evaluation of hamamatsu h13974 large sensitive area flat

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This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication. Citation information: DOI 10.1109/TRPMS.2018.2797323, IEEETransactions on Radiation and Plasma Medical Sciences

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Evaluation of Hamamatsu H13974 Large SensitiveArea Flat Panel PMT Array for use in Small

Animal Imaging and ScintimammographyEleftherios Fysikopoulos, Maria Georgiou and George Loudos

Abstract—In this work we report for the first time on theperformance of the Hamamatsu H13974 flat panel PMT array.This new photodetector is based on four Hamamatsu H12700APosition Sensitive Photomultiplier Tubes in a 2× 2 arrangementproviding large sensitive area (100 × 100 mm2) for use inscintimammography or/and medium size mammals imaging.We acquired raw images from three pixellated scintillators forpossible Positron Emission Tomography (PET) and Single PhotonEmission Computed Tomography (SPECT) applications (BGOwith 2 × 2 × 6 mm3, CsI:Tl with 2 × 2 × 4 mm3 and CsI:Nawith 1.45 × 1.45 × 6 mm3 pixel sizes) irradiated with 511 keV(68Ga) and 140 keV (99mTc) respectively. Different couplingschemes involving optical grease and optical glasses of differentthicknesses were used in order to find the scheme that betterresolves the crystal elements of the scintillators especially inthe dead areas of the flat panel PMT array. We present theresults of investigating the basic properties, including energyresolution, intrinsic spatial resolution, detection efficiency andsensitivity uniformity. The scintillation material, the couplingscheme and the analog readout circuits are the main factors thataffect the performance of the detector. The proper combinationof these factors should be carefully chosen so as to build a highperformance imaging system.

Index Terms—Position sensitive gamma-ray detector, Multianode photomultiplier, Detector performance evaluation, Smallanimal imaging, Scintimammography.

I. INTRODUCTION

NUCLEAR medicine, the major family of molecular imag-ing modalities, is based on detecting nuclear radiation

emitted from the body after introducing a radioactive tracerinside. These techniques are revolutionizing the way we study,visualize and investigate complex biochemical phenomena,diagnose diseases, approach design, and assess therapy nonin-vasively. A number of dedicated imaging systems have beendeveloped for use on small animal imaging for the evalu-ation of new radiopharmaceuticals or small human organs,including scintimammography and sentinel node detection. Insuch applications, it is important to use a detector that canproduce images at sufficiently high resolution and sensitivityand in the same time be cost-efficient compared to generalpurpose clinical systems. Many of these systems are basedon position sensitive photomultiplier tubes (PSPMTs) sincethey meet these requirements and they can be purchased at arelatively low cost. High resolution nuclear medicine cameras

E. Fysikopoulos and M. Georgiou are with BET Solutions R&D, Athens,Greece, e-mail: [email protected].

G. Loudos is with the Department of Biomedical Engineering, Technolog-ical Educational Institution of Athens, Athens, Greece.

based on PSPMTs have been extensively used the last twodecades for in vivo studies with small animals [1]-[6] and inscintimammography applications [7]-[10].

Molecular imaging of small laboratory animals, such asmice, rats and rabbits, plays an important role in biologicalstudies of physiology and pathology as well as in drugdesign and evaluation. Biodistribution studies including ex vivotissue counting or autoradiography are used to study vari-ous radiopharmaceuticals, biomolecules, and disease models.However, they require a large number of animals, which intro-duces restrictions in the design of the experimental protocolsand raises economical and ethical issues. Nuclear medicinetechniques (single photon emission computed tomography(SPECT), positron emission tomography (PET)) allow the invivo investigation of complex biochemical phenomena [11]-[13], by providing multiple recordings for the same animal,over a period of days or even weeks [14].

Breast cancer is the most common cancer in women world-wide, with nearly 1.7 million new cases diagnosed in 2012(second most common cancer overall) [15]. Although radionu-clide imaging offers additional advantages to standard X-raymammography, conventional gamma cameras are inadequatefor scintimammography, since they suffer from poor sensitivityfor lesions with diameter less than 1 cm, especially for tumorsclose to the chest wall [16], [17]. Breast-specific nuclearmedicine cameras (either single photon emission mammogra-phy (SPEM) [18] or positron emission mammography (PEM)[19]) improve sensitivity and spatial resolution of conventionalscintimammography, and provide accessibility to posterior andmedial areas of the breast.

In the beginning of the previous decade the commercialavailability of square PSPMTs and mainly Hamamatsu H8500allowed the construction of larger field of view (FOV) camerasbased on the combination of two or more PSPMTs [2], [9],while methods for recovering signals in the dead area betweenthe PSPMTs were developed [20]-[22]. This approach offersa variety of different arrangement schemes overcoming thelimited choise in camera active area when using a single large-area PSPMT (i.e. Hamamatsu round R3292 PSPMT) [2].

Hamamatsu has very recently developed H13974-00-1616flat panel PMT Array which provides a large sensitive area(100× 100 mm2) ideal for imaging of medium size animals,such as rats or/and rabbits and for use in scintimammogra-phy applications. Furthermore, the large size of the PSPMTallow its application in broader non medical applications. Thearray is based on 4 square PSPMTs (H12700A) in a 2 × 2

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arrangement. The H12700 series is a direct replacement forthe H8500 since the dimensional outlines and anode pinoutare identical. H13974-00-1616 flat panel PMT array offers acomplete solution, compensating the crucial technical issueof minimizing the effects of the dead spaces between theindividual PMTs in an array. In other words, the H13974provides a high quality, ready to use large sensitive area flatpanel PMT array.

The aim of this study is to evaluate for the first time theperformance of this recently commercially available flat panelarray for use in SPECT and PET applications. Performancemetrics include energy resolution, detection efficiency andsensitivity uniformity. Different scintillators of small pixel sizewere used to study successful identification of single pixels,especially over the dead areas of the PSPMT.

II. MATERIALS AND METHODS

A. Photodetector

Hamamatsu’s H13974-00-1616 flat panel PMT array isbased on 4 square PSPMTs (H12700A) in a 2 × 2 arrange-ment and total dimensions 105.7 × 105.7 × 32.5 mm3. TheH12700A Multi-Anode photomultiplier tube is a novel 64-channel, 51.7×51.7 mm2 square pixelated device produce byHamamatsu. It consists of a 8× 8 pixel matrix (6× 6 mm2)able to detect single photons, amplifying the signal through a10-stage dynode chain. The PSPMT is sensitive in the rangeof 300 nm to 650 nm, with a maximum quantum efficiencyof ∼33% at about 400 nm [23]. Among the multi-channelphotodetectors, the H12700A stands out for its large effectiveactive area (48.5 × 48.5 mm2) while the very small inactiveborder around the device and the MaPMT square cross-sectional geometry allows for a close packing ratio (∼87%)[23], making it ideal for use in multiple tube detector schemesby minimizing the dead area corresponding to gaps betweenthe tubes (the separation between the edges of photocathodesof adjacent tubes is only about 3.2 mm). Further minimizationis provided through the H13974 series, in which the dead space(gap) is only 3 mm according to the company, critical belowin comparison to previous published in-house solutions basedon H8500 PSPMTs [3], [4] and also below a potential in-house solution based on H12700 series PSPMTs. Moreover,the array is equipped with a high voltage adjustment circuitfor each PSPMT, allowing the use of just one high voltagecomponent for the whole array. Also the chosen PSPMTs havesimilar gain and anode non-uniformity eliminating the need touse an anodes gain variations compensation circuit [24]. Thesefeatures, together with the nominal low dark counts contribu-tion and the moderate cross-talk between neighbouring pixels,make the H13974-00-1616 flat panel PMT array particularlysuitable for applications in nuclear medicine where relativelylarge field of view is required. High voltage was set in therecommended range, at -860 V.

B. Front end electronics

The array comes along with a prefabricated readout elec-tronic circuit, capable for resolving scintillator crystal elementsin dead areas [22]. The 256 output signals of the four PSPMTs

were reduced to 4 position signals (Xa, Xb, Yc, Yd) througha discretized position-sensitive readout circuit (DPC). DPCconsists of a string or array of resistors that divide the chargebetween low impedance collection op-amps. This simple tech-nique provides positional charge division. Four custom pre-amplifiers were implemented at the end of the resistive chainto shape the position signals, taking into account the analogto digital conversion (ADC) sampling rate [25]. The X andY position can be determined by equations (1) and (2). Asummed signal from the four position signals was used toprovide the signal which is proportional to photon’s energy(3).

X =(Xa +Xb)− (Yc + Yd)

Xa + Yc +Xb + Yd(1)

Y =(Xa + Yd)− (Yc +Xb)

Xa + Yc +Xb + Yd(2)

E = Xa +Xb + Yc + Yd (3)

C. Data acquisition electronics

The four position signals are continuously sampled by a 12bit octal-channel free running ADC with up to 65 Msps sam-pling rate [26]. A Field Programmable Gate Arrays (FPGA)based data acquisition system was developed in a Spartan-6FPGA LX150T Development Board (Avnet, Inc) [27], for trig-ger and digital signal processing of the acquired samples. Thisplatform also contains a 128 MB DDR3 SDRAM componentmemory and a gigabit Ethernet, which are used in the currentimplementation for temporary data storage and transmission.More details about the data acquisition method that was usedcan be found in a previous study of our group in [28].

D. Scintillation detection and optical coupling

Three different pixellated scintillator arrays were used, inorder to evaluate the photodetector under 140 keV (99mTc)and 511 keV (68Ga) irradiation for SPECT and PET ap-plications respectively. Different coupling schemes involvingoptical grease BC-630 (Saint Gobain, France) and borosilicateglasses (Przisions Glas & Optik, PGO, Germany) of differentthicknesses were used in order to find the scheme that bestresolves the crystal elements of the scintillators in the deadareas of the flat panel PMT array.

• BGO with pixel size 2 × 2 mm2, septa 0.25 mm, 6 mmthickness (5 mm + 1 mm reflector) and 50×50 mm2 totaldimensions (Hilger Crystals, UK) for PET applications.The scintillator was coupled to the photodetector withoptical grease. Optical glasses of 1 mm, 2 mm and 3 mmthickness along with optical grease were used.

• CsI:Tl with pixel size 2× 2 mm2, septa 0.22 mm, 4 mmthickness (3 mm + 1 mm reflector) and 12.7 cm diameter(Hilger Crystals, UK) for SPECT applications. Opticalglasses of 1 mm, 2 mm, 3 mm and 4 mm thickness alongwith optical grease were used.

• CsI:Na with pixel size 1.45 × 1.45 mm2, septa 0.25mm, 9.3 mm thickness (5 mm + 1 mm reflector + 3.3





























































































































































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Fig. 1. (First row): Raw images acquired with the discrete 2 × 2 × 6mm3 pixellated BGO scintillator array under 68Ga excitation. The image in the firstcolumn was obtained by coupling the scintillator to the PSPMT detector, using only optical grease, the second by using 1 mm glass window the third using2 mm and the last 3 mm. (Second row): Normalized energy spectrum and 3 horizontal profiles (upper 2 PSPMTs - center dead area - lower 2 PSPMTs) ofthe image acquired with 2 mm glass window which provides the best better imaging performance.

mm optical glass) and 50 × 100 mm2 total dimensions(Amcrys, Kharkov, Ukraine) for SPECT applications. Thescintillator was coupled to the photodetector using onlyoptical grease and optical grease plus glass of 1 mmthickness.

E. Imaging Performance

The performance evaluation was based on previously pub-lished literature [2], [7], [29]. Energy resolution, detectionefficiency and sensitivity uniformity for the detector weremeasured with no collimator in place. A 50 uCi 99mTc pointsource was used to provide 140 keV irradiation. Similarly, a100 uCi 68Ga point source was used for 511 keV irradiation.The corresponding point source was placed 500 mm above thedetector in each set-up and each irradiation lasted 3 hours, inorder to acquire several thousands counts per pixel to minimizethe statistical error.

The full width at half maximum (FWHM) of the photopeakof the normalized energy spectrum was used to computeenergy resolution. Individual energy spectra were obtained foreach crystal in the array. For each individual crystal spectrum,the location of the photopeak was identified. The spectrawere then scaled so that their photopeaks were centered ona common channel, and summed. The resulting spectrum isreferred to as the normalized energy spectrum.

Detection efficiency was measured on the timed images,using a ±20% energy window. The ratio between the recordedcounts and the decay-corrected source emissions provides anestimate of the expected efficiency of the detector.

The sensitivity uniformity was calculated using the sameenergy window. Using the raw flood image a grid that mapseach crystal pixel was determined and the values in eachcrystal pixel were summed leading to a flood matrix. This floodimage was then smoothed processed with a 9-point smoothingfilter with the following weightings [1 2 1; 2 4 2; 1 2 1],

according to previous published quality control proceduresfor the evaluation of detectors non-uniformity [30], [31]. Themean value and standard deviation of the pixel counts in thecentral 75% of the image area were calculated. The ratiobetween the standard deviation and the mean was reportedas the detector sensitivity uniformity.

III. RESULTS

The results of the measurements are summarized in Table I.Figure 1 presents the raw images acquired with the discrete

TABLE IIMAGING PERFORMANCE

Scintillator En. Res. Det. Eff. Sen. Unif.

BGO glass 2mm 18.8 % 31 % 4 %

BGO glass 3mm 19.3 % 34.6 % 5.3 %

CsITl glass 3mm 19.6 % 41 % 12 %

CsITl glass 4mm 19.3 % 46 % 15 %

2 × 2 × 6 mm3 pixellated BGO scintillator array under68Ga excitation. The image in the first column was obtainedby coupling the scintillator to the PSPMT detector entrancewindow, using only optical grease, the second by using 1 mmultraviolet light transmitting glass window, the third using 2mm and the last 3 mm window, respectively. The same opticalgrease was used among all coupling surfaces. It can be seenthat using 2 and 3 mm glass windows one can resolve thecrystal elements of the scintillator in the dead areas of the flatpanel PMT array. The best imaging performance is achievedwith the 2 mm glass window (Table I). The normalized energyspectrum for this scheme along with 3 horizontal profiles(upper 2 PSPMTs - center dead area - lower 2 PSPMTs) areillustrated in the lower part of figure 1. Each scintillator pixelcorresponds to 19 pixels of the raw image.





























































































































































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Fig. 2. (First row): Raw images acquired with the discrete 2 × 2 × 4mm3 pixellated CsI:Tl scintillator array under 99mTc excitation. The image in thefirst column was obtained by coupling the scintillator to the PSPMT detector, using 1 mm glass window, the second using 2 mm, the third using 3 mm andthe last 4 mm. (Second row): Normalized energy spectrum and 3 horizontal profiles (upper 2 PSPMTs - center dead area - lower 2 PSPMTs) of the imageacquired with 3 mm glass window which provides the best imaging performance.

The average peak to valley ratio was measured equal to 2.5.We observed that as the glass thickness increases, the curveof the crystal elements is getting wider. Especially, an averageincrease of 0.2 mm in the FWHM is measured while using3 mm glass (1.42 mm) compared to coupling with 2 mmglass window (1.22 mm). As the effect on system’s spatialresolution needs to be considered, when uniformity at rawimages is ensured and central crystal elements are clearlyresolved, further increase of the thickness of the couplingmaterial should be avoided. The normalized energy resolutionwas approximately 18.8% FHWM at 511 keV. The average in-trinsic spatial resolution measured as described above, at threedifferent regions including a part of the dead area, was 1.22mm. The detection efficiency at 511 keV was measured equalto 31%. Finally, detector response non-uniformity (standarddeviation divided by mean) was ∼ 4%.

Figure 2 presents the raw images acquired with the discrete2 × 2 × 4 mm3 pixellated CsI:Tl scintillator array under99mTc excitation. The image in the first column was obtainedby coupling the scintillator to the PSPMT detector entrancewindow, using 1 mm ultra transmitting glass window, thesecond using 2 mm, the third using 3 mm and the last 4 mm.The same optical grease was used among all coupling surfaces.It can be seen that using 3 and 4 mm glass windows one canresolve the crystal elements of the scintillator in the dead areasof the flat panel PMT array. The best imaging performanceis achieved with the 3 mm glass window (Table I). Thenormalized energy spectrum for this scheme along with 3horizontal profiles (upper 2 PSPMTs - center dead area - lower2 PSPMTs) are illustrated in the lower part of figure 2. Eachscintillator pixel corresponds to 19 pixels of the raw image.The average peak to valley ratio was measured equal to 2.14.Once again, we observed that as the glass thickness increases,

the curve of the crystal elements is getting wider. Especially,an average increase of 0.14 mm in the FWHM is measuredwhile using 4 mm glass (1.44 mm) compared to couplingwith 3 mm glass window (1.3 mm). As the effect on system’sspatial resolution needs to be considered, when uniformity atraw images is ensured and central crystal elements are clearlyresolved, further increase of the thickness of the couplingmaterial should be avoided. The normalized energy resolutionwas approximately 19.6% FHWM at 140 keV. The detectionefficiency at 140 keV was measured equal to 41%. Finally,detector response non-uniformity (standard deviation dividedby mean) was 12%.

Figure 3 presents the raw images acquired with the discrete1.45×1.45×6 mm3 pixellated CsI:Na scintillator array under99mTc excitation. This scintillator has an intrinsic optical glasswindow 3.3 mm thick. The image in the first column wasobtained by coupling the scintillator to the PSPMT detectorentrance window, using only grease and the second using a 1mm thick extra optical glass window. It can be seen that forboth coupling schemes the crystal elements of the scintillatorin the central dead area of the flat panel PMT array couldnot be resolved. The normalized energy spectra for the centralarea of one PSPMT (Figure 3) for the two schemes along with3 horizontal profiles (upper 2 PSPMTs - center dead area -lower 2 PSPMTs) are illustrated in the lower part of figure 3.Each scintillator pixel corresponds to 15 pixels of the rawimage. The average peak to valley ratio for the two schemeswas measured equal to 1.44 and 1.23 respectively. It can beseen that clearer visualization of scintillator crystal elementswas achieved using only grease. The energy resolution wasmeasured equal to 26.6% and 28.9% respectively.

Previous work of our group contains the development ofa SPECT prototype system based on 4 H8500 PSPMTs in a





























































































































































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Fig. 3. (First row): Raw images acquired with the discrete 1.45× 1.45× 6mm3 pixellated CsI:Na scintillator array under 99mTc excitation. The scintillatorhas an intrinsic optical glass window 3.3 mm thick. The image in the first column was obtained by coupling the scintillator to the PSPMT detector, usingonly grease and the second using a 1 mm thick extra optical glass window and optical grease. (Second row): Normalized energy spectrum and 3 horizontalprofiles (upper 2 PSPMTs - center dead area - lower 2 PSPMTs) of the image acquired using only grease. (Third row): Normalized energy spectrum and 3horizontal profiles (upper 2 PSPMTs - center dead area - lower 2 PSPMTs) of the image acquired using the extra 1 mm thick optical glass window.

2 × 2 arrangement coupled via a 4 mm thick optical glassto a CsI:Na pixelated scintillator (1.5 × 1.5 × 6 mm3 ,Hilger UK). The 256 output signals of the four PSPMTswere reduced to 4 position signals (Xa, Xb, Yc, Yd) througha discretized position-sensitive readout circuit (DPC), similarto the circuit provided withe H13974-00-161. In that case, allcrystal elements are clearly distinguished as it is presented infigure 4 that illustrates the raw flood image acquired under99mTc excitation and the corresponding profile in one centralrow. Thus, we conclude that the results presented in figure 3are due to the low quality of the scintillator and not due tothe photodetector performance.

IV. DISCUSSION AND CONCLUSIONS

The H13974-00-1616 flat panel PMT array was evaluatedwith different scintillation schemes, and imaging performance,including energy resolution, detection efficiency and sensi-tivity uniformity are reported. This detector is proposed asan alternative to single large area PSPMTs, for examplethe R3292 from Hamamatsu that is no more commerciallyavailable, usually used for applications that require relativelylarge field of view.

The main objective of this study is to examine the behaviourof this recently commercially available photodetector, usinga simple four channel readout, in areas corresponding togaps between the tubes in which it may exhibit degradedperformance. We observed that the performance is stronglydepended on the scintillation material. In the case of BGOscintillator, the central crystal elements that corresponds to the

dead area of the detector are well distinguished when using a 2mm thick optical glass along with optical grease for coupling.The peak to valley ratio measured in the central row is equalto 2.3. On the other hand in the case of CsI:Tl scintillator theoptimal optical coupling (3 mm thick optical glass along withoptical grease) provides relatively poor peak to valley ratioin the central row equal to 1.75. However, the mean peakto valley ratio measured in three rows including the centralone is acceptable in both schemes (2.5 and 2.15 respectively).Moreover, despite the fact that these two scintillators havethe same pixel size, the light emitted from the central crystalelements is not fully collected in the case of the CsI:Tl result-ing in poor uniformity in contrast with the BGO scintillator.This result is possibly explained by the matching factor of thequantum efficiency of the detector with the emission spectrumof the BGO (maximum at 480 nm) and the CsI:Tl (maximumat 550 nm). Worse performance is observed for the CsI:Nascintillator. The central crystal elements are lost even withoptical glass 4 mm thick along with optical grease, while theenergy resolution is found to be greater than 25%. We believethat this result is due to the production method of the pix-elated scintillator and consequently its intrinsic performance.Theoritically, the emission spectrum of the CsI:Na (maximumat 420 nm) is matching better with the detector’s maximumquantum efficiency (maximum at 400 nm) than that of theCsI:Tl. Furthermore the pixel size is slightly smaller in thecase of CsI:Na, so we expected to distinguish the centralcrystal elemets using the appropriate coupling scheme. Onthe other hand, in a previous work of our group, the central





























































































































































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Fig. 4. (First row): Raw image acquired with discrete 1.5 × 1.5 × 6mm3

pixellated CsI:Na scintillator array under 99mTc excitation. The image wasobtained by coupling the scintillator to the 4 H8500 PSPMTs in 2 × 2arrangement, using 4 mm thick optical glass window and optical grease.(Second row): Horizontal profile in two central rows.

crystal elements of a CsI:Na pixellated scintillator producedby another company are well distinguished using 4 H8500PSPMTs in a 2× 2 arrangement for photodetection (figure 4).Taking into account that the two detectors are similar interms of photodetection and analog readout electronics, westrongly believe that similar performance may be obtainedwith the H13974 and a CsI:Na pixellated scintillator producedfollowing different method other than the one used for thescintillator used in the evaluation. Finally, it worth to mentionthat another group has reported that the minimum pixel sizethat they could detect with a 2×2 H8500 array is 2 mm. In thatstudy a LYSO pixelated scintillator (2×2 pixel size) is coupledto the photodetector using 3 mm and 4 mm borosilicate opticalglasses [32].

The analog circuit that reduces the anodes of the detector tofour position signals plays also significant role in the accuratedetermination of all crystal elements of the scintillator. De-pending on the values of the resistors and in combination withthe coupling scheme, not only the dead areas of the detectorcan be fully recovered, but also the intrinsic spatial resolution

can be increased when the peak to valley ratio is high. Thenumber of position signals of the final readout system and thenoise of the analog circuits affects the final performance ofthe detector [2], [4], [33].

The overall conclusion of this study is that the H13974-00-1616 flat panel PMT array is suitable for the developmentof large field of view nuclear imaging cameras for applica-tions in medium size animals imaging, such as rats or/andrabbits and in scintimammography. H13974-00-1616 offers acomplete solution, compensating the crucial technical issueof minimizing the effects of the dead spaces between theindividual PMTs in an array. The dead space (gap) of theH13974 series is only 3 mm, critical below in comparisonto in-house solutions based on H8500 PSPMTs or belowpotential custom solutions based on H12700 series. The arrayis equipped with a high voltage adjustment circuit for eachPSPMT, allowing the use of just one high voltage componentfor the whole array. Moreover, there is no need for an anodesgain variations compensation circuit, as the chosen PSPMTshave similar gain and anode non-uniformity. The scintillationmaterial, the coupling scheme and the analog readout circuitsare the main factors that affect the performance of the detectorin areas corresponding to gaps between the tubes. The propercombination of these factors should be carefully chosen so asto build a high performance gamma camera.

ACKNOWLEDGMENTS

The authors would like to thank Hamamatsu Photonicsfor providing the H13974 flat panel PMT array and for theexcellent collaboration.

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evaluation of hamamatsu h13974 large sensitive area flat

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