance with a carrier
; the so-called carrier-free radioactive substance,
where the radionuclide is contained in all molecules of the
substance, is difficult to prepare and is used only for special
purposes.
Application of
radiopharmaceuticals
For their own use in nuclear medicine (diagnostic or
therapeutic), radiopharmaceuticals are administered
to the body, most often intravenously, or orally or by inhalation
(as mentioned above). For each type of scintigraphic examination,
a certain optimal amount of a given
radiopharmaceutical is determined , expressed in units of activity
*) [MBq] of the radionuclide bound in the radiopharmaceutical -
with normalization to body proportions, usually the patient's
weight.
*) Radiopharmaceuticals cannot be dosed
according to their weight [mg] - the weight of the "active
substance", as is usual for drugs. The weights of applied
radiopharmaceuticals are immeasurably small (often even on the
verge of chemical provability). This is due to the high specific
activity of the radionuclides used with a short half-life. The
only way to dose radiopharmaceuticals is through the applied
activity in [MBq]. The activity of the
radiopharmaceutical for application is measured in a
metrologically calibrated activity meter (§2.3, section " Well ionization activity meters ") . Guideline values ??of
the recommended applied activity for various types of
radiopharmaceuticals are given in the table in §5.7 " Radiation exposure during radiation diagnosis and
therapy ", passage "Radiation
dose to patients from radionuclide examinations .
"
Other features of
radiopharmaceuticals, such as sterility , apyrogenita
content of excipients and other non-radioactive components, as
well as details for the preparation of specific
radiopharmaceuticals are important for medical use, their
discussions, however, are beyond our physically and methodically
conceived treatise ...
4.9.
Clinical scintigraphic diagnostics in nuclear medicine
4.9.0.
Common general principles of clinical scintigraphy
General ideas of scintigraphic diagnostics were presented at the
beginning of §4.1 (section " Role and definition of
scintigraphy ; nuclear medicine ") . In the next text of
chapter 4, the physical principles of gamma imaging,
physical-electronic implementation of various scintigraphic
methods and computer processing of scintigraphic data were
discussed in detail. Here we will supplement this physical part
with some specific clinical applications . But
first we will make a few general remarks :
The human organism (as well as the organisms of all higher animals) is a very complex system both in its anatomical
structure and, above all, in the diversity of its internal
functions and metabolism. Imaging methods
provide some valuable ways to "look" inside this
complex system. Their basic output result is generally a brightness
modulated image : the brightness of each element of the
image is determined by the physical or chemical characteristics
of the corresponding site in the tissue. In X-ray imaging
, it is the absorption coefficient given by the density,
thickness and composition of the tissue. For sonography
, the acoustic impedance is displayed given the density,
elasticity and dissipation viscosity of the tissue. Nuclear magnetic
resonance displaysthe density of resonant nuclei
(especially hydrogen) and relaxation parameters
dependent on the binding of atoms in the tissue. Scintigraphy
shows the distribution of a radioindicator , showing the
movement of the radiopharmaceutical and its biochemical uptake,
rearrangement and excretion due to local metabolic and functional
processes (at the "molecular"
level) .
Most imaging modalities - X-ray
diagnostics, magnetic resonance, sonography - provide images of
the anatomical structure of tissues and organs -
their size and shape, placement, density inhomogeneities. On the
other hand, scintigraphic images (planar, SPECT, PET) have in
principle functional ones character. They do not
show any "real-palpable" objects, no morphology, but
capture the distribution - passage,
accumulation, excretion - of specific radiolabeled substances.
The degree of local accumulation of the radioindicator depends on
the intensity of local metabolic and functional processes, which
is reflected in the luminance modulation of the corresponding
sites (pixels) of the scintigraphic image. If the
radiopharmaceutical enters the examined tissue through the
bloodstream, not only the function but also the degree of blood
circulation - perfusion of this tissue or organ
can be assessed from the rate of uptake of the radioindicator (§4.9.4, part " Scintigraphy of myocardial perfusion " and §4.9.8, part " Perfusion brain scintigraphy ") .
Functional imaging of the
distribution of the radioindicator in the relevant tissues,
organs and lesions then serves primarily for diagnostic
purposes, but can also be an important starting point for
radiotherapy, in nuclear medicine for biologically
targeted radionuclide therapy :
Fig.4.9.1 Scintigraphic diagnostics and
radionuclide therapy in nuclear medicine
Static scintigraphy
Static scintigraphic images capturing radiolabel uptake are
usually evaluated visually , but semi-quantitative
assessments can also be performed using radiotracer accumulation
ratios in relevant areas of interest (ROI) or tissue background.
This results in certain relative numerical values - indices
*). A series of scintigraphic images is evaluated both visually
and quantitatively using computational algorithms (see below " Radiopharmaceutical uptake
" and " Dynamic scintigraphy ") .
*) Some software also allow comparison with databases
of normal patients and determination of certain so-called
scores - agreed quantitative parameters to facilitate
decision-making between normal and pathological findings.
Thus, radionuclide gammagraphy
depicts functional and metabolic processes in tissues and organs,
not their anatomical or morphological structure. However,
scintigraphic images can provide certain information about morphology
indirectly - by deriving from the representation of the
distribution of a radioindicator in the functional tissue of a
certain organ or in tumor tissue. Non-functional tissue areas
and, in general, areas where the radioidicist does not penetrate
are not displayed. It is often useful to supplement and combine
("merge") scintigraphic functional images with
anatomical X-ray images (especially SPECT or PET with CT) - to
perform a functional-anatomical correlation (discussed above in §4.6 "Relationship
scintigraphy and other imaging techniques ", the" Mergers
images of PET and SPECT with CT and NMRI ") . This can lead to more
accurate and comprehensive diagnosis.
Disorders of function often precede
disorders of anatomical structures - especially if the disorder
is caused by altered molecular biochemical processes at the
cellular level. Using radioisotope nuclear medicine techniques
can therefore pathological changes often reveal earlier
than other diagnostic procedures - even before structural changes
are visible. Scintigraphy has although a smaller spatial
resolution than X-ray or MRI, but thanks register individual
photons gi s very sensitive to subtle changes in the distribution
of the radiopharmaceutical due to metabolic abnormalities.
Accumulation
( uptake ) of radiopharmaceuticals
Rate of uptake - the accumulation
of applied radiopharmaceuticals - in the examined tissues and
organs is an important indicator of physiological or pathological
function . This parameter, also called uptake ( up = take , take = take -> uptake
= absorption ) can be determined on
the scintigraphic image based on the measured number of
pulses in the marked area of interest (ROI) of the
investigated structure, which is basically directly proportional
accumulated activity. The number of photons emitted from the
accumulated radioactivity lesion or body towards the detector
camera is influenced by three factors :
- Is reduced by photons absorbed or dissipated
in the tissue located between the examined organ and the body
surface (for gamma 140keV it is about 50%
6cm soft tissue) .
- It can be increased by photons coming from a
radiopharmaceutical collected in the surrounding tissues before
and after the organ under investigation - body background
.
- The most important factor: Radioactivity in the
examined organ is determined by the applied activity
and the share of accumulation of this total activity in the given
organ(this proportion corresponds to the
functional status of the organ) . The
measured number of pulses is then directly proportional to the activity
, acquisition time and sensitivity of the gamma
camera .
To determine the percentage of
accumulation, we must first convert the activity
applied to the patient in [MBq] to the corresponding number
of pulses and acquisition time detected by the gamma
camera. This can be done either by multiplying the camera
sensitivity coefficient (§4.5,
passage " Sensitivity (detection efficiency) of
the gamma camera ") for a given radionuclide and collimator (if we have measured it in advance) , or ad hocby capturing an image of the applied
activity (syringe) with the radiopharmaceutical (subtracting the activity of the "empty"
syringe after administration and correcting for the half-life of
the radiopharmaceutical and the acquisition times) . Furthermore, it is desirable to make a correction
for the absorption (attenuation) of gamma radiation in
the tissue layer between the measured organ and the body surface (§4.3, passage " Adverse effects of SPECT and their
correction ", point " Absorption
of gamma radiation ") .
Determination of the percentage
accumulation may be useful especially for assessing the accumulation
capacity of the thyroid gland (§4.9.1
" Thyrological diagnostics
", especially before radioiodine treatment - § ..., passage
"...") ,relative renal
function (§4.9 .., " Nephrological
diagnostics ") and accumulation of radiopharmaceutical in tumors
during biologically targeted radionuclide therapy
(§3.6, part " Radioisotope therapy ") . In these images of
tumor lesions (especially on tomographic
PET images with 18 FDG) , a standardized value of SUV
accumulation is often used to assess tumor viability (described in more detail above in §4.2, section "
Quantification of positive
lesions on gammagraphic images - SUV
") .
Specificity of
scintigraphic diagnostics
The specificity of
scintigraphic methods may be lower and higher
than for other modalities, depending on the mechanism of
pharmacokinetics of the radioindicator used. E.g. in skeletal
scintigraphy (§4.9.7 "Skeletal scintigraphy ") bone metastases are seen before X-ray or MRI,
but cannot be distinguished from focal changes caused by other
mechanisms of increased osteoblastic activity (inflammation,
fractures) - high sensitivity but low
specificity . In scintigraphic methods mapping the
targeted and specific binding of a radiopharmaceutical to the
cells of the examined tissue (tumor scintigraphy, perfusion
scintigraphy of the myocardium or brain, receptor diagnostics),
scintigraphic diagnostics can be highly specific
.
Dynamic scintigraphy
In addition to imaging and localizing structures in which a
certain radiopharmaceutical accumulates physiologically or
pathologically, it is sometimes important to assess the temporal
dynamics of this accumulation or the passage of a
radioindicator through the examined organs. This is the task of dynamic
scintigraphy , capturing the time course of the
distribution of the radioindicator - individual phases of the
passage of the radiopharmaceutical through the examined organ -
using a series of sequential images examined
areas, scanned sequentially at selected time intervals *). The
acquisition times of the images are chosen with respect to the
speed of the studied event; they range from tenths of a second
(for fast events such as cardiology) to tens of seconds or
several minutes (kidneys, liver). The total scanning time is
determined by the duration of the investigated event, ranging
from one minute to one hour. The obtained series of images can be
evaluated both visually (to observe the passage,
accumulation and leakage of the radio indicator
in various places), but above all quantitatively
. In the pictures, we draw regions of interest
(ROI) of significant structures, from which we construct curves
of the time course of radiopharmaceutical distribution.
We can then perform mathematical
ou analyzes of
curves -
for significant points and sections of curves
various time
intervals, ratios, integrals and other quantities are determined,
relevant functions are interpolated by
least squares method , rate coefficients of increase or decrease
of radioactivity are calculated , curves are derived,
integrated, filtered, deconvolution, etc. - according to the used
mathematical model of the investigated process. By this mathematical analysis of the time dependence
curves of the indicator radioactivity in the relevant tissues and
organs, we can obtain diagnostically important quantitative
parameters of their function, both total and regional.
*) A special type of dynamic scintigraphy isphase
scintigraphy of the cardiac cycle - ECG gated
ventriculography . It is not created by simple sequential
imaging, but by periodic recording and synchronous composing of a
large number of consecutive images in different corresponding
phases of the cardiac cycle - a detailed dynamic scintigraphy of one
representative cardiac cycle is created (described in detail
in §4.4 "Gated phase scintigraphy ").
Functional parametric
images
When evaluating dynamic scintigraphy, we can obtain various
quantitative parameters of the function, not
only total, but also regional, local . In the
extreme case, we can imagine that every pixel -
the pixel (i, j) of the image matrix will be considered as a
small elementary area of interest (microROI).
From this microROI we can construct a curve of the
time course of the radio indicator distribution at the
corresponding place and mathematically process it
using a certain model: eg interpolate an exponential or other
suitable function and determine a certain diagnostically telling
dynamic parameter - eg velocity coefficient, slope gradient,
increase or decrease half-life. radioactivity. The value of this
locally calculated parameter is then stored in the same
localized element (i, j) new image matrices, which we
declare in the computer's memory. This is done for all pixels of
the image of the examined area. By such processing of time curves
from all microROI, ie from all pixels (i, j) of
the image, we get a new artificial image, which
no longer expresses the measured numbers of pulses on the
scintigram, but shows the distribution of a
diagnostically important parameter of the investigated
function - functional parametric image , clearly
mapping the local distribution of the function of the
investigated organ. Thus, a parametric image in individual
pixels, instead of the number of pulses, contains a number-value
that characterizes a functional parameter.
Functional parametric images are
most often used in the evaluation of gated ventriculography -
heart rate image, paradoxical image, Fourier images of phase and
pulsation amplitude in individual places of the heart chamber
(see below " Gated ventriculography ", or in more detail in " Radionuclide
ventriculography " of the book
"Ostnucline"). Special parametric images are provided
by the so-called factor analysis (described
in the works of M. ámal and H. Trojanová ........) .
Clearance
Important global parameter that can be obtained by dynamic
scintigraphy is called. clearance (Eng. Clear = clean ).
It is the amount (volume) of blood or plasma in the body that is
"purify" of a certain monitored substance per unit
time. After a single intravenous administration of the
appropriate radioindicator, we first monitor its accumulation and
then its leakage from blood or plasma. The rate
of decrease in plasma activity concentration *) depends on the
elimination performance of the examined organs - kidneys or
liver; with impaired function of the relevant elimination organs,
the clearance value decreases. In dynamic scintigraphy, we
construct a curve (histogram) from the area of ??interest outside
the elimination organs, preferably from the precordium or lung
area. This elimination curve represents the time course
of the radioindicator concentration in the blood / plasma. We
then evaluate this curve from the " blood pool
" withcompartmental analyzes - we
interpolate bi- or multi-exponential functions
using the least squares method . The elimination curve is
basically created by the composition of two exponential
functions. The first, faster exponential, is a manifestation of a
relatively rapid balancing of the radioindicator concentration
between the blood (intravascular) volume and the interstitial
environment (by filtering plasma from blood
capillaries into tissue fluid to transfer nutrients and oxygen to
cells)- dilution into the entire
distribution area of ??the radiopharmaceutical. The second, more
gradual component, is a reflection of the organ's own clearance
by the elimination organ (kidney or liver). By interpolating the
biexponential function, we mathematically separate both
components, while the velocity coefficient in the later and
slower exponential indicates the required value of plasma
clearance . If the radioindicator used is excreted only
in the organ under investigation, it is also the organ
clearance value . The value obtained by this calculation
is relative [sec. -1 ]; the actual (absolute) clearance in [ml./sec.] is
obtained by multiplying the velocity coefficient by the value of
the total distribution volumethe indicator used
in the organism. The elimination curve of the blood pool can be
further combined with organ curves and used for the so-called deconvolution
analysis , by which we obtain the so-called transit
functions and transit times of the radiolabel
passage through the elimination organs and their parts (eg
parenchyma and hollow kidney system).
*) Plasma clearance was previously measured
by sampling methods : after iv
administration of the radioindicator, blood samples were taken at
certain time intervals and their plasma activity was measured
using scintillation detectors. Using simplified compartmental
analysis, clearance values were calculated from them (originally
by graphical analysis on semi-logarithmic paper).
Combination of diagnostics and therapy -
teragnostics
Teragnostics is generally a treatment strategy that specifically
combines diagnostics with therapy. New diagnostic
imaging methods, especially molecular imaging in
nuclear medicine, make it possible to integrate
individual (personalized) diagnostics and targeted therapy (or prevention) of serious diseases into a common
field, for which the name teranostics or teragnostics (created by composing names: therapy +
diagnostics => teragnostika
expires. Theranostics ) .
It's a kind of "diagnostic of therapy
. "Analysis of the diagnosis and possible treatment for an
individual patient to specify whether the selected therapy will
be effective even before its commencement. A further assess
response to performed by a therapy.
For teranostics in nuclear medicine
is optimal when the atom diagnostic gamma radionuclide(such as 99m Tc or 68 Ga) has a similar chemical
coordination of electrons in the valence shell as a therapeutic
beta or alpha radionuclide( 90 Y, 177 Lu, or 227 Th). This allowing the same biochemical vector
molecule to be used for scintigraphic diagnostics as
well as for subsequent radioisotope therapy - just labeled with
another radionuclide. The diagnostic drug is then focused on the same
molecular target as the therapeutic radiopharmaceutical, which
allows to determine in advance the optimal applied
activity and estimate the effectiveness of treatment -> theranostics
:
Fig.4.9.2 Principle of teragnostics in
scintigraphic diagnostics and biologically targeted radionuclide
therapy in nuclear medicine.
Above: The same ligand-targeted
biomolecule is labeled first with a diagnostic radionuclide
(gamma or positron) and then with a therapeutic radionuclide
(beta or alpha) using a suitable chelator. Bottom:
Use of the resulting diagnostic radiopharmaceutical for
scintigraphy, or a therapeutic radiopharmaceutical for
biologically targeted radionuclide therapy.
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. We will first label the
relevant biologically targeted substance with a diagnostic
gamma-radionuclide, apply low activity and
perform scintigraphic imaging. In case of successful uptake in
target tissues (and sufficiently low
unwanted uptake in healthy tissues - in critical organs), we label the same substance with therapeutic beta- or
alpha-radionuclide and apply high activity to
the patient (determined individually based
on scintigraphy) . It can be almost
certainly assumed that this therapy will be successful..!
..
Teranostic
radionuclides
Radionuclides suitable for theragnostics can be basically of
three types :
1 . One radionuclide
with mixed radiation beta - + gamma, beta - + beta +
, alpha + gamma, or alpha + beta + , which is used to
label the relevant biologically targeted radiopharmaceutical.
Gamma or positron emission allows scintigraphic imaging of planar
/ SPECT or PET. The emitted beta electrons or alpha particles
cause the radiobiological therapeutic effect of the desired
destruction of pathological cells in the target tissue where the
radiopharmaceutical has been taken up - teranostics. Such a
radionuclide that is capable at the same time to
enable diagnosis and therapy, we can (working)
call it a " monotheranostic
radionuclide ".
The best known example of such a
" monoteranostic " radionuclide is the
classical radioiodine 131
I , whose gamma radiation of energy
364keV allows scintigraphy (planar or
SPECT) , while beta electrons -
can exert a therapeutic effect - with significantly higher
applied activity. It has been used for decades in radioisotope
diagnostics + thyroid therapy (hyperthyroidism
and metastases of differentiated cancer, see below §4.9.1 ), although the name
"teranostika" was not introduced at that time. More
recently, the iodine-131-labeled monoclonal antibody tositumomab ( Bexxar ) has been used to
treat lymphomas.
In principle, some other
"monoteranostic" mixed radiation radionuclides, such as
lutetium 177 Lu, can be used for teranostic purposes . So far, the
experimental teranostic radionuclide is terbium 149
Tb with mixed alpha-beta + radiation :
annihilation radiation from positrons + can be used for PET diagnosis , emitted alpha particles
cause a therapeutic effect(Concomitant
gamma radiation is not very suitable here for teranostics; the
main problem of 149-Tb is the complex conversion scheme with a
number of secondary radionuclides - see 149 Tb ) .
2 . Two radioisotopes of the same element ,
one of which emits gamma photons or positrons beta + for scintigraphic
diagnostics, the second electron emitting isotope beta -
or alpha-particles for therapeutic effect. If we label the same
biochemical with these two different isotopes, we get a
diagnostic radiopharmaceutical for scintigraphy and a therapeutic
radiopharmaceutical that will have identical biochemical
properties to achieve successful teranostics
.
An example is iodine 123
I for scintigraphy and iodine 131 I for therapy(used mainly in
thyrology). So far, the experimental pairs
of teranostic radionuclides are the positron isotope
64 Cufor
imaging PET and the beta-isotope
67 Cufor
therapy, similarly 64 Sc/ 67 Sc, or the pair 86
Y for PET diagnostics and
90 Yfor beta
therapy.
3. Two different radioisotopes of different
elements, one of which emits gamma photons or beta+positrons for scintigraphy,
the second radionuclide is a beta - or alpha emitter
for therapy. Each of these radionuclides is chelated to the same
biochemical substance to give the appropriate diagnostic and
therapeutic radiopharmaceutical. Their identical biochemical
properties are no longer 100% ensured here, they
can be influenced by various chelators and complex chemical bonds
- they need to be carefully verified ... In the
positive case, the teranostic approach will also be successful
. A new interesting method of 100% ensuring identical
pharmacokinetics of diagnostic and therapeutic
radiopharmaceuticals are the radiohybrid theranostic
radiopharmaceuticals described below .
This method of teranostics is performed,
so far mostly experimentally, using several radionuclides, for a
number of monoclonal antibodies with the help of various
chelators. For scintigraphic imaging of planar / SPECT / PET, the
radionuclides 99m-Tc, 111-In, 18-F, 68-Ga, 89-Zr, 124-I are used,
which are combined with the beta-radionuclides 90-Y, 177-Lu, or
with alpha-radionuclides 212-Bi, 227-Th, 225-Ac - in each case
the same monoclonal antibody. Recently, for
example, the combination (68Ga / 177Lu) -PSMA J591 or (68Ga /
225Ac) -PSMA-617 in metastatic prostate ca appears to be
promising.
Radiohybrid teranostic
radiopharmaceuticals
A newly developed interesting "trick" is to bind two
chelators to a targeted ligand molecule simultaneously
( hybridly ) with two required atoms ,
yet non-radioactive . For example, natural fluorine
19
F and natural lutetium natLu (consisting of 97.4% 175 Lu and 2.6% 176 Lu) . If we then add to such a preparation the appropriate radioactive
isotope - either 18 F or 177 Lu , it is labeled by the mechanism of isotope
exchange with one or the other radionuclide. As needed,
either 18 F for diagnosis (lutetium remains inactive) or 177
Lu for therapy (here
again fluorine remains inactive) . It also
automatically ensures the identical pharmacokinetics of
the substance labeled with the diagnostic and
therapeutic radionuclide in the sense of point 3. above, as
the two molecules are atom- chemically identical
, differing only in isotopically. It is an ideal feature for teranostics
. It is currently being tested on the PSMA 18 F / 177 Lu
.
Fig.4.9.3. Principle of radiohybrid teranostics. Above:
Ligand vector biomolecule with two bound chelators with
non-radioactive atoms (here natural fluorine and lutetium). Middle:
Binding of radioactive fluorine atoms 18 F or lutetium 177 Lu by isotope
exchange with inactive atoms. Bottom: Use
of the resulting diagnostic radiopharmaceutical for scintigraphy
(PET) or a therapeutic radiopharmaceutical for biologically
targeted radionuclide therapy.
Radionuclide
examinations in nuclear medicine
In nuclear medicine, a number of methods have been developed for radionuclide
examinations of various tissues and organs in order to
determine their normal or pathological conditions.
In the beginning of the field of
nuclear medicine (60s-80s of the 20th century), sample
methods were often used - blood or
plasma samples taken from patients after application of a
radioindicator were measured using detectors (mostly cavity
scintillation counters). From the measured activity of these
samples (in relation to the applied activity), the parameters of
function - clearance, distribution volumes - of the respective
radiopharmaceuticals in the monitoring organs or blood
circulation were determined using mostly empirical methods,
dilution principles, etc. Alternatively, the concentration of the
radiolabel in the body was measured using scintillation probes
aimed at organs of interest (eg kidneys, heart). During the
80s-90s and the first decade of the 21st century. these methods
were gradually abandoned. These were often "blind"
methods; from today's point of view, they provided less accurate
and less reliable results with greater laboriousness, with the
possibility of significant individual errors ;
they are already mostly abandoned .
Now the simple "equation" is that :
current (and
future) nuclear
medicine = scintigraphy + radionuclide therapy with open radionuclides .
The preparation of
the patient before the scintigraphic examination depends
on the examined process, the radiopharmaceutical used, the state
of health and the patient's previous medication. Above all,
before the examination, with a certain time interval, it is
necessary to discontinue such drugs that would adversely affect
the biodistribution of the radiolabel used or the function of the
examined organ (for thyroid examination
they are iodine preparations, for cardiac nitrates,
beta-blockers, diuretics, cardiotonics).
Prior to non-thyrological examinations, Chlorigen is given over
time to block the thyroid gland. Prior to the actual
scintigraphic examination, an appropriate radiopharmaceutical
must be prepared at the nuclear medicine institution (§4.8 " Radionuclides
and radiopharmaceuticals for scintigraphy ") , prepared in a syringe
with a specific activity optimized individually for individual
patients (usually based on body weight) *) . The activity of the radiopharmaceutical for
application is measured in a metrologically calibrated activity
meter (§2.3, section " Well ionization activity meters
") .
*) Guideline values of the
recommended applied activity for different types of
radiopharmaceuticals are given in the table in §5.7 "Radiation exposure during radiation diagnosis and
therapy ", section"
Radiation exposure of patients from radionuclide examinations
".
The time course of
scintigraphic examination depends mainly on whether it is static
or dynamic gammagraphy, what is the course of the examined
process in the organism and how fast is the pharmacokinetics of
radioindicator used. In static scintigraphy, the
application of the radiopharmaceutical is usually performed
off-camera (in the application room), the volume and speed of the
application do not matter, the actual imaging is taken with a
certain time interval , it is necessary to wait
until the radio indicator is sufficiently absorbed - it can even
be in 2 hours (e.g for skeletal
scintigraphy or for PET imaging of 18 FDG accumulation in tumor lesions) .
In dynamic scintigraphy, the radio
indicator is applied directly below the scintillation camera, the
field of view of which is set to the patient's examination area,
whereas dynamic imaging starting immediately
with the application. For dynamic scintigraphy of rapid
events (such as blood flow through
the atria and ventricles during angiocardiography, or monitoring
the dynamics of the perfusion phase in the brain or kidneys) it is necessary to perform a so-called bolus
application : rapid application of a radioindicator with
high activity in a small volume of about 0.5 ml . - compact
bolus, with immediate start of dynamic scintigraphy with
a sufficiently high frame rate (one or more
frames / sec.) .
Here we will briefly describe some
more important methods of scintigraphic diagnostics . In the
introduction to the individual areas of scintigraphic
diagnostics, we will first present a brief outline of the structure
and biological function of the examined tissue or organ
and its most common pathologies , based on which
we will analyze methodological approaches to the
diagnosis of relevant disorders and diseases. For each specific
method, we state itsmedical purpose used radiopharmaceuticals
, design testing, and finally the processing
and evaluation of samples and results scintigrams normal
and pathological (all listed Scintigraphic
images were acquired and evaluated at the Clinic of Nuclear
Medicine University Hospital Ostrava) .
Since there are a number of scintigraphic methods in nuclear
medicine, we have divided this topic into several numbered subchapters
according to the investigated organs, systems or separate issues
:
4.9.1.
Thyreological radionuclide diagnostics
The thyroid gland
is located in the front part of the neck and, despite its small
dimensions (approx. 5 ´ 7 cm), it is a relatively important organ, intervening
in a number of processes in the whole organism. Thyroid function
is closely linked to iodine metabolism in the
body. Sodium iodide NaI penetrates thyroid cells by transport via
the Na / I symporter (which is a 37 kDa transmembrane
glycoprotein) - the "iodine pump". Within thyroid
cells, iodine binds in the thyroglobulin molecule to form
moniodo- and diiodo-tyrosine. Their combination then produces thyroid
hormones - triiodothyronine (T3) and tetraiodothyronine
(thyroxine T4). Therefore, radioiodine is also efficiently taken
up by the thyroid gland. Thyroid hormones enter cells in the body
and are involved in regulating a number of metabolic processes in
the body. They affect the transport conditions on cell membranes
for the entry of sugars and amino acids into cells. T3 binds to
the corresponding T3 receptors on the surface of mitochondria and
thus regulates intracellular metabolism. It also binds to
T3-responsive domains in nuclear DNA and initiates mRNA
production for proteosynthesis in cells.
Thyroid pathology
The most common functional pathologies of the thyroid
gland are :
- hyperthyroidism
- increased thyroid function with excessive production of
hormones (T3, T4).
- hypothyroidism
- decreased thyroid function.
- functional autonomy - independence of
the function of certain areas in the thyroid gland on regulatory
mechanisms.
A common morphological
disorder is an enlarged thyroid gland or goiter
, which can be diffuse or nodular . Depending
on the function, the goiter may be eufunctional, hyperfunctional
or hypofunctional. Nodes - areas of increased
density, occur quite often in thyroid tissue. And not only one
node - unifocal , but also more nodes - multifocal disability.
In terms of function, they may have the same function as the
surrounding tissue, or they may be hyper- or hypofunctional.
.........
The most serious thyroid disease is
its cancer - thyroid cancer . From a
histological point of view, there are 3 basic types of thyroid
tumors :
- Differentiated adenocarcinoma, which is
further divided into follicular, papillary and mixed. Follicular
carcinomas (15-30% of all thyroid malignancies) are mostly
unifocal and spread mainly through the bloodstream. Papillary
and mixed cancers are the most common (30-70%), they are mostly
multifocal and metastasize mainly through the lymphatic system.
Differentiated thyroid carcinomas retain iodine
accumulation and are therefore successfully treated with
131 I radioiodine . (§3.6, section " Radioisotope
therapy ", passage
"") .
- Medullary carcinoma (approximately
5-10% of the incidence) is based on parafollicular
C-cells, it spreads mainly hematologically and its treatment is
more difficult than in differentiated ones. It often does not
respond to radioiodine .....
- Undifferentiated ( anaplastic
) carcinoma (approximately 10% of the incidence) originates from
follicular cells, metastasized by blood and lymphatic routes. It
tends to be quite aggressive with invasion of surrounding tissues
and the formation of more distant metastases. Its treatment is
difficult and usually unsuccessful (it does not respond to
radioiodine).
From a general point of view, the
issue of cancer is discussed in more detail in §3.6 " Radiotherapy",
radionuclide therapy, especially in the section" Radioisotope
therapy ", not only cancer, but
also, for example, hyperthyroidism.
Radioisotope
diagnosis of the thyroid gland is the oldest
method of nuclear medicine (first tested in
1938) . This is due to the strong ability
of the thyroid gland to accumulate iodine - and
thus even the radioiodine, the radioactive
isotopes. Previously performed only simple accumulation
tests with radioiodine 131I, whose amount in the thyroid gland was measured single
gamma-probe, was examined what percentage of the applied amount
of the radioiodine to captures in the thyroid gland. Later, scintigraphic
methods were introduced .
The
thyroid accumulation test
is now performed only before radionuclide
radioiodine therapy to determine the applied activity. The
patient is given about 0.5-1 MBq of radioiodine
, orally in the form of a solution of sodium radioiodide. After
its absorption from the GIT, iodine ions are taken up by the
functional tissue of the thyroid gland (or
even by metastases of differentiated thyroid gland) . After 6 or 24 hours, the captured activity in the
thyroid gland is measured either by a simple collimated
radiometric probe (see Fig.2.4.3 b
in §2.4, passage " Scintillation probe ") or by a gamma
camera (with marking and
quantification of ROI on the thyroid image)
and compared with the activity of the administered
radiopharmaceutical - the result is the percentage of
radioiodine taken up *). Normal values are about 5-15% in 6 hours
and 10-30% in 24 hours. This measurement can also be performed repeatedly
over several days to determine the dynamics of gradual leakage
( clearance ) of radioiodine from thyroid.
*) The results may be skewed by some drugs
containing iodine, which saturates the uptake mechanisms and
reduces the accumulation of radioiodine - these must be
discontinued!
Measurement of
radioiodine accumulation in the thyroid gland, as well as its
clearance - effective half-life, is important for
individual determination of the required applied activity of
radioiodine to achieve optimal therapeutic effect in
hyperthyroidism and autonomic adenomas -
see §3.6, section " Therapy of thyroid
gland with radioiodine 131 I ", section " Individually applied activity - Marinelli
equation " .
Thyroid
scintigraphy
Purpose:
To show the distribution of functional thyroid tissue
in the primary diagnosis, its location, shape and size and reveal
possibly. anomalies - areas of increased or decreased function in
thyroid tissue, finding ectopic thyroid tissue. Furthermore,
demonstration of functional properties of tactile nodes and
functional autonomy. In combination with laboratory determination
of T3, T4, TSH levels, assessment of hyper- or hypothyroidism
. In thyroid cancer therapy, scintigraphy is used to demonstrate
residual accumulating tissue and to detect distant accumulating metastases
of differentiated thyroid cancer.
Radiopharmaceuticals:
The basic method consists in the oral administration of radioiodine
131
I in the form of sodium iodide
Na 131 I.
Due to the higher radiation exposure (radiation b which is to
diagnose unusable), however, 131 I has the primary diagnosis is not used, it is replaced
by an isotope of iodine 123 I , or 99
m Tc
-sodium pertechnetate, which also scavenges and accumulates in
the cells of the thyroid iodide similarly, but unlike therefrom
does not bind to thyroglobulin and does not enter into other
metabolic reactions. 131 I (application approx. 10MBq) is used only in patients with proven thyroidopathy
before radioiodine therapy.
99mTc-MIBI, 99mTc-Tetrofosmin are also used to detect less
differentiated tumor tissue and metastases . It is further used
to image medullary carcinoma (which contains somatostatin
receptors)111 In-pentetreoid, further 123,131 I-MIBG.
Process :
For scintigraphy of the thyroid gland, as a small organ, we use a
high-resolution collimator using ZOOM, or a pinhole
collimator (which increases the projection of the thyroid gland
over a larger usable area of the camera's scintillation
detector). After iv application of 99m Tc-pertechnetate (approx. 100-200 MBq), a planar image
is captured in the AP projection after about 30 minutes. For
better morphological orientation, it is advisable to take a
picture with markers using a point "pointer".
Evaluation :
The normal image of the thyroid
gland has a like "butterfly" almost symmetrical shape,
with an approximately homogeneous distribution of the
radiopharmaceutical in both lobes of the parenchyma. The
pathological picture shows an inhomogeneous distribution
with "cold" nodes of reduced function or
"hot" nodes of increased thyroid tissue function.
Possibly. functional autonomy of hot nodes can be
determined by repeated suppression scintigraphy of the
thyroid gland after several days of administration of
triiodothyronine. In functional autonomy, the accumulation of the
radiopharmaceutical in the "hot" deposit does not
change, while in the other parenchyma (paranodular tissue), due
to hormonal suppression, the accumulation decreases or disappears
significantly (§3.9 " Quantitative thyroid
scintigraphy " in OSTNUCLINE) .
Normal thyroid scintigram |
![]() |
Whole-body scintigraphy after radioiodine therapy of the thyroid gland . Multiple accumulating deposits in both lung wings - metastases and thyroid gland - appeared . Under favorable circumstances, these malignant foci can be successfully eradicated by radioiodine therapy. |
Hyperfunctional node in the right lobe of glands |
||
Nodular goiter with unfunctional node on the left |
||
Examples of typical images of thyroid
scintigraphy (the pictures were taken by MD. V.Dedek, PhD., KNM Ostrava) |
When searching for functional metastases
of differentiated thyroid carcinomas, it is appropriate
to use whole-body scintigraphy with radioiodine (application approx. 100-200 MBq) .
Thyrological diagnostics in nuclear medicine can then be followed
by radionuclide therapy of the thyroid gland -
treatment of hyperthyroidism, autonomic adenoma, thyroid
carcinoma (see §3.6 " Radiotherapy
", section " Radioisotope therapy with open emitters ") . Afrer therapy is
followed by control scintigraphy of the thyroid
gland or metastases at certain time intervals .
Parathyroid
scintigraphy (parathyroid glands)
The parathyroid glands are 2 pairs of
small formations located on the back of both lobes of the thyroid
gland. They are the parathyroid hormone- producing
glands that affect the calcium content - it releases it from the
bones into the blood.
Purpose: Using scintigraphy, we try to show the
hyperfunctional parenchyma of enlarged
parathyroid glands (usually their adenoma), which by their
increased production of parathyroid hormone adversely affect the
turnover of calcium in the body.
Radiopharmaceuticals: As there are no
radioindicators that are selectively taken up in the parathyroid
glands, the cationic complexes 99m Tc-MIBI and 99m Tc-Tetrofosmin are used, but they also accumulate in
the thyroid parenchyma.
Design and evaluation: Displaying small
parathyroid glands against the background of much larger thyroid
tissue is not easy. We can help in two ways:
- Two-phase scintigraphy using faster leaching
of 99m
Tc-MIBI from thyroid tissue than from the parathyroid gland
affected by enlargement or adenoma. In about 15-30 minutes after
iv application of the radiopharmaceutical (approx. 700-800MBq) we
take the first image, the next in 2-3 hours.
- Subtraction scintigraphy performing computer
image reading. The image of the thyroid gland taken after the
application of 99m Tc-pertechnetate alone (approx. 200MBq) is subtracted
from the "summation" image [of the thyroid gland +
parathyroid glands], taken after the subsequent application of 99mTc-MIBI or 99m Tc-Tetrofosmin.
Both images must be captured under identical conditions, without
changing the position. After subtracting the image of the thyroid
gland tissue in the "summation" image, the image of the
parathyroid body itself remains.
4.9.2. Nephrolological radionuclide
diagnostics
The urinary excretory system, formed by the system [kidneys -
ureters - bladder - urethra], is collectively called the uropoietic
system .
The kidneys *) are mainly used to filter
blood (which is supplied to the kidneys by the renal
arteries ), which removes metabolic products and other
unnecessary or harmful substances from the body, which are then
discharged as urine out of the body. Waste products - catabolites
- of nitrogen metabolism (urea, creatinine), acid catabolites,
water and electrolytes are thus removed , thus maintaining a
stable internal environment. The kidneys also have a regulatory
function- ensure homeostasis of the organism - water, salts,
minerals, acid-base balance, participate in maintaining blood
pressure. The basic building block and functional unit of the
renal parenchyma is the nephron . The human
kidney contains about 800,000 to 1.5 million nephrons.
*) The kidneys are called ren in
Latin , and nephros in Greek - hence the synonym for renography
= nephrography in the examination methods .
The nephron begins with a ball
of capillaries called the glomerulus , where the
branching renal artery supplies blood. In the glomeruli, the
basic clearance function of the kidneys takes place - glomerular
filtration , which is a process of ultrafiltrationblood
plasma under pressure across the glomerular membrane. The
microporous structure of the glomerular wall (which prevents the
flow of plasma proteins larger than about 100 kDa) and the
electrostatic barrier of the glomerular membrane (negative charge
of polyanionic macromolecules of the membrane and negative charge
of most plasma proteins prevent the transfer of even smaller
proteins with molecular weight above about 60) . The glomerular
ultrafiltrate is primary urine , which is essentially
plasma without cells and large molecules of plasma proteins; more
than 150 liters are made in her kidneys a day. In addition to
waste metabolic products, it contains a number of substances and
nutrients (such as glucose) that should remain in the body.
The glomerular filtrate enters
a hollow canal of the nephron, called the tubule.
Here, tubular resorption takes place , during
which part of the substances from the glomerular filtration is
reabsorbed and returned to the blood, leaving through the
vascular bed around the tubules through a drainage vessel from
the kidney. Most water, glucose, amino acids, minerals return to
the blood by resorption. This maintains homeostasis - the balance
of water and salt in the body. In addition to resorption, tubular
secretion also occurs here - tubule cells actively take
up some substances from the blood (eg creatinine) and transport
them to the tubular cavities, ie to the urine. The total amount
of substance excreted in the urine is given by the sum of:
(glomerular filtration) - (tubular resorption) + (tubular
secretion).
After passing through the tubular system,
definitive secondary urine is formed, which consists of
water with dissolved urea, sodium chloride and a small amount of
other substances; about 1.5-2 liters per day is excreted. The
tubules open into thicker collecting ducts, funnel-shaped calyxes
and finally into the renal pelvis ( pelvis renalis ) of
the hollow kidney system, from where urine flows through the ureters
into the bladder. From there, after releasing the sphincter, it
flows out of the body through the urethra.
Kidney and urinary
tract pathology
The kidneys are relatively often
affected by inflammatory and infectious diseases. Pyelonephritis
is a bacterial purulent inflammation of the kidneys (pelvis )
parenchyma), which can be acute or chronic and can lead to
deterioration in renal function if repeated or prolonged. Glomerulonephritis
is an inflammatory disease affecting mainly the glomeruli in the
kidneys, which can occur after infections (especially
streptococcal), autoimmune processes and other causes. Some
kidney diseases can lead to deterioration of renal
function , which can be irreversible (nephron loss), in
extreme cases can result in kidney failure
.......
Very common kidney and urinary tract
involvement is lithiasis ( urolithiasis
) - " kidney stones ", formed by the accumulation and
increased concentration of mineral salts that crystallize
in the urinary tract, especially in the pelvis of the
kidneys or in the bladder. The most common are stones from
calcium oxalate or uric acid. They can grow to various sizes,
from small particles of "sand" to larger stones (> 1
cm), which can block the outflow of urine from the kidney. This
ureteral obstruction causes congestion in the
hollow system of the kidney, which also burdens the parenchyma
(which must filter against pressure); with prolonged obstruction,
renal function is irreversibly impaired.
When renal function is
impaired, glomerular filtration is reduced and thus waste
products are retained in the body and tubular resorption is
reduced, and the absorption of electrolytes and water is
impaired. This also affects blood pressure and can lead to
disorders of acid-base balance, or. and hematopoiesis.
The renal parenchyma is
affected by cystic (often
polycystic) disease. Renal tumors , such as Grawitz's
tumor , are relatively uncommon .................
The
kidneys and their excretory functions have already become a
suitable object for radioisotope diagnostics in the beginning of
the field of nuclear medicine.
Until the 1970s and 1980s, radioisope
renography was one of the most frequent nuclear medical
examinations . After application of the nephrotropic
radioindicator 131 I-hippuran with two collimated scintillation renographic
probes (see Fig.2.4.3 b
in §2.4, passage " Scintillation probe "), attached to the
patient's back in the localization of the kidneys, sensed the
course of radioactivity in the kidneys. The electrical signal
from the detectors was fed to a double recording recorder, the
pen of which plotted the so-called nephrographic curves
on paper . From the shapes of the nephrographic curves it was
possible to deduce various pathological conditions and disorders
of renal function, as well as urine outflow from the kidneys.
Semi-quantitative analysis of nephrographic curves was sometimes
performed. However, it was only an approximate "blind"
examination, without the possibility of regional assessment.
Replacing isotope nephrography with dynamic scintigraphy
has significantly refined diagnostics and provided much
more comprehensive informations :
Dynamic
renal scintigraphy
Purpose :
It is the most important method of nuclear nephrology. It is used
for a comprehensive assessment and quantitative analysis of
perfusion and excretory function of the kidneys
(and their parts), clearance and drainage - the dynamics of urine
outflow from the kidneys. It can also provide certain information
about the morphology of the kidneys, which, however, is derived
from the display of the radiolabel distribution in the functional
tissue of the kidneys (parenchyma) and from the outflow or
accumulation of the radiolabel in the hollow system.
Radiopharmaceuticals:
- 99m
Tc-DTPA (diethylenetriaminepentaacetic acid), which is
excreted by passive ultrafiltration in the glomeruli (not
resorbed in the tubules). In addition to functional imaging, it
is also suitable for the determination of glomerular filtration
by plasma clearance.
- 99m
Tc-MAG3 (mercaptoacetyltriglycine), which binds to
transport plasma proteins and is excreted by tubular secretion
(almost non-filtered in glomeruli). It provides contrasting
images of the functional renal parenchyma, captures the dynamics
of urine outflow from the kidneys well, and can be used to
determine the effective renal plasma flow (ERPF) by plasma
clearance. For this purpose, use is sometimes also
ortojodhippuran labeled with
123 I (f
131 I), which is excreted from about
80% by tubular secretion to about 20% by glomerular filtration.
Execution :
Under the scintillation camera, set in the rear projection on the
kidney area, approx. 200MBq of radio indicator is applied and
dynamic acquisition is started immediately in short time
intervals: perfusion phase approx. 1s / frame -
100 frames, followed by functional phase approx.
10-30s / frame, total acquisition time 30 minutes. The images are
stored sequentially in the computer's memory. In the case of
visible retention in the hollow system, a diuretic
is applied in about 15 minutes - a substance that increases
diuresis (intensity of urinary excretion) by affecting transport
in various parts of the nephron; furosemide is the most
commonly used .
Processing :
By observing the sequences of images, we can visually
observe the entry of the radioindicator into the kidney, its
uptake in the parenchyma, transport into the calyx-pelvic system
and outflow of the ureter into the bladder. We can recognize or.
abnormality. Forquantitative analysis on
suitable summation images we mark the areas of interest
(ROI): the area of ??the bloodstream (area around the heart),
left and right kidneys and their parts (parenchyma, hollow
system), tissue background (on which correction is performed -
subtraction). From these areas, the computer then creates curves
("histograms") of the time course of activity in these
places. Nephrographic curves are created from
the ROI of the kidneys , from which the time and speed parameters
of reaching the maximum, speed or half-life are determined. A diuretic
test is important in case of delayed excretion or
retention of the radiolabel in the kidney : If the decrease in
the nephrographic curve did not occur even after the application
of the diuretic, this indicates obstructive hydronephrosis
.
By differentiating the
activity of the parenchyma and the pelvis, it is possible to
decide whether the pathology of the nephrographic curve is caused
by a functional disorder of the indicator in the parenchyma or
changes in the outflow (dilatation of the hollow system or
obstruction of the urinary tract in the kidney) - this is exactly
done by the so-called deconvolution analysis of
nephrographic curves.
The curve from the area of
interest ROI of the bloodstream represents the time changes
(especially the decrease) of the concentration of the
radioindicator in the blood (plasma). The rate of decrease in
plasma activity concentration depends on the elimination ability
- clearance - of the kidneys, with impaired
renal function, the rate of clearance is reduced.
Mathematical analysis and complex evaluation of dynamic functional scintigraphy of kidneys - MAG3 | |
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Evaluation:
After intravenous administration of the radioindicator,
the kidneys of the usual shape, size and placement are
displayed, without focal changes. The nephrographic curve
of the left kidney has a normal course, on the curve of
the right kidney we observe a slowdown of drainage and
retention, disappearing after diuretics. Conclusion: Visual
evaluation of sequential images and quantitative analysis
of nephrographic curves indicate good function of
both kidneys, rapid transit through the parenchyma and
free drainage of the hollow system. In the right kidney,
a slight slowing of dilatation- type
drainage . |
Here are examples of evaluation almost normal á and distinctly pathological â dynamic renal scintigraphy.
Mathematical analysis and complex evaluation of dynamic functional scintigraphy of kidneys - MAG3 | |
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Evaluation:
After intravenous administration of the radioindicator, a
well-accumulating left kidney of the usual shape and size
was displayed, without focal changes. The right kidney
appears late as markedly hypofunctional
and inhomogeneous - only the narrow margin of the
functional parenchyma around the markedly dilated
excavated hollow system with significant retention
is preserved . The nephrographic curve of the left kidney
has a physiological course. The nephrogram of the right
kidney has a markedly flat shape with a low functional
segment, the curve has a permanently ascending course,
unresponsive to the application of a diuretic in the 17th
minute. Conclusion: Visual evaluation of sequential images and quantitative analysis of nephrographic curves indicate good left kidney function, but severely hypofunctional right kidney with marked renal parenchymal atrophy. Left renal drainage physiological, right obstructive drainage disorder , no response to diuretic. Global kidney function is almost normal due to age. Signature: MUDr. Jozef Kubinyi, Ph.D. |
The mathematical procedure for the analysis of
dynamic scintigraphy of the kidneys is described in detail in
§3.4 " Dynamic scintigraphy of the kidneys " of the book "OSTNUCLINE - Mathematical
analysis and computer evaluation of functional
scintigraphy".
Renovascular
hypertension - captopril
test
Elevated blood pressure is a disorder that can seriously endanger
health, especially vascular complications. It is usually a
primary hypertensive disease , but high blood pressure
can also be caused secondarily, by a disease of some other
organs. This is often secondary to nephrogenic hypertension
in kidney disease such as pyelonephritis or glomerulonephritis.
Here, nuclear nephrology can also be used in differential
diagnosis. A specific case is renovascular
hypertension - increased blood pressure caused by
insufficient perfusion of the kidneys (their ischemia) due to
stenosis of the renal artery. This retains water and sodium in
the body, as even otherwise healthy kidneys cannot sufficiently
fulfill their function. The RAAS renin-angiotensin-regulatory
system is activated: renin , produced to an increased
extent in the ischemic kidney, is converted to angiotensin II
by the action of a conversion enzyme., which maintains
the blood flow of the glomerulus at the required value by
increasing the pressure in the glomerulus. However, this
physiological compensation by systemic action on the arterioles
and an increase in aldosterone levels leads to an undesirable
increase in blood pressure. Inhibition of angiotensin converting
enzyme (ACE) by a suitable drug can block this regulatory
mechanism. Such a short-acting ACE inhibitor is captopril
, which can be used here for diagnostic purposes.
We therefore apply captopril
before starting dynamic scintigraphywhich, by inhibiting ACE,
reduces tone in vas efferens and reduces glomerular
filtration. The secretion of DTPA by glomerular and MAG3 tubular
cells is slowed down, so that the originally normal nephrographic
curves become pathological - a decrease in glomerular filtration
and thus a slowing down of the transport of the
radiopharmaceutical by the renal parenchyma. By comparing the
kidney curves from native dynamic scintigraphy without captopril
with scintigraphy after captopril, we can reveal the renovascular
origin of hypertension.
Dynamic scintigraphy of the
transplanted kidney
The principle and methodological procedure are basically
analogous to the above-mentioned dynamic scintigraphy of the
kidney. In addition to the assessment of the clearance
function of the transplanted kidney, it is important to
assess in particular its perfusion , acute
tubular necrosis *) and the risk of rejection
, drainage of the graft and ureter, detection or
complications and anomalies in transplantation (such as urinome
).
*) Somewhat misleading name " acute tubular necrosis -
ATN" means a delayed onset of perfusion and renal graft
function after transplantation, depending on the duration of cold
ischemia kidneys in the time interval between removing the graft
donor and transplant recipients. Severe ATN may result in
rejection.
In contrast to the aforementioned dynamic renal
scintigraphy is different projections - camera in the AP
projection, the AP includes the area including the iliac
arteries, the transplanted kidney and the bladder, and the Hilson
perfusion index and rate are determined from the area of
interest: blood pool, illiaca artery, transpl. Analysis of
dynamic scintigraphy of the transplanted kidney is described in
§3.5 " Dynamic scintigraphy of the transplanted kidney" OSTNUCLINE books.
Mathematical analysis and complex evaluation of dynamic functional scintigraphy of a transplanted kidney | |
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Evaluation:
After intravenous administration of a radioindicator, the
abdominal aorta and iliac artery are imaged in the usual
way, followed by a well-perfused
transplanted kidney. In the further course, the
radiolabel is well concentrated in the kidney, then
excreted into the bladder quickly enough. Conclusion: Visual evaluation of sequential images and quantitative analysis of the curves of the passage of the radio indicator indicate good perfusion and function of the graft , rapid transit through the parenchyma and free drainage of the hollow system. There are no signs of incipient rejection. |
Static
scintigraphy of the kidneys
Purpose: Using this simpler method le to obtain
functional-morphological information about the distribution
of the functional parenchyma in the kidney, derived
about the shape, size and location of the kidneys, sometimes
structural changes. Separate renal function
can be determined - % share of left and right kidney in total
function (renal functional symmetry test).
Radiopharmaceuticals: Labeled substances are
used as radiopharmaceuticals that are taken up by the renal
parenchyma but do not pass into the urine. The most widely used
is 99m
Tc-DMSA (dimercaptosuccinic acid), which is taken up in proximal
tubule cells.
Design: After iv application approx. 100MBq 99mTc-DMSA is performed
in about 2 hours by its own static scintigraphy, mostly in 4
projections, the most important of which is the projection of PA
and AP.
The evaluation is mostly visual , the
size, shape and placement of the kidneys, the distribution of
functional tissue are evaluated. Computer processing is performed
to determine the separated function (corrected for the absorption of radiation g from various
deep-seated kidneys) is described in §3.7
" Static
scintigraphy of the kidneys " of
the book OSTNUCLINE.
Radionuclide urowlowmetry and cystography
Purpose: Used to examine the
dynamics of micturition, determination of bladder volume, bladder
residue, evacuation rate, detection of vesicoureteral reflux.
Execution: A patient whose bladder
is filled with a radioactive solution is urinated in front of a
camera detector, while dynamic scintigraphy of the ureter and
bladder area is scanned, at a frequency of about 1 frame / sec.
There are two methods of filling the bladder, direct and indirect
cystography. The indirect method consists in the
application of a nephrotropic radiopharmaceutical (usually 99m Tc-MAG3, approx.
200MBq), after which a normal dynamic scintigraphy of the kidneys
is performed, during which the bladder is filled with renal
function. When direct cystography with a
radioindicator (approx. 50MBq) fills the bladder through the
catheter.
Evaluation: In addition to the visual assessment of a series of
images of the micturition, we mark the areas of interest of the
bladder, ureters and tissue background, from which we create
curves of the time course of radioactivity, especially the urodynamic
curve . By their computer analysis we can quantify the
course of micturition - determine the duration and speed
of micturition , bladder residue . We
can visually assess and quantify the regurgitation of
urine from the bladder to the ureters or. up to the pelvis of the
kidneys - vesicoureteral or vesicorenal
reflux . ....
Computer analysis of dynamic uroflowmetry is described in §3.6
" Radionuclide
uroflowmetry " of the book
OSTNUCLINE.
Mathematical analysis and complex evaluation of dynamic uroflowmetry | |
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Evaluation:
After the start of dynamic scintigraphy, micturition soon
begins, during which there is a sufficiently rapid
emptying of the bladder with a low residue.
Vesico-ureteral reflux, more pronounced on the right, is
well visible in scintigraphic images of the emptied
bladder and urinary tract. Conclusion: Radionuclide uroflowmetry shows normal micturition flow and low bladder residue, but shows vesico-ureteral reflux . |
4.9.3 Diagnosis of the gastrointestinal
tract - liver and bile ducts, pancreas, spleen,
esophagus and stomach
Liver scintigraphy
The liver is an important organ in
which important metabolic, detoxification and elimination
processes take place; they are incorporated into the digestive
tract and also into the reticuloendothelial system
(RES). They are one of the largest organs, weighing about 1.5 kg
in humans, are located in the right diaphragmatic arch of the
abdominal cavity. The liver parenchyma consists mainly of
polygonal liver cells - hepatocytes (60%) and
reticuloendothelial Kupfer cells (15%). Then there are
the hepatic stellate It cells, Pit cells and
the walls of a large number of blood vessels in the blood and
intrahepatic bile ducts.
Hepatocytes
take up various substances from the plasma, transform
them and then excrete them into the bile , which
leaves the intrahepatic pathways (via the gallbladder) through
the ductus choledochus to the intestinal tract. Liver cells are
significantly involved in a number of metabolic and synthetic
processes :
- Carbohydrate
metabolism - liver cells take up glucose from portal blood
and convert it to lipids or glycogen, conversion of lactate and
alanine to glucose - Lipid metabolism - fatty acid synthesis and
oxidation, glycerol formation , phospholipids, cholesterol,
lipoproteins ...
- Amino
acid metabolism -
- Synthesis of plasma proteins -
- Detoxification
function - It is mainly the detoxification of ammonia
(formed during the decomposition of amino acids). Ammonia is
converted to urea and glutamine, which are then excreted in the
kidneys. Furthermore, some foreign molecules (especially
hydrophobic, which cannot be excreted by the kidneys) are
oxidized by cytochrome and excreted in bile or plasma (from where
they are then removed in the kidney).
Red blood cells have a limited lifespan (about 120 days), after
which they are taken up in the spleen and liver and sequestered.
Iron is separated from hemoglobin, which is used to synthesize
new hemoglobin, while the remaining component (heme) bound to the
hemopexin protein is phagocytic by Kupfer cells and converted to
bilirubin. It binds to the blood protein albumin and is taken up
by hepatocytes in the liver. binds to glucuronic acid. This
produces conjugated bilirubin , which the liver cells
secrete into the bile , which then drains into the small
intestine.
The liver also has a number of other functions - the production
of hormones and their degradation or inactivation, cholesterol
degradation, part of hematopoiesis, storage functions of lipids
and glycogen.
Hepatocytes secrete water, salt ions,
acids, cholesterol, phospholipids and bilirubin - liver bile
- into the bile capillaries, which gradually coalesce into the
bile ducts. Bile collects in the gallbladder - a
sac-shaped "reservoir" on the bile, from where it is
controlled and drained through the bile ducts - ductus cysticus -
ductus choledochus - into the duodenum (duodenum) and from there
into the intestinal tract (small intestine), where it is involved
in fat digestion. . Kupfer
cells are fixed macrophages that phagocytose
bacteria, foreign proteins, persistent erythrocytes and some
other cells. Pit cells are large granular
lymphocytes with significant cytotoxic activity, which by their
phagocytic ability cooperate with Kupfer cells. It cells
(lipocytes) contain a large amount of lipids.
The blood circulation
of the liver is about 1.5 liters / minute and has two
components :
1. The hepatic artery supplies blood rich in
oxygen (20% of the blood circulation), nourishes the liver
parenchyma.
2. Functional circulation from the portal vein
- vena portae (approximately 80% of the hepatic
circulation) brings blood containing absorbed nutrients from the
digestive tract, as well as various products of cell metabolism.
The vena portae branches into veins flowing through the
portobiliary space between the hepatocytes, and the blood then
flows through the liver veins into the inferior vena cava.
Pathology of the liver
and bile ducts
The liver has a large functional reserve and the ability to
regenerate. However, the liver can be damaged due to excessive
exposure to toxic substances , hepatotoxins (such as
alcohol), inflammatory and infectious diseases (hepatitis A, B,
C). These lesions can result in nodular remodeling, fibrosis, and
gradual disappearance of the liver parenchyma - liver
cirrhosis associated with liver failure and vascular
complications - by portal hypertension and portosystemic shunts
(see "spleen" below).
Pathologies
of the bile ducts , especially cholelithiasis
, are relatively common - stones in the gallbladder, which are
caused by increased concentration and decreased bile solubility.
Gallstones can cause inflammation and clog the bile ducts,
preventing the outflow of bile from the liver - obstructive
jaundice occurs, manifested as jaundice caused by the
accumulation of bilirubin in the plasma.
Liver cancer
- the primary liver tumor is hepatocellular carcinoma
(hepatoma). Much more common are secondary liver involvement with
metastases from other tumors (most often breast
or colorectal ca). A benign tumor of the liver is hemangioma
.
Radionuclide
diagnostics of the liver primarily uses the functions of
hepatocytes and Kupfer cells. After administration of the
radiopharmaceutical, which is a bloodstream hepatopcyty uptake,
we dynamic liver scintigraphy - cholescintigrafií
investigate the function of liver and biliary tract. By applying
a radiopharmaceutical that is taken up (phagocytosed) in Kupfer
cells, we can show the distribution of the parenchyma by static
scintigraphy of the liver and thus (indirectly) obtain
information about the morphology of the liver.
Dynamic
liver scintigraphy - cholescintigraphy
Purpose:
It is used for comprehensive assessment and quantitative analysis
of hepatic excretory function
(and its parts), clearance and drainage - dynamics of bile
formation and outflow through intrahepatic pathways
into the gallbladder, gallbladder evacuation, duodenal and
intestinal tract. It can also detect duodeno-gastric reflux. In
addition, it may provide some information on the morphology of
the liver, which, however, is derived from the display of the
radiolabel distribution in the functional tissue of the liver
(parenchyma) and from the outflow or accumulation of the
radiolabel in the bile ducts.
Radiopharmaceuticals:
Iminodiacetic acid (IDA) derivatives labeled 99m Tc - HIDA, EHIDA
Execution:
After iv application of approx. 100-200 MBq of hepatotropic
radiopharmaceutical while lying in the front projection, dynamic
scintigraphy is immediately started at a frame rate of approx.
20-30 sec./frame. To event. stimulate gallbladder emptying, a
cholekinetic stimulus - cholecystokinin or a fatty diet
(chocolate) is given during the examination in about 30 minutes.
Dynamic scintigraphy is scanned for 60 minutes. If the radio
indicator does not appear in the intestines after this time,
further still images are recorded in 2 and 4 hours.
Evaluation:
By visual evaluation of images of different phases of
radioindicator distribution in the liver and passage through the
bile system, we assess the distribution of hepatocytes in the
liver parenchyma, bile duct morphology, gallbladder deposition
and size, its emptying and bile drainage into the intestinal
tract, event. duodenal-gastric reflux.
On the relevant dynamic images, we mark the regions of
interest: bloodstream, whole liver, liver parenchyma, ductus
choledochus, intestinal tract, gallbladder (if seen) and stomach
(if duodenal-gastric reflux is suspected). From the curves
of the time course of radioindicator concentration in
these areas of interest, we evaluate and quantify hepetocellular liver
function - clearance and rate of radiopharmaceutical
extraction by hepatocytes from the bloodstream, biliary
outflow dynamics, including determination of gallbladder
ejection fraction , duodeno-gastric reflux rate.
Furthermore, we can construct transit functions
from them and determine the transit times of passage of
the radioindicator through the whole liver and liver parenchyma.
Mathematical analysis and complex evaluation of dynamic functional scintigraphy of the liver and bile ducts - cholescintigraphy | |
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Evaluation:
After intravenous administration of the radioindicator,
the liver of the usual shape and size is imaged in a
timely manner. The liver parenchyma does not show focal
changes. Bile ducts can be differentiated from 10
minutes, from 13 minutes. the gallbladder begins to fill.
In 30.min. a cholekinetic stimulus was administered. In
the next course, we observe a rapid passage of the
radioindicator through the biliary system with a smooth
outflow through the ductus choledochus into the intest.
tract. Conclusion: Visual evaluation of sequential images and quantitative analysis of liver curves indicate good hepatocellular function , rapid transit through the liver parenchyma and free drainage of the biliary system, without signs of biliary obstruction. The gallbladder has a good filling and evacuation function. |
Mathematical analysis and computer evaluation of dynamic cholescintigraphy is described in detail in §3.10 " Dynamic liver scintigraphy " of the book "OSTNUCLINE".
Static scintigraphy of the liver
Purpose: To obtain indirect information about the morphology of
the liver - detection of diffuse involvement and detection of
focal liver lesions.
Radiopharmaceuticals: Sn-colloid
(or sulfur colloid) labeled with 99m Tc, which is rapidly taken up from the bloodstream by
Kupfer cells after application.
Execution: After iv application of approx. 150 MBq 99m Tc-Sn colloid, in
15 min. performs scintigraphy in the front, back and front. right
side projection. For better imaging of lesions deposited deeper
in the parenchyma, it is advisable to perform SPECT imaging.
Evaluation:
In addition to the placement, shape and size of the liver, we
visually assess the distribution of the radiopharmaceutical in
the parenchyma on planar or tomographic images. Liver lesions are
usually accompanied by decreased Kupfer cell density, which
results in decreased radiopharmaceutical accumulation. The
finding is non-specific : local reductions
(cold deposits) may be caused by cysts, abscesses or tumors
(metastases), diffuse involvement (hepatomegaly,
uneven distribution in the parenchyma, increased accumulation in
extrahepatic RES - spleen and bone marrow) may be caused by
hepatitis, cirrhosis, metabolic disorders, malignancies.
Scintigraphy
of hepatic hemangiomas
Hemangioma is a benign mesenchymal tumor of blood
vessels. A cavernous type of hemangioma often occurs in
the liver . It is a highly vascularized structure that has a
higher proportion of blood - and thus a higher
concentration of erythrocytes - than the surrounding
tissue. After application of radionuclide-labeled erythrocytes,
hemangiomas appear as "hot" deposits of increased
radioactivity deposition than in the surrounding tissue.
Purpose: The examination is used to detect cavernous hemangiomas
in the liver and to distinguish them from other structures (such
as primary or metastatic tumors).
Radiopharmaceuticals: Autologous erythrocytes labeled with 99m
Tc - labeled either in vitro , but more often in
vivo using Sn-pyrophosphate (applied 20 min. before
application of 99m TcO 4 ).
Execution: Simultaneously with the iv application of the
radiopharmaceutical, dynamic scintigraphy of the perfusion phase
*) is started in a projection in which the best imaging of
suspicious deposits is assumed. Total dynam. shooting about 2
min., frame rate 2-3 s./frame. After 40-60 minutes, static
scintigraphy of the liver area is performed, preferably in SPECT
tomography .
*) Note:
Previously performed dynamic scintigraphy of the perfusion
phase in hemangiomas it is based on the fact that in the
hemangioma there is an increased blood pool at a relatively
slower blood flow compared to the vessels of the surrounding
tissue. However, monitoring of the perfusion phase has been shown
to have little clinical benefit and is therefore generally not
performed .
Evaluation: On static planar
scintigrams and on reconstructed tomographic SPECT sections, we
look for deposits of increased deposition of
labeled erythrocytes, which indicate the presence of hemangiomas.
In SPECT scintigraphy, the limit of detection of hemangiomas is
about 1 cm.
Distinguishing
hemangiomas from other units suspected of malignancy is important
for primary tumor diagnosis. Biopsies should not be
performed on hemangiomas as high blood vessels,
as there is a risk of bleeding .
Pancreas
scintigraphy
The pancreas is a small but
metabolically and endocrine important organ located in the
abdominal cavity in the duodenum , just below the liver.
The exocrine component, which opens into the
duodenum, produces digestive enzymes -
pancreatic lipase for the breakdown of fats, alpha amylase for
the breakdown of starch, proteases for the breakdown of proteins.
Under normal circumstances, digestive enzymes are inactive
form after their formation inside the pancreas (otherwise they would damage - "digest" -
pancreatic tissue, pancreatitis would occur ) , only when they reach the duodenum are they
activated and can begin to perform their digestive function.
Endocrine part (whose
cells are arranged in the islets of Langerhans) produces pancreatic hormones - insulin
(regulates blood sugar levels), glucagon , somatostatin
, pancreatic polypeptide.
Pancreatic Pathology
The most common disease associated with pancreatic is diabetes
- diabetes mellitus caused by tissue damage islets of
Langerhans, where insulin is formed. Inflammation of the
pancreas, pancreatitis , is caused by the retention
of digestive enzymes in the pancreas, which remain inside,
are prematurely activated and "self-digest" damage the
pancreatic tissue, causing swelling and an inflammatory reaction.
Acute pancreatitis is caused by sudden obstruction
of the pancreatic duct to the duodenum, usually by bile
stones. Chronic pancreatitis, caused by a slow
outflow of pancreatic enzymes into the duodenum, has a milder and
longer-lasting course. In more severe cases of necrotizing
hemorhagic pancreatitis , proteolytic pancreatic enzymes can
enter the bloodstream and cause toxic effects in various tissues
and organs. A very serious disease is (adeno) carcinoma
of the pancreas, which often metastasizes the whole body
through the lymphatic system.
Dynamic scintigraphy of the
pancreas
The pancreas is difficult to
access for functional diagnosis. For scintigraphic
examination of the pancreas the radiopharmaceutical 75
Se-selenomethionine H 2 C-S- (CH 2 ) 2 CH (NH 2 ) COOH labeled with the radionuclide selenium 75
Se was developed (physical
properties in §1.4 Radionuclides ", passage" Se-75
") . The intake of amino acids
in the pancreas is a reflection The similarity between selenium
and sulfur is so close that the substitution of selenium
instead of sulfur in the methionine molecule leads to an
analogue that has all the metabolic properties of an amino acid,
including incorporation into proteins, and is therefore
efficiently taken up by the pancreas in
digestive enzyme production.
The first
attempts at radioisotope examination of the pancreas in the 1960s
with collimated probe detection ("blind"
measurement - a priori to no avail ...) and
then static scintigraphy with a motion gammagraph and camera
without computer acquisition, were able to assess only gross
pancreatic abnormalities. Valid scintigraphy of the pancreas
brought valid results :
Dynamic
scintigraphic examination of the pancreas can be useful
for early detection of pancreatic exocrine dysfunction, retention
or obstruction of the excretory tract - in diabetic patients,
pancreatitis, cancer.
Place
the gamma camera equipped with a ME or HE collimator, the
analyzer window set to a 264keV peak, slightly obliquely above
the liver and pancreas area. After iv application approx. 100kBq
/ kg 75 Se-selenomethionine
with the gamma camera are buying dynamic sequential
images the liver and pancreas at intervals of 5-10
minutes. for 60-120 minutes. Selenomethionine is also taken up
non-specifically in the liver (where
proteosynthesis also takes place) , so that the image of
the pancreas is often displayed in interfering background against
the radioactivity of the liver. To eliminate this disturbing
background, 99m Tc-colloid is
sometimes applied at the end of the examination (with
the patient's position unchanged) , which is specifically
taken up by the liver. The resulting scintigraphic images of the
liver are then subtracted from the 75 S-selenomethionine
images . By this gradual subtraction of images the image
of the liver is suitably suppressed and better separation and
visibility of the pancreas is achieved.
The regions
of interest (ROI) of suitable parts of the pancreas are
then marked on scintigraphic images , from which dynamic
curves of the time dependence of selenomethionine
u accumulation are formed . The curves are mostly evaluated
visually (but at our workplace we also developed a
program for their quantitative processing) . In
physiological cases, the curve after the initial rapid increase
reaches a peak after 20-30 min. from the
application, followed by a slower decline. Reduced
functionof the pancreas on the curve manifests itself in
a flat shape, with a later onset and slowing down of the rate of
increase. In this case, the pancreas is displayed less clearly on
scintigraphic images. Retention of digestive
enzymes within the pancreas is manifested by a later and slower
decline; with more severe pathology, the peak and decline do not
appear at all, the curve still has a slowly increasing trend.
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Dynamic scintigraphy of the pancreas.
Left: The area of interest (ROI) of the appropriate part of the pancreas is marked on the summary scintigraphic image. Middle: Typical normal and pathological curves of the time course of selenomuthionine uptake in the pancreas. Right: Different responses of the curves to cholecystokinin application. |
In such pathological cases, in about 60.-90.
minute iv applies pancreozymin , more commonly called cholecystokinin
(also stimulates gallbladder contraction)
, 1u / kg, which stimulates the secretion of pancreatic
enzymes (and possibly also secretin ,
which, among other things, potentiates the effect of
cholecystokine and also has trophic effects on the pancreas)
. It is analogous to the above cholekinetic stimulus in dynamic
cholescintigraphy, or the application of a diuretic (furosemide) in dynamic renal scintigraphy.
Depending on the response of the curve to this pancreatic
stimulator, it is possible to distinguish the decrease
in function, parenchymal pancreatic damage or obstruction .....
Pancreatic
scintigraphy was relatively infrequent, diagnostic yield
relatively low, with a significant percentage of indeterminate
findings. It was performed mainly in the 70s-80s. Pancreatic
cancer is now visualized by CT imaging and static PET / CT
scintigraphy.
Static scintigraphy
of the pancreas with 99m
Tc- interleukin-2 is also tested
for imaging of chronic inflammatory changes in type 1 autoimmune
diabetes, to identify patients with pancreatic inflammation. It
makes it possible to detect an increased incidence of activated T
cells - the degree of lymphocytic infiltration even in smaller
inflammatory processes of insulitis, in the early period for the
treatment of immunotherapies.
Scintigraphy of the spleen and dynamic
splenoportography
The spleen (Latin lien ,
Greek splén ) is a somewhat
"mysterious" organ located in the abdominal cavity,
near the stomach. Phylogenetically, the spleen probably developed
as an organ of hematopoiesis . However, it retains this
function only in the prenatal period (until about the 6th month
of fetal development), then it is taken over by the bone marrow.
After birth, the spleen functions only as a "filtering"
organ with a large number of macrophages, retaining
microorganisms from the blood and obsolete or damaged ("worn
out") blood cells - sequestration of erythrocytes .
It is also immune , produces antibodies and
immunocompetent cells, has phagocytic ability (RES system).
The weight of the spleen is
about 100-200g, gradually decreasing with age. In some diseases,
however, the spleen enlarges - spenomegaly . Mild
splenomegaly (weight up to 500g) can also occur during
infections. Moderate splenoagaly (500-1000g) may accompany acute
leukemia, malignant lymphoma, polycythemia and more. Rarely,
severe splenomegaly (weight> 1000 g) occurs, eg in chronic
myeloid leukemia, ...
Anatomically, the spleen
belongs to the reticuloendothelial (RES) and hematopoietic
system. The blood supply to the spleen takes place through the portal
circulation ( vena portae*), which is significantly
associated with the liver. Therefore, we have placed
scintigraphic methods related to the spleen in the context of
liver scintigraphy.
*) The portal vein drains
blood from the organs of the abdominal cavity - from the
intestines, lower esophagus, stomach, spleen, pancreas - to the
liver.
Pathology of the spleen and portal pathways
One of the pathologies of the spleen is the above-mentioned splenomegaly
. Portal hypertension is increased blood
pressure in the basin of the portal vein ( vena portae
). The portal vein block is most often intrahepatic due to liver
cirrhosis. Instead of the portal vein, the blood then flows
through the created "connectors" - portosystemic
short circuits - into the systemic circulation (into the
basin of the inferior vena cava). There is a development collateral
flow, overloading of the veins creates varices .
Blood from the digestive system bypasses the liver, so it is not
detoxified, which can lead to damage to some tissues (e g brain).
Increased destruction and sequestration of erythrocytes
in the spleen is manifested in hemolytic anemia (see " Half-life of erythrocytes and
localization of their destruction ") . This often leads to
splenomegaly. In this case, splenectomy is recommended
to normalize the blood count .
Static scintigraphy
of the spleen
Purpose: Imaging of the functional
tissue of the spleen to determine its shape, size and placement,
including possibly inhomogeneities. It can also be used to
visualize and mark the spleen for application before dynamic
splenoportography.
Radiopharmaceuticals: 99m Tc-labeled
autologous heat-damaged erythrocytes are used for selective
imaging of the spleen and are replicated to the patient.
Radiocolloids , mainly 99m Tc-sulfur-colloid, whose larger colloidal particles are
taken up in the reticuloedothelium , are used to image the
reticuloendothelial system of the spleen (+ liver) . spleen and
liver system.
Design: After application of about
100-200MBq of the above radio indicator, planar scintigraphic
images in the front, back, and left side projections are taken in
about 20-30 minutes. For a more detailed distinction of
pathological structures, we can add the display of SPECT.
Evaluation: In the pictures we assess the shape, size and placement
of the spleen; the size of the spleen can be estimated from the
dimensions of the scintigraphic image using empirical methods. By
observing the homogeneity of the distribution of the
radiopharmaceutical or the presence of focal changes, we can
infer abscesses, cysts, splenomegaly, hematomas or tumors of the
spleen.
Dynamic
splenoportography
Purpose: Examination of blood flow
through the portal stream and detection of portosystem
shunts. Splenoporography is one of the less frequent
scintigraphic examinations, now it is almost abandoned ...
Radiopharmaceuticals: 99m Tc pertechnetate.
Execution: The application of the
radioindicator is performed intrasplenically
with a thin needle in a small volume (up to 10 ml.) And fast
enough (bolus) so that the phase of the first flow is well
expressed. At the same time, we will launch the acquisition of
dynamic scintigraphy in the front projection - 60 images after 2
seconds, which captures the flow of the radio indicator through
the portal and system streams.
Evaluation: Visually evaluate
images capturing individual phases of passage: spleen ® v.lienalis ® v.portae ® liver ® systemic circulation.
Under normal circumstances, after itrasplenic application, the
radiolabel passes rapidly through the v.lienalis and v.portae into
the liver, where the flow slows down appropriately in the capillary
bed, then flows through the hepatic veins and inferior vena cava
into the heart and lungs, and then into the systemic circulation.
In the presence of shunts (connectors) of the
portal and systemic flow, part of the radioindicator passes out of the liver and reaches the heart prematurely . In addition to these portosystemic shorts, we can also
consider an event in a series of images. obstruction of the
v.lienalis or v.portae.
To quantify the dynamics of the flow of the
radioindicator through the portal and systemic streams, we mark
the relevant areas of interest: vena lienalis, liver, heart +
lungs, from which we create curves of the time course of the
passage of the radioindicator. From these curves we can quantify
the flow dynamics . For curves from the v.lienalis, liver,
and heart regions, the time of arrival of the
radiolabel, the time of maximum, the steepness (gradient) of
increase (the ascending section intersects the linear
function) and the half-life of the radiolabel (the
descending section interpolates the exponential function) are
determined.
The procedure for computer evaluation of dynamic
splenoportography is described in §3.16 " Dynamic
splenoportography " of
the book "OSTNUCLINE".
Evaluation of dynamic splenoportography | |
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Evaluation:
After intrasplenic application of 99m-Tc, we observe on
scintigrams a fast flow of the radioindicator through the
v.lienalis and v.portae to the liver, without obvious
portosystemic short circuits. After the usual slowing
down in the capillary bed of the liver, the radiolabel
flows out through the inferior vena cava into the heart
and lungs. Conclusion: Visual evaluation of sequential scintigrams as well as quantitative analysis of the curves indicate normal flow conditions in the portal stream, without obvious portosystemic shunts. |
Here are examples of the evaluation of normal á and significantly pathological â radionuclide splenoportografie.
Evaluation of dynamic splenoportography | |
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|
Evaluation:
After intrasplenic application of 99m-Tc, we observe on
scintigrams a fast flow of the radioindicator through the
v.lienalis to caudal shunts. The liver is practically
invisible. Long-term retention of the radioindicator in
the spleen. Through caudal protocaval shunts, the radio
indicator flows into the heart and lungs. Conclusion: Visual evaluation of sequential scintigrams and quantitative analysis of the curves shows significant caudal portosystemic shunts , with no apparent flow through the portal vein to the liver. |
Scintigraphy
esophageal and gastric
Esophagus used to swallow food (solid and liquid
phase) from the mouth to the stomach , where
there is a first stage digestion of food. Under a physiological
state, the swallowed bite is actively transported by the peristalsis
of the esophagus to the stomach, where it arrives in
about 7 seconds. The motility disturbances of the
esophagus may occur due to stenosis of the esophagus,
innervation disturbances, ........... food passage is then
decelerated and is irregular. A disorder of the lower esophageal
sphincter causes part of the food to return from the stomach back
to the esophagus - gastroesophageal reflux
develops .
..........
Dynamic esophageal
scintigraphy - swallowing act
Purpose: Assessment of esophageal
motility, its patency, course of swallowing and detection of the
presence and severity of gastro-oesophageal reflux disease.
Radiopharmaceuticals: 99m Tc Sn-colloid or 99m Tc-DTPA
Procedure: Orally administer about
50 MBq of radioindicator mixed with about 10 ml. water (or fruit
juice) and immediately (better in advance) we start a fast
dynamic scintigraphy sitting in the front projection - 120 images
after 0.5 sec. (captures the passage through the esophagus) and
then about 60 images after 30 sec. (captures gastric evacuation
or late reflux). During the examination we can possibly. perform
compression on the epigastrium or Valsava maneuver to provoke
gastroesophageal reflux.
Evaluation: First, we visually
observe images of the passage of the radioindicator through the
esophagus, its distribution in the stomach and then its gradual
evacuation to the intestinal tract. Under normal circumstances, a swallowed
bite is rapidly transported to the stomach with the help of
esophageal peristalsis, so that the passage of the radioindicator
through the esophagus must be sufficiently rapid
and smooth ,
without temporary or permanent retention. In various pathological
conditions such as achalasia, disorders of patency (narrowing of
the lumen of the esophagus - tumor, external oppression of the
esophagus, etc.), or disorders of esophageal innervation,
disorders after operations on the esophagus, the passage through
the esophagus slows down . In
scintigraphic images, we then see slowed down or uneven passage through the
esophagus, which may be accompanied by retention
of the
radioindicator in some parts of the esophagus. However, only more
pronounced abnormalities are seen in the scintigraphic images ; more detailed and sensitive analysis and
quantification of the esophageal passage is performed on the curves from individual parts of the esophagus
and on special mathematical constructions - transport
function
and condensed image .
A common
pathology is gastroesophageal reflux, when due to
insufficiency of the lower esophageal sphincter, part of
the gastric contents return to the esophagus, ie
abnormally oriented movement against physiological
direction of food passage. Regurgitation manifests itself
in the relevant scintigraphic images as the presence of a
radioactive deposit, especially in the lower third of the
esophagus (reflux can extend to higher esophageal levels - more
detailed and sensitive analysis of the presence and location of
reflux is performed on curves from individual parts of the
esophagus) . Reflux can occur either passively
(spontaneously, under native conditions), or it can be caused by increased
pressure in the stomach (appropriate compression of
the stomach area) - then it is active reflux.
Analysis and computer
evaluation of dynamic scintigraphy of esophageal swallowing
function and gastric evacuation is described in detail in §3.20
" Dynamic
scintigraphy of the esophagus and stomach " of the book "OSTNUCLINE".
Mathematical analysis and complex evaluation of dynamic esophageal scintigraphy - swallowing act | |
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|
Evaluation:
After oral administration of the radioindicator, we
observe in scintigraphic images first a rapid passage of
the upper and middle part of the esophagus, then a
somewhat slowed passage of the distal part of the
esophagus. Once the stomach is reached, most of the
radiotracer returns to the middle stage of the esophagus,
where it retains for about 20 seconds before proceeding
to the stomach. This abnormal movement of the swallowed
radio indicator is particularly evident in the condensed
image and the transport function. Conclusion: Visual evaluation of sequential images and quantitative analysis of radiolabicure passage curves indicate good patency and motility of the upper esophagus, while severe passage pathology with marked gastro-oesophageal reflux was observed in the middle and lower esophagus (probably related to cardiac incidence ). |
Dynamic scintigraphy of
evacuation of the stomach and small intestine
Purpose: Monitoring the rate of
evacuation of food from the stomach to the intestine, or. small
intestinal transport rates.
Radiopharmaceuticals: 99m Tc Sn-colloid or 99m Tc-DTPA, which we
use to describe solid or liquid food.
Execution: A patient sitting in
front of the camera is given a small bite of solid food
orally , marked approx. 50 MBq 99m Tc. We run dynamic scintigraphy, in which we scan the
stomach area for about 1.5 hours at a frequency of 1 frame per
minute. If we want to investigate the evacuation of liquid
food, dynamic scintigraphy of the stomach may follow
dynamic scintigraphy of the swallowing act of the esophagus
(described above). However, a solid diet is more representative
for assessing gastric evacuation. In addition, scintigraphic
examination of small bowel transport may follow,
in the form of a slow dynamic study or sequential still images;
it is taken for several hours as needed. It is advisable to take
a control image of the abdominal cavity the next day (after 24
hours).
Evaluation: On the images of the stomach, we mark the ROI, from
which we generate a curve of the time course of the activity.
This curve begins with a flat arm corresponding to a phase in
which the ingested diet does not leave the stomach (or leaves it
only very slowly). This is followed by a different rapid decrease
in activity, capturing the evacuation from the stomach to the
intestine. We evaluate half of the evacuation of the
stomach T 1/2 - the time required for the activity in the stomach to
decrease by half (the descending part of
the curve interpolates the exponential function, from the rate
coefficient of which we determine T 1/2 ) . Normal values ??of the
half-life of gastric evacuation in the case of solid food are in
the range of 60-90 min., In the case of liquid food approx. 30-40
min. If scintigraphy of the small intestine followed, we
determine in the sequential images the time since ingestion, for
which the activity first appears in the initial wide part of the
large intestine ( caecum ). It is the so-called oro-caecum
time , whose normal values are about 2-5 hours. As the
half-life of gastric evacuation also affects the rate of small
bowel transport, it must also be taken into account.
4.9.4 Nuclear cardiology
The heart (Latin cor
. , Greek cardia ) is a hollow
muscular organ that, with its regular contractions, functions as
a pump that drives blood circulation throughout
the body. This ensures the transfer of respiratory gases,
nutrients and metabolic waste products. Cardiology
deals with the structure, function and diseases of the heart .
The heart of higher organisms, especially mammals and humans,
consists of several anatomical and functional parts :
- Cardiac cavities and supply vessels
Non-oxygenated blood (passed through the organism) is supplied to
the heart through hollow veins - upper and lower, which
connect to the venous canal in front of the heart
. During the flow through the heart, the blood passes through 4
cavities , which are separated from each other by valves
, preventing the backflow of blood. Blood flows from the venous
canal into the right atrium . From there, it
enters the right ventricle through a tricuspid
valve . The right ventricle, with its contractions, expels blood
through the " crescent " valve into
the lungs - the main arteries of the pulmonary
circulation. As it passes through the lungs ,
the blood is oxygenated. Oxygenated blood flows from the lungs
through the pulmonary veins into the left atrium
and from there through the bicuspid valve (also called mitral
for resemblance to the shape of a bishop's miter) into the left
ventricle. With the contractions of the left ventricle,
blood is expelled through the aortic valves into the aorta
, whereby oxygenated blood enters the main arterial circulation -
it passes through individual tissues and organs, releases oxygen,
transports nutrients, receives metabolic products and returns to
the heart via venous system.
The "pumping" of
blood takes place by alternating the phases of systole and
diastole of the heart chamber. In systole , the
heart chamber contracts and blood flows from the heart chambers
into the arteries. During the relaxation phase - diastole
- the muscles of the ventricles weaken and the heart fills with
blood with passive pressure. Each systole expels about 70 ml from
the heart. blood (so-called stroke volume ). The amount
of water that the chamber pumps per minute is called
volum minute heart - cardiac output
.
- Heart valves
act as one-way valves that allow blood to flow in only
one direction, while closing in the opposite direction and
blocking flow. There are 4 valves in the heart: - A double-
valve (mitral) valve between the left atrium and the left
ventricle; - Tricuspid (tricuspid) valve between the
right atrium and the right ventricle; - Aortic valve at
the interface of the left ventricle and aorta; - Crescent pulmonary
valve in the right ventricle in the lung. In order for the valve
to function properly, a sufficiently large opening for blood flow
must be created when it is opened, and when it is closed, it must
fit snugly to prevent blood flow back. A common disorder is the insufficiency
of the valves, when part of the expelled blood returns - the
so-called regurgitation , during which the heart must
then pump it again. This reduces the pumping efficiency and the
heart is overloaded.
- Heart muscles
The driving element of pumping is the heart muscle - the
myocardium , which drives the heart's pumping
activity with its regular contractions. It is a transversely
striated, highly powerful muscle. They are made up of cardiac
cells by cardiomyocytes. The strongest heart
muscle is in the left ventricle, which must expel blood into the
great circulation under considerable pressure.
- The
vascular supply of the
heart
In order for the heart muscle to work, it needs oxygen and
nutrients. The vascular supply of the heart muscle with
oxygenated blood is provided by two coronary coronary
arteries emanating from the aorta. They branch into a
network of vessels that surrounds the myocardium and resembles a
wreath in shape.
When some sections of the coronary arteries are narrowed (mainly
due to atherosclerotic plaques or embolizations), the vascular
supply of the heart muscle is reduced - ischemic heart
disease . Ischemic necrosis - myocardial
infarction - occurs after 20-40 minutes with complete
closure of the artery, in which irreversible death of the heart
muscle occurs in the basin of a closed vessel.
- Control of cardiac activity
The contraction of the heart muscle is stimulated by electrical
impulses . The control of heart activity is largely autonomous
- electrical stimuli for myocardial contraction are generated and
conducted in the heart wall, in the cardiac conduction system
. The main source of excitement is the sinoarthritic node
- a cluster of cells in the wall of the right atrium near the
venous canal. This node is affected by the autonomic (vegetative)
nervous system from the cardioregulation center in the brainstem,
in the elongated spinal cord (hypothalamus). The signal is
divided into two Tawar arms in the interventricular
septum, right and left, which faces the myocardium and spreads
excitement along the walls of the ventricles. These electrical
excitation signals can be sensed using an ECG.
Cardiovascular
pathology
Ischemic heart disease consists of a narrowed lumen of
the coronary arteries of the myocardium due to atherosclerosis,
which results in impaired perfusion of the heart muscle.
Severe reduction in perfusion is manifested by angina
pectoris , complete closure leads to myocardial
infarction . ......
Defects of the
heart valves consist either of a narrowing (stenosis) or
of their insufficiency , especially of the mitral or
aortic valve. This leads to backflow - regurgitation ,
which reduces the efficiency of the heart's pumping function.
Valves can be affected as part of birth defects, but also as an
acquired disability in infectious endocarditis.
Heart rhythm
disorders , also called arrhythmias ,
can be caused by a disturbance in the production of electrical
arousal, or a disturbance in the propagation of arousal. More
severe arrhythmias are corrected using a pacemaker.
Disorders of
myocardial contractility - hypokinesia, akinesia,
asynchrony, dyskinesia (or aneurysm) .....
Intracardiac shunts
are openings - defects - in the heart wall (septum) between the
ventricles or atria. ......
In connection with the
above-outlined function of cardiac activity and its disorders, cardiological
diagnostics performs in three basic directions :
1. Acoustic diagnostics of cardiac echoes of
systolic-diastolic function using a stethoscope and diagnostics
of electrical activity of the heart using
electrocardiography ECG. They are the oldest cardiological
methods. They are now being approached by ultrasound
sonography .
2. Diagnosis of central hemodynamics
- measurement of blood flow through the heart cavities and large
vessels, detection of intracardiac shunts and heart valve
insufficiency, including assessment of their severity.
3. Diagnosis of myocardial perfusion
- ischemic heart disease, ischemic myocardial viability ....
Nuclear medicine can offer cardiology four
diagnostic circuits :
- Methods examining systolic-diastolic function
of the heart as "pumps" can demonstrate
overall and regional impairment of heart wall motility or
synchronization with electrical activity of the heart, determine
the overall "performance of the heart pump". It is a
equilibrium ECG-gated ventriculography and SPECT of the
myocardium.
- Examination of central hemodynamics -
blood flow through the heart cavities and large vessels. After
the application of the bolus of the
radioindicator, it is possible to monitor the dynamics of blood
flow through large vessels, filling of atria and ventricles,
including the detection of incacardiac shunts ,
to determine the cardiac output,
cardiopulmonary volume , flow times and other important
hemodynamic parameters.
- Examination of the regional blood flow of the
myocardium , at rest and under load, allows to diagnose ischemic
heart disease , its location and severity.
- Verification of myocardial viability
damaged by ischemia. It is important for planning revascularization
procedures (by-pass, angioplasty) - revascularization only
makes sense in the case of a viable myocardium (which is perhaps
only temporarily hibernated by ischemia), not in the
case of an already unviable (necrotic) myocardium.
Equilibrium
gated ventriculography
Purpose:
It is a dynamic scintigraphic method that provides comprehensive
information about the activity of the heart as a pump of
blood circulation. Changes in activity in individual cardiac
compartments - chambers and atria during their pulsation during
the cardiac cycle are displayed. Because the radioindicator is
evenly and stably "mixed" in the bloodstream, changes
in activity - and thus in the emitted radiation g - are directly
proportional to changes in the volume of the
ventricles and atria during pulsation. We can determine the
hemodynamic functional parameters of the heart chambers, display
the regional kinetics of the heart walls, determine the
regurgitation fraction of the left ventricle (along with radiocardiography).
The method is useful for assessing the impact of ischemic heart
disease or myocardial infarction, or cardiomyopathy, on
ventricular function. It is also used to determine the
cardiotoxicity of cytostatics in chemotherapy of cancer.
In continuous dynamic imaging,
the number of pulses accumulated during one cycle would be too
low to imaging the shapes and sizes of the ventricles and
determine their volume changes. Therefore, the ECG gating
technique is used : in addition to scintigraphic pulses, an
electrical ECG signal is also captured from the
camera , which appropriately controls (triggers, gates,
"gates") the acquisition process. Gating pulses are
derived from the R-wave of the ECG and synchronize
periodic storage of scintigraphic images in defined areas of
computer memory. Gradual addition of corresponding images from
individual cardiac cycles creates the resulting set of images,
which represents the phase dynamic scintigraphy of
one " representative " cardiac
cycle , created by synchronous summation of several
hundred continuous cycles (described in
detail in §4.4 "Gate Phase Scintigraphy ") . By computer evaluation
of this phase scintigraphy, we can then assess the pulsation
of the walls of the ventricles and atria and create volume
curves. during the cardiac cycle, from which we can
determine a number of quantitative parameters of
systolic-diastolic heart activity.
Radiopharmaceuticals:
A radioindicator should be used that is maintained at a
sufficiently long concentration and does not leak from the
bloodstream. They are 99 m Tc-labeled red blood cells , which can
be prepared in two ways: 1. Laboratory in vitro
from a blood sample, wherein the labeled erythrocytes 99 m Tc is reinjected
back to the patient. 2. In vivo , in which the
patient is first administered the dissolved tin salt and after
about 15 minutes the required 99mTc-pertechnetate activity is administered . Tin ions Sn 2+ will allow the
binding of technetium to red blood cells in the circulation.
Execution:
After application of approx. 400 MBq of radiopharmaceutical is
scanned by a camera detector, aimed at the heart area in the LAO
projection at an angle of about 35-50 °, so that the ventricular
septum is approximately perpendicular to the plane of the
detector. The patient has supplied ECG
(cardiomonitor) electrodes , the R-wave output
of which is connected to the camera's synchronization circuit.
Let's wait for the heart rate to stabilize. We save in the ECG-gated
mode so that the heart cycle is divided into about 16-32 images,
reserved in the memory of the acquisition computer. We record
about 500-800 heart cycles, eliminating cycles with premature or
delayed R-wave. We can perform the examination at rest and in
ergometric or pharmacological load.
Evaluation:
In the simplest case, the regional mobility of
the left ventricular wall can be assessed visually by means of cinematographic
projection of individual images of a representative cardiac cycle
in rapid succession (which makes the movement of the heart
sections visible - "heartbeat"). It can be assessed
semiquantitatively using the contour method, where the
evaluation program plots the contour of the chamber
in the end-diastole ED and end-systole ES into a single image .
On the mutual relation of these contours we can recognize
disorders - hypokinesia, akinesia, dyskinesia ( Fig.3.1.2 in " Radionuclide ventriculography
") . Furthermore, we can construct parametric
images distribution of a certain parameter in the organ.
The simplest is the heart stroke image (image of
heart rate volume - difference of ED-ES images) and paradox
image created by subtraction of ES-ED images, in which
the atria and pathologically dyskinetic areas of the ventricle
(which are not emptied in systole but enlarged) are
physiologically imaged. Furthermore, it is possible to compile an
image of the ejection fraction , resulting from the
image of heart rate by dividing by the image of the ventricular
end-diastole (ED-ES) / ED. The most accurate analysis of cardiac
cycle dynamics is provided by regional Fourier analysis
using sine and cosine functions with certain amplitudes with
phases in each pixel. We obtain two parametric images: the
amplitude image, each site of which is proportional to
the intensity of pulsation (local heart rate volume) and the
phase image , expressing the time-phase shift (delay) of
the onset of myocardial contraction at a given site compared to
the arrival of the ECG R-wave - Fig.3.1.3
, 3.1.4 in " Radionuclide ventriculography " .
On the ED and ES images of the
representative cycle, we mark (manually or with the help of
mathematical algorithms, including taking into account parametric
images) the areas of interest (ROI) of the left
ventricle, or right ventricle and area of tissue background. The
computer program creates a chamber volume
curve from which important hemodynamic
parameters are calculated: ejection fraction, heart
rate, cardiac output, end-diastolic and residual volume of the
ventricle, ejection and filling velocities of the ventricle - see
pictures below. These parameters are especially important for the
left ventricle; for the right ventricle, which has a less regular
shape, accurate determination is more difficult.
Mathematical analysis and computer evaluation of
radionuclide ventriculography is described in detail in §3.1
" Radionuclide
ventriculography " of the book
"OSTNUCLINE".
Mathematical analysis and complex evaluation of radionuclide ventriculography | |
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Evaluation:
We do not observe regional disorders of heart
wall motility on phase scintigraphic images of the
cardiac cycle, nor on Fourier images of phase and
amplitude. Conclusion:
Visual evaluation of images of individual phases of the
cardiac cycle and quantitative analysis of cardiac
dynamics indicate good global and local
contractility of the walls of the left ventricle. |
Here are examples of the evaluation of normal á heavily pathological â radionuclide ventriculography.
Mathematical analysis and complex evaluation of radionuclide ventriculography | |
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|
Evaluation:
The following regional disorders of left
ventricular wall motility are observed
in phase scintigraphic images of the cardiac cycle and in
Fourier images of phase and amplitude : segment
hypokinesia: Virtually
all except posterolat. Conclusion:
Visual evaluation of images of individual phases of the
cardiac cycle and quantitative analysis of cardiac
dynamics indicate a severe disorder of
left ventricular wall contractility with extensive
hemodynamically significant apical dyskinesia. Extremely
reduced ejection fraction of dilated LV. |
Dynamic
Bolus Angiocardiography
Purpose:
Rapid dynamic scintigraphy of the passage and dilution of a
radioactive bolus through the right heart, lungs, and left heart,
which provides information about the dynamics of blood
flow in cardiac sections and large vessels. By analyzing
this dynamics, we can obtain quantitative parameters of chamber
function, their volume parameters, detection and quantification
of intracardiac short circuits . Together with
radionuclide ventriculography, the regurgitation fraction of
the left ventricle can be determined .
Radiopharmaceuticals:
Ordinary 99m Tc-pertechnetate or better 99mTc-DTPA can be used for dynamic angiocardiography alone.
If gated ventriculography is subsequently performed (eg to
quantify regurgitation), 99m Tc-labeled erythrocytes must be used as described above
for ventriculography.
Embodiment:
Application of tracer (about 400 to 800 MBq) are done under a
camera directed at the area of the heart and lungs in the right
oblique slope detector cameras 30-45 ° (in
this projection optimally separates and distinguishes images
chambers, atrial, pulmonary artery, and aorta) . It is applied to the shortest possible vascular
distance to the heart - to the anecubital vein of the right hand,
to the subclavian vein or to the internal jugularis. Since this
is a dynamic scintigraphy of fast action, it is necessary to
perform a so-called bolus application (lat. bolus = bite ): fast single
application of a radioindicator with high volume activity in a
small volume of approx. 0.5 ml., in a short time approx. 1 s. (with rinsing of several ml of physiological solution
using a three-way valve) , with immediate
start of dynamic scintigraphy with a sufficiently high frame rate
(approx. 4 frames / sec.) . We scan the fast phase for approx. 60-100 sec., Then a
slower sensing of the equilibrium phase can follow (10 sec. / frame for approx. 5 min.) .
Evaluation:
By visual inspection of sequential images of the passage of the
bolus through the cardiac circulation, we can qualitatively
assess or. abnormalities, especially premature circulation and
recirculation, which could be caused by heartbeats. Then we mark
the regions of interest (ROI) and create curves
distribution of the radio indicator from the right and left
ventricles, right atrium, lungs and possibly aorta.
To detect and quantify
the LP shunt, we use the time course of radioactivity in
the lungs - pulmogram . Under normal
circumstances, on this pulmogram, in addition to the sharp peak
of the first flow, after about 30-50 sec. appears only a low
broad peak of systemic recirculation, caused by the return of the
radioindicator through the systemic circulation back to the
heart. However, if an LP short circuit is present, another premature
recirculation peak will appear soon after the peak of
the first flow(or in the case of a small short circuit, only the
extension of the descending arm of the curve), caused by
recirculation by shorting the passed blood from the left
ventricle to the right ventricle and then to the lungs. By
mathematically decomposing the pulmogram into the curve of the
first pass, system recirculation and short-circuit recirculation,
we can quantify the magnitude of the shunt using
the ratio of short-circuit flow and lung flow Q z / Q p (without shuntt is close to 0) or ratio of lung flow and
system flow Q p / Q s (without shunt is approaching 1)
- for details see " Bolus radiocardiography ", Figures 3.2.3 and 3.2.4. From the right and left ventricular curves, we can
determine the average transit time of the central circulation and
the cardiopulmonary blood volume. If a peak of premature
circulation is present on the aortic curve, the right-left
shunt can be detected and quantified by decomposition of the
curves .
If the sensed phase
equilibrium curve analysis of the left ventricle can be (a combination of the dilution and Stewart-Hasmiltonova
principle - see " Bolus radiokardiografie "
fig. 3.2.1 and 3.2.2 ) provide cardiac output. By analyzing
the curve from the right ventricle, the ejection fraction
of the right ventricle can in principle be determined . The combination
of bolus angiocardiography and gated ventriculography makes it
possible to quantify regurgitation in heart valves - to determine
the regurgitation fraction .
Mathematical analysis and computer evaluation of
radionuclide angiocardiography is described in detail in §3.2
"Bolus radiocardiography " of the book "OSTNUCLINE".
Mathematical analysis and complex evaluation of radionuclide angiocardiography | |
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|
Visual
assessment: After
intrajugular bolus injection of the radiolabel, the
unexposed cavities of the right heart are imaged,
followed by filling of the unexpanded lung and pulmonary
artery, which are emptied reasonably rapidly into the
normally configured cavities of the left heart and aorta.
During the passage of the bolus through the left heart,
we do not observe a premature occurrence of a
radioindicator in the right heart and lungs Conclusion: In the visual evaluation of sequential
scintigrams of the passage of the bolus through the
cardiac circulation, nor in the quantitative analysis of
circulation curves, we do not observe pathological
changes in central hemodynamics. |
Here is the final protocol for the evaluation of angiocardiography without a short circuit á (but with regurgitation) and the intermediate results in the evaluation of a patient with a marked left-right shunt â .
Mathematical analysis of LP shunt in the evaluation of radionuclide angiocardiography | |
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|
By decomposing the pulmogram curve into primary circulation, systemic recirculation and premature recirculation, we observe a hemodynamically significant left-right shunt . A comparison of the bolus passage curves through the right atrium and ventricle indicates a short circuit at the level of the ventricular septum. |
Myocardial
perfusion scintigraphy
Purpose:
It is a non-invasive method for the assessment of regional
myocardial perfusion and the effect of coronary stenosis for
blood supply to the heart muscle in the relevant basin - at rest
and during physical or pharmacological stress. It can be used to
detect ischemic heart disease, its location, extent and degree of
myocardial damage, assessment of myocardial cell viability. Radiopharmaceuticals:
The basic requirement is that the radioindicator is efficiently
taken up by myocardial cells at the first flow and fixed in them
(without redistribution) for the duration of scintigraphy. In
this case, the displayed distribution of the radiopharmaceutical
in the heart muscle is proportional to the "supply" -
the regional flow in the coronary artery. However, this
distribution also depends on the functional state of the
cardiomyocytes: the accumulation of the radiolabel does not occur
in areas where the heart cells are necrotic (or are replaced by
connective tissue after infarction). Thus, the distribution of
the radiolabel in the myocardium is proportional to the regional
blood flow through the myocardium and the viability of myocardial
cells.
Previously, mainly thallium, 201 Tl-
chloride , was used, as an analogue of potassium. It
enters myocytes via the cell membrane mostly through the active
process of the Na / K ATP (adenosine
triphosphate) system, partly also by
passive diffusion. .... The disadvantage of thallium is the low
energy of X and gamma radiation, which causes poorer resolution
and significant absorption in the tissue; also a considerable
radiation load (to which abundant Auger
electrons also contribute) . Approx. 100
MBq 201 Tl
is applied.
Note: In the near future, however, a
partial "renaissance" of thallium can be expected in
connection with the introduction of special semiconductor
CZK cameras in nuclear cardiology. These gamma cameras
have a higher detection efficiency for low-energy photon
radiation of about 73keV 201 Tl .
Currently, 99m Tc- labeled radioindicators
are mainly used for scintigraphy of myocardial perfusion and
assessment of its viability , which provide significantly higher
quality images at lower radiation exposure. They are mainly 99m
Tc-isonitriles - non-polar lipophilic complexes that
enter myocardial cells by passive transport and bind in their
cytoplasm or mitochondria. The radiopharmaceutical accumulates,
depending on the blood supply, in healthy viable cells, while in
cells damaged (eg due to ischemia) or even dead and replaced by
scar fibrous tissue, no accumulation occurs. The distribution of
the radioindicator at individual sites of the myocardium is then
proportional to the regional blood flow through the myocardium
and the viability of myocardial cells.
The concentration of these 99m Tc-isonitrile radiopharmaceuticals in myocytes remains
stable for several hours after iv administration and thus shows
the immediate perfusion situation of the
myocardium at the time of application, eg during exercise. The
most common perfusion radiopharmaceuticals are 99m
Tc-MIBI (2-methoxyisobutyl-isonitrile)
and 99m Tc-Tetrofosmin (diphosphine
complex) . Approx. 500-800 MBq 99m Tc MIBI or
tetrofosmin is applied.
PET
radiopharmaceuticals for perfusion myocardial scintigraphy
Myocardial scintigraphy using positron
emission tomography (PET) is performed relatively
infrequently. Firstly, for less widespread and expensive PET
instrumentation, but mainly due to the difficult
availability of suitable positron radionuclides and
radiopharmaceuticals (discussed above in
§4.8., Section " Radionuclides and
radiopharmaceuticals for PET
") .
The simplest PET radioindicator for
perfusion is "labeled water
", in which ordinary oxygen 16 O is replaced by a positron radionuclide 15 O - ie water H
2 15 O. After application, it passes
through free diffusion through capillaries and cell membranes, so
that the distribution of radioactivity, measured by PET, is
proportional to blood flow. However, due to the high
concentration of radioindicators in the bloodstream, myocardial
imaging is not very contrasting. Another perfusion radioindicator
is ammonia 13 NH 3 labeled with nitrogen-13. Thanks to the high first
extraction (80%) and linear uptake according to blood flow in the
myocardium, it provides quality images. PET scintigraphy with
these short-lived radionuclides 15 O (T 1/2 = 2min.) And 13 N (T 1/2 = 10min.) Is bound to centers with
a cyclotron and is used mainly for research purposes.
The short-term positron radionuclide rubidium
82
Rb is used somewhat more often
for scintigraphy of myocardial perfusion by the PET
method due to the fact that it can be obtained from 82 Sr / 82 Rb generator in the
workplace . It is applied in the form of chloride 82 RbCl, behaving as an
analogue of potassium (similar to thallium 201 Tl mentioned
above; compared to thallium, however, rubidium-82 provides better
quality scintigraphic images at lower radiation exposure) .
The most common PET
radiopharmaceutical fluoro-deoxyglucose 18
FDG although it is mainly used for tumor imaging, as it
is a glucose analogue, it is taken up in the myocardium depending
on perfusion, ischemia and viability (see
" Metabolic scintigraphy, myocardial viability
imaging " below) .
Execution:
The examination can be performed at rest or under load. At rest
conditions, the distribution of blood flow in the myocardium is
usually homogeneous, minor or moderate perfusion disorders do not
manifest. If the coronary stenosis is not greater than about 90%,
the resting flow through the myocardium is sufficient to ensure
normal myocardial metabolism; we get a normal perfusion
scintigram of the myocardium, despite possibly. presence of
ischemic heart disease. Disorder of myocardial perfusion, even
without exercise, can only be observed in severe coronary artery
disease or after myocardial infarction.
Diagnostic susceptibility to perfusion disorders only becomes
apparent during stress testing, when the demand
for oxygen supply to the heart tissue increases - for coronary
blood flow. Normal coronary arteries respond to this by
visodilation and a corresponding increase in coronary flow, which
is reflected in an increased concentration of radioindicator in
perfusion scintigraphy. However, pathologically narrowed coronary
vessels are not capable of this (if
possible, they are already dilated by compensatory mechanisms
even at rest) , the load has only a small
effect on their flow. In the basin of the coronary artery,
affected by hemodynamically significant stenosis, the load shows
relatively lower perfusion than in the surrounding parts of the
myocardium - on the scintigram this place appears as a perfusion
defect in the myocardium, or at least as a reduction in
radiopharmaceutical distribution.
Therefore, the application of the radiopharmaceutical is
performed under exercise, either physical (usually a bicycle
ergometer) or pharmacological - application of vasodilators
(dipyridamole, adenosine, dobutamine) ......
We perform our own
scintigraphy in about 10 minutes after application. Prior to the
introduction of SPECT tomographic scintigraphy, scintigraphic
scintigraphy was performed planarly (in LAO
30 °, 60 ° projection) , but SPECT
tomographic scintigraphy provides better differentiation of
individual parts of the myocardium (advantages
of tomographic scintigraphy over planar were discussed above in
§4.3 " Tomographic scintigraphy " ). The most commonly used
is a two-detector SPECT camera *), whose detectors are set to an
angle of 90 °. It is taken by approx. 32-64 projections at a
total angle of 180 ° around the patient - from the right front
oblique projection of RAO 45 ° to the left rear oblique
projection of LPO 45 °. An ECG-gated myocardial SPECT
is performed with R-wave synchronization of the
electrocardiogram, analogously as described above for gated
ventriculography, we read about 500 cycles.
*) Single-purpose special types of cardiological cameras
optimized for myocardial scintigraphy have also been developed .
At some workplaces, perspective semiconductor CZK cameras
are beginning to be used for SPECT myocardium (§4.2., Part " New and alternative physical and technical
principles of scintillation cameras
", passage "Semiconductor multidetector
cameras ", Fig.4.2.10 on
the right). When using them, it is
sufficient to apply less than half of the usual activity, while
reducing the examination time.
If
stress perfusion scintigraphy is normal, resting scintigraphy is
no longer necessary.
Evaluation:
On coronary tomographic images , sagittal and transverse
sections, the myocardium is displayed *) in the form of more or
less closed "rings" or "rolls" or
"horseshoes", on which we can visually
evaluate the distribution of radioactivity - assess the size,
number and location of perfusion defects and
hypoperfusions.
* ) In the pictures we can clearly see only the thicker wall of
the left ventricle, not the thinner wall of the right ventricle.
For the semiquantitative
evaluation is often used structure so-called polar maps
(slang " bull's eye ") : coronal slices perpendicular to the short axis of the
left ventricle and the noise are transformed to each other so
that they form concentric circular profile with a tip in the
center. The brightness (or color) of the resulting circular area
is modulated by the different concentration of the radio
indicator shown in the summation of the individual layers. These
profiles are stored as concentric rings in a new circular image.
This gives a clear normalized polar map of the distribution of
activity in the myocardial wall - a map of regional
perfusion of the myocardium (in
polar coordinates centered in the apex) ,
which is divided into segments corresponding to
the basin of individual coronary arteries(according
to international recommendations, 17 segments are used) . The relative perfusion values in the individual
segments are compared with the corresponding values in normal
patients stored in the normal database . This
results in relative indices - the so-called scores
, which help determine the severity of the myocardial perfusion
disorder (and possibly the corresponding
risk) .
Visual images and polar maps are evaluated in
scintigraphy under load and at rest
. In the case of ECG-gated SPECT myocardium, in addition to
perfusion, we can also determine the ejection fraction of
the left ventricle.
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Evaluation of exercise
+ resting scintigraphy of myocardial perfusion with 4DM SPECT program. We observe a significant reduction in perfusion in the basal segment of the lower wall, partially reversible (results from a comparison of polar maps at rest and load). (scintigraphic
images were taken by |
Metabolic scintigraphy,
imaging of myocardial viability
Perfusion scintigraphy alone (showing hypoperfusion sites), or
analysis of general and regional contractile systolic function
(or dysfunction), may not be able to assess the maintenance of
myocardial tissue viability, which may be temporarily hibernated
due to hypoperfusion. To assess the function and condition of the
myocardial tissue itself , myocardial cells, it
may be useful to visualize metabolism glucose,
fatty acids or amino acids. Appropriate biochemically active
compounds labeled with radionuclides, especially positron
isotopes of biogenic elements, are used for this purpose. These
imaging methods have the ability to independently (separately or
simultaneously) assess blood flow and metabolism. Areas of
myocardium with reduced flow but preserved metabolism can thus be
recognized - revascularization may be useful here .
In clinical practice, PET
examination of myocardial glucose metabolism
using 18 F-FDG (fluorodexyglucose) is used.
Under normal circumstances of a well-perfused viable myocardium,
myocytes gain the necessary energy mainly by beta-oxidation of fatty
acids. However, in ischemia, anerobic glycolysis is becoming
more important as a source of energy for the work of cardiac
cells. After iv administration, 18 F-FDG accumulates to an increased extent in the
ischemic, but viable, areas of the myocardium, while it does not
accumulate in the normally perfused areas (because glucose is not
used to obtain energy there). FDG also does not accumulate in the
non-viable or necrotic myocardium because metabolism does not
take place there at all. By comparing between regional blood flow
and metabolism, we get information about normal, hibernating and
necrotic myocardium. In the case of ischemic defects with preserved
viability, it is useful to perform revascularization
surgery, after which there is a "revival" of
these sites and often an improvement in myocardial contractility
and an increase in overall cardiac performance.
In particular for research and
experimental studies, imaging of the distribution of free labeled
fatty acids such as 123 I- labeled BMIPP penta- and hexadecanoic acid
derivatives for SPECT is used to assess metabolism . For PET, it
is 11
C-palmitate for fatty acid metabolism and 11 C-acetate (which is incorporated
into the Krebs cycle) for oxidative
metabolism. ........
Receptor
scintigraphy of the myocardium
The spread of excitations in the myocardium and the regulation of
coronary flow are determined by the function of the sympathetic
nervous system in the heart. This function depends, among other
things, on the distribution of receptors ,
especially noradrenaline receptors. The distribution of these
receptors can be visualized by 123
I -labeled MIBG , which binds
to them. ...........
4.9.5. Pulmonary scintigraphy
(nuclear pneumology)
The lungs (Latin pulmo
, Greek pneumo ) are a key
respiratory organ in higher animals and humans. They exchange
gases - especially oxygen and carbon dioxide *), between
blood and air. The chain [ventilation of the pulmonary alveoli ® diffusion of
gases through the alveolocapillary membranes ® perfusion of the
lungs ® blood circulation] transports oxygen
from the air to the cells of tissues and organs and removes
carbon dioxide from the tissues into the atmosphere.
*) The main source of energy for cells
during metabolism is oxidation(especially
glucose), which produces "energetic" molecules such as
ATP (adenosine triphosphate), the "waste" products are
mainly water and carbon dioxide. Thus, gas exchange is required
for the metabolism to function - oxygen supply and carbon dioxide
removal. In higher organisms, this gas exchange is not sufficient
by passive diffusion, but takes place by active
respiration (respiration, ventilation) of air from the
atmosphere through the lungs (in fish by oxygen exchange from the
water in the gills).
The lungs have a spongy
consistency, they consist of more than 300 million alveoli
- alveoli , which are small hollow thin-walled
sacs (about 150 m m in diameter ). Their wall - the alveolocapillary
membrane - is formed by one layer of thin cells,type
I pneumocytes (wall thickness is about 1 m m). The total area
of ??the alveolocapillary membrane is about 60m 2 . There are also
thicker type II pneumocytes in the alveolar wall , which
produce substances that reduce surface tension (surfactant) and macrophages
, which phagocytose dust and foreign particles.
Respiratory gases diffuse through the
membrane of the alveoli in the direction of pressure and
concentration gradients; depends on the partial pressure of these
gases in the inhaled air and in the non-oxygenated blood flowing
in the capillaries around the alveoli. Oxygen has a lower partial
pressure in deoxygenated blood and therefore passes through the
membrane from the alveoli to the blood. Carbon dioxide, on the
other hand, has a higher partial pressure in the venous blood and
therefore passes from the capillaries through the membrane into
the air in the alveoli, from where it is exhaled.
Anatomically, the lungs are a pair of
organs - the left and right lungs. They are divided into lobes
(3 lobes have the right lung, 2 lobes have the left). The lobes
are further divided into bronchopulmonary segments ,
each of which has its own air and blood supply.
Air
is fed into the lungs (and discharged) bronchi (
bronchi ), which are the cartilage tube walls that are
inside the lungs manifold branches into increasingly finer tube
until the alveoli. Airflow in the lungs - respiration
- takes place alternately cast ( inspirium ) and
exhalation (expirium). When inhaling, the contraction of the
intercostal muscles and the diaphragm increases the volume of the
thoracic cavity, and new air is drawn into the lungs by the
airways due to the negative pressure. During exhalation, the air
used is blown out of the airways into the atmosphere by the
passive pressure when the chest is contracted.
Blood
circulation of the lungs - pulmonary perfusion - starts the
pulmonary artery ( pulmonary artery
), an artery emanating from the right
ventricle and delivering non-oxygenated blood. In the lungs, they
branch many times except for the capillaries
that surround the alveoli. Here, oxygen diffuses into the blood
and carbon dioxide into the alveoli. The vessels carrying the
oxygenated blood connect in the pulmonary veins, which open into
the left atrium. From there, the left ventricle pushes oxygenated
blood through the aorta into the bloodstream, distributing it
throughout the body. Deoxygenated blood, which is also enriched
with carbon dioxide, then leads through the venous system to the
right atrium, from where the right ventricle returns it to the
lungs for further oxygenation and CO 2 depletion - small cardiac circulation ( pulmonary,
cardiopulmonary ), continuing to the left atrium. Oxygenated
blood is then pumped through the left ventriclelarge blood
circulation - systemic . Normally, almost the same amount of
blood flows through the pulmonary circulation as the systemic
circulation, but under lower pressure.
Pathology of the lungs
and respiratory system:
Pulmonary edema (
"emphysema"), as a result of pulmonary hypertzenze,
elevated pressure in the pulmonary circulation, left ventricular
failure, .........
Pulmonary embolism is embolism pulmonary arteries due to
thrombosis of the venous system or the right heart. It leads to hypoperfusion
or aperfusion of certain parts of the lungs. Sudden obstruction
of the middle lung branches is called a pulmonary infarction.
Bronchial asthma - spasm and bronchial constriction
associated with dyspnea ........
Pneumoconiosis is a
ventilation disorder caused by fibrosis due to prolonged
inhalation of dust, such as silica ( silicosis ),
popularly called "dusting of the lungs". .......
Inflammatory and infectious diseases , tuberculosis
...........
Tumors - primary lung cancer, metastases of other lung
tumors ....
Three
basic conditions must be met for the order functioning of the
respiratory system: 1. Good pulmonary perfusion;
2. Good ventilation of the pulmonary alveoli; 3.
Proper function of the alveolar membrane. Methods of nuclear
medicine also focus on the diagnosis of these components of the
respiratory system.
Pulmonary perfusion
scintigraphy
Purpose:
Imaging the distribution of capillary perfusion in the lung
parenchyma, revealing regional perfusion defects due to
embolization or other involvement of the pulmonary arterial bed
(compression by inflammatory or tumor foci, pleural effusion,
.....).
Radiopharmaceuticals:
99mTc-MAA (labeled macroaggregate of
albumin, or albumin microspheres).
Execution:
After iv application of approx. 200 MBq of radioindicator, static
scintigraphy of the lung area is performed in 4-6
projections, the most important of which are PA and AP
projections.
Evaluation:
On the images we visually evaluate the homogeneity of the
distribution of the radio indicator in the lung wings and
possibly local or segmental defects that would
indicate hypoperfusion due mainly to embolization
. Regional quantification of relative perfusion
of individual parts and segments of the lung can also be
performed .
Pulmonary ventilation
scintigraphy
Purpose:
To visualize the distribution of alveolar ventilation of the lung
parenchyma, to reveal regional ventilation defects for
peripheral airway patency - due to conjugation or other
disorders. It can also be used to assess the contribution of the
function of individual parts of the lungs to the overall
respiratory function (when deciding on surgery).
Radiopharmaceuticals:
- Inert radioactive gases: krypton 81m
Kr , xenon 133 Xe;
- Radioactive aerosols marked 99m Tc.
Procedure:
Scintigraphic examination of pulmonary ventilation can be
performed in two ways :
- Inhalation of
radioactive aerosol
Before scintigraphic examination of pulmonary ventilation, the
patient breathes for about 10 minutes. air with aerosol 99m
Tc-DTPA (activity in the nebulizer
approx. 1000 MBq) , the particles of which
are trapped in the alveoli. Then we store static images of the
lungs in individual projections under the camera.
- Inhalation of radioactive gas
Ventilation scintigraphy of the lungs is more preferably
performed by inhalation of radioactive inert gas - krypton
81m
Kr (activity in the generator
approx. 5 GBq) or xenon 133 Xe (if appropriate respiratory equipment
is available , see below) , with simultaneous scintigraphic
scanning.
81m Kr is obtained from
the generator 81 Rb / 81m Kr . The principle of this generator
is in the left part of Fig. .... The parent rubidium 81 Rb is fixed in the
solid phase in a small column, through which a stream of elution
air is passed by means of a fan (air
pump with adjustable power) . Through the radioactive decay of
rubidium-81, the continuously released daughter gas krypton 81m Kr is entrained by
the passing air and led into a breathing mask ,
from which the patient inhales a mixture of air and radioactive 81m Kr. One -way
valves are included in the circumference of the breathing
mask, further upstream of the outside air mixing valve
to ensure free breathing. The exhaled air is led into an extinction
vessel (volume approx. 30 liters) , from which, due to the very short half-life of 81 m Kr, practically
non-radioactive air emerges.
During this examination of pulmonary
ventilation , inhaled air with a trace content of
radioactive 81 m Kr enters the pulmonary alveoli , while the
emitted radiation of 191keV gamma is scanned by a gamma
camera . The scintigraphic image of the site of reduced
activity shows areas of the lung with impaired
ventilation , where krypton-81m, and thus no air, does not get
(either at all or reduced).
Fig .... Generator 81 Rb / 81m Kr for
scintigraphy of pulmonary ventilation.
Left: Principle of generator operation. Middle:
One of the design arrangements of the Rb-Kr generator. Right:
Disintegration scheme 81 Rb and 81m Kr; in
the black field is the scintillation spectrum of gamma radiation 81m
Kr ..
Evaluation:
On the images we visually evaluate the homogeneity of the
distribution of the radio indicator in the lung wings and
possibly defects that would indicate ventilation
disorders due to eg conjugation . Regional
quantification of relative ventilation of individual
parts and segments of the lungs can also be performed .
Dynamic ventilatory
scintigraphy of the lungs
is a relatively complex method for the analysis of
respiratory function of the lungs . The patient's
breathing is connected to a closed circuit spirometer
, into which approximately 300 MBq 133 Xe is applied and at the same time dynamic lung
scintigraphy is scanned in the PA projection.
A more detailed description of
the design and evaluation of dynamic ventilation scintigraphy is
in §3.11b " Dynamic lung scintigraphy (133-Xenon
ventilation) " of the
OSTNUCLINE book. Due to the considerable complexity and
instrumental complexity of the method, as well as the difficult
availability of 133 Xe, dynamic ventilation scintigraphy is practically no
longer performed.
Combined perfusion +
ventilatory pulmonary scintigraphy
For a more comprehensive assessment and differential diagnosis of
pulmonary pathologies, it is useful to compare perfusion and
ventilatory static scintigraphy and to correlate the
corresponding images in individual projections. In addition to
separate pulmonary perfusion (with 99m Tc MAA) and separate pulmonary ventilation (with 99m Tc-aerosol, or with
gaseous 81m
Kr), simultaneous combined perfusion + ventilation
scintigraphy is often performed - application of 99m Tc-MAA and then
alternating inhalation of 81m Kr , with simultaneous scintigraphy (with the analyzer window set alternately to 140keV 99m Tc and 190keV 81m Kr). All these scintigraphies are performed in a number of
different projections - the basic ones are AP and PA, as well as
oblique LPO, RPO, sometimes even lateral LL, RR. The resulting
scintigraphic studies then have 4, 6, 8, or 12 images.
Computer evaluation of multistatic scintigraphy of combined pulmonary perfusion and ventilation | |
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Visual
evaluation: When evaluating scintigraphic images of lung ventilation in all projections, we observe a homogeneous distribution of the radioindicator in both lung wings - without a defect . In scintigraphic images of lung perfusion, we observe an inhomogeneous distribution of the radioindicator in both lung wings with several areas of hypoperfusion , without segmental defects Conclusion: |
The procedure for computer evaluation of perfusion and ventilation lung scintigraphy is described in §3.11a " Static lung scintigraphy (perfusion, ventilation) " of the "OSTNUCLINE" book.
4.9.6 Scintigraphic diagnostics in oncology.
Scintigraphy of inflammation.
Views tumor tissue is one of the most important
methods of scintigraphic diagnostis. It serves both for primary
tumor diagnosis (in combination with other imaging methods),
helps in the planning of radiotherapy, allows
you to assess the response to treatment and the
success of therapy, during longer-term follow-up helps to
determine in time or. recurrence of cancer.
These individual aspects are discussed in §3.6 " Radiotherapy ", section " Diagnosis
of cancer ". In
scintigraphic tumor diagnosis we can use the following options :
- Imaging the tumor as a defect in functional
tissue
after application of radiopharmaceuticals that accumulate
physiologically in healthy functional tissue but are little or
not absorbed in the tumor tissue. It then appears on the
scintigram as a "cold lesion" (photopenic lesions). It
is a non-specific method ( it is not
possible to decide whether the defect is caused by a malignant or
benign structure) and is relatively insensitive(Cold
lesions are shown with less contrast and are less difficult to
see in images than "hot" deposits). It can be used
secondarily in liver or kidney scintigraphy. In lung
scintigraphy, we can observe local reductions or outages of
perfusion or ventilation due to oppression of blood vessels and
bronchi by lung tumors and metastases. In thyroid scintigraphy,
the functional parenchyma absorbs radioisotopes of iodine and 99m Technecistan well,
while malignant tumors and benign dysfunctional adenomas appear
as "cold" focal defects in the background of the
parenchyma (see scintigraphic images in
§4.9.1 " Thyrological diagnostics " ).
- Scintigraphy of tumor metabolic
activity
depicting altered (usually increased) tumor cell metabolism, or
tumor-induced tissue metabolism in the immediate vicinity of
tumors. PET imaging of the distribution of 18
F-fluoro-deoxyglucose (FDG), or 18 F-3-fluoro-3-deoxy-thymidine (FLT), or 18 F-fluoroquinoline is
currently most commonly used for gamma-visual imaging of tumor
metabolic activity . FDG is taken up in viable tumor cells and,
unlike glucose, does not exit the cells, thus providing contrasting
images of viable and proliferating tumor lesions .
Semi-quantitative assessment of the metabolic activity
of the depicted lesions is often performed using the SUV
value (defined and discussed above in
§4.1, section " Scintigraphic
image quality - detectability of lesions
", passage " Quantification
of positive lesions on gammagraphic images - SUV ") . PET scintigraphy should
be performed in hybrid combination with X-ray CT imaging
to obtain anatomical images in correlation with scintigraphically
imaged tumors. However, the increased accumulation of FDG is not
only specific for tumors, it can also be in inflammatory foci (see " Scintigraphy of inflammation
" below) .
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Example
of PET / CT scintigraphy with 18 FDG in a patient with lymphoma. ( PET / CT images were
taken by |
. -
Imaging of
non-specific binding of the radiolabel to tumors
Some radiopharmaceuticals show increased uptake in tumors and
metastases, and the mechanisms of these uptake are not well known
and are not specific to the tumor species. Increased
proliferation and metabolic activity of tumor cells or increased
permeability of tumor neoangiogenesis capillaries may contribute
to higher accumulation. Such a radiopharmaceutical is 67 Ga-citrate, which
appears to increase uptake into viable tumor tissues of
lymphomas, melanomas, lung cancer and the detection of their
metastases. However, it also accumulates in inflammatory lesions (see " Scintigraphy of inflammatory
deposits " below) . A
very important method of cancer diagnosis is skeletal
scintigraphy (see §4.9.7 below ) using osteotropic
radiopharmaceuticals, which allows very early detection of
skeletal metastases.
- Scintigraphy of special tumor
receptors
Some tumors contain specific receptors on cell membranes or in
the cytoplasm, mainly for peptides. Using these radiolabeled
peptides, the concentration of their receptors and thus the
respective tumor site can be scintigraphically imaged. It mainly
uses imaging of somatostatin receptors in
neuroendrocrine tumors, lung carcinomas, medullary thyroid
carcinoma, carcinoid, ... The most commonly used
radifarmaceuticals 111 In-pentetreoid and 99mTc-depreotide. Estrogens (such as estradiol) labeled
with 18 F
or 123 I
are rarely used for the diagnosis of hormone-dependent breast
cancer .
- Immunoscintigraphy
is based on the highly specific nature of antigen-antibody
immunological reactions. A monoclonal antibody (§3.6, passage " Monoclonal antibodies ") , labeled with a suitable
radionuclide, binds selectively to the appropriate tumor
marker (antigen) after application , after which the
relevant tumor can be imaged by scintigraphy. ........
scintigraph. pictures ........ ................ In more detail
" Diagnosis of cancer " .....
Scintigraphic diagnostics can
be useful in monitoring the function of healthy
tissues and organs in connection with oncological treatment.
Radiotherapy not only affects the target tumor tissue, but can
adversely affect organs located in the irradiated area - radiotoxicity
. Systemic chemotherapy has side effects on many tissues and
organs - chemotoxicity . For the uncomplicated course of
oncological treatment, it is useful to monitor the function of
critical tissues and organs during and after therapy.
Scintigraphic diagnostics such as radionuclide ventriculography , myocardial perfusion , and dynamic
renal scintigraphy may also contribute
to this .
Scintigraphy
of inflammatory foci
In addition to tumor lesions, scintigraphy is also used in the
diagnosis of inflammatory foci . Active
inflammation is accompanied by hyperemia, increased metabolic
turnover, increased cellular fluid, and migration of leukocytes
to the site of inflammation. Several procedures are used for
scintigraphy of inflammatory deposits :
- Scintigraphy
with 67 Ga-citrate, which first binds to
circulating transferrin in plasma after iv administration. Due to
hyperemia and increased capillary permeability, gallium bound to
the transport protein penetrates into the extravascular space in
the inflammatory locus and is then bound to lactoferrin and taken
up intracellularly in the lysosomes of neutrophilic leukocytes.
However, gallium citrate is also taken up in some tumors (as
mentioned above).
- Scintigraphy
after administration of 99m Tc-exametazime, HMPAO, or 111 In-oxine- labeled autologous leukocytes in
vitro , which then migrate and increase in concentration
sites of inflammation.
- Scintigraphy
using in vivo labeled antigranulocytes
monoclonal antibodies such as 99m Tc-sulesomab and 99m Tc-besilesomab that bind to leukocytes. Leukocytes,
including those so labeled, are increasingly concentrated at the
site of ongoing inflammation. These monoclonal antibodies of
murine origin may be at risk of undesired immunogenicity
upon repeated administration , leading to false negative results (immune anaphylactic reactions are unlikely in
diagnostic use due to small trace amounts of the substance
administered, in contrast to therapeutic applications) - discussed in §3.6, passage " Monoclonal antibodies ").
![]() |
Scintigram of a patient with left
knee TEP inflammation using the 99mTc-besilesomab radioindicator. (scintgraphic images were
taken by |
- Scintigraphy using labeled somatostatin
analogues such as 111 In-pentetreoid and 99m Tc-depreotide. They bind to structures containing
somatostatin receptors - in inflammation they are lymphocytes and
macrophages. Otherwise, of course, they also accumulate in some
tumors (especially neuroendocrine).
- PET
scintigraphy after application of 18 F-fluorodeoxyglucose (FDG), which
accumulates in inflammation due to the increased glucose
metabolic turnover that accompanies inflammation. However, FDG is
also taken up in viable tumors - which is the main use of FDG.
- 3-phase
scintigraphy of the skeleton in musculoskeletal
inflammation, showing, inter alia, regional hyperemia
accompanying osteomyelitis - infectious purulent
inflammation of the bone and bone marrow (see
the section " Dynamic (3-phase) scintigraphy of the
skeleton " below in §4.9.7) .
4.9.7. Skeleton scintigraphy
The skeletal system - the skeleton - serves
primarily as a mechanical support and reinforcement of the
organism. In conjunction with the muscles that are
attached to the bones by tendons, it also mediates the body's locomotor
activity . In addition, some bones in their cavities - bone
marrow - provide space for hematopoietic tissues . Bone
is composed of cells of osteoblasts , osteocytes
, osteoclasts , fibrous and amorphous intercellular
mass. Osteoblasts
are cells of roughly cubic shape with numerous protrusions
through which they are in contact with each other. It is realized
in osteoblast organelles
bone metabolism . Osteoblasts
produce collagen fibers and an amorphous (proteoglycan)
intercellular mass. They also produce alkaline phosphatase
enzymes, which cause bone mineralization . Osteoblasts
are present in bone especially where bone tissue is formed or
rebuilt. During development, osteoblasts gradually lose their
organelles and protrusions and turn into elongated osteocytes
.
Osteoclasts
are larger "breaking" cells that, by producing acid
phosphatase and collagenase, release bone minerals and disrupt
bone structure. This frees up space for new bone formation - they
help growth processes and bone remodeling.
Intercellular
bone mass it consists of bundles
of collagen fibers cemented by an amorphous proteoglycan mass,
which is mineralized in the finished bone tissue . The
mineral component consists of microscopic crystals of calcium
phosphate - hydroxyapatite , whose needles
are bound to collagen fibers. It is this mineralization
that contributes to bone strength and hardness.
Overall, bone consists of an average of
60% minerals, 28% organic matter (of which about 4% fat) and 12%
water. The reduced content of the mineral component leads to
porous and less solid bone - the so-called osteoporosis
.
Skeletal pathology
The most common disorders here are fractures -
arising from a mechanical force exceeding the mechanical strength
of the bone (in case of previous bone damage, eg osteoporosis,
the pathological fracture can be caused by a small
force).
Inflammatory bone diseases - osteomyelitis ....
Degenerative bone disease ....
Tumor bone changes
can be primary , arising from bone tissue - osteosarcoma
, chondrosarcoma , large cell osteoblastoma .
However, secondary tumors are more common - metastases of
other tumors to the skeleton; most often ca breast or prostate.
Scintigraphic
diagnostics of the skeleton focuses primarily on the detection of
pathological conditions related to changes in bone metabolism
. This is especially important in tumor diagnosis during early
detection or. metastatic infiltration into the skeleton and in
inflammatory and degenerative bone diseases.
Static scintigraphy of
the skeleton
Purpose:
This is a functional examination that shows the distribution and
changes in bone metabolism and the distribution of bone
reconstruction in order to reveal pathological processes of
tumor, inflammatory or degenerative.
Radiopharmaceuticals:
As osteotropic radiopharmaceuticals , labeled phosphate
complexes are used for single photon scintigraphy (planar,
SPECT) , which bind mainly to the surface of hydroxiapatite
crystals, partly also to calcium phosphate. The most commonly
used is 99m Tc-MDP (methylene diphosphonate), or 99m Tc-HDP
(hydroxymethylenediphosphonate - oxidronate). 18F-sodium
fluoride (NaF) is used for PET imaging , which reacts
with hydroxiapatite and is incorporated into the bone mineral
component in the form of fluoroapatite.
Execution:
After iv application of approx. 500-800 MBq 99m Tc-MDP, the radiopharmaceutical gradually accumulates
in metabolically functional bones, depending on regional blood
flow and osteoblastic activity of bone tissue. Scintigraphic
imaging is performed about 2-4 hours after application (when only about 5-3% of the administered activity
persists in the bloodstream, so we can find a sufficiently
contrasting image of the skeleton with a low body background) . When using 18 F-fluoride, about 100-300 MBq is applied and PET
imaging is performed as early as 1 hour after injection(fluoride has a faster absorption in the bones and a
faster breakdown of unbound radioindicator from the blood) . Prior to the examination, the patient is urinated so
that the accumulated activity in the bladder does not cross-image
the pelvic bones. Planar images of either individual parts of the
skeleton are taken in suitable projections - targeted
scintigraphy , or it is better to first obtain whole-body
planar scintigraphy in the projection of PA and AP with
a sliding movement of the bed with the patient under the camera.
Planar scintigraphy can be supplemented by emission tomography of
SPECT "ingested" sites. When using 18 F-fluoride, of
course, PET scanning is performed. For a more accurate assignment
of displayed pathological lesions to anatomical structures, it is
appropriate to perform X-ray CT imaging with fusion
SPECT / CT or PET / CT images on hybrid instruments; it then allows to specify the spatial location of active
deposits, whether the lesion is really in the skeleton and how.
extends beyond the bone .
Evaluation:
The displayed distribution of the radiopharmaceutical in the
skeleton is not uniform, but is given by regional vascularization
and the intensity of bone osteogenesis. In places with more
intensive metabolism, there is a higher accumulation of
radiopharmaceuticals. Physiologically, these are sites with
growing bone in children, bone regeneration after fracture,
.......
Pathological changes in the skeleton are usually manifested by
locally increased accumulation of radiopharmaceuticals - " hot
deposits " osteoblastic, less often decreased
accumulation (" cold lesions""-
osteolytic processes without surrounding osteoblastic reaction).
Symptoms of pathology may also be an overall diffuse
increase in radiopharmaceutical uptake in metabolic bone
diseases, bone marrow involvement, or generalized massive tumor
infiltration - so-called" superscan
", excessively intense and contrasting skeletal imaging.
Pathological finding on the skeleton does not direct information
about the etiology of increased osteoblastic activity. This may
be inflammation, traumatic changes, tumors.
The importance of bone
scintigraphy, however, is early detection of cancer bone disease
- primary, but most often bone metastases . This
place, acting as osteoblastic lesions, are shown
as significant deposits (solitary and multiple) of increased
accumulation of osteotropic radiopharmaceuticals. .
Left:N ormal skeletal scintigram of the skeleton Right:
Multiple metastases (ca. breast) to the skeleton
Evaluation of skeletal scintigraphy is basically visual , but it is also possible to perform a quantitative analysis of radiolabel uptake in selected osteoblastic foci (expressed, for example, as SUV - Standardized Uptake Value ) and based on this to assess disease progression and response to treatment. The procedure for computer evaluation of static skeletal scintigraphy is described in §3.12 " Static skeletal scintigraphy " of the book "OSTNUCLINE".
Dynamic (3-phase)
scintigraphy of the skeleton
Purpose:
It is a combination of dynamic and static scintigraphy of a
selected area of the skeleton. It is used to assess the perfusion
of bones and surrounding tissues, to assess regional hyperemia
due to osteomyelitis, bone trauma, arthritis, bone tumors,
inflammatory lesions of the surrounding tissues. ....
Execution:
After rapid iv application of osteotropic radiopharmaceutical 99m Tc-MDP, dynamic
scintigraphy of the examined area is started immediately
. Scintigraphy is performed in 3 phases :
1. The perfusion (angiographic) phase ("flow")
detects the regional blood flow of the monitored part of the
skeleton and surrounding tissues, 2sec./frame for about 5
minutes.
2. Tissue phase " blood pool"
capturing the transition of the radiopharmaceutical from the
blood vessels" into the extracellular space of soft tissues
and bones. We take about 10 images / 30 sec.
3. The skeletal late phase is a common static
scintigram after 2-3 hours, the same as the static skeletal
scintigraphy
Evaluation:
The procedure for computer evaluation of dynamic 3-phase skeletal
scintigraphy is described in §3.13 " Dynamic (3-phase)
scintigraphy of the skeleton of the
"OSTNUCLINE" book.
Computer evaluation of dynamic 3-phase scintigraphy of the skeleton | |
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|
Visual
evaluation: After intravenous administration of 99m-Tc-diphosphonate, we first observe in the dynamic phase a good and symmetrical perfusion of the thigh areas, the left knee and tibia show increased perfusion with foci of (probably peristatic) hyperemia. In the further course in the skeletal phase, we observe deposits of increased accumulation of the radioindicator in the head of the left fibula, in honey. condyle of the left tibia and in the later. parts of the knee joint. Conclusion: |
4.9.8 Scintigraphic diagnostics in neurology
- CNS
The central nervous system (CNS) of vertebrates
consists of the brain and spinal cord . The
brain ( encephalon ) is the central governing
organ of the nervous system of all higher animals and humans. It
controls and controls most bodily functions, as well as
perception, thinking, memory, emotional and cognitive functions.
In humans, the brain has a volume of about 1500 cm 3 and weighs an average
of 1350 grams. Very complex processes in the brain, as the
central organ of the central nervous system, take place on
several levels. There are mainly two basic groups of cells in the
brain - neurons and glia .
Neurons,
nerve cells, are the basic functional units of nervous tissue.
They are highly specialized, fully differentiated effector cells
that receive, transmit and process special chemical-electrical
signals (information), which allows the body to respond to
external stimuli. Numerous protrusions emerge from the basic cell
- the body of the neuron (it contains the
nucleus) - so-called dendrites (shorter
protrusions that are centrifugal, transmit excitations to the
neuron) and axons (long protrusions that are
centrifugal, used to transmit excitations out of the neuron).
There are a large number of chemically controlled ion
channels and electro-chemically controlled signal channels
in the neuronal membrane . Individual neutrons are using synapsesfunctionally
interconnected into complex neural networks .
Synaptic connections of neurons,
used to transmit excitations, take place between the nerve
endings of one neuron and the entrance membrane of another
neuron. Synapses have a presynaptic part , a synaptic
cleft and a postsynaptic part . The arousal created
by the irritation of a neuron first spreads across the axon in
the form of an electric charge. A special chemical substance - a neurotransmitter
( nerve mediator) is eliminated from the synaptic
vesicles in the nerve ending - the terminal - due to an
electrical signal from the respective precursors.), which causes
a synaptic potential on another neuron *). Neorotransmitters are
low-molecular substances, providing unidirectional transmission
of excitation from one controlling nerve cell (presynaptic) to
another nerve cell (controlled, postsynaptic), which can again be
a nerve or muscle cell, or. glandular. These substances, when
secreted from the presynaptic cell, diffuse through the synaptic
cleft to the membrane of the postsynaptic cell, where they