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Source: STANDARD HANDBOOK OF BIOMEDICAL ENGINEERING AND DESIGN
CHAPTER 27
NUCLEAR MEDICINE IMAGING
INSTRUMENTATION
Mark T. Madsen
University of Iowa, Iowa City, Iowa
27.1 INTRODUCTION 27.1
27.2 SCINTILLATION CAMERAS
27.3 SPECT SYSTEMS 27.14
27.4 SUMMARY 27.20
REFERENCES 27.20
27.2
27.1 INTRODUCTION
Nuclear medicine is a diagnostic imaging modality that is used to obtain clinical information about
most of the major tissues and organs of the body. Diagnostic information is obtained from the way
the tissues and organs process radiolabeled compounds (radiopharmaceuticals). The radiopharmaceutical is typically administered to the patient though an intravenous injection. The radiopharmaceutical
is carried throughout the body by the circulation where it localizes in tissues and organs. Images of
these distributions are acquired with a scintillation camera. Ideally, the radiopharmaceutical would go
only to abnormal areas. Unfortunately, this is never the case and the abnormal concentration of the
radiotracer is often obscured by normal uptake of the radiopharmaceutical in the surrounding tissues.
Images of higher contrast and better localization can be obtained with tomographic systems designed
for nuclear medicine studies (SPECT systems). These are described in detail below.
The imaging of radiotracers in the body presents special challenges that are unique. The flux of
gamma rays available for imaging is orders of magnitude less than that used in x-ray radiography or
x-ray computed tomography (CT). In addition, the high energy of the gamma rays makes detection
more difficult. As a result, the images produced in nuclear medicine studies are much noisier and
have worse spatial resolution. In order to appreciate these problems and how they affect the design of
nuclear medicine imaging devices, we will briefly review the physics of gamma ray interactions.1
The intensity of gamma rays traveling through material is gradually reduced by absorption or
scattering. This loss of gamma rays is referred to as attenuation and is described by the exponential
equation
(27.1)
where I0 = initial intensity
†††††††I(x) = intensity of rays after traveling a distance x through the material
††††††††††µ = linear attenuation coefficient of the material
27.1
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27.2 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
Over the range of gamma ray energies used in radionuclide imaging, the two primary interactions
that contribute to the attenuation coefficient are photoelectric absorption and Compton scattering.
Photoelectric absorption refers to the total absorption of the gamma ray by an inner shell atomic
electron and is the primary interaction in high-Z materials such as sodium iodide (the detector
material used in the scintillation camera) and lead. In low-Z materials such as body tissues, its
contribution to attenuation is relatively small. Compton scattering occurs when the incoming gamma
ray interacts with a loosely bound outer shell electron. A portion of the gamma ray energy is imparted
to the electron and the remaining energy is left with the scattered photon. The amount of energy lost
in the event depends on the angle between the gamma ray and scattered photon. Compton scattering
is the dominant interaction in body tissues.
High attenuation is desirable in detecting and shielding materials. Ideally materials used for these
purposes would absorb every gamma ray. In the body, attenuation is very undesirable, but
unfortunately, unavoidable. Attenuation reduces the intensity of gamma rays available for detection
and scattered radiation that reaches the detector causes a significant loss of contrast.
27.2 SCINTILLATION CAMERAS
The scintillation camera is the primary imaging instrument used in nuclear medicine and is often
referred to as a gamma camera. 2 The scintillation camera is a position-sensitive gamma ray imager.
Although the entire field of view is available for detection, it processes one event at a time. The
spatial resolution is approximately 10 mm and it yields a count rate of 200 to 300 cpm/µCi in the
field of view (cpm = counts per minute). The field of view covers a large portion of the body and is
typically 40 ◊ 50 cm, although other sizes are available.
The rectilinear scanner was the first practical nuclear medicine imaging device and it was still in
use through the 1970s. It was invented by Benedict Cassen in 1950. The rectilinear scanner used the
detected count rate of a radiation detector to control the brightness of a small light bulb masked to
expose a small area of a film. The movement of the radiation detector and the bulb were linked so
that, as the detector moved in a raster pattern over the patient, the bulb tracked a corresponding path
over the film. The developed film revealed the internal distribution of the radiopharmaceutical
sampled by the scanning probe. While the invention of the rectilinear scanner was a major step
forward, it had several shortcomings that were inherent in its design. Because it sampled only one
small area at a time, image acquisitions took a long time (10 to 20 minutes). In addition, only static
imaging was possible. Dynamic studies, such as those that followed the progression of a tracer
through the body, required an imaging system with a large field of view where all areas are equally
sampled, i.e., a scintillation camera.
The first scintillation camera was developed by Hal O. Anger in 1958.3 Although this system was
very crude, it contained the fundamental components of all future designs: NaI(Tl) as the primary
detector and weighted signals from an array of photomultiplier tubes to determine the location of
detected events. Table 27.1 gives typical performance values for a modern scintillation camera.
The gamma rayñsensitive element of the scintillation camera is a large, thin piece of NaI(Tl).
Although the crystals originally had a circular cross section, most scintillation cameras now use a
rectangular crystal with dimensions as large as 40 ◊ 50 cm. The thickness of NaI(Tl) in most
conventional cameras is 9.5 mm, but in systems that are used for coincidence detection, the crystal
may be twice as thick. NaI(Tl) is a scintillator; It converts gamma ray energy into visible light. The
amount of light generated is directly proportional to the absorbed energy. NaI(Tl) is very efficient at
this and the absorption of one 140-keV gamma ray will yield 5000 visible light photons. There are
a number of advantages associated with NaI(Tl) in addition to its high light output. It efficiently
absorbs 140-keV gamma rays (with a photopeak efficiency of 85%) and it has a moderate energy
resolution. Energy resolution is an important property since it provides the means to discriminate
against scattered radiation. Gamma rays that undergo scattering within the patient degrade the quality
of images. However, scattered gamma rays necessarily have less energy than unscattered gamma rays
and can be selectively eliminated on that basis. Another positive feature of NaI(Tl) is that it can be
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.3
TABLE 27.1 Scintillation Camera Specifications
manufactured in many shapes and sizes. There are disadvantages though. NaI(Tl) actively absorbs
water vapor from the air and loses its transparency. It must be hermetically sealed, and loss of this seal
results in irreparable damage. Another disadvantage is that the persistence of the scintillation is long
enough that it limits the count rate that the crystal can accurately handle. Most nuclear medicine
imaging is performed far below this limit. However, some first pass studies do result in significant
count rate losses from this limit. The biggest problem is encountered in coincidence imaging.
Converting the gamma ray energy to visible light is only part of the battle. In order for the
information from the scintillation to be useful, it has to be converted into an electronic signal. This is
accomplished with a photomultiplier tube (Fig. 27.1). The photomultiplier tube is a vacuum tube with
a photoemissive surface called the photocathode. Visible light hitting this surface knocks off
electrons. These electrons are accelerated to an electric terminal called a dynode. The first dynode has
FIGURE 27.1 Photomultiplier tube. The photomultiplier tube converts the scintillation into an electronic pulse preserving the linear relationship between the magnitude of the scintillation and the energy of the interaction. The location of a source
can be inferred from the magnitude of the signal change. The relationship between
the signal magnitude and source position is nonlinear, and positioning errors occur
both when the source is far from the PMT and when it is directly under it. The ideal
response can be approximated with the use of a light pipe.
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27.4 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
a potential approximately 100 volts higher than the photocathode, and the electrons hit it with
enough force to knock off about 4 more new electrons. The next dynode is another 100 volts higher,
so the process is repeated. The same process occurs over a total series of nine dynodes resulting in a
signal amplification of 1,000,000. Proportionality is maintained throughout this amplification so that
the size of the electron pulse is directly proportional to the energy deposited by the gamma ray.
A scintillation camera needs to be able to record gamma ray events over a large area. This requires
uniform sampling by an array of photomultiplier tubes (PMTs). The PMTs are arranged in a close
packed array that covers the entire surface of the NaI(Tl) crystal (Fig. 27.2). The PMTs used in
scintillation cameras are usually 2 or 3 in across, so that as many as 120 PMTs may be used. PMTs
have been manufactured in variety of different cross sections in order to maximize their areal
coverage. Circular, hexagonal, and square tubes have all been used. The signals obtained from the
PMTs will be used to determine two important properties about the gamma ray interaction: where did
it occur and how much energy was deposited? At first blush, it may seen that even 2-in PMTs are too
coarse to determine the event location. However, we will see that magnitude of the PMT output is
fairly sensitive to the location of a source.
FIGURE 27.2 Photomultiplier tube array. The photomultiplier tubes are
arranged in a close-packed array to cover the back surface of the NaI(Tl)
crystal.
If a PMT is mounted to an NaI(Tl) crystal and the signal output is plotted as a source is moved
from left to right, the result shown by the solid line (Fig. 27.1) will be obtained. When the source is
positioned far from the PMT, the signal is weak and the location of the source is not certain. When
the source is directly under the PMT, the signal is strong, but the dependence on the position is
modest. However, when the source is just to the left or right of the PMT, the signal change with
source location is large. Over this region, the location of the source can be accurately tracked. With
only a single PMT, we could not tell if the source was on the left or right (or front or back), but by
considering the signals from other surrounding PMTs, that can be determined. The main problem is
what to do about the poor response near the center of PMT. What can be done to transform the
measured PMT output into the ideal signal shown in Fig. 27.1? The early solution to this was to
displace the PMTs from the NaI(Tl) crystal with a light pipe.
A light pipe is a transparent material such as lucite or quartz that is interspersed between the crystal
and PMT array. The displacement of the PMTs causes the light from the scintillations to spread out,
yielding a more favorable signal output. To achieve even better results, some manufacturers have
used special mask patterns on the front surface of the light pipe along with sculptured grooves in the
back. These efforts pay off in more accurate positioning of the event locations; however, there is a
price to pay. The effort to scatter the light also results in higher light losses, leading to increases in
statistical fluctuations and ultimately degradations in spatial resolutions. In the early gamma cameras,
there was no way around this dilemma. However, digital electronics provides the capability of
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.5
nonlinear mapping, which greatly reduces the demands put on the light pipe. This has allowed light
pipes to be made very thin and even eliminated.
Figure 27.3 shows a schematic of an analog scintillation camera. The scintillation light from an
absorbed gamma ray is transmitted through to the photomultiplier tube array. The energy of the event
is determined by summing all the photomultiplier tube signals. As will be seen, the energy signal is
used both for scatter discrimination and for normalizing the position signals. The position of the
event is determined by summing weighted outputs from each photomultiplier tube. These weighting
factors are determined by the location of the photomultiplier tube in the array. Separate weighting
factors are used for determining the x and y signals. This process is referred to as Anger logic, since
it is the scheme developed by Hal Anger in the first scintillation camera. In the initial designs, literally
all the photomultiplier tubes participated in the energy and position signal summations. It was
subsequently found that the signals from photomultiplier tubes located far from the event contributed
mostly noise. In modern designs, the photomultiplier tube signal must exceed a threshold before it is
included in the sum. Another point that should be made is that all the processing is performed on each
detected event. The decision to include the event as a valid count is not made until the end of the
processing when the pulse height analysis is done. If the event falls within the selected energy
window, the normalized x and y signals are available for either an analog display or digital storage.
FIGURE 27.3 Analog scintillation cameras. The signals from each photomultiplier tube are sampled to determine both the position and energy of the detected
event. Separate weighting factors are used for the x and y signal determinations. The
energy signal is used to normalize the position signals and discriminate against
scattered radiation.
The position signals determined from summing the weighted PMT signals vary with the brightness
of the scintillation, which itself depends on the energy absorbed in the crystal. This means that an
object imaged with a high-energy gamma ray like I-131 (364 keV) will be magnified when compared
to the same object imaged with Tc-99m (140 keV). This magnification is a concern even when only
one gamma ray energy is imaged because of the finite energy resolution of the scintillation camera
system. The pulse heights from the absorptions of identical gamma rays vary enough to cause slight
minifications and magnifications, ultimately degrading spatial resolution. The solution to this problem
is to normalize the position signals with the measured energy signal. This removes the image size
dependence with energy, thereby improving spatial resolution and allowing the simultaneous imaging
of more than one radionuclide without distortion. This feature is the primary component for
guaranteeing good multiwindow spatial registration.
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27.6 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
As has been previously noted, gamma rays that are scattered within the patient have distorted
spatial information and degrade image contrast. Because scattered gamma rays necessarily lose
energy, they can be selectively avoided by accepting only events that have pulse heights
corresponding to the primary gamma ray energy. The pulse height analyzer provides this capability.
A ìwindowî is centered to cover 15 to 20 percent of the photopeak. All energy pulses that meet this
criterion generate a logic pulse that indicates to the system that a valid event has occurred. This output
enables the recording of the x and y position information in a computer or display.
The precision of locating gamma ray events by a scintillation camera is referred to as intrinsic spatial
resolution. The original Anger camera had very poor intrinsic spatial resolution (~12 mm full-widthhalf-maximum). With the improvement in electronics and pulse processing methods, the spatial
resolution of the scintillation camera has improved steadily, approaching 3 mm in modern systems.4
These improvements include: better-quality, low-noise PMTs; improvements in the Anger logic
electronics including signal thresholding; smaller PMTs, and improved PMT quality control. The most
recent improvement has been the replacement of most of the analog processing with digital electronics.
With improvements in the speed of digitization electronics and decreases in component costs, the
trend in scintillation cameras has been to digitize the PMT signals (Fig. 27.4).5 The analog-to-digital
converters assign a numeric magnitude to the PMT signals. All subsequent determinations of energy
and positions can then be done by computer algorithms that can accurately model the nonlinear
behavior of the PMT signals with source position. The success of this approach has allowed the
reduction and even total elimination of the light pipe. This moves the PMTs closer to the scintillation,
thereby improving both the precision of the position determination (i.e., spatial resolution) and the
energy resolution.
FIGURE 27.4 Digital scintillation cameras. By digitizing the output of the
PMTs, the analog weighting electronics can be replaced by a nonlinear positioning algorithm. This allows a more accurate correction, culminating
with the elimination of the light pipe.
Once there is an x and y coordinate that locates a valid event, this information has to be stored as
image data. Although it is possible on some scintillation camera systems to store the individual
coordinates sequentially (referred to as list mode acquisition), most systems store the information
directly in histogram or matrix mode. With this method, an array of computer memory, typically 128
◊ 128 or 256 ◊ 256, is reserved for each image frame. The matrix elements or pixels are initially set
to 0. The coordinates for each event point to one of the pixels and this pixel is incremented by 1.
When the acquisition-stopping criteria are met, the image is complete. The information in the matrix
is either gray-scale or color encoded to display the image data. The entire process is shown
schematically in Fig. 27.5. A gamma ray originating in the patient is absorbed in the NaI(Tl) crystal.
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.7
FIGURE 27.5 Scintillation camera. The scintillation camera processes each
detected event to determine the x, y, and energy. If an event falls within the
selected energy range, the memory location pointed to by the x and y coordinates is incremented. This process continues until the stopping criteria
(number of counts or acquisition time) are met.
The light from the scintillation is sampled by the PMT array, which determines both the x and y
coordinates of the event and its energy. If the energy signal falls within the window of the pulse
height analyzer, the x and y coordinates are used to increment the appropriate pixel. This process is
repeated for every detected event.
In order to form images with a scintillation camera, a collimator must be placed in front of the
NaI(Tl) crystal. The collimator (Fig. 27.6) is the image-forming aperture of the camera system, and
it is necessary for the imaging process. The collimator projects the gamma ray pattern originating in
the patient onto the NaI(Tl) crystal. It does this by selectively absorbing gamma rays. The collimator
is a close-packed array of holes in a lead plate. Most often the holes are parallel, but fanbeam
converging and diverging collimators are available. Gamma rays whose trajectory takes them through
a hole get to interact with the NaI(Tl). All the others are absorbed. The design of collimators depends
on the gamma ray energy and the ever-present trade-off between count sensitivity and spatial
resolution.6,7 Collimators used for imaging Tc-99m typically have holes that are 1 to 1.5 mm across
and are 20 to 40 mm thick. Typical collimator design parameters are given in Table 27.2.
Although a collimator is necessary for the formation of images, it represents the limiting factor in
the both count sensitivity and spatial resolution of the scintillation camera. Because of the bruteforce
absorption approach to forming images with collimators, they are very inefficient. Less than 1 in
5000 gamma rays that hit the front surface of the collimator get through to the crystal. To improve
the count sensitivity, the collimator hole size could be increased and the hole length shortened.
Unfortunately, these changes degrade the spatial resolution. The spatial resolution of the collimator is
constrained by the geometry of the holes and is typically in the range of 6 to 8 mm at 10 cm when
used with Tc-99m. This is the dominant factor in determining the overall system resolution, since the
intrinsic spatial resolution is in the range of 3 to 4 mm.
One very important property to remember about collimators is that the spatial resolution gets
worse as the source-to-collimator distance increases. This is illustrated in the set of phantom images
that were acquired from 5 to 30 cm from the collimator surface. To obtain the best-quality images,
spatial resolution comes at the price of count sensitivity; therefore it is crucial to keep the collimator
as close to the patient as possible.
The modern scintillation camera has improved performance because of improvements in the
components and electronics. The availability of digital electronics has allowed the elimination of the
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27.8 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
FIGURE 27.6 Collimation. The collimator is the image-forming aperture of the scintillation
camera. It projects an image of the radionuclide distribution onto the NaI(Tl) crystal by brute
force absorption of all gamma rays except those whose trajectory takes them through the holes.
The collimator is also the limiting factor of both spatial resolution and count sensitivity. The
spatial resolution significantly degrades with source-to-collimator distance.
light pipe, which improves both energy and spatial resolution. However, this requires additional
corrections because of the nonlinear response of the PMT array to the scintillations. If a collimated
point source were focused on a portion of the NaI(Tl) crystal located exactly on a photomultiplier
tube center, the energy spectrum would be distinctly different than one that was acquired from a
point between two tubes (Fig. 27.7). This difference reflects the efficiency for collecting all the
TABLE 27.2 Collimator Specifications
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.9
scintillation light. When the source is located directly on a photomultiplier tube, more of the
scintillation is sampled by the photomultiplier tubes, and the pulses are therefore somewhat larger on
the average than at other points. This position-dependent shift in the energy spectrum causes an
overall loss in energy resolution. It also means that portions of the crystal will be accepting
proportionately more scattered radiation. The solution to this problem is to locally sample the energy
spectra and regionally adjust the energy window for each area. Typically the camera field of view is
divided into a 64 ◊ 64 matrix and energy window adjustments are made for each of the 4096 regions.
FIGURE 27.7 Energy correction. Because of the spatial arrangement of the
PMTs, the magnitude of the energy signal varies with location, degrading the
overall energy resolution. This problem is overcome by setting multiple local
energy windows across the field of view. This energy correction does not improve uniformity, but it does remove the dependence of the scintillation camera on scatter conditions.
Figure 27.7 shows the effect of energy correction when the scintillation camera is exposed to a
uniform flux of gamma rays. First, it should be noted that both the corrected and uncorrected images
are highly nonuniform and are not adequate for imaging. The energy correction simply makes sure
that each region of the crystal is contributing valid photopeak events to the image. This results in only
a subtle improvement in uniformity at this stage. However, it makes the subsequent corrections more
robust, since there will be much less dependence on the effects of scattered radiation, which can vary
over a large range, depending on the imaging situation.
Because of the nonlinear response of the photomultiplier tubes, detected events are not correctly
positioned using Anger logic alone. This mispositioning of events has a profound effect on field
uniformity. 8 The parameter that quantifies how well-detected events are positioned is called spatial
linearity. The optimization of spatial linearity requires the acquisition of an image from a welldefined
distribution (Fig. 27.8). Typically this is accomplished with a highly precise rectangular hole pattern
that is placed directly on the NaI(Tl) crystal. A distant point source of radioactivity is used to project
an image of the hole pattern onto the scintillation camera. The image of this pattern appears similar
to the image on the left with distortions caused by the mispositioning of events. Because the actual
and measured location of the holes is known, regional displacements to the x and y coordinates can
be calculated for each hole. Displacements for regions in between holes that are not directly sampled
are interpolated at a very high sampling frequency (1024 ◊ 1024). This information is stored and is
available as a lookup table. This measurement is usually done by the vendor at the factory and may
be repeated several times a year. When a valid event is detected, the initial x and y coordinates are
modified by the appropriate displacements that are read from the lookup table. Using this approach,
events can be accurately positioned to better than 0.5 mm. The improvement in spatial linearity has
a profound effect on field uniformity. Both images show the response of the scintillation camera to a
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27.10 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
uniform flux of gamma rays. With spatial linearity correction, the field becomes uniform to within
±10 percent of the mean image counts. This level of uniformity is adequate for most conventional
imaging.
FIGURE 27.8 Spatial linearity correction. Residual positioning errors
are corrected by imaging a precision hole phantom. A correction factor
table is generated with the appropriate x and y offsets to reposition events
to their correct location. The application of spatial linearity correction
has a profound effect on image uniformity.
There are still some residual nonuniformities that exist in the scintillation camera even after energy
and spatial linearity correction have been applied. These can be further reduced by applying
uniformity correction (Fig. 27.9). Typically, a high count flood is acquired and a map of the
nonuniformities is stored in a memory buffer. During acquisition, the number of valid events that is
acquired is modulated by this map to ensure uniformity. With this additional correction, the field
uniformity can be reduced to within ±3 percent of the mean image counts. It should be noted that
field uniformity can be degraded by a number of factors, including the energy of the gamma ray.
The most crucial factor for a system that is operating properly is the setting of the energy window.
Figure 27.9 illustrates the dependence of uniformity with the energy window setting. Some
scintillation cameras are more forgiving than others, but all show more nonuniformity when the energy
window is displaced from the center of the photopeak.9 There is often a gamma ray energy dependence
as well. Most scintillation cameras are optimized for the best performance for the 140-keV gamma rays
of Tc-99m. In some systems, uniformity significantly degrades at other gamma ray energies.
Photomultiplier tubes are relatively unstable components. Their performance changes as they age
and is also sensitive to variations in temperature and humidity. In order for the energy, spatial
linearity, and uniformity corrections to remain valid, there must be some way of maintaining the
photomultiplier tubes at a constant operating point. Most scintillation camera systems have
photomultiplier tube stabilization firmware that dynamically adjusts the photomultiplier tubes in
response to a known reference signal. Some vendors use a constant-output light-emitting diode inside
the electronics housing that flashes 10 times per second. The individual photomultiplier tube signals
from these calibration flashes are sensed by electronics that can compare the measured output to the
desired value, and then make appropriate adjustments to maintain the operating point. Another
approach uses the ratio between the count rates in a photopeak and scatter window to maintain
constant photomultiplier tube response. Photomultiplier tubes are also very sensitive to magnetic
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.11
FIGURE 27.9 Uniformity correction. Nonuniformities in the field that remain after spatial linearity
correction are corrected from the acquisition of a high count reference flood image. In many cameras, the
field uniformity degrades when the energy window is not centered on the photopeak.
fields, and changes in the orientation of the camera with respect to the earthís magnetic field are
enough to cause measurable changes in field uniformity.10 To reduce this effect, each photomultiplier
tube is individually surrounded with mu-metal shielding.
The emission of gamma rays from a radioactive source has an exponential distribution. This
means that for any particular event rate, short intervals between events occur much more often than
long intervals. Because the scintillation light persists for a finite time, there will eventually be light
emitted from more than one event as the count rate increases. In a scintillation camera, pulse pileup
becomes evident at count rates as low as 20,000 counts per second (cps), and it gets increasingly
worse as the count rate increases. Since information about pulse height becomes compromised, the
performance of the scintillation camera degrades at high count rates. In most conventional and
SPECT imaging, the count rate is low enough that pulse pileup is not a major concern. However, high
count rates are encountered in some first-pass studies, and it is the primary problem in coincidence
imaging. Corrections can be made for pulse pileup, since the physics of scintillations is well known
(Fig. 27.10). Pileup can be detected on the basis of pulse height analysis. If the pileup is the result of
only two events, the contribution to the second pulse from the first can be accurately estimated and
subtracted, thus preserving both events. Multiple pileups can be identified and discarded. The count
rate performance of scintillation cameras has improved dramatically in recent years because of the
demands of coincidence imaging. When a conventional scintillation camera is recording count data at
120,000 cps, it is losing about 20 percent of the valid events. This loss increases with increasing count
rate. In addition to the loss of sensitivity, both uniformity and spatial resolution get progressively
worse as the count rate increases. Finally, it should be recognized that Anger logic will produce
artifacts at very high count rates. ìVirtualî sources will appear midway between real sources because
of the signal averaging.11 This problem has been addressed in coincidence systems by using maximum
likelihood estimation instead of the conventional Anger logic.
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27.12 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
FIGURE 27.10 Pulse pileup. The persistence of the scintillation causes pulse pileup when the event rate
exceeds 50,000 counts/s. Techniques can be applied to recover some of the lost information. This is especially important for coincidence imaging. Note also that the Anger positioning algorithm causes ìghostî
images when the count rate becomes high.
The performance of scintillation cameras is often specified in terms of the spatial resolution.
Spatial resolution is a measure of image blur and it is often specified by the width of the pointspread
function. On a perfect imaging system, the image of a point has no dispersion so that all the counts
fall at the same point. On a real imaging system the point is blurred and a count profile through it has
a gaussian shape. The width of this count profile at the half-maximum level (FWHM) is a commonly
used method for specifying the spatial resolution. The imaging system will more or less blur every
point in a similar way, leading to a loss of contrast. There are two spatial resolution parameters that
are of interest with scintillation cameras. The intrinsic spatial resolution describes how precisely the
event location is determined when a gamma ray interacts with the crystal. The extrinsic or system
spatial resolution combines the effect of the collimation with the intrinsic resolution. The intrinsic
spatial resolution varies from about 3 mm FWHM on systems with thin crystals and 2-in PMTs to
about 4.5 mm on systems with thicker crystals and 3-in PMTs.
The system spatial resolution depends on a variety of factors including the gamma ray energy,
collimation, the source to collimator distance, and the intrinsic spatial resolution. Higher-energy
gamma rays require thicker septal walls, which limits the resolving power of the collimator. As
indicated above, there is a strong dependence on the distance that the source is from the collimator.
The collimator and intrinsic resolution combine in quadrature like the sides of a right-angle triangle
to yield the system resolution. The collimator resolution is generally more than 50 percent larger than
the intrinsic resolution and therefore is the dominant factor. The system resolution at 10 cm ranges
from 6.5 mm for an ultrahigh-resolution collimator to about 9.5 mm for a general-purpose
collimator. It should be noted that there is nearly a factor of 3 loss of count sensitivity with the
ultrahigh-resolution collimator compared to the general-purpose collimator.
Scintillation cameras have the capability of simultaneously imaging different-energy gamma rays.
Most scintillation cameras handle at least three and many can handle six or more energy windows. It
is important that there is no significant distortion of the images obtained at the different energies. The
parameter that monitors the correspondence between images acquired at different energies is referred
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.13
to as multiwindow spatial registration. The multiwindow spatial registration should be less than 3 mm
on a modern scintillation system.
In the future there will be further improvements in scintillation camera technology. 12 These
include photomultiplier tube replacements, new scintillators, and solid-state alternatives.
Photomultiplier tubes are expensive, bulky, and susceptible to drifting. Avalanche photodiodes are a
possible replacement for photomultiplier tubes (Fig. 27.11). The photodiodes use solid-state
components that are very compact and rugged. Photomultiplier tubes still have superior performance
specifications and are better matched for NaI(Tl). But considerable progress has been made with the
photodiode approach and special-purpose devices that use this technology are beginning to appear.
Figure 27.11 shows a schematic for a miniature camera module that uses avalanche photodiodes. The
scintillator used in this application is CsI(Tl). The scintillation light emitted for CsI(Tl) is better
matched to the properties of the photodiode. Because of the compactness offered by such designs,
scintillation cameras can be designed in novel ways for specific applications. One potential
application is breast cancer imaging, where several of these devices could be arranged around the
breast to collect SPECT data more efficiently than bulky conventional scintillation cameras.
Solid-state detectors directly convert the absorbed gamma ray energy into collection of electric
charge and do not need photomultiplier tubes. Since the PMT is a bulky and expensive component,
this represents a significant breakthrough. Cadmium zinc telluride is an attractive solid-state detector13
(Fig. 27.12). It can be manufactured in a pixelated array and has comparable gamma ray detection
FIGURE 27.11 PMT replacements. PMTs are an expensive and bulky component of the
scintillation camera. Future designs may incorporate avalanche photodiodes.
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
27.14 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
FIGURE 27.12 Solid state imagers. Although NaI(Tl)-based scintillation cameras will be the choice for the near future, improvements in solid-state detectors
such as CdZnTe may lead to competitive imaging systems.
efficiency to NaI(Tl) at 140 keV. Another advantage is that the energy resolution with CdZnTe is
nearly a factor of 2 better than that of NaI(Tl). However, several production problems need to be
solved before CdZnTe becomes a viable replacement for NaI(Tl).
27.3 SPECT SYSTEMS
Single-photon-emission computed tomography (SPECT) produces tomographic images of the internal distribution of radiopharmaceuticals. 14,15 It is most commonly used in the diagnosis of coronary
artery disease and in tumor detection. Projection images collected by one or more scintillation
cameras are mathematically reconstructed to obtain the tomographic slices. Most clinical SPECT
studies are qualitative with simplistic corrections for attenuation and scattered radiation. Quantitative
SPECT requires corrections for attenuation, scatter, and spatial resolution, although these have not
been routinely implemented in the past because of their computational load. SPECT instrumentation
has evolved to include coincidence imaging of positron-emitting radiopharmaceuticals, specifically
18
F fluorodeoxyglucose.
A SPECT system consists of one or more scintillation cameras mounted to a gantry that can
revolve about a fixed horizontal axis (the axis of rotation) 16ñ19 (Fig. 27.13). SPECT studies are
usually acquired over a full 360∞ arc, although myocardial perfusion studies typically use only data
from the 180∞ arc that minimizes tissue attenuation. SPECT acquisitions are performed with the
scintillation camera located at preselected angular locations (step-and-shoot mode), or in a
continuous rotation mode. In the step-and-shoot mode, the detector rotates to each angular position
and collects data in a preselected frame duration while the detector is motionless. In the continuousrotation mode, the study duration is selected and the rotation speed is adjusted to complete the orbit
during this time. Projections are collected as the detector rotates and are binned into 60 to 120
frames over 360∞.
It is crucial to maintain close proximity to the body as the detector rotates about the patient to
achieve the best possible spatial resolution. Although a number of different approaches have been
used to accomplish this, the most common method moves the detectors radially in and out as a
function of rotation angle. Myocardial perfusion studies are the most commonly performed SPECT
procedures. Because the heart is located in the left anterior portion of the thorax, gamma rays
originating in the heart are highly attenuated for views collected from the right lateral and right
posterior portions of the arc. For this reason, SPECT studies of the heart are usually collected using
the 180∞ arc that extends from the left posterior oblique to the right anterior oblique view.20 This
results in reconstructed images with the best contrast, although distortions are often somewhat more
pronounced than when 360∞ data are used.21 Because of the widespread use of myocardial perfusion
imaging, many SPECT systems have been optimized for 180∞ acquisition by using two detectors
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.15
FIGURE 27.13 SPECT acquisition. One or more scintillation
cameras collect images at typically 60 to 120 angles around a
360∞ orbit. The scintillation camera acquires projection images
from a large volume simultaneously.
arranged at ~90∞ (Fig. 27.14). This reduces the acquisition time by a factor of 2 over single detectors
and is approximately 30 percent more efficient than triple detector SPECT systems. Positioning the
detectors at 90∞ poses some challenges for maintaining close proximity. Most systems rely on the
motion of both the detectors and the SPECT table to accomplish this.
The heart is continually moving during the SPECT acquisition, and this further compromises
spatial resolution. Because the heart beats many times per minute, it is impossible to directly acquire
FIGURE 27.14 SPECT system configurations. Although a single scintillation camera can be
used to acquire SPECT data, multiple detectors improve the overall sensitivity. Two detectors
arranged at either 180 or 90∞ are the most common configuration.
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
27.16 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
a stop-action SPECT study. However, since the heart motion is periodic, it is possible to obtain this
information by gating the SPECT acquisition. 22 In a gated SPECT acquisition, the cardiac cycle is
subdivided and a set of eight images spanning the ECG R-R interval is acquired for each angular
view. These images are acquired into predetermined time bins based on the patientís heart rate, which
is monitored by the ECG R wave interfaced to the SPECT system. As added benefits of gating, the
motion of the heart walls can be observed and ventricular volumes and ejection fractions can be
determined. 22,23
Although most SPECT imaging samples more or less static distribution of radionuclides, some
SPECT systems can perform rapid sequential studies to monitor tracer clearance. An example of this
is determination of regional cerebral blood from the clearance of 133 Xe.24 Multiple 1-minute SPECT
studies are acquired over a 10-minute interval. When one acquisition sample is completed, the next
begins automatically. In order to minimize time, SPECT systems that perform these studies can
alternately reverse the acquisition direction, although at least one SPECT system utilizes slipring
technology so that the detectors can rotate continuously in the same direction.
In order to produce accurate tomographic images, projection data representing the line integrals
of activity in the internal distribution have to be acquired. This information is not directly available
because of tissue attenuation. Simple attenuation correction methods can be used in regions of the
body such as the abdomen or head where the tissue density is more or less uniform. However,
compensation for attenuation in the thorax requires an accurate attenuation map for each
tomographic plane. This is especially important for myocardial perfusion studies, since the artifacts
resulting from tissue attenuation mimic the patterns caused by coronary artery disease. In recent
years, all the SPECT manufacturers have offered systems that can perform transmission measurements
along with the emission studies. These systems use the scintillation camera to detect the transmission
of gamma rays from an external source.17,25,26 Several different configurations are available. Most use
a line source that is translated across the camera field of view at each angular stop in much the same
way as a first-generation CT scanner (Fig. 27.15). Typically a source such as Gd-153 or Ba-133 is
FIGURE 27.15 Attenuation correction. To obtain accurate SPECT results, corrections must be made for
tissue attenuation. In regions where there are large variations in tissue density such as the thorax, this
requires an independent transmission study. This shows one possible configuration where line sources
of Gd-153 are translated across the field of view to collect the transmission data. This information is
reconstructed to obtain a crude CT image of the thorax to correct myocardial perfusion studies. (Courtesy of GE Medical Systems.)
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.17
used that has a different gamma ray energy than those used in the emission study so that the studies
can be acquired simultaneously (with appropriate correction for cross talk). Other approaches use
multiple stationary line sources to collect this information. At least one vendor has an x-ray tube and
separate detectors to obtain a moderate-quality CT scan.27 In each configuration, the transmission data
are collected and reconstructed to yield the attenuation maps. This information can be incorporated
into an iterative algorithm to effect the correction.
In spite of the energy discrimination available on all SPECT systems, Compton scattered radiation
still accounts for about 30 to 40 percent of the acquired counts in SPECT imaging. Scattered radiation
decreases contrast and can impact other corrections. For example, when attenuation correction is
applied without also correcting for scattered radiation, the count density in the heart walls near the
liver may be overenhanced. SPECT systems in the future may resort to other detectors that have
substantially better energy resolution than that of NaI(Tl), but for now, scatter compensation routines
must be employed. Scatter correction has been performed in several different ways.15,28ñ33 The easiest
to implement is the subtraction method, where information is simultaneously acquired into a second
energy window centered below the photopeak in the Compton scatter region of the energy spectrum.
After establishing an appropriate normalization factor, the counts from the scatter window are
subtracted from the photopeak window. The corrected projections are then used in the reconstruction
algorithm. The disadvantage of this approach is that it increases noise and it is difficult to establish an
accurate normalization factor. To accommodate this type of correction (and also to image differentenergy gamma rays), SPECT systems allow the simultaneous acquisition from multiple energy
windows. The number of energy windows available varies for each manufacturer, although every
system is capable of imaging from at least four energy windows.
27.3.1 SPECT Image Reconstruction
The details of SPECT image reconstruction are beyond the scope of this article. However, the
demands of image reconstruction do impact the features required by the computer. The CPU must be
fast enough and have enough memory to accommodate the entire SPECT data set. This is well within
the capability of home PCs. A typical SPECT study is less than 5 Mbyte and the current processor
speeds approaching 1 GHz are fast enough to render reconstructions using either filter backprojection
or optimized iterative algorithms in an acceptable time (less than 10 minutes).
27.3.2 SPECT System Performance
Typical performance specifications for SPECT imaging systems are summarized in Table 27.3. As
with conventional planar imaging, the scintillation cameras, and the associated collimation are the
primary factors affecting the performance. SPECT spatial resolution is nearly isotropic with an
FWHM of 8 to 10 mm for brain imaging where the detectors can get close to the radioactive source.
TABLE 27.3 SPECT System Performance
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
27.18 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
The spatial resolution degrades to 12 to 18 mm for body imaging because the detectors can not be
positioned as close. The components of SPECT spatial resolution and their relative importance can be
identified from the equation shown below:
As before, R int and R col represent the intrinsic and collimator resolution components. R filter is the
FWHM of the smoothing kernel required to yield an acceptable reconstruction. The intrinsic spatial
resolution is the least important factor in this calculation, since it is usually a factor of 2 or more
smaller than the other components. The trade-off between spatial resolution and count sensitivity is
explicit in this equation. Decreasing R col to improve spatial resolution will often require R filter to
become larger to compensate for increased noise.
27.3.3 SPECT/PET Hybrid Systems
The primary reason for the success of nuclear medicine imaging is the availability of
radiopharmaceuticals that provide crucial diagnostic information. For cancer diagnosis and followup,
18
F flourodeoxyglucose ( 18F FDG) is an exquisite imaging agent for a wide variety of malignancies
including lung, colon, breast, melanoma, and lymphoma. Because 18F is a positron emitter that yields
very high energy x-rays (511 keV) when the positron combines with a free electron, it can not be
imaged well on conventional SPECT systems. The thin NaI(Tl) crystals have a low efficiency for
detection at this energy (less than 10 percent). Also, the collimators designed for the 511-keV photons
have low count sensitivity and poor spatial resolution. The collimation can be eliminated if coincidence detection is used. Annihilation photons from positrons are always colinear. This feature can be
exploited to count only events that are simultaneously detected by opposed detectors. Two opposed
scintillation cameras with their collimators removed and the addition of coincidence electronics will
turn a SPECT system into a PET tomograph (Fig. 27.16).34,35 The efficiency for coincidence detection
equals the product of the individual efficiencies, so that the coincidence efficiency is about 1 percent
for conventional scintillation cameras. This is actually much higher than the efficiency with collimators. Coincidence efficiency can be improved by using thicker NaI(Tl) crystals, and all the vendors
have done this. However, thicker crystals degrade the intrinsic spatial resolution when the scintillation
FIGURE 27.16 Two or more opposed scintillation cameras can be used as coincidence detection systems for PET imaging. Valid events are established
when the two detectors record events within the 10- to 15-ns timing window.
Graded filters and lead septa are placed in front of the detectors to limit scattered radiation. (Courtesy of GE Medical Systems.)
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
NUCLEAR MEDICINE IMAGING INSTRUMENTATION 27.19
cameras are used for conventional (i.e., collimated) studies. A solution to this problem, called
StarBrite, has been recently introduced. Thick NaI(Tl) crystals (25 mm) have slots machined in the
PMT side of the detector to prevent the scintillation light from scattering throughout the crystal to
maintain good spatial resolution. There are other concerns, primarily the count rate capability. With
most SPECT imaging studies, there are essentially no count rate losses resulting from the time it takes
to process each event. The amount of radioactivity that can be safely administered and the low
sensitivity of the collimation keep the count rate in the range that the electronics can easily handle.
However, when the collimators are removed for coincidence imaging, the NaI(Tl) crystals are exposed to count rates above the capacity of the conventional electronics. Much effort over the past
decade has been devoted to increasing the count rate capability of the scintillation cameras. In the
early 1990s, the maximum observed counting rate for a scintillation camera was in the 200,000 to
400,000 count/second range. This rate has been extended to over 1,000,000 counts/second by
shortening the integration time on the pulses and implementing active baseline restoration. Because
the light is proportional to the energy deposited in the crystal, one can shorten the pulse integration
time without extreme degradation. Even with this substantial improvement in count rate, the maximum activity the system can handle is about 3 mCi. Typical performance values for a coincidence
scintillation camera system are given in Table 27.4.
TABLE 27.4 SPECT/PET Hybrid System Performance
In addition to specialized electronics, other measures have been taken to help reduce the count
rate burden. One example of these is a graded absorber placed in front of each detector to help
reduce the scattered radiation component. 36 Because scattered radiation has lower energy than the
unscattered photons, low-energy scatter will be preferentially absorbed by lead, since the
photoelectric cross section is inversely proportional to the cube of the gamma ray energy. If only lead
is used, the lead characteristic x-ray resulting from the absorption will be detected. Additional layers
of tin, copper, and aluminum will absorb the respective characteristic x-rays. This graded filter causes
a significant reduction in the low-energy scattered radiation and, since the scintillation cameras
process every event, also reduces the overall count rate burden.
Even though the thin NaI(Tl) crystals have low intrinsic efficiency, the uncollimated detectors
present a large solid angle to the annihilation photons. Maximum sensitivity is achieved when all
coincidences, even those at large angles, are accepted. This makes the camera sensitivity extremely
sensitive to the source location. Sources located near the central axis of the detectors have a large
solid angle, while those at the periphery can interact only with a very small portion of the detectors.
In addition, the scatter component increases to well over 50 percent when large coincidence
acceptance angles are used. This problem has been addressed by using lead slits aligned perpendicular
to the axial direction of the system to restrict the angular extent of the coincidences. While this
reduces the sensitivity of the imaging system, it also reduces the scatter component to less than 30
percent and limits the solid angle variability.
The intrinsic spatial resolution of the hybrid systems is comparable to that of the dedicated PET
systems with a FWHM of 4 to 5 mm. However, the count sensitivity is at least an order of magnitude
lower. This, along with the maximum count rate constraint, guarantees that the coincidence camera
data will be very count poor and therefore require substantial low-pass filtering when reconstructed.
As a result the quality of the reconstructed images is perceptibly worse than the dedicated PET
images. In head-to-head comparisons, it has been found that the hybrid systems perform well on
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NUCLEAR MEDICINE IMAGING INSTRUMENTATION
27.20 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
tumors greater than 2 cm in diameter located in the lung.37ñ39 Tumors smaller than these and those
located in high-background areas are detected with a much lower sensitivity. These results are
important since they provide a guide where the application of the coincidence camera will be useful.
In a conventional scintillation camera, Anger logic is used to determine the location of an
interaction. The tacit assumption in this approach is that only one event is occurring at a time. At the
high count rate encountered in coincidence imaging, multiple interactions are likely, and when this
occurs, the events are improperly located somewhere between the two true locations. Improved
algorithms have been developed that can identify multiple hits and that use a maximum likelihood
calculation to correctly determine event locations.
The projection data collected by the coincidence cameras require correction for random
coincidences, scatter, and attenuation if accurate tomographic images are to be obtained. Random or
accidental coincidences occur when two unrelated photons are detected. These random coincidences
increase rapidly with the count rate and give rise to a nondescript background that has to be
subtracted from the projections. Typically randoms are either monitored in a separate time window or
are calculated from the singles count rate and are subtracted. Scatter correction is sometimes ignored
in coincidence PET, or maybe monitored by a separate energy window and subtracted as discussed
for SPECT imaging.
Accurate reconstruction of PET data requires correction for attenuation since the degree of
attenuation for coincidence imaging is very high, approaching values of 100 or more. Attenuation
correction requires information about the transmission of the annihilation radiation through the body
at the coincidence lines of response. When attenuation correction is ignored, severe artifacts are seen
in the reconstructed images. As with SPECT imaging, a separate transmission study with an external
source, typically Cs-137, must be acquired to provide the attenuation map used in the correction.
27.4 SUMMARY
SPECT imaging is expected to play a continuing important role in medical imaging. Future improvements in SPECT instrumentation are likely to include new detectors and collimation schemes. The
coincidence scintillation cameras will also continue their evolution with the addition of more cameras
and multidetector levels optimized for SPECT and coincidence imaging. Reconstruction algorithms
will evolve as new techniques are developed and as the performance of the computer expands. In
spite of the importance of instrumentation, the primary motivating factor in SPECT imaging will
continue to be the creation and implementation of new radiopharmaceuticals. While SPECT will
continue to be highly utilized for myocardial perfusion imaging, SPECT use in tumor imaging will
probably experience the largest growth. Applications will include treatment planning for internal
radiation therapy as well as diagnostic studies.
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27.22 DESIGN OF MEDICAL DEVICES AND DIAGNOSTIC INSTRUMENTATION
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