A small animal PET based on GAPDs and charge signal transmission approach for hybrid

PET-MR imaging

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2011 JINST 6 P08012

PUBLISHED BY IOP PUBLISHING FOR SISSA

RECEIVED: June 29, 2011ACCEPTED: August 9, 2011

PUBLISHED: August 24, 2011

A small animal PET based on GAPDs and chargesignal transmission approach for hybrid PET-MRimaging

Jihoon Kang,a,b Yong Choi,a,1 Key Jo Hong,a Wei Hu,a,b Jin Ho Jung,a

Yoonsuk Huha,b and Byung-Tae KimbaDepartment of Electronic Engineering, Sogang University,1 Shinsu-Dong, Mapo-Gu, Seoul 121-742, Republic of Korea

bDepartment of Nuclear Medicine, Samsung Medical Center,Sungkyunkwan University School of Medicine,50 Ilwon-Dong, Gangnam-Gu, Seoul 135-710, Republic of Korea

E-mail: ychoi.image@gmail.com

ABSTRACT: Positron emission tomography (PET) employing Geiger-mode avalanche photodiodes(GAPDs) and charge signal transmission approach was developed for small animal imaging. Ani-mal PET contained 16 LYSO and GAPD detector modules that were arranged in a 70 mm diameterring with an axial field of view of 13 mm. The GAPDs charge output signals were transmittedto a preamplifier located remotely using 300 cm flexible flat cables. The position decoder circuits(PDCs) were used to multiplex the PET signals from 256 to 4 channels. The outputs of the PDCswere digitized and further-processed in the data acquisition unit. The cross-compatibilities of thePET detectors and MRI were assessed outside and inside the MRI. Experimental studies of thedeveloped full ring PET were performed to examine the spatial resolution and sensitivity. Phantomand mouse images were acquired to examine the imaging performance. The mean energy and timeresolution of the PET detector were 17.6% and 1.5 ns, respectively. No obvious degradation onPET and MRI was observed during simultaneous PET-MRI data acquisition. The measured spatialresolution and sensitivity at the CFOV were 2.8 mm and 0.7%, respectively. In addition, a 3 mmdiameter line source was clearly resolved in the hot-sphere phantom images. The reconstructedtransaxial PET images of the mouse brain and tumor displaying the glucose metabolism patternswere imaged well. These results demonstrate GAPD and the charge signal transmission approachcan allow the development of high performance small animal PET with improved MR compatibil-ity.

KEYWORDS: Gamma camera, SPECT, PET PET/CT, coronary CT angiography (CTA); Multi-modality systems

1Corresponding author.

c 2011 IOP Publishing Ltd and SISSA doi:10.1088/1748-0221/6/08/P08012

mailto:ychoi.image@gmail.comhttp://dx.doi.org/10.1088/1748-0221/6/08/P08012

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Contents

1 Introduction 1

2 Materials and methods 22.1 System description 22.2 Performance measurement of the PET detector modules outside MRI 32.3 Characterization of the cross-compatibility of PET detector module and MRI 42.4 Phantom and small animal imaging of the full ring PET 4

2.4.1 Spatial resolution and sensitivity 42.4.2 Phantom images 52.4.3 Mouse images 5

3 Results 63.1 Performance of the PET detector modules outside the MRI 63.2 Characterization of the cross-compatibility of PET detector module and MRI 73.3 Phantom and small animal imaging of the full ring PET 8

3.3.1 Spatial resolution and sensitivity 83.3.2 Phantom images 83.3.3 Mouse images 9

4 Discussion 9

5 Conclusion 11

1 Introduction

Positron emission tomography (PET) has attracted considerable interest for the non-invasive visu-alization of small animals for various preclinical studies. Currently, combined positron emissiontomography and computed tomography (PET-CT) dedicated to small animal imaging is commer-cially available [15] and has proven to be a valuable imaging tool providing a fused image of highresolution anatomical and quantitative functional information. Moreover, combined PET and mag-netic resonance imaging (PET-MRI) has been proposed for simultaneous functional and morpho-logical images [68]. Extensive studies have been carried out by several research groups to developMR-compatible PET based on photomultiplier tube (PMT) using optical fiber technology [912],and avalanche photodiodes (APDs) using RF shielding technology [13, 14].

Recently, the next generation photosensor Geiger-mode avalanche photodiodes (GAPDs) [15],also called a solid state photomultiplier (SSPM) [16], silicon photomultiplier (SiPM) [17], multipixel photon counters (MPPC) [18] and micro-pixel avalanche photodiode (MAPD) [19], was de-veloped. GAPDs consist of a densely packed matrix with many microcells (100010000) rang-ing from 55 to 100100 m2 size, and each microcell operates independently in a Geiger mode

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as on/off switch for the photons. The amplitude of output pulse is proportional to the total numberof fired microcells, reflecting the number of absorbed photons. GAPDs have been studied activelyas a PET photosensor owing to their several key properties [20]. Compared to PMT, GAPDs havecompactness and insensitivity to high magnetic fields. This allows GAPD to be located inside theMR bore without the need for optical fibers to integrate hybrid PET-MRI. In contrast to APDs,GAPDs have high gain (106) and low excess noise factor (1.1), allowing operation with a sim-ple preamplifier. Their fast response time (< 1 ns) allows a high true-to-random ratio as well asthe development of a hybrid PET-MRI with time-of-flight (TOF) capability in human whole-bodyapplication. Further advantages over PMT and APDs are the low operating voltage (< 100 V)and high uniformity (< 20%) among the pixels. In addition, the fabrication costs can be reducedsubstantially because it can be manufactured using a standard Metal Oxide Semiconductor (MOS)production process [21].

The utility of the charge signal transmission approach, which relays the charge signal fromthe photosensor to the remotely located preamplifiers for PET signal transmission, was recentlyreported [22]. This detector concept has several potential merits because it allows the placementof amplifier units at a safe distance for integrated PET-MR scanner, can decrease the space re-quirements to insert a PET scanner into the restricted MR bore, minimize the mutual interferencebetween PET and MRI, and eliminate the need for placing RF shielding materials close to fieldof view of the MR scanner. Moreover, it can reduce the deterioration by temperature-related per-formance changes, which result from the local heat production generated in the amplifier unit, inthe PET system based on semiconductor photosensor. A previous study verified that there was noconsiderable degradation in PET detector performance, such as photopeak position of the 511 keV,energy resolution and time resolution, even though the PET charge signal was transferred via longcables (300 cm). On the other hand, the scope of our previous study was limited to the performancecharacterization of the PET detector that did not use the channel reduction circuits to acquire to-mographic image from a PET scanner.

The aim of this study was to develop a small animal PET scanner based on GAPDs and thecharge signal transmission approach. The performance of the PET detector modules was evaluatedand the cross-compatibility between the PET detector module and MRI was assessed. Quantitativeanalysis of the full ring prototype PET was performed and tumor mouse images were acquired.

2 Materials and methods

2.1 System description

The full-ring PET contained 16 detector modules arranged in a ring, 70 mm in diameter with anaxial field of view (FOV) of 13 mm (figure 1). Each PET detector was comprised of a 44 lutetiumyttrium oxyorthosilicate array (LYSO array, Sinocera, Shanghai, China) with an individual crystalsize of 3 3 10-mm, arranged with a 3.3 mm pitch. All crystal elements were polished andseparated with white epoxy except for the photosensor face. The crystal block was coupled directlyto a 3-side buttable GAPD array (SPMArray2, SensL, Cork, Ireland). Each pixel of the GAPDarrays had a 2.85 2.85-mm active area and 3,640 microcells of the 35 35-m. The feasibilityof these GAPDs for the development of PET has been reported elsewhere [2326].

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(a) (b)

Figure 1. Schematic diagram (a) and electronic components (b) of small animal PET. (a): The animal PETcontained 16 LYSO and GAPD detector modules arranged in a 70 mm diameter ring with an axial field ofview (FOV) of 13 mm. (b): The GAPD outputs were transmitted to the amplifier units and further processedin a PDC multiplexing 64 GAPD pixel output to one analog pulse signal and 6 bit position information.

The 3-meter flexible flat cables (FFCs, New Grand TECH, Shenzhen, China) were used totransmit the charge signal from the PET detector to remotely located amplifier units, which con-sisted of a 16 channel trans-impedance preamplifier (TIA) and a bias regulation circuit.

The TIA converted the GAPD outputs to differential voltage signals providing high gain(103) without introducing additional noise and temperature-related gain instability. The biasregulation circuit could adjust the bias voltage (28.532.5 V) finely at 8 mV intervals using aprogrammable digital potentiometer (AD5231, Analog Devices, MA, USA) and it was possible toprovide optimal operating conditions for the GAPDs.

The position decoder circuits (PDC) [27] that were capable of multiplexing the 64 GAPDpixel output (4 PET detectors 16 channels/detector) to one analog pulse signal and 6 bit positioninformation were used. In addition, the PDC contained the gain adjustable circuits to achieve gainhomogeneity for all channels of the PET detectors. The 4 PDCs were used to multiplex all the 256electrical PET signals (16 PET detectors 16 channels/detector), which simplified the PET systemdesign by reducing the required ADC number and analog output lines from 256 to 4 channels.Four PDC output signals were fed into the data acquisition (DAQ) unit using 10-meter twist-pairedcables and co-axial cables to process the interaction position and analog signal, respectively.

The DAQ unit (Lyrtech, Quebec, Canada) consisted of free-running analog to digital convert-ers (ADC) and a field programmable gate array (FPGA). The analog signals of the PDCs weredigitized at a 105 MHz sampling rate and a 14-bit vertical resolution in the -1.25+1.25 V range.The digitized signals were processed further by FPGA to calculate the accurate energy and timeinformation [28]. After signal processing, the output data containing the pulse energy, arrival timeand position information were recorded in the 128 MB RAM in list mode format.

2.2 Performance measurement of the PET detector modules outside MRI

A 200-kBq 22Na point source placed centrally between the paired PET detectors was used to ir-radiate the LYSO-GAPD arrays. The energy and time spectra were acquired at room temperaturewithout additional cooling of the PET detectors. The energy and time resolution were calculatedas the full width at half maximum (FWHM) of the Gaussian distribution plot. The lower energy

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threshold was set to 350 keV and the coincidence time window was 4 ns. The variation of thecount uniformity and photopeak position of the 16 channels flood histogram were calculated as theratio of the standard deviation to the average value.

2.3 Characterization of the cross-compatibility of PET detector module and MRI

The performance measurements of the PET detectors were also repeated in 7-T MRI (BrukerBioSpec, Ettlingen, Germany) to characterize the effect of MRI on the PET detector module. Apair of LYSO-GAPD PET detectors was inserted inside the MRI bore between the RF coil andgradient coils. The GAPD outputs were transmitted to the amplifier units using 3-meter FFCs. ThePET electronics were positioned outside of the 5-Gauss line (1.5-meters away from the magnetisocenter) in the MR to minimize the mutual interference between PET and MRI. In this study, noelectromagnetic shielding was introduced to protect the PET components from the MR gradientand RF field.

A uniform cylindrical phantom (30 mm diameter, 100 mm length) filled with a copper sul-fate solution was imaged to examine the effect of the PET detector modules on MR images. TheCuSO4-filled phantom was placed isocentrically inside the RF-coil, and the transaxial images wereacquired with and without a pair of LYSO-GAPD PET detectors inside the MR bore. The standardMR imaging sequences, including the Gradient echo (TR = 205 ms, TE = 6 ms, FA = 15 degree),Spin echo T1 (TR = 419 ms, TE = 8 ms) and Spin echo T2 (TR = 3,000 ms, TE = 75 ms) wereused in this study. The ParaVision software (Bruker BioSpec, Ettlingen, Germany) was used toacquire and process the MR data, such as data acquisition, analysis, reconstruction and visualiza-tion. Representative three transaxial MR image slices with a 5 mm interval in the axial directionwere analyzed quantitatively. A region of interest (ROI) was drawn at the center of the phantomimage enclosing approximately 80% of the phantom and the uniformity and signal to noise ratio(SNR) were calculated from the ROI. The experimental tests were repeated five times to minimizethe measurement errors. In each measurement, the MR phantom was re-positioned and the RF-coilwas re-inserted and re-tuned.

2.4 Phantom and small animal imaging of the full ring PET

The list-mode data of the full ring PET were rebinned using a single slice rebinning (SSRB) methodand the valid events were sorted into the 2D sinogram, which had 43 samples in the transverse di-rection and 79 angular samples. The transverse sampling distance was 1.5 mm near the CFOV.The sinogram was normalized for the detector efficiency, and these normalization factors were es-timated from a direct inversion of the sinogram acquired with a uniform cylindrical phantom filledwith a 18F-FDG solution. The missing data caused by the effect of the gaps was compensatedfor by a nearest neighbor 1-D interpolation in the radial direction [29]. The PET images were re-constructed by a 2D filtered backprojection (2D FBP) using a Hanning filter with a cutoff at theNyquist frequency.

2.4.1 Spatial resolution and sensitivity

The spatial resolution of the prototype PET was measured using a glass capillary tube with aninner diameter of 0.5 mm. The line source was placed at five radial offset locations of 0, 10, 15,

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Figure 2. Experiment setup for the mouse imaging study. The prototype PET was installed with a heatingpad, anesthesia and respiration monitoring systems for the in-vivo study.

20 and 25 mm in the FOV. Each line source was filled with approximately 370-kBq of a 18F-FDGsolution. The radial profiles of each location were fitted with Gaussian profiles and its width wasmeasured by the FWHM. The coincidence sensitivity of the prototype PET was measured using apoint source with an inner diameter of 0.5 mm. A 200-kBq 22Na point source was located preciselyat the center of the ring and the PET data was acquired for 1 min. The data recorded in list modeformat was sorted using four different energy windows ranging from 10% (460560 keV) to40% (300710 keV). A coincidence time window of 4 ns was applied.

The system sensitivity was calculated as the number of detected events divided by the numberof decays by positron emissions that were predicted to have occurred during the acquisition period.

2.4.2 Phantom images

Two cable lengths (10 cm and 300 cm FFCs), connecting the GAPD arrays to the amplifier units,were used to evaluate the effect of the cable length on the PET image. Custom-made hot-spherephantoms, 60 mm in diameter, were used to examine the imaging performance. The sphere diame-ters were 3, 4, 5, 6 and 7 mm and the center-to-center distance between the spheres was twice theirdiameter. The phantoms were filled with a 20 MBq 18F-FDG solution. The PET imaging data wasacquired for 10 minutes.

2.4.3 Mouse images

An in vivo rodent imaging study was performed in accordance with the protocols approved bythe Samsung Biomedical Research Institute (SBRI) in Korea. A male mouse with a tumor inthe right thigh was injected with 100-MBq of 18F-FDG through the tail vein and imaged after 1hour of radiotracer uptake. The mouse was placed on a carbon animal bed with a heating pad thatmaintained a temperature of 37 C. During the PET scan, the mouse was anesthetized by isofluraneinhalation and respiration was monitored. The PET imaging data was acquired for 10 minutes/bedat 2 different bed positions. Figure 2 shows the experimental setup.

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(a) (b)

Figure 3. Representative energy spectrum (a) and time spectrum (b) acquired outside MRI. The energy andtime resolution of the PET detector were 16.4% and 1.5 ns, respectively.

(a) (b)

Figure 4. Flood histogram (a) and plots of the energy resolution and photopeak position (b) of the 4 4LYSO and GAPD array. The changes in the count uniformity in the flood histogram and the photopeakposition were 3.3% and 3.9%, respectively.

3 Results

3.1 Performance of the PET detector modules outside the MRI

Figure 3 shows the representative energy and time spectrum of the PET detector. The mean energyresolution was 17.6%, ranging from 16% to 19%. The time resolution obtained was 1.5 ns for 511-keV photons. Figure 4.a shows the flood histogram of the data acquired from the PET detector.The variation of the count uniformity was 3.3%. The variation of the photopeak position of theLYSO-GAPD detector was 3.9% and was improved after gain adjustment of the PDC (figure 4.b).

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(a) (b)

Figure 5. Energy spectra (a) and Time spectra (b) acquired both outside and inside the MRI. No obviousperformance degradation of the PET detectors was observed.

Figure 6. Quantitative analysis of the MR phantom images in terms of the SNR (left) and uniformity (right):gradient echo (first row), T1 weighted spin echo (second row) and T2 weighted spin echo (third row).

3.2 Characterization of the cross-compatibility of PET detector module and MRI

The energy and time spectra were acquired simultaneously for approximately 5 min, whereas MRimaging was performed using three different sequences, as shown in figure 5. No obvious per-formance degradation of the PET detectors was observed, as measured by the energy and timeresolution. In addition, the photopeak position and coincidence count rate were similar regardlessof whether the PET detectors had been operated outside or inside the MRI.

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Figure 7. Spatial resolution of the developed PET. The measured radial resolution ranged from 2.8 mm to4.1 mm.

Table 1. Sensitivity at the CFOV for the 4 different energy window settings.

Energy window Total counts Sensitivity (%)

460560 keV (10%) 11,396 0.29

410615 keV (20%) 18,894 0.48

350650 keV (30%) 25,843 0.65

300710 keV (40%) 31,050 0.78

There were no significant artifacts or distortions observed in the MR phantom images. Figure 6shows the calculated values of the uniformity and SNR for three MR sequences. There was nomajor loss caused by inserting the PET detector modules in the SNR and uniformity of MR images.

3.3 Phantom and small animal imaging of the full ring PET

3.3.1 Spatial resolution and sensitivity

Figure 7 shows the measured phantom images and radial profiles for 5 different radial offsets acrossthe useful FOV. The radial resolution was 2.8 mm FWHM at the CFOV, which increased to 4.1 mmwith a 20 mm offset.

Table 1 lists the sensitivity at the CFOV for different energy window settings. The sensitivitywas corrected for a branching ratio of 22Na (0.906). The system sensitivity increased approximatelyfour times as the energy window was changed from 460560 keV to 300710 keV. The prototypePET had a peak sensitivity of 0.65% in the standard energy window of 350650 keV.

3.3.2 Phantom images

Figure 8 shows transverse phantom images and line profiles acquired using the 10 cm (a) and300 cm (b) FFCs connecting GAPDs and preamplifiers. The 3 mm diameter was clearly resolvedin the hot-sphere phantom images. As expected from previous studies [22], there was no degrada-tion of the PET image quality caused by employing a long cable (300 cm) from the GAPDs to thepreamplifiers used for PET signal transmission.

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(a) (b)

Figure 8. Hot-sphere phantom images and line profiles were acquired using the 10 cm (a) and 300 cm (b)FFCs between GAPDs and preamplifiers. There was no degradation of the PET image quality caused by thelong cable (300 cm) from the GAPDs to the preamplifiers in the PET signal transmission.

3.3.3 Mouse images

A PET mouse image was acquired to demonstrate the potential of this prototype system for an invivo study. Figure 9 shows the reconstructed transaxial PET images of a mouse. The mouse brainand tumor displaying the glucose metabolism patterns were well imaged.

4 Discussion

A small animal PET based on the GAPDs and a charge signal transmission approach was devel-oped. In this system, the charge signals of GAPDs were transmitted to the amplifier units positionedfar away from the PET detectors or outside the MR bore using long transmission cables. The initialPET images of the phantom and tumor mouse were obtained in this study. In addition, simultane-ous PET and MR data was acquired with no deterioration in the performance of both PET detectormodules and MRI. The concept proposed in this study might provide several technical advantages.

It is feasible to develop a PET system based on a semiconductor photosensor with no obviousloss of PET performance. Unfortunately, APDs have not achieved the traditional PMT perfor-mance. The low internal gain of the APD (< 102) requires the use of sophisticated low noisepreamplifiers [30]. The overall time resolution of the imaging device is reduced by the signifi-cantly inferior timing properties of the APDs due to the slow rise time, high time jitter, high noisefactor, and low signal-to-noise ratios [21]. This results in a wider coincidence time window and an

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Figure 9. In vivo study of the prototype PET. The PET imaging data was acquired for 10 min at the differentbed positions and the mouse brain (top) and tumor (bottom) displaying the glucose metabolism patterns werewell imaged.

increased number of random coincidence counts, which will degrade the image quality of PET. Onthe other hand, GAPD can overcome these technical challenges. Simple electronics can be used inthe development of PET based on GAPD. In addition, GAPD can provide good PET performance,such as energy resolution (figure 3. (a)), time resolution (figure 3. (b)), count uniformity (figure 4.(a)) and long-term stability [25]. The inherent characteristics of the GAPDs could improve the PETimage performance by providing a high SNR and high true-to-random coincidence rates.

In addition, the energy resolution of the 511-keV photopeaks was 17% on average for allcrystals, which is better than the 25% [31] and 27% [32] energy resolutions obtained for pre-viously reported animal PET scanners based on GAPD and LGSO. This may be the result of thedetector configuration, where the individual LYSO crystal was coupled one to one to a separatepixel of the GAPD and the output signals from each PET detector module were processed inde-pendently using the PDC, which could eliminate the performance degradation caused by opticalcrosstalk through the light guide inserted between the crystal block and photosensor array in theconventional detector configuration based on the Anger type channel reduction circuit.

As shown in figure 8, the developed PET system using 300 cm long cables produced phantomimages without noticeable quality degradation. Moreover, the charge signal transmission approachcould minimize the deterioration by temperature changes. A potential technical hurdle is that thePET performance degrades gradually by the heat generated from the amplifiers unit in the PET sys-tem based on semiconductor photosensor [22]. Previous studies using a semiconductor photosensor

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employed a costly cooling system or gain control system to minimize the temperature-related per-formance variations of the photosensor [3336]. This detector concept, photosensors and amplifiersfabricated in the separated housing box, could minimize the temperature-related performance vari-ations of the photosensor, such as the gain, PDE and changes in dark currents, which could affectthe PET performance, such as altered photo-peak position, decreased count rate, degraded energyresolution and time resolution.

In terms of the MR compatibility, the developed PET system only requires the placement ofnon-magnetic PET components, scintillation crystals and GAPD arrays without preamplifiers andsubsequent electrical circuits, inside the MR bore. The PET system requires only low DC powersupplies with low current consumption (32 V, < 2 mA) inside the MR bore. Moreover, this systemdoes not introduce conducting materials, such as RF shielding materials and electronic circuitsinside the MR bore, which can produce eddy currents and heat from the high power switchinggradient coil [3739]. The free scalability of the PET geometry, such as transaxial and axial FOV,are an additional advantage by locating the amplifier units at low fringe field areas, which coulddecrease the space requirements to insert a PET scanner into the restricted MR bore.

This study had some limitations. First, the spatial resolution was relatively poor compared toother systems based on the Anger logic circuit, which was caused by the one to one coupling ofan individual crystal with a separate pixel of large area GAPDs ( 33mm2). Although the PETspatial resolution was poor, these results demonstrate that it was feasible to develop small animalPET using GAPDs and charge signal transmission approach technologies. Moreover, the coinci-dence detection efficiency was improved by a factor of 2 when the crystal length was increasedfrom 5 mm to 10 mm at the expense of parallax error. A sensitivity of 0.7% was acquired usingthe developed PET, even though it had a shorter axial FOV and larger transaxial FOV than otherprototype systems [12-14, 40]. Second, simultaneous PET-MR images were not acquired becauseit was difficult to use a radioisotope in a MR imaging site. On the other hand, the feasibility ofhybrid PET-MRI with this system design was observed to some extent (figure 5 and 6). Neverthe-less, further studies will be needed to improve the spatial resolution using smaller size GAPDs andacquire simultaneous PET-MR images using a range of MR imaging sequences.

5 Conclusion

A small animal PET was developed based on GAPDs and the charge signal transmission approach.The PET detector performance and cross-compatibility were examined. Quantitative analysis ofthe full ring prototype PET was performed and the tumor mouse images were acquired success-fully. These results demonstrated that the GAPD and charge signal transmission approach allowthe development of high performance small animal PET with improved MR compatibility.

Acknowledgments

This research was supported by the Converging Research Center Program through the National Re-search Foundation of Korea (NRF) funded by the Ministry of Education, Science and Technology(2011K000715), and by the Technology Innovation Program funded by the Ministry of KnowledgeEconomy (10030029), Republic of Korea.

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http://dx.doi.org/10.1118/1.3495683http://dx.doi.org/10.1088/0031-9155/54/11/015http://dx.doi.org/10.1088/0031-9155/54/11/015http://dx.doi.org/10.1088/1748-0221/6/05/P05012http://dx.doi.org/10.1088/1748-0221/6/07/P07003http://dx.doi.org/10.1016/j.nima.2010.04.028http://dx.doi.org/10.1016/j.nima.2010.04.028http://dx.doi.org/10.1016/j.nima.2010.05.047http://dx.doi.org/10.1016/j.nima.2010.05.047http://dx.doi.org/10.1109/TNS.2005.858236http://dx.doi.org/10.1109/TNS.2005.858236http://dx.doi.org/10.1109/23.983270http://dx.doi.org/10.2967/jnumed.110.079707http://dx.doi.org/10.1088/0031-9155/55/19/013http://dx.doi.org/10.1073/pnas.0711622105http://dx.doi.org/10.1088/0031-9155/55/7/003http://dx.doi.org/10.1088/0031-9155/55/7/003http://dx.doi.org/10.1016/j.neuroimage.2009.10.036http://dx.doi.org/10.1016/j.neuroimage.2009.10.036http://dx.doi.org/10.1088/0031-9155/56/9/015http://dx.doi.org/10.1002/jmri.1880050115http://dx.doi.org/10.1002/jmri.1880050115http://dx.doi.org/10.1118/1.2132571http:/dx.doi.org/10.1053/j.semnuclmed.2008.02.001http://dx.doi.org/10.1109/TNS.2009.2015448http://dx.doi.org/10.1109/TNS.2009.2015448

IntroductionMaterials and methodsSystem descriptionPerformance measurement of the PET detector modules outside MRICharacterization of the cross-compatibility of PET detector module and MRIPhantom and small animal imaging of the full ring PETSpatial resolution and sensitivityPhantom imagesMouse images

ResultsPerformance of the PET detector modules outside the MRICharacterization of the cross-compatibility of PET detector module and MRI Phantom and small animal imaging of the full ring PETSpatial resolution and sensitivityPhantom imagesMouse images

DiscussionConclusion