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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5255_Библиотеки_им_академика_М_И_Перельмана
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112
Radiation Detection in vivo and gamma camera imaging
Components of thyroid probe system
Crystal
For medium energies a NaI(T1) crystal of 50 mm diameter and 25 mm thick is normally
used in a thyroid probe (1). About 70% of the incident gamma rays from
131
I are absorbed
by 25 mm of NaI(Tl) crystal (2). Thicker crystal improve the sensitivity.
Shielding
Lead Shielding is provided around the detector to reduce its response to the environmental
radiation. Shielding is extended to the back of the detector. About 12.5 mm lead thickness is
used for the shielding (2).
Collimator
A simple lead collimator cylindrical in shape but with a given solid angle is mounted in
front of the crystal to allow photons to strike the detector emitted from a given mass/volume
of tissue. The collimator is said to be a flat field collimator as its response is more or less
uniform (± 10%) at or beyond a specified distance from the detector. The field of view at the
working distance of 15-20 cm should be preferably between 12 cm to 15 cm in diameter.
The shielded and collimated detector is called probe. The probe is usually mounted in a
support that is adjustable at various positioning of the neck for counting measurements.
Associated electronics
The associated electronics in an in-vivo counting system for gamma-radiation measurement
is the same as discussed in in vitro detection. The modern systems provide digital display of
acquired counts.
Application of thyroid probe
Thyroid probe is mainly used to measure the uptake of a radioiodine (
It is routinely used in a nuclear medicine facility to estimate the radioiodine uptake by
thyroid. The required activity is administered to the patient and exactly same amount of
activity (by volume) is placed in a neck phantom. The phantom is positioned at a working
distance (about 15-20 cm) from the detector and counts with and without (background
counts) standard activity for a preset time are acquired. The counts are measured normally
at 2 h, 24 h, and 48 h or at any other time as suggested by the physician. In the same way,
patient’s neck counts and background counts at thigh level are counted under similar geometry
and orientation. While taking thigh counts for background, it should be ensured that the
bladder is voided which otherwise adds undesirable counts in the measurement. The
background counts are subtracted from the measurements to get net counts. The percent
uptake can be calculated using following equation:
131
I) by thyroid gland.
Radiation Detection in vivo and gamma camera imaging
Percentageuptake
100
Net neck counts of the patient
=
tan
Net s dard counts
113
Quality control tests for Thyroid Probe
Acceptance and reference tests
The instrument should first be physically inspected before commissioning for all expected
components in good condition. It should then be subjected to the quality acceptance testing.
The tests like energy calibration, energy resolution, detector efficiency, precision, linearity
of energy response, linearity of activity response are performed in the same manner as for a
spectrometer.
Radiation leakage from the lead shielding
The detector shielding should be checked for a leakage of radiation from a commonly used
radionuclide. It can be accomplished by keeping radionuclide source near but outside the
detector shielding. The count rate is recorded at different locations should not exceed B +
3B above the background. Where B is the background count rate.
Flat field response of the collimator
A point source of
The counts are acquired with the point source at various positions across the field of view
and an iso-response curve is plotted. From the iso-response curve, the working distance and
field of view for routine and specific clinical conditions can be established.
Routine checks
Following are the routine checks, which need to be carried out on daily basis before the
instrument is put to use:
1. Check for the mechanical safety of collimator and probe mounting.
2. Check the photopeak for a standard source of
location. This may be used for efficiency test also.
3. Repeat step 2 for
Check for background count rate under operating conditions for the radionuclide in use.
131
I is used to check the response of the collimator at different distances.
137
Cs by pacing it at a given and fixed
131
I.
Radionuclide Imaging
Rectilinear Scanner
Rectilinear scanner was the first imaging device in nuclear medicine that was introduced in
114
Radiation Detection in vivo and gamma camera imaging
early 1950 by Benedict Cassen. It could provide, for the first time, a two dimensional image
of radioactivity distribution in an organ. In this imaging device, a shielded and collimated
NaI(Tl) detector head was made to move over the organ in a rectilinear pattern line by line.
Focusing collimator was used to measure the counts from a focal point in the organ. The
detector head along with the collimator used to move along a line to measure the point-bypoint counts, which were translated on to a paper as dots with the help of a mechanical
taper. The detector then moved sideways and recorded counts from another line. This would
continue till the entire organ is scanned. The gamma ray coming from the organ of interest
interacts with the crystal and produce scintillation, the intensity of, which is proportional to
the energy absorbed in the crystal. The electrical signals are generated and processed exactly
in the same manner as mentioned in thyroid probe system. This instrument has faded away
after the development of gamma camera and is not described here in detail.
Gamma camera imaging
Basic principle
Hal O. Anger developed the first clinically successful gamma camera (Anger camera) in late
1956. The first camera had a 4-inch sodium iodide crystal coupled to a close-packed array of
7 PMTs. Scintillation light (hence the name scintillation camera) from an absorbed gamma
photon spread over several tubes, and centroiding logic (“Anger” logic) calculated the position
of each event within the crystal. The use of several PMTs to determine the exact location of
scintillation event with his logic circuit proved to be a revolutionary idea in imaging the
tracer distribution in the body. The camera could image organs like heart, brain, liver,
kidneys etc. with better resolution than the rectilinear scanner and in much shorter time. The
dynamic imaging could also be easily started with anger camera for evaluating the functional
status of the body organs. The first commercial Anger camera made by Nuclear–Chicago
Corporation (8-inch scintillator and 19 small PMTs) was delivered to Ohio State University
Hospital in 1962. Later other commercial companies also came forward to manufacture the
Anger camera system.
The modern scintillation cameras are available with a large field of view (LFOV) either
in circular(~400mm dia) or rectangular(~540 × 400 mm) form with Nal(Tl) detector with 3/
8” (~1cm) thickness. Detectors with 5/8” thickness are also available but 3/8” thick crystals
are commonly used for detecting gamma rays from the most commonly used radionuclide
99m
(
Tc) for imaging. The detector, being hygroscopic, is hermetically sealed with thin
aluminum sheet from all sides except the back, which is sealed with light guide. The PMTs
are coupled to light guide with the help of a silica glue. The light guide and silica glue are
transparent so that the scintillations can be viewed by array of photomultiplier tubes (PMTs).
The front face of the crystal is covered by the lead collimator. The entire detector head is
shielded by lead to protect it from undesirable background radiation. The detector assembly
allows photons to pass through the collimator only from the front face in a specified direction.
Radiation Detection in vivo and gamma camera imaging
115
The lead collimator also provides physical protection to the detector. With a parallel hole
collimator the photons are allowed to pass and strike the detector perpendicularly. The block
diagram of gamma camera along with the associated components is shown in figure 2.
Figure 2: Block diagram of Gamma Camera
When a photon interacts with the crystal, a light flash (scintillation) is produced at a
given location in the crystal. The PMT, which is nearest to the point of interaction, will get
maximum light and those, which are farther, will receive lesser amount of light. Thus the
cathode of a PMT (photo cathode) receives light that varies inversely with the distance of
each PMT from the point of interaction. Photocathode converts the light into electrons and
dynodes of PMT multiply them, as has been explained in a chapter on ‘radiation detectors’.
A large number of electrons strike the anode of PMT, which are passed through a resistor or
capacitor to convert them into a voltage signal. Though there is huge electron multiplication
by the PMT but the signal is still too week and needs amplification. This is done with a preamplifier and linear amplifier. However, the modern pre-amplifiers are quite capable of
amplifying the signal to a desired voltage.
Logic circuitry (Anger logic)
This logic circuit determines the location of the event (scintillation) on the crystal. After
amplification, the signals are sent simultaneously to X and Y and Z pulse-arithmetic circuits
(Figure 3). The amplitude of X and Y signals are made in proportion to the spatial coordinates
of the original scintillation by assigning a weighting factor to each PMT depending upon its
location from the center (origin). The weighting factor of the PMT away from the central
point is more than that the near one. For PMTs on the right side of the center, the signals
will be X+ and they are X- for those, on the left side. Similarly Y+ and Y- are assigned to
points, which are above and below the central point in Y direction respectively. Four signals
(X+, X-, Y+ and Y-) are generated from an event in crystal from various PMTs. They finally
116
k
( )
X X
( )
Y Y
X X Y Y
Radiation Detection in vivo and gamma camera imaging
result in three signals; two denote the location (X and Y) and the third represents the amount
of light produced in the crystal or the energy of photon actually absorbed in the detector.
Figure 3: Block diagram of Anger logic for determining the spatial coordinates of the event
produced in the detector
As different energy photons transfer different amount of energy to the crystal, the
output of PMTs will vary in proportion to photon energy. To ensure that the coordinate
signals remain within the field of view of the cameras, it is essential to normalize the X and
Y signals by dividing them with the Z signal. In the Z circuit, the pulses summed to provide
a signal that is proportional to the total energy absorbed in the crystal during photon
interaction. The Z signal is passed on to the pulse height analyzer (PHA) and if this falls
within the pre-selected window then only the coordinate signals (X and Y) are allowed to
record the event on the monitor or to store it in the computer memory. Mathematically the
three signals are shown as follows:
X =
Y =
Z
k
Z
Z =
Where k is a constant. The X and Y signals are positive or negative depending upon their
relative position from the center of the crystal. The most important characteristic of the
positioning signal is that they should exhibit exact linearity with distance from the event
production relative to the crystal center.
Radiation Detection in vivo and gamma camera imaging
117
Digitization of analogue signals
In modern systems (digital cameras) the X, Y and Z signals are digitized with the help of
analogue to digital converter (ADC). The ADCs are normally 8,9,10,12 bit which divide the
X and Y values into 28(256), 29(512), 210(1024), 212(4096) equal parts and each part for a
detector of 40 cm diameter will correspond to 40/256(1.6mm), 40/512(0.8mm) 40/1024
(0.4mm), and 40/4096(0.1mm) respectively. It should be noted that the intrinsic spatial
resolution of NaI(Tl) detector in gamma camera with associated electronic components is at
the most 2-3 mm, therefore 8 bit ADC may be adequate for digitizing X and Y signals in
256×256 matrix. The SPECT acquisition is done mostly in (lower) 128×128 matrix. For
multi channel analyzer (MCA) it is always better to have more channels (12-bit ADC or so).
The speed of ADC should be high enough to handle the high-count rate achievable with
modern systems.
Display
A cathode ray tube can be used to display the counts in the form of point light flash at
appropriate location corresponding to their X and Y coordinates. These events are randomly
generated at various locations in the crystal. The CRT should be able to represent large
number of events per unit time (> million events per minute). These signals are digitized
with the help of an analogue to digital converter (ADC) for further image processing and
display (Figure 4). For displaying the stored data as an image, look up tables are used. The
entire range of counts in various pixels in the matrix can be divided into various gray levels
to form the image. Counts above a selected value may be given white color and below the
background count level may be black.
Figure 4: Schematic representation of digitization of X and Y signals
The counts in between are divided into various shades of gray in increasing or decreasing
order. The upper and lower threshold may be selected/changed depending upon specific
requirement. In addition to gray scales, various color codes are also available and are used
for different level of counts.
118
( )
D L H
Radiation Detection in vivo and gamma camera imaging
Collimators
In gamma camera imaging, collimator plays important role. Without a collimator image
formation is not possible as the entire detector may be flooded with photons from all points
in the object. Collimators allow photons to pass through them in a given direction and
interact with the detector to produce scintillations. For a point in the object (emitting photons)
there is a location on the detector, which is recorded in the computer memory. Thus there is
one to one correspondence between a point in the object and a point (location) on the
detector and then a point in the image matrix. The following types of collimators are
generally used in gamma camera imaging.
1. Parallel hole collimator
2. Diverging collimator
4. Pin hole collimator
5. Focusing collimator
Parallel hole collimator
This is one of the most commonly used collimator in
gamma camera imaging. There are tens of thousands of
parallel holes (drilled perpendicularly) in the collimator
separated by lead septa. The diameter, length of the hole
and septa thickness determine the sensitivity and
resolution of the collimator (Figure 5).
As per IAEA document (3), the collimator resolution
is given by
RC=
e
L
e
(1)
Figure 5: Parallel hole collimator
Where Rc is the resolution of the collimator
D is the diameter of the holes
Lc is the effective length of the holes
H is the distance of collimator from the source
It should be noted that the higher the value of Rc the worse is the collimator resolution.
The numerical value of the Rc should be small for better resolution. It can be seen from
equation-1 that Rc increases (resolution worsens) with increase in hole diameter ‘D’ and
decrease in effective length of the collimator ‘Le’. The value of Rc also increases with H,
though not proportionately. It is therefore desired to keep the patient as close to the collimator
Radiation Detection in vivo and gamma camera imaging
( ( ))
2
(32 (2.5 0.2))
119
as possible to improve the resolution. Collimator resolution may be improved by decreasing
the hole diameter and increasing the effective length of the collimator holes. However, this
adversely affects the sensitivity of the collimator. The collimator sensitivity is the fraction
of gamma rays that is allowed to pass through it. As per IAEA document (3), it is expressed
mathematically as:
4
C D
S =
L D T
e
2
(2)
Where S is the sensitivity of the collimator and T is the septa thickness. C is a
constant determined by the shape of hole in the collimator. For hexagonal, circular and
square holes the values of C are taken as 0.069, 0.063 and 0.080 respectively (3).
When collimator is to be used for imaging with energetic gamma rays, septa thickness
is required to be increased accordingly. Septal thickness ranges from about 0.2 mm for low
energy radionuclides, such as
99m
Tc to approximately 1.5 - 2 mm for
131
I gamma rays and for
other isotopes with energies above 400 keV (3).
For a collimator with hole diameter of 2.5 mm and effective hole length of 32 mm,
septa thickness of 0.2mm and C=0.069 (hexagonal holes), the sensitivity of the collimator
using equation-2 will be
4
S =
0.069 2.5
S = 0.00036 or 0.036%
This example shows that only 1 gamma ray out of 2800, which have fallen on the
collimator, can pass through it to strike the crystal. The values of the collimator parameters
used above may slightly vary from manufacturer to manufacturer. The collimator sensitivity
always remains low because of the lead septa in the path of the photons traveling towards
the detector.
In imaging we need better resolution and high sensitivity, therefore the design of a
collimator is always a trade off between resolution and sensitivity for a given energy range.
From equations given above, it can be realized that increase in sensitivity degrades the
resolution. The parameters which determine resolution and sensitivity of a parallel hole
collimator are the hole diameter D, the effective hole length Le, the thickness of the septa T,
the distance of the source from the collimator H and the shape of the hole. The effective
hole length Le is slightly less than the physical hole length to take into account the penetration
of the gamma rays through edges of the septa at either end. When Le is not known, the
physical length can be used for calculation purposes.
The overall system resolution (Rs) can be calculated as:
120
s
R
2 2
( )
R R
Radiation Detection in vivo and gamma camera imaging
=
c i
Where Ri is the intrinsic resolution of detector.
Pinhole Collimator
Pinhole collimator is used to image small organs such as the thyroid. It has a single
aperture hole at a distance of 20-30 cm from the detector (Figure 6). The hole diameter can
be varied from 2mm to 8 mm using inserts with different hole diameters. Image of the
object is inverted on the detector due to pinhole principle but most of the manufacturers
incorporate a software that displays the image in right order. The image is magnified if
placed near the hole but the magnification decreases with the distance from the collimator.
They are not the frequently used collimators these days as zoom acquisition with parallel
hole collimator is available.
Converging and diverging collimators
In both these collimators, holes converge to a point rather than being parallel to each
other (Figure 6). In converging collimator the holes converge towards a point (focal point)
in front of the collimator. The image gets magnified if placed between the collimator and
focal point. In diverging collimators the holes converge in opposite direction i.e. towards a
point behind the collimator. The image of the object therefore gets minified. They were
quite useful to overcome the limited field of view of gamma camera in early days when
LFOV detectors were not available. Both these collimators are hardly used today in nuclear
medicine.
Figure 6: Parallel hole collimator (Top left), Diverging collimator (Top right), Pin hole
collimator (bottom left3) and Converging collimator (bottom right)
Radiation Detection in vivo and gamma camera imaging
121
Focusing collimators
A long bore fan beam collimator was designed for
SPECT imaging of head (4). The geometric
efficiency of such a fan beam collimator was 55%
higher than that of a parallel hole collimator having
same hole length and spatial resolution. With further
improvement in collimator design (Figure 7), the
application of fan beam collimator has gained
popularity in SPECT imaging of small organs. Fan
beam collimator is one of the most popular focusing
collimators, which has been extensively used for
brain SPECT. It improves the sensitivity by 1.5 to 2
Figure 7: Cross sectional view of fan
beam collimator
fold over a parallel hole collimator of comparable
resolution.
Quality Control of Gamma Camera
General Aspects
Gamma camera is the most important instrument in nuclear medicine, which is used both for
planar and tomographic imaging. These systems have many sensitive electronic components/
circuits, which needs to be checked and corrected for their optimum performance. Similarly,
the associated computer system should also be checked during acceptance testing and then
periodically. The mechanical movements particularly during rotation of the detector head in
SPECT studies need proper evaluation, calibration and then routine/periodic QC tests.
The terms quality assurance, quality control and calibration should not be confused.
Quality assurance refers to all aspects of a procedure, which contribute to the quality of the
results obtained. In nuclear medicine, this term includes all aspects such as proper organisation
of services, routine quality control of equipment and radiopharmaceuticals, other components
such as submission of requests for the radionuclide procedure, the explanation to the patient
by nuclear medicine staff (do’s and don’t) including radiation safety procedures, preparation
and administration of radiopharmaceuticals, safety against radiation hazards and accidents
caused by faulty equipment, the scheduling of patients for investigations, reporting and
finally evaluation of results. Similarly, the quality assurance of nuclear medicine
instrumentation will include all aspects of evaluation for the performance of various
equipment, computer hardware and software, hard copy device, image quality assessment
and finally their organisation and implementation to provide best quality image. The term
quality control basically refers to the evaluation of all those parameters, which are related to
the performance of the instrument. Calibration basically refers to the adjustments of
parameters or values for optimum performance of the equipment. For example energy
calibration in gamma camera refers to place the window centrally on the photopeak for
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