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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 $\times \,\, 5$ mm $\times \,\,15$ mm Lu2 x Gd2(1– x )SiO5:Ce crystal elements and two dual-photocathode PMTs. Each photocathode was 8 mm $\times \,\,18$ mm in size. We set a $9 \times 10$ crystal element array on the two PMTs and tried to identify the elements by the $2\times 2$ 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 $2 \times 5$ crystal element array. Then, the array of each layer was increased to $9\times 9$ 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 ${17 \times 17 \times 17}$ array fabricated by the laser engraving technique. $ {4 \times 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.

Sensitivity and spatial resolution of positron emission tomography (PET) scanners can be improved by using thicker scintillation crystals with depth-of-interaction (DOI) encoding. Subsurface laser engraving (SSLE) can be used to segment crystals of a scintillation detector in order to fabricate a DOI detector. We previously applied SSLE to crystal bars of 3 × 3 × 20 mm3and 1.5 × 1.5 × 20 mm3and developed two dual-ended detectors with DOI segments of 3 mm and 1.5 mm, respectively. To further improve the DOI resolution, our SSLE detector design can be used with smaller pitch crystal bars, making them excellent detector candidates for small animal PET scanners with submillimetre resolution. In the present study, three small crystal bars of 1 × 1 × 20 mm3, 2 × 1 × 20 mm3, and 2 × 1 × 40 mm3were laser engraved to 12, 20 and 40 segments, respectively, by applying SSLE in their height directions. The segmented crystal bars were characterised in three prototype detector arrangements. First, the 1 × 1 × 20 mm3crystal bars were characterised in an 8 × 8 crystal array designed for DOI encoding along crystal height in a conventional small animal PET design. Second, a 4 × 8 crystal array of 2 × 1 × 20 mm3crystal bars was characterised for using the DOI information for crystal interaction positioning along the axial axis of a small animal PET scanner. Finally, the third part of the study was performed on a single 2 × 1 × 40 mm3crystal bar with 40 segments to investigate the feasibility of DOI estimation in longer crystals for application in a system with extended axial length. We evaluated the capability of segment identification and energy resolution of theses detectors. The 3D position maps of the detectors were obtained using the Anger-type calculation and the crystal identification performance was evaluated for each detector. Clear segment separation was obtained for the crystal arrays with 12 (segment pitch of 1.67 mm) and 20 (segment pitch of 1 mm) segments. Mean energy resolutions of 8.8% ± 0.4% and 9.6% ± 0.8% at 511 keV were obtained for the segments in the central regions of the 8 × 8 array with 12 segments and the 4 × 8 array with 20 segments, respectively. Clear segment identification was found to be difficult for the detector with 40 segments, especially for the segments at the middle of the crystal. Energy and interaction positioning characterisation results suggest that both prototype detectors with 12 and 20 segments are well suited for small animal PET scanners with high spatial resolution.

Detectors with depth-of-interaction (DOI) information are utilized in positron emission tomography (PET) scanners to improve the sensitivity of radiation detection and the uniformity of spatial resolution. We recently developed a series of dual-ended detectors using crystal bars segmented by applying the subsurface laser engraving (SSLE) technique to 7, 13 and 20 DOI segments. It is crucial to achieve the submillimeter level spatial resolution for our detector as the narrow crystal bars are highly fragile after applying the SSLE technique. In this work, we focused on optimizing the surface area of the SSLE-induced layer to the narrow crystal bars in order to resolve the issue of fragility while maintaining sufficient crystal segment separation. The SSLE layers were induced to the middle of the cross section of the five crystal bars with dimensions of 1.5 × 1 . 5 × 20 mm3 with a distance of 0.1, 0.2, 0.3, 0.4 or 0.5 mm from the two opposite lateral edges of each crystal bar. All crystals were segmented into four DOI segments and the performances of the five sets of dual-ended DOI detectors consisting of one crystal bar were evaluated from the viewpoints of optimum crystal segment identification and energy resolution while mitigating the fragility issue. The 3D position maps of the detectors were obtained by the Anger-type calculation, and the crystal identification performance was evaluated. Clear separation of the segments with a distance-to-width ratio (DWR) larger than 7 was obtained for the detectors in which the distances of the SSLE-induced layer from the two opposite lateral edges of the crystal were 0.1, 0.2 and 0.3 mm. The DWR for the other two detectors was less than 4.4. Average energy resolutions of 11.8 ± 0.2%, 10.6 ± 0.2% and 8.8 ± 0.2% were obtained for the detectors in which the SSLE layers were induced with 0.1, 0.3 and 0.5 mm distances from the two opposite lateral edges of the crystal, respectively. There was a compromise between energy resolution and crystal segment identification. The optimum surface area of the SSLE-induced layer was accepted as a trade-off between crystal segment identification and fragility. The crystal segmented by the SSLE-induced layer at the middle of the cross section with a distance of 0.3 mm from the two opposite lateral edges was considered to be a good candidate with acceptable performance and fragility.

We have developed a four-layer depth of interaction (DOI) detector with single-side photon readout, in which segmented crystals with the patterned reflector insertion are separately identified by the Anger-type calculation. Optical conditions between segmented crystals, where there is no reflector, affect crystal identification ability. Our objective of this work was to improve crystal identification performance of the four-layer DOI detector that uses crystals segmented with a recently developed laser processing technique to include laser processed boundaries (LPBs). The detector consisted of 2 × 2 × 4 mm 3 LYSO crystals and a 4 × 4 array multianode photomultiplier tube (PMT) with 4.5 mm anode pitch. The 2D position map of the detector was calculated by the Anger calculation method. At first, influence of optical condition on crystal identification was evaluated for a one-layer detector consisting of a 2 × 2 crystal array with three different optical conditions between the crystals: crystals stuck together using room temperature vulcanized (RTV) rubber, crystals with air coupling and segmented crystals with LPBs. The crystal array with LPBs gave the shortest distance between crystal responses in the 2D position map compared with the crystal array coupled with RTV rubber or air due to the great amount of cross-talk between segmented crystals with LPBs. These results were used to find optical conditions offering the optimum distance between crystal responses in the 2D position map for the four-layer DOI detector. Crystal identification performance for the four-layer DOI detector consisting of an 8 × 8 array of crystals segmented with LPBs was examined and it was not acceptable for the crystals in the first layer. The crystal identification was improved for the first layer by changing the optical conditions between all 2 × 2 crystal arrays of the first layer to RTV coupling. More improvement was observed by combining different optical conditions between all crystals of the first layer and some crystals of the second and the third layers of the segmented array.

A small animal positron emission tomography (PET) scanner employs a scintillation crystal with a small cross section to improve the spatial resolution. However, the identification of crystals becomes more difficult as the cross section of the crystal is reduced. In particular, the identification of edge crystals is more difficult than the central crystals. The aim of this study is to optimize a light guide design of a small animal PET detector to improve the resolvability of the edge crystals using GATEv6.2 optical simulation. The small animal PET module consists of an 11 × 11 array of pixelated LYSO crystals (0.92×0.92×10.0 mm3), a 1 mm thick light guide, and a 4 × 4 array SiPM (Hamamatsu Photonics, Japan; S13361-3050NE-04; each channel has an effective area of 3 × 3 mm2). The optical characteristics of LYSO and the SiPM were taken into account for the GATEv6.2 optical simulation. In order to improve the crystal resolvability in the edge and corner regions, four air slits were inserted into the light guide. The flood map was obtained with various slit depths of 0.2, 0.5, and 0.8 mm, respectively. The identification of the corner crystals could be improved by using a combination of a light guide thickness of 1 mm, and a slit depth of 0.5 mm without compromising energy resolution. In future, the optical simulation results will be validated by an experimental measurement.

X’tal cube is the cubic depth of interaction (DOI) PET detector which our research group developed. In this work, aiming to get higher sensitivity, we developed the long rectangular shape X’tal cube (long-XC) by extending the cubic X’tal cube structure in one direction. We verified performance of this long-XC and also studied detector parameters for optimization. The same as the X’tal cube, the crystal block of the long-XC is composed of a 3D array of cubic scintillation crystal elements. Reflectors are not inserted between these crystal elements. The scintillation light then spreads without being obstructed by reflectors and is detected by multiple numbers of the multi-pixel photon counters (MPPCs) coupled on all six sides of the crystal block. For crystal element identification, a simple Anger-type calculation is used. In this study, we arranged 3.0 mm $\times 3.0$ mm $\times 3.0$ mm LGSO crystal elements in a $6\times 6\times 14$ array for the long-XC. In a previous study, we had already confirmed that for the X’tal cube consisting of a $6\times 6\times 6$ array of the same crystal elements and 54 MPPCs, identification of all 216 crystal elements was possible and the average energy resolution for all the elements was about 11 %. The long-XC contains more than twice the number of the crystal elements but less than twice the number of the MPPCs compared to the previous X’tal cube. The detector parameters investigated with the long-XC were: the number of MPPCs on both sides in the extended direction (edge MPPCs); the MPPC type, the MPPCs of $25\,\,\mu \text {m}\,\,\times 25\,\,\mu \text {m}$ or $50\,\,\mu \text {m}\,\,\times 50\,\,\mu \text {m}$ pixel sizes; the material between the crystal elements, an air gap or an optical glue having a closer refractive index to that of LGSO than air has; and the MPPC signals used in the Anger-type calculation. Results of the crystal element identification performance showed that reducing the number of the edge MPPCs caused performance degradation only at the part near the edge. For the MPPC type, the $50~\mu \text {m}$ type was better than the $25~\mu \text {m}$ type, and for the material, air was much better than the optical glue. We found that the choice of MPPC signals for the Anger-type calculation was effective in the optical glue condition. For the long-XC in the air gap condition and using the $50~\mu \text {m}$ type MPPCs, we observed it had good performance and there was no significant degradation at the central part which is far from the edge MPPCs. For irradiation of 662 keV gamma-rays, we measured approximately 11 - 13 % energy resolution for each crystal element and there was only a small difference in light outputs between crystal elements at the central part and at the edges. These results at the central part suggested the possibility of further extension of the long-XC for higher sensitivity.

Recently, various types of PET-MRI systems have been developed by a number of research groups. However, almost all of the PET detectors used in these PET-MRI systems have no depth-of-interaction (DOI) capability. The DOI detector can reduce the parallax error and lead to improvement of the performance. We are developing a new PET-MRI system which consists of four-layer DOI detectors positioned close to the measured object to achieve high spatial resolution and high scanner sensitivity. As a first step, we are investigating influences the PET detector and the MRI system have on each other using a prototype four-layer DOI-PET detector. This prototype detector consists of a lutetium yttrium orthosilicate crystal block and a 4 × 4 multi-pixel photon counter array. The size of each crystal element is 1.45 mm × 1.45 mm × 4.5 mm, and the crystals are arranged in 6 × 6 elements × 4 layers with reflectors. The detector and some electric components are packaged in an aluminum shielding box. Experiments were carried out with 3.0 T MRI (GE, Signa HDx) and a birdcage-type RF coil. We demonstrated that the DOI-PET detector was normally operated in simultaneous measurements with no influence of the MRI measurement. A slight influence of the PET detector on the static magnetic field of the MRI was observed near the PET detector. The signal-to-noise ratio was decreased by presence of the PET detector due to environmental noise entering the MRI room through the cables, even though the PET detector was not powered up. On the other hand, no influence of electric noise from the PET detector in the simultaneous measurement on the MRI images was observed, even though the PET detector was positioned near the RF coil.

We have developed a 4-layer depth-of-interaction (DOI) detector which consists of four layers of scintillation crystal arrays and a position sensitive photomultiplier tube (PS-PMT). To control the behavior of scintillation light in each DOI crystal array, some reflectors between crystals are removed so that all crystal responses in the four layers are expressed in one two-dimensional (2D) position histogram by implementing an Anger-type calculation of the PS-PMT signals. Since the method utilizes spread of scintillation light through the boundary of the crystals with no reflectors, positioning performance in the 2D position histogram depends on the crystal dimensions, the crystal surface finish, and gap materials. In this work, we propose an adjustment method for crystal identification performance of the 4-layer DOI detector by selecting the optimal refractive index for the gap materials between the crystals in which the reflectors are removed. We fabricated single-layer detectors and 4-layer detectors using LYSO scintillators and evaluated the crystal identification performance, while varying the refractive index of the gap materials (by using different optical cements). As a result, we found the 2D position histogram changes as a function of the refractive index of the optical cement between crystals and the optimal refractive indices can be determined for the detector using the LYSO crystals. We concluded that the crystal response of the 2D position histograms can be adjusted and crystal identification performance can be optimized by changing the refractive index of the gap materials.

We are developing a PET-MRI system which consists of PET detectors integrated with the head coil of the MRI in order to realize high spatial resolution and high sensitivity in simultaneous measurements. In the PET-MRI system, the PET detectors which consist of a scintillator block, photo-detectors and front-end circuits with four-layer depth-of-interaction (DOI) encoding capability are placed close to the measured object. Therefore, the proposed system can achieve high sensitivity without degradation of spatial resolution at the edge of the field-of-view due to parallax error thanks to the four-layer DOI capability. In this paper, we fabricated a prototype system which consists of a prototype four-layer DOI-PET detector, a dummy PET detector and a prototype birdcage type head coil. Then we used the prototype system to evaluate the performance of the four-layer DOI-PET detector and the reciprocal influence between the PET detectors and MRI images. The prototype DOI-PET detector consists of six monolithic multi-pixel photon counter (MPPC) arrays (S11064-050P), a readout circuit board, two scintillator blocks and a copper shielding box. Each scintillator block consists of four layers of Lu1.8Gd0.2SiO5:Ce (LGSO) scintillators and reflectors are inserted between the scintillation crystals. The dummy detector has all these components except the two scintillator blocks. The head coil is dedicated to a 3.0 T MRI (MAGNETOM Verio, Siemens) and the two detectors are mounted in gaps between head coil elements. Energy resolution and crystal identification performance of the prototype four-layer DOI-PET detector were evaluated with and without MRI measurements by the gradient echo and spin echo methods. We identified crystal elements in all four layers from a 2D flood histogram and energy resolution of 15–18% was obtained for single crystal elements in simultaneous measurements. The difference between the average energy resolutions and photo-peak positions with and without MRI measurements was lower than 0.3 percentage points and 1.7% for all layers. The results indicated that these performances were sufficient as PET detectors for the proposed PET-MRI system and there was no influence from the MRI measurements on the PET imaging in the simultaneous measurements. The signal-to-noise ratio of the MRI image and static magnetic field of the MRI were also evaluated with and without measurements of the PET detectors. The maximum decrease of the static magnetic field due to the LGSO scintillators was approximately 1 ppm. The signal-to-noise ratio decreased from 242.80 without the PET detector to 44.948 in simultaneous measurements.

We are developing a novel PET detector with 3D isotropic resolution called a crystal (X'tal) cube. The X'tal cube detector consists of a crystal block all 6 surfaces of which are covered with silicon photomultipliers (SiPMs). We have developed a prototype detector with 3D isotropic 1 mm resolution. On the other hand, when the X'tal cubes are arranged to form a PET scanner, insensitive inter-detector gaps made by the SiPM arrays should not be too wide, or, better yet, they should be removed. Reduction of the number of SiPMs will also be reflected in the production costs. Therefore, reducing the number of faces to be connected to the SiPMs has become our top priority. In this study, we evaluated the effect of reducing the number of SiPMs on the positioning accuracy through numerical simulations. Simulations were performed with the X'tal cube, which was composed of a 6 × 6 × 6 array of Lu2x Gd2(1-x)SiO5:Ce crystal elements with dimensions of (3.0 mm)(3). Each surface of the crystal block was covered with a 4 × 4 array of SiPMs, each of which had a (3.0 mm)(2) active area. For material between crystal elements, we compared two: optical glue and an air gap. The air gap showed a better crystal identification performance than did the optical glue, although a good crystal identification performance was obtained even with optical glue for the 6-face photodetection. In conclusion, the number of photodetection faces could be reduced to two when the gap material was air.

We have developed a three-dimensional (3-D) position-sensitive radiation detector named X'tal cube, which can be applied to a PET detector. The X'tal cube is composed of a scintillation crystal block and multi-pixel photon counters (MPPCs). The crystal block is segmented three-dimensionally into small cubes by optical discontinuity (3-D segmentation) and no reflector is inserted inside. Scintillation light originating in a segment then spreads three-dimensionally so that the MPPCs are set on all six surfaces of the crystal block to detect the light. Regarding the 3-D segmentation of the crystal block, we have already succeeded in getting the segmentation inside of a monolithic scintillator block by a laser processing technique instead of the general way, arranging small cubic scintillator elements into a 3-D array. We have confirmed that utilizing the laser processing technique not only eliminates the difficulty of handling the small scintillator elements but also improves detector performance. As a new trial, we considered fabrication of the crystal block by stacking the scintillator plates which were segmented two-dimensionally by the laser processing technique (2-D segmentation). Plates are also easier to handle and for the laser processing, 2-D segmentation is simpler than 3-D segmentation. In this study, we prepared the X'tal cube with the scintillator plates (Plate-XC) and evaluated its performance to confirm its technical feasibility. For the Plate-XC, we segmented 18 mm × 18 mm × 2.0 mm LYSO plates two-dimensionally by laser processing so as to make a 9 × 9 array of 2.0 mm × 2.0 mm segments. The crystal block was composed of 9 stacked LYSO plates without using coupling material but with air gaps. The Plate-XC showed sufficient crystal identification performance when 662 keV gamma-rays were irradiated. Furthermore, to understand the characteristics of the Plate-XC, we also analyzed the scintillation light distribution in the crystal block. Results indicated that light spread from outer segments was influenced by the segment boundary conditions, air gaps or laser-processed gaps, while the spread from the center segment did not seem to have such an influence. Regarding energy performance, we obtained around 10% energy resolution for the outer segments as well as for the center segment.

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 mm 3 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 mm 3 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.

Silicon photomultipliers (SiPMs) provide many advantages for PET detectors, such as their high internal gain, high photon detection efficiency and insensitivity to magnetic fields. The number of detectable scintillation photons of SiPMs, however, is limited by the number of microcells. Therefore, pulse height of PET detectors using SiPMs is saturated when large numbers of scintillation photons enter the SiPM pixels. On the other hand, we previously presented a depth-of-interaction (DOI) encoding method that is based on the light sharing method. Since our encoding method detects scintillation photons with multiple readout pixels, the saturation effect can be suppressed. We constructed two prototype four-layer DOI detectors using a SiPM array and evaluated their performances. The two prototype detectors consisted of four layers of a 6×6 array of Lu 2(1− x ) Y 2 x SiO 5 (LYSO) crystals and a SiPM (multi-pixel photon detector, MPPC, Hamamatsu Photonics K.K.) array of 4×4 pixels. The size of each LYSO crystal element was 1.46 mm×1.46 mm×4.5 mm and all surfaces of the crystal elements were chemically etched. We used two types of MPPCs. The first one had 3600 microcells and high photon detection efficiency (PDE). The other one had 14,400 microcells and lower PDE. In the evaluation experiment, all the crystals of the detector using the MPPC which had the high PDE were clearly identified. The respective energy and timing resolutions of lower than 15% and 1.0 ns were achieved for each crystal element. No saturation of output signals was observed in the 511 keV energy region due to suppression of the saturation effect by detecting scintillation photons with several MPPC pixels by the light sharing method.

The X'tal cube is our original PET detector, which is being developed to achieve isotropic 3D positioning detectability. The X'tal cube is based on a 3D segmented crystal block for which all surfaces are covered with photo-detectors. Previously, instead of our initial approach of gluing segmented pieces of crystals, we successfully constructed a crystal block segmented by laser processing, and we developed the X'tal cube with the laser-processed 3D square grids of 2 mm length. In this paper, we extend the laser processing to 3D square grids of 1 mm length. The volume of a 1-mm crystal segment is 1/8 of that of a 2-mm crystal segment. We also evaluate imaging resolution performance with a newly developed one-pair prototype system to simulate a ring-type scanner, while our previous reports were limited to development and evaluation of a single detector. In particular, we compare 1-mm X'tal cube detectors to 2-mm X'tal cube detectors. The one-pair prototype system consisted of two X'tal cubes, two rotating stages, and a 192-channel data acquisition system. Each X'tal cube consisted of the LYSO cubic crystal block of 18 × 18 × 18 mm3 in which the 3D grids of 1 mm pitch were fabricated by internal laser processing. The 4 × 4 arrays of multi pixel photon counters were optically coupled to each surface of the crystal block. The detector positions were automatically controlled to simulate a ring-type PET with a 14.6 cm diameter. Data were collected for all assumed detector positions and then a sinogram was obtained. The data were reconstructed using filtered backprojection. The average spatial resolution of the 2-mm X'tal cube was 1.9 mm full width at half maximum (FWHM) over the field-of-view (FOV). On the other hand, the average spatial resolution of the 1-mm X'tal cube was 1.3 mm FWHM over the FOV. By applying deconvolution with the assumption that the point source was a Gaussian function of 1.0 mm FWHM, we estimated the average spatial resolution of the 2-mm X'tal cube and the 1-mm X'tal cube as 1.6 mm FWHM and 0.83 mm FWHM, respectively. The average spatial resolution of the 1-mm X'tal cube was improved approximately 48% compared to the 2-mm X'tal cube. In conclusion, we confirmed the potential of the X'tal cube for uniform and high resolution imaging.

A conventional PET scanner has a 15–25 cm axial field-of-view (FOV) and images the whole body using about six bed positions. We proposed our original OpenPET geometry which can extend the axial FOV with a limited number of detectors. For an alternative approach, we then designed an axially extendable multiplex cylinder (AEMC) PET scanner to provide high versatility for clinical and research studies using a silicon-photomultiplier (Si-PM)-based DOI detector. Since Si-PMs have high gain like PMTs and a compact design, the Si-PM-based detector is expected to enable various new detector arrangements. The AEMC-PET scanner consists of four independent and laminated detector rings using a four-layer depth-of-interaction (DOI) detector. This PET scanner can extend the axial FOV as each stacked detector ring can be slid aside. When this scanner is used for the four-layer DOI detector, its minimum axial FOV is 24 cm and its maximum crystal thickness is 3 cm. On the other hand, the axial FOV can be extended to 96 cm while keeping a continuous axial FOV, but the crystal thickness must be 1/4 of 3 cm. In addition, using the OpenPET geometry, the axial FOV can be extended to 168 cm. In this work, we studied the feasibility of the PET scanner with a variable axial FOV using Monte Carlo simulation. For the 180-cm line source simulation, the 96-cm axial FOV was obtained with twofold better sensitivity compared to the 24-cm axial FOV. Also, we showed that the AEMC-PET scanner had a continuous axial FOV to 168 cm. We expect the AEMC-PET scanner will provide high versatility such as for measuring trace whole-body uptakes while keeping the continuous axial FOV and the scan time for static images will be reduced with a limited number of detectors., 2012IEEE NSS&MIC

One of the challenging applications of PET is implementing it for in-beam PET, which is an in situ monitoring method for charged particle therapy. For this purpose, we have previously proposed an open-type PET scanner, OpenPET. The original OpenPET had a physically opened field-of-view (FOV) between two detector rings through which irradiation beams pass. This dual-ring OpenPET (DROP) had a wide axial FOV including the gap. This geometry was not necessarily the most efficient for application to in-beam PET in which only a limited FOV around the irradiation field is required. Therefore, we have proposed a new single-ring OpenPET (SROP) geometry which can provide an accessible and observable open space with higher sensitivity and a reduced number of detectors than the DROP. The proposed geometry was a cylinder shape with its ends cut at a slant, in which the shape of each cut end became an ellipse. In this work, we developed and evaluated a small prototype of the SROP geometry for proof-of-concept. The SROP prototype was designed with 2 ellipse-shaped detector rings of 16 depth-of-interaction (DOI) detectors each. The DOI detectors consisted of 1024 GSOZ scintillator crystals which were arranged in 4 layers of 16×16 arrays, coupled to a 64-channel FP-PMT. Each ellipse-shaped detector ring had a major axis of 281.6 mm and a minor axis of 207.5 mm. For the slant mode, the rings were placed at a 45-deg slant from the axial direction and for the non-slant mode (used as a reference) they were at 90 deg from the axial direction with no gap. The system sensitivity measured from a 22 Na point source was 5.0% for the slant mode. The average spatial resolutions of major and minor axis directions were calculated as 3.8 mm FWHM and 4.9 mm FWHM, respectively for the slant mode. This difference resulted from the ellipsoidal ring geometry and the spatial resolution of the minor axis direction degraded by the parallax error. Comparison between the slant mode and the non-slant mode showed no remarkable difference in terms of the sensitivity distribution and spatial resolution performance. Therefore, we concluded that the SROP geometry has a good potential as an open geometry especially suitable for in-beam imaging.

The X'tal cube is a depth-of-interaction (DOI)-PET detector which is aimed at obtaining isotropic resolution by effective readout of scintillation photons from six sides of the crystal block. The X'tal cube is composed of a 3D crystal block with isotropic segments. Each face of the 3D crystal block is covered with a 4×4 array of multi-pixel photon counters (MPPCs). Previously, in order to fabricate the 3D crystal block efficiently and precisely, we applied a sub-surface laser engraving technique to a monolithic crystal block instead of gluing segmented small crystals. A dense arrangement of multiple micro-cracks carved by the laser beam works efficiently as a scattering wall for the scintillation photons. The X'tal cube with the laser-processed block showed excellent performance with respect to crystal identification and energy resolution. In this work, for characteristics comparison between the laser-processed block and the conventional segmented array block, we made the laser-processed block and two types of segmented array blocks, one with air gaps and the other with glued segmented small crystals. All crystal blocks had 3D grids of 2 mm pitch. The 4×4 MPPC arrays were optically coupled to each surface of the crystal block. When performance was evaluated using a uniform irradiation of 511 keV, we found that the X'tal cubes with the laser-processed block could easily achieve 2 mm 3 uniform crystal identification. Also, the average energy resolution of each 3D grid was 11.1±0.7%. On the other hand, the glued segmented array block had a pinched distribution and crystals could not be separated clearly. The segmented array block with air gaps had satisfactory crystal identification performance; however, the laser-processed block had higher crystal identification performance. Also, the energy resolution of the laser-processed block was better than for the segmented array blocks. In summary, we found the laser-processed X'tal cube had excellent crystal identification performance and high energy performance due to its effective dense arrangement of micro-cracks.

We are developing a three-dimensional (3D) position-sensitive detector with isotropic spatial resolution, the X'tal cube. Originally, our design consisted of a crystal block for which all six surfaces were covered with arrays of multi-pixel photon counters (MPPCs). In this paper, we examined the feasibility of reducing the number of surfaces on which a MPPC array must be connected with the aim of reducing the complexity of the system. We evaluated two kinds of laser-processed X'tal cubes of 3 mm and 2 mm pitch segments while varying the numbers of the 4 × 4 MPPC arrays down to two surfaces. The sub-surface laser engraving technique was used to fabricate 3D grids into a monolithic crystal block. The 3D flood histograms were obtained by the Anger-type calculation. Two figures of merit, peak-to-valley ratios and distance-to-width ratios, were used to evaluate crystal identification performance. Clear separation was obtained even in the 2-surface configuration for the 3 mm X'tal cube, and the average peak-to-valley ratios and the distance-to-width ratios were 6.7 and 2.6, respectively. Meanwhile, in the 2 mm X'tal cube, the 6-surface configuration could separate all crystals and even the 2-surface case could also, but the flood histograms were relatively shrunk in the 2-surface case, especially on planes parallel to the sensitive surfaces. However, the minimum peak-to-valley ratio did not fall below 3.9. We concluded that reducing the numbers of MPPC readout surfaces was feasible for both the 3 mm and the 2 mm X'tal cubes.

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 (12)C 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-to-noise 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.

The single-ring OpenPET (SROP), for which the detector arrangement has a cylinder shape cut by two parallel planes at a slant angle to form an open space, is our original proposal for in-beam PET. In this study, we developed a small prototype of an axial-shift type SROP (AS-SROP) with a novel transformable architecture for a proof-of-concept. In the AS-SROP, detectors originally forming a cylindrical PET are axially shifted little by little. We designed the small AS-SROP prototype for 4-layer depth-of-interaction detectors arranged in a ring diameter of 250 mm. The prototype had two modes: open and closed. The open mode formed the SROP with the open space of 139 mm and the closed mode formed a conventional cylindrical PET. The detectors were simultaneously moved by a rotation handle allowing them to be transformed between the two modes. We evaluated the basic performance of the developed prototype and carried out in-beam imaging tests in the HIMAC using (11)C radioactive beam irradiation. As a result, we found the open mode enabled in-beam PET imaging at a slight cost of imaging performance; the spatial resolution and sensitivity were 2.6 mm and 5.1% for the open mode and 2.1 mm and 7.3% for the closed mode. We concluded that the AS-SROP can minimize the decrease of resolution and sensitivity, for example, by transforming into the closed mode immediately after the irradiation while maintaining the open space only for the in-beam PET measurement.

We have developed a new three-dimensional (3D) position sensitive radiation detector, called the X'tal cube. The X'tal cube is composed of a scintillation crystal block and a number of multi-pixel photon counters (MPPCs) which are coupled on all six sides of the block. The block is segmented into cubes and no reflector is used between the segments. Scintillation light originating in a crystal segment accordingly propagates three-dimensionally along the alignment of the crystal segments and is efficiently detected by the MPPCs. The X'tal cube could be used as the detector element of a PET system, for instance. We constructed two prototypes of the X'tal cube and evaluated their performance using gamma-ray sources. The crystal block of each prototype is composed of a 3D array of Lu2xGd2(1-x)SiO5: Ce (LGSO, x = 0.9) crystal segments. Each crystal volume is 3.0 mm x 3.0 mm x 3.0 mm. MPPCs of a 3.0 mm x 3.0 mm active area are coupled to each surface of the crystal block. In this work, we show that all crystal segments are identified by a simple Anger-type calculation performed on the MPPC signals for both prototypes. The X'tal cube provides high spatial resolution in three dimensions regardless of the incident angle of the radiation.

The OpenPET and its real-time imaging capability have great potential for real-time tumor tracking in medical procedures such as biopsy and radiation therapy. For the real-time imaging system, we intend to use the one-pass list-mode dynamic row-action maximum likelihood algorithm (DRAMA) and implement it using general-purpose computing on graphics processing units (GPGPU) techniques. However, it is difficult to make consistent reconstructions in real-time because the amount of list-mode data acquired in PET scans may be large depending on the level of radioactivity, and the reconstruction speed depends on the amount of the list-mode data. In this study, we developed a system to control the data used in the reconstruction step while retaining quantitative performance. In the proposed system, the data transfer control system limits the event counts to be used in the reconstruction step according to the reconstruction speed, and the reconstructed images are properly intensified by using the ratio of the used counts to the total counts. We implemented the system on a small OpenPET prototype system and evaluated the performance in terms of the real-time tracking ability by displaying reconstructed images in which the intensity was compensated. The intensity of the displayed images correlated properly with the original count rate and a frame rate of 2 frames per second was achieved with average delay time of 2.1 s.

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 × 16 × 16 array of lutetium gadolinium oxyorthosilicate (LGSO) crystals 0.993 × 0.993 × 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 × 4 array of MPPCs), each having a sensitive area of 3.0 × 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.

Conventionally, block detectors, which consist of a two-dimensionally segmented scintillator array with inserted reflectors, are often used for PET. On the other hand, PET detectors with a monolithic block have been investigated because they are expected to offer higher resolution than do segmented crystal arrays. However, previous reports focused on detectors dedicated as small-animal PET, and the thickness was not good enough to stop 511-keV radiation. We developed a PET detector that uses a large and thick monolithic LYSO and 64-channel PS-PMT. When the LYSO was covered with reflectors, the spatial resolution, which was 3 mm FWHM at the center, rapidly became worse at the edge. We eliminated the loss of spatial resolution by replacing the reflectors with black paper, but the light output was decreased. Therefore, we concluded that spatial resolution and light output were in a trade-off relationship due to the edge effect of scintillation light.

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 (11)C 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 (18)F-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.

A conventional PET scanner has a 15–25 cm axial field-of-view (FOV) and images a whole body using about six bed positions. An OpenPET geometry can extend the axial FOV with a limited number of detectors. The entire whole-body PET scanner must be able to process a large amount of data effectively. In this work, we study feasibility of the fully 3D entire whole-body PET scanner using the GATE simulation. The OpenPET has 12 block detector rings with the ring diameter of 840 mm and each block detector ring consists of 48 depth-of-interaction (DOI) detectors. The OpenPET has the axial length of 895.95 mm with five parts of 58.95 mm open gaps. The OpenPET has higher single data loss than a conventional PET scanner at grouping circuits. NECR of the OpenPET decreases by single data loss. But single data loss is mitigated by separating the axially arranged detector into two parts. Also, multiple coincidences are found to be important for the entire whole-body PET scanner. The entire whole-body PET scanner with the OpenPET geometry promises to provide a large axial FOV with the open space and to have sufficient performance values. But single data loss at the grouping circuits and multiple coincidences are limited to the peak noise equivalent count rate (NECR) for the entire whole-body PET scanner.

Avalanche photodiodes (APDs) are being used as photodetectors in positron emission tomography (PET) because they have many advantages over photomultipliers (PMTs) typically used in PET detectors. We have developed a PET detector that consists of a multi-pixel APD and a 6×6×4 array of 1.46×1.46 mm 2 ×4.5 m LYSO crystals for a small animal PET scanner. The detector can identify four-layer depth of interaction (DOI) with a position-sensitive APD coupled to the backside of a crystal array by just an optimized reflector arrangement. Since scintillation lights are shared among many pixels by the method, weaker signals in APD pixels far from the interacting crystals are affected by noise. To evaluate the performance of the four-layer DOI detector with the APD and the influence of electrical noise on our method, we constructed a prototype DOI detector and tested its performance. We found, except for crystal elements on the edge of the crystal array, all crystal elements could be identified from the 2D position histogram. An energy resolution of 16.9% was obtained for the whole crystal array of the APD detector. The results of noise dependence of detector performances indicated that the DOI detector using the APD could achieve sufficient performance even when using application-specific integrated circuits.

We have proposed an ldquoOpenPETrdquo geometry, which consists of two axially separated detector rings, each of axial length W. A long and continuous field-of-view (FOV) including a 360-degree open gap G between two detector rings can be imaged through iterative image reconstruction. In addition to providing stress-less PET scanning and simultaneous PET/CT, the OpenPET is expected to lead to realization of in-beam PET. The OpenPET also extends the axial FOV with a limited number of detectors. However, the axial FOV is limited to 3 W because the maximum limit of G to obtain the axially continuous FOV is W. In this paper, therefore, we propose an alternative geometry to extend axial FOV even more without increasing the number of detectors. The proposed geometry consists of multiple detector rings separated by multiple gaps. By optimizing each width of the gaps based on a new concept of multiplex geometry of the OpenPET, the axial FOV can be theoretically increased to an unlimited extent without increasing the number of detectors. The multiplex OpenPET geometry was compared with the standard OpenPET and the conventional PET using numerical simulation data and experimental data. The results show that similar reconstructed images are obtained by three geometries. The proposed geometry is expected to help realize an affordable entire body PET scanner that enables whole body dynamic imaging.

One trend in positron emission tomography (PET) instrumentation over the last decade has been the development of scanners dedicated to small animals such as rats and mice. Thicker crystals, which are necessary to obtain higher sensitivity, result in degraded spatial resolution in the peripheral field-of-view (FOV) owing to the parallax error. On the other hand, we are developing the jPET-D4, which is a dedicated human brain PET scanner that has a capability for depth-of-interaction (DOI) measurement. Although its crystal width is about twice that of commercially available small animal PET scanners, we expect the jPET-D4 to have a potential for small animal imaging by making full use of the DOI information. In this article, we investigate the jPET-D4's potential for small animal imaging by comparing it with the microPET Focus220, a state-of-the-art PET scanner dedicated to small animals.The jPET-D4 uses four-layered GSO crystals measuring 2.9 mm x 2.9 mm x 7.5 mm, whereas the microPET Focus220 uses a single layer of LSO crystals measuring 1.5 mm x 1.5 mm x 10.0 mm. First, the absolute sensitivity, counting rate performance and spatial resolution of both scanners were measured. Next a small hot-rod phantom was used to compare their imaging performance. Finally, a rat model with breast tumors was imaged using the jPET-D4.Thanks to the thicker crystals and the longer axial FOV, the jPET-D4 had more than four times higher sensitivity than the microPET Focus220. The noise equivalent counting-rate performance of the jPETD4 reached 1,024 kcps for a rat-size phantom, whereas that of the microPET Focus220 reached only 165 kcps. At the center of the FOV, the resolution was 1.7 mm for the microPET Focus220, whereas it was 3.2 mm for the jPET-D4. On the other hand, the difference of resolution became smaller at the off-center position because the radial resolution degraded faster for the microPET Focus220. The results of phantom imaging showed that the jPET-D4 was comparable to the microPET Focus220 at the off-center position even as the microPET Focus220 outperformed the jPET-D4 except for the peripheral FOV.The jPET-D4 human brain PET scanner, which was designed to achieve not only high resolution but also high sensitivity by measuring DOI information, was proven to have a potential for small animal imaging.

One way to speed-up iterative image reconstruction is by parallel computing with a computer cluster. However, as the number of computing threads increases, parallel efficiency decreases due to network transfer delay. In this paper, we proposed a method to reduce data transfer between computing threads by introducing an intra-thread update. The update factor is collected from each slave thread and a global image is updated as usual in the first K sub-iteration. In the rest of the sub-iterations, the global image is only updated at an interval which is controlled by a parameter L. In between that interval, the intra-thread update is carried out whereby an image update is performed in each slave thread locally. We investigated combinations of K and L parameters based on parallel implementation of RAMLA for the jPET-D4 scanner. Our evaluation used four workstations with a total of 16 slave threads. Each slave thread calculated a different set of LORs which are divided according to ring difference numbers. We assessed image quality of the proposed method with a hotspot simulation phantom. The figure of merit was the full-width-half-maximum of hotspots and the background normalized standard deviation. At an optimum K and L setting, we did not find significant change in the output images. We also applied the proposed method to a Hoffman phantom experiment and found the difference due to intra-thread update was negligible. With the intra-thread update, computation time could be reduced by about 23%.

In a conventional PET scanner, coincidence events are limited by the energy window for detection of photoelectric events. In contrast, Compton scatter events occur both in a patient and in detector crystals. Scatter events within the patient cause scatter coincidences, but scatter events within crystals have useful information for an activity distribution. The PET scanner with an extended energy window has higher sensitivity but higher scatter fraction than the PET scanner with a default energy window. In this work, we develop a scatter reduction method using a depth-of-interaction (DOI) detector. The DOI detector can make an upper layer act as an absorber of patient scatter, thus shielding a lower layer. Therefore, Compton scatter events with an interaction at the lower layer mainly have more crystal scatter events than patient scatter events. Our method proposes a different energy window for each layer. The energy window of the upper layer is limited to the region of photopeak events, the same as done for the conventional PET scanner. The energy window of the lower layer is extended to the region of crystal scatter events and photopeak events. We tested our proposed technique for a whole-body PET scanner using GATE simulations. Simulation results show the DOI-PET scanner can provide more true events while keeping a low scatter fraction using our scatter reduction method. We conclude this method is promising for obtaining high image quality using DOI information and energy information.

Depth-of-interaction (DOI) detectors are considered to be advantageous for time-of-flight positron emission tomography (TOF-PET) because they can correct timing errors arising in the scintillation crystals due to a propagation speed difference between annihilation radiation and scintillation photons. We experimentally measured this timing error, using our four-layer DOI encoding method. The upper layers exhibited the larger timing delays due to the longer path lengths after conversion from annihilation radiation into scintillation photons that traveled by zigzag paths at a speed decreased by a factor of the refractive index (n). The maximum timing delay between the uppermost and the lowermost layers was evaluated as 164 ps when n=1.47. A TOF error correction was demonstrated to improve the timing resolution of the four-layer DOI detector by 10.3%, which would increase the effective sensitivity of the scanner by about 12% comparison with a non-DOI TOF-PET scanner. This is the first step towards combining these two important fields in PET instrumentation, namely DOI and TOF, for the purpose of achieving a higher sensitivity as well as a more uniform spatial resolution.

The long patient port of a PET scanner tends to put stress on patients, especially patients with claustrophobia. It also prevents doctors and technicians from taking care of patients during scanning. In this paper, we proposed an 'open PET' geometry, which consists of two axially separated detector rings. A long and continuous field-of-view (FOV) including a 360 degrees opened gap between two detector rings can be imaged enabling a fully 3D image reconstruction of all the possible lines-of-response. The open PET will become practical if iterative image reconstruction methods are applied even though image reconstruction of the open PET is analytically an incomplete problem. First we implemented a 'masked' 3D ordered subset expectation maximization (OS-EM) in which the system matrix was obtained from a long 'gapless' scanner by applying a mask to detectors corresponding to the open space. Next, in order to evaluate imaging performance of the proposed open PET geometry, we simulated a dual HR+ scanner (ring diameter of D = 827 mm, axial length of W = 154 mm x 2) separated by a variable gap. The gap W was the maximum limit to have axially continuous FOV of 3W though the maximum diameter of FOV at the central slice was limited to D/2. Artifacts, observed on both sides of the open space when the gap exceeded W, were effectively reduced by inserting detectors partially into unnecessary open spaces. We also tested the open PET geometry using experimental data obtained by the jPET-D4. The jPET-D4 is a prototype brain scanner, which has 5 rings of 24 detector blocks. We simulated the open jPET-D4 with a gap of 66 mm by eliminating 1 block-ring from experimental data. Although some artifacts were seen at both ends of the opened gap, very similar images were obtained with and without the gap. The proposed open PET geometry is expected to lead to realization of in-beam PET, which is a method for an in situ monitoring of charged particle therapy, by letting the beams pass through the gap. The proposed open PET geometry will also allow simultaneous PET/CT measurements of the same PET FOV as the CT FOV, in contrast to the conventional PET/CT where each FOV is separated by several tens of centimeters.

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 32×32×4 crystal (4096 crystals, each 1.46 mm×1.46 mm×4.5 mm) and a reflector, is mounted on the 256ch FP-PMT. In this study, we evaluated the spatial resolution of reconstructed images with the evaluation system of two four-layer DOI detectors which consist of 32×32×4 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.

Depth-of-interaction (DOI) measurement is essential for PET to reduce the parallax error, which results in improved spatial resolution while maintaining the sensitivity. One of the ultimate goals for improvement of DOI detectors is achieving their isotopic spatial resolution. For this purpose, we have developed an isotropic 3D position-sensitive PET detector named the X'tal cube detector. The X'tal cube detector is composed of a scintillation crystal block segmented into cubes and arrays of multi pixel photon detectors (MPPCs) which cover all surfaces of the crystal block. Segment responses are made as result of the 3D Anger-type calculation and are drawn in a 3D position histogram. In our previous study, although we achieved (0.8 mm)3 isotropic spatial resolution in most of the detector, we were not able to discriminate the outermost responses and the second outermost responses. This was because of a small mismatch of the size between the MPPC array (13 × 13 mm2) and crystal block (13.6 × 13.6 × 13.6 mm3), a dead space, 0.5 mm width gap, existed at the edge. In this study, we improved both the crystal identification performance at the edge and got 10% higher volume resolution than that of the previous detector without any degradation. To reduce the gap between the MPPC array and the scintillator crystal block, the scintillation crystal block used was a 13.1 × 13.1 × 13.1 mm3 monolithic LYSO having 3D segmentation in a 17 × 17 × 17 array of 0.77 × 0.77 × 0.77 mm3 grids fabricated by laser processing. Each face of the crystal block was coupled to a monolithic 4 × 4 array of MPPCs of 13 × 13 mm2. We irradiated 511 keV annihilation gamma rays from two 22Na sources to the detector and acquired the data with NIM and CAMAC. According to the profile of the response, 17 segments in a row were clearly identified. The peak-to-valley ratio was 4.4 ± 0.6 at the central.

We are developing a PET system integrated with a birdcage RF-coil for PET-MRI in order to realize both high sensitivity and high spatial resolution of the PET image by using the 4-layered depth-of-interaction (DOI) PET detector. We constructed a full-ring prototype system and evaluated performances, especially imaging performance, of the prototype system in simultaneous measurement. The prototype system consists of eight four-layer DOI-PET detectors and a prototype birdcage RF-coil developed for the proposed system. The PET detectors consist of six monolithic multi-pixel photon counter array (S11064-050P), a readout circuit, fourlayer DOI scintillator arrays and a shielding box made of 35 μm thick copper foil. The crystal array consists of 2.0 mm x 2.0 mm x 5.0 mm LYSO crystals arranged in 38 x 6 x 4 layer. The RF-coil has eight coil elements and the eight PET detectors are positioned at each element gap. The diameter of the RF-coil elements is 261 mm. We conducted performance tests of the prototype system with a 3.0 T MRI (MAGNETOM Verio). Only the PET detectors, the RF-coil and the cables were in an MRI room during measurements. A data acquisition system and power supplies for the MPPCs and preamplifiers were outside the MRI room and connected to all the detectors through a penetration panel. As a result, the spatial resolutions of a Na-22 point source in the PET image were lower than 1.6 mm in whole the FOV due to the DOI capability. In addition, the influence of the simultaneous measurements on the PET performance is negligible. On the other hand, the SNR of the phantom image in the magnitude images was degraded from 259.7 to 209.4 due to noise contamination from the power supplies.

We are developing a high-performance brain PET scanner, jPET-D4, which provides 4-layer depth-of-interaction (DOI) information. The scanner is designed to achieve not only high spatial resolution but also high scanner sensitivity with the DOI information obtained from multi-layered thin crystals. The scanner has 5 rings of 24 detector blocks each, and each block consists of 1024 GSO crystals of 2.9 mm/spl times/2.9 mm/spl times/7.5 mm, which are arranged in 4 layers of 16/spl times/16 arrays. At this stage, a pair of detector blocks and a coincidence circuit have been assembled into an experimental prototype gantry. In this paper, as a preliminary experiment, we investigated the performance of the jPET-D4's spatial resolution using the prototype system. First, spatial resolution was measured from a filtered backprojection reconstructed image. To avoid systematic error and reduce computational cost in image reconstruction, we applied the DOI compression (DOIC) method followed by maximum likelihood expectation maximization that we had previously proposed. Trade-off characteristics between background noise and resolution were investigated because improved spatial resolution is possible only when enhanced noise is avoided. Experimental results showed that the jPET-D4 achieves better than 3 mm spatial resolution over the field-of-view.

Previously, we proposed a new depth of interaction (DOI) encoding method and proved that it worked successfully with four-layered Gd/sub 2/SiO/sub 5/ crystals for a small animal positron emission tomography (PET) detector. We are now planning to develop a small animal PET scanner, jPET-RD (for rodents with DOI detectors), which has both high resolution and high sensitivity by the use of a DOI detector with a 32/spl times/32/spl times/4 crystal array. The scintillator for the detector will be Lu/sub 2(1-x)/Y/sub 2x/SiO/sub 5/ (LYSO). In this work, we evaluated performance of a DOI detector composed of four layers of a 12/spl times/12 LYSO (Lu: 98%, Y: 2%) crystal array by irradiating 511 keV gamma rays uniformly. The new encoding method was used for crystal identification. The size of each crystal was 1.46 mm/spl times/1.46 mm/spl times/4.5 mm. The crystal block was coupled to a 256-channel flat panel position sensitive photomultiplier tube, which has 16/spl times/16 multi anodes at intervals of 3.04 mm. As we expected, all crystals are expressed on a single two-dimensional position histogram without overlapping. Energy resolution of all events is 21.8% and time resolution of all events is 0.69 ns in FWHM. When layers are counted from the top, the energy resolutions of the first, second, third, and fourth layer events are 11.6%, 12.3%, 13.3%, and 19.1% and the time resolutions are 0.60ns, 0.59ns, 0.60ns, and 0.66ns, respectively.

We found that Bi/sub 4/Ge/sub 3/ O/sub 12/ (BGO) scintillator can be elements of a four-layer depth of interaction (DOI) detector and it was proved with a 12/spl times/12/spl times/4 array of BGO crystals in dimensions of 2.9 mm/spl times/2.9 mm/spl times/7.5 mm coupled to a 256-channel flat panel position sensitive photomultiplier tube. Appropriate reflector insertion in the array makes all crystal identification possible on one position histogram. Despite the large refractivity and small light output of BGO, the four-layer BGO detector showed no significant variation in the full energy peaks among all crystal elements. When no optical grease was used in the construction of the BGO DOI-block and irradiated with gamma-rays from /sup 137/Cs, a top layer crystal has 80% of light output relative to the bottom layer. The obtained two-dimensional position histogram by the irradiation was clear enough to allow identification of the crystals of interaction. Profiles of the histogram show peak-to-valley ratio of 1.9:1 for the top layer crystals and larger ratio for other layer crystals in the experiment.

In a PET scanner, a line of response is drawn between the two scintillators detecting the pair of annihilation radiation. However, the non-collinearity of the two photons limits the special resolution. The slight deviation from 180 degrees is known to be 0.47 degrees in the full width half maximum (FWHM) for pure water; however, the value may be changed by chemical and physical conditions for the positrons in vivo. The aim of this study is an experimental determination of the angular correlation of the annihilation radiation from a human body. \nUsually, the angle is measured by using plural position sensitive detectors such as gamma-cameras; however, this method cannot be used when the radiation source is specially distributed and the acquisition time is limited. The authors overcame this problem by measuring the energy distribution instead of the angular distribution; the two distributions can be converted each other by the conservation laws of momentum and energy. \nThe photopeak due to in vivo [18F]-FDG was measured by using a germanium detector with the energy resolution of 1.17 keV at 511 keV (0.23%), and the Gaussian distribution was determined to be 2.68 keV in FWHM. Therefore, the energy deviation due to the momentum of electron-positron pair before annihilation was calculated to be 2.41 keV in FWHM. This value is corresponding to 0.54 degree in the angular distribution and about 15% larger than that had obtained for pure water. \nBy the convolution of the annihilation radiation non-collinearity into the positron range, the physical limit of spatial resolution of FDG-PET with ideal radiation detectors can be estimated: 0.5 mm for a small animal PET with a 10-cm diameter, 1.2 mm for a human brain PET with a 40-cm diameter, and 2.1 mm for a human whole-body PET with a 80-cm diameter., World Congress on Medical Physics and Biomedical Engineering 2006

PET detectors based on light sharing, which can obtain spatial resolution smaller than the photo-detector size with a simple calculation to estimate the center of gravity, have often been studied, but the difficulty in identifying inter-crystal scattering (ICS) has not been resolved yet. In particular, ICS identification is strongly required to make full use of a minutely segmented crystal block such as used in the X'tal cube. The maximum likelihood (ML) method that calculates the likelihood function with response functions of photo-detector signals is suitable for the X'tal cube. However, the method does not take ICS into account. Therefore, in this paper, we developed a new method taking ICS into account for the X'tal cube detector. The proposed method was based on the ML method, and unknown positions were expanded to two to estimate scattering position and absorbed position for the ICS event. We evaluated performance of the proposed method using simulated data. We simulated the X'tal cube composed of the 6 × 6 × 6 array of (3 mm)3 LGSO crystals. Each surface of the crystal block was covered with 4 × 4 MPPCs, each with a 3 × 3 mm2 active area. As the results, the proposed method promises high ICS event identification, but crystal identification performance of the proposed method is almost equal to previous methods.

The OpenPET is our original idea that realizes the world's first open-type 3D PET scanner for PET-image guided particle therapy such as in situ dose verification and direct tumor tracking. Even with a full-ring geometry, the OpenPET has an open gap between its 2 detector rings through which the treatment beam passes. Following our initial 2008 proposal, we developed a small prototype in 2010 to show a proof-of-concept. Now, we report the development of a prototype whole-body OpenPET. The key technology which enabled the OpenPET realization is our original, 4-layered depth-of-interaction detector. In order to measure a radiation from the limited activity produced though fragmentation reactions, Zr-doped GSO (GSOZ), which contains less natural radioactivity, was chosen for the scintillators instead of Lu-based scintillators although timing performance was compromised. In order to compensate for the limited light yield, on the other hand, we used 64-channel flat-panel PMTs with a super-bialkali photocathode, which had a 30% higher quantum efficiency. In order to enable stable in-beam PET measurement even under high background radiations, voltage divider circuits were designed to provide 5 times higher linearity. Additionally, to avoid severe radiation damage, we did not use gain control ASICs in the front-end circuits, and position analyzer circuits were placed with a 15-m cable extension. The prototype consists of 2 detector rings, and each detector ring has 2 sub-rings of 40 detectors. Each detector consists of 16 × 16 × 4 array of GSOZ (2.8 × 2.8 × 7.5mm3). The portable gantry has a compact design; each detector ring has a 940 mm outer diameter and 171 mm thickness for the detector inner bore of 640 mm diameter and 113 mm thickness. The system was tested with a carbon beam irradiation at a clinical intensity. Phantom images were obtained by applying a GPGPU-based, list mode iterative reconstruction algorithm with geometrical detector response modeling.

Depth-of-interaction measurement is essential for PET to reduce the parallax error, resulting in improved spatial resolution with maintaining sensitivity. One of ultimate goals of improvement in DOI detectors is achieving isotopic spatial resolution. For this purpose, we have developed an isotropic 3D position sensitive PET detector named X'tal cube. The X'tal cube is composed of a scintillation crystal block segmented into cubes. Arrays of multi pixel photon detectors (MPPCs) cover all surfaces of the crystal block. In our previous study, we achieved 1.0 mm isotropic spatial resolution. In this study, we finally succeeded to develop the X'tal cube with (0.8 mm)3 isotropic segments. The scintillation crystal block was a 13.6 × 13.6 × 13.6 mm3 monolithic LYSO. 3D segmentation of 17 × 17 × 17 array of 0.8 mm × 0.8 mm × 0.8 mm grids was made by the laser processing. We used 96 MPPCs (S12642-0404PB-50(X)). For each face of the crystal block, monolithic 4 × 4 array of MPPCs with about 0.2 mm gaps was coupled. The position histograms were made by the 3D Anger-type calculation from uniform irradiation data, which showed good identification of each segments. We obtained 9.0% energy resolution for the central part of the crystal block. In conclusion, we succeeded to develop a prototype PET detector with the world's finest segments of (0.8 mm)3 grids.

A high-performance brain PET scanner, jPET-D4, which provides four-layer depth-of-interaction (DOI) information, is being developed to achieve not only high spatial resolution, but also high scanner sensitivity. One technical issue to be dealt with is the data dimensions which increase in proportion to the square of the number of DOI layers. It is, therefore, difficult to apply algebraic or statistical image reconstruction methods directly to DOI-PET, though they improve image quality through accurate system modelling. The process that requires the most computational time and storage space is the calculation of the huge number of system matrix elements. The DOI compression (DOIC) method, which we have previously proposed, reduces data dimensions by a factor of 1/5. In this paper, we propose a transaxial imaging system model optimized for jPET-D4 with the DOIC method. The proposed model assumes that detector response functions (DRFs) are uniform along line-of-responses (LORs). Then each element of the system matrix is calculated as the summed intersection lengths between a pixel and sub-LORs weighted by a value from the DRF look-up-table. 2D numerical simulation results showed that the proposed model cut the calculation time by a factor of several hundred while keeping image quality, compared with the accurate system model. A 3D image reconstruction with the on-the-fly calculation of the system matrix is within the practical limitations by incorporating the proposed model and the DOIC method with one-pass accelerated iterative methods.

A 256 channel flat panel position sensitive photomultiplier tube (256ch FP-PMT) was evaluated for application to a depth of interaction (DOI) PET detector composed of small crystal elements. The 256ch FP-PMT has 52 /spl times/ 52 mm opening area with 89% useful area. Its 256 anodes are placed at a 3.04 mm interval. To estimate the FP-PMT performance, light spread functions of anodes at central and peripheral regions of a prototype 256ch FP-PMT were measured with a Gd/sub 2/SiO/sub 5/ (GSO) crystal sized 1.42 /spl times/ 1.42 /spl times/ 4.5 mm. The crystal cross section area is about one quarter of one anode area of the FP-PMT. It was found that light spread functions of peripheral anodes were similar to the functions of central anodes and also the functions measured with a 2.9 /spl times/ 2.9 /spl times/ 7.5 mm GSO crystal which has almost the same dimensions of cross-section as the anode area. Transit time fluctuation among the 256ch FP-PMT anodes was estimated by measuring time resolutions of a LuYSiO/sub 5/ (LYSO) crystal at some central and peripheral positions. The LYSO crystal has the same dimensions as the smaller GSO crystal. BaF/sub 2/ was used for a reference detector. The obtained distributions show about the same time resolution between the central and peripheral positions and the full width at half maximum is (366 /spl plusmn/ 15) ps. The transit time fluctuation is /spl plusmn/ 106 ps. Two-dimensional position histograms of a 32 /spl times/ 32 GSO crystal array and two layers of 9 /spl times/ 9 GSO crystal arrays are obtained by uniform gamma-ray irradiation. The crystal dimensions are 1.42 /spl times/ 1.42 /spl times/ 4.5 mm so that the 32 /spl times/ 32 array covers all the useful area and indicates variation of crystal identification performance over the useful area. The 9 /spl times/ 9 /spl times/ 2 array is coupled to peripheral region. The resultant histograms confirm that a 256ch FP-PMT has enough capability for identifying crystals of this size even on the periphery and in DOI arrangement.

This paper describes a depth-of-interaction recognition method that can be applied to radiation detectors for positron emission tomography (PET). The proposed method uses optical filters to differentiate wavelength distribution of scintillation photons depending on crystal depth positions. Then the crystal of interaction can be identified with a multi-anode photo-detector whose sensitive surface is partly covered with optical filters. A validation experiment was performed with a position sensitive photomultiplier. As a result, it was proved that the crystal depth of interaction was clearly identified for photo-absorption events.

We are now planning to develop a positron emission tomograph dedicated to small animals, such as rats and mice, which meets the demand for higher sensitivity. For the positron emission tomography (PET), we newly designed a four-layer depth of interaction (DOI) detector, which is constructed of all the same crystal elements. In this DOI detector, appropriate insertion of the reflector controls the behavior of scintillation photons by inserting the reflectors between crystal elements so that the DOI information of four layers can be extracted from one two-dimensional (2-D) position histogram made by Anger-type calculation. As a preliminary experiment, we measured crystal identification performance of the DOI detector, which consists of four layers of a 16 /spl times/ 16 crystal array using Gd/sub 2/SiO/sub 5/ crystals with Ce concentration of 0.5 mol %. Each crystal is in dimensions of 1.42 mm /spl times/ 1.42 mm /spl times/ 4.5 mm. A crystal block is optically coupled to a 256-channel flat panel position sensitive photomultiplier tube whose opening area is 52.0 mm /spl times/ 52.0 mm. We obtained sufficient positioning performance for this four-layer DOI detector on the 2-D position histogram. We concluded that it would be a promising device to realize a small animal PET scanner with high sensitivity and high resolution.