Search

A whole-body positron emission tomography scanner must be equipped with many detectors so that they are often composed of low-cost photomultiplier tubes (PMTs), that is, large anode-type PMTs. The general detector structure has a scintillation crystal element array coupled on an array of the large anode PMTs, and for crystal element identification, the method is needed to get sufficient spread of scintillation light for distribution among distant PMT anodes. Besides the common method of using a light guide, some methods have been proposed for better element identification performance. In this paper, we introduce a new method, in which the scintillation light spread is promoted not only by removing reflectors between crystal elements but also by restricting light exit to the PMTs by placing additional reflectors at the bottom of the crystal element array. Because the additional reflectors are parallel to the PMT surface, we call them parallel reflectors. We verified our method with a detector consisting of the 2.45 mm × 5 mm × 15 mm Lu2xGd2(1−x)SiO5:Ce crystal elements and two dualphotocathode PMTs. Each photocathode was 8 mm × 18 mm in size. We set a 9×10 crystal element array on the two PMTs and tried to identify the elements by the 2×2 PMT signals. Detector performance was evaluated with 137Cs point sources (662-keV gamma rays). The results showed that despite a significant decrease in light output at the boundary of the two PMTs, the method made crystal element identification possible. We measured energy resolutions of 13.1% and 17.8% for the elements on the PMT photocathode area and the PMT boundary, respectively. Our method is applicable to the depth-of-interaction detector consisting of multilayer crystal element arrays. By inserting parallel reflectors between the layers, we would be able to control the path of scintillation light originating in the upper layer. The effect was first examined in a basic study with two layers of a 2×5 crystal element array. Then, the array of each layer was increased to 9×9 for performance evaluation. It was difficult to identify all 162 crystal elements with four PMT signals; however, results suggest the possibility of scintillation light path control in each layer separately by the parallel reflectors and the potential for better performance by more precise adjustment of parallel reflectors.

The depth-of-interaction detector which can identify gamma-ray detection position along the depth direction is essential for a PET scanner to reduce parallax error. The X’tal cube we have developed is composed of a scintillation crystal block segmented into cubes. Six arrays of multipixel photon counters (MPPCs) cover six surfaces of the crystal block. We developed the new X’tal cube consisting of 0.77 mm cubic segments. A (13.1 mm)3 LYSO crystal block had 3-D segmentation in a 17 × 17 × 17 array fabricated by the laser engraving technique. 4 × 4 MPPC arrays with a size of (13 mm)2 were used. Active area of each MPPC was (3.0 mm)2. To evaluate segment identification performance, we irradiated with gamma-rays from a 22Na source and obtained a 3-D position histogram made by the 3-D Anger-type calculation with all MPPC signals. The histogram showed 17 segments in a row can be identified clearly. Further, to measure its intrinsic spatial resolution with incident angles of 0◦ and 45◦, gamma-rays in a fan-beam were irradiated and scanned onto the X’tal cube. Spatial resolutions of 0.74 ± 0.10 mm and 0.89 ± 0.10 mm were obtained by both scan experiments, respectively.

Depth of interaction (DOI) information is indispensable to improve sensitivity and resolution of positron emission tomography (PET) systems. We recently developed two dual-ended detectors with DOI resolutions of 3 mm and 1.5 mm using crystal bars of 3×3×20 mm3 and 1.5×1.5×20 mm3, respectively, segmented by applying the subsurface laser engraving (SSLE) technique to 7 and 13 DOI segments. However, the spatial resolution of our detector needs to be improved for a small animal PET scanner with submillimeter resolution. In this work, we developed two types of dual-ended detectors; one consisted of crystal bars of 1×1×20 mm3 (with 1.67 mm DOI resolution) with readout along the radial direction and the other consisted of crystal bars of 2×1×20 mm3 (with segment pitch of 1 mm) with the readout along the axial direction. The crystal bars were segmented in the height direction by applying the SSLE technique. We evaluated the performances of two detectors from the viewpoint of the best crystal identification; the first was composed of one 8×8 crystal array made of 1×1×20 mm3 crystal bars (with 12 segments and segment pitch of 1.67 mm) and the second one was composed of one 4×8 crystal array made of 2×1×20 mm3 crystal bars (with 20 segments and segment pitch of 1 mm). The 3D position maps of the detectors were obtained by the Anger type calculation and the crystal identification performance was evaluated for the crystal arrays. Clear separation was obtained for both crystal arrays with 12 and 20 segments with segment resolutions of 1.67 mm and 1 mm, respectively. The energy resolutions were also evaluated for both crystal arrays. The energy resolutions of 8.0% ± 0.6% and 9.0% ± 0.7% were obtained for the segments at the middle of the 8×8 and 4×8 crystal arrays, respectively., The 2019 IEEE Nuclear Science Symposium (NSS) and Medical Imaging Conference (MIC)

We have developed the X’tal cube PET detector which ‘ is composed of 96 multi pixel photon counters (MPPCs) and a (13.1 mm)3 LYSO scintillator segmented into the 17 x 17 x 17 array. Each segment size is (0.77 mm)3. Six 4 x 4 MPPC arrays cover all six surfaces of the LYSO scintillator. The Anger calculation from 96 MPPC signals is used to determine the segment detecting a gamma-ray. We confirmed that the X’tal cube achieved sub-millimeter isotropic spatial resolution. However, we also observed inter crystal scattering (ICS) events as background events. There is a concern that ICS events degrade contrast of PET images for the high spatial resolution PET scanner. \n We proposed a method to discriminate the ICS events and applied it to the X’tal cube. If the photo absorption event occurs at one segment, the distribution of spread light on each MPPC is the same for every event. On the other hand, even if the result of the Anger calculation is the same as the photo absorption event, the distribution of spread light by the ICS events differs from that of the photo absorption event because scintillation light is generated at different segments. To discriminate the ICS events, first using a uniform irradiation experiment, we obtained the mean and standard deviation of pulse height peak of each MPPC for the segment to set the threshold for the discrimination of the ICS events. Then for each detected event, each MPPC pulse height was measured and the chi-square value was calculated. If the chi-square value exceeded the threshold, the event was eliminated as an ICS event. Next, to prove that our method can eliminate the ICS events, we obtained the response function by a fan-beam scanning experiment. We compared the results before and after applying our method. At the central part of the X’tal cube, full width half maximum of the ICS distribution in the response function was improved from 4.0 mm to 2.1 mm and over 50% of the ICS events were eliminated., IEEE NSS-MIC2018

The accumulation of induced radioactivity within in-beam PET scanner scintillators is of concern for its long-term clinical usage in particle therapy. To estimate the effects on OpenPET which we are developing for in-beam PET based on GSOZ (Zi doped Gd2SiO5), we measured the induced radioactivity of GSO activated by secondary fragments in a water phantom irradiation by a 12C beam with an energy of 290 MeV u−1. Radioisotopes of Na, Ce, Eu, Gd, Nd, Pm and Tb including positron emitters were observed in the gamma ray spectra of the activated GSO with a high purity Ge detector and their absolute radioactivities were calculated. We used the Monte Carlo simulation platform, Geant4 in which the observed radioactivity was assigned to the scintillators of a precisely reproduced OpenPET and the single and coincidence rates immediately after one treatment and after one-year usage were estimated for the most severe conditions. Comparing the highest coincidence rate originating from the activated scintillators (background) and the expected coincidence rate from an imaging object (signal), we determined the expected signal-tonoise ratio to be more than 7 within 3 min and more than 10 within 1 min from the scan start time. We concluded the effects of scintillator activation and their accumulation on the OpenPET imaging were small and clinical long-term usage of the OpenPET was feasible. Keywords: in-beam PET, GSO scintillator, particle therapy, openPET, induced radioactivity, secondary fragments (Some figures may appear in colour only in the online journal)

In-beam positron emission tomography (PET) can enable visualization of an irradiated field using positron emitters (β+ decay). In particle therapies, many kinds of secondary particles are produced by nuclear interactions, which affect PET imaging. Our purpose in this work was to evaluate effects of secondary particles on in-beam PET imaging using the Monte Carlo simulation code, Geant4, by reproducing an experiment with a small OpenPET prototype in which a PMMA phantom was irradiated by a 11C beam. The number of incident particles to the detectors and their spectra, background coincidence for the PET scan, and reconstructed images were evaluated for three periods, spill-time (beam irradiation), pause-time (accelerating the particles) and beamoff time (duration after the final spill). For spill-time, we tested a background reduction technique in which coincidence events correlated with the accelerator radiofrequency were discarded (RF gated) that has been proposed in the literature. Also, background generation processes were identified. For spilltime, most background coincidences were caused by prompt gamma rays, and only 1.4% of the total coincidences generated β+ signals. Differently, for pause-time and beam-off time, more than 75% of the total coincidence events were signals. Using these coincidence events, we failed to reconstruct images during the spill-time, but we obtained successful reconstructions for the pausetime and beam-off time, which was consistent with the experimental results. From the simulation, we found that the absence of materials in the beam line and using the RF gated technique improved the signal-to-noise ratio for the spill-time. From an additional simulation with range shifter-less irradiation and the RF gated technique, we showed the feasibility of image reconstruction during the spill-time.

Positron emission tomography (PET) has become a popular imaging method in metabolism, neuroscience, and molecular imaging. For dedicated human brain and small animal PET scanners, high spatial resolution is needed to visualize small objects. To improve the spatial resolution, we are developing the X'tal cube, which is our new PET detector to achieve isotropic 3D positioning detectability. We have shown that the X'tal cube can achieve 1 mm3 uniform crystal identification performance with the Anger-type calculation even at the block edges. We plan to develop the X'tal cube with even smaller 3D grids for sub-millimeter crystal identification. In this work, we investigate spatial resolution of a PET scanner based on the X'tal cube using Monte Carlo simulations for predicting resolution performance in smaller 3D grids. For spatial resolution evaluation, a point source emitting 511 keV photons was simulated by GATE for all physical processes involved in emission and interaction of positrons. We simulated two types of animal PET scanners. The first PET scanner had a detector ring 14.6 cm in diameter composed of 18 detectors. The second PET scanner had a detector ring 7.8 cm in diameter composed of 12 detectors. After the GATE simulations, we converted the interacting 3D position information to digitalized positions for realistic segmented crystals. We simulated several X'tal cubes with cubic crystals from (0.5 mm)3 to (2 mm)3 in size. Also, for evaluating the effect of DOI resolution, we simulated several X'tal cubes with crystal thickness from (0.5 mm)3 to (9 mm)3. We showed that sub-millimeter spatial resolution was possible using cubic crystals smaller than (1.0 mm)3 even with the assumed physical processes. Also, the weighted average spatial resolutions of both PET scanners with (0.5 mm)3 cubic crystals were 0.53 mm (14.6 cm ring diameter) and 0.48 mm (7.8 cm ring diameter). For the 7.8 cm ring diameter, spatial resolution with 0.5*0.5*1.0 mm3 crystals was improved 39% relative to the (1 mm)3 cubic crystals. On the other hand, spatial resolution with (0.5 mm)3 cubic crystals was improved 47% relative to the (1 mm)3 cubic crystals. The X'tal cube promises better spatial resolution for the 3D crystal block with isotropic resolution.

We are developing a novel, general purpose isotropic-3D PET detector X'tal cube which has high spatial resolution in all three dimensions. The research challenge for this detector is implementing effective detection of scintillation photons by covering six faces of a segmented crystal block with silicon photomultipliers (SiPMs). In this paper, we developed the second prototype of the X'tal cube for a proof-of-concept. We aimed at realizing an ultimate detector with 1.0 mm^3^ cubic crystals, in contrast to our previous development using 3.0 mm^3^ cubic crystals. The crystal block was composed of a 16 x 16 x 16 array of lutetium gadolinium oxyorthosilicate (LGSO) crystals 0.993 x 0.993 x 0.993 mm^3^ in size. The crystals were optically glued together without inserting any reflector inside and 96 multi-pixel photon counters (MPPCs, S10931-50P, i.e. six faces each with a 4 x 4 array of MPPCs), each having a sensitive area of 3.0 x 3.0 mm^2^, were optically coupled to the surfaces of the crystal block. Almost all 4096 crystals were identified through Anger-type calculation due to the finely adjusted reflector sheets inserted between the crystal block and light guides. The reflector sheets, which formed a belt of 0.5 mm width, were placed to cover half of the crystals of the second rows from the edges in order to improve identification performance of the crystals near the edges. Energy resolution of 12.7% was obtained at 511 keV with almost uniform light output for all crystal segments thanks to the effective detection of the scintillation photons.

The OpenPET geometry is our new idea to visualize a physically opened space between two detector rings. In this paper, we developed the first small prototype to show a proof-of-concept of OpenPET imaging. Two detector rings of 110 mm diameter and 42 mm axial length were placed with a gap of 42 mm. The basic imaging performance was confirmed through phantom studies; the open imaging was realized at the cost of slight loss of axial resolution and 24% loss of sensitivity. For a proof-of-concept of PET image-guided radiation therapy, we carried out the in-beam tests with 11C radioactive beam irradiation in the heavy ion medical accelerator in Chiba to visualize in situ distribution of primary particles stopped in a phantom. We showed that PET images corresponding to dose distribution were obtained. For an initial proof-of-concept of real-time multimodal imaging, we measured a tumor-inoculated mouse with 18F-FDG, and an optical image of the mouse body surface was taken during the PET measurement by inserting a digital camera in the ring gap. We confirmed that the tumor in the gap was clearly visualized. The result also showed the extension effect of an axial field-of-view (FOV); a large axial FOV of 126 mm was obtained with the detectors that originally covered only an 84 mm axial FOV. In conclusion, our initial imaging studies showed promising performance of the OpenPET.

The jPET-D4 is a brain positron emission tomography (PET) scanner that we have developed to meet user demands for high sensitivity and high spatial resolution. For this scanner, we developed a four-layer depth-of-interaction (DOI) detector. The four-layer DOI detector is a key component for the jPET-D4, its performance has great influence on the overall system performance. Previously, we reported the original technique for encoding four-layer DOI. Here, we introduce the final design of the jPET-D4 detector and present the results of an investigation on uniformity in performance of the detector. The performance evaluation was done over the 120 DOI crystal blocks for the detectors, which are to be assembled into the jPET-D4 scanner. We also introduce the crystal assembly method, which is simple enough, even though each DOI crystal block is composed of 1,024 crystal elements. The jPET-D4 detector consists of four layers of 16 × 16 Gd2SiO5 (GSO) crystals and a 256-channel flat-panel position-sensitive photomultiplier tube (256ch FP-PMT). To identify scintillated crystals in the four-layer DOI detector, we use pulse shape discrimination and position discrimination on the two-dimensional (2D) position histogram. For pulse shape discrimination, two kinds of GSO crystals that show different scintillation decay time constants are used in the upper two and lower two layers, respectively. Proper reflector arrangement in the crystal block then allows the scintillated crystals to be identified in these two-layer groupings with two 2D position histograms. We produced the 120 DOI crystal blocks for the jPET-D4 system, and measured their characteristics such as the accuracy of pulse shape discrimination, energy resolution, and the pulse height of the full energy peak. The results show a satisfactory and uniform performance of the four-layer DOI crystal blocks; for example, misidentification rate in each GSO layer is

We are developing a small animal PET scanner, "jPET-RD" to achieve high sensitivity as well as high spatial resolution by using four-layer depth-of-interaction (DOI) detectors. The jPET-RD is designed with two detector rings. Each detector ring is composed of six DOI detectors arranged hexagonally. The diameter of the field-of-view (FOV) is 8.8 cm, which is smaller than typical small animal PET scanners on the market now. Each detector module consists of a crystal block and a 256-channel flat panel position-sensitive photomultiplier tube. The crystal block, consisting of 32 x 32 x 4 crystal (4096 crystals, each 1.46 mm x 1.46 mm x 4.5 mm) and a reflector, is mounted on the 256ch FP-PMT. \nIn this study, we evaluated the spatial resolution of reconstructed images with the evaluation system of two four-layer DOI detectors which consist of 32 x 32 x 4 LYSO (Lu: 98%, Y: 2%) crystals coupled on the 256ch FP-PMT by using RTV rubber. The spatial resolution of 1.5 mm was obtained at the center of the FOV by the filtered back projection. The spatial resolution, better than 2 mm in the whole FOV, was also achieved with DOI while the spatial resolution without DOI was degraded to 3.3 mm.