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sample and the standard. However, when the absolute activity of a radioactive sam-
ENERGY (keV)
ple needs to be determined, then the detection efciency of the counter must be
measured for the γ-ray energy of interest using a standard of the radioactive sample
of known activity. The photopeak efciency is determined from the count rate of the
standard at the appropriate PHA setting divided by the disintegration rate from the
known activity of the standard. The efciency correction can then be applied to the
count rates of samples of unknown radioactivity when counted at the same setting
as the standard to give the absolute activity. For absolute activity, the photopeak
efciency must be determined for each photon energy.
When multiple γ-rays, either from a single radionuclide or from many radionuclides, are present in a radioactive sample, then the energy spectrum becomes complicated by the overlapping of different photopeaks and also by Compton
contributions from the high-energy photons to the low-energy photopeaks. The latter contributions are termed the spillover, or crosstalk contributions.
Figure 8.12 illustrates an energy spectrum of the 140-keV peak of
keV peak of
131
I, in which the Compton contribution from the 364-keV peak to the
99m
Tc and 364-
140-keV peak is shown. Corrections must be made for this spillover to the 140-keV
peak. This is accomplished by counting a sample of pure
131
I in both 140-keV and
364-keV discriminator settings and determining the percentage of spillover from
the ratio of the counts in the 140-keV photopeak to those in the 364-keV phoopeak.
140 keV
364 keV
Fig. 8.12 A combined spectrum of the 140-keV γ-ray of
ted line under the 140-keV photopeak is the spillover, or crosstalk, contribution from the 364keV photon
99m
Tc and 364-keV γ-ray of
131
I.The dot-
8.7.3 Effects ofSample Volume
The sample volume affects the counting efciency of well counters. As the sample
volume for a given activity is increased, more radiations are lost through the opening of the well without interacting in the detector, and hence, the counting efciency
drops. Therefore, correction factors should be determined for different sample volumes and applied to the measured activity.
Well counters are available with automatic sample changers having provisions of
counting as many as 500 samples. Most counters are programmable with computers
and provide printouts with various information on counting. The major advantage of
the well counter is its high detection efciency due to increased geometric efciency, which approaches almost 100% depending on the volume of the sample.
The detection efciency of a well counter decreases with increasing photon energy
and decreasing detector size. Typically, the overall detection efciency is close to
100% for 140-keV photons of
99m
Tc and 30–90% for 364-keV photons of
131
I,
depending on the detector size.
8.8 Thyroid Probe
The thyroid probe is a counter commonly employed to measure the uptake of
123
I in the thyroid gland after the oral administration of a
131
I-NaI or
123
I-NaI capsule.
131
I or
It consists of a NaI(Tl) detector, 5cm in diameter by 5cm in thickness, and other
associated electronics, as in a well counter. The operation of the probe is similar to
that of a well counter.
One of the differences between the well counter and the thyroid probe is that the
latter requires a collimator, which limits the eld of view on the thyroid. The collimator is a 20- to 25-cm long cylindrical barrel made of lead and covers the detector
as well as the PM tube (Fig.8.13). This reduces the background activity from the
γ-radiations from areas outside the thyroid reaching the detector.
Fig. 8.13 A schematic diagram of a thyroid probe. PHA pulse-height analyzer
The efciency of a thyroid probe varies inversely with the square of the distance
AB
100
between the detector and the thyroid. The probe is initially calibrated for photon
energies in the same manner as the well counter using the 662-keV γ-ray energy of
137
Cs, and then discriminator settings are set for the 364-keV γ-ray of
131
I.Attenuation
of photons in the thyroid tissues reduces the overall detection efciency of the probe.
Photons scattered within the thyroid gland by Compton scattering may reach and
interact in the detector because they originate in the eld of view and are not stopped
by the collimator thickness. These scattered photons, however, are excluded from
the total measured counts by selecting the appropriate lower and upper discriminator settings on the PHA for the 364-keV γ-ray of
131
I.
8.8.1 Thyroid Uptake Measurement
In the thyroid uptake test, a
(0.37–0.55MBq) of
131
I is measured in a lucite thyroid phantom at a xed distance
using the thyroid probe and the settings for 364-keV photons of
131
I-NaI capsule containing about 10–15 μCi
131
I.The thickness
and composition of the lucite phantom are equivalent to those of the patient’s neck.
This count is considered as the standard count. The capsule is then administered to
the patient orally, and the thyroid count is obtained at the same distance as the standard count 24h after administration. The room background count is taken to subtract from the standard count, and the thigh count is taken as background to subtract
from the thyroid count. The thyroid uptake is then calculated as follows:
uptake
CD
(8.9)
where A is the thyroid count, B is the thigh count, C is the standard count corrected
for 24-h decay, and D is the room background. Identical procedures are employed
123
with
I-NaI using ∼300 μCi (11.1MBq). At times, the 6-h thyroid uptake also is
determined depending upon the clinical judgement of the physicians.
8.9 Questions
Describe the mechanism of γ-ray interaction in the NaI(Tl) detector. In γ-ray counting, why is NaI(Tl) commonly chosen as the detector?
1. (a) Describe the operation of a photomultiplier (PM) tube.
(b) What is the typical high voltage applied to the PM tube?
(c) What are the photocathodes commonly made of?
(d) How many photoelectrons are emitted from the photocathode for each keV
of photon energy?
2. (a) Ideally, a photopeak should appear as a line in a γ-ray spectrum. Indicate
different factors that contribute to the broadening of the photopeak.
(b) A photopeak is due to only photoelectric effect of γ-rays, or due to all
γ-rays that deposit full energy in the detector. True or false?
3. (a) Describe the function of a pulse-height analyzer (PHA).
(b) Do the following factors affect the size (pulse height) of the photo-
peak pulses?
(i) Gain of the amplier.
(ii) High voltage of PM tubes.
(iii) Distance between the source and the detector.
(iv) Light photons produced in the detector.
4. In a γ-ray spectrum, describe the origins of the following:
(a) Backscatter peak.
(b) Compton valley.
(c) Characteristic K x-ray peak.
(d) Iodine escape peak.
(e) Sum peak.
5. (a) Describe the principles of a liquid scintillation counter.
(b) What is a scintillation solution and how does it work?
(c) What is the purpose of using a secondary solute to the scintillation solution?
(d) What are the most common solvents for liquid scintillation counting?
(e) Can you count 3H (β− energy=0.018MeV) and 14C (β− energy=0.156MeV)
in the same sample using a liquid scintillation counter equipped with
three PHAs?
6. (a) Dene the energy resolution of a detector.
(b) For a given detector, the energy resolution of low-energy photons is poorer
than that of high-energy photons. True or false?
(c) For NaI(Tl) detectors, the energy resolution should be less than 10% for the
662-keV photon of
7. (a) A point source of
137
Cs. True or false?
99m
Tc is placed 10cm away from a NaI(Tl) detector that
has a diameter of 20cm. Calculate the geometric efciency.
(b) What would be the geometric efciency if the source were placed on the
surface of the detector?
8. (a) Explain the dead time and pulse pileup of a counter.
(b) What is the distinction between the paralyzable and nonparalyzable
systems?
(c) What are the typical dead times for Geiger–Müller (GM) counters and
NaI(Tl) counters?
9. (a) Describe the energy calibration of a NaI(Tl) well counter.
(b) Why does the count rate differ from the disintegration rate of a sample of a
radionuclide?
(c) How does the sample volume affect the count rate?
10. What are the spillover or crosstalk contributions in a spectrum of several γ-rays?
How would you correct for them?
11. A radioactive sample has two γ-ray photons of 130- and 120-keV energies. If a
NaI(Tl) crystal has an energy resolution of 10% at 125keV, could the two photons be detected as separate photopeaks?
12. Both gas-lled detectors and semiconductor detectors operate by ionization of
atoms by radiation. Why do semiconductor detectors give better energy resolution than gas-lled detectors?
13. A patient is given orally a 10-μCi
131
I-NaI capsule. Before administration, the
count rate of the capsule in a thyroid phantom is 297,000cpm. The 24-hour
count rate of the patient’s thyroid is 168,000cpm. If the room background is
200 cpm and the patient’s thigh count rate is 1000cpm, calculate the thyroid uptake.
14. High-activity sources such as radiopharmaceutical dosages and x-ray exposure
outputs are better measured with ionization chambers than GM counters and
NaI(Tl) well counters. Why?
15. Which of the following counters can detect individual events of the radiation
interacting with the detector?
(a) Ionization chamber.
(b) GM counter.
(c) NaI(Tl) well counter.
16. What type of Compton scattering causes the Compton edge of a γ-ray spectrum?
17. Discuss the properties of newer detectors. Explain why LSO is preferably used
in PET cameras.
Suggested Readings
Bushberg JT, Seibert JA, Leidholdt EM Jr, Boone JM. The Essential Physics of Medical Imaging.
3rd ed. Philadelphia: Lippincott Williams & Wilkins; 2011.
Cherry SR, Sorensen JA, Phelps ME. Physics in Nuclear Medicine. 4th ed. Philadelphia:
W.B.Saunders; 2012.
Cradduck TD.Fundamentals of scintillation counting. Semin Nucl Med. 1973; 3:205–223.
Hendee WR, Ritenour ER. Medical Imaging Physics. 4th ed. NewYork: Wiley-Liss; 2002.
Hine GJ. Sodium iodide scintillators. In: Hine GJ, eds. Instrumentation in Nuclear Medicine.
NewYork: Academic Press; 1967; I: 95–117.
Knoll GF. Radiation Detection and Measurement. 4th ed. NewYork: Wiley; 2010.
Peng CT, Horrocks DL, Alpen EL, eds. Liquid Scintillation Counting. NewYork: Academic Press;
1980; I, II.
Gamma Camera
9.1 Gamma Camera
The gamma or scintillation camera is an imaging device that is most commonly
used in nuclear medicine. It is also called the Anger camera in honor of Hal
O.Anger, who invented it in the late 1950s. Gamma cameras detect radiation from
the entire eld of view simultaneously and therefore are capable of recording
dynamic as well as static images of the area of interest in the patient. Various designs
of gamma cameras have been proposed and made available, but the Anger camera
with a single crystal is by far the most widely used. Although many sophisticated
improvements have been made on the gamma cameras over the years, the basic
principles of the operation have essentially remained the same.
9
9.1.1 Principles ofOperation ofaGamma Camera
The gamma camera usually consists of several components: a detector, a collimator,
PM tubes, a preamplier, an amplier, a pulsed-height analyzer (PHA), an X-,
Y-positioning circuit, and a display or recording device. A schematic diagram of a
gamma camera is illustrated in Fig.9.1, and a commercial gamma camera is shown
in Fig.9.2. The detector, PM tubes, and ampliers are housed in a unit called the
detector head, which is mounted on a stand. The head can be moved up or down to
position on the patient. The X-, Y-positioning circuits, PHA, and some recording
devices are mounted on a console. In the past, the cameras were operated by switches
and dials on the console. Nowadays, much of the operation of the camera is performed by a computer built in it. The computer is run by appropriate software in
conjunction with a keyboard, a mouse, and a video monitor. High voltage, window,
and photopeaks are all set by the operator’s choice of parameters. Acquisition of the
data and processing of the data are carried out by the computer. Although stationary
cameras are permanently installed at desired locations, portable gamma cameras are
© The Author(s), under exclusive license to Springer Science+Business Media, LLC,
part of Springer Nature 2025
G. B. Saha, Physics and Radiobiology of Nuclear Medicine,
https://doi.org/10.1007/978-1-0716-4816-2_9
119
120
9 Gamma Camera
Amplifier
Preamplifier
PM Tubes
NaI(Tl)
Collimator
Source
Fig. 9.1 A schematic diagram of a gamma camera. PHA pulse-height analyzer, PM
Photomultiplier tube
Fig. 9.2 A typical
single-head gamma
camera. (Courtesy of
Siemens Medical Solutions
USA, Inc.)
X,Y Positioning
Summing Circuit
Y
X
Circuit
X
Display Scaler
Z
Z
Storage
PHA
Z
Z
Y
mounted on wheels, for use in situations requiring movement of the camera from
room to room, such as to the patient’s bedside. Mobile cameras are installed in
wheeled vans such that they can be moved to places where gamma cameras are not
available for nuclear medicine studies.
The operational principles of a gamma camera are identical to those of solid
scintillation counters described in Chap. 8. Basically, γ-rays from a source interact
with the NaI(Tl) detector, and light photons are emitted. The latter strike the photocathode of PM tubes, and a pulse is generated, which is then amplied by an amplier and sorted out by a PHA.Finally, the pulse is positioned by an X-, Y-positioning
circuit on the recording device or stored in the computer, corresponding to the location of γ-ray interaction in the detector (see later).
9.1 Gamma Camera
121
The functions of PM tubes, preamplier, amplier, PHA, and recording devices
are the same as described in Chap. 8, and therefore, only essential features pertaining to the use of gamma cameras are highlighted.
9.1.2 Detector
NaI(Tl) crystals are the common detectors used in gamma cameras. In older cameras, circular detectors of a dimension of 25–35cm in diameter were used, whereas
in modern cameras, most manufacturers use rectangular detectors of the size of
approximately 45×60cm, which provide a practical eld of view of 40×55cm.
Although the thickness of the detector used varies from 0.63 to 1.84cm, the optimum thickness of 0.95cm is used by many manufacturers in the current cameras.
Thinner crystals (0.64cm) are sufcient for low-energy radionuclides such as
99m
Tc, and
123
I and used in portable gamma cameras for cardiac studies.
NaI(Tl) crystals are hygroscopic, and absorbed water causes color changes that
reduce light transmission to the PM tubes. For this reason, these crystals are hermetically sealed in aluminum containers. Also, the entrance and side of the crystals
are coated with a reective substance (e.g., magnesium oxide) so that light photons
are reected toward the photocathode of the photomultiplier tube.
Increasing the thickness of a detector increases the probability of complete
absorption of γ-rays and hence the sensitivity (dened in Chap. 10) of the camera.
However, the probability of multiple Compton scattering also increases in thicker
detectors, and therefore the X, Y coordinates of the point of γ-ray interaction can be
misplaced. This results in poor resolution of the image of the area of interest. For
this reason, thin NaI(Tl) detectors are used in gamma cameras, but this decreases
the sensitivity of the camera, because many γ-rays may escape from the detector
without interaction.
201
Tl,
9.1.3 Collimator
In gamma cameras, a collimator is attached to the face of the NaI(Tl) detector to
limit the eld of view so that γ-radiations from outside the eld of view are prevented from reaching the detector. Collimators are normally made of material with
high atomic number and stopping power, such as tungsten and lead, of which lead
is the material of economic choice in nuclear medicine. Two methods are employed
to make collimators—the cast method and foil method. In the cast method, holes are
made through a single slab of lead, whereas in the foil method, corrugated lead
strips are glued together, through which holes are made. They are designed in different sizes and shapes and contain one or many holes to view the area of interest.
Collimators are primarily classied by the type of focusing, although other classications are also made based on septal thickness and the number of holes.
Depending on the type of focusing, collimators are classied as parallel-hole, pinhole, converging, and diverging types; these are illustrated in Fig.9.3. Pinhole collimators are made in conical shape with a single hole and are used in imaging small
122
Pinhole collimator
Converging collimator
Fig. 9.3 Different designs
of collimators
9 Gamma Camera
Parallel hole collimator
Diverging collimator
organs such as the thyroid glands to provide magnied images. Converging collima-
tors are made with tapered holes converging to an outside point and are employed
to provide magnied images when the organ of interest is smaller than the size of
the detector. Images are magnied by converging collimators. Diverging collima-
tors are constructed with tapered holes that are divergent outward from the detector
face and are used in imaging organs such as lungs that are larger than the size of the
detector. The images are minied with these collimators.
Parallel-hole collimators are made with holes that are parallel to each other and
perpendicular to the detector face and have between 4000 and 46,000 holes depending on the collimator design. These collimators are most commonly used in nuclear
medicine procedures and furnish a one-to-one projected image. Because pinhole
and converging collimators magnify and the diverging collimators minify the image
of the object, some distortion occurs in images obtained with these collimators.
Because large eld of view (LFOV) cameras are readily available now, diverging
collimators are not used in routine nuclear medicine studies.
Parallel-hole collimators are classied as high-resolution, all-purpose, and highsensitivity type, or low-energy, medium-energy, or high-energy type, depending on
the resolution and sensitivity they provide in imaging. High-sensitivity collimators
are made with smaller thickness than all-purpose collimators, whereas highresolution collimators are thickest of all. These characteristics are discussed in
detail in Chap. 10.
Several collimators are available that are designed for some specic purposes.
Fan-beam collimators are designed with holes that converge in one dimension but
are parallel to each other in the other dimension. These collimators are primarily
used for imaging smaller objects and hence magnify the images. Cone-beam collimators are similar to fan-beam collimators and magnify the images except that the
holes are designed such that they converge in two dimensions.
In earlier collimators, the holes were made circular, but current designs have
square, hexagonal, or even triangular holes with uniform thickness of lead around
the opening. These collimators provide better spatial resolution than the circularhole ones.
9.1 Gamma Camera
123
9.1.4 Photomultiplier Tube
As in scintillation counters, PM tubes are essential in gamma cameras for converting the light photons in the NaI(Tl) detector to a pulse. Instead of one PM tube, an
array of PM tubes are mounted on the back of the detector with optical grease, or in
some instances, using lucite light pipes between the detector and the PM tubes. In
older cameras, the number of PM tubes used were 19 or 37, whereas most manufacturers utilize more than 100 PM tubes in modern cameras. In modern gamma cameras, square or hexagonal PM tubes are used for better packing. The output of each
PM tube is used to dene the X, Y coordinates of the point of interaction of the γ-ray
in the detector by the use of an X-, Y-positioning circuit (see later) and also is
summed up by a summing circuit to form a pulse known as the Z pulse. The Z pulse
is then subjected to pulse-height analysis and is accepted if it falls within the range
of selected energies.
9.1.5 X-, Y-Positioning Circuit
Each pulse arising out of the γ-ray interaction in the NaI(Tl) detector is projected at
an X, Y location on the image corresponding to the X, Y location of the point of
interaction of the γ-ray. This is accomplished by an X-, Y-positioning circuit in conjunction with the PM tubes and a summing circuit. Figure9.4 illustrates the principles of X, Y positioning of pulses arising from γ-ray interactions in the detector
employing seven PM tubes. All PM tubes are connected through capacitors to four
output leads representing four directional signals, X+, X−, Y+, and Y−. The
Fig. 9.4 Arrangement of
seven photomultiplier
(PM) tubes to produce X
and Y pulses for the X, Y
location of the γ-ray
interaction in the detector.
+
X
, X−, Y+, and Y− pulses
are obtained by summing
the output of all PM tubes
weighted by capacitors for
the location of each PM
tube in relation to the site
of γ-ray interaction.
(Adapted from Anger
1958:27)