Technology challenges in small animal PET imaging

ARTICLE IN PRESS

Nuclear Instruments and Methods in Physics Research A 527 (2004) 157–165

*Corresp

brooke, Q

+819-820-6

E-mail a

(R. Lecomt

0168-9002/$

doi:10.1016

Technology challenges in small animal PET imaging

Roger Lecomtea,b,*aUniversit!e de Sherbrooke, Sherbrooke, Qu!e., Canada

bMetabolic and Functional Imaging Centre, CRC/CHUS, Sherbrooke, Qu!e., Canada

Abstract

Positron Emission Tomography (PET) is a non-invasive nuclear imaging modality allowing biochemical processes to

be investigated in vivo with sensitivity in the picomolar range. For this reason, PET has the potential to play a major

role in the emerging field of molecular imaging by enabling the study of molecular pathways and genetic processes in

living animals non-invasively. The challenge is to obtain a spatial resolution that is appropriate for rat and mouse

imaging, the preferred animal models for research in biology, while achieving a sensitivity adequate for real-time

measurement of rapid dynamic processes in vivo without violating tracer kinetic principles. An overview of the current

state of development of dedicated small animal PET scanners is given, and selected applications are reported and

discussed with respect to performance and significance to research in biology.

r 2004 Elsevier B.V. All rights reserved.

PACS: 87.57.�s; 87.62.+n; 87.58.Fg

Keywords: Positron emission tomography; Small animal imaging; Detector technology; Scanner performance

1. Introduction

Positron Emission Tomography (PET) is apowerful tool for research into molecular path-ways and genetic processes with animal models ofhuman diseases. Being non-invasive, PET can beused repeatedly in the same animal to monitor theprogression of disease and to investigate theresponse to therapy under controlled conditions.Imaging rats and mice with PET is a challengingtechnological undertaking, due to the high resolu-tion and high sensitivity that are required for

onding author. Universit!e de Sherbrooke, Sher-

u!e., Canada. Tel.: +819-346-1110X14608; fax:

490.

ddress: [email protected]

e).

- see front matter r 2004 Elsevier B.V. All rights reserve

/j.nima.2004.03.113

visualizing complex functions taking place intotiny tissue structures. The logistics of quantitativefunctional imaging, aiming to extract basic bio-chemical and pharmacokinetic data in a reprodu-cible way, also imposes special requirements thatare not common to clinical PET imaging.

2. Technical requirements of small animal PET

Clinical PET imaging in a 70 kg human subjecttypically requires 5–15mCi of activity andachieves a spatial resolution of 6–10mm (B1mlvolumetric), providing sufficient details for thestudy of the brain and heart function [1], and forthe detection and grading of tumors [2]. With thefirst generation dedicated small animal PET

d.

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R. Lecomte / Nuclear Instruments and Methods in Physics Research A 527 (2004) 157–165158

scanners reaching B2mm full-width-at-half-maxi-mum (FWHM) resolution [3–14], it is possible tomeasure regional radiotracer uptake in the mainorgans of B300g rats [12–22]. In cardiac studies,using similar injected activity as currently used inhumans, images of nearly comparable definition andquality can be obtained (see Fig. 1) [12,20–22]. Thisrepresents a gain of 4–5 in linear resolution, orB100 in volumetric resolution, relative to clinicalscanners. In 20–30g mice, radiotracer biodistribu-tion and global organ dynamic uptake can bemeasured [14,23–27] (see Fig. 2), but a further gainof at least 2 in linear resolution, or about 10 involumetric resolution, that is, B1mm FWHM or1ml, would be required to achieve the same imagedefinition. At the same time, the tissue radioactivityconcentration would have to be increased or the

Fig. 1. 18FDG PET images of normal (top) and diseased

(bottom) rat heart obtained with the Sherbrooke small animal

PET scanner. Images reoriented along the heart main axes:

horizontal long axis (HLA: left), short axis (SA: center), vertical

long axis (VLA: right). From [1].

Fig. 2. 18FDG PET image of a 22 g mouse heart obtained with

the Sherbrooke small animal PET scanner. Images as measured

along the coronal (left), sagittal (center) and transaxial planes

(right).

scanner sensitivity improved grossly by the samevolumetric factor, in order to achieve a similarquantitative accuracy per image voxel [28].

2.1. Spatial resolution

The reconstructed FWHM image resolution inPET can be nicely described by an empiricalformula proposed by Moses and Derenzo [29,30]:

FWHM ¼ a

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiðd=2Þ2 þ b2 þ r2 þ ð0:0022DÞ2

qð1Þ

where d is the detector size, b the position decodingaccuracy, r the positron range and D the detectordistance, used here to describe the annihilationphoton noncollinearity (0.5D tan 0.25�=0.0022D).The factor 1.1oa o1.3 depends on the algorithmused to reconstruct the image and typically takes avalue of a=1.2 when using filtered backprojectionreconstruction with a ramp filter.The first two terms under the square root form

what is referred to as the intrinsic detectorresolution, consisting of the geometric and decod-

ing resolution factors. The intrinsic resolution iscommonly measured by moving a thin annihila-tion line source between opposite coincidentdetectors separated by a short distance, while thedecoding resolution can be estimated from thehistogram obtained by directly irradiatingthe detector array. The positron range in tissuehas an ‘‘inverse cusp’’ like distribution consistingof a narrow spike and long exponential tails.Whereas the actual FWHM is very small [31], theeffective FWHM taking into account the broad-ening due to the entire range distributionr=2.35 rms provides a more realistic estimate ofthe image blurring resulting from positron range.Table 1 reports values for the most commonpositron emitters. The contribution from noncol-linearity to resolution is significant (2.1mm) inclinical whole-body scanner having a diameterDE80 cm, but is almost negligible (p0.33mm) ina small dedicated scanner for rats and miceimaging (Dp15 cm).The published resolution data of several existing

tomographs have been analyzed to remove theeffects of noncollinearity and positron range(or effective source size used for resolution

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R. Lecomte / Nuclear Instruments and Methods in Physics Research A 527 (2004) 157–165 159

measurements) and were plotted as a function ofcrystal size (Fig. 3). The scanners can be groupedinto three different categories, depending on thetype of coupling between scintillators and photo-detectors and on the positioning algorithm used toidentify the crystal of interaction:

* The first category, mostly consisting of ‘‘block’’type detectors [32], uses light-sharing andAnger-like logic to decode the crystal ofinteraction [3–5,33,34]. The positioning accu-racy calculated as the quadratic differencebetween the measured intrinsic resolution andthe geometric resolution (dashed line in Fig. 2)amounts to bE2mm on the average.

Table 1

Positron range data for common positron emitters

Isotope Max b+

energy (MeV)

FWHM (mm) Effective

FWHMa

(mm)

11C 0.96 0.13 0.9213N 1.2 0.17 1.3515O 1.7 0.28 2.418F 0.64 0.13 0.5464Cu 0.65 0.13 0.5468Ga 1.9 0.31 2.882Rb 3.4 0.42 6.1

aEffective FWHM: r=2.35 rms.

Fig. 3. Intrinsic spatial resolution of several existing tomographs as a

the type of position decoding scheme used to identify the crystal of i

* The second category uses a direct coupling ofcrystals to the photodetectors and attempts toavoid as much as possible the spread ofscintillation light between crystals [7–10,35–37].Position decoding is performed using a chargedivision readout scheme of the signals fromposition-sensitive or multiple-anode PMTs. TheHIDAC system [11], which uses lead converterswith 0.4mm diameter holes to detect photonsand thin wires to sense the charge can beclassified into this category. On the average,the decoding resolution is consistently estimatedto bE1mm for these systems.

* The third category uses true ‘‘pixel’’ detectors(crystals individually coupled to photodetectorsand independent electronic signal processingchannels) and achieves an intrinsic resolutionthat is exactly equal to the geometric resolution(bE0mm) [6,11,38,39].

These data suggest that there is some inherentloss of positioning accuracy associated with alltypes of decoding schemes. The positioningaccuracy can be improved by limiting the lightexchange between individual crystals, but chargespread within the multiplying structure of thePMT and the readout circuit still adds someuncertainty in crystal identification. The othersource of positioning error that cannot be avoided

function of detector size. Data have been labelled according to

nteraction.

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0.0001

0.001

0.01

0.1

1

0 8 16 24 32 40 48 56 64

Projection bin

Rel

ativ

e in

ten

sity

MeasuredDetector ResponseDetector ScatterObject Scatter

BGO 3mmSepta 0.8 mmObjet 11 cm� E

th =344 keV

~ 40%

~10%

Fig. 4. Relative contributions from the detector geometric

response, the detector scatter and object scatter for a line source

at the center of an 11-cm diameter tissue-equivalent object,

when measured with arrays of 3-mm wide individual BGO

detectors.


with position decoding schemes is the occurrenceof multiple interaction events depositing a fractionof the incident photon energy in more than onecrystal. The event coordinates are then inaccu-rately decoded at an intermediate position betweenthe crystals where the interactions have occurred.While such multiple interaction events also lead tomispositioning in pixel detectors, the resultingdegradation of FWHM resolution is practicallynegligible due to the exponential-shaped distribu-tion of such scatter events (see Fig. 4 below), andthe effect can potentially be minimized by sophis-ticated multiple event processing [38,40].Assuming that the positron range and annihila-

tion photon non-collinearity can be minimized insmall animal PET by using radioisotopes with lowenergy positrons (e.g., 18F, 64Cu) and a smalldiameter ring geometry, the image resolutiondegradation would still amount to

1:2

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi0:542 þ 0:0022� 150ð Þ2

qE0:7 mm:

Hence, to achieve a reconstructed image resolutionof 1.0mm, an intrinsic detector resolution con-tribution not exceeding 0.7mm would also berequired. From the data shown in Fig. 3, it wouldappear that only pixel detectors with a size of lessthan 1.2mm would have the ability to provide anoverall spatial resolution of 1mm in reconstructedimages. However, it can be assumed that thepositioning accuracy scales down with the size of

individual crystals and can be improved by specialcrystal surface treatment and assembly. Recentresults obtained by Correia et al. [41] and Miyaokaet al. [42] have shown that an image resolution of1mm is achievable with 0.8mm wide crystalassembled with cleverly designed crystal interfaceswhich reduce the positioning accuracy tobB0.4mm.

2.2. Parallax error

Another serious problem that arises in high-resolution systems designed for small animal ima-ging is the parallax error due to penetration of the511keV in the detectors. One could compensate forthis effect by redefining an effective ring diameter inthe image reconstruction algorithm to take intoaccount the average interaction depth in thedetector. However, this correction is inaccurate.One simple way to minimize the effect is to use alarge system diameter relative to the imaging field-of-view (FOV), but the reduced sensitivity, degradedresolution due to non-collinearity and high costmake this solution unacceptable. It is generally morefavorable to trade off some detection efficiency withshorter detectors in a smaller size system [8,12,43].Alternatively, the parallax error can be somewhatattenuated and its effect spread over the entire FOVby using planar detector arrays rather than the usualcylindrical geometry [34].Several different solutions have been explored to

measure the depth-of-interaction (DOI) withindetectors in order to correct for the parallax error:pulse height discrimination (PHD) [44], depthencoding by light sharing [45,46], signal ratio[47,48], pulse shape discrimination (PSD) [49]. Allhave attractive features and inherent limitations.For instance, the PHD technique is simple toimplement but sensitive to scatter; the light codingrequires delicate crystal cutting and preparationthat is difficult to implement in large scaleproduction; the signal ratio method provides acontinuous DOI measurement and is insensitive toscatter, but is expensive because two photodetec-tor readouts per crystal are needed; and the PSDapproach must use scintillators with differentdecay times, whose other characteristics may notbe optimal for PET. Nevertheless, the PSD

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Table 2

Detection efficiency parameters at 511 keV for common

scintillators

NaI BGO GSO LSO LuAP

Decay time (ns) 230 300 60 40 18

Density (g/cm3) 3.67 7.13 6.71 7.35 8.34

Effective Z 50 73 58 65 65

1/m at 511 keV (mm) 25.9 11.2 15.0 12.3 11.0

PE (%)a 18 44 26 34 32

PE 2 (%)b 3.2 19 6.8 12 10

aPhotoelectric absorption probability:

PE = 100�sp/(sp+sc).bCoincident photoelectric absorption probability.


technique has been demonstrated with PMT andAPD readouts using up to four different crystals[50–53], and was already implemented in a fewsystems [13,34,54,55].

2.3. Sensitivity

Sensitivity in small animal PET is important toenable fast dynamic image acquisition with short(Bsec) time frames in pharmacokinetic studies. Itis also crucial to overcome biological limitations inreceptor studies with high-affinity, low-specificactivity radiolabelled ligands [56–58]. Scannersensitivity in PET can be reported as the pointsource coincidence detection efficiency at thecenter of the FOV (in units of cps/mCi or cps/MBq) or as the scanner count rate per unitradioactivity concentration (cps/mCi/cc or cps/MBq/cc) in standard flood sources, representativeof the subjects to be imaged. However, thesemeasurements provide no information on thesignal-to-noise ratio that can be achieved inimages, nor on the maximum event rate thescanner can process. A more meaningful para-meter is the effective sensitivity given by the noiseequivalent count rate [59]

NEC ¼T2

T þ S þ aR¼ T �

T

T þ S þ aRð2Þ

which is the rate of true coincidences T weightedby the ratio of trues to total events measured bythe scanner, including scatter (S) and random (R)event rates (1oao2 depends on random correc-tion). As the registered event rates are affected bydead time in the detector and processing electro-nics, so is the NEC rate which will eventually reacha maximum value that is a characteristics of thescanner for a given source geometry and distribu-tion. The maximum NEC rate should occur at apoint that is well beyond the typical tissue activityconcentrations encountered in small animal ex-periments, which can reach a few tens mCi/cc(B1MBq/cc) in activation and blood flow studies.This is best achieved with highly pixelateddetectors that restrict dead time solely to thedetector channel being activated.One factor of great importance that is

often overlooked in high-resolution PET is the

probability of Compton scatter in the detectors. Itseffect on spatial resolution and sensitivity can beappraised from Fig. 4, where the relative contribu-tions of geometric detector response, detectorscatter and object scatter distributions wereisolated for 3mm wide BGO scintillators [60]. Itis worth noting that the detector scatter fraction,which is already substantial with 3mm BGO, willdramatically increase with smaller size crystals andlower Z materials, where the photoelectricinteraction probability can be significantly lower.Table 2 summarizes the efficiency parameters forthe most popular scintillators. The individualdetector photoelectric fraction, which turns outto be approximately the same as the photoelectricabsorption probability (PE) in the tiny crystalsbeing considered for the next generation smallanimal PET scanners [36,41,61,62], becomes oneof the most critical parameter for sensitivity, sinceit appears squared in the trues rate and to thefourth power in NEC rate. It is particularlydisturbing to observe that the coincident photo-electric absorption probability on first interactionin the faster lutetium based and GSO scintillatorsis only slightly more than one half to one third thatof BGO, respectively. Obviously, the photoelectricfraction is effectively increased in block detectorswith position decoding readout schemes, but at theexpense of some inherent loss of resolution, as wehave seen.In order to maximize NEC rate, it is equally

important to minimize the contributions fromobject scatter and random coincidence events.

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0.001

0.010

0.100

1.000

10.000

-60 -50 -40 -30 -20 -10 10 20 30 40 50 600

Axial P osition ( m m )

Fra

ctio

n (%

)

Scatter without collimator

Scatter with collimator

Random without collimator

Random with collimator

FieldCollimator

350 keV

Fig. 5. Comparison of the scatter and random probabilities for

a point source in an 11-cm diameter tissue-equivalent object

moved along the axis of the Sherbrooke animal PET scanner

with and without the field collimators in place. The field

collimators in this system extend 87.5mm inside the 310-mm

diameter ring of detectors. The asymmetry on the right-hand

side of the peak with the field collimators in place is due to a

notch at the end of the collimator to accommodate a rotating

line source [63].


The scatter and random fractions from within theimaging FOV are expected to be low, consideringthe small size of rats and mice relative to the meanfree path of 511 keV photons in tissue and therelatively small amount of radioactivity ending upwithin the target organ in most studies. Therefore,in a well-collimated 2-D system where the detectorsolid angle can be tightly limited to the imagingFOV, it is possible to achieve fairly high NECrates in spite of very low detection efficiency.However, in the absence of adequate shielding, theprobability of scatter and random events fromactivity outside the FOV might be appreciable.Fig. 5 illustrates the dramatic increase, by an orderof magnitude or more, of both the scatter andrandom probabilities that would result in theabsence of field collimators in the Sherbrookeanimal PET scanner [63]. The situation would beeven worse without thick annular lead shieldspreventing direct exposure of the detectors to theactivity outside the FOV. These data stress theimportance of proper shielding against activityoutside the FOV for maximizing image signal-to-noise ratio.Minimizing scatter and random contributions in

a large axial FOV system with a compact geometry

is a difficult task involving complex tradeoffs,which may dramatically affect the design, perfor-mance and cost of the scanner. There is nowsufficient published data on small animal scannerswith various geometries to draw some generalconclusions on the design criteria to be favored forspecific applications (see Table 3). Notwithstand-ing the differences in NEC measurements andother system characteristics, it appears that highNEC rates, and hence high image signal-to-noiseratio, is not necessarily correlated with systemabsolute detection efficiency. In situations wheresensitivity is the most critical issue, such as is thecase in receptor binding studies with high affinityligands, a large axial FOV, 3D imaging system isclearly the best choice. On the other hand, inkinetic studies with non-specific tracers where alarge fraction of the injected activity is circulatingand eventually ends up in non-target organs, awell-collimated, reduced axial FOV geometry hasdefinite advantages to enhance image signal-to-noise in the region-of-interest. In either case, thereare simple measures that can and should beapplied to minimize contamination from out-of-field activity, such as the use of a properly shieldedseringe, masking of the point of injection, andspecial shielding of hot non-target organs, such asthe bladder, kidneys or liver whenever possible.Using high detection efficiency, low dead time

detectors in an efficient 3D geometry (small systemdiameter, large axial FOV) increases the trues rateT and NEC. Similarly, using fast detectorsefficiently shielded against out-of-field activity willdecrease S and R and improve NEC.

3. Methodological issues in small animal PET

The logistics of small animal PET is basicallythe same as for clinical PET. Therefore, it requiressimilar facilities in terms of staff support, access toradiotracers and image analysis capability. Still,there are a number of practical considerations thatmust be addressed to fully exploit the potential ofsmall animal PET imaging for biological studies:anesthesia, physiological and pharmacologicalconstraints on radiotracer administration, mea-surement of input function, subject repositioning

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Table 3

Comparison of detection efficiency and peak NEC rates for different small animal PET scanners

Scanner FOV (cm+� cm) Source (cm+� cm) Efficiency (cps/mCi) Peak NEC (kcps) Ref

Sherbrooke 12� 1.05 11� 3.5 200 61 [6]

MicroPET/UCLA 11.25� 1.8 6� 4 208 18 [8]

MicroPET/P4 (Concorde) 19� 7.8 5.1� 10.8 529 85 [64]

3.7� 6.8 290

RatPET 2D 11.5� 5 5.2� 5 736 10 [5]

3D 1600 41 [65]

A-PET 12.8� 12.8 5.1� 10.9 851 300 [66]

3.7� 6.8 400

HIDAC 10� 21 6� 6 480 50 [11]


for serial studies, etc. These issues have beendiscussed in recent papers [14,56,67–70]. However,relatively little attention has been given to thespecific requirements on scanner design andoperation that must be satisfied in implementingcomplex biological imaging protocols. Examplesof such practical issues relate to the acquisitionof kinetic data, physiologic gating, the monitoringof the animal’s vital signs and the registration ofthese physiologic data with image data, and thefeasibility of intervention during imaging. Flexible,programmable acquisition protocols must beavailable to the user in order to satisfy the widevariety of experimental conditions encountered inbiological studies. It must also be realized that thePET image data is often only one component ofthe biological investigation that must be correlatedwith other physiologic or metabolic data collectedconcurrently. In the acquisition of rapid kineticdata for blood flow studies for instance [20,22], it isquite difficult to synchronize the image acquisitionwith the bolus injection of the radiotracer and withthe input function obtained by drawing bloodsamples. Considering the shorter time scale ofphysiological processes in small animals, accurateestimation of the kinetic model rate constants iscritically dependent on an exact timing betweenthese three measurements. Another area of appli-cations where synchronicity is important relates toreal-time follow-up of intervention. Monitoringthe early effects of therapy to study the mechan-isms of action of a drug or to optimize treatmentdelivery requires accurate timing and should avoidloss of data between change of acquisition modes.

The utility of cardiac gating has been demon-strated for the study of left ventricular function inseveral small animal models [22,71–75]. Thecardiac cycle in anesthetized and diseased animalsare often subject to instabilities provoking varia-tions of the heartbeat duration and arrhythmia[22]. Such irregular cycles are usually compensatedfor by dropping the last frame of the gated seriesor by rejecting bad beats on the fly until stableconditions come up again. To make the mostefficient use of available data and shorten acquisi-tion time, it would be advantageous to collect alldata for post-acquisition processing. The samealso applies to respiratory gating [76], which wouldbe useful in high resolution imaging of the thorax.Ideally, all relevant data should be registered

together with the PET image data, using the sametime stamps, so that external data as well as anyfortuitous or provoked change of the animal’sphysiologic state can be related to the PET images.In practice, the associated data could as well serveto monitor the physiologic state of the subjectduring an imaging session and, if necessary, beused to discard unreliable image data a posteriori.Such real time monitoring of the animal conditionis important to ensure the validity of the acquiredPET images and contributes to improve thereproducibility of PET investigations.Such capabilities call for list-mode data acquisi-

tion, where data from other sources can beinterleaved into the PET image data stream[77,78]. One key advantage of the list-modedata acquisition is the ability to re-sample ordiscard dynamic image data as desired after the

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measurement, thus avoiding the necessity todetermine the imaging sequence beforehand andto synchronize image acquisition with externalinformation.

4. Conclusion

Driven by the stringent requirements of mole-cular imaging in small animal models, the PETtechnology has substantially evolved during thelast decade. The image quality currently achievedin rats using the available dedicated small animalscanners withB10 ml resolution is grossly compar-able to human images obtained with contempor-ary clinical systems. Current developments indetector technology, scanner design and imagingmethodology can be expected to provide 1 mlvolumetric resolution with sufficient signal-to-noise for dynamic mouse imaging in the nearfuture.

Acknowledgements

The author wishes to acknowledge the contribu-tions of the Sherbrooke PET Group in preparingthis paper. This work was supported in part by theCanadian Institutes of Health Research underGrant MOP-15348 and by research grants of theNatural Sciences and Engineering Research Coun-cil of Canada.

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