Nuclear Tomographic Imaging—Single Photon and Positron Emission Tomography (SPECT and PET)

The formation of projection images in nuclear medicine was discussed in the previous chapter. A nuclear medicine projection image depicts a two-dimensional projection of the three-dimensional activity distribution in the patient. The disadvantage of a projection image is that the contributions to the image from structures at different depths overlap, hindering the ability to discern the image of a structure at a particular depth. Tomographic imaging is fundamentally different—it attempts to depict the activity distribution in a single cross section of the patient.

There are two fundamentally different types of tomography: conventional tomography, also called geometric or focal plane tomography, and computed tomography. In conventional tomography, structures out of a focal plane are not removed from the resultant focal plane image; instead, they are blurred by an amount proportional to their distances from the focal plane. Those close to the focal plane suffer little blurring and remain apparent in the image. Even those farther away, although significantly blurred, contribute to the image, thereby reducing contrast and adding noise. In distinction, computed tomography uses mathematical methods to remove overlying structures completely. Computed tomography requires the acquisition of a set of projection images from at least a 180° arc about the patient. The projection image information is then mathematically processed by a computer to form images depicting cross sections of the patient. Just as in x-ray transmission imaging, both conventional and computed tomography are possible in nuclear medicine imaging. Both single photon emission computed tomography (SPECT) and positron emission tomography (PET) are forms of computed tomography.

19.1 FOCAL PLANE TOMOGRAPHY IN NUCLEAR MEDICINE

Focal plane tomography once had a significant role in nuclear medicine, but is seldom used today. The rectilinear scanner, when used with focused collimators, is an example of conventional tomography. A number of other devices have been developed to exploit conventional tomography in nuclear medicine. The Anger tomoscanner used two small scintillation cameras with converging collimators, one above and one below the patient table, to scan the patient in a raster pattern; a single scan produced multiple whole-body images, each showing structures at a different depth in the patient in focus. The seven-pinhole collimator was used with a conventional scintillation camera and computer to produce short-axis images of the heart, each showing structures at a different depth in focus. The rectilinear scanner and the Anger tomoscanner are no longer produced. The seven-pinhole collimator, which never enjoyed wide acceptance, has been almost entirely displaced by SPECT.

The gamma camera itself, when used for planar imaging with a parallel-hole collimator, produces a weak tomographic effect. The system’s spatial resolution decreases with distance, causing structures farther from the camera to be more blurred than closer structures. Furthermore, attenuation of photons increases with depth in the patient, also enhancing the visibility of structures closer to the camera. This effect is perhaps most clearly evident in planar skeletal imaging of the body. In the anterior images, for example, the sternum and anterior portions of the ribs are clearly shown, whereas the spine and posterior ribs are barely evident.

19.2 SINGLE PHOTON EMISSION COMPUTED TOMOGRAPHY

19.2.1 Design and Principles of Operation

SPECT generates transverse images depicting the distribution of x- or γ-ray-emitting nuclides in patients. Standard planar projection images are acquired from an arc of 180° (most cardiac SPECT) or 360° (most non-cardiac SPECT) about the patient. Although these images could be obtained by any collimated imaging device, the vast majority of SPECT systems use one or more camera heads that revolve around the patient. The SPECT system’s digital computer then reconstructs the transverse images using either filtered backprojection or an iterative reconstruction method, which are described later in this chapter, as does the computer in an x-ray CT system. Figure 19-1 shows a variety of SPECT systems.

SPECT was invented by David Kuhl and others in the early 1960s, about 10 years before the invention of x-ray CT by Hounsfield (Kuhl and Edwards, 1963). However, in contrast to x-ray CT, most features of interest in SPECT images were also visible in planar nuclear medicine images and SPECT did not come into routine clinical use until the late 1980s.

Image Acquisition

The camera head or heads of a SPECT system revolve around the patient, acquiring projection images. The head or heads may acquire the images while moving

(continuous acquisition) or may stop at predefined evenly spaced angles to acquire the images (“step and shoot” acquisition). If the camera heads of a SPECT system produced ideal projection images (i.e., no attenuation by the patient and no degradation of spatial resolution with distance from the camera), projection images from opposite sides of the patient would be mirror images, and projection images over a 180° arc would be sufficient for transverse image reconstruction. However, in SPECT, attenuation greatly reduces the number of photons from activity in the half of the patient opposite the camera head, and this information is greatly blurred by the distance from the collimator. Therefore, for most non-cardiac studies, such as bone SPECT, the projection images are acquired over a complete revolution (360°) about the patient. However, most nuclear medicine laboratories acquire cardiac SPECT studies, such as myocardial perfusion studies, over a 180° arc symmetric about the heart, typically from the 45° right anterior oblique view to the 45° left posterior oblique view (Fig. 19-2). The 180° acquisition produces reconstructed images of superior contrast and resolution because the projection images of the heart from the opposite 180° have poor spatial resolution and contrast due to greater distance and attenuation. Although studies have shown that the 180° acquisition can introduce artifacts (Liu et al., 2002), the 180° acquisition is more commonly used than the 360° acquisition for cardiac studies.

▪ FIGURE 19-1 A. SPECT/CT system with two scintillation camera heads in a fixed 180° orientation and a non-diagnostic x-ray CT system for attenuation correction and anatomic correlation. The x-ray source is on the right side of the gantry and a flat-panel x-ray image receptor is on the left. (Photo credit: Emi Manning, UC Davis Health System.)

▪ FIGURE 19-1 (Continued) B. Technologist moving the upper camera head closer to the patient for SPECT imaging. (Photo credit: Emi Manning, UC Davis Health System.) Dual head, variable angle SPECT/CT camera with heads in the 90° orientation (C) for cardiac SPECT and in the 180° orientation (D) for other SPECT or whole-body planar imaging. E. Dual head, fixed 90° SPECT camera for cardiac imaging. (© Siemens Healthineers 2019. Used with permission.) F. Single head SPECT camera, with head in a position for planar imaging. (Courtesy Emi Manning, UC Davis Health System.)

▪ FIGURE 19-2 180° cardiac orbit.

SPECT projection images are usually acquired in either a 64² or a 128² pixel format. Using too small a pixel format reduces the spatial resolution of the projection images and of the resultant reconstructed transverse images, due to too large physical pixel dimensions. (A zoom factor of ˜1.5 is often employed for 64² cardiac SPECT, to reduce the pixel dimensions, thereby improving spatial resolution.) When the 64² format is used, typically 60 or 64 projection images are acquired and, when a 128² format is chosen, 120 or 128 projection images are acquired. Using too few projections creates radial streak artifacts in the reconstructed transverse images.

The camera heads on older SPECT systems followed circular orbits around the patient while acquiring images. Circular orbits are satisfactory for SPECT imaging of the brain, but cause a loss of spatial resolution in body imaging because the circular orbit causes the camera head to be many centimeters away from the surface of the body during the anterior and perhaps the posterior portions of its orbit (Fig. 19-3). Modern SPECT systems provide non-circular orbits (also called “body contouring”) that keep the camera heads in close proximity to the surface of the body throughout the orbit. For some systems, the technologist specifies the non-circular orbit by placing the camera head as close as possible to the patient at several angles, from which the camera’s computer determines the orbit. Other systems perform automatic body
contouring, using sensors on the camera heads to determine their proximity to the patient at each angle.

▪ FIGURE 19-3 Circular (A) and body-contouring (B) orbits.

In brain SPECT, it is usually possible for the camera head to orbit with a much smaller radius than in body SPECT, thereby producing images of much higher spatial resolution. In many older cameras, a large distance from the physical edge of the camera head to the useful portion of the detector often made it impossible to orbit at a radius within the patient’s shoulders while including the base of the brain in the images. These older systems were therefore forced to image the brain with an orbit outside the patient’s shoulders, causing a significant loss of resolution. Most modern SPECT systems permit brain imaging with orbits within the patient’s shoulders, although a patient’s head holder extending beyond the patient table is generally necessary.

Transverse Image Reconstruction

After the projection images are acquired, they are corrected for non-uniformities and for center-of-rotation (COR) misalignments. (These corrections are discussed below, “Quality Control in SPECT.”) Following these corrections, transverse image reconstruction is performed using either filtered backprojection or iterative methods.

As described in Chapter 10, filtered backprojection consists of two steps. First, the projection images are mathematically filtered. Then, to form a particular transverse image (also known as a slice), simple backprojection is performed of the row of each projection image corresponding to that transverse image. For example, the fifth row of each projection image is backprojected to form the fifth transverse image. A SPECT study produces transverse images covering the entire field of view (FOV) of the camera in the axial direction from each revolution of the camera head or heads.

Mathematical theory specifies that the ideal filter kernel, when displayed in the spatial frequency domain, is the ramp filter (Fig. 19-4). (The spatial frequency domain is discussed in Appendix G, Convolution and Fourier Transforms.) However, the actual projection images contain considerable statistical noise, from the random nature of radioactive decay and photon interactions, due to the relatively small number of counts in each pixel. If the images were filtered using a ramp filter kernel and then backprojected, the resultant transverse images would contain an unacceptable amount of statistical noise.

In the spatial frequency domain, statistical noise predominates in the highfrequency portion. Furthermore, the spatial resolution characteristics of the gamma camera cause a reduction of higher spatial frequency information that increases with the distance of the structure being imaged from the camera. To smooth the projection images before backprojection, the ramp filter kernel is modified to “roll-off” at
higher spatial frequencies. Unfortunately, this reduces the spatial resolution of the projection images and thus of the reconstructed transverse images. A compromise must therefore be made between spatial resolution and the statistical noise of the transverse images.

▪ FIGURE 19-4 Typical filter kernels used for filtered backprojection. The kernels are shown in frequency space. Filter Kernel A is a Butterworth filter of fifth order with a critical frequency of 0.20 Nyquist and Filter Kernel B is a Butterworth filter of fifth order with a critical frequency of 0.30 Nyquist. Filter Kernel A provides more smoothing than Filter Kernel B. A ramp filter, which provides no smoothing, is also shown.

Typically, a different filter kernel is selected for each type of SPECT study; for example, a different kernel would be used for Tc-99m HMPAO brain SPECT than would be used for Tc-99m sestamibi myocardial perfusion SPECT. The choice of filter kernel for a particular type of study is determined by the amount of statistical noise in the projection images (mainly determined by the injected activity, collimator, and acquisition time per image) and their spatial resolution (determined by the collimator and distances of the camera head(s) from the organ being imaged). The preference of the interpreting physician regarding the appearance of the images also plays a role. Projection images of better spatial resolution and less quantum mottle require a filter with a higher spatial frequency cutoff to avoid unnecessary loss of spatial resolution in the reconstructed transverse images, whereas projection images of poorer spatial resolution and greater quantum mottle require a filter with a lower spatial frequency cutoff to avoid excessive quantum mottle in the reconstructed transverse images. Although the SPECT camera’s manufacturer may suggest filters for specific imaging procedures, the filters are usually empirically optimized in each nuclear medicine laboratory. Figure 19-5 shows a SPECT image created using three different filter kernels, illustrating too much smoothing, proper smoothing, and no smoothing.

Filtered backprojection is computationally efficient. However, it is based upon the assumption that the projection images are perfect projections of a three-dimensional object. As discussed in the previous chapter, this is far from true in gamma camera imaging, mainly because of attenuation of photons in the patient, the inclusion of Compton scattered photons in the image, and the degradation of spatial resolution with distance from the collimator.

In SPECT, iterative reconstruction methods are increasingly being used instead of filtered backprojection. In iterative methods, an initial activity (typically uniform) distribution in the patient is assumed. (Alternatively, an activity distribution created by simple backprojection could be used.) Then, projection images are calculated
from the initial assumed activity distribution, using a model of the imaging characteristics of the gamma camera and the patient. The calculated projection images are compared with the actual projection images and, based upon this comparison, the assumed activity distribution is adjusted. This process is repeated several times, with successive adjustments to the assumed activity distribution, until the calculated projection images approximate the actual projection images (Fig. 19-6).

▪ FIGURE 19-5 SPECT images created by filtered backprojection. The projection images were filtered using the filter kernels shown in Figure 19-4. The image on the left, produced using Filter Kernel A, exhibits a significant loss of spatial resolution. The image in the center was produced using Filter Kernel B, which provides a proper amount of smoothing. The image on the right, produced using the ramp filter, shows good spatial resolution, but excessive statistical noise.

As was stated above, in each iteration, projection images are calculated from the assumed activity distribution. The calculation of projection images can incorporate the system resolution point spread function (PSF) of the gamma camera, which takes into account the decreasing spatial resolution with distance from the camera face. If a map of the attenuation characteristics of the patient is available, the calculation of the projection images can include the effects of attenuation. Furthermore, the PSF can be modified to incorporate the effect of photon scattering in the patient. Alternatively,
modeling scatter within the photopeak based on either secondary energy window images (dual- or triple-energy-window method) or projection using the photopeak transverse images and attenuation and material density maps (effective scatter source estimation) is now more commonly applied. If all this is done, iterative methods will partially compensate for the effects of decreasing spatial resolution with distance, as well as attenuation and photon scattering in the patient. Iterative reconstruction can be used to produce higher quality tomographic images than filtered backprojection, or it can be used to produce images of similar quality to those produced by filtered backprojection, but with less administered activity or shorter acquisition times.

▪ FIGURE 19-6 Flowchart for iterative reconstruction. In some implementations, iterative reconstruction is performed for a specified number of iterations, instead of being terminated when a sufficiently good approximation is achieved.

Iterative methods are computationally less efficient than filtered backprojection. However, the increasing speed of computers, the small image matrix sizes used in nuclear imaging, and the development of computationally efficient algorithms, such as the ordered-subset expectation maximization method (Hudson and Larkin, 1994), have made iterative reconstruction feasible for SPECT. Since iteratively reconstructed SPECT transverse images will contain substantial noise due to relatively poor counting statistics, three-dimensional spatial filtering is commonly applied after reconstruction for noise reduction.

Attenuation Correction in SPECT

Radioactivity whose x- or γ-rays must traverse long paths through the patient produces fewer counts, due to attenuation, than does activity closer to the surface of the patient adjacent to the camera. For this reason, transverse slices of a phantom with a uniform activity distribution, such as a cylinder filled with a well-mixed solution of radionuclide, will show a gradual decrease in activity toward the center (Fig. 19-7, on the left). Attenuation effects are more severe in body SPECT than in brain SPECT.

Approximate methods are available for attenuation correction. One of the most common, the Chang method, presumes a constant attenuation coefficient throughout the patient (Chang, 1978). Approximate attenuation corrections can overcompensate or undercompensate for attenuation. If such a method is to be used, its proper functioning should be verified using phantoms before its use in clinical studies.

These methods are only appropriate for filtered backprojection reconstruction. Furthermore, attenuation is not uniform in the patient, particularly in the thorax, and these approximate methods cannot compensate for non-uniform attenuation.

▪ FIGURE 19-7 Attenuation correction. On the left is a reconstructed transverse image slice of a cylindrical phantom containing a well-mixed radionuclide solution. This image shows a decrease in activity toward the center due to attenuation. (A small ring artifact, unrelated to the attenuation, is also visible in the center of the image.) In the center is the same image corrected by the Chang method, using a linear attenuation coefficient of 0.12 cm^-1, demonstrating proper attenuation correction. On the right is the same image, corrected by the Chang method using an excessively large attenuation coefficient.

In the 1990s, several manufacturers provided SPECT cameras with sealed radioactive sources (commonly containing Gd-153, which emits 97 and 103-keV γ-rays) to measure the attenuation through the patient. The sources were used to acquire transmission data from projections around the patient. After acquisition, the transmission projection data were reconstructed to provide maps of tissue attenuation characteristics across transverse sections of the patient, similar to x-ray CT images. Finally, these attenuation maps were used during an iterative SPECT image reconstruction process to provide attenuation-corrected SPECT images.

The transmission sources were available in several configurations. These included scanning collimated line sources that were used with parallel-hole collimators, arrays of fixed line sources used with parallel-hole collimators, and a fixed line source located at the focal point of a fan-beam collimator.

The transmission data were usually acquired simultaneously with the acquisition of the emission projection data because performing the two separately can pose significant problems in the spatial alignment of the two data sets and greatly increases the total imaging time. The radionuclide used for the transmission measurements was chosen to have primary γ-ray emissions that differed significantly and were lower in energy from those of the radiopharmaceutical. Separate energy windows were used to differentiate the photons emitted by the transmission source from those emitted by the radiopharmaceutical. However, scattering of the higher energy emission photons in the patient and in the detector caused some cross-talk in the lower energy window.

Major manufacturers of nuclear medicine imaging systems now provide systems combining two camera heads capable of planar imaging and SPECT and an x-ray CT scanner, with a single patient bed. These systems have supplanted systems with radioactive transmission sources, and are referred to as SPECT/CT systems. In SPECT/CT systems, the x-ray CT attenuation image data can be used to correct the radionuclide emission data for attenuation by the patient. This is discussed in more detail later in this chapter.

Attenuation correction using radioactive transmission sources and x-ray CT-derived attenuation maps has been extensively studied in myocardial perfusion SPECT, where attenuation artifacts can mimic perfusion defects. These studies have shown that attenuation correction reduces attenuation artifacts and produces modest improvement in diagnostic performance when the studies are read by experienced clinicians (Hendel et al., 2002; Masood et al., 2005). However, other studies have shown that attenuation correction can cause artifacts, particularly when there is spatial misalignment of the emission data with respect to the attenuation maps determined from the transmission information. Furthermore, a period of transition is required, for even experienced clinicians to retrain themselves to interpret attenuation-corrected images. Therefore, it remains common for SPECT myocardial perfusion imaging to be performed without attenuation correction, although that may change with the increasing implementation of cardiac SPECT/CT.

Generation of Coronal, Sagittal, and Oblique Images

The pixels from the transverse slices may be reordered to produce coronal and sagittal slices. For cardiac imaging, it is desirable to produce oblique images oriented either parallel (vertical and horizontal long-axis images) or perpendicular (short-axis images) to the long axis of the left ventricle. Because there is considerable anatomic variation among patients regarding the orientation of the long axis of the left ventricle, the long axis of the heart must be determined before the computer can create the oblique images. This task is commonly performed manually by a technologist, although the software on most systems is now capable of correct automatic reorientation of myocardial perfusion images, with operator verification and override.

▪ FIGURE 19-8 Fan-beam collimator.

Collimators for SPECT

Most SPECT is performed using parallel-hole collimators. However, specialized collimators have been developed for SPECT. The fan-beam collimator, shown in Figure 19-8, is a hybrid of the converging and parallel-hole collimator. Because it is a parallel-hole collimator in the y-direction, each row of pixels in a projection image corresponds to a single transaxial slice of the subject. In the x-direction, it is a converging collimator, with spatial resolution and efficiency characteristics superior to those of a parallel-hole collimator (see Fig. 18-12). Because a fan-beam collimator is a converging collimator in the cross-axial direction, its FOV decreases with distance from the collimator. For this reason, the fan-beam collimator is mainly used for brain SPECT; if the collimator is used for body SPECT, portions of the body are excluded from the FOV, which can cause artifacts, called “truncation artifacts,” in the reconstructed images. One manufacturer offers a variable-focal-length converging collimator for cardiac imaging, where the focal length increases from the center outward, ending up as a parallel-hole collimator at the edge (to eliminate truncation artifacts), in both the transaxial and axial directions. Along with a heart-centric camera head orbit, an approximate factor of four improvement in sensitivity in the region of the heart is achieved (Fig. 19-9).

Multihead SPECT Cameras

To reduce the limitations imposed on SPECT by collimation and limited time per view, camera manufacturers provide SPECT systems with two camera heads that revolve around the patient (Fig. 19-1) and, in the past, SPECT systems with three heads were commercially available from at least two manufacturers. The use of multiple camera heads permits the use of higher resolution collimators, for a given level of quantum mottle in the tomographic images, than would a single head system. However, the use of multiple camera heads poses considerable technical challenges for the manufacturer. It places severe requirements upon the electrical and mechanical stability of the camera heads. In particular, the X and Y offsets and X and Y magnification factors of all the heads must be precisely matched throughout the rotation about the patient. Today’s multihead systems are very stable and provide high-quality tomographic images for a variety of clinical applications.

Multihead gamma cameras are available in several configurations. Double-head cameras with opposed heads (180° head configuration) are good for head and body SPECT and whole-body planar scans (Fig. 19-1A and B). Triple-head, fixed-angle cameras are good for head and body SPECT, but less suitable for whole-body planar scans because
of the limited width of the crystals. Double-head, variable-angle cameras are highly versatile, capable of head and body SPECT and whole-body planar scans with the heads in the 180° configuration and cardiac SPECT in the 90° configuration (Fig. 19-1C and D). (The useful portion of the crystal does not extend all the way to the edge of a camera head. If the two camera heads are placed at an angle of exactly 90° to each other, both heads cannot be close to the patient without parts of the patient being outside of the FOVs. For this reason, one manufacturer provides the option of SPECT acquisitions with the heads at a 76° angle to each other, as well as zoom from the corner where the two heads meet instead of in the center of the two detectors’ FOVs.)

▪ FIGURE 19-9 A. Illustration of Siemens Healthineers’ IQ•SPECT SMARTZOOM collimator. The focal length increases from the center outward in both the radial and axial directions. B. Illustration of IQ•SPECT cardiocentric orbit acquisition. (Adapted with permission from Siemens Healthineers.)

Multielement Detector SPECT Cameras

SPECT systems that employ a multitude of small or curved detectors and alternative scanning techniques, as opposed to conventional rotating gantry SPECT scanners with two or three large FOV detectors, are now commercially available. Spectrum Dynamics Medical’s Veriton is a general-purpose (energy range 40-220 keV) scanner that has twelve detectors equally spaced over 360° around the patient. Each detector consists of a 6-mm-thick rectangular CZT crystal with 16 (transaxial dimension) × 128 (axial dimension), 2.46 mm × 2.46 mm detector elements, and an integrated parallel-hole tungsten collimator. Scanning of each (31.5 cm axial) SPECT FOV is achieved by a combination of detector swivel, rotation, and auto-contouring, with an increase in volume sensitivity on the order of 3 times compared to a conventional large FOV, dual-detector rotating gantry SPECT with LEHR collimation (Fig. 19-10). A scanner dedicated to cardiac SPECT, the NM530c from GE Healthcare, contains nine 8 cm × 8 cm and 5-mm-thick CZT crystals with 2.46 mm × 2.46 mm elements, equally spaced along a stationary L-shaped gantry. Strategic orientation of each detector and pinhole collimation allows all views to be acquired simultaneously without detector motion, resulting in an approximate fivefold increase in counting efficiency compared to conventional NaI(Tl)-LEHR (Fig. 19-11A and B). Spectrum Dynamics Medical also developed a dedicated cardiac CZT SPECT scanner (D-SPECT) Cardio, with nine 4 cm × 16 cm and 6 mm thick crystals, with 2.46 mm × 2.46 mm elements and parallel-hole tungsten collimation, in an L-shaped gantry, where scanning of the heart is achieved by translation and swivel of each detector, resulting in an approximate eightfold increase in sensitivity (Fig. 19-11C and D). A

third, NaI(Tl)-based dedicated cardiac SPECT scanner is CardiArc Inc.’s CardiArc, which employs three adjacent curved crystals and an array of PMTs. “Slit-hole” scanning is performed, whereby one series of lead sheets with horizontal gaps remains stationary while a curved lead sheet with six vertical slits rotates back and forth in electronic synchrony with six corresponding regions of the crystals, resulting in a sensitivity gain of about four.

▪ FIGURE 19-10 Veriton-CT multiple detector SPECT/CT system. View into the gantry from the SPECT side, showing the twelve CZT detector modules equally spaced over 360° retracted (left). Illustration of body contoured acquisition (right). (With permission from Spectrum Dynamics Medical.)

▪ FIGURE 19-11 Dedicated multiple CZT detector cardiac SPECT scanners. A. Discovery NM530c. B. Illustration of Discovery NM530c myocardial perfusion imaging (MPI) SPECT acquisition. (A and B, used with permission of GE Healthcare.) C. D-SPECT Cardio. D. Illustration of D-SPECT Cardio MPI SPECT acquisition (C and D, with permission from Spectrum Dynamics Medical.)

▪ FIGURE 19-11 (Continued)

19.2.2 Performance

Spatial Resolution

The spatial resolution of a SPECT system can be measured by acquiring a SPECT study of a line source, such as a capillary tube filled with a solution of Tc-99m, placed parallel to the axis of rotation (AOR). The National Electrical Manufacturers
Association (NEMA) has a protocol for measuring spatial resolution in SPECT. This protocol specifies a cylindrical plastic water-filled phantom, 22 cm in diameter, containing three line sources (Fig. 19-12, on the left) for measuring spatial resolution. The full widths at half maximum (FWHMs) of the line sources are measured from the reconstructed transverse images, as shown on the right in Figure 19-12. A ramp filter is used in the filtered backprojection so that the filtering does not reduce the spatial resolution. The NEMA spatial resolution measurements are primarily determined by the collimator used. The tangential resolution for the peripheral sources (typically 7 to 8 mm FWHM for low-energy high-resolution parallel-hole collimators) is superior to both the central resolution (typically 9.5 to 12 mm) and the radial resolution for the peripheral sources (typically 9.4 to 12 mm).

These FWHMs measured using the NEMA protocol, while providing a useful index of ultimate system performance, are not necessarily representative of clinical performance, because these spatial resolution studies can be acquired using longer imaging times and closer orbits than would be possible in a patient. Patient studies may require the use of lower resolution (higher efficiency) collimators than the one used in the NEMA measurement to obtain adequate image statistics. In addition, the filters used before backprojection for clinical studies cause more blurring than do the ramp filters used in NEMA spatial resolution measurements. The NEMA spatial resolution measurements fail to show the advantage of SPECT systems with two or three camera heads; double and triple head cameras will permit the use of higher resolution collimators for clinical studies than will single head cameras. Finally, the NEMA protocol is not applicable to, nor does it reflect, spatial resolution for iterative reconstruction SPECT, in particular, with system resolution, attenuation, and scatter compensations applied.

▪ FIGURE 19-12 NEMA phantom for evaluating the spatial resolution of a SPECT camera (left). The phantom is a 22-cm-diameter plastic cylinder, filled with water and containing three Co-57 line sources. One line source lies along the central axis and the other two are parallel to the central line source, 7.5 cm away. A SPECT study is acquired with a camera radius of rotation (distance from collimator to AOR) of 15 cm. The spatial resolution is measured from reconstructed transverse images, as shown on the right. Horizontal and vertical profiles are taken through the line sources and the FWHMs of these LSFs are determined. The average central resolution (FWHM of the central LSF) and the average tangential and radial resolutions (determined from the FWHMs of the two peripheral sources as shown) are determined. (Adapted by permission of the National Electrical Manufacturers Association, Performance Measurements of Scintillation Cameras, 2001.)

Spatial resolution deteriorates as the radius of the camera orbit increases. For this reason, brain SPECT produces images of much higher spatial resolution than does body SPECT. For optimal spatial resolution, the SPECT camera heads should orbit the patient as closely as possible. Body-contouring orbits (see above, “Design and Operation”) provide better resolution than do circular orbits.

Comparison of SPECT to Conventional Planar Gamma Camera Imaging

In theory, SPECT should produce spatial resolution similar to that of planar gamma camera imaging. In clinical imaging using filtered backprojection, its resolution is usually slightly worse. The camera head is usually closer to the patient in conventional planar imaging than in SPECT. The spatial filtering used in SPECT to reduce statistical noise also reduces spatial resolution. The short time per view of SPECT may mandate the use of a lower resolution collimator to obtain adequate numbers of counts. On the other hand, iterative reconstruction SPECT is capable of having better resolution than planar, especially when compensation for system resolution is applied.

In planar nuclear imaging, radioactivity in tissues in front of and behind an organ or tissue of interest causes a reduction in contrast. Furthermore, if the activity in these overlapping structures is not uniform, the pattern of this activity distribution is superimposed on the activity distribution in the organ or tissue of interest. As such, it is a source of structural noise that impedes the ability to discern the activity distribution in the organ or tissue of interest. The main advantage of SPECT over conventional planar nuclear imaging is improved contrast and reduced structural noise produced by eliminating counts from the activity in overlapping structures. SPECT using iterative reconstruction can also partially compensate for the effects of the scattering of photons in the patient and collimator effects such as the decreasing spatial resolution with distance from the camera and collimator septal penetration. When attenuation is measured using sealed radioactive transmission sources or an x-ray CT scanner, SPECT can partially compensate for the effects of photon attenuation in the patient.

19.2.3 Quality Control in SPECT

Even though a technical quality control program is important in planar nuclear imaging, it is critical to SPECT. Equipment malfunctions or maladjustments that would not noticeably affect planar images can markedly degrade the spatial resolution of SPECT images and produce significant artifacts, some of which may mimic pathology. Upon installation, a SPECT camera should be tested by a medical physicist. Following acceptance testing, a quality control program should be established to ensure that the system’s SPECT performance remains comparable to its performance at acceptance.

X and Y Magnification Factors and Multienergy Spatial Registration

The X and Y magnification factors, often called X and Y gains, relate distances in the object being imaged, in the x and y directions, to the numbers of pixels between the corresponding points in the resultant image. The X magnification factor is determined from an image of two point sources placed against the camera’s collimator a known distance apart along a line parallel to the x-axis:

The Y magnification factor is determined similarly but with the sources parallel to the y-axis. The X and Y magnification factors should be equal. If they are not, the projection images will be distorted in shape, as will be coronal, sagittal, and oblique images. (The transverse images, however, will not be distorted.) Multielement gamma camera (discussed in the previous chapter) do not suffer from such magnification errors, as detector element and total detector dimensions are identical between detectors, and Anger logic is not used for event positioning.

The multienergy spatial registration, described in the previous chapter, is a measure of the camera’s ability to maintain the same image magnification, regardless of the energy of the x- or γ-rays forming the image. The multienergy spatial registration is not only important in SPECT when imaging radionuclides such as Ga-67 and In-111, which emit useful photons of more than one energy, but also because uniformity and AOR corrections, to be discussed shortly, determined with one radionuclide will only be valid for others if the multienergy spatial registration is correct. (As discussed in the previous chapter, multienergy spatial registration is not applicable to multielement gamma cameras).

Alignment of Projection Images to the Axis of Rotation (COR Calibration)

The AOR is an imaginary reference line about which the head or heads of a SPECT camera revolve. If a radioactive line source were placed on the AOR, each projection image would depict it as a vertical straight line near the center of the image; this projection of the AOR into the image is called the COR. The location of the COR in each projection image must be known to correctly calculate the three-dimensional activity distribution from the projection images. Ideally, the COR is aligned with the center, in the x-direction, of each projection image. However, there may be misalignment of the COR with the centers of the projection images. This misalignment may be mechanical; for example, the camera head may not be exactly centered in the gantry. It can also be electronic or a digital setting. The misalignment may be the same amount in all projection images from a single camera head, or it may vary with the angle of the projection image or along the AOR. There are actually four other possible misalignments of a SPECT detector besides x-direction shift: axial tilt (discussed later in this chapter), detector-to-detector axial shift, yoke swivel, and axial swivel with respect to the AOR.

If a COR misalignment is not corrected, it causes a loss of spatial resolution in the resultant transverse images. If the misalignment is large, it can cause a point source to appear as a tiny “doughnut” (Fig. 19-13). (These “doughnut” artifacts are not seen in clinical images; they are visible only in reconstructed images of point or line sources. The “doughnut” artifacts caused by COR misalignment are not centered in the image and so can be distinguished from the ring artifacts caused by non-uniformities.) gamma cameras manufacturers provide software to assess and correct the effects of
COR misalignment. The COR alignment is assessed by placing a point source, several point sources, or a line source in the camera’s FOV, acquiring a set of projection images, and analyzing these images using the SPECT system’s computer. If a line source is used, it is placed parallel to the AOR.

▪ FIGURE 19-13 Center-of-rotation (COR) misalignment in SPECT. Small misalignments cause blurring (center), whereas large misalignments cause point sources to appear as “tiny doughnut” artifacts (right). Such “tiny doughnut” artifacts would only be visible in phantom studies and are unlikely to be seen in clinical images.

The SPECT system’s computer corrects the COR misalignment by shifting each clinical projection image in the x-direction by the proper number of pixels prior to filtered backprojection or iterative reconstruction. When a line source or multiple points along the AOR are used for calibration, a COR correction can be derived and applied separately for each transverse slice. If the COR misalignment varies with the camera head angle, instead of being constant for all projection images, it can only be corrected if the computer permits angle-by-angle corrections. Separate assessments of the COR correction must be made for different collimators and dual-head configurations (e.g., 180° and 90°), and, on some systems, for different camera zoom factors and image formats (e.g., 64² versus 128²). The COR correction determined using one radionuclide will only be valid for other radionuclides if the multienergy spatial registration is correct.

Uniformity

The uniformity of the camera head or heads is important; non-uniformities that are not apparent in low count daily uniformity studies can cause significant artifacts in SPECT. The artifact caused by a non-uniformity appears in transverse images as a ring centered about the AOR (Fig. 19-14).

Multihead SPECT systems can produce partial ring artifacts when projection images are not acquired by all heads over a 360° arc. Clinically, ring artifacts are most apparent in high-count density studies, such as liver scans. However, ring artifacts may be most harmful in studies such as myocardial perfusion in which, due to poor counting statistics and large variations in count density, they may not be recognized and thus lead to misinterpretation.

The causes of non-uniformities were discussed in the previous chapter. As mentioned in that chapter, modern gamma cameras have digital circuits using lookup tables to correct the X and Y position signals from each interaction for systematic position-specific errors in event location assignment and the Z (energy) signal for systematic position-specific variations in scintillator light collection or CZT charge
generation efficiency. However, these correction circuits cannot correct non-uniformity due to local variations in detection efficiency, such as dents or manufacturing defects in the collimators.

▪ FIGURE 19-14 Image of a cylinder filled with a uniform radionuclide solution, showing a ring artifact due to a non-uniformity. The artifact is the dark ring toward the center.

If not too severe, non-uniformities of this latter type can be largely corrected. A very high-count uniformity image is acquired. The ratio of the average pixel count to the count in a specific pixel in this image serves as a correction factor for that pixel. Following the acquisition of a projection image during a SPECT study, each pixel of the projection image is multiplied by the appropriate correction factor before COR correction and filtered backprojection or iterative reconstruction. For the high-count uniformity image, at least 30 million counts should be collected for 64² pixel images and 120 million counts for a 128² pixel format. These high-count uniformity images are typically acquired weekly to monthly. Correction images must be acquired for each camera head and collimator. For cameras from some manufacturers, separate intrinsic correction images must be acquired for each radionuclide. The effectiveness of a camera’s correction circuitry and use of high-count flood correction images can be tested by acquiring a SPECT study of a large plastic cylindrical container or a SPECT performance phantom filled with a well-mixed solution of Tc-99m and examining the transverse images for ring artifacts. However, this testing will not assess parts of the collimator or camera face outside the projected image of the container or phantom.

Camera Head Tilt

The camera head or heads must be aligned with the AOR; for most types of collimators, this requires that faces of the heads be exactly parallel to the AOR. If they are not, a loss of spatial resolution and contrast will result from out-of-slice activity being backprojected into each transverse image slice, as shown in Figure 19-15. The loss of resolution and contrast in each transverse image slice will be less toward the center of the slice and greatest toward the edges of the image. If the AOR of the camera is aligned to be level when the camera is installed and there is a flat surface on the camera head that is parallel to the collimator face, a bubble level may be used to test for head tilt. Some SPECT cameras require the head tilt to be manually adjusted for each acquisition, whereas other systems set it automatically. The accuracy of the automatic systems should be periodically tested. A more reliable method than the bubble level is to place a point source in the camera FOV, centered in the axial (y) direction, but
near the edge of the field in the transverse (x) direction. A series of projection images is then acquired. If there is head tilt, the position of the point source will vary in the y-direction from image to image. Head tilt can be evaluated from viewing a cine of the projection images.

▪ FIGURE 19-15 Head-tilt. The camera head on the left is parallel to the AOR, causing the counts collected in a pixel of the projection image to be backprojected into the corresponding transverse image slice of the patient. The camera head on the right is tilted, causing counts from activity outside of a transverse slice (along the grey diagonal lines of response) to be backprojected into the transverse image slice (orange colored vertical slice).

▪ FIGURE 19-16 Flangeless deluxe Jaszczak phantom for testing SPECT systems, a product of Data Spectrum Corporation. (Courtesy of Data Spectrum Corporation.) A soluble radioactive material, typically labeled with Tc-99m, is introduced into the phantom and mixed in the water until it is uniformly distributed. The acrylic plastic spheres and rods are not radioactive and are “cold” objects in the radioactive solution.

SPECT Quality Control Phantoms

There are commercially available phantoms (Fig. 19-16) that may be filled with a solution of Tc-99m or other radionuclide and used to evaluate system performance. These phantoms are very useful for the semiquantitative assessment of spatial resolution, image contrast, and uniformity, although the small sizes of most such phantoms allow uniformity to be assessed only over a relatively small portion of the camera’s face. They are used for acceptance testing of new systems and periodic testing, typically quarterly, thereafter. Table 19-1 provides a suggested schedule for a SPECT quality control program.

19.2.4 Dual Modality Imaging—SPECT/X-ray CT Systems

Several manufacturers provide imaging systems incorporating two gamma camera heads capable of planar imaging and SPECT and an x-ray CT system with a single patient bed. Some of these systems have very simple x-ray CT systems that provide x-ray CT information for attenuation correction of the SPECT information and image co-registration only, whereas others can produce diagnostic quality CT images.

The advantages and disadvantages of using an x-ray CT system for attenuation correction and image co-registration are the same as in the case of the PET/CT systems described later in this chapter. The x-ray CT system acquires the attenuation information much more quickly than does a system using radioactive sealed sources and the attenuation information has less statistical noise, but the linear attenuation coefficients for the energies of the γ-rays must be estimated from the CT attenuation information. The methods for doing this are similar to those used in PET/CT, which are discussed later in this chapter. Artifacts can occur when the calculation produces the wrong attenuation coefficient. This can be caused by high atomic number material in the
patient, such as metal objects and concentrated contrast material. Furthermore, the SPECT and x-ray CT information are not acquired simultaneously and patient organ motion can result in the misregistrations of, and therefore artifacts in, SPECT image information, the same as in PET/CT (discussed later in this chapter).

TABLE 19-1 RECOMMENDED SCHEDULE FOR ROUTINE QUALITY CONTROL TESTING OF A PLANAR OR SPECT GAMMA CAMERA

TEST	FREQUENCY	NOTES
Set and check energy discrimination window(s)	Before first use daily	Point (intrinsic) or planar (extrinsic) source of radionuclide
Extrinsic or intrinsic low-count uniformity images of all camera heads	Before first use daily	5-10 million counts, depending upon effective area of camera head
Cine review of projection images and/or review of sinogram^a (S)	After each clinical SPECT study	Check for patient motion
Visual inspection of collimators for damage	Daily and when changing collimators	If new damage found, acquire a new high-count uniformity calibration image
High count-density extrinsic or intrinsic uniformity images of all camera heads (S)	Monthly	30 million counts for 64² images and 120 million counts for 128²
Spatial resolution check with bar pattern	Weekly	Cycle week-to-week between 0°, 90°, 180° and 270° orientation of bar pattern
Center of rotation (S)	Weekly to monthly	Point or line source(s), as recommended by manufacturer
Efficiency of each camera head	Quarterly, semiannually, or annually
Reconstructed cylindrical phantom uniformity (S)	Annually	Cylindrical phantom filled with Tc-99m solution
Point source reconstructed spatial resolution (S)	Annually	Point source
Reconstructed SPECT phantom (S)	Quarterly	Using a phantom such as the one shown in Figure 19-16
Pixel size check	Annually	Two point sources
Head-tilt angle check (S)	Quarterly	Bubble level or point source
Extrinsic uniformity images of all collimators not tested above	Annually	Planar source. High-count density images of all collimators used for SPECT
Multienergy spatial registration	Annually	Ga-67 point source
Count rate performance	Annually	Tc-99m source
The results of these tests are to be compared with baseline values, typically determined during acceptance testing. If the manufacturer recommends or an accrediting body specifies additional tests or more frequent testing, these recommendations or specifications should take precedence. For tests with multiple frequencies listed, it is recommended that the tests be performed initially at the higher frequency, but the frequency be reduced if the measured parameters prove stable. Tests labeled (S) need not be performed for cameras used only for planar imaging.
^aA sinogram is an image containing projection data corresponding to a single transaxial image of the patient. Each row of pixels in the sinogram is the row, corresponding to that transaxial image, of one projection image.

A common use of SPECT/CT is myocardial perfusion imaging. In myocardial perfusion imaging, non-uniform attenuation, particularly by the diaphragm and, in women, the breasts, can cause apparent perfusion defects in the SPECT images. Attenuation correction using the CT information has been reported to improve
diagnostic accuracy by compensating for these artifacts. However, spatial misregistration between the heart in the CT and SPECT images can cause artifacts (Goetze et al., 2007). General SPECT-only imaging is being supplanted by dual-modality imaging, due to the mainstreaming and thus proliferation of SPECT/CT systems. An important quality assurance step in SPECT/CT is to verify the alignment of the SPECT and CT image information in every clinical examination.

19.3 POSITRON EMISSION TOMOGRAPHY

PET generates images depicting the distribution of positron-emitting nuclides in patients. Nearly all PET systems manufactured today are coupled to x-ray CT systems, with a single patient bed passing through the bores of both systems, and are referred to as “PET/CT” systems. Figure 19-17 shows a PET/CT system. This section discusses PET imaging systems; their use in PET/CT systems is discussed later in this chapter.

In a typical PET system, several rings of detectors surround the patient. PET scanners use annihilation coincidence detection (ACD) instead of collimation to obtain projections of the activity distribution in the subject. The PET system’s computer then reconstructs the transverse images from the projection data, as does the computer of an x-ray CT or SPECT system. Modern PET scanners are multislice devices, permitting the simultaneous acquisition of many transverse images over a preset axial distance. The clinical importance of PET today is largely due to its ability to image the radiopharmaceutical fluorine-18 fluorodeoxyglucose (FDG), a glucose analog used for locating malignant neoplasms, differentiating malignant neoplasms from benign lesions, staging patients with malignant neoplasms, monitoring the response to therapy for neoplasms,
differentiating severely hypoperfused but viable myocardium from scar, and other applications. However, other positron-emitting radiopharmaceuticals have been approved for use and their clinical use is increasing; these are discussed later in this chapter.

19.3.1 Design and Principles of Operation

Annihilation Coincidence Detection

Positron emission is a mode of radioactive transformation and was discussed in Chapter 15. Positrons emitted in matter lose most of their kinetic energy by causing ionization and excitation. When a positron has lost most of its kinetic energy, it interacts with an electron by annihilation, as shown on the left in Figure 19-18. The entire mass of the electron-positron pair is converted into energy equal to 1.02 MeV (E = mc², where c is the speed of light and m is the combined mass of the electron and positron), which appear as two 511-keV photons that are emitted in nearly opposite directions. In solids and liquids, positrons travel only very short distances (see Table 19-3) before annihilation.

If both photons from an annihilation interact with detectors and neither photon is scattered in the patient, the annihilation occurred near the line connecting the two interactions, as shown on the right in Figure 19-18. Circuitry within the scanner identifies pairs of interactions occurring at nearly the same time, a process called ACD. The circuitry of the scanner then determines the line in space connecting the locations of the two interactions, which is known as a line of response (LOR). Thus, ACD establishes the trajectories of detected photons, a function performed by collimation in SPECT systems. However, the ACD method is much less wasteful of photons than collimation. Additionally, ACD avoids the degradation of spatial resolution with distance from the detector that occurs when collimation is used to form projection images.

True, Random, and Scatter Coincidences

A true coincidence is the nearly simultaneous interaction with the detectors of emissions resulting from a single nuclear transformation. A random coincidence (also called an accidental or chance coincidence), which mimics a true coincidence, occurs when emissions from different nuclear transformations interact nearly simultaneously with the detectors (Fig. 19-19). A scatter coincidence occurs when one or both of the photons from a single annihilation are scattered, and both are detected (Fig. 19-19). A scatter coincidence is a true coincidence because both interactions result from a
single positron annihilation. Random coincidences and scatter coincidences result in misplaced coincidences because they are assigned to LORs that do not intersect the actual locations of the annihilations. They are therefore sources of noise, whose main effects are to reduce image contrast and increase statistical noise.

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