THE UNIVERSITY OF NEW MEXICOHEALTH SCIENCES CENTER

COLLEGE OF PHARMACYALBUQUERQUE, NEW MEXICO

Correspondence Continuing Education Courses For Nuclear Pharmacists and

Nuclear Medicine Professionals

VOLUME IX, NUMBER 3

Recent Advances in Nuclear MedicineImaging Technology

By:

David S. Binns, ANMT Peter MacCallum Cancer Institute

East Melbourne, VictoriaAUSTRALIA

The University of New Mexico Health Sciences Center College of Pharmacy is approved by theAmerican Council on Pharmaceutical Education as a provider of continuing pharmaceuticaleducation. Program No. 039-000-01-002-H04. 4.0 Contact Hours or .40 CEUs.

Approved for continuing Nuclear Pharmacy education by the Florida Pharmacy Association,Florida Program No. PSN2003-001.

PLEASE NOTE:This lesson is worth point 4 CEUs (4 credit hours). You must complete &return two answer sheets in order to obtain full credit. Questions 1-30 mustbe completed on first answer sheet, questions 31-40 must be completedon the 2nd answer sheet as questions 1-10.

Recent Advances in Nuclear MedicineImaging Technology

By:

David Binns, ANMT

Coordinating Editor and Director of Pharmacy Continuing EducationWilliam B. Hladik III, MS, RPh

College of PharmacyUniversity of New Mexico Health Sciences Center

Managing EditorJulliana Newman, ELS

Wellman Publishing, Inc.Albuquerque, New Mexico

Editorial BoardGeorge H. Hinkle, MS, RPh, BCNP

Jeffrey P. Norenberg, MS, RPh, BCNPLaura L. Boles Ponto, PhD, RPh

Timothy M. Quinton, PharmD, MS, RPh, BCNP

Guest ReviewerSteve Meikle, PhD

PET and Nuclear Medicine DepartmentRoyal Prince Alfred Hospital

Camperdown, NSWAUSTRALIA

While the advice and information in this publication are believed to be true and accurate at press time, the author(s),editors, or the publisher cannot accept any legal responsibility for any errors or omissions that may be made. The

publisher makes no warranty, express or implied, with respect to the material contained herein.

Copyright 2001University of New Mexico Health Sciences Center

Pharmacy Continuing EducationAlbuquerque, New Mexico

RECENT ADVANCES IN NUCLEAR MEDICINEIMAGING TECHNOLOGY

STATEMENT OF OBJECTIVES

Upon completion of this material, the reader should be able to:

1. Understand the physics and selection criteria of scintillation detectors.

2. Understand the critical design issues in the development of gamma and PET cameras.

3. Discuss the relative merits of different imaging systems and how these have driven clinical

acceptance.

4. Explain the technology of various gamma and PET camera innovations that have helped

improve imaging performance.

5. Understand future developments in gamma and PET camera technology.

COURSE OUTLINE

I. INTRODUCTION

II. BASIC PRINCIPLES OF DETECTION

A. Before Detection – Photons ThroughMatter

B. Detector Geometry

C. Scintillation Detectors

D. Detector Properties

1. Background

2. Why NaI(T1) for SPECT?

3. Why BGO for PET?

4. Semi-Conductor Detectors

5. Tomographic Reconstruction

a. Filtered Back Projection (FBP)

b. Iterative Methods

c. Accelerated Methods-OrderedSubsets

III. INSTRUMENTATION

A. Gamma Camera Design

1. Introduction

2. Gantry

3. Collimation

4. Multi-Headed Cameras

5. Attenuation Correction

B. PET Camera Design

1. Introduction

2. Detector Selection, Design and Properties

3. 2D vs. 3D Acquisition

4. Attenuation and Scatter

5. Partial Ring BGO PET Cameras

6. Dedicated NaI(T1) PET Scanners

7. Time-of-Flight PET Detectors

8. Multi-Modality Cameras – Marriage of CT and PET

9. Future Developments in PETScanners

C. Hybrid SPECT/PET Gamma Camera Systems

1. Introduction

2. Design Considerations

3. Future Developments in Hybrid Scanners

IV. CONCLUSION AND THE FUTURE

V. REFERENCES

VI. QUESTIONS

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INTRODUCTIONThe development of radiopharma-

ceuticals and the evolution of nuclearmedicine instrumentation often go hand inhand. There are many examples throughoutthe history of nuclear medicine wherediscoveries in one technology have drivendiscoveries in the other. Much of mainstreamradiopharmaceutical development hasfocused on labeling 99m-Tc to variouscompounds. 99mTc wasn’t selected becauseof its chemical properties; it was selected inpart because its physical or nuclearproperties elegantly matched the detectorproperties of the Anger gamma camera. Thedevelopment of 99mTcHMPAO in the mid-1980’s saw specialized fan beam collimatorsenter the market capable of high qualityneurological SPECT. Fluorine-18 bonescanning, once in-vogue when rectilinearscanners were in general use, is know beingrevisited with the increasing number of PETscanners.

Nuclear medicine instrumentation isfundamentally associated with physics and tohave an appreciation of this topic requires atleast a minimum knowledge of thesefundamentals. The physics presented in thisdocument is by no means replete. It merelyrepresents the basics of what is required toform a good understanding of nuclearmedicine instrumentation and imaging, nowand in the near future.

In the 1990s nuclear medicinecontinued to play an expanding role inhealthcare. The growth rate in nuclearmedicine services in most western countriesis equal to or greater than most otherdiagnostic modalities and it is increasinglybeing recognized in many settings as amodality of choice in the cost effectiveclinical management of patients.

Despite this growth in technology,few recent important innovations in nuclearmedicine instrumentation have been adoptedby manufacturers, utilized by nuclear

medicine service providers, or applied to thepatient. The vast majority of nuclearmedicine scans are still performed by theAnger gamma camera. The intrinsicperformance specifications of the gammacamera have remained virtually unchangedover the last decade. The refining of theseinstruments continues but not inrevolutionary terms.

There has been a remarkable but notsurprising improvement in the performance ofassociated computing power. Processing timeshave been reduced, but this has been matchedby increasingly complex reconstruction anddisplay packages. Digital networktechnology and standardization of imageformats has allowed connectivity betweenimaging devices, their associated work-stations, and hospital wide area networks.

Digitization of the detection systemhas made the gamma camera more easilycalibrated and stable. Except for fan-beamcollimation in neurological SPECT, therehave been few new collimator designs thathave reached production.

Contemporary gamma camerasystems are much improved than those of adecade ago. Gantry designs are moreergonomic for patient and operator comfort,safety, and throughput. Gantries andscanning beds are now highly complex andspecially engineered to perform SPECT andother studies with a high degree of precision.Body contouring (where the gantry detectsthe boundary of the patient’s body) allowsfast set-up times and reduced detector topatient distances.

Measured attenuation correctionSPECT studies show promise, especially innuclear cardiology, but still needs to gainwidespread clinical acceptance. Theintroduction of 99mTechnetium myocardialagents and gated myocardial perfusionSPECT has reduced the problems associatedwith attenuation artifacts that reduce201-Thallium specificity.

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The inter-dependence of resolutionand sensitivity is a major factor in gammacamera system performances. Invariably acompromise is made between the two. Onesimple, albeit expensive, way of making aquantum leap in system sensitivity withoutcost to resolution is to employ multipledetectors.1 Increasingly dual and tripleheaded cameras are being employed in bothSPECT and planar imaging. The increasedsensitivity can be utilized to reduce statisticalnoise, reduce acquisition times, improveresolution, or any of these in combination.

Arguably, the most revolutionaryinnovation of the gamma camera in the 1990swas the development of the coincidencegamma camera allowing PET to beperformed part-time. This involvedsignificant changes to the detectorelectronics and logic but meant that a cameracould be operated in PET or SPECT mode.Thus, the ability to perform PET wasavailable with the purchase of an optioncosting in the hundreds of thousands ofdollars rather than the millions needed fordedicated PET. It also meant the cost couldbe amortized by performing SPECT whenthere was limited availability of PET tracers.

Much debate has ensued about thequality of gamma camera PET versusdedicated PET. There is much clinicalevaluation of gamma camera PET takingplace around the world. This is a very youngtechnology with much research anddevelopment of instrumentation taking place,particularly regarding the use of differentscintillation detectors more suited to PET.

PET has had a history in medicine foralmost as long as single photon nuclearmedicine imaging.2 During the 1980’s andearly 1990s it saw slow proliferation,primarily in neurological and myocardialresearch, but found few viable clinicalapplications. The major restraint in manyinstances was the high cost of PET scanning,which diluted its cost effectiveness. With the

expansion of PET in oncology many centersdeveloped more clinically based protocolsand higher patient throughput, and a generalchange in focus away from biological,physiological, and pharmacological research.New centers opened which shared cyclotronfacilities and had smaller infrastructures,which rationalized the service and reducedcost in most instances. PET scans stillcomprise a small percentage of total nuclearmedicine services, although this percentageis growing steadily.

From a technology point of view PETstill offers the potential for radicalinnovation. It doesn’t have the restraint thatcollimators place on single photon imaging(Table 1), nor is it as technologically mature.This is reflected at current scientificconferences associated with imagingtechnology, which show an unprecedentednumber of papers covering new technologiesassociated with PET instrumentation. Someof the most important of these new directionsare new scintillation detectors, depth-of-interaction and new reconstructionalgorithms.

The diagnostic power of aninvestigation results fundamentally from itsability to contrast normal and diseased tissue.The resultant changes in radioactivedistribution need to be of a suitablemagnitude that can be differentiated by theimaging device.

This concept is often referred to as“signal to noise ratio” of an image, and it isthe objective of Nuclear Medicine imagingequipment to maximize this value.Interestingly, it is similar to the “target tobackground” parameter frequently used inevaluating the performance ofradiopharmaceuticals. Increasing the “signalto noise ratio” can be achieved by eitherimproving contrast (usually by improvingspatial resolution) or by decreasing noise(through increased sensitivity).1

It is important to appreciate the

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limitations of nuclear medicine imagingdevices when assessing the accuracy of atest. To do this it is necessary to understandthe processes and steps involved in detectingthe signal coming from the patient. Innuclear medicine, this signal is comprised ofphotons emitted as the result of a randomevent — radioactive decay.

BASIC PRINCIPLES OF DETECTION

Before Detection – Photons ThroughMatter

Gamma rays emitted from a decayingnucleus are monoenergetic (some isotopeshave multiple but discrete photon energies).Those escaping from the patient withoutundergoing Compton scatter will emergewithout reduction in energy. As photonstravel through greater distances of bodymatter more will undergo collisions with theouter shell electrons of atoms comprising thismatter. Following this collision the photons

path will change direction and it will proceedwith a reduction of energy – the greater thescatter angle the greater the loss of energy.The equation that describes the relationshipbetween the scatter angle (q) (i.e., The angleof the change in direction followingcollision) and the change in photon energy isdescribed below:

EY2 = EY1 / [ 1 + ( EY1 / 511 ) x ( 1-cosθ ) ] keV

—Where EY1 is the pre-collision and EY2 isthe post-collision photon energies (keV)

Figure 1 shows the effect scatter hason the energies of the photons of severalradionuclides. Energy windows are normallydefined as a percentage of the peak so the y-axis shows the percentage energy loss atvarious scatter angles. Significant scatter canoccur of photons at low energies, yet still beaccepted by the energy window of theimaging device.

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Figure 1. This graph shows the relationship between the angle of scatter of a single photon and itspercentage loss in energy for 201-T1 (70 keV), 99m-Tc (140 keV), 131-I (360 keV) and 18-F (511 keV).Also shown is the 10% loss in energy that relates to a 20% (i.e., ±10%) energy window typicallyused for NaI(T1) detectors. Thallium-201 can undergo scatter of up to 75 degrees and still beaccepted. Note that energy resolution for NaI(T1) is approximately 8% (FWHM), and hence thereis a degree of inaccuracy in the measurement of a photon’s energy. This means that a scatteredphoton that loses greater than 10% of its energy may still be included in the acceptance window.

The ability of an imaging detector toaccurately discriminate the energies ofscattered and unscattered photons is crucialto its performance.

Ultimately, especially in deeperstructures, fewer photons will emergeunscathed. It is desirable to disregard thesescattered photons. This results in reducedsignal from areas where photons must travelthrough a greater depth of tissue to reach thedetector. Indeed many photons especiallythose undergoing multiple Comptonscattering will be stopped by thephotoelectric effect. Hence, there isattenuation of signal from deeper structures.

Detector GeometryOnce photons leave the body their

probability of interacting with the detector isproportional to the geometry that the detectorpresents to the source. In the first instancethis relates to the solid angle of coverage ofthe detector, and secondly any collimationthat may be present.

In single photon imaging the majorinfluencing factor on the sensitivity of adetector is the collimator placed in front of it.

Collimators are a necessary butinefficient approach to obtaining positionalinformation from photons approaching thedetector. Only photons travelingperpendicular (usually) to the collimator willbe allowed to pass through to the detector. Nophotons, even the vast majority ofunscattered ones, are allowed to pass —typically as few as one in 3000 photons(0.033% for a low energy all-purposecollimator), will reach the detector. Systemsensitivity can be improved by designing acollimator that will admit more photons, butthis is usually at the cost of resolution.

In PET scanning (assuming there isno septa) and there is a ring of detectors witha radius of 100cm and 15cm axial field ofview, the geometry of a centrally place pointsource is around 10%. In PET scanning twoof these photons need to be detected incoincidence to determine an event. However,from a geometric point of view, PET has ahuge advantage over single photon imagingwhen it comes to the detection of events.Table 1 shows the relative efficiencies ofSPECT and PET at the various steps of thedetection process.

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Table 1. Approximate Respective Sensitivities at Different Stages of the Detection Process for a Gamma Camera and PET Scanner.

Gamma Camera PET (99m-Tc) (18-F) Decay • Relative gammas emitted 1,000,000 2,000,000 • Body absorption 250,000 1,000,000 Detector Geometry • Photons reaching detector 100 100,000 Detection • Detection of all events 75 50,000 • Energy discrimination 25 25,000 • (PET) coincidence detection 250 Events Recorded 25 250

Note that collimation causes the greatest loss of sensitivity in single photon scanners while in the PET scanner the inefficient step is the appropriate pairing of single photons into coincidences.

Scintillation DetectorsThe detection of a photon emerging

from the patient relies on the photon’sinteraction with the detector, and conversionof its electromagnetic energy into anothermeasurable form. Almost all nuclearmedicine imaging instrumentation achievesthis by utilizing inorganic scintillationcrystals that release light photons followingabsorption of the incident radiation. The light

photons travel through the crystal and aresubsequently detected by photo multipliertubes (PMT), and ultimately converted to anelectrical signal, the amplitude of which isproportional to the energy deposited in thecrystal by the incident photon. The energiesof individual photons need to be assessed tomake sure that only those that have emergedunscattered and deposit all their energy in thecrystal are accepted (Figure 2).

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(i) Gamma camera schema (ii) Spectrum Figure 2. Detection of gamma rays emitted from a patient. Only radiation traveling perpendicular to the crystal can pass through the holes of the collimator and reach the scintillation crystal (c, d, e, f, g, h, i).• a outside detector solid angle. • b attenuated by collimator • c and d two photons irresolvable by detector • e partially absorbed by detector, excluded by energy discrimination • f scattered within crystal but totally absorbed. • g small scatter angle – accepted by energy discrimination • h large scatter angle – rejected by energy discrimination • i ideal photon – no body scatter and photoelectric absorption in detector • j no interaction with crystal • k totally internally absorbed – multiple Compton then photoelectric absorption • l totally internally absorbed – photoelectric absorption (i) The energy properties of detected radiation related to their path direction and interaction

with the patient, collimator, or crystal. Note that although c, d, f, and i are unscattered by the body and fully absorbed by crystal, their recorded energies vary due to relatively poor precision of the gamma camera in determining photon energy.

Detector Properties

BackgroundThe two general types of nuclear

medicine imaging are conventional nuclearmedicine (planar and SPECT) and PET.These modalities have quite differentrequirements for detector properties andconfiguration.

Most conventional (i.e., singlephoton) nuclear medicine imaging focuseson imaging photons under 200 keV withdetector count-rates under 20 k counts persecond (cps). PET, conversely, images 511keV annihilation photons, and due to theabsence of collimation needs to detect thesephotons at system count-rates measured inthe millions of counts per second.3 PET alsorequires good stopping power for 511 keVphotons and good energy resolutionespecially in systems acquiring in 3D mode.

There is also current interest inhybrid SPECT and PET systems that couldperform each modality without compro-mising the other.

A number of physical properties canbe used to evaluate scintillation detectors;these are shown below in Table 2.

Over the last two decades, detector

scintillators for single photon and PETimaging have been almost exclusivelydeveloped for one modality or the other.Gamma cameras have used Sodium Iodidedoped with thallium (NaI(Tl)) and BismuthGerminate (BGO) has been used for PET.Each scintillators type has properties whichmakes them the most suitable for theirrespective imaging, but not optimal for theother. Table 3 shows the properties of anumber of different detector scintillators.

Why NaI(Tl) for SPECT?The selection of NaI(Tl) as the

scintillator of choice for gamma cameras hasstood the test of time. Although it hasrelatively poor stopping power (density andeffective atomic number), it can achievegood efficiency for gamma rays with energyless than 200 keV with a 12 mm crystalthickness.

The attractive property of NaI(Tl) isits good light output per interaction (Table3[f ]). A gamma camera detector scintillatorconsists of a single large crystal coupled to70-120 PMTs. When scintillation occurs theposition of the emitted photon is calculatedby looking at the relative electronic signalproduced by all the PMTs of the detector.

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Table 2. Relating desirable characteristics to physical properties of scintillation detectors.

Desirable Characteristic Detector Property Purpose

High detection efficiency High atomic number and density

Improved sensitivity

Short duration of scintillation Short light decay constant

a. High count-rate capability b. Good timing precision

(hence less randoms in PET)

High detectable light output proportional to energy of absorbed gamma

High primary photon yield Improved energy resolution

Good light transmission of appropriate wavelength to match PMT efficiency

Crystal transparency, Emission Wavelength

Improved sensitivity

Easily manufactured and processed Availability, Strength? Hygroscopic

Reasonable cost; durability

Sufficient photons are released by NaI(Tl) toallow this to be achieved.

When the electronic outputs of all thePMTs are added, the magnitude isproportional to the amount of light produced,which in turn is proportional to the energyabsorbed by the crystal. The good lightproduction properties of NaI give thismeasurement good precision, which resultsin good energy resolution.

In single photon imaging, the count-rates normally encountered by the detector,which is shielded by a collimator, aresufficiently low that the relatively longphoton decay time doesn’t cause much deadtime (the minimum time it takes a detectorsystem to correctly record two events,typically 20 microseconds).

NaI(Tl) crystals remain the optimal

choice for gamma cameras doing singlephoton imaging. As collimator rather thancrystal properties limit system resolution,there is little incentive to develop a bettercrystal. Much of any improvement inintrinsic resolution is negated by the effect ofcollimation.1 Typically, detector (intrinsic)resolution is around 3 mm and collimatorresolution 8mm. The calculation for systemresolution is as follows:

Rsystem = √ (Rdetector2 + Rcollimator2)

– Where Rsystem = resolution of systemRdetector = resolution of detectorRcollimator = resolution of collimator

Rsystem using the above values wouldbe 8.5 mm. An improvement of Rdetector from

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Table 3. Properties of Various Inorganic Scintillators That Have Been Used in Gamma and PET cameras.1,35,34,36

a. Density (g/cm3) 3.67 7.13 7.40 4.54 6.71 4.88

b. Effective Atomic Number 51 75 65 34 59 53

c. Attenuation coefficient @ 511 keV (cm-1)

0.328 0.901 0.820 0.226 0.667 0.437

d. Attenuation coefficient @ 140 keV (cm-1)

2.38 12.2 10 1.30

e. Light Decay Constant (ns) 230 300 40 70 60 0.8

f. Relative Emission Intensity 100 15 75 120 30 12

g. Emission Wavelength (nm) 410 480 420 440 430 220

h. Index of Refraction 1.85 2.15 1.82 1.8 1.85 1.49

i. Energy Resolution 8 12 12 <8 8 10

j. Half value layer @140 keV (mm)

2.9 0.58 0.69 5.3

k. Half value layer @511 keV (mm)

20.8 7.6 8.3 18.0 10.4 15.8

L. Percentage 511 keV efficiency 25 mm (12.5 mm)

56 (34) 90 (68) 87 (64) 43 (25) 81 (57) 66 (42)

m. Manufactured cost ($US / cc)

5 15 50 25

3 mm to 1 mm would result only in animprovement of Rsystem to 8.0 mm. NaI(Tl)crystals can be manufactured relativelyinexpensively. They are hygroscopic andtherefore need to be hermetically sealed.

Why BGO for PET?BGO has good stopping power for

511 keV photons and achieves goodsensitivity. However, it has a relatively poorlight photon yield and the photons are of awavelength not ideally suited to PMTs. Thisresults in a relatively poor energy resolutioncompared to say NaI(Tl). This is one reasonBGO isn’t constructed into larger crystalwith multiple PMTs — it doesn’t havesufficient light output in such a configurationto determine the position of an event usingAnger logic.

BGO PET cameras have many crystalelements (up to 20,000) arranged radiallyaround the patient. These elements are theminimum resolving units and are typically 6mm x 6 mm and 30 mm deep. In early

models, each element was attached to a PMT,but it was difficult to pack PMTs closetogether. Contemporary BGO PET camerasgenerally use “detector modules” similar tothose shown in Figure 3. Typically, 64elements are partially cut into a single BGOblock.3 Following the absorption of a 511keV photon by one of these elements, light ischanneled along the cell until it reaches theend of the saw cuts, at which point it spreadsand is observed by the four PMTs. Therelative ratio of light measured by the fourPMTs indicate which detector elementabsorbed the photon, and thus identify theposition of the incident 511 keV photon.

This type of design, with smalldetectors, allows the camera to handle thehigh count-rates observed in PET. Dead timeis determined by, and restricted to, the signalcoming from an individual detector (or inthis case a module), so with smaller detectorsthe measured count-rates will be lowerresulting in less dead time compared to largerdetectors.

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Figure 3. Diagram of a BGO PET Camera Detector Module. The module contains 8 x 8 crystal elements and four PMTs. The elements are separated by saw-cuts of varying depths. This allows the element where the event occurred to be located by comparing the ratio of light detected in each of the four PMTs.

Semi-Conductor DetectorsOne of the problems associated with

scintillation crystals is the inefficient way inwhich an incident gamma ray is recorded asan event. It requires absorption, scintillation,and conversion to an electrical signal by thePMT. The high statistical fluctuations thatoccur in this signal leads to poor energyresolution. As stated previously, good energyresolution enables a system to reject photonsthat are likely to have been scattered.

Semi-conductor detectors offer thepromise of producing greater signal to noiseper radiation interaction and therefore betterenergy resolution compared to gammacameras.1

As a gamma ray passes through thesemi-conductor material, it creates “electron-hole pairs” and produces ionization. Themotion of “electron-hole pairs” in an appliedelectric field produces an electrical signal,which is proportional to the energy of thephoton absorbed. The detector performancedepends on the charge collection efficiencyand its ability monitor leakage the gamma-induced current is much higher than theleakage current.

From a theoretical point of view, theimproved signal to noise ratio of the semi-conductor detectors means a potential energyresolution of around 1%. There are problems;however, with the gamma induced electricalcharges being trapped in the semi-conductor,which degrade the detector performance. Thecurrent energy resolution is approximately3%, which should improve with furtherdevelopment. One of the challenges isdeveloping sufficiently fast and complexelectronics to process simultaneously thesignal from large numbers of detectors.

An associated benefit of semi-conductor detectors is their smaller weightand size when compared to scintillationdetectors with PMTs.

Currently the semi-conductorCdZnTe (or CZT) is showing most promise.

It has a relatively high effective atomic num-ber, resulting in good stopping power. A 7 mmthick CZT detector has equivalent stoppingpower to 10 mm of NaI(Tl) at 140 keV.

Semi-conductor detectors forimaging purposes will consist of one detectorper image element or pixel. The size of thiselement will directly influence intrinsicresolution, so to be comparable in resolutionwith Anger gamma cameras, they will needto be approximately 3 mm x 3 mm in crosssection. This would mean that, for a relativelysmall square shaped detector with a FOV30 cm x 30 cm, 104 semi-conductors wouldneed to be integrated.

At present, it is difficult to estimatethe costs of such systems, as mass productionof semi-conductor detectors and integratedcircuitry will greatly improve cost efficiency.Currently there is one system producedcommercially,4 with a relatively small FOV(20 cm x 20 cm), 3 mm x 3 mm detectorelements, and an impressive 250,000 cpsmaximum count-rate. The weight of thedetector head is only 25kg and is about 8cmdeep, including collimator. An equivalentAnger gamma camera would weighapproximately 150 kg and be 40 cm deep.

Tomographic Reconstruction

Filtered Back Projection (FBP)Tomographic reconstruction takes

data acquired radially and creates a transaxialslice that estimates the distribution ofradioactivity within this region. FBP is thelong established technique for converting thedata acquired around the patient in SPECTand PET into slice or transaxial data. FBPwas originally adapted from the first CTalgorithms developed in the mid 1970s, andit has continued to be the only techniqueavailable from most manufacturers until thelate 1990s.

During FBP, data acquired at aparticular address is back projected acrossthe image space as a line of counts whose

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position is determined by the angle andaddress at which it was acquired. The FBPalgorithm assumes that a detected eventcould have originated from anywhere alongthe line perpendicular to the collimator.Following the completion of all the backprojections acquired at all angles the originalpoint source is now represented by thesuperimposition of all the back projections.In Figure 4 it can be noted that there arecounts deposited in the image where theyshouldn’t be and there is blurring of the pointsource. This blurring is theoretically removedby the use of a ‘Ramp’ Fourier filter. It doesthis by isolating the different frequencies thatare present in the blurring and reducing themappropriately. The ‘Ramp’ filter is the inverseshape of the frequencies that comprise thisblurring effect.

The main problem with FBP is thatnoise can lead to significant artifacts. In lowcount studies, statistical noise can lead tostreaking of the data. It is particularlyobvious in the low count areas surroundinghot structures and makes interpretation ofthese areas difficult.

FBP does not take into account theeffects of attenuation, scatter and the depthdependency of collimator resolution; itassumes that data is represented by simpleprojection ray sums.

Iterative MethodsIterative, or statistical, techniques

achieve reconstruction by repetitivelyrefining the reconstructed data. It does thisby ‘guessing’ how the reconstructed datawould appear to the detector and comparingit to what the acquired data actually lookslike. The difference between the two is seento be the error in the reconstructed data andappropriate modifications are made. Aftermuch iteration of approximations, thereconstructed and acquired data will(hopefully) converge.

The attraction of this technique is that

the various factors that affect the signal

before it reaches the detector can be modeled

into the reconstruction process.1 In theory,

the effects of scatter, attenuation, and depth

effect on resolution can be calculated and

compared to what the detector ‘sees’ and be

corrected.

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Figure 4. Acquisition and reconstruction of tomographic data into transaxial slices using FBP. The ‘Ramp’ Fourier filter is essential in eliminating the superimposition of ray sums created during the back-projection process.

Accelerated Methods – Ordered SubsetsThe knowledge to perform iterative

reconstructions in tomography has beenaround for some time. However, thecomputational power required to effectivelyrun the algorithms clinically has not beenavailable. For this reason alone, thecomparatively computation-efficient FBPhas been the preferred method.

The ordered subsets-expectationmaximization (OSEM) method is an iterativereconstruction technique, which performsremarkably efficiently when compared toconventional iterative methods.5 In OSEM, asubset of projections is used per iterationinstead of the complete set. It uses an orderlyselection of projections per iteration, so thatmaximum new information is provided.6

Reconstruction times are proportional to thenumber of projections processed per iterationso if only four projections per subset are usedinstead of 128, OSEM will performapproximately 32 times faster. Surprisingly,although OSEM uses only a fraction of theavailable projection data per iteration, therate of convergence between measured andobserved data is approximately equivalent tothat of the conventional iterative methods.

The majority of SPECT and PETvendors currently supply OSEM and its useis widespread. Its speed of reconstruction isslow compared with FBP (probablycomparable to FBP five years ago), althoughit is an attractive alternative to FBP. As fastercomputers become available more complexmodeling of attenuation, scatter, resolution,and noise should improve the accuracy ofreconstruction.

INSTRUMENTATION

Gamma Camera Design

IntroductionAs mentioned previously the

fundamental detector system of the gammacamera has remained essentially unchanged

through the 1990s. The gamma camera as asystem has improved greatly as an imagingsystem. Much effort has been spentincreasing study throughput. Quality controlprocedures have been streamlined. Pro-cessing is faster, but it is also more complex.Hospital-wide computer networking allowson-line, remote review and archiving ofstudies, and integration into hospitalinformatics means greater accessibility.

GantryThe complexity of the gamma camera

gantry is mainly because of its requirementto perform SPECT. Detectors with associatedshielding and collimation weigh severalhundred kilograms and they need to bemoved closely and safely, around or along thepatient. The orbit, especially in SPECT, mustfollow an exact path. Rotation of the heavygantry with the requirement for preciseangular movements results in complexacceleration and deceleration control usingstepper motors. As well, patient safetynecessitates a system that detects resistanceto detector movement and touch sensitivesurfaces. These requirements mean thatmuch of the detector motion is undercomputer processor control.

Collimator resolution is depthdependent. Scan quality is improved if thedistance between patient and collimator iskept to a minimum. Most current systems usepatient-sensing technology, commonlyphoto-diodes, which either predetermine anorbit at set-up, or move detectors to a closeposition while scanning. This necessitatesfail-safe mechanisms in respect to patientsafety.

There have also been improvementsin the manufacture of SPECT scanning bedsor pallets. These pallets are commonlysupported at one end (diving board style) andare relatively narrow to allow the detector torotate laterally and posteriorly around thepatient. They must be able to rigidly support

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heavy patients while being relatively radio-translucent. Fiberglass has been replaced bycarbon fiber and Kevlar as the materials ofchoice.

CollimationBackground. Collimator design has

remained relatively static over the lastdecade. The vast majority of collimatorsremain of the parallel-hole type.

The manufacturing of collimators hasbeen refined and their quality in mostinstances fulfils the more stringent criteriafor SPECT imaging. SPECT requires lowvariation in hole angulation. This was aconcern in the early days of SPECT whencollimators not specifically manufactured forthe purpose were being used. Mostcollimators are currentlyproduced using a castingtechnique that has improvedrobustness and uniformitywhen compared to foil-typecollimators.7,8

As previously men-tioned, collimator selection isa compromise between reso-lution and sensitivity. Mostinstalled gamma camerashave at least two collimatorsto choose from, or more if thecamera is to be used forimaging radionuclides withhigher energy photons.Several physical parametersneed to be considered whenselecting a collimator. Theseare depicted in Figure 5.

Sensitivity = (0.069 x 2.54)/(32 x (2.5+0.2))2

= 0.00036

~ One photon per 2800

Resolution @5 cm depth = (2.5 x (32+50))/32

= 10.3 mm

Collimator sensitivity and resolutionare defined as follows:

Collimator sensitivity = Khole shape * d4

______________(L * (d + s))2

Collimator resolution = d * (L + H)____________ L

(Where H = source distance fromcollimator, and Khole shape is an equationconstant determined by different shapedholes. For hexagonal holes, its value is 0.069,0.080 for square holes, and 0.057 for circularholes).38

It is important to note from these twoequations that sensitivity to activity (in the

absence of attenuation) remains independentof scanning distance (H), whereas in parallel-hole collimators resolution is depthdependent.

Recent Innovations. While the aboveequations hold true for parallel-holecollimators, it is possible to enhance systemresolution with the use of convergent

13

Figure 5. Cross-section through cast collimator with hexagonal holes.

collimators. Convergent collimators arecollimators where the holes point to a focus.The characteristics of these collimators aresuch that (i) as an object moves away fromthe detector it will appear to be magnified onthe detector, (ii) sensitivity actually increasesas the object moves away from the detector,and (iii) resolution at depth is enhanced whena suitable amount of convergence isemployed. The distance to the point ofconvergence or focus is dependent on theapplication and the size of the detector but istypically 50cm. The main disadvantage of theconvergent collimator is that the field of viewgets smaller as the detector gets further fromthe patient. For this reason the applicationsfor convergent collimators are limited,especially where truncation of the object bythe FOV is intolerable (i.e., SPECT).

Typically, tomographic reconstruc-tion is performed by reconstructingindividual transaxial slices. These slices canthen be ‘stacked’ to produce a volume ofdata. Reconstruction in tomography iscomplicated when the acquisition andsubsequent back projection of data is notperpendicular to the axis of rotation (i.e., theback-projection path travels through multipletransaxial slices).

The fan-beam collimator (Figure 6A,6D) has been developed for small FOVSPECT studies.7,9 It combines improvedsensitivity-resolution performance by beingconvergent in the transaxial plane whilecomplying with reconstruction limitations bybeing parallel in the axial plane. SPECTreconstruction requires that the sides of thebody being imaged are not truncated by the FOV,so the fan-beam collimator use is restricted tobrain and pediatric work. The software usedto reconstruct parallel-collimated studiesdoes not need to be modified greatly toreconstruct fan-beam studies.

Cone-beam collimators (Figure 6B,6D) are axially and transaxially convergent.In theory, they have excellent performance

characteristics with very high sensitivity andresolution,10,7 but they are difficult toimplement and reconstruction is complex.

Multi-Headed CamerasIn a survey11 of new gamma cameras

installed in Australia in 1994-1996, 52%were dual headed, 16% triple headed, and

14

Figure 6. Design and FOV for parallel hole(A and C), fan-beam (A and D) and cone-beam (B and D) collimators in axial (A andB) and transaxial (C and D) planes. Con-vergent collimators have a reduced FOV thatmakes them suitable for small objects only(neurology and pediatrics); otherwise, FOVtruncation of body may occur (see D). Cone-beam collimators that converge in the axialplane, result in more-complex tomographicreconstruction since the photon’s pathintersects multiple transaxial planes (B).

32% single headed. These figures will besimilar to other markets. Multi-headedgamma cameras are a simple yet oftenexpensive solution to system insensitivity.This approach reduces scanning time and/orimproves scan quality.

Many multi-headed detectors arebased on fixing multiple large field of view(LFOV) detectors on the one gantry. Thesesystems are best suited to wide area surveysin both SPECT and planar imaging, but notall multi-headed gamma cameras have beendesigned in accordance with this philosophy.Systems optimized for specific types ofimaging have also emerged. MyocardialSPECT requires a relatively small axial FOV.One manufacturer determined that the areasof the detector imaging superior and inferiorto the heart (or other small organs) werebeing wasted and these parts of the detectorcould be better utilized as a second detectorat the same axial level. This camera (GEOptima) effectively doubled the sensitivity ofthe single headed LFOV alternative, butdidn’t add greatly to the price.

Most early model dual-headedgamma cameras had fixed opposed detectors.These systems are ideal for 360° SPECTimaging and simultaneous anterior andposterior planar whole-body sweeps.However many centers perform 180°-SPECTwhen performing Thallium-201 myocardialstudies. In this situation, imaging with fixedopposed gamma cameras has no benefitsover a single headed gamma camera. Mostdual-headed systems currently on the marketare available in a variable angleconfiguration with detectors able to be set up90° to each other for 180° SPECT. For thistype of acquisition, the gantry rotatesthrough 90° to achieve the most efficientutilization of both detectors.

Multi-headed gamma camerasrequire more complex geometric calibrationas each detector needs to work in unison withthe others.

Attenuation CorrectionBackground. The characteristic

attenuation of photons by the patient’s bodymakes it difficult to appreciate thedistribution of radiotracer distribution. Noother single factor restricts SPECT’s abilityto achieve absolute quantification ofradiotracer distribution.

In many clinical studies, allowancefor the pronounced effects of attenuation isrequired through knowledge of the patient’sbody habitus; without it the accuracy of thestudy suffers. One of the main purposes forattenuation correction is accuratereconstruction of myocardial perfusionstudies. The heart position means that it isparticularly prone to regional differences dueto attenuation. It sits not only adjacent tolung tissue and mediastinum, but also at thebase of the thoracic cavity where there is alarge change in attenuation when movinginferiorly.

Corrections for attenuation intomography are available and they are of twomain types:• Those that assume the attenuative nature of

the body (first-order method), and• Those that derive the attenuative nature of

the body (transmission method).First Order Attenuation Correction

(Chang). Chang’s method is used widely andis the most easily implemented.

The method calculates a restorativefactor for individual pixels within atransaxial slice based on their relativeposition within the body.7 Its failing is that itassumes a homogeneous attenuationenvironment within the body. This is asignificant issue, especially in areas such asthe chest, in which there is bone, soft tissue,air, and lung. Thus, it is useful only whenthere is relative homogeneity in attenuation(e.g., abdomen, pelvis and brain). In areas ofthe body in which counts are different due toregional differences in attenuation (e.g., inmyocardial perfusion studies), the relative

15

differences will be maintained followingcorrection.

The first step is to determine theboundary of the body. The depths ofindividual pixels within transaxial slices arecalculated at each projection angle. Theattenuation correction factor applied to apixel is the inverse of the average of theattenuation factors obtained at eachprojection angle.

Measured Transmission AttenuationCorrection. Some of the early research intoSPECT attenuation correction wasperformed using co-registered CT andSPECT studies. Attenuation coefficientswere extrapolated from the CT Hounsfieldvalues.12,13 There were many problems,including the accurate alignment of CT andSPECT studies, and the errors in theapproximating of linear attenuationcoefficients from CT values. Interestingly,current work in progress from onemanufacturer is hoping to use a combinedSPECT and CT system for this application.

An accurate attenuation map of bodytissue can be determined by performing atransmission study akin to CT but using aradioactive source. Early techniques used aflat flood source of radioactivity mounted tothe detector and on the opposite side of thepatient (Figure 7A).14 It showed the potentialfor measured transmission attenuationcorrection (MTAC) but it added greatly to thetime of the study.

Some of the factors that need to beaddressed in providing accurate and effectiveMTAC are:• Source selection – ideally a source with

similar photon energy to the emissionradionuclide is used. This allows for themost accurate attenuation coefficientcalculation, and means that the detectorperformance is matched between emissionand transmission studies.

• Signal separation – The acquisition ofboth signals (emission and transmission)

must be separable and cross-contaminationminimized. Signal separation can beachieved by doing the transmission studyon the patient before radionuclideadministration. However, this isn’tpractical in many types of studies wherethere is a prolonged uptake phase (e.g.,bone scanning) or where other factorsdetermine the administration of theradionuclide (e.g., myocardial stressstudies). A common method of signalseparation is to use a transmission sourcewith a photon whose energy is similar tothe emission photon but separable byenergy discrimination. Simultaneousacquisitions of transmission and emissionsignals can be performed but down-scatterfrom the upper to lower energy windowneeds to be allowed for. Another method isto dedicate one detector of a multi-detectorcamera to acquiring the transmission data.

• Exposure to source – Ideally the sourceshould be shielded when not in use.Exposure to patient and operator should bekept to a minimum. The half-life should besuitably long so as not to require frequenthandling or costly replacement.

Moving Line Source. One commonapproach is to use the “moving line source”technique for MTAC (Figure 7B). From apractical point of view, it would be optimal toperform emission and transmission studiessimultaneously, but as mentioned above thesignals must be separable and notcompromised by downscatter.

With this technique, a line sourcecollimated to only project photons throughthe patient perpendicular to the detectorcollimator sweeps across the FOV at eachSPECT projection frame. A position de-coder attached to the line source tells thedetector where at a particular time the linesource is and extracts from the gammacamera’s signal the line of image data(around 50 mm wide) that reflects thetransmission signal. Other areas of the

16

gamma camera’s image that don’t containtransmission source counts are stored asemission data.

The advantage of this system is thatthe transmission signal can be confined to atarget area leaving other areas of the FOVunaffected for collection of emission data.The strength of the transmission signal canbe increased so that the contribution to it byemission signal is trivial.

A further improvement may begained by combining fan beam and movingsource technologies, by moving a pointsource along the focal line of a fan-beamcollimator (Figure 7D).15

Multi-Detector Systems. In multi-detector gamma cameras, one of thedetectors may be dedicated toacquiring transmission data. Movingline source techniques are difficult toimplement when there are opposingdetectors. A solution is a fan-beamcollimator with a line source positionedat its foci (Figure 7C).

This is easily implemented ontriple headed gamma cameras, and haseven been developed on gammacameras with opposing detectors usingan asymmetric fan-beam system.

Dedicated PET Camera Design

IntroductionPET involves the detection of

the 511 keV annihilation photons thatare emitted in near opposite directionsfollowing positron decay. Interestingly,as PET devices improve in resolutionsome of the characteristics of positrondecay are more often seen as theresolving limit of this modality.16

PET imaging assumes that iftwo annihilation photons are recordedin coincidence on opposite sides of thepatient then the nucleus from which thepositron undergoing annihilationoriginated must be on a straight line

between the two points of detection. This lineis known as the line of response (LOR). Twoproperties impact on this assumption. Firstly,the positron has energy, which it loses duringcollision with the orbital electrons ofsurrounding tissue. Once it has lost most ofits energy, it annihilates with an electron toform a positronium. Its range depends on theenergy of its emission (ß+ Emax), which isnuclide dependent. The estimated effect onPET imaging resolution is approximately 0.2mm FWHM for 18-F (ß+ Emax = 0.64Mev)and 4.5 mm for Rb-82 (ß+ Emax = 3.35MeV).

Secondly, annihilation photons aren’temitted in exactly opposite directions.Photon non-colinearity amounts to

17

Figure 7. Comparison of the geometry of various transmission configurations (top row) for attenuation correction in SPECT.15 The bottom row shows the relative sensitivity (T) for the source activity stated. Downscatter from the emission isotope (99m-Tc, 140 keV) into the transmission window (153Gd, 100 keV) was calculated (Ed), and related to the transmission counts for each geometry. Best results are achieved when transmission sensitivity (T) is highest and relative emission downscatter (Ed) is lowest. High transmission sensitivity has the added benefit of reduced source activity and therefore inexpensiveer purchase and replacement costs. (Reprinted with permission of the Society of Nuclear Medicine from: Beekman FJ, et al.Half-fanbeam collimators combined with scanning pointsources for simultaneous emission-transmission imaging.J Nucl Med.1998; 39:1996-2003.

approximately 180° ± 0.25°. The effect ofnon-colinearity on PET imaging resolutiondepends on the separation of the detectors. Ina typical camera, the resolution loss is about2.1 mm. The small-diameter detector ring ofsmall animal PET scanners means that thiseffect is reduced.17

Detector Selection, Design and PropertiesHistorically PET cameras have used

BGO crystals as small discrete detectorsin a ring. The size of each crystal elementis the major deter-minant of systemresolution. For a 6mm square detectorelement the geo-metric resolution is3 mm in the centerof the FOV.18

The best in-trinsic resolution inPET detector ring isachieved at thecenter of the FOV.The rate of deteri-oration in resolutionaway from thecenter depends onthe diameter of thedetector ring. Photonsemitted further awayfrom the center ofthe FOV are morelikely to hit thecrystal elements atan angle (Figure 8),which produces agreater positionaluncertainty due to aparallax error. Notethat as the ringdiameter increasesthere is less degra-dation in resolutionas you move furtheraway from the

center, although the non-colinearity errorincreases.

Research is being undertaken todetermine where, along the length of acrystal, a photon is absorbed. This is knownas “depth of interaction.”19,16 Estimatessuggest that if the depth of interaction couldbe calculated accurately to one-third of thedepth of the crystal; parallax error, asdescribed in Figure 8, could almost beeliminated.18 Depth of interaction cal-culations can be performed with the use of

18

(i)

(ii)

Figure 8. PET ring detector. (i) Detector elements, which are typically30 mm deep and 6 mm square, are mounted radially around the detector.Photons traveling from the center of the FOV will hit the elementparallel with its longest dimension. As the distance increases from thecenter of the FOV, a parallax error occurs which affects resolution. Theparallax error is greater in scanners with smaller radii and hencedegradation of resolution is greater. (ii) Loss of resolution away fromthe center of the FOV is caused by parallax error. Two photons fromseparate sources (A) can be resolved because they are detected inadjacent elements. Two photons from separate sources cannot beresolved if they are detected in the same element (B). The parallax errorincreases as the distance increases from the center of the FOV.Determining the depth of interaction within in the crystal element canreduce the error.

light measuring photo-diodes at the patientend of the crystal, which determine in whichof the elements within a detector module anevent has occurred. This is then compared tothe amount of light seen by a single PMT atthe opposite end and from their ratio, thedepth of interaction within the crystal iscalculated. Detector modules like the oneshown in Figure 9 are being developed toallow depth of interaction calculation.

The ability to calculate depth ofinteraction will enable the diameter of thedetector ring to be reduced, decreasing thenumber of detectors (and cost). It will alsoreduce the deterioration in resolution causedby annihilation photon non-colinearity.

The recent development of LutetiumOxyorthosilicate (LSO) crystals may impacton detector selection. LSO has manyproperties that make it more suitable for PETthan BGO. Of particular importance are

more efficient light photon production and ashorter light decay constant. This shouldresult in better energy resolution that willreduce the contribution of scatter. As well,increases in the amount of light produced perscintillation should mean that each detectormodule could be cut into smaller elements,as the signal will be sufficiently strong toallow more accurate determination of itsposition of interaction.

There is also potential, due to thehigh light output, for an LSO detector to bemanufactured in larger detectors usingsimilar logic to the gamma camera. Such adetector would need to be able to handle veryhigh count-rates and this should be possiblewith the short light decay constant of LSO.

2D vs. 3D AcquisitionFor many years, PET scanners

acquired images in 2D mode using ringshaped septa to restrict coincidences to oneof many transverse planes. In doing so, thesecoincidences were acquired and recon-structed in the one transverse section.Researchers recognized the inefficiency ofthis mode of acquisition and attempted toincrease sensitivity by removing the septa(Figure 10).

When septa were removed, there wasa dramatic increase in sensitivity (500%),unfortunately accompanied by a dramaticincrease in the scatter fraction (from 10% to35%). In 2D acquisitions scatter contributionis mainly confined to photons that remain inthe same transaxial plane following collision,as otherwise they will be absorbed by thesepta. In 3D acquisitions, photons scatteredin any direction have an unimpeded path tothe detector.

The increased sensitivity of 3Dacquisitions is attractive for a number ofreasons.37 A large proportion of patients arevolunteers in whom the radiation dose isquite consequential. Ethical approval oftendemands restrictions in the number of scansperformed and the dose administered to these

19

Figure 9. PET detector module with matrixof photodiodes for calculating ‘depth ofinteraction’. The module uses a singleconventional PMT (not shown) at theopposite and rear ends of crystal elements.The crystal element where scintillationoccurs is identified by a single photodiodewithin the matrix. Ratio of signals betweenphotodiode and PMT reflects “depth ofinteraction.”

groups. Many studies are also undertakenusing tracers with low uptake and on whomdynamic studies result in short acquisitionintervals. 3D acquisitions are often employedin such studies to improve image quality byreducing statistical noise.

Problems arising from 3D PETacquisitions include:• The increased scatter fraction requires

more extensive and vigilant correction. • The detectors are no longer shielded from

“out of field” activity. Hot structures suchas the brain and bladder can contributesignificantly to random and scatterfractions.

• In 2D PET, there is consistent sensitivity toactivity along the axial plane. In 3D PET,the axial acceptance angle for coincidencedecreases closer to the ends of the FOV.Axial normalization is required to correct

for this reduced res-ponse, however, reducedsampling means thatthere is greater noise.When acquiring multiplebed position whole-body studies, 3D PETsystems typicallyoverlap by half the axialFOV to achieve constantsam-pling and noisealong the axial plane.

• Coincidences detectedare no longer confinedto a single transaxialplane. The LOR of acoincidence can intersectmany transaxial planes.This requires eitherapproximation of thedata to a transaxial slice(single slice rebinning)or complex 3D recon-struction algorithms notdissimilar to thoserequired for conebeamcollimated SPECT.

In 3D scanners, scatter contribution isprimarily eliminated through energydiscrimination. The commonly used BGOscintillator has relatively poor energyresolution and therefore a future replacementscintillator should have better energyresolution to be able to take advantage of 3Dtechnology.

Attenuation and Scatter CorrectionBackground. Despite the penetrating

nature of 511 keV photons when comparedto the low energy photons used in themajority of conventional nuclear medicinestudies, PET scanning is subject topronounced attenuation by the patient. PETrequires the detection of two single 511 keVphotons in coincidence; the attenuation effecton those two single photons is thereforecompounded. The half layer value for 511 keV

20

Figure 10. Axial cross-section through PET scanners in 2D and 3D modes.

photons in water is approximately 7 cm. Thismeans that 511 keV photons emitted from apoint in the middle of a water-filled cylinderwith radius of 7 cm have a 50% probabilityof exiting the body without being attenuated.However, the overall probability of detectingboth photons in coincidence is only 25%. Ina cylinder with 14 cm radius (which is notdissimilar to the average human torso) theprobability of remaining unattenuated is6.25% (0.25 squared). The equivalent value

for 140 keV photons in SPECT would be 12%.Figure 11 shows that the effect of attenuationon 511 keV coincidence imaging is more pro-nounced than that on 201-Tl SPECT imaging.

The established method forperforming transmission correction in PETcameras uses positron emitting rod sourcesof relatively long half-life isotopes fortransmission sources. These are rotated 360°around the patient and the density ofrecorded coincidences reflects attenuation inthe patient. The algorithm that reconstructstransmission data uses the same principles asCT reconstruction and the resultant imagesare comprised of attenuation coefficients.Germaniun-68 (T1/2 =270 days) sources aretypically used and are retractable whenperforming an emission study.

Post-Injection Transmission Studies.As emission and transmission signals hittingthe detectors are both 511 keV photons, earlyPET scanners needed to perform thetransmission study prior to doseadministration to the patient and the patientneeded to remain stationary duringtransmission, uptake and emission periods.For studies with long uptake times, this wasquite difficult and resulted in inefficient useof camera resources.

The ability to perform post-injectiontransmission studies became available withthe development of a technique that wascapable of predicting geometrically whethera coincidence originated from the patient ortransmission source. It achieved this byrecording whether a LOR intersects with theposition of the transmission source as itrotates at the moment of detection.22 If aphoton doesn’t meet these criteria it isrejected (Figure 12i). Some emissioncoincidences may accidentally intersect withthe source but their contaminatingcontribution is negligible (<1%). Thistechnique has been refined to the point wheretransmission and emission studies can beperformed simultaneously.

21

Figure 11. Left column shows the attenuation correction images for attenuation applied to different radioisotopes imaged with SPECT and PET of a phantom. Right column shows count profiles graphs. Actual restorative values for correction of attenuation are displayed in the y-axis. Central maximum restorative factors for Tl201, Tc99m, F18-SPECT, and F18-PET are 10,6,3 and 20 respectively. (Figure provided courtesy Dale Bailey, Hammersmith Hospital, UK.)

Singles Transmission. A limitation ofPET systems that operate close to theirmaximum count-rate capability is dead timeloss due to the rod source passing close to thedetectors during transmission scanning. Thiscan lead to significant image degradation ifthe source is too radioactive.

An alternative is to use a ‘singles’source, usually Cesium-137 (T1/2 = 30 years,gamma=660 keV), which is collimated toproject a fan beam of radiation through thepatient towards the detectors opposite, and isretracted when not in use. Shielding protectsdetectors adjacent to the source and hence amuch greater activity can be used when com-pared to positron sources (Figure 12ii[a]). Apseudo-coincidence is formed between thesource and the detected event.23 The trans-mission signal can be differentiated from theemission signal by energy discrimination(Figure 12ii[b]) so that post-injectiontransmission scanning can be performed.

An additional benefit is that due tothe long half-life of Cesium-137 there is norecurrent expense of source replacement.

Transmission Image Segmentation.One issue with transmission attenuationcorrection is that it is achieved through anarithmetic operation between two sets ofimages that contain noise,1 which leads to acorrected image whose noise component ishigher than that of either of the emission ortransmission images. Transmission imagescontain pixel attenuation coefficient data.Segmentation is a technique that reducesnoise dramatically in areas of uniformattenuation, by giving all pixels the sameattenuation coefficient. Typically, all softtissue equivalent values are segmented, whileothers such as bone and lung use themeasured value.

Scatter Correction. In 3D PET of thechest, it has been estimated that 50% or moreof detected coincidences have undergonesome Compton scatter within the body.24 Thiscompares to a scatter fraction of 10-15% in2D PET with septa.

As with SPECT one of the majorfactors contributing to acquiring scatteredphotons is detector energy resolution. Poorenergy resolution leads to increased scattercontribution. Shielding can also reducescatter. The septa rings used in 2D PETacquisition drastically reduce scatter, butthey also reduce system sensitivity by afactor of five.

Correction for scatter contribution isusually performed by measuring the counts

22

Figure 12. Transmission attenuation correctionin PET. (i) Post-injection technique with rotating

positron rod source. LOR is deemed to befrom the transmission source (‘t’) if itintersects with it. A small percentage(<1%) of emission (‘e’) LORs will beincorrectly recorded as transmission LORs.

(ii) (a) Singles attenuation correction withrotating collimated Cesium-137 point source.(b) Separation of transmission andemission photons by energy discrimination.

outside the body. Techniques can determinethe amount of counts contributed by scatterwithin the patient by measuring andextrapolating counts from outside the patient.

New techniques are being developedwhich model the scatter component in theemission image based on the transmissionimages and subtract its contribution.

Partial Ring BGO PET ScannersThe development of 3D acquisition

techniques for PET scanning meant thatcounts achieved were sufficiently great tocontemplate reducing the number of

detectors. The primary motivation was tomanufacture a PET scanner at greatlyreduced cost (Table 4) (Siemens ECAT ART,CTI PET Systems, TN).

A scanner20 was developed whichcomprised two banks of opposed detectors,each spanning 82.5°. No septa were installedso acquisitions were restricted to 3D.Thedetectors were BGO modules and identical tothose used in a full ring PET system capableof 2D and 3D acquisitions (Siemens ECATEXACT, CTI PET Systems, TN, USA). Thedetectors are fixed on a slip ring that rotates

23

Table 4. Comparison of Commercially available PET systems. Each example is approximately representative of market competitors 1,20

Dedicated Full ring

BGO*

Dedicated Partial ring

BGO†

Dedicated Full ring NaI(Tl) ‡

Coincidence Gamma Camera§

Crystal dimensions Width x length X depth (mm)

3.9x8.3x30

6.5x6.5x20

500x300x25

510x380x15.9

No. of crystals (or elements)

12K 4.2K 6 2

Detector diameter (mm)

93 82 90 66

Transaxial Resolution (mm)

4.8 6.3 5.8 4.8

Scatter fraction (%) 3D(2D)

35(10) 32 27.8 26

Sensitivity (cps/Bq/ml) 3D(2D)

27.0 (5.4) 7.5 12.4 3.7

Crystal volume (cc) 11800 3570 22000 3081

Axial FOV (cm) 15 16.2 25 34

Approx. cost US $ 2,000,000 1,300,000 1,200,000

120,000 (Coincidence

detection option)

400,000 (Gamma camera)

*GE Advance (GE Medical Systems Inc) †Siemens Exact ART (CTI/Siemens) ‡ ADAC C-PET (ADAC Laboratories) §ADAC MCD (ADAC Laboratories)

at 30 rpm, and during acquisition, the ringrotates continuously.

Comparing the systems found thatthe 3D partial ring system had approximatelyone third the sensitivity of the full-ringsystem in 3D mode, but twice the sensitivityof the full-ring system in 2D mode.

Further cost savings are made in thismodel by using BGO crystals with a depth of20 mm (cf. 30 mm in most high-end PETsystems).

Dedicated NaI(Tl) PET ScannersDespite its relatively poor 511 keV

photon stopping power, one group persistsusing NaI(Tl) for dedicated PET scanners(UGM/PennPET) with very good results.25

The advantage of NaI(Tl) is that it has verygood light output per interaction, and it isinexpensive.

One of the unique features of thissystem is that it uses large crystal detectorsof gamma camera proportions, in contrast toBGO PET scanners where small crystalelements are the smallest resolving unit.Resolution is not therefore restricted to thephysical dimension of these elements.

The PennPET scanner consists of sixsuch detectors with PMTs arranged aroundthe patient, with an axial FOV of up to250 mm. The geometry is akin to a 6-headedstationary gamma camera with crystals fixedas close as possible to each other. Crystalthickness is 25 mm, which is approximatelytwice as thick as those used by conventionalgamma cameras. This gives it an efficiencyof 56% for 511 keV photons and 32% incoincidence, which is well short of BGOperformance in comparison at 90% and 81%,respectively.

In has achieved good performanceand a niche in the market by addressing someof the problems associated with NaI(Tl) incoincidence mode.

A major hurdle and remaininglimitation is the maximum countrate

achievable. The detector operates exclusivelyin 3D mode without shielding betweenpatient and detector. In SPECT the detectoris protected from the vast majority ofradiation emerging from patient by thecollimator and dead time loss is rarely anissue. Typically, clinical count rates rarelyexceed 10 Kcps and dead time loss becomesan issue when count rates exceed 50 Kcps.The PennPET scanner typically operates at500 Kcps per detector or 3 Mcps for thewhole system. These rates are typicallyobserved in the 250 mm axial FOV scanneran hour following 3 mCi (111MBq) of FDG.High count rates are achieved through twoinnovations, which are a departure fromAnger logic that has serviced conventionalgamma cameras for 30 years:1. When scintillation occurs in a gamma

camera the light produced is detected andmeasured by all PMTs of the detector.During this time (dead time), the system iseffectively paralyzed from detectingfurther events. The PennPET scannercircumvents detector wide dead time bydividing each detector in 3 overlapping“virtual” detectors that can operateindependently (Fig.13A). Determinationof the position of an event is restrictedto the closest PMT and surrounding sixPMTs.

2. Clipping the electrical signal from thePMT after 150nsec further reduces deadtime. Sufficient light is produced by theNaI(Tl) crystal to forego the signal in thelatter section of the pulse (Figure 13B).Accurate energy calculation is obtainedby integrating the area under the earlysection of the pulse.

These two innovations increase by anorder of magnitude the maximum operatingcount rate of these detectors compared toAnger type systems. The technology of theUGM/PennPET system was subsequentlyapplied to dual headed gamma cameras, andthe manufacture of the first commercial

24

hybrid PET/SPECT gamma camera (ADACLaboratories, MCD).

Count rate limitations make thisscanner unsuitable for short half-life studiesthat typically scan following bolusadministration of greater than 30mCi, but arewell suited to relatively steady state studieslike those with FDG.

This group has further improvedsystem performance with the use of sixcurved detectors, which are formed into acircular arrangement.26,27

Time-of-Flight PET DetectorsWhen 511 keV photons are detected

on opposite sides of an object within adefined time window, they are deemed toform a coincidence. This coincidence isrecorded as a LOR and the nucleus fromwhich the event originated is predicted tohave arisen from somewhere along this line.

Normally the coincidence window is5 ns to 15 ns. This window is long enough toallow for sampling errors in the timing ofevents, but short enough to exclude randomcoincidences. The detector property thatimpacts most on the accuracy of timing of anevent is the photon decay constant.

If the timing of detected events can bedetermined with sufficient accuracy then the

position along the LOR from wherethe annihilation photons originatedcan be calculated by looking at thedifferent arrival times at thedetectors. In a perfect time-of-flightsystem a coincidence of events wouldultimately be recorded as a count in avoxel and obviate the need fortomographic reconstruction. Currenttechnology allows for a minimumtiming resolution of around 0.3ns,which equates to around 5cm alongthe LOR. This amount of timingresolution would still requireacquired data to undergotomographic reconstruction albeit ina constrained manner.39

Barium Fluoride (BaF2) hasappropriately short photon decayconstant (0.8 ns), but it suffers fromrelatively poor light output, awavelength not suited toconventional PMTs and poor photonstopping power.

Time-of-flight detection maybe possible with LSO. It has a photondecay constant of 40 nanosec, whichis much longer than BaF2, however,event timing of sufficient accuracymay be possible to allow time-of-flight detection.40

25

Figure 13. NaI(Tl) detector modifications. Singledetector zoned into separate “virtual” detectors thatoperate independently. Photon scintillation processedby closest PMT and six surrounding PMTs only. Deadtime effectively restricted to zone and not whole crystal.A. PMT signal is clipped after 150 nsec and energyintegrated under early section of pulse, thus reducingdead time.

Multi-Modality Cameras – Marriage of CTand PET

Much of the research into the clinicalefficacy of PET has identified that improvedaccuracy of diagnosis is markedly improvedby combined observation of PET (functionalimaging) and CT (structural imaging), asopposed to each modality in isolation. Toimprove the combining of modalities, digitalco-registration of functional and structuralimages is pursued by many clinicians;however, this is difficult to achieve with anyaccuracy due to the studies being acquiredduring different sessions under differentconditions and unavailability of images in adigital format.

A system that combines highperformance PET and spiral CT in the onescanner system should be on the marketsoon.28 It is essentially a PET and CT scannerbolted together with a common scanning bed.It will provide the facility for accurate co-registration and potentially allow CT tosupply PET with transmission images forcorrection of scatter and attenuation. It willalso allow PET data to be combined with CTfor radiotherapy planning and includefunction as a basis for treatment rather thanstructure alone.

Such systems do not acquire theirdata of the same tissue simultaneously.Scans are usually acquired end-to-end albeitin the same session. There is still potentialfor mis-registration of function and structuredue to patient movement or changes in theanatomy of the patient (eg. filling bladder ormoving small bowel).

In the future a scanner may bedeveloped which is designed as a combinedscanner which can simultaneously acquirefunctional and structural data rather than asseparate units bolted togeter.

Future Developments in PET ScannersMuch of the future of PET scanning

will rely on the selection of a replacement for

BGO. BGO is the current crystal of choice inhigh-end dedicated PET scanners; however,over the last few years LSO has emerged as areplacement with more desirable physicalproperties. Its greater and faster photonproduction properties, compared to BGO,should allow the construction of smallercrystal elements and therefore better systemresolution, and allow accurate ‘depth ofinteraction’ calculation. Despite this, itsenergy resolution in a standard configurationisn’t any better than BGO. As mentionedpreviously energy resolution is an importantproperty in the control of scattercontribution, especially in 3D scanners.

The GSO scintillator offers goodenergy resolution and may be less expensivethan the LSO. A dedicated PET camera usingthis detector material is about to released onthe market.

Tomographic reconstruction of 3Dacquisitions is being improved with severaltechniques now available that perform true3D reconstructions, as opposed toapproximation (or rebinning) of data into 2D.

The physical joining of PET scannersto CT (or even MRI) should improve theaccuracy of diagnosis in many clinical areas,and assist in surgical and radiotherapyplanning. Although not completely realisedyet, the potential for accurate attenuation andscatter modeling from acquired CT datashows promise.

Hybrid SPECT/PET Gamma CameraSystems

IntroductionCurrently most manufacturers are

offering coincidence based gamma camerasystems capable of performing PET.Although performance specifications areinferior to dedicated PET scanners and thefuture holds promise for what is a youngtechnology.

The ability to perform both SPECTand PET on the one scanner is economically

26

attractive. PET scanners usually experienceconsiderable down time due to unavailabilityof dose, especially in departments withouton-site cyclotron facilities. Hybrid camerasare available for SPECT imaging at thesetimes.

Capital investment is appreciably lessthan for dedicated PET (Table 4). Thecoincidence detection option for most ofthese types of cameras is approximatelyUS$120K on top of the gamma camera costof US$400K. This compares favorably withthe cost of high end dedicated PET ofapproximately US$2000K.

There has been considerable debateon the merits of hybrid and dedicated PETscanners and a great deal of analysis of theirrespective physical performance.29 Scanquality of hybrid scanners is less than that ofdedicated PET scanners. This impacts mostin the detection of small volume disease orwhere there is low contrast, especially inareas of the body subjected to moreattenuation (e.g. abdomen). Unfortunately,access to PET imaging is limited in manyhealth regions. Often the debate isn’t whethera study should be performed on a hybrid ordedicated PET camera, but whether any PETscan is available to the patient. In thissituation and with knowledge of theirlimitation, hybrid scanners may play a role insome applications.30

Some consensus has been reachedfollowing clinical testing. Hybrid camerasshow best performance in areas of the bodythat aren’t subjected to great amounts ofattenuation. It shows best results in head,neck, and thorax. In a comparison of hybridand dedicated PET in detecting suspectedbody malignancies,31 gamma-camera-basedPET found 78% of all lesions seen ondedicated PET. This value equates to 83% inthe thorax but only 67% for subphreniclesions. In the 10 lesions missed by hybridPET 8 were <1.5 cm and 2 cm were deep inthe abdomen. Interestingly the resolution

specifications of the two scanners showed thehybrid scanner had 5 mm and dedicatedscanner 6.5 mm, while sensitivity was afactor of 12:1 in favor of dedicated PET.

These results are reflected in a paperby Shreve et al where coincidence gammacamera PET detected 55% (66/105) oflesions detected by dedicated PET in a widerange of malignancy types and locations.32 Ofmediastinal lesions >1.5 cm in diameter itdetected 94% (15/16), but only detected 30%of these lesions where the diameter was lessthan 1.5cm. The hybrid camera performedwell in other areas apart from the abdomenwhere it only detected 23% (6/26) of lesions.

Hybrid scanners usually operate in anopen 3D mode and are subject to the effectsof out of field activity. Resolutionmeasurements are similar to dedicated PETscanners but hybrid scanners suffer fromlower sensitivity and higher scatter fractions.Low detector sensitivity is exacerbated byincomplete or partial geometry.

Design ConsiderationsOne of the main restrictions to the use

of gamma cameras for PET has beenhandling the very high count-rateexperienced following the removal ofcollimators. Typically, the detector count-ratefor PET is 50 times higher than that forroutine SPECT.

Higher count-rate capability isachieved by the techniques described inpreviously. These include pulse clipping ofthe PMT signal, and calculation of theposition of detected events that are restrictedto local PMTs. These two innovations allowdetector count rates of greater than 500Kcpsto be achieved and were originally adaptedfrom the PennPET NaI(Tl) scanner into thefirst commercially available coincidencegamma camera.33 Other manufacturers haveintroduced various innovations to process thehigh count-rates arising from detectors andPMTs.

27

There are also issues related to theability of a crystal whose thickness isoptimized for lower energy photons to detect511 keV photons in coincidence with anyefficiency. Increasing crystal thicknessimproves efficiency in PET mode butdegrades the intrinsic resolution (andtherefore system resolution) in SPECTmode. As PET relies on the detection ofevents in coincidence improving theefficiency of a single detector has a squareeffect on coincidence efficiency. Doublingefficiency in singles photon mode has aquadrupling in coincidence efficiency.

As stated in previously, systemresolution is defined by intrinsic resolutionand collimator resolution. Degradation inintrinsic resolution due to increasing crystalthickness has a small effect on systemresolution, but many argue as to what istolerable.

Future Developments in Hybrid ScannersLSO. There has been a degree of

excitement amongst the nuclear medicinecommunity with the possible use of LSO inhybrid gamma cameras. Its superior stoppingpower and fast light photon properties wouldmake it an excellent replacement for NaI(Tl)when operating in PET mode acquisitions. Ithas superior sensitivity to NaI(Tl) at half thecrystal thickness.

A problem with LSO is that it isinherently radioactive (containing naturallyoccurring Lu-176), with emissions in theSPECT photon range.34 To counter thisproblem a thin layer of a second type ofcrystal suited to SPECT (YSO or NaI(Tl))could be optically coupled to the front of anLSO crystal. This veneer of crystal isdesigned to detect lower energy SPECTphotons, while the majority of 511 keVphotons will penetrate this layer to bedetected by the LSO layer. In SPECT mode,scintillations in the front crystal can becharacterized and differentiated from the

scintillations occurring because of photonsemitted from within LSO due toradioactivity. When operating in PET modethe front layer would be penetrated by thevast majority of 511 keV photons, leavingthem to be absorbed by the LSO layer.

A factor in the success of hybridgamma cameras is economics. LSO is bycomparison much more expensive thanNaI(Tl) (see Table 3(m)) and this will impactgreatly on the overall cost of the scanner. InSPECT mode, especially for photons withenergy <200 keV, there is unlikely to besufficient improvement in image quality tojustify the increased cost over NaI(Tl).

Slotted Integral Crystals. Poorsensitivity due to poor detector efficiency isone of the major limitations of hybridsystems in PET mode. Increased crystalthickness improves the efficiency ofdetecting 511 keV photons, but thickercrystals degrade intrinsic resolution inSPECT mode. Resolution degradation resultsfrom greater spread of light photonsfollowing scintillation. Spreading of lightover a greater area decreases the accuracywith which the origins of that light, photonabsorption and scintillation, can becalculated. Crystals for SPECT systems aregenerally 3/8'' to 1/2'' thick.

As stated in previously, doubling theefficiency of detecting single events resultsin a four-fold improvement in detectingcoincidences.

A way of improving the resolutioncharacteristics of thicker crystals is tocrisscross the PMT surface with thin slots ofa depth approximately half the total crystalthickness (Fig.14). The slots, which areapproximately 1/2'' apart, are designed toreduce the radial spread of light followingscintillation, and channel the lights towardsthe PMTs.

An additional benefit is animprovement in image quality whenscanning high-energy photons in singlephoton mode (e.g., I-131).

28

CONCLUSIONS AND THE FUTUREMany readers of this manuscript will

have noticed the current dominance of PETin this topic. Despite its age, PET isundergoing radical change in almost everyaspect of the modality, from the basicdetector through to clinical applications.Future developments in instrumentation willfocus on new scintillation materials, theimplementation of depth of interactionanalysis, and improved energy resolution for3D mode of acquisition. This should allowcameras of exceptional resolution (2mm) toemerge at the high end of the market,potentially married to CT or even MRIscanners. Other markets will ensure thathybrid PET/SPECT camera technologycontinues to develop.

New detector materials and designsbeing investigate for hybrid PET/SPECTcameras may produce spin-offs that aremanufactured into, and improve conventionalgamma cameras. Due to collimation, it isunlikely that any improvement in the intrinsic

performance of the detector will significantlyimpact on system performance.

Single photon development continuesstrongly. Increasingly attenuation correctionsystems that use transmission sources will beused clinically. Multi-detector camerascontinue to be popular as a way of improvingcamera sensitivity. Reconstruction tech-niques are becoming increasingly complex,facilitated by increasing computer power.Analysis of results, especially in the area ofgated myocardial perfusion imaging has alsobenefited from this power.

Many departments are now integratedinto hospital wide area networks and“intranets” that allow images, and not juststudy reports, to be available to referringpersonnel.

AcknowledgementsI would like to thank Ms. Ailsa Cowie

for extensive manuscript revision. Thanksalso to Dale Bailey of the MRC CyclotronUnit in Hammersmith UK for his illustrationused in Figure 11.

29

Figure 14. Slotted NaI(Tl) crystal showing reduced light spread compared to unslotted crystal of equivalent thickness, and hence improved intrinsic resolution.

A. Top view of crystal B. Slotted 1” crystal C. Plain 1” crystal with greater light

spread, and equivalent sensitivity. D. Medium crystal (5/8”) with less

light spread but less sensitivity to high-energy photons.

E. Thin crystal (3/8”) with even less light spread and therefore excellent resolution, but poor sensitivity to medium and high-energy photons.

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31

QUESTIONS

1. Which one of the following statements isincorrect?

a. The basic design of the NaI detectorhas remained unchanged for the last10 years.

b. Increased computing power isallowing more complexreconstruction and analysis innuclear medicine and PET.

c. Measured attenuation correction isgaining clinical acceptance innuclear medicine.

d. Multiheaded gamma cameras wereintroduced to reduce radio-pharmaceutical costs.

2. Photons lose a portion of their energywhen they undergo Compton scatter.Which one of the following options iscorrect? Scattered photons are rejectedduring acquisition by which of thefollowing processes?

a. Timing delayb. Energy window discriminationc. The Doppler effectd. Coincidence windowing

3. Which of the following physicalproperties of scintillation detectorseffects energy resolution most?

a. High stopping powerb. High photon yieldc. Refractive indexd. Density

4. A photon with energy of 140keVundergoes Compton scatter and itsdirection is changed by 35 degrees. Thephoton’s post collision energy will be:

a. 103 keVb. 113 keVc. 123 keVd. 133 keV

5. Typically, only around 1 in 10000photons from the patient will reach thegamma camera detector. Which one of

the following best explains thisinefficiency?

a. Only a small section of the patient isin the field of view of the detector.

b. Attenuation of photons by the body.c. Decay of the isotope between

administration and imaging.d. Attenuation by the collimator.

6. When comparing the system sensitivityof PET and SPECT which one of thefollowing is correct?

a. PET has greater sensitivity becauseit doesn’t need collimation.

b. PET has greater sensitivity purelybecause for each decay 2 photons areemitted.

c. PET has greater sensitivity becausethe higher photon energy means lessattenuation.

d. PET has greater sensitivity becauseit has 360 degrees of detectorsaround the patient.

7. A gamma camera has detector (intrinsic)resolution of 4.0 mm and collimatorresolution of 10.0mm. If detectorresolution was reduced to 2.0mm, thensystem resolution would change from:

a. 14 mm to 12 mmb. 40 mm to 20 mmc. 12.2 mm to 11.8 mmd. 10.8 mm to 10.2 mm

8. Sodium Iodide NaI(Tl) has virtually beenthe exclusive choice for scintillationdetectors in gamma cameras for 30 years.Which of the following is incorrect?

a. NaI(Tl) has excellent stopping powerwhich gives it high sensitivity tomedium and high energy photons.

b. NaI(Tl) has a good light photonyield which gives it good energyresolution.

c. Despite being hygroscopic, it can bemanufactured relatively cheaply.

d. NaI(Tl) has a good light photonyield which allows it to bemanufactured into a large crystal andcoupled to multiple PMTs.

32

9. Which one of the following statementsrelating to BGO PET detectors isincorrect?

a. BGO has relatively poor energyresolution.

b. BGO has excellent photon stoppingpower.

c. BGO is manufactured into blockdetectors rather than large singledetectors.

d. BGO is a likely replacement for NaIas a gamma camera detector.

10. Despite having many superior physicalproperties to NaI, LSO is unlikely tobecome a material of choice in SPECTgamma cameras. Which one of thefollowing explanations is incorrect?

a. NaI has better stopping power.b. LSO is very expensive.c. The effect of collimation on extrinsic

resolution dilutes any improvementin intrinsic resolution.

d. LSO is inherently radioactive.

11. Select the correct statement. Semicon-ductor detectors:

a. Have worse energy resolution thanscintillator/PMT type detectors.

b. Will be much heavier and larger thanscintillator/PMT type detectors of acomparable FOV.

c. Are constructed in a matrix ofindividual detectors, with each cellbeing the minimum resolvable unit.

d. Are used in most current gammacameras.

12. Which statement relating to filtered backprojection tomographic reconstruction isincorrect?

a. Can produce serious artifacts in lowcount studies, in particular, ‘starring’of counts from ‘hot’ structures.

b. The ‘Ramp’ filter corrects for the blur-ring effect caused by statistical noise.

c. Has only limited capability ofcorrecting for attenuation and scatter.

d. It is the most commonly usedreconstructive algorithm, even today.

13. Which statement relating to iterativetomographic reconstruction is incorrect?

a. It has become increasingly utilizedwith the availability of fastercomputing.

b. Iterative techniques achieve recon-struction by comparing acquired datato what has been reconstructed in theprevious iteration. The reconstructeddata is modified to minimize anydifference.

c. Iterative reconstruction should onlybe used in ‘high count’ studies.

d. The OSEM iterative technique isable to achieve reconstruction at anaccelerated rate compared toconventional iterative reconstruction.

14. What is the sensitivity of a collimator(with hexagonal holes) with thefollowing properties ______?

Hole length (L) = 35 mmSeptal thickness (s) = 0.2 mmHole diameter (d) = 2.2 mm

a. One photon per 4b. One photon per 43c. One photon per 436d. One photon per 4365

15. Using the collimator properties definedin Question 7, what would be its resolu-tion at a scanning depth of 100mm?

a. 7.5 mmb. 8.5 mmc. 9.5 mmd. 10.5 mm

16. Which one of the following statementsregarding fan-beam collimators isincorrect?

a. SPECT studies can be reconstructedwith the same algorithm used forparallel hole collimators.

b. They are parallel in the axial plane.c. They are convergent in the transaxial

plane.d. Their FOV is smaller than parallel

hole collimators and therefore areonly suitable for neurology andpediatric SPECT.

33

17. Which one of the following statementsregarding cone-beam collimators isincorrect?

a. The holes in cone-beam collimatorsall converge to a focal point.

b. Reconstruction of SPECT studies iscomplex due to the data having beenacquired obliquely through transaxialplanes.

c. Increased sensitivity is due to thefocused geometry of the collimator.

d. The collimator is suitable for allSPECT studies.

18. Multi-detector gamma cameras havesome advantages over single detectorsystems. Which of the followingstatements is false?

a. Multi-detector gamma cameras canreduce acquisition times.

b. Multi-detector gamma cameras canimprove resolution.

c. Multi-detector gamma cameras canreduce attenuation.

d. Multi-detector gamma cameras canimprove sensitivity.

19. Which of the following statementspertaining to SPECT attenuationcorrection is false?

a. First order attenuation correctionassumes homogeneous attenuationwithin the body and thereforeachieves good correction ofmyocardial studies.

b. A transmission source should beselected with a photon whose energyis similar to that of the emissionphoton so that accuratedetermination of attenuationcoefficients can be achieved.

c. A transmission source should bewell shielded when in use.

d. Moving transmission sources, eitheras point or line sources, allowregional concentration of photons,thus reducing the percentage ofcontamination from emissionphotons.

20. Which one of the following does noteffect the resolution of a PET scanner:

a. Positron rangeb. Uptake timec. Annihilation photon non-colinearityd. Detector element size

21. Which of the following statements isincorrect? “Positron range”:

a. Is the distance a PET camera needsto be to a cyclotron to make ifeconomically viable.

b. Is the average distance a positrontravels from the nucleus before itannihilates with an electron.

c. Degrades PET resolution.d. Is related to the Emax of the isotope

used.

22. Which of the following statements isincorrect? In PET scanning, “Anni-hilation photon non-colinearity”:

a. Results in reduced sensitivity.b. Degrades PET resolution by about 2

mm in a human scanner.c. Is caused by the annihilation photons

not traveling in exactly oppositedirections.

d. Has greatest effect when thediameter of PET detectors isincreases.

23. Which one of the following statements iscorrect? In PET scanning, “Depth ofInteraction” analysis:

a. Is the calculation of the point withinthe patient where radioactive decayoccurs.

b. Is used to determine coincidencetiming in ‘time-of-flight’ scanners.

c. Helps to correct for detector parallaxerrors in coincidences occurringaway from the center of the FOV.

d. Determines the position of a rotatingtransmission source.

34

24. Which one of the following scintillationdetector materials has the best stoppingpower for 511 keV photons?

a. NaI(Tl)b. BGOc. LSOd. GSO

25. Which one of the following scintillationdetector materials has the best photonyield?

a. NaI(Tl)b. BGOc. LSOd. GSO

26. Which one of the following scintillationdetector materials has the best photondecay constant for PET?

a. NaI(Tl)b. BGOc. LSOd. GSO

27. Which combination of properties belowwould you select as belonging to an idealPET detector?

a. high stopping power, low photonyield and long photon decayconstant.

b. high stopping power, high photonyield and long photon decayconstant.

c. low stopping power, high photonyield and short photon decayconstant.

d. high stopping power, high photonyield and short photon decayconstant.

28. Using the combined properties selectedin the questions 27, which of thefollowing materials would make the bestPET detector?

a. NaI(Tl)b. BGOc. LSOd. YSOe. BaF2

29. Which of the following statements isincorrect? When acquiring PET studiesin 3D mode

a. There is an increased scatter fractionwhen compared to 2D acquisitions.

b. Sensitivity increases towards thecenter of the axial FOV.

c. The detectors are not shielded fromout-of-field activity.

d. There is approximately a 5 folddecrease in sensitivity compared to2D acquisitions.

30. Which one of the following statementsconcerning attenuation in PET isincorrect?

a. Attenuation of 511keV photons bysoft tissue negligible and thereforecorrection only needs to beperformed in large patients.

b. The attenuation effect in PET isconsiderable despite the use of511keV photons due to bothannihilation photons needing to bedetected to form a coincidence.

c. The effect of attenuation in PET isapproximately equivalent to thatsuffered by 201Tl imaging usingSPECT.

d. The effects of attenuation is greatesttowards the middle of the body.

31. Select the one incorrect statement. Post-injection PET attenuation correction:

a. Allows the transmission scan to beacquired after the patient has beeninjected.

b. Are achieved by differentiatingbetween photons arising from thepatient and the transmission source.

c. Improves throughput of the scanner.d. Can only be performed before the

emission study.

35

32. Select the one incorrect statement below.Singles PET attenuation correction:

a. Allows a transmission source ofincreased activity to be employedthus reducing acquisition time.

b. Separation of emission andtransmission gamma rays is achievedby energy discrimination.

c. Longer half-life sources can beutilized thus reducing frequency andcost of replacement.

d. Cannot be acquired once a patienthas been administered activity.

33. Select the one correct answer below.Image segmentation of transmission data:

a. Allows for more accurate calculationof attenuation coefficients.

b. Removes noise from areas ofequivalent attenuation in thetransmission image by applying auniform attenuation coefficientvalue.

c. Corrects for subtle differences inattenuation coefficients in thepatient.

d. Corrects for the effects of scattercontribution.

34. Which of the following statementsrelating to the Penn-PET scanner withSodium Iodide detectors is incorrect?

a. Detector dead time is reduced bylogically dividing each detector intoa number of smaller “virtual”detectors.

b. Detector dead time is reduced byusing only a central PMT and thesurrounding 6 PMTs to determinethe position of a scintillation event.

c. Detector dead time is reduced byclipping the output signal fromindividual PMTs after a finite periodof time.

d. Detector dead time is reduced byusing large integral detectors.

35. Which one of the following statements isincorrect concerning “Partial ring BGOPET scanners”?

a. The scanner only operates in 3Dacquisition mode.

b. The scanner has twice the sensitivityof a full ring BGO PET scanner.

c. Manufacturing costs are reduced byusing approximately half the numberof detectors arranged in twoopposing banks.

d. Manufacturing costs are reduced byusing detectors with reduced depth.

36. Which one of the following statements isincorrect concerning Hybrid PET/SPECT gamma cameras?

a. The high count-rates experiencedwhen imaging in PET mode can benegated by installing collimators.

b. Increasing crystal thickness toimprove sensitivity to 511 keVphotons reduces intrinsic resolution.

c. Open detector geometry results in alarge contribution from out-of-fieldactivity.

d. The camera has relatively poorsensitivity due to partial geometryand the use of detectors whosethickness is not optimal for 511 keV.

37. Which one of the following statements isincorrect concerning Hybrid PET/SPECTgamma cameras?

a. They are not suitable for short half-life isotope studies because of thevery high transient count-ratesexperienced.

b. They are economically attractive dueto small capital cost compared todedicated PET cameras.

c. They are economically attractive asthey can be used as a gamma camerawhen isotope may not be available.

d. They have equivalent performancecharacteristics compared todedicated PET cameras.

36

38. When selecting a scintillation detectormaterial for “time-of-flight” PETscanning the most important physicalproperty is:

a. Costb. Densityc. Light photon decay constantd. Stopping power

39. Which of the following statementsregarding combined PET/CT scanners isincorrect?

a. Attenuation correction of the PETemission data can be calculated fromthe CT data.

b. Improved diagnostic accuracy maybe achieved by precise co-registration of structure (CT) andfunction (PET).

c. Patients being scanned will need tohold their breath between scans.

d. PET and CT scans are acquired atseparate times in the one session.

40. The use of slotted NaI(Tl) detectorsimprove efficiency in PET mode andresolution in SPECT mode of hybridgamma cameras. Which statement belowbest describes how this is achieved?

a. The slots in the crystal stop 511keVphotons scattering throughout thecrystal.

b. The slots allow gamma photonstraveling perpendicular to thedetector surface to reach the rear ofthe crystal.

c. Resolution is maintained in a thickercrystal by reducing the radial spreadof light photons within the crystalvia the use of slots.

d. Higher sensitivity in PET mode isachieved by the crystal havingvariable thickness.

37