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Instrumentation for Single-Photon Emission Imaging

  • Pat ZanzonicoEmail author
Living reference work entry

Abstract

This chapter reviews the underlying physical principles of single-photon emission tomography (SPECT) and the design and operation as well as the capabilities and limitations of SPECT scanners used clinically and preclinically [1–4]. This is the “companion” chapter to the earlier chapter on positron emission tomography (PET) instrumentation; the material in that chapter on the basic principles of radiation detectors, iterative image reconstruction, and multimodality imaging applies as well to SPECT.

Keywords

SPECT Scintillation detectors Ionization detectors Gamma cameras Filtered back-projection Cardiac imaging 

Glossary

Glossary

ACF

Attenuation correction factor

APD

Avalanche photodiode

ASIC

Application-specific integrated circuit

COR

Center of rotation

CsI(Tl)

Thallium-doped cesium iodide

CT

X-ray computed tomography

CZT

Cadmium zinc telluride

ESSE

Effective scatter source estimation

eV

Electron volt

FBP

Filtered back-projection

FL

Focal length

FOV

Field of view

FWHM

Full-width half-maximum

GAP

General all-purpose

keV

Kilo-electron volt (103 eV)

LEAP

Low-energy all-purpose

LEHR

Low-energy high-resolution

LEHS

low-energy high-sensitivity

LOR

Line of response

MeV

Mega-electron volt (106 eV)

MR

Magnetic resonance

MRI

Magnetic resonance imaging

NaI(Tl)

Thallium-doped sodium iodide

PET-MR

Positron emission tomography-magnetic resonance

PET

Positron emission tomography

PET/CT

Positron emission tomography/computed tomography

PMT

Photomultiplier tube

RF

Radiofrequency

ROI

Region of interest

SAD

Source-to-aperture distance

SPECT-MR

Single-photon emission computed tomography-magnetic resonance

SPECT

Single-photon emission computed tomography

SPECT/CT

Single-photon emission computed tomography/computed tomography

TEW

Triple-energy window

Introduction

This chapter reviews the underlying physical principles of single-photon emission tomography (SPECT) and the design and operation as well as the capabilities and limitations of SPECT scanners used clinically and preclinically [1, 2, 3, 4]. This is the “companion” chapter to the earlier chapter on positron emission tomography (PET) instrumentation; the material in that chapter on the basic principles of radiation detectors, iterative image reconstruction, and multimodality imaging applies as well to SPECT.

SPECT, like PET, is now a mature but continually improving technology. Among the most noteworthy refinements in recent years have been the application of solid-state ionization detector (i.e., semiconductor)-based gamma cameras, the introduction of novel, application-specific detector and collimator configurations (most notably in nuclear cardiology), and the improved quantitative accuracy.

Radionuclides for SPECT

Implicit in the suitability of a radionuclide for in vivo imaging is that it emits insufficient abundance radiations penetrating enough to escape from the body and interact with external detectors. These emissions include γ-rays and characteristic x-rays (“single photons”) used for SPECT and planar gamma camera imaging generally or the 511-keV annihilation γ-rays associated with positron (β+) decay and used for PET. In addition, imaging radionuclides ideally should emit few or no non-penetrating radiations, that is, particulate radiations such as β particles and electrons (excluding, of course, positrons necessary for signal generation in the case of PET). Such particulate radiations typically have ranges in tissue of the order of 1 mm or less and thus cannot escape from the body and be detected externally. These contribute to the radiation dose to tissues and organs without providing any imageable signal. (High-energy (≥ ~1-MeV) beta particles (e.g., those emitted by the pure beta-particle emitter yttrium-90) produce a small but imageable amount of bremsstrahlung (“braking radiation” x-rays) as they slow down in tissue. Bremsstrahlung imaging is not widely used, however, and produces rather poor-quality images. Further, an imaging radionuclide ideally should have a physical half-life comparable to the time required for the administered radiotracer to localize in the tissue of interest. This will provide sufficient time for it to localize in that tissue while still retaining a near-maximal imaging signal, followed by elimination of the radioactivity by a combination of physical decay and biological clearance. The radiation dose to patients and individuals around them and the potential problem of radioactive contamination are thereby minimized. The pertinent physical properties of single-photon radionuclides for imaging are summarized in Table 1 [1, 2, 5, 6].
Table 1

Physical properties of single-photon-emitting radionuclides used in gamma camera imaging and SPECT [1, 2, 5, 6]

Radionuclide

Physical half-life

Decay mode

Energies of imageable x- and γ-rays (MeV)

Abundance imageable x- and γ-rays (%)

x- and γ-ray-to-β particle, conversion electron, and Auger electron energy-per-decay ratioa

Production

Gallium-67

3.26 days

Electron capture

0.0933

40

4.5

Cyclotron

0.185

24

0.300

16

Technetium-99m

6.01 h

Isomeric transition

0.140

89

7.8

Generator (Molybdenum-99)

Indium-111

2.83 days

Electron capture

0.172

90

12

Cyclotron

0.247

94

Iodine-123

13.2 h

Electron capture

0.159

84

6.1

 

Iodine-131

8.04 days

β decay

0.364

82

2.0

Reactor

aAs discussed, the ratio of the energy emitted per decay in the form of β particles, conversion electrons, and Auger electrons (i.e., non-penetrating particulate radiations) to that emitted in the form of x- and γ-rays should be high for a radionuclide for gamma camera imaging

For planar imaging as well as SPECT, a radionuclide emitting x- and γ-rays with energies of 100–200 keV and with an abundance of 100% (i.e., one x- or γ-ray emitted per decay) and with minimal emission of particulate radiations and of higher-energy x- and γ-rays is ideal. The abundance of imageable x- or γ-rays relative to particulate radiations is expressed by the x- and γ-ray-to-β particle, conversion electron, and Auger electron energy-per-decay ratio in Table 1; the higher the value of this parameter, the better a radionuclide will be for imaging. With a mean free path in soft tissue of the order of 10 cm and in NaI(Tl) of less than 0.5 cm, 100- to 200-keV photons provide adequate penetrability through tissue but are low enough in energy to be efficiently collimated (see below) and stopped in relatively thin scintillation and even ionization detectors, yielding optimum-quality images with reasonably low radiation doses. The absence of higher-energy x- and γ-rays (i.e., with energies in excess of several hundred keV) is important because such radiations cannot be efficiently collimated and detected. Yet, they may undergo scatter in the patient and/or detector hardware and contribute, even with energy discrimination, mispositioned, and otherwise spurious counts to the image. Based on the foregoing criteria (Table 1), 99mTc is a near-ideal radionuclide for gamma camera imaging, emitting only a 140-keV γ-ray and little particulate radiations. Iodine-131 (131I), on the other hand, emits a relatively high-energy, difficult-to-collimate 364-keV γ-ray and abundant β particles.

Gamma Camera Design

Developed in the late 1950s by Hal Anger, the gamma camera (Fig. 1), also known as the scintillation or Anger camera , has long been the predominant imaging device in nuclear medicine [1, 3]. Its large detector area allows simultaneous and therefore rapid data acquisition over a large area of the body. Gamma camera crystals, almost universally, are composed of a plate of NaI(Tl) and vary in thickness from 6.35 to 25.4 mm. A 9.53-mm-thick crystal thus provides a reasonable balance between sensitivity and resolution and is the most widely used for general gamma camera imaging. About 95% of the 140-keV photons from 99mTc are absorbed in a 9.54-mm crystal. For clinical applications, gamma camera crystals are most commonly rectangular in shape and ~50 × 60 cm in area for general-purpose imaging.
Fig. 1

Basic design of a conventional gamma camera, consisting of a multi-hole collimator, a thin large-area NaI(Tl) crystal, a two-dimensional array of PMTs and associated electronics (high-voltage power supply, preamplifier, and amplifier), position logic circuitry, energy discriminator, and image display. Note that there are actually two position logic circuits – for the determination separately of the x- and y-positions of the scintillation within the crystal. Note further that the output signal from each PMT is actually split into three parts, one (the z pulse) for the determination of the energy of the incident radiation as well as one each for the determination of its x- and y-positions. In current-day gamma cameras, the output signal from each PMT is digitized, and the position of each event is determined with computer software. The left inset shows a photograph of the two-dimensional PMT array backing the crystal in a typical rectangular field-of-view gamma camera. The right inset shows a drawing of a portion of a parallel-hole collimator, identifying the dimensions – aperture diameter, septal thickness, and septal length – of such a collimator. The “desirable” events (arrows labeled “1”) are unscattered (i.e., photopeak) photons traveling in a direction parallel or nearly parallel to the axes of the apertures and thus yielding correctly positioned counts in the gamma camera image. “Undesirable” events include (1) scattered as well as unscattered photons traveling in a direction oblique to the axes of the apertures (2) and thus eliminated by attenuation by one or more collimator septa; septal penetration (3), that is, unscattered photons traveling in a direction oblique to the axes of the apertures yet passing through the septa and yielding mispositioned counts; scatter (4), that is, photons undergoing Compton scatter within the patient and either eliminated by energy discrimination or not eliminated and yielding mispositioned counts (From Refs. [1, 3] with permission)

Once the incident radiation passes through the collimator (see below), it strikes, and may produce a scintillation within, the crystal. The resulting light signal is distributed among a two-dimensional array of photomultiplier tubes (PMTs) backing the crystal, the light intensity varying inversely with the distance between the scintillation and the respective PMT : the farther the PMT is from the scintillation, the less light it receives, and the smaller is its output pulse. This inverse relationship is the basis of the Anger position logic circuitry for determining the precise location of a scintillation within the crystal. In the older gamma cameras, the x- and y-coordinates were calculated by analog circuitry, that is, using matrices of resistors. In current models, this is done by digitizing the output signal from each PMT and using digital electronics.

The gamma camera collimator, almost always comprised of lead, “directionalizes” the incoming radiation (Fig. 1). Any radiation traveling at an oblique angle to the axes of the holes (apertures) will strike the lead walls (septa) between the holes and not reach the crystal. As a result, only radiations traveling perpendicular to the crystal surface pass through the apertures and contribute counts to the resulting image. A certain fraction of photons striking the septa will nonetheless pass through them and reach the crystal; this phenomenon, which degrades image quality, is known as septal penetration . Most collimators used clinically are parallel-hole collimators, with the apertures and septa parallel to one another. In addition, single-aperture pinhole collimators, most commonly used for thyroid imaging because of their pronounced magnifying effect, are available as well (Fig. 2a). Pinhole collimators, however, suffer from low sensitivity, limited field of view, and geometric distortion, since image magnification varies with both source-aperture distance and lateral position in the field of view, and therefore are rarely used for clinical imaging other than for the thyroid (normally a small, relatively flat organ). For preclinical (i.e., mouse and rat) imaging, multi-aperture pinhole collimation is now the standard approach, since it combines the magnification effect (and improved resolution) of pinhole collimation with the greater sensitivity afforded by multiple apertures (Fig. 2b).
Fig. 2

Pinhole collimation for gamma cameras. (a) A single-aperture pinhole collimator. Pinhole collimation provides a magnification effect and thereby improved “effective” resolution. For example, if a source were placed at distance of 3 cm from the aperture (i.e., source-to-aperture distance (SAD) = 3 cm) of a pinhole collimator having a focal length of 9 cm (i.e., focal length (FL) = 9 cm) fitted to a detector with an intrinsic spatial resolution of 3 mm (i.e., FWHM = 3 mm), the magnification is FL/SAD = 9/3 cm = 3, and the effective resolution therefore improves to FWHM/(FL/SAD) = 3 mm/(9/3 cm) = 3 mm/3 = 1 mm (Photograph courtesy of Gamma Medica (Northridge, CA)) (b) A multi-aperture pinhole collimator, now the standard approach to preclinical (i.e., mouse and rat) SPECT imaging. As shown, however, each of the apertures produces an image, and the resulting multiple images – one per aperture – overlap. The apertures are angled with respect to one another to minimize this overlap. If larger-area detectors are used, this angulation can be increased and therefore the image overlap further minimized and image quality improved (Courtesy of Bioscan (Washington, DC))

Gamma camera collimators are “rated” with respect to photon energy and resolution/sensitivity. Low-energy, or “technetium,” collimators, including “low-energy all-purpose (LEAP)” (or “general all-purpose (GAP)”), “low-energy high-resolution (LEHR),” and “low-energy high-sensitivity (LEHS)” collimators, are designed to image radionuclides emitting x- and γ-rays less than 200 keV in energy. These include 99mTc (photopeak energy: 140 keV) as well as thallium-201 (201Tl) (68–80 keV) and iodine-123(123I) (159 keV). Medium-energy, or “gallium,” collimators are designed for radionuclides emitting x- and γ-rays 200–300 keV in energy, including gallium-67 (67Ga) (93, 185, and 300 keV) as well as indium-111 (111In) (172 and 247 keV). High-energy, or “iodine,” collimators are designed to image radionuclides emitting x- and γ-rays greater than 300 keV in energy, including 131I (364 keV). In progressing from low- to medium- to high-energy collimation, the collimators are made longer and the septa thicker in order to interpose more lead between the subject and the crystal. This is done in order to maintain septal penetration (i.e., expressed as the percent of counts in an image attributable to photons penetrating the septa) at or below an acceptably low level, typically set at 5%. This, in turn, reduces the overall fraction of emitted x- and γ-rays reaching the crystal. To compensate, at least in part, for the resulting lower sensitivity, the apertures are made wider in progressing from low- to medium- to high-energy collimators. This, however, degrades spatial resolution by dispersing the counts passing through each aperture over a large area of the crystal (Figs. 3 and 4). Overall, therefore, gamma camera images are progressively poorer in quality for radionuclides emitting low- versus medium- versus high-energy x- and γ-rays. For each energy rating (and as indicated above for low-energy collimators), collimators may also be further rated as “general purpose” (or “all purpose”), “high resolution,” or “high sensitivity.” High-resolution collimators have narrower apertures (and therefore lower sensitivity), and high-sensitivity collimators have wider apertures (and therefore coarser resolution), respectively, than general-purpose collimators. In instances where a radionuclide emits multiple photons, it is the highest-energy photon which dictates the collimator to be used. For example, a medium-energy collimator must still be used to image gallium-67 (photon energies: 93, 185, and 300 keV) even if only the two lower-energy (i.e., the 93 and 185 keV) photons are used for imaging.
Fig. 3

(a, b) In gamma camera imaging with parallel-hole collimation, the 3D volume subtended by a point source of radiation is a cone, with the area of the detector over which radiations from the source are dispersed increasing, and therefore the spatial resolution worsens, with increasing source-to-collimator distance (Adapted from Ref. [40] with permission)

Fig. 4

Collimator spatial resolution, FWHMcollimator, for a parallel-hole collimator with thickness t, aperture diameter d, septal thickness a, and source-to-collimator face distance b. The collimator is fitted to a detector whose midplane is at a distance c from the back surface of the collimator (From Ref. [9] with permission)

Solid-state ionization (i.e., semiconductor)-based gamma cameras utilizing cadmium zinc telluride (or CZT) are now commercially available [7, 8]. CZT detectors are more expensive than detectors based on NaI(Tl) crystal coupled with PMTs in the conventional Anger camera but offer several important advantages, including improved energy resolution (typically resulting in a 30% reduction of the scatter counts in the photopeak energy window) and superior intrinsic spatial resolution [7, 8]. The mass density of CZT is higher than that of NaI(Tl) (5.8 vs. 3.7 g/cc), but due to considerations of the higher cost and in order to maintain excellent energy resolution, CZT detectors are thinner (~5 mm) than the typical 3-/8-in. (9.5-mm) NaI(Tl) crystals. The intrinsic sensitivities of CZT and NaI(Tl) are therefore comparable. However, the compact form factor of CZT detectors (Fig. 5) allows for novel detector geometries and, in turn, high-sensitivity collimation.
Fig. 5

Typical configuration of a CZT-based semiconductor detector . ASIC, application-specific integrated circuit (From Ref. [8] with permission)

The full-width half-maximum (FWHM) spatial resolution , FWHM system, of gamma cameras is determined by a combination of physical and instrumentation factors. Intrinsic resolution, FWHM intrinsic, is the component of spatial resolution contributed by the crystal and associated electronics and is related to statistical fluctuations in pulse formation; typical values are of the order of 5 mm. These statistical fluctuations include variations in the production of light photons resulting from x- or γ-ray interactions in the crystal and in the number of electrons emitted by the photocathode and the series of dynodes in the PMTs. Collimator, or geometric, resolution, FWHMcollimator, represents the major contribution to system resolution and is determined by the collimator design (Figs. 3a and 4). The spatial resolution of a parallel-hole collimator, the most common type, is determined by the geometric radius of acceptance of each aperture, as illustrated in Fig. 4:
$$ F W H{M}_{\mathrm{collimator}}=\frac{d\left({t}_e+ b+ c\right)}{t_e} $$
(1)
where d is the diameter of the collimator aperture, b the source-to-collimator face distance, c the collimator back-to-mid-crystal distance, and t e = t − 2/μ the effective thickness of the collimator where t is the actual collimator thickness and μ is the radiation energy-dependent linear attenuation coefficient of the collimator material [9]. As indicated by Eq. 1, collimator resolution is improved (i.e., lowered) by reducing the diameter d of the collimator aperture, the source-to-collimator face distance b, and the collimator thickness t. System resolution is further degraded by the contributions of septal penetration resolution, FWHMpenetration, and by scatter resolution, FWHMscatter. The spatial resolution of a gamma camera system, FWHMsystem, can be obtained by combining the resolution of the respective components of the system [9]:
$$ F W H{M}_{\mathrm{system}}=\sqrt{FWH{M_{\mathrm{intrinsic}}}^2+ FWH{M_{\mathrm{collimator}}}^2+ FWH{M_{\mathrm{penetration}}}^2+ FWH{M_{\mathrm{scatter}}}^2} $$
(2)

In clinical practice, the collimator resolution generally dominates the system resolution.

SPECT Data Acquisition

Although there are many possible combinations of detector number, geometry, and motion that can acquire the necessary projection data, rotating gamma camera-based SPECT systems are the most common [4]. Nowadays, two-detector systems predominate clinically and two- to four-detector systems preclinically. The basic SPECT imaging paradigm includes acquisition of planar projection images from multiple angles around the subject, correction of the acquired data for nonuniform scanner response and possibly other signal-degrading effects, and mathematical reconstruction of thin (several-millimeter-thick) transverse tissue section images [4]. The raw data are acquired as a series of discrete planar images at multiple angles about the longitudinal axis of the patient (Fig. 6). The number of counts recorded in each projection-image pixel represents the ray sum, or line integral, of the sampling line perpendicular to and extending from the detector through the subject. The following are typical SPECT acquisition parameters: 20–30 min for data acquisition; ~60 to ~120 projection images at ~6 to ~3° angular increments, respectively; and a 180 or 360° rotation for cardiac or noncardiac studies, respectively. An angular increment in excess of 6° between successive projection images will result in prohibitive undersampling artifacts in the reconstructed images. Because of the length of time (20–30 min) required for an acquisition, dynamic SPECT and whole-body SPECT remain largely impractical at the current time. It should be emphasized that SPECT images can in principle be quantitative in absolute terms, with voxel values representing the local activity concentration [10, 11, 12]. However, in contrast to PET, this is often not the case in routine practice because of the confounding effect of scatter and attenuation as well as limited count statistics. Accurate correction for these effects remains more challenging in SPECT than in PET [10, 11].
Fig. 6

(a) The basic data acquisition paradigm in rotating gamma camera SPECT. Photographs of a dual-detector gamma camera, with (b) the two detectors in opposed positions, as routinely used for a 360° rotation and general (noncardiac) SPECT and (c) the two detectors perpendicular to each other, as routinely used for a 180° rotation and cardiac SPECT (with projection images acquired from approximately right anterior oblique to left posterior oblique). The advantage of such two-detector systems is that two projection images can be acquired simultaneously and, in first order, either the acquisition time halved or the acquired counts doubled (From Ref. [4] with permission)

Although PET offers important advantages over SPECT (i.e., generally better spatial resolution, higher sensitivity, and more accurate activity quantitation), SPECT offers the capability of multi-isotope imaging. Because different SPECT radionuclides emit x- and γ-rays of different energies, multiple isotopes and therefore multiple radiotracers can be imaged simultaneously using distinct, isotope-specific photopeak energy windows.

Data “Corrections”

Uniformity correction. Even optimally performing SPECT scanners exhibit some nonuniformity of response [4, 13, 14]. Among the thousands of pixels in a SPECT projection image, slight variations in detector thickness, light emission or coupling properties, electronics performance, etc. result in slightly different measured count rates for the same activity. In principle, such nonuniform response can be corrected by acquiring data from a uniform flux of γ-rays and normalizing to the mean count rate from all the pixels in SPECT. This “uniformity map” corrects for the nonuniform count rate of the individual pixels or lines of response (LORs) to thereby yield a pixel-by-pixel uniformity correction. The effects of nonuniformity of scanner response and of the correction for nonuniform response are illustrated in Fig. 7 [14].
Fig. 7

Uniformity correction and its effects. The left panel shows gamma camera image of a uniform source of radioactivity without the uniformity correction applied. The pattern of PMTs is apparent in this grossly nonuniform, uncorrected image. In contrast, the corrected image in the right panel is uniform

For gamma camera imaging, a correction table may be acquired using either a uniform flood source placed on the detector or a point source placed sufficiently far (typically 1–2 m) from the uncollimated detector to deliver a uniform photon flux. Acquisition of the data required for uniformity correction is somewhat problematic in practice because of statistical considerations: tens of millions must be acquired to avoid possible “noise”-related artifacts in the uniformity correction table.

Deadtime correction. In SPECT, deadtime count losses are generally minimal at diagnostic administered activities. While real-time correction for deadtime count losses is routinely applied in PET, this is generally not the case in SPECT. Count rates encountered in PET are much higher than in SPECT – in part because of the use of electronic rather than absorptive collimation – and therefore accurate deadtime correction is more critical in PET.

Center-of-rotation misalignment correction. In rotating gamma camera SPECT, the location of the projection of the center of rotation (COR) on the projection-image matrix must be constant [14, 15, 16]. If the mechanical and electronic CORs are aligned, the pixel location of the projection of the COR onto the projection-image matrix will be the same for all projection images, and, for all such images, the counts in each pixel will then be projected across the appropriate row of pixels in the tomographic image matrix. If, however, the mechanical and electronic CORs are not aligned, the pixel location of the COR will vary among the projection images, and the counts in each projection-image pixel will be projected across different locations in the tomographic image matrix, and blurred images will result (Fig. 8a). In today’s SPECT systems, COR misalignment may be easily measured, and corrections are created and automatically applied using the system’s software (Fig. 8b). In contrast, PET scanners typically utilize fixed rings of detectors and thus do not suffer from COR misalignment.
Fig. 8

(a) Center-of-rotation (COR) misalignment and resulting image-blurring artifacts in rotating gamma camera SPECT. The degree of blurring is related to the magnitude of the spatial misalignment of the mechanical and electronic CORs. A misalignment as small as 3.2 mm (or ½ a pixel for a 64 × 64 image matrix) can produce perceptible blurring in SPECT images, with the blurring substantially worse for a misalignment of 6.4 mm (1 pixel) (Adapted from Ref. [15] with permission). Note that for a cross-sectional image of a line source, COR misalignment blurs the expected point into a full or partial circle depending on the position of the source in the FOV: if it is at or near the center of the FOV, the line source appears as a full circle in cross section; if it is near the periphery of the FOV, it appears as a partial circle. (b) COR misalignment can be measured and corrected based on acquiring a 360° circular SPECT study of a 99mTc point source and constructing graphs of the x- and y-positions (perpendicular and parallel to the axis of rotation, respectively) of the position of the maximum-count pixel in each projection image versus angular position. The x- and y-position-versus-angle graphs should be a sinusoidal curve and a straight line, respectively. The angle-by-angle deviation between the x-position on the best-fit sine curve and the x-position of the actual maximum-count pixel thus yields a correction table indicating the offset by which each projection image must be shifted at each angular position to align the CORs. Alternatively, the average of the offsets may be used at each angular position (Adapted from Ref. [14, 15] with permission)

Scatter correction. Scatter results in generally diffuse background counts in reconstructed images, reducing contrast and distorting the relationship between image intensity and activity concentration [10, 17, 18]. Several methods, characterized as energy distribution- or spatial distribution-based methods, have been proposed for SPECT scatter correction [12]. However, only the dual-energy window-based or triple-energy window (TEW)-based estimation and the effective scatter source estimation (ESSE) method have been implemented clinically. In multi-energy window-based methods [19] (Fig. 9), the number of scattered photons in the photopeak energy window is estimated based on counts in one or more scatter windows. In the TEW method , for example, the scatter is estimated as the area of the trapezoid beneath the line joining the two adjacent narrow scatter windows; for radionuclides emitting a single x- or γ-ray, the upper scatter window can be omitted. For each projection-image pixel i,j, the number of scatter counts in the photopeak window, (C i,j )scatter, is calculated as:
Fig. 9

Dual-energy windows (shaded boxes) superimposed on the energy spectrum of unscattered and scattered events for a patient-sized phantom filled with 99mTc. In contrast to 99mTc, a radionuclide such as 131I emits a higher-energy γ-ray which will produce scattered events in the photopeak (unscattered-event) energy window in addition to those from scatter of the primary “imaging” γ-ray. For such radionuclides, the dual-energy window method may be generalized to the triple-energy window (TEW) method by using a second scatter energy window (indicated by the dashed-line box) immediately above the photopeak window (Adapted from Ref. [40] with permission)

$$ {\left({C}_{i, j}\right)}_{\mathrm{scatter}}=\left\{\frac{{\left({C}_{i, j}\right)}_{\mathrm{lower}}}{W_{\mathrm{lower}}}+\frac{{\left({C}_{i, j}\right)}_{\mathrm{upper}}}{W_{\mathrm{lower}}}\right\}\frac{W_{\mathrm{photopeak}}}{2} $$
(3)
where (C i,j )lower and (C i,j )upper are the counts in the lower and upper scatter energy windows, respectively, for projection pixel i,j and W photopeak, W lower, and W upper are the widths (in keV) of the photopeak, lower scatter, and upper scatter windows, respectively. Typically, 20% photopeak and ~5% scatter energy windows are used. Because of the use of such narrow scatter energy windows, the TEW estimate of scatter counts tends to be noisy. In the ESSE method [20], a model-derived scatter function is used to calculate and correct for the scatter in the projection data based on the estimate of the activity distribution. An effective scatter source is calculated, and the attenuated projection of that source yields the scatter component of the SPECT projections. Other, more sophisticated and presumably more accurate approaches to scatter modeling include Monte Carlo simulation-based methods [21, 22, 23] and analytic methods using the Klein-Nishina formula for Compton scattering [24]. These methods are computationally intensive, however, and have been limited to use in research settings but likely will become clinically practical as computing power continues to increase.

Attenuation correction. As in PET, correction for the attenuation of the γ-rays as they pass through tissue is generally the largest correction in SPECT. The correction factors for a 99mTc SPECT scan of the brain (roughly equivalent to a 10-cm diameter water-equivalent cylinder) are ~2, for example. The magnitude of the correction depends on the energy of the γ-rays, the thickness of tissue(s) that the γ-rays must travel through, and the attenuation characteristics of the tissue(s).

Like scatter corrections, attenuation corrections in SPECT are not yet as well developed or as reliable as those in PET – because, for single photons, the attenuation correction factor (ACF) also depends on the depth of the source [4]. For many years, if attempted at all, SPECT attenuation correction factors were calculated (as in Chang’s first-order correction and the Sorenson method [25, 26]) based on the assumptions – neither of which is generally accurate – that the body is a uniform medium with a single, well-defined value of μ and that the body’s contour is known. At one time, manufacturers incorporated long-lived radioactive sources (such as gadolinium-153 (153Gd)) into SPECT scanners to perform attenuation correction. As part of the SPECT procedure, a shutter opens at each projection-image angle to expose a highly collimated line source, and a transmission image is acquired. The transmission images thus acquired are reconstructed into an ACF map for correction of the SPECT study. The introduction of hybrid SPECT/CT scanners (see below) is resulting in more practical and more accurate CT-based attenuation correction in SPECT [27].

Image Reconstruction

In emission tomography, the emission data correspond to the projected sum of counts (or line integrals) at various angles about the axis of the scanner. The full set of 2D projection data is usually represented and stored as a two-dimensional matrix in polar coordinates (distance r, angle ϕ) known as a “sinogram” [28] (see Fig. 14 in the chapter entitled “Instrumentation for Positron Emission Tomography”).

There are two basic classes of image reconstruction methods, analytic [4, 28] and iterative [29]. For PET, iterative reconstruction is used almost exclusively nowadays and was reviewed in the earlier chapter on PET instrumentation. For SPECT, on the other hand, analytic image reconstruction (i.e., filtered back projection [4, 28]) is still widely used, though the use of iterative methods is increasing. The basic procedure for image reconstruction by filtered back-projection (FBP) is as follows: each angular projection is Fourier transformed from real to frequency (or “k”) space; the projection is filtered in frequency space using a ramp filter (to enhance high and suppress low spatial frequencies and thereby minimize blurring artifacts) (Figs. 10ac); the filtered projection is inverse Fourier transformed from frequency back to real space; and the filtered projection data in real space are uniformly distributed, or back projected, over the reconstructed image matrix [4, 28]. To compensate for the high noise levels resulting from the relatively low counts in most nuclear imaging studies, low-pass, or apodizing, filters (known as Shepp-Logan, Hann, etc.) are used in place of the ramp filter to eliminate those spatial frequencies above a specified cutoff frequency (Fig. 10dg). In this way, the high spatial frequencies characteristic of statistical noise are eliminated (or at least minimized) in the reconstructed images.
Fig. 10

(a) Graphical representation of the ramp and two other mathematical filters used in reconstruction of tomographic images. Note that the cutoff spatial frequencies of all three filters are set to the maximum frequency that can be imaged with the system being used to reduce noise amplification artifacts in the reconstructed images. Note further that the Shepp-Logan and Hann filters “roll off” gradually with increasing spatial frequency to further reduce noise amplification and other artifacts. (b) Cross section through a phantom containing high (white), intermediate (light gray), and low (dark gray) activity concentrations. (c) Tomographic image of phantom reconstructed by “simple” back-projection, that is, with no mathematical filtering of the projection images. Note the resulting blurring and other artifacts in the reconstructed image. (dg) Tomographic images of phantom reconstructed by filtered back-projection using a Shepp-Logan filter with cutoff frequencies equal to 1, 0.8, 0.6, and 0.2 times of the maximum spatial frequency that can be imaged with the system. Note the reduction in noise (i.e., graininess or mottle) but coarser delineation of edges (i.e., poorer resolution) as the cutoff frequency is reduced. For clinical SPECT studies, which generally have limited numbers of counts, cutoff frequencies lower than the maximum imageable frequency are used routinely (Adapted from Ref. [40] with permission)

Quantitation

Once the SPECT emission data have been corrected for dead time, system response (by uniformity correction), scatter, and attenuation, the count rate per voxel in the reconstructed tomographic images is proportional to the local activity concentration [10, 17, 18]. In routine practice, however, scatter and attenuation corrections are generally less accurate for SPECT than for PET (as previously noted) and reconstructed SPECT images less quantitatively accurate than PET images. Further, acquired SPECT data generally include far fewer counts than do PET data because of the use of absorptive collimation for SPECT and electronic collimation for PET, increasing the statistical uncertainties in SPECT.

Another consideration which adversely effects the quantitative accuracy is partial-volume averaging: for a source which is “small” relative to the spatial resolution of the imaging system, the image-derived activity or activity concentration in such a source underestimates the actual value, and the smaller the source the greater the degree of the underestimation [10, 30]. The dimensions of a source must be at least two times the full-width half-maximum (FWHM) spatial resolution of the imaging system to avoid such underestimation of the activity or activity concentration [30]. Since the source size dependence of the partial-volume effect can be measured (e.g., by phantom studies), the underestimated activity or activity concentration can be corrected if the source dimensions can be independently determined (e.g., by CT) [31]. Another key factor which limits quantitative accuracy is, of course, subject motion.

As noted in the earlier chapter on PET instrumentation, under favorable circumstances (i.e., for a high-count study of a large, stationery source region with a high source-to-background ratio), the quantitative accuracy of the PET is of the order of 10% [30]. A comparable accuracy, ~10%, is achievable for SPECT with state-of-the-art attenuation and scatter corrections [32]. With routinely available corrections, however, the quantitative accuracy of SPECT generally exceeds 20% [33].

Multimodality Imaging: SPECT/CT and SPECT-MRI

SPECT/CT scanners are, of course, now commercially available [27]. The paradigm for SPECT/CT scanners is similar to that of PET/CT in that the SPECT and CT gantries are separated and the SPECT and CT scans are acquired sequentially, not simultaneously. In such devices, the ~1-m separation of the SPECT and CT scanners is more apparent because the rotational and other motions of the SPECT detectors effectively preclude encasing them in a single housing with the CT scanner. Multimodality imaging devices for small animals (i.e., rodents) – PET/CT, SPECT/CT, and even SPECT-PET/CT devices – are now commercially available as well.

Unlike PET-MR devices, neither clinical nor preclinical SPECT-MR scanners are commercially available at the present time. Investigational preclinical SPECT-MR devices have been developed, however, and proof-of-principle studies have been performed (Fig. 11) [34, 35, 36, 37]. As an alternative to scintillation detectors coupled to either PMTs or APDs, prototype SPECT-MR scanners have utilized solid-state (i.e., semiconductor) ionization detectors [34, 35, 36, 37]. However, when a semiconductor detector is placed in a magnetic field, electron-hole pairs created by the absorption of x- and γ-rays are subject to the so-called Lorentz force . As a result, when such a detector is placed in any orientation other than parallel or antiparallel to the magnetic field, electrons traveling toward the anode will undergo a shift in their detected position (Fig. 12). For the prototype “MRSPECT” system described by Hamamura et al., for example, a mean “Lorenz shift” of 1.4 mm was measured [34, 35]. In this system, the correction for the Lorentz shift was performed prior to SPECT image reconstruction by shifting the nuclear projection data to their proper locations; this pixel-by-pixel correction was derived by imaging of a uniform flood source, analogous to derivation of gamma camera uniformity corrections generally. The detector elements were coupled to an application-specific integrated circuit (ASIC) readout board, and the detector-ASIC board housing and cables were wrapped with a fine copper mesh for RF shielding.
Fig. 11

Proof-of-principle single-photon scintigraphy-MR imaging study in a mouse. (a) Diagrammatic representation of the system for simultaneous single-photon scintigraphy using a CZT ionization detector and MR imaging. Note that this is not a SPECT (i.e., tomographic) system but rather a planar imaging system. (b) Static non-contrast T1-weighted MR image (upper left panel). Frame from dynamic 99mTc-sestamibi study (upper middle panel). “Fusion” of T1-weighted MR and 99mTc-sestamibi images. Since the 99mTc-sestamibi image is a planar image (upper right panel), the MR and scintigraphic images are not truly fused. (c) Frame from a dynamic MR study prior to injection of gadopentetate dimeglumine contrast (lower left panel). Frame from a dynamic contrast (gadopentetate dimeglumine)-enhanced MR study (lower middle panel). Kidney ROIs extracted from the static non-contrast T1-weighted MR image (lower right panel). (d) Renal signal-versus-time curves following the 99mTc-sestamibi and gadopentetate dimeglumine injections (Adapted from Ref. [37] with permission)

Fig. 12

Diagrammatic illustration of the Lorentz shift. After the interaction of an incident gamma ray with an ionization detector such as CZT, the resulting electron is deflected by a distance Δx due to the Lorentz force. Since the exact depth of the interaction of the gamma ray in the detector is statistical and can occur anywhere along the thickness of the element, there is a range of possible deflected distances Δε (Adapted from Ref. [34] with permission)

Commercial Devices

Clinical. All of the major manufacturers of medical imaging equipment, including General Electric, Philips, Siemens, and Toshiba, currently market clinical SPECT devices. SPECT scanners continue to mainly utilize dual NaI(Tl) detectors (most commonly ~10 mm in thickness) in an opposed configuration; many commercial systems allow rotation of the detectors to an angle of 90° with respect to one another for cardiac studies. SPECT system spatial resolution (FWHM) with low-energy, high-resolution (LEHR) collimation is typically 7–8 mm for 99mTc at a source-to-collimator distance of 10 cm.

Conventional dedicated cardiac SPECT scanners generally employ two scintillation detectors fixed at a right angle to one another and, in some models, radioactive sources for transmission imaging-based, patient-specific attenuation correction [7, 8]. Cardiac imaging typically involves acquisition of projection data over only a 180° orbit (instead of the 360° orbit used for general SPECT imaging) – from right anterior oblique to left posterior oblique – because of the left, anterior position of the heart within the thorax. By positioning the two detectors at a right angle to one another, both detectors can simultaneously acquire usable data over the 180° orbit; if the two detectors were in the usual opposed position, only one of the two detectors would actually acquire usable data. Dedicated cardiac SPECT scanners with two scintillation detectors fixed at a right angle to one another include the Philips CardioMD™ and the Siemens CCam™.

Siemens has also introduced so-called “cardio-focal” (SmartZoom™) field-upgradeable collimators for high-sensitivity cardiac imaging with conventional dual-head gamma cameras (Fig. 13) [7, 8]. Also known as “astigmatic” collimation , with such a converging/diverging collimator, the converging collimation magnifies the center of the field of view (FOV) in both the axial and transaxial directions, while the diverging collimation minifies the periphery of the FOV to provide coverage of the entire body and thereby avoid truncation artifacts commonly seen with fanbeam collimators when imaging the torso. Such cardio-focal collimators can increase sensitivity for cardiac imaging by approximately twofold compared to that with high-resolution parallel-hole collimator.
Fig. 13

Data acquisition geometry with the Siemens SmartZoom™ cardio-focal (diverging/converging) collimator. It provides higher sensitivity in the heart region by selective maximal magnification of that region (From Ref. [7] with permission)

The Digirad Cardius 3 XPO™ cardiac scanner employs three pixilated CsI(Tl) scintillation detectors and images patients in an upright seated position (Fig. 14) [7, 8]; one- and two-detector Cardius systems are available as well. It uses fanbeam collimators in the three-detector configuration coupled to photodiode detectors. The two outer detectors are positioned at a 67.5° angle relative to the central detector. Data are acquired by rotating the chair (rather than detectors) by 67.5°, yielding an acquisition arc of 202.5°. Its SPECT spatial resolution (FWHM) is 11.0 mm (radius of rotation: 20 cm) and sensitivity 160 cpm/μCi for 99mTc and LEHR collimation. The Digirad Cardius X-ACT™ system augments the Cardius 3 XPO™ with a fanbeam x-ray source for CT imaging. The CardiaArc HD-SPECT™ [38] also images patients in an upright seated position. It is comprised of three curved stationery NaI(Tl) crystals backed by three rows of 20 PMTs, with a thin curved lead collimator (the so-called aperture arc) having six vertical slots for horizontal collimation and rotating slowly back and forth every 10 s. The CardiaArc™ quotes a spatial resolution of 3.6 at an 82-mm source-to-detector distance to 7.8 mm at a 337-mm source-to-detector distance.
Fig. 14

Photographs of the Digirad Cardius 3 XPO™ system. The patient is seated upright, and the chair rotates to achieve the required angular sampling (From Ref. [7] with permission)

Two vendors have introduced CZT-based cardiac SPECT systems [7, 8]. The Spectrum Dynamics D-SPECT™ utilizes pixilated CZT detector arrays mounted in nine vertical columns, with four detectors in each column, placed in a 90° gantry geometry (Fig. 15). Each column consists of an array of 1,024 CZT elements (2.46 × 2.46 × 5-mm thick) arranged in a 16 × 64 element array with a size of 40 × 160 mm, achieving 16 cm of coverage of the thorax. Each detector array is fitted with square aperture, high-sensitivity, parallel-hole collimators such that the dimensions of each hole are matched to the size of a single detector element. Spectrum Dynamics also recently introduced a less expensive six-detector column system. In the other CZT-based camera design found in the General Electric Discovery NM 530c™ and Discovery NM/CT 570c™, a curved array of 19 5-mm-thick CZT-pixilated detector arrays each with a focused pinhole collimator is used (Fig. 16). Each detector array consists of four detectors, each with 246 CZT detector elements (2.46 × 2.46 × 5-mm thick) arranged in a 16 × 16 element array. In these two systems, the detector gantry does not rotate or otherwise move [7, 8]. The remarkably high sensitivity of these systems is not related to the intrinsic sensitivity of the CZT detectors but rather the high-sensitivity geometric arrangement and collimation of the detectors.
Fig. 15

(a) Detector configuration (top view) of the Spectrum Dynamics D-SPECT™ system with nine angularly directed detector blocks. (b) Photograph of the system. The patient may be seated upright or the chair reclines so that the patient may lie nearly supine. (c) One of the nine detector blocks. (d) One of the CZT detectors (From Ref. [7] with permission)

Fig. 16

(a) Detector configuration (top and side views) of the General Electric NM 530c™ system with nine detector block arrays. (b) Photograph of the system. The patient lies supine. (c) The pinhole collimation. (d) One of the CZT detectors (Adapted from Ref. [7] with permission)

The foregoing dedicated cardiac scanners provide better spatial resolution and higher sensitivity than conventional gamma cameras, as illustrated in Fig. 17 [7, 8, 39]. These systems provide up to a sevenfold increase in sensitivity and twofold improvement in reconstructed image resolution compared with those of conventional dual-head systems. Cardiac imaging times with these scanners thus range from 2 to 3 min for systems with solid-state detectors to 4 min on scintillation detector systems with cardio-focal collimation [7, 8, 39]. The design and performance parameters of other commercially available clinical SPECT scanners are summarized in Table 2.
Fig. 17

Sensitivity in % (left axis) and reconstructed image resolution in mm (right axis) of dedicated cardiac SPECT systems compared to those of a conventional dual-head system. DSPECT: Spectrum Dynamics DSPECT™ system; NM: GE 530c/570c™ system; IQSPECT: Symbia system with dual-head Siemens SmartZoom™ cardio-focal (diverging/converging) collimators; Cardius: Cardius Digirad X∙ACT™ system with fanbeam collimators; Conventional: conventional dual-headed SPECT with low-energy high-resolution collimators (From Ref. [7] with permission)

Table 2

Design and performance parameters of commercially available clinical SPECT scannersa,b

 

General Electric Discovery NM/CT 670c™

General Electric Infinia Hawkeye™

Philips Brightview™

Philips CardioMD™

Siemens Symbia TruePoint™

Siemens CCam™

Toshiba TCam™

Detector material

NaI(Tl)

NaI(Tl)

NaI(Tl)

NaI(Tl)

NaI(Tl)

NaI(Tl)

NaI(Tl)

Number of detectors

2

2

2

2

2

2

2

Angle between detectorsc

Variable – 90–180°

Variable – 90–180°

Variable – 90–180°

Fixed – 90°

Variable – 90–180°

Fixed – 90°

Variable – 90–180°

Crystal thickness (in)

3/8

3/8

3/8

3/8

3/8

1/3

3/8

Uniformity CFOV IU (%)d,e

3.0

3.0

2.5

2.6

2.9

2.5

3.7

Energy resolution (%)

9.8

9.8

9.4

9.4

9.6

9.4

10

Intrinsic spatial resolutione (FWHM in mm) – 99mTc

3.8

3.8

3.7

3.7

3.8

3.7

3.9

System spatial resolution @ 10 cme (FWHM in mm) – 99mTc with LEHRf collimation

7.4

7.4

7.2

7.7

8.7

7.6

7.4

System sensitivitye (cps/MBq) – 99mTc with LEHRf collimation

72

72

74

86

91

82

91

SPECT spatial resolutione (FWHM in mm) – 99mTc with LEHRf collimation

9.9

9.9

9.8

Not available

11

Not available

Not available

aSPECT scanners are currently marketed as “SPECT-only” scanners or as multimodality devices with either “diagnostic-grade” or “nondiagnostic-grade” CT scanners; the latter are used only for attenuation correction and anatomic orientation

bExtracted from the marketing literature of the respective manufacturers

cGamma cameras in which the angle between the two detectors may be varied from 90 to 180° are general-purpose (including cardiac) imaging devices; those in which the two detectors are fixed at a 90° angle with respect to one another are dedicated cardiac imaging devices

dCFOV IU: central field-of-view integral uniformity, ((maximum counts per pixel – minimum counts per pixel)/(maximum counts per pixel – minimum counts per pixel) 100% for a high-count “flood” image

eGamma camera and SPECT scanner performance parameters and their measurement are described in references [41, 42, 43]

f LEHR low-energy high-resolution

Preclinical. A number of preclinical SPECT devices are commercially available. The various preclinical SPECT scanners are rather similar in design, generally utilizing multiple NaI(Tl) scintillation detectors fitted with multi-aperture pinhole collimators. The superior spatial resolution of preclinical versus clinical SPECT scanners, ~1 versus ~10 mm, is due to the magnification effect afforded by the pinhole collimation (Figure 2b). Of course, this is achieved at the cost of lower sensitivity, though the use of multi-aperture collimators and multiple (up to four) detectors as least partially mitigates the reduction in sensitivity. Preclinical devices are currently marketed as “SPECT-only” scanners, as multimodality devices with integrated CT scanners (typically conebeam devices with flat-panel detectors) or even as trimodality PET-SPECT/CT systems. The design and performance parameters of commercially available preclinical SPECT scanners are summarized in Table 3.
Table 3

Design and performance parameters of commercially available preclinical SPECT scannersa,b

 

Mediso NanoSPECT™

Carestream Albira™

Gamma Medica Triumph™

MI Labs Vector™c

Siemens Inveon™

Detectors

NaI(Tl)

NaI(Tl)

CZT

NaI(Tl)

NaI(Tl)

Number of detectors

4

2

Up to 4

Not available

Up to 4

Parallel-hole collimation?

Yes

No

Yes

No

Yes

Pinhole collimation?

Yes

Yes

Yes

Yes (Tungsten)

Yes

Number of apertures per pinhole collimator

9

Up to 5

“Multiple”

5 rings of 15 apertures each

Up to 5

SPECT spatial resolution (FWHM in mm) – 99mTc with pinhole collimation

1.2

0.8

0.4

0.35

1.0–1.5

Sensitivity (cps/MBq 99mTc) – with pinhole collimation

1,640

1,000

Up to 6,500

Up to 1,600

400–600

aPreclinical scanners are currently marketed as “SPECT-only” scanners or as multimodality devices with integrated CT and/or PET scanners. As noted, the Carestream Albira, Gamma Medica Triumph LabPET Solo, and Siemens Inveon are available as trimodality PET-SPECT/CT systems

bExtracted from the marketing literature of the respective manufacturers

cThe MI Labs Vector also functions as a PET scanner by collimated single-photon (i.e., noncoincidence) detection of the 511-keV annihilation gamma rays accompanying positron emission

Concluding Remarks

Although the spatial resolution of SPECT (~1 and ~10 mm for preclinical and clinical devices, respectively) is excellent by historical standards for this modality, it remains relatively coarse compared to that of such high-resolution “anatomic” imaging modalities as CT and MRI (0.1–1 mm). Nonetheless, the distinctive and important advantages of radionuclide-based imaging – high detection sensitivity, “image ability” of non-perturbing doses of radiotracers, quantitative accuracy, and a vast array of radiotracers – ensure that SPECT (particularly in combination with CT and, potentially, MRI) will remain an invaluable molecular imaging modality in clinical practice and in clinical and preclinical research.

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Copyright information

© Springer International Publishing Switzerland 2016

Authors and Affiliations

  1. 1.Department of Medical PhysicsNew YorkUSA

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