Chapter 4: Cardiac SPECT and PET Instrumentation

The changing needs of cardiologists and patients have led to an unprecedented evolution in nuclear cardiology instrumentation. Modern instrumentation has a wide range of collimation, attenuation correction (AC) options, and many of the newest systems no longer relying on the Anger gamma camera design. Solid-state detector technologies and new aperture designs are commercially available and in widespread use. Hybrid SPECT and PET systems now have transmission imaging options for AC using x-ray CT, radionuclide, or novel fluorescence x-ray sources.

The changes in instrumentation have been driven by revolutionary changes in the delivery of health care. In particular, there has been a shift from office-based nuclear cardiology services to hospital-based imaging. According to MedAxiom, respondents reported a dramatic shift in providers from private practices to integrated practices. In 2008, only 11% of MedAxiom respondents reported being in an integrated group (e.g., integrated into a hospital environment). By 2012, almost 60% of respondents reported practicing within an integrated practice.¹ The integration of cardiovascular imaging into the hospital setting has significantly altered the kinds of instrumentation available to the cardiologist. This access has provided clinicians the capability for routine AC, neuronal imaging, absolute blood flood, sarcoid imaging, amyloid imaging, metabolic imaging (glucose and fatty acid tracers) and other advanced molecular imaging techniques. This chapter will describe the instrumentation that is commonly used in nuclear cardiology and explain some of the important advances in instrumentation that can improve image quality and reduce radiation exposure.

Anger Camera

Hal Anger, with his team at the University of California, Berkeley had worked with the rectilinear scanners and pin-hole camera designs. In 1957, this group changed the design to allow for a uniform magnification and resolution across the field of view with increased system sensitivity. The first "Anger" scintillator camera consisted of a 4-in sodium iodine crystal optically coupled to 7 photomultiplier tubes (PMTs).^2,3

Gamma camera design has evolved significantly since 1957 with improvements in resolution, sensitivity, and energy discrimination, however the conceptual design of the Anger camera is remarkably unchanged. The "Anger" gamma camera consists of (see Fig. 4-1):

• Focusing, multi-hole collimator: The focusing collimator is placed in front of the other camera components. The image is formed by excluding photons not traveling along a particular path, for example, perpendicular to the camera face or holes focusing on a point or line.

• Scintillation crystal: A material that will luminesce with optical photons when excited by ionizing radiation.

• PMT array: Electronics for capturing the photons produced by the scintillator crystal and converting those photons into an electronic pulse.

• Analog-to-digital conversion hardware: A digitization board converts the analog pulse-height information into a position and energy for each photon event.

Collimation, Imaging Efficiency, and Image Quality

Unlike optical photons, gamma photons cannot be focused with either reflection or refraction. "Focusing" for gamma rays is typically performed by preferentially accepting only those photons traveling in a particular direction. The use of collimation dramatically reduces photon sensitivity by absorbing almost 99.99% of all photons entering the camera. The most common collimator used in SPECT is a parallel-hole collimator. These collimators function by using an array of parallel holes in a block of attenuating material such as lead or tungsten (see Fig. 4-2). There are some less common geometry collimators that are specific to various cardiac application such as fan beam collimation that focus on a line at a fixed distance from the collimator and variable distance collimator.⁴

The use of collimation instead of the lenses that are used in conventional photography, leads to some interesting differences in nuclear imaging:

1. The use of collimation results in a loss of resolution with increasing source distance from the detector surface (see Fig. 4-3).⁵

2. Sensitivity is constant with distance. Specifically, the number of counts/second recorded is independent of distance, so long as the loss of resolutions does not result in counts missing the detector entirely.

3. The field of view for parallel-hole collimation does not change with distance.

The resolution of a gamma camera depends on the diameter and length of the holes in the collimating material. "High-resolution" collimators tend to have a longer hole length, smaller hole diameter, and thinner septa compared with "general-purpose" or "all-purpose" collimators.⁶

Parallel-hole collimation is generally favored but most nuclear laboratories because of the consistency of the resolution with patient positioning, camera rotation, and protocol variation.⁷

Fan-beam geometry allows radiation emitted from a source at a specific distance to pass through the collimator without interaction and reach the crystal surface.⁸ A hybrid design combining both fan-beam properties in the center of the field of view and a gradual approach to parallel-hole design toward the edge of the field of view has been implemented on a dual detector Anger-based SPECT system for cardiac SPECT.^4,9 This design improves the geometric component of system sensitivity by a factor of 2 over parallel geometry and minimizes truncation of projections near the edge of the field of view.

Image Formation

When a gamma photon is absorbed by the scintillation crystal, a burst of secondary optical photons is then released. An array of PMTs, positioned behind the crystal, detects and converts the light pulse into an electric pulse that is digitized.

The digitization process is a mathematical weighting of the pulses from all PMTs detecting the scintillation produces X, Y (spatial location), and Z (energy) values. The energy value is tested by a discrimination circuit that will exclude the signal if it is not within the permitted energy range (referred to as the "energy window").

Energy resolution has an important role in differentiating between scatter and photopeak photons. When gamma rays scatter off of electrons in the media, they release some of their energy to the electrons (referred to as the recoil electron). As a result of scattering of the electrons, the energy of the scattered photon is reduced. This process of photons elastically scattering off of electrons in the media is referred to as Compton scattering. Compton scattering degrades the images quality by reducing the contrast resolution and spatial resolution of the image.

System sensitivity is another key imaging performance measure expressed in units of counts/min/μCi (cpm/μCi) of Tc-99m. It is dependent on the length and diameter of the holes in the collimator. Higher system sensitivity collimators allow more photons to pass through the collimator at the expense of spatial resolution. Conversely, higher-resolution collimators allow fewer photons to pass through the collimator by restricting the acceptance angle of the collimator holes. This reduction in acceptance angle improves the spatial resolution of the system.

Tomographic Gantry

To produce tomographic images of a patient, a series of projections around the patient is needed. The most common approach is to mount one or more Anger cameras on a rotating gantry. Early designs for the gantry involved a single camera head that could be rotated around the patient on a fork mount. Because these systems were also used for planar imaging, the camera head could be rotated around two axes.¹⁰ With the increase use of SPECT and the challenge of maintaining an accurate center of rotation led to near universal use of the single axis of rotation gantry for cardiac studies.

The most common design in use today for cardiac SPECT is a dual-headed gamma camera with 90 degrees of separation between each camera head (see Fig. 4-4). To complete a tomographic study, the camera only needs to complete a 90-degree arc, and camera heads rotate simultaneously, thus saving image acquisition time. This geometry offers an excellent combination of a wide field of view, scanning efficiency, and simplicity. Most systems also offer the option to acquire a noncircular orbit to minimize the distance between the camera and the patient. This has the effect of improving the resolution by up to 1.5 to 2.5 mm.¹¹

Solid State, Non-Anger Imaging

Several solid-state detector-based systems are now available commercially.^12–14 These systems have a crystal layer placed over an array of semiconductor units that collect the scintillation light directly forming a net electronic charge in each unit that is proportional to the incident scintillation light. Because of the more direct coupling of crystal and detectors, the energy and location are more easily determined. Common solid-state crystals are cadmium-zinc telluride (CZT) and cesium-iodide (CsI) due to their physical and optical emission properties.¹⁵ A lead or tungsten collimator is also used over the crystal surface for localization of incident photons as with the Anger approach. In comparison to Anger systems, solid-state detectors have superior energy resolution compared with the NaI(Tl) crystals. Thus, solid-state detectors intrinsically have greater contrast resolution. System sensitivity of the solid-state detectors is modestly (10–15%) lower than Anger systems.¹⁶ However, this decrease in acquisition efficiency is offset by the use of multiarrays of detectors encompassing greater angular range for detection about the patient that improves the efficiency for detection.

Multi-Detector System

More recently, several alternatives to the Anger gamma camera have been introduced. One such system utilizes nine independent scanning detectors to concentrate the imaging on a small field of view, centered on the heart¹² (D-SPECT, Spectrum Dynamics, Haifa, Israel). The advantage of this system is the ability of the scanner to concentrate the acquisition on the myocardium. This has the effect of preferentially improving the system's SPECT sensitivity in the region of the heart (see Fig. 4-5). The D-SPECT system does not have the same quality control procedures as a conventional Anger SPECT system. Daily quality control (QC) consists of three QC steps: (1) global homogeneity, (2) regional homogeneity, and (3) field of view.¹⁷ These must be performed to insure optimal image quality.

Several studies have been performed to investigate the potential for this scanner for ultra–low-dose, stress first imaging.¹⁸ In a study of 284 patients (208 with coronary angiography, 76 low likelihood) populations were split into a 128 standard stress dose (10.2 ± 0.5 mCi) patients and 156 low stress dose (5.9 ± 1.2 mCi). In that study there was no significant difference between sensitivity, specificity, and accuracy between the low dose and standard dose studies (86.1%, 76.6%, and 81.4%; and 90.6%, 78.1%, and 84.4%, respectively, p = ns). This implied that for those patients that only required a stress study, the radiation dose was as 1.7 ± 0.3 mSv.

Multi Pinhole

In the multipinhole aperture approach,^13,19 individual "pinholes" are placed strategically on the surface of a solid lead structure on each detector such that photons may only pass through the apertures to reach the underlying detectors (see Fig. 4-6). This approach permits the images to be acquired without gantry rotation as with conventional systems. The system sensitivity (per unit activity) compared with conventional Anger-based systems is approximately five times that of conventional SPECT imaging.

One such system consisting of 19 pinholes has become available to cardiologists (Discovery NM 530c, General Electric). This system projects onto 4 solid-state CZT pixilated detector. This system has the advantage that there are no moving parts and is 5 to 10 times more sensitivity than conventional Anger systems.²⁰

Sensitivity and specificity of the Discovery NM 530c were studied in a population of 160 patients imaged myocardial perfusion patients using a weight-adjusted rest/stress Tc-99m-sestamibi study. In that study, the automated, computer results showed higher specificity (59–67% vs. 27–60%) and lower sensitivity (71–72% vs. 79–93%) than the visual reads.²¹ This would suggest that the appearance of these images may be different than conventional images, however the differences are predictable. It is likely that with experience, visual interpretation should approach computer-generated results.

AC is employed in cardiac SPECT to compensate for image artifacts that can result from absorption of photons by the soft tissue of the patient.^22–24 AC requires an anatomical map as additional information for the reconstruction algorithm to compensate for this loss of signal. Two transmission imaging approaches commonly used to measure the patient anatomy are as follows:

• Radionuclide sources: These sources can acquire transmission data simultaneously or sequentially with the emission scan while the x-ray–based approaches acquire images only in sequential mode.^8,22 These approaches typically have a very low radiation dose (0.03 mSv).²⁵

• X-ray–based CT systems: X-ray CT images also have spatial resolution that is on the order of 10 times greater than the emission images. X-ray CT images can be temporally gated with ECG signals for AC while radionuclide source methods are generally not acquired with ECG gating.²³

Line Source Attenuation Correction

Scanning line sources using Gd-153 (100 keV, T_1/2 = 273 days) are the most commonly used radionuclide source method for SPECT AC.^8,22 These systems use an external source of radiation to project gamma rays through the patient to the detector on the far side of the patient. To create a transmission reconstruction of the attenuating medium, two scans are needed: a blank scan of the line source without the patient present, and a scan with the patient present. The ratio of the scan divided by the blank scan is then used to create a patient-specific map of the attenuation.

One advantage line source AC has over CT-based AC is that it can be performed simultaneous to the emission study. This improves laboratory efficiency and removes the possibility of misregistration between the transmission and emission datasets. Another advantage of using a line source AC is the elimination of common CT artifacts such as breathing artifacts, metal artifacts, etc.

The greatest challenge of using line source AC is acquiring sufficient counts to perform a transmission reconstruction. It is important for quality of AC imaging to maintain the strength of the radionuclide sources by periodic replenishment as recommended by the manufacturer.²⁶ As the sources decay, the acquisition time must be extended proportionally to maintain image quality, which can decrease patient convenience and laboratory efficiency. Compromised quality of the attenuation map may result from image noise otherwise, introducing artifacts in the attenuation maps resulting in compromised accuracy or artifacts in the corrected images.²⁷

CT-Based Attenuation Correction

The high spatial resolution of x-ray CT images has significant implications for AC quality. CT-based AC is similar to line source AC in that a patient-specific map of the patient's anatomy is used to compensate for the influence of soft tissue attenuation. Though this process on its surface is similar to line source AC, there are several important differences:

1. X-ray CT-based transmission imaging is by necessity performed sequentially to the emission study. This has a negative impact on laboratory efficiency. This also can introduce artifacts that are a result of misregistration between the transmission and emission datasets.

2. X-ray CT-based images are recorded in Hounsfield units. An algorithm is needed to translate these units into attenuation coefficients.

3. X-ray CT-based AC typically has higher radiation dose for the patient.²⁸ Every effort must be made to minimize the radiation from the CT portion of the study.

Proper AC requires the attenuation map images be aligned in all dimensions with the SPECT images.²⁹ As the CT images may be acquired over a single phase, and the emission images over multiple phases, this is not always easily performed. As a result, the attenuation map boundaries of the CT images are much sharper compared with the radionuclide images in the same patient. Misalignment by as little as 1 cm may produce artifacts with severity similar to true perfusion defects. The linear attenuation coefficients (or Hounsfield units), which comprise the attenuation map pixel values, vary greatly in the thorax compared with other regions of the body where lung tissue values for 140 keV (Tc-99m) are almost 10 times less than those for the myocardium. Whenever CT-based AC is applied, images must be carefully inspected for misregistration and corrected as necessary.³⁰

One application that has been suggested is to use diagnostic quality CT in conjunction with myocardial perfusion to assess the presence of coronary calcium in patients with no known history of CAD. In a study of 1,126 asymptomatic patients with a quantitative coronary calcium score, it was demonstrated that the use of coronary calcium assessment in conjunction with myocardial perfusion assessment independently improved the prediction of future cardiac events.³¹ There have also been smaller studies that suggest that the combination of SPECT with CT angiographic result could provide some added clinical value, additional research is needed to establish optimal clinical scenarios and justification for the additional procedure and radiation dose.³²

In conventional parallel-hole SPECT imaging, multiple projection images of the patient are taken at uniformly spaced angles to obtain a three-dimensional (3D) image. Iterative algorithms, described below, may include additional information about the physics of the imaging process, including spatial resolution, noise, and attenuation properties to improve the accuracy of reconstruction.^33–35

Filtered Backprojection Reconstruction

The physics of the image formation process is modeled mathematically by a projector matrix. A common model for this process is the Radon transform.³⁶ All the activity along each line of sight is integrated onto a single (ideal) point on the detector surface.

The 3D reconstruction is made by taking the Fourier transform of each projection dataset and filtering that data using a linear filter that suppresses lower-frequency modes when compared to higher frequency modes (the "ramp filter," see Fig. 4-7). Noise filtering can also be applied to the projections prior to reconstruction, during reconstruction, or postreconstruction. ECG-gated projections, separated into bins over the cardiac cycle, are independently reconstructed to produce the dynamic functional representation of the heart over the cardiac cycle.

Filtered-backprojection (FBP) reconstruction is being replaced by iterative reconstruction technique. Iterative reconstructions can model the noise in the data, resolution, scatter, and attenuation. For these reasons, FBP reconstruction should be avoided when iterative techniques are available.

Iterative Image Reconstruction

Iterative algorithms use a stepwise approach and have a more comprehensive and mathematically acceptable framework for tomographic imaging.^37,38 They provide a statistically "most probable" image that would produce the projection data, given the projection model including the physics of imaging. Iterative SPECT reconstruction begins with an initial estimate of the activity distribution, usually a uniform map. At each step of the iteration sequence, an update of the present estimate on each transverse plane is calculated. The most common strategy for performing iterative reconstruction are the maximum likelihood expectation–maximization (MLEM) algorithm³⁷ and the ordered subsets expectation–maximization (OSEM) model³⁸ (see Fig. 4-8). MLEM reconstruction requires the specification of the number of iterations, while OSEM requires specification of both number of iterations and subsets (a subgroup of the projection images).

Iterative algorithms can be used to reconstruct the attenuation map from transmission images, or to reconstruct attenuation-corrected images where the attenuation map is provided as an input to the MLEM or OSEM algorithm. The attenuation information is integrated into the reconstruction and creates a set of weights for compensating for the attenuation in the image (see Fig. 4-9). When applied to x-ray CT acquisitions, this approach removes the severe artifacts that result from metals or other highly dense materials used in devices such as ICDs for CT.^39,40

Reconstruction Algorithms for Rapid SPECT Acquisitions

Software approaches for image reconstruction have been recently described and clinically validated for "fast" image acquisition protocols with "one-half" to "one-fourth" the time of conventional acquisitions.^33–35 These algorithms differ somewhat in modeling detail by manufacturer, but they are based on the principle described by Muehllehner that when system resolution improves, the signal-to-noise value increases at each point, as the blur is reduced and the counts are more efficiently utilized.⁴¹ The various methods have the common attributes of combined distance-dependent spatial resolution compensation and optimized statistical noise modeling internal to the iterative reconstruction algorithm. These new algorithms offer an opportunity to reduce radiation dose to the patient and/or decrease image acquisition time.

Photopeak Scatter and Nonstationary Spatial Resolution

AC for SPECT is performed using iterative reconstruction, but it does not correct for all of the artifacts that are present in SPECT. The different commercial approaches differ by what processes are included, and how they are balanced in the modeling of the image formation process. Compton scattering of photons, partial volume effects, and depth-dependent spatial resolution are factors that are also important to correct for physical limitations of SPECT.⁴² Compton-scattered photons detected in the photopeak window reduce the contrast resolution or "blur" the image as a result of redirection by the scattering process. This increases the "background" or surrounding pixel values locally, but may actually lead to false-negative regions depending on the management by the reconstruction algorithm. Tracer uptake in the liver and other sub diaphragmatic structures can significantly influence the image values in the heart, typically the inferior regions where blurring into surrounding pixels occurs.⁴³ Using techniques to minimize uptake prior to imaging is important for Tc-99m–based agents because of the affinity of sub diaphragmatic structures.⁴⁴ Extracardiac activity adjacent to the myocardium can cause improper scaling of the image suppressing the other regions of the myocardial image. This has the potential for an artifactual decrease in anterior wall intensity.⁴⁵ Information for estimating the amount of scatter in the photopeak can be collected in a separate energy window simultaneously to subtract a fraction of these data from the photopeak image. Most scatter correction methods with conventional cardiac SPECT are implemented in combination with AC methods utilizing the attenuation map or modeling to estimate scatter interaction and perform the corrections during image reconstruction.⁴⁶

Basics of PET

The first application of positron annihilation medical imaging was first explored in by Brownell et al. for imaging brain tumors.⁴⁷ This technique was later expanded to tomographic imaging in 1971.⁴⁸ Myocardial perfusion PET with ¹³N ammonia was one of the first practical implementations of PET in the myocardium.^{49 13}N ammonia produces high-quality myocardial perfusion images with favorable responses to changes in blood flow.^49–50 The short half-life of ¹³N requires a nearby cyclotron for production, thereby limiting the number of sites that can use it. A more practical approach employs a generator to produce the radionuclide. ⁸²Rb was studied in 1986 as a generator-produced potassium analog for the detection of myocardial infarction.^51–52 Using an ⁸²Sr generator (T_1/2 = 28 days), a user could be provided with an on-site generator every 4 weeks. More recently, these generators can be supplied for up to 6 weeks, depending upon available camera technology.^53,54

Metabolic cardiac imaging with PET has been of great interest because of the relative simplicity of using PET radionuclides with metabolically active molecules. ¹⁸F-labeled deoxyglucose (FDG) (a glucose analog) has long been used for imaging myocardial viability^55,56 and has been also used for imaging of ischemic activation⁵⁷ and sarcoid imaging.⁵⁸ As a metabolic agent, FDG is trapped in cells that utilizes predominately glucose metabolism. This can be present in hibernating myocardium, severely ischemic myocardium, infections and inflammation as well as in tumors. In addition, metabolic agents have been explored for fatty acid metabolism, cardiac denervation, and the detection of amyloid plaques.^59–61

PET Scanners

PET photon events consist of the following⁶²:

• True pairs: Actual coincidence events where two unscattered 511-MeV photons produced by the same positron annihilation and are detected by the system. These events are used for producing the tomographic images.

• Photon scatter events: Coincidence events where two photons produced by the same positron annihilation are recorded by the system, but at least one of the photons has been scattered by electrons in the patient. These events degrade image contrast.

• Random events: Coincidence events where two photons produced by different positron annihilations are recorded by the system within the timing gate of the system. These events degrade image contrast.

• Prompt gamma: Additional photons that are produced during the nuclear decay that produces the positron. Given the timing, these photons can appear as coincidence photons. With Rb-82, 13% of all decays produce and additional 776-keV prompt gamma. When these photons are not accounted for, the over-correction for scatter can reduce specificity from 90% to 22%.⁶³

A coincidence "event" occurs when two 511-keV photons resulting from positron decay and annihilation, emitted nearly 180 degrees apart, are detected "simultaneously" within a small (4–12 nanoseconds) timing window.⁶² Most PET scanners consist of a system of concentric rings of multiple-detector blocks around a central axis (see Fig. 4-10). Each of these detectors contains a scintillator crystal and a small array of photo-multiplier tubes. The physical properties of the scintillator crystal determine much of the performance of the PET scanner. Ideally, a scintillator crystal has a high stopping power (to reduce the depth the gamma rays penetrate before being recorded), high light output per event (needed to differentiate photon energies), and a rapid light curve decay (to minimize the time needed to record the photon event and have the crystal ready to receive the next event). Response time of the system describes how effective a system can differentiate random events from true coincidence. To accomplish this, a coincidence timing window of <10 nanoseconds is needed to differentiate true pairs from random coincidence event. As challenging as that may appear, most modern PET scanners have coincidence windows of <5 nanoseconds, making it possible to not only differentiate the difference between randoms and trues, but utilize the differences in the arrival times of the two annihilation photons to localize the annihilation event along the line of site.

The intrinsic sensitivity of PET is typically much higher than SPECT due to the fact that localization of the annihilation event does not require the use of collimation. Despite this fact, a large number of PET imaging systems use a system of septa to reduce scatter and randoms and thereby improve the image quality of the final reconstructed images.

Recent enhancements in processing techniques have enabled fully 3D imaging (septa removed) with Rb-82.⁶⁴ Removing the septa significantly improves the sensitivity of the imaging system, potentially increasing system sensitivity by a factor of two to five times.⁶⁵ The American Society of Nuclear Cardiology recommends reducing the infused dose of Rb-82 from 50 mCi using two-dimensional (2D) imaging to 30 mCi in 3D. Similar dose reduction is also recommended for FDG imaging.⁶⁶

Attenuation Correction of PET

In contrast with SPECT, AC is almost always applied in the processing of cardiac PET. The geometry of PET makes it possible to estimate attenuation without knowing the depth of the annihilation event in the patient. Unlike SPECT, the attenuation of a true event is the sum of the attenuation both 511-keV photons; specifically the total attenuation along the line of response (see Fig. 4-11). Therefore, the correction that is applied is independent of where the annihilation event took place along a line of response. The result is a high-resolution final image with greatly reduced attenuation and scatter artifacts (see Fig. 4-12).

Radionuclide Source for Attenuation Correction

Radionuclide sources for transmission imaging used for AC are Germanium-68 (Ge-68, 511 keV) or Cesium-167 (Cs-137, 630 keV). Both are contained in rotating rod sources that can be extended and retracted for image acquisition as needed throughout the study. Such cameras that utilize radioactive line sources are called "dedicated" PET systems. Because the source photons are positrons, there is no need to translate the attenuation coefficients from the transmission energy to the emission energy. Another advantage of dedicated PET is the attenuation map is acquired while the patient is free breathing. This reduces misregistration and breathing artifacts when compared to PET/CT.^67,68 Finally, the radiation dosage from a dedicated PET scanner is typically less than an x-ray–based CT (though dosage-reducing strategies can be employed to minimize x-ray radiation dosage). One possible limitation of dedicated PET is some systems require long transmission acquisition scans (4–8 minutes). Newer reconstruction methods with newer instrumentation reduce the number of counts required for image reconstruction, reducing the acquisition time to 60 to 90 seconds.⁴⁰

CT Source for Attenuation Correction

PET-CT systems can employ a conventional multi-slice CT system, capable of acquiring high-resolution, diagnostic quality data and/or low dose, AC specific x-ray tubes. CT-derived attenuation maps have the advantage of being able to acquire a high signal-to-noise transmission image in a short period of time. Despite this, CT-specific artifacts can make PET/CT more challenging than line source AC (see Fig. 4-13).⁶⁷ Patient motion and breathing introduce the greatest challenges. Several approaches have been proposed such as end expiration breath holding⁶⁸ and shallow free breathing. Cine/CT has been demonstrated to be both robust and easy to implement.⁶⁹ By using several passes over the diaphragm, an average position for the diaphragm can be obtained. Though this does reduce breathing artifacts, it requires rescanning the same tissue areas several times to cover the entire respiratory cycle, thus increasing potential radiation dose. Lowering kVp and tube current can help reduce the dose, however these improvements still result in CTAC scans of over 2 mSv.

Radiation exposure for the CT portion of a PET/CT scan can be a significant contributor to the overall radiation of the patient study and can be as high as 8 mSv when dose optimization is not used.⁷⁰ Optimizing the acquisition setting is crucial for obtaining the best transmission study at the lowest possible radiation exposure⁷¹:

• Optimize the CT scanner acquisition parameters. The x-ray tube voltage (kVp, which determines the mean energy of the x-rays produced by the x-ray tube) and tube current (mA which is proportional to the photon output of the x-ray tube). These should be less than 100 kVP and 10 mAs for most applications

• When possible, prospective ECG triggering should be used.

• If helical scanning is required, dose modulation should be employed.

• Field-of-view settings for obtaining the AC scan should be confined to the cardiac region only without truncating the heart.

• Cine-CT free breathing protocols should be avoided unless the PET/CT system is designed by the manufacturer to perform this scan at radiation dose comparable to either ECG triggering or helical scanning with dose modulation.

One of the most significant sources of artifact in cardiac PET is misregistration of the transmission and emission datasets in space. When a patient moves between these two scans, the AC cannot be properly applied without adjusting for the patient movement (Fig. 4-14). This artifact is thought to be significant in as many as 40% of all PET/CT studies, thus if software is not available to align the images, a substantially lower specificity will occur.⁷² Imaging guidelines recommended that all cardiac PET studies be routinely examined for misregistration and corrected whenever possible.²

By far the most common technique for misregistration correction is a rigid shift of the transmission and emission datasets.⁷³ The user can interactively visualize an overlay of the transmission and emission data and move one of the datasets relative to the other until a satisfactory positioning is obtained. These offsets can then be used to re-reconstruct the tomographic data.

Scatter Correction

When the sinogram data are acquired in the 2D mode (septa extended), the scatter fraction defined as the ratio between the number of scatter events and the total number of coincidence events (scattered and unscattered) is from 10% to 15%, depending on the scanner geometry, the energy window, and the patient sizes.⁷⁴ With septa retracted (no septa in the 3D mode), the scatter fraction can even increase to 30% to 50%, additionally depending on the maximum ring difference and span of the 3D data acquisition. Currently, most commercialized PET scanners offer options for scatter correction for the demand of imaging accuracy. For 2D PET imaging, implementation of nonstationary convolution–subtraction method has been used.⁷⁴ The methods for 3D scatter correction are primarily relied on physical models for simulating the scatter process of the entire images.^74–76 The 3D scatter simulation uses the activity images obtained from reconstructing uncorrected emission sinogram as the initial input of activity distribution to estimate the scatter component. The attenuation map functions to produce 3D scattering probabilities for each LOR of scatter events, based on the Klein–Nishina formula.^77,78

Randoms Corrections

Random events (randoms) are a result of unrelated photon events being recorded within the timing acceptance window of the system. Random events degrade image contrast, however because random photons are from unrelated annihilation events they only reduce overall image contrast. Hence, most randoms correction algorithm are implemented by subtracting a constant from the entire image.⁷⁹

Protocols for Cardiac PET Perfusion Acquisition

The American Society of Nuclear Cardiology has developed imaging guidelines for the development of cardiac PET imaging protocols and should be consulted when developing clinical protocols.⁶⁶ For a detailed explanation of protocols, see Chapter 10. Almost all cardiac PET perfusion studies are performed using vasodilator stress (although dobutamine may also be used). Patients should refrain from using of any product containing caffeine and other substance that could interfere with the vasodilator. For specific requirements, practitioners should refer the package insert of the vasodilator and national guidelines.⁸⁰

For most studies, this single infusion site can be used for both the vasodilator and the ⁸²Rb, particularly dipyridamole or regadenoson. A notable exception is when adenosine is used as the vasodilator. Because of the short half-life of adenosine, it is impractical to switch between the low flow rate adenosine to the high flow rate ⁸²Rb without bolusing the adenosine into the heart or interrupting the flow of adenosine. Stress testing using ¹³N ammonia is considerably less restrictive than Rb-82, allowing for vasodilator and exercise testing.

Prior to the study, technologists will need to apply ECG patches for both the 12-lead cardiac monitoring system and 3-lead ECG triggering system for the scanner. Once patients have been prepared, they are positioned supine in the scanner system. Because AC is applied in all cardiac PET studies, it is possible to image patients with an arm down, however, it is important that the arm is immobilized and is free from the infusion activity.

The first stage of imaging is the transmission study, either acquired using CT or line source. For PET/CT systems, a low dosage chest planar image for positioning followed by a low dosage chest CT for AC will then be acquired of the mediastinal region. Can you explain how you use the map for positioning? For dedicated PET systems, a line source transmission study is acquired for both positioning and AC. How do you position? The transmission studies must be repeated if patient positioning is not correct.

The most common PET radiotracer for perfusion studies is Rubidium-82. Generally for this tracer, the resting perfusion study typically follows the transmission study. Delivery of the ⁸²Rb using a generator cart requires strict adherence to the manufacturer's recommendations for dose delivery, QC, and general operations. Laboratories using an ⁸²Rb generator should obtain training from the manufacturer prior to using the generator cart. To minimize the chance of patient motion between the emission and transmission study, it is important to begin the emission study as soon as practical. Most imaging protocols recommend a delay between the end of the infusion to the start of imaging of 90 to 120 seconds. Studies that measure flow however, require that the dynamic flow study be started prior to the infusion.^81–83 The stress protocol immediately follows the rest acquisition.

These imaging studies can be acquired either using a list mode (inclusion of all photon pair events) or a frame mode. In principle, these two modes will yield similar results, however list mode acquisitions do offer some additional flexibility to create multiple studies from a single data acquisition. Practically, frame mode and list mode acquisitions do not appear to be different either quantitatively or visually.⁸⁴

Reconstruction

Because AC is applied prior to reconstruction, either FBP or iterative reconstruction can be used for creating 3D tomograms. Early PET systems utilized an FBP algorithm for reconstruction. More recently modern reconstruction algorithms for PET utilize an iterative, ordered subsets/expectation maximization algorithm because of the improved noise and image quality properties of the final reconstructed volumes.^37,38

Iterative reconstruction algorithms use a stepwise updating algorithm to "search" for a source distribution that could produce the projection data observed. It relies on a model (the projector), of the transport of photons through the patient to the camera. In principle, this projector can be used to model any physical process in the photon transport, thus giving it considerably more flexibility than the FBP algorithm.

The iterative reconstructions PET in principle are similar to the reconstruction algorithms used in SPECT except the projection model is different. For a standard 2D acquisition, straight forward application of MLEM or OSEM is sufficient. If data are acquired in 3D, additional steps must be applied to take into account the oblique planes of data.

The simplest technique for reconstructing 3D data is to use the Fourier rebinning algorithm.⁸⁵ This algorithm uses the geometric properties of the oblique planes to create a translation of the 3D data into traditional 2D planes.

Time-of-Flight PET

Localization of the point of annihilation using the 511-keV photon pair to a line through the patient is improved upon by using "time-of-flight" (TOF) measurement capabilities.⁵¹ In this approach, additional information about the location of the annihilation event is obtained by measuring the (very small) difference in arrival times of the two coincidence photons. With conventional coincidence detection, the small difference in arrival times is ignored. TOF imaging requires very precise detector and electronic circuitry and calibration given the speed of light and the very short distances that the photons travel. Information provided by the difference in arrival times is included in reconstruction algorithms that model this modified probability information to improve spatial resolution and system sensitivity compared with conventional coincidence detection.⁸⁶ It may also be used to improve separation of scattered photons from true coincidence events and system contrast resolution.

An understanding of the physics and related technical principles for conventional and evolving SPECT is essential for obtaining the highest-quality clinical utilization in nuclear cardiology. From image interpretation to laboratory management, the clinician will be called upon to understand these principles in daily routine. Basic physics, statistics, mathematics, and radiation physics all play a vital role in daily service that provides the optimal patient care, provides the required and regulatory requirements for accreditation as regionally required, and sustains the competitiveness of the specialty.

Nuclear cardiology imaging instrumentation currently offers a great diversity of options. Current systems are the result of continued efforts to develop hardware/detector configurations and reconstruction software that have addressed specific challenges to imaging the heart. The integration of CT options has enhanced conventional perfusion imaging, permitted new protocols through AC, and broadened the scope of services for diagnosing and managing cardiovascular disease with noninvasive techniques. Notably, PET myocardial perfusion imaging is poised to leverage its strengths in nontraditional laboratory environments also in response to pressure for improved efficiency and accuracy of performing studies.