Transcription
ASNC IMAGING GUIDELINES FOR NUCLEARCARDIOLOGY PROCEDURESSingle photon-emission computed tomographyThomas A. Holly, MD,a Brian G. Abbott, MD,b Mouaz Al-Mallah, MD,cDennis A. Calnon, MD,d Mylan C. Cohen, MD, MPH,e Frank P. DiFilippo, PhD,fEdward P. Ficaro, PhD,g Michael R. Freeman, MD,h Robert C. Hendel, MD,iDiwakar Jain, MD,j Scott M. Leonard, MS, CNMT, RT(N),a Kenneth J. Nichols,PhD,k Donna M. Polk, MD, MPH,l and Prem Soman, MD, PhDm1. INTRODUCTIONThe current document is an update of an earlierversion of single photon emission tomography (SPECT)guidelines that was developed by the American Societyof Nuclear Cardiology. Although that document wasonly published a few years ago, there have been significant advances in camera technology, imagingprotocols, and reconstruction algorithms that promptedthe need for a revised document. This publication isdesigned to provide imaging guidelines for physiciansand technologists who are qualified to practice nuclearcardiology. While the information supplied in this document has been carefully reviewed by experts in thefield, the document should not be considered medicaladvice or a professional service. We are cognizant thatSPECT technology is evolving rapidly and that theserecommendations may need further revision in the nearfuture. Hence, the imaging guidelines described in thispublication should not be used in clinical studies untilthey have been reviewed and approved by qualifiedphysicians and technologists from their own particularinstitutions.From Northwestern University,a Chicago, IL; Warren Alpert MedicalSchool of Brown University,b Providence, RI; Henry Ford Hospital,cDetroit, MI; MidOhio Cardiology & Vascular Consultants,dColumbus, OH; Maine Cardiology Associates,e South Portland, ME;Cleveland Clinic,f Cleveland, OH; University of Michigan,g AnnArbor, MI; St. Michael’s Hospital,h University of Toronto, Toronto,ON; University of Miami Miller School of Medicine,i Miami, FL;Drexel University College of Medicine,j Newton Square, PA; LongIsland Jewish Medical Center,k New Hyde Park, NY; HartfordHospital,l Hartford, CT; UPMC Cardiovascular Institute,m Pittsburgh, PA.Unless reaffirmed, retired, or amended by express action of the Boardof Directors of the American Society of Nuclear Cardiology, thisImaging Guideline shall expire as of May 2015.Reprint requests: Thomas A. Holly, MD, Northwestern University,Chicago, IL.J Nucl Cardiol1071-3581/ 34.00Copyright Ó 2010 by the American Society of Nuclear Cardiology.doi:10.1007/s12350-010-9246-y2. INSTRUMENTATION QUALITY ASSURANCEAND PERFORMANCEThe proper choice of equipment to acquire clinicaldata and a well-designed quality assurance (QA) program are both essential requirements for optimizingdiagnostic accuracy and ensuring consistent, high-quality imaging. The following guidelines are intended toprovide an appropriate means of assessing equipmentfunction in conjunction with nuclear cardiology imaging. Because the optimal manner in which to performspecific tests varies considerably between models ofimaging equipment, this document is not intended toreplace manufacturers’ recommendations.For decades, the design of SPECT cameras hasremained essentially unchanged, consisting of one ormore large-area scintillation or Anger cameras [singleNaI(Tl) crystal with large photomultiplier tubes (PMTs)and parallel-hole collimator]. The QA and performanceof these cameras are well understood, and detailedguidelines are presented in this document. However, inrecent years, novel and dedicated cardiac SPECT cameras have emerged with significantly different detector,collimator, and system designs. Because this technologyis rapidly evolving, detailed guidelines for QA testing donot yet exist for these new designs. Until detailedguidelines become available, the user is responsible forassessing and following vendor recommendations, andmust devise acceptable QA standards specific to thescanner design. To assist in this endeavor, the basicdesign concepts and their implications on performanceand QA are summarized in this document.2.1. DetectorsDetectors are the heart of a SPECT system and areresponsible for collecting the high-energy photonsemitted by the patient, estimating the photon energy andlocation of interaction, and generating count data forsubsequent image reconstruction. The ability to performthese duties depends on their design, materials, andelectronics. Energy resolution, camera sensitivity, and
Holly et alSingle photon-emission computed tomographyspatial resolution are the primary variables that dictatethe performance of a SPECT detector.2.1.1. Scintillation camera (anger camera). The majority of SPECT systems are based onAnger camera technology where one or more camerasrotate about the body of the patient. Anger camerasconsist of a single crystal, which absorbs incidentgamma photons and scintillates or emits light inresponse, and in back of which are banks of photomultiplier tubes and electronics to compute gamma rayenergy and the location of scintillation within the crystal. Anger camera NaI(Tl) crystals are typically 1/4 to3/8 in. thick, although they may be as thick as 5/8 in.The thicker the crystal, the greater the sensitivity of theAnger camera, because of the increased probability thata gamma ray passing through the crystal will interact.However, the thicker the crystal, the greater the spreadof the emitted light photons produced from the scintillation, and the less precise the computation of gammaray interaction location resulting in poorer intrinsicresolution of the camera.2.1.2. Scintillation crystals (pixelated). Anarray of scintillation crystals is an alternative to thesingle-crystal Anger camera design. A large number ofsmall crystals (e.g., 6 mm CsI(Tl) cubes) are coated withreflective material and packed into an array. Anadvantage of this pixelated design is that the scintillationlight is much more focused than in an Anger camera andcan be detected by a photodiode array instead of conventional PMTs, thereby making the detector muchmore compact. A possible concern for pixelated detectors is that their less efficient light collection maydegrade energy resolution. Pixelated detectors arecapable of very high counting rates due to their isolatedlight pulses and have been used in first-pass cardiacscintigraphy.2.1.3. Semiconductor/solid-state detectors. Recently, alternative SPECT systems have beenintroduced, some based on small solid-state detectormodules. In solid-state detectors, gamma rays areabsorbed into the semiconductor material which directlygenerates electron-hole pairs which are pulled to the endplates through an applied electric field. The collectedcharge from the electron-hole pair is used to determinethe location and energy of the gamma ray. One suchsolid-state detector is made of cadmium zinc telluride(CZT), and SPECT devices utilizing these detectorshave been reported to provide improved count sensitivity, superior energy resolution, and finer spatialresolution.1,2 The magnitude of improvement that hasbeen reported consists of simultaneously acquiring threeto ten times more counts,3 with over two times betterspatial resolution than the Anger camera.2 The smallsize of solid-state modules has made a number ofJournal of Nuclear Cardiologyinnovative detector designs possible. Some devices nowhave static arrangements of CZT crystals, and the onlymoving part is a collimator array. Other arrangements ofCZT detectors have each detector module equipped withits own pinhole collimator, for which there are nomoving parts other than a mechanism to move thedetector as close as possible to the patient.2.2. Energy ResolutionImage quality is affected by energy resolution.Ideally, images would be comprised of only primaryphotons emitted by the decay of a radionuclide, not thesecondary scattered photons. The better the energy resolution of an imaging device, the more successful thediscrimination of primary from scattered photons, andthus, the better the image contrast and the more accuratequantitation of the amount and distribution of theradionuclide in vivo. In the context of myocardial perfusion imaging, systems with better energy resolutionhave the ability to discriminate areas of hypoperfusionfrom those of neighboring normal perfusion throughenhanced image contrast.A symmetric energy window of 20% centeredaround the 140 keV energy peak of the emitted photonsis standard for Tc-99m. With the improved energy resolution of many modern cameras, a 15% window can beused with little loss of primary gamma rays, with someimprovement in image contrast. The lower energy andgreater width of the Tl-201 photopeak require a widerenergy window. An energy window setting of 30% isappropriate for the 70 keV peak of Tl-201, and a 15%energy window is appropriate for the 167-keV peak. Theabsolute energy calibration can be unreliable at the lowenergy of Tl-201, falling at slightly different energypositions in the spectrum of different cameras. Thus,energy peak and window settings should be establishedfor each individual camera based on the energy spectrumdisplay.4 For most modem Anger cameras, the energyresolution is expected to be 9-10% for the 140-keVphotons from Tc-99m, and 15-17% for the 72-keVphotons from Tl-201.5 For CZT detectors, an energyresolution of 5 to 6% has been reported for Tc-99m.1,22.3. Spatial ResolutionSpatial resolution quantifies the size of the smallestobject that can be resolved reliably, and is oftenexpressed as the full-width at half-maximum (FWHM)of a point spread function. For projection data acquiredwith an Anger camera, the total resolution depends onthe intrinsic resolution and the collimator resolution.The intrinsic resolution is typically 3.5-4.0 mm forTc-99m and is a function of the crystal and the position
Journal of Nuclear CardiologyHolly et alSingle photon-emission computed tomographylogic circuitry used to compute the location of thedetected photon. The collimator has the largest impacton the spatial resolution. For a high resolution collimator, the spatial resolution is typically *10 mm for aTc-99m point source placed at 10 cm from the collimator face. The resolution varies for differentcollimators depending the length and diameter of thecollimator holes. Collimators with smaller and longerholes will have better spatial resolution but will havelower sensitivity.For all SPECT imaging devices, system spatialresolution is also influenced by the image reconstructionprocess. The most commonly used reconstruction algorithm is filtered backprojection which involves areconstruction kernel (e.g., a ramp function in the spatialfrequency domain) and a postfilter which is usuallydesigned to suppress image noise. In the spatial frequency domain representation, a noise smoothing filterhas a high frequency roll-off whose effect is to suppresshigh-frequency image noise especially for count-poorstudies at the expense of sacrificing spatial resolution. Aroll-off at higher spatial frequency results in betterspatial resolution, but often entailed images of a highernoise or lower signal-to-noise ratio.6 Recently, moresophisticated algorithms have been employed for imagereconstructions that model the physics and characteristics of the imaging system involved in the imageformation process. These include the statistical maximum-likelihood expectation-maximization (ML-EM)algorithms,7 and the ordered-subset expectation-maximization (OS-EM) algorithms.8For new detector designs based on multiple smallsolid-state detectors, tomogram formation is implemented through the use of ML-EM or OS-EMalgorithms that not only generate the tomogram, but alsomake it possible to correct for estimated amounts ofscattered radiation, using information in multiple energywindows.9,10 Furthermore, MLEM and OSEM algorithms enable incorporation of detailed knowledge aboutthe manner in which detector response varies with thedistance away from each detector.11 These algorithmsalso enable attenuation compensation using attenuationmaps obtained from scanning line sources or CT imagesin modem SPECT/CT system.2.4. Detector SensitivityThe image quality of scintigrams is largely determined by the signal-to-noise ratio. The greater thesensitivity of the device, the more counts are available,and the higher the signal-to-noise ratio. The greater thenumber of counts that can be acquired, the more certainone can be as to the correct three-dimensional (3D)distribution of count concentrations within the body.The amount of radionuclide that can be injected perstudy is limited by the amount of radiation to which thepatient can be exposed during a diagnostic procedure.Thus, given that there is an upper limit to the amount ofan isotope that can be injected routinely, it is desirable tohave the most sensitive device available to form theclinical images.2.5. Count Rate LimitationsAt very high photon detection rates, the cameraelectronics can have difficulty analyzing every photon.The missed counts, often referred as dead time effects,can affect the estimated distribution of the radiotracer inthe body. For current cardiac radiotracers, protocols andimaging systems, dead time effects are not an issue.2.6. CollimationSPECT image reconstruction requires that theincident direction of each acquired count be known. Anexternal collimator is used to do so, by absorbing photons outside a range of incident angles as specified bythe collimator design. By limiting the number ofdetected photons to a specific direction or small range ofdirections, the collimator allows the detection of only avery small fraction of the photons emitted by the patient(see Table 1). However, since a range of incident anglesis allowed to pass through the collimator, there is a lossof resolution. An inherent trade-off in the design ofTable 1. Typical performance parameters for low-energy (\150 keV) collimatorsCollimator typeUltra-high resolutionHigh resolutionAll/general purposeHigh sensitivityResolution (FWHM at 10 410-410-410-4Relative efficiency.3.61*2.1Note: A relative efficiency of 1 corresponds approximately to a collimator efficiency of 1.7 9 10-4, whereas efficiency is definedas the fraction of gamma rays and x-rays passing through the collimator per gamma ray and x-ray emitted by the source.
Holly et alSingle photon-emission computed tomographySPECT collimators is resolution vs sensitivity. It ispossible to improve the spatial resolution of a collimatorby further restricting the range of incident angles, butthis is at the expense of reducing sensitivity. Angercameras require collimators to localize the site withinthe patient from which the gamma ray photons areemitted.Because of this resolution vs sensitivity trade-off,the collimator is perhaps the most significant componentof the SPECT camera affecting image quality. MostSPECT cameras for nuclear cardiology utilize conventional parallel-hole collimation. However, it has longbeen recognized that significant improvements inSPECT performance are possible with alternative collimator designs. Coupled with advances in detector andcomputing technology in recent years, new collimatordesigns have become practical for nuclear cardiology.2.6.1. Parallel-hole collimation. The structure of a parallel-hole collimator is similar to that of ahoneycomb, consisting of a large number of narrowchannels separated by thin septa. The range of incidentangles passing through the collimator depends on thechannel width and length. The septa must be thickenough to absorb photons of the desired energy in orderthe accept photons that are incident on the collimatorwithin a range of directions. Collimators are often categorized as low-, medium-, or high-energy depending onthe photon energy for which the collimator is designed.Various designs of low-energy collimators are classifiedas high-resolution, all-purpose, high-sensitivity, etc.,depending on the collimator’s angle of incidence.For Anger cameras, parallel-hole collimation isstandard. The low-energy, high-resolution collimator isusually best for Tc-99m, although some ‘‘all-purpose’’collimators give excellent results. Imaging with Tl-201is usually best with the low-energy, medium-resolution(all-purpose) collimators because count statisticsbecome limiting when using high-resolution collimators.The difference in medium- and high-resolution collimators is usually that the collimator depth (i.e., length ofthe collimator hole) is greater in high-resolution collimators. Medium- and high-resolution collimators havesimilar near-field resolution, but high-resolution collimators maintain good resolution at a greater distancefrom the collimator face. The difference is moreimportant in SPECT imaging, where the distance frompatient to collimator is greater.5The selection of which collimator to use is important to resultant image quality. A confounding aspect ofthis selection is that collimators with the same name(e.g., ‘‘general purpose’’) vary in performance from onemanufacturer to another. Table 1 gives approximatevalues for collimator specifications. The user shouldrefer to specific imaging protocols for appropriateJournal of Nuclear Cardiologycollimator selection. It is important to perform periodicassessment of collimator integrity, as failure to detectand correct for localized reduced sensitivity can generate uniformity-related artifacts, including reconstructionartifacts.12,13The trend is to trade off some of the additionalcounts that would be obtained with multiple detectorsequipped with general-purpose collimators for higherresolution imaging by using higher-resolution collimators. For newer reconstruction algorithms that model thecollimator blur, some degree of resolution recovery isfeasible. With such algorithms, the all-purpose collimator may offer improved image quality.14 Dependingon the scanner configuration, reconstruction algorithm,radionuclide dose, and patient population, the optimalcollimator choice may vary between laboratories.2.6.2. Sweeping parallel-hole collimation. A variant of the standard parallel-hole collimatoris a sweeping configuration, where an array of smalldetectors is arranged around the patient instead of one tothree large-area Anger cameras. Each small detector isconfigured with a parallel-hole collimator, whichrestricts its field-of-view (FOV) over a small volume. Toview the entire patient, each detector pivots about itsown axis, sweeping its FOV like a searchlight across theentire imaging volume. With a sufficient number ofthese small pivoting detectors, a sufficient degree ofangular sampling is achieved to reconstruct high-qualityimages. The sweeping FOV of the pivoting detectorsprovides the flexibility to control the amount of time thatis spent imaging regions of the imaging FOV. Byspending more time imaging the myocardium and lesstime imaging the rest of the chest, data collection ismore efficient and could lead to reduced scan time ordose. Although the concept of sweeping parallel-holecollimation is not new,15 recent advances in detectordesign have led to its practical implementation.16 Withsolid-state CZT detectors, a compact stationary scannerdesign is feasible.2.6.3. Converging collimation (fixed andvariable fan and conebeam). With large-areaAnger cameras and parallel-hole collimation, only asmall fraction of the detector area is utilized in imagingthe myocardium. Converging collimators allow more ofthe crystal area to be used in imaging the heart, magnifying the image, and increasing sensitivity.5Converging collimators with a short fixed focal distancehave the potential for truncating portions of the heartand/or chest, especially in large patients. This truncationmay generate artifacts. Currently, there is a limitedclinical use of fixed-focus fan beam and cone beamcollimators in nuclear cardiology.Another converging collimator design that hasreached a commercial stage is the variable-focus cone-
Journal of Nuclear Cardiologybeam collimator.17 The center of this collimator focuseson the heart to increase the myocardial counts acquired.The remainder of the collimator has longer focal lengthto improve data sampling over the entire chest andthereby reduces truncation artifact.2.6.4. Multi-pinhole collimation. Collimation requires the mapping of a point on the detector to anangle of incidence. Conventional parallel-hole andconverging hole collimators use an array of narrowchannels of radiation-absorbing material to specify theangle of incidence. A fundamentally different approachis pinhole collimation where a small pinhole aperturesurrounded by radiation-absorbing material allowsphotons to pass through. The line connecting the point ofemission of the gamma photon and the pinhole aperturespecifies the direction of photon incidence on thedetector.The sensitivity and resolution characteristics ofpinhole collimators differ markedly from those of parallel-hole collimation. The sensitivity of a parallel-holecollimator is almost constant as a function of distancefrom the collimator. However, the sensitivity of a pinhole collimator depends on the inverse square of thedistance from the pinhole aperture and becomes quitelarge at small distances. It also increases as the square ofthe pinhole diameter with concurrent loss in spatialresolution. The magnification factor of a pinhole isdetermined by the ratio of the pinhole-to-detector distance relative to the pinhole-to-source distance. Withincreased magnification factors, the detector intrinsicresolution has decreased effect on the total system spatial resolution which approaches that of the pinholecollimator. In summary, a pinhole collimator is mosteffective for imaging a small object placed close to thepinhole aperture, and is clinically used most often forthyroid imaging. Pinhole collimators have receivedmuch attention recently for small animal SPECT imaging. Multi-pinhole collimators are effective in increasingthe sensitivity and angular sampling, provided that the ofoverlapping pinhole projections are avoided throughinter-pinhole shielding, or in cases with small amount ofoverlap, special image reconstruction algorithms areused to minimize the effect on image quality.Multi-pinhole collimation was used in the earlydays of nuclear cardiology to provide multiple angularviews (and in some cases, limited angle tomography) ofthe myocardium.18 With advances in iterative imagereconstruction, multi-pinhole collimators have made acome-back in nuclear cardiology. Fully tomographicmyocardial SPECT imaging has been demonstratedusing multi-pinhole collimators with both Angercameras19 and small area CZT detectors.20 The characteristics of pinhole collimation offer the potential ofboth improved spatial resolution and sensitivityHolly et alSingle photon-emission computed tomographycompared to conventional pinhole collimation withAnger cameras.2.7. System DesignThe sections above have discussed the various typesof detectors and collimators used in nuclear cardiology.The SPECT imaging system is built around these components to optimize the imaging protocols, patienthandling, and flexibility of the scanner and the overallimage quality. A complete understanding of the systemdesign is necessary to establish a proper QA program.2.7.1. Multi-purpose SPECT. A common system design for cardiac SPECT is a multi-purposeSPECT scanner with one or more large-area Angercameras. The detectors are attached to a gantry whichrotates the Anger cameras and adjusts their distancerelative to the center-of-rotation. Projection images areacquired as the gantry rotates the detectors in a continuous or step-and-shoot mode. An advantage of thisSPECT system design is that it is compatible withgeneral nuclear medicine scans (planar and SPECT) aswell as cardiac SPECT.Single-head cameras were used widely for cardiacimaging initially. Adding more detectors is beneficial,since doubling the number of detectors doubles theacquired counts, if all other variables remain fixed. Forcardiac SPECT studies in which a 180-degree orbit(right anterior oblique [RAO] to left posterior oblique[LPO] or LPO to RAO) is recommended, the preferredconfiguration is to have two detectors separated by 90 as they rotate around the heart. A potential concern withthe 90 configuration is whether there is truncation ofthe SPECT projections in the region between the twodetectors. Many dual-detector systems also offer a moreacute configuration (e.g., 80 ) and a longer scan range(e.g., 100 ) to avoid possible truncation near the heart.The various mechanical configurations and detectormotions of multi-purpose SPECT systems have thepotential for misalignment errors and must be includedin the laboratory’s QA program (e.g., center-of-rotation[COR] correction, multi-detector registration).2.7.2. Dedicated cardiac. As nuclear cardiology grew in volume, the market expanded for SPECTscanners designed solely for cardiac imaging instead offor all nuclear medicine procedures. Many scannersemerged with smaller-area Anger camera detectors in afixed 90 configuration, with significantly reduced cost,weight, and space requirements. However, care in patientpositioning is more critical when using these dedicatedcardiac SPECT systems than the multi-purpose SPECTsystems with large-area detectors, to avoid truncating theheart and the associated artifacts. Otherwise, QA is similar to the multi-purpose SPECT systems.
Holly et alSingle photon-emission computed tomographyThe growth of dedicated cardiac SPECT systemsalso led to the adoption of novel detector and collimation technologies as described above. Because of thewide variety of designs and their novelty, their QArequirements may differ from the standard Anger camera with parallel-hole collimation. Thus a detailedquality assurance summary for each design is not presented in this document, though areas where the QA orperformance may differ are noted. Each laboratoryshould understand these differences and consult with thevendor’s user manuals to develop its own program untilsuch standards and guidelines are developed.2.7.3. Patient configuration. Because of thevariety in SPECT scanner designs, the patient handlingsystem can differ markedly between SPECT scanners.Although this does not impact camera quality assurancedirectly, the differences in patient configuration and thepotential effect on study interpretation should beunderstood.The conventional patient handling system holds thepatient horizontally on a level table as the detectorsrotate around the patient. The patient typically is imagedin a supine position, although some laboratories routinely image in the prone position.21 Some dedicatedcardiac SPECT systems instead hold the patient in areclined position or upright position. The differentdirection of gravity relative to the patient may affect theexpected patterns of attenuation and the degree andnature of patient motion.Some new dedicated cardiac systems have detectorsand/or collimators that move internally to the scannerand have no visible moving components. Other systemsmay not require detector or collimator motion at all andare entirely stationary. However, some dedicated cardiacSPECT cameras differ in that the detectors are stationaryand that the patient is rotated in the upright position. Therotating chair of these cameras must be included in acenter-of-rotation assessment of the QA program.2.7.4. Data sampling. A range of angular projection images is needed to reconstruct tomographicimages. The way that SPECT systems acquire the multiple projection images differs significantly. Forconventional parallel-hole collimators and Anger cameras or similar detectors, recommended guidelines existfor the minimum number of angular samples.For other designs, the topic of data sampling maynot be so straightforward. Truncation with small FOVdetectors may cause artifacts particularly at the edge ofthe FOV. Cone-beam and pinhole geometries have lessdata completeness at distance further away from thecentral transaxial image slice and show increased imageartifacts. Some focused collimator designs may havesufficient sampling in the cardiac region though notover the rest of the torso. Proper patient positioning isJournal of Nuclear Cardiologyimportant to ensure that the heart is in the fully sampledregion.Although each vendor has recommended cardiacimaging protocols for their cameras, the user shouldvalidate these protocols for their own use with QAphantoms if similar validation work has not yet beenpublished. Studies with a full-scale anthropomorphicphantom would be appropriate for this validation tosimulate imaging under actual conditions.2.7.5. SPECT systems with photon sourcesfor attenuation correction (AC). Currently, thereare two types of transmission tomographic imagingsystems for acquiring patient-specific attenuation mapsthat can be used to correct SPECT images for photonattenuation. The first type, referred to in these guidelinesas transmission computed tomography (TCT), uses asealed radioactive source (e.g., Gd-153) with the standard collimated scintillation detectors used for SPECTimaging. The second type of transmission imaging system uses an x-ray tube in conjunction with a computedtomography (CT) detector. The primary differencesbetween these classes of transmission imaging systemsinclude the type of radiation used and the photonemission rate that dictates the quality-control (QC)protocols that are required. The radiation from radioactive source used in TCT is monoenergetic gamma-raysand those used in x-ray CT is polychromatic x-ray froman x-ray tube. Also, the flux of x-ray photons from atypical x-ray tube is much higher than a conventionalsealed source used in TCT, CT images can be acquiredon the order of seconds to a few minutes depending onthe x-ray tube strength.The number of photons from a sealed radioactivesource used to form the transmission image in TCT is onthe same magnitude as that from the emission source butmuch less than that from a t
ASNC IMAGING GUIDELINES FOR NUCLEAR CARDIOLOGY PROCEDURES Single photon-emission computed tomography Thomas A. Holly, MD,a Brian G. Abbott, MD,b . Unless reaffirmed, retired, or amended by express action of the Board of Directors of the American Society of Nuclear Cardiology, this Imaging Guideline shall expire as of May 2015. Reprint .
