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We developed a human-scale single-ring OpenPET (SROP) system, which had an open space allowing us access to the subject during measurement. The SROP system consisted of 160 4-layer depth-of-interaction detectors. The open space with the axial width of 430 mm was achieved with the ring axial width of 214 mm and the ring inner diameter of 660 mm. The detectors were axially shifted to each other so that the detector ring was aligned along a plane horizontally tilted by 45° against the axial direction. The system was developed as a mobile scanner to be used not only in clinical positron emission tomography (PET) rooms but also in charged-particle therapy treatment rooms as well as animal experiment rooms. Almost uniform spatial resolution better than 3 mm throughout the entire field of view (FOV) was realized with an iterative image reconstruction method. Peak absolute sensitivity was 3.1%, and there was a region with sensitivity better than 0.8% for a length of more than 700 mm. An in-beam imaging experiment conducted at the heavy ion medical accelerator in Chiba showed that the system was operable even at the highest beam intensity available for heavy-ion therapy. In addition, we conducted entire-body monkey dynamic imaging utilizing the long region inside the gantry by positioning a monkey along the direction having the longest FOV tilted by 45° against the axial direction. We concluded the developed system has a capability to realize versatile PET applications by utilizing its wide-open space and mobility in addition to high spatial resolution with sufficiently good sensitivity. -9mm]Please consider rephrasing the sentence “We concluded the developed system” for clarity.

To improve positron emission tomography (PET) image quality, depth-of-interaction (DOI) information and time-of-flight (TOF) information are key technologies. In this work, we developed the DOI-TOF detector based on our original single-ended readout scheme with the continuous layered gadolinium fine aluminum garnet (GFAG) array. The size of each GFAG crystal is $1.45\times 1.45\times 20$ mm3. The multipixel photon counter (MPPC) used as a photo sensor has a surface area of $3.0\times 3.0$ mm2. One pair of crystals coupled with a partial optical window is arranged across two MPPCs. Boundaries without the partial optical window are covered with optical reflectors. Each pair is coupled with paired MPPCs of different patterns. Crystal identification is obtained from paired MPPCs of different patterns and output rates. By limiting the spread of scintillation light, the proposed DOI detector based on the local centroid calculation was expected to improve crystal response, including the edge. For performance comparison, we prepared three discrete layers of the GFAG array with the same total thickness. For the continuous layer, the DOI resolution, the energy resolution and the timing resolution of the pair of detectors are 4.7 mm, 14%, and 402 ps, respectively. Their respective values are 29%, 20%, and 33% better than those with the three discrete layers.

In this article, we developed the novel single-end readout depth-of-interaction (DOI) detector with quadrisected crystals comparable in size to a single photo sensor. The proposed DOI detector consists of 784 gadolinium aluminum gallium garnet (GAGG) crystals, which are arranged in four layers of $14\times14$ arrays coupled with an $8\times8$ multipixel photon counter (MPPC) array. The size of each GAGG crystal is $1.45\times 1.45\times4.5$ mm3. Also, each MPPC has a surface area of $3.0\times3.0$ mm2. One pair of four-layered crystals coupled with a partial optical window is arranged across two MPPCs. Boundaries without the partial optical window are covered with optical reflectors. Light sharing is done between a pair of four-layered crystals optically coupled at one end. Each pair is coupled with paired silicon photomultipliers of different patterns. Crystal identification from each pair is obtained from the output rate of two MPPCs. From 511-keV uniform irradiation, responses of all crystal elements in the four layers can be separated clearly on a 2-D position histogram after applying the Anger calculation. Also, DOI information can be mapped to each cluster on curved lines. The energy resolution for all crystals is 9.1% after event reduction of intercrystal scatter.

The quality and uniformity of positron emission tomography (PET) images can be improved significantly using a PET scanner, which has both time-of-flight (TOF) and depth-of-interaction (DOI) capabilities. We have already developed a DOI dual-end detector using segmented crystal bars by applying a subsurface laser engraving (SSLE) technique, and the DOI was determined by the ratio of the detected light between two readouts using the Anger calculation. In this study, we investigated the influence of the number of DOI segments on the performance of DOI identification and the coincidence timing performance of the detector. The detector consisted of a single lutetium fine silicate (LFS) crystal bar (3 × 3 × 20 mm3) with various numbers of DOI segments that were made by applying the SSLE technique and Hamamatsu silicon photomultiplier (SiPM) modules. The maximum number of DOI segments was six, and the SiPM module included one hundred forty-four 4 mm × 4 mm SiPM readouts. The coincidence resolving time (CRT) of each DOI segment was obtained from the side irradiation of the dual-end detector. All DOI segments of the detector with 2, 3, 4, 5 or 6 DOI segments were clearly identified, and average energy resolutions of 9.8 ± 0.5% and 12.5 ± 1.4% were obtained at the 511 keV photo peak for the detectors with 2 DOI and 6 DOI segments, respectively. The minimum and maximum estimated CRT of 180 ± 6 ps and 236 ± 6 ps were obtained for the detectors with 2 DOI and 6 DOI segments, respectively. Insignificant differences were observed between the CRT values of different segments of one detector. Greater CRT values were obtained for detectors with larger DOIs. The results of this study prove that there is a high potential for segmented crystal bars using the SSLE technique as a good candidate for PET scanners with TOF and DOI capabilities, which can significantly improve the quality of PET images.

Previously, we had developed a four-layered depth-of-interaction (DOI) PET detector based on the light sharing method. Reflectors , which were inserted in every two lines of crystal segments and shifted differently for each layer, projected 3D crystal positions onto a 2D position histogram without any overlapping after applying the Anger-type calculation. The best crystal separation we ever obtained was for the four-layered 32 × 32 array of LYSO crystals sized at 1.45 × 1.45 × 5 mm3. However, assembling small crystals tended to cost a lot, and fine tuning of the front-end circuit was required to get fine crystal identification. In this paper, therefore, we proposed a more practical four-layered DOI detector. Its key concept was that the crystals in the top layer, which have the highest detection efficiency, are the biggest contributors to the PET spatial resolution. We applied two new ideas: (1) using quarter size crystals only for the first (top) layer and (2) inserting a thin light guide between the first and the second layers of the crystal array. In the developed prototype detector, we used 24 × 24 LYSO crystals of quarter size (1.4 × 1.4 × 5.0 mm3) in the first layer and the other layers were 12 × 12 arrays of crystals of 2.8 × 2.8 × 5.0 mm3. For better crystal identification of small crystals in the first layer, we optimized the optical condition between crystals by using an optical adhesive and air. Also, the thin light guide of 0.5 mm thickness was inserted between the first and the second layers for improvement of crystal identification of the first layer. With the appropriate insertion of the light guide, all crystals of the first layer were identified as well as the crystals in the other layers. Our developed four-layered DOI detector showed good potential for high spatial resolution without a large increase in the number of crystals.

Brain PET, which has led research in molecular imaging and diagnosis of brain cancer, epilepsy and neurodegenerative disorders, is being spotlighted again to promote earlier diagnosis of dementia with the advent of amyloid and tau tracers. To meet this demand, in this paper, we developed a brain-dedicated PET imaging device with a hemispherical detector arrangement, which provides comparable sensitivity with fewer detectors than conventional cylindrical geometries. The introduction of the time-of-flight (TOF) measurement capability was a key point for the development to get a gain in the image signal-to-noise ratio. Currently, whole-body PET scanners with around 200-400 ps coincidence resolving time (CRT) are commercially available. In order to obtain the same TOF gain which can be obtained with 400 ps CRT for a 30 cm diameter object, 267 ps CRT will be required for a 20 cm diameter object such as the human head. In this work, therefore, we aimed at developing a TOF brain-dedicated PET prototype with the hemisphere detector arrangement and the CRT faster than 267 ps. The detector was composed of a 12 × 12 lutetium fine silicate (LFS) array coupled with a 12 × 12 multi-pixel photon counter (MPPC) array. Each LFS crystal with a size of 4.14 × 4.14 × 10 mm3 was individually coupled to a separate MPPC. Singles list-mode data from each detector were stored, and coincidences were identified using a coincidence-detection software algorithm. The CRT of 245 ps was finally achieved as the system average after a fine timing correction. For image reconstruction, we implemented the list-mode TOF-OSEM. For a small rod phantom, rods of 3 mm diameter were clearly separated. Also, images of the 3D Hoffman brain phantom, which demonstrated clear contrast between gray and white matter, supported the effect of TOF information.

Recently, silicon photomultipliers (SiPMs) have become established as photo-sensors for PET scanners. Commercially available SiPM arrays are typically about 3–4 cm in size. In order to develop larger detectors with an Anger calculation, multiple SiPM arrays need to be combined. This configuration requires the use of a light guide and has inactive areas between SiPM arrays. If not addressed, crystal identification degrades not only for peripheral crystals but also for crystals to bridge the gap between the SiPM arrays. On the other hand, we have developed the crosshair light-sharing (CLS) PET detector, which is based on a single-ended readout scheme with roughly quadrisected crystals comparable in size to a SiPM, with depth-of-interaction (DOI) capability. The CLS PET detector based on the local centroid calculation was expected to mitigate the edge effect similar to what the one-to-one coupling detector is able to do. In this work, we developed the axially long CLS PET detector with two multipixel photon counter (MPPC) arrays. The proposed long CLS PET detector consisted of 30 × 14 gadolinium fine aluminum garnet (GFAG) arrays coupled to two 8 × 8 MPPC arrays without the light guide. From 511-keV uniform irradiation, responses of all crystal elements could be separated clearly on a 2D position histogram after applying the Anger calculation. Pulse heights of additional crystals to bridge the gap were 18% lower than those of crystals at the center of the MPPC array, but similar to that of the peripheral crystals based on non-uniformity of the MPPC array. On the other hand, energy resolutions of additional crystals to bridge the gap were almost the same to those of other crystals. The long CLS PET detector promises improvement of the packing fraction while keeping performance.

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.

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.

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 previously proposed the world's first open-type geometries which we named OpenPET. For in-beam PET, careful designing and testing of detectors are important because secondary particles generated in a target degrade detector performance in the OpenPET geometries. As well, we have developed small prototypes and identified their promising potential for in-beam PET and succeeded in visualizing a 3D distribution of beam stopping positrons inside a phantom. In this work, we developed a whole-body dual ring OpenPET (WBDROP) with a singles-based data acquisition (DAQ) system. This scanner design has two separated rings. Each ring has two detector rings of 40 detector blocks, and each detector block consists of a 4-layered depth-of-interaction (DOI) detector. The ring diameter is 660 mm, and the distance between separated rings, which can be controlled, is 94 mm. Each DOI detector consists of 1024 GSOZ crystals which are arranged in four layers of 16 × 16 arrays, coupled to a 64-ch flat panel position sensitive photomultiplier tube. In order to increase scalability and maintain the possibility for future development of complicated coincidence algorithms to reject background signals, our designed DAQ system has no coincidence circuit and uses only singles list-mode DAQ circuits. The DAQ system makes a judgment on coincidence based on software, where the singles list-mode data include 6-bit energy markers and 500-ps tags for timing information. The system sensitivity measured from a 22Na point source was 4.4%. The average spatial resolutions were calculated as about 3.5 mm FWHM. For the singles count rate, the dead time at 250 MBq was about 8%. The singles-based DAQ system had significant performance in singles list-mode data acquisition. For in-beam PET, imaging performance of the WBDROP was confirmed through phantom experiments. Phantom study results with 11C and 10C beam irradiations of about 3 Gy showed that the beam stopping position in the target could be measured with a precision of better than 2 mm. The WBDROP promises high performance for not only conventional PET studies but also in-beam PET measurements.

High sensitivity and high spatial resolution dedicated brain PET scanners are in high demand for early diagnosis of neurodegenerative diseases and studies of brain functions. To meet the demand, we have proposed the helmet-chin PET geometry which has a helmet detector and a chin detector. Our first prototype scanner used 54 4-layer depth-of-interaction (DOI) detectors. The helmet detector of the scanner had three detector rings with different radii arranged on a surface of a hemisphere (with a radius of 126.5 mm) and a top cover detector. Therefore, in this study, for our next development, we propose a spherical arrangement, in which the central axis of each detector points toward the center of the hemisphere, and we optimize the size of the detector crystal block to be arranged on the helmet detector. We simulate the spherical arrangement with the optimized crystal block size and compare its imaging performance with the multi-ring arrangement, which has a similar detector arrangement to that of our first prototype. We conduct Monte Carlo simulation to model the scanners having the 4-layer DOI detectors which consist of LYSO crystals. A dead space of 2 mm is assumed on each side of the crystal blocks such as for wrapping. The size of the crystal block is varied from 4×4 mm2 to 54×54 mm2 while fixing the thickness of the crystal block to 20 mm. We find that the crystal block sized at 42×42 mm2 has the highest sensitivity for a hemispherical phantom. The comparison of the two arrangements with the optimized crystal blocks show that, for the same number of crystal blocks, the spherical arrangement has 17% higher sensitivity for the hemispherical phantom than the multi-ring arrangement. We conclude that the helmet-chin PET with the spherical arrangement constructed from the crystal block sized at 42×42×20 mm3 has better imaging performance especially at the upper part of the brain compared to the multi-ring arrangement while keeping similar spatial resolution throughout the FOV.

There is a growing interest in developing brain PET scanners with high sensitivity and high spatial resolution for early diagnosis of neurodegenerative diseases and studies of brain functions. Sensitivity of the PET scanner can be improved by increasing the solid angle. However, conventional PET scanners are designed based on a cylindrical geometry, which may not be the most efficient design for brain imaging in terms of the balance between sensitivity and cost. We proposed a dedicated brain PET scanner based on a hemispheric shape detector and a chin detector (referred to as the helmet-chin PET), which is designed to maximize the solid angle by increasing the number of lines-of-response in the hemisphere. The parallax error, which PET scanners with a large solid angle tend to have, can be suppressed by the use of depth-of-interaction detectors. In this study, we carry out a realistic evaluation of the helmet-chin PET using Monte Carlo simulation based on the 4-layer GSO detector which consists of a 16 × 16 × 4 array of crystals with dimensions of 2.8 × 2.8 × 7.5 mm3. The purpose of this simulation is to show the gain in imaging performance of the helmet-chin PET compared with the cylindrical PET using the same number of detectors in each configuration. The sensitivity of the helmet-chin PET evaluated with a cylindrical phantom has a significant increase, especially at the top of the (field-of-view) FOV. The peak-NECR of the helmet-chin PET is 1.4 times higher compared to the cylindrical PET. The helmet-chin PET provides relatively low noise images throughout the FOV compared to the cylindrical PET which exhibits enhanced noise at the peripheral regions. The results show the helmet-chin PET can significantly improve the sensitivity and reduce the noise in the reconstructed images.

Myocardial perfusion imaging (MPI) with PET plays a vital role in the management of coronary artery disease. High sensitivity systems can contribute to maximizing the potential value of PET MPI; therefore, we have proposed two novel detector arrangements, an elliptical geometry and a D-shape geometry, that are more sensitive and more compact than a conventional large-bore cylindrical geometry. Here we investigate two items: the benefits of the proposed geometries for cardiac imaging; and the effects of scatter components on cardiac PET image quality. Using the Geant4 toolkit, we modeled four time-of-flight (TOF) PET systems: an 80 cm diameter cylinder, a 40 cm diameter cylinder, a compact ellipse, and a compact D-shape. Spatial resolution and sensitivity were measured using point sources. Noise equivalent count rate and image quality were examined using an anthropomorphic digital chest phantom. The proposed geometries showed higher sensitivity and better count rate characteristics with a fewer number of detectors than the conventional large-bore cylindrical geometry. In addition, we found that the increased intensity of the scatter components was a big factor affecting the contrast in defect regions for such a compact geometry. It is important to address the issue of the increased intensity of the scatter components to develop a high-performance compact cardiac TOF PET system.

The crosshair light-sharing (CLS) PET detector is our original depth-of-interaction (DOI) detector, which is based on a single-ended readout scheme with quadrisected crystals comparable in size to a photo-sensor. In this work, we developed 32 CLS PET detectors, each of which consisted of a multi-pixel photon counter (MPPC) array and gadolinium fine aluminum garnet (GFAG) crystals, and we developed a benchtop prototype of a small animal size PET. Each GFAG crystal was 1.45 × 1.45 × 15 mm3. The MPPC had a surface area of 3.0 × 3.0 mm2. The benchtop prototype had two detector rings of 16 detector blocks. The ring diameter and axial field-of-view were 14.2 cm and 4.9 cm, respectively. The data acquisition system used was the PETsys silicon photomultiplier readout system. The continuous DOI information was binned into three DOI layers by applying a look-up-table to a 2D position histogram. Also, energy and timing information was corrected using DOI information. After the calibration procedure, the energy resolution and the coincidence time resolution were 14.6% and 531 ps, respectively. Imaging test results of a small rod phantom obtained by an iterative reconstruction method showed clear separation of 1.6 mm rods with the help of DOI information.

Silicon photomultipliers (SiPMs) have become a standard photodetector to be coupled with scintillators in PET, but the SiPM saturation is limiting the performance achievable with the detectors employed, in particular the high energy resolution which is necessary for whole gamma imaging (WGI). The concept of WGI combines PET and a Compton camera by inserting a scatterer detector ring into a PET ring. Not only typical SPECT radionuclides such as 99m Tc (140 keV), but also unusual positron emitters such as 89Zr (909 keV) and 44Sc (1157 keV) can be imaging targets. For better spatial resolution in Compton imaging, the scatterer detector requires better energy resolution for a wide range of deposited energies. The use of bright scintillators such as GAGG is essential, but the SiPM saturation may prevent full use being made for such bright scintillators. We expected that inserting a thick light guide between GAGG and SiPM could spread scintillation photons to surrounding SiPMs and eliminate the saturation effect. On the other hand, the thicker the light guide becomes, the greater the number of scintillation photons that may be absorbed. Therefore, in this paper, we investigated the relationship between the light guide thickness and the energy resolution. A 22 × 22 array of GAGG crystals (0.9 × 0 . 9 × 6 mm3 each) was optically coupled to the 8 x 8 multi-pixel photon-counter (MPPC) array (3 × 3 mm2 pixel, 50 × 50 μ m 2 sub-pixel) via a light guide, for which thickness was changed from 0 (i.e., without the light guide) to 8 mm. Using point sources with different energies (133Ba, 22Na and 137Cs), we compared crystal identification performance, linearity of the output signal and energy resolution. Increasing the light guide thickness gradually degraded crystal identification performance but improved linearity of the output signals. Energy resolution at 81 keV constantly deteriorated with increasing light guide thickness. Energy resolutions at 356, 511 and 662 keV were improved with increasing light guide thickness to a certain value after which they deteriorated; the thicknesses at which deterioration started were 2.0 mm, 3.0 mm and 4.0 mm, respectively, for the energy resolutions at 356, 511 and 662 keV. We found that the optimum light guide thickness for the target energy range was 2.0 mm, and for this thickness, energy resolution values were 22.0% at 81 keV, 7.6% at 356 keV, 8.3% at 511 keV and 8.2% at 662 keV.

The purpose of this work is to develop a validated Geant4 simulation model of a whole-body prototype PET scanner constructed from the four-layer depth-of-interaction detectors developed at the National Institute of Radiological Sciences, National Institutes for Quantum and Radiological Science and Technology, Japan. The simulation model emulates the behaviour of the unique depth of interaction sensing capability of the scanner without needing to directly simulate optical photon transport in the scintillator and photodetector modules. The model was validated by evaluating and comparing performance metrics from the NEMA NU 2-2012 protocol on both the simulated and physical scanner, including spatial resolution, sensitivity, scatter fraction, noise equivalent count rates and image quality. The results show that the average sensitivities of the scanner in the field-of-view were 5.9 cps kBq−1 and 6.0 cps kBq−1 for experiment and simulation, respectively. The average spatial resolutions measured for point sources placed at several radial offsets were 5.2± 0.7 mm and 5.0± 0.8 mm FWHM for experiment and simulation, respectively. The peak NECR was 22.9 kcps at 7.4 kBq ml−1 for the experiment, while the NECR obtained via simulation was 23.3 kcps at the same activity. The scatter fractions were 44% and 41.3% for the experiment and simulation, respectively. Contrast recovery estimates performed in different regions of a simulated image quality phantom matched the experimental results with an average error of -8.7% and +3.4% for hot and cold lesions, respectively. The results demonstrate that the developed Geant4 model reliably reproduces the key NEMA NU 2-2012 performance metrics evaluated on the prototype PET scanner. A simplified version of the model is included as an advanced example in Geant4 version 10.5.

We proposed a concept of whole gamma imaging (WGI) that utilizes all detectable gamma rays for imaging. An additional detector ring, which is used as the scatterer, is inserted in the field-of-view of a PET ring so that single gamma rays can be detected by the Compton imaging method. In particular, for the non-pure positron emitters which emit an additional gamma ray almost at the same time, triple gamma imaging will be enabled; localization on each line-of-response (LOR) is possible by using the Compton cone of the additional gamma ray. We developed a prototype to show a proof of the WGI concept. The diameters of scatterer ring and PET ring were set as 20 cm and 66 cm, respectively. For Compton imaging of the 662-keV gamma ray from a 137Cs point source, spatial resolution obtained by the list-mode OSEM algorithm was 4.4 mm FWHM at the 8 cm off-center position and 13.1 mm FWHM at the center position. For PET imaging of a 22Na point source, spatial resolution was about 2 mm FWHM at all positions. For the triple gamma imaging, 5.7 mm FWHM (center) and 4.8 mm FWHM (8 cm off-center) were obtained for the 22Na point source just by plotting the intersecting points between each LOR and each Compton cone of the 1275-keV gamma ray. No image reconstruction was applied. Scandium-44 was produced as a practical candidate of the non-pure positron emitters, and 6.6 mm FWHM (center) and 5.8 mm FWHM (8 cm off-center) were obtained in the same manner. This direct imaging approach which neither requires time-consuming event integration nor iterative image reconstruction may allow in vivo real-time tracking of a tiny amount of activity. Our initial results showed the feasibility of the WGI concept, which is a novel combination of PET and Compton imaging.

Time-of-flight (TOF) is now a standard technology for positron emission tomography (PET), but its effective use for small diameter PET systems has not been studied well. In this paper, we simulated a brain-dedicated TOF-PET system with a hemispherical detector arrangement. We modeled a Hamamatsu TOF-PET module (C13500-4075LC-12) with 280 ps coincidence resolving time (CRT), in which a 12 x 12 array of multi pixel photon counters (MPPCs) is connected to a lutetium fine silicate (LFS) crystal array of 4.1 x 4.1 mm2 cross section each, based on one-to-one coupling. On the other hand, spatial resolution degradation due to the parallax error should be carefully addressed for the small diameter PET systems. The ideal PET detector would have both depth-of-interaction (DOI) and TOF capabilities, but typical DOI detectors that are based on light sharing tend to degrade TOF performance. Therefore, in this work, we investigated non-DOI detectors with an appropriate crystal length, which was a compromise between suppressed parallax error and decreased sensitivity. Using GEANT4, we compared two TOF detectors, a 20 mm long non-DOI and a 10 mm long non-DOI, with a non-TOF, 4-layer DOI detector with a total length of 20 mm (i.e., 5 x 4 mm). We simulated a contrast phantom and evaluated the relationship between the contrast recovery coefficient (CRC) and the noise level (the coefficient of variation, COV) for reconstructed images. The 10 mm long non-DOI, which reduces the parallax error at the cost of sensitivity loss, showed better imaging quality than the 20 mm long non-DOI. For example, the CRC value of a 10 mm hot sphere at COV = 20% was 72% for the 10 mm long non-DOI, which was 1.2 times higher than that of the 20 mm long non-DOI. The converged CRC values for the 10 mm long non-DOI were almost equivalent to those of the non-TOF 4-layer DOI, and the 10 mm long non-DOI converged faster than the non-TOF 4-layer DOI did. Based on the simulation results, we evaluated a one-pair prototype system of the TOF-PET detectors with 10 mm crystal length, which yielded the CRT of 250 ± 8 ps. In summary, we demonstrated support for feasibility of the brain-dedicated TOF-PET system with the hemispherical detector arrangement.

Time-of-fligfrt (TOF) PET improves image quality and quantitative accuracy compared with a conventional PET system. While it is true that TOF works better for larger objects, recent improvement in timing resolution has encouraged application of TOF to brain-dedicated PET systems. Thus, as the second prototype of the helmet-type PET, we have developed a new TOF brain-dedicated PET prototype using detector modules of 12×12 lutetium fine silicate (LFS) crystals (4.1×4.1×10 mm3) connected to a 12×12 (4 mm pitch, 144-ch) MPPC array. In this paper, we investigated another detector module, which has the same outer size but smaller crystals: 16×16 LFS of 3.1×3.1×10 mm3 size coupled to 16×16 (3 mm pitch, 256-ch) MPPC array. The 3-mm 256-ch module showed an energy resolution of 12.0%. For the coincidence response function, the 3-mm module showed a better full width at half maximum (FWHM) of 1.9 mm compared with the 4-mm 144-ch module (2.4 mm). The FWHM was improved by 21%. The coincidence resolving time obtained by the 3-mm 256-ch module was 241 ps. The energy resolution and coincidence resolving time were almost the same between the two modules. The 3-mm 16x16 MPPC-based TOF-PET module can be expected to improve spatial resolution without compromising the energy resolution and coincidence resolving time compared with the 4-mm 12×12 module.

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.

We have proposed a new concept of whole gamma imaging (WGI), which is a combined Compton-PET system by inserting a scattering detector ring inside a PET detector ring. In addition to positron emitters, WGI can visualize single gamma emitters based on the Compton imaging method. Besides, for triple-gamma emitters such as 44Sc that emits a pair of 511 keV photons and a 1157 keV single gamma-ray almost at the same time, a direct imaging method would be possible; the source position will be calculated as an intersection point between a surface of the Compton cone and a line-of-response. In 2017, we succeeded in developing the world’s first WGI prototype. However, limited sensitivity for Compton imaging degraded the system sensitivity for triple-gamma emitters. Therefore, in this work, we simulated the WGI system to optimize the scatterer detector for higher sensitivity. Using GEANT4, we modeled the WGI geometry, in which a scatter ring (GAGG crystals, 20 cm inner diameter) was inserted into a PET ring (GSOZ crystals, 66 cm inner diameter). In order to improve the sensitivity of triplegamma imaging and reduce parallax error, we used 4-layer depth-of-interaction (DOI) detectors as the scatterer. Total thickness of the scatterer was increased from 6 mm to 18 mm. However, thick scatterer also increased detection efficiency of 511 keV photons in the scatterer. Therefore, we improved the triple-gamma detection method in order to use the scatterer detection events of 511leV for triple-gamma imaging. Events in the scatterer were identified as positron annihilation and Compton scattering of the 1157 keV gamma-ray based on their energy information. The simulation results showed that (1) the sensitivity of the WGI using the 18-mm thick GAGG scatterer with the new triple-gamma detection method was 2.7 times higher than that of our first prototype (using 6-mm thick GAGG), and (2) the use of 4-layer DOI kept the position resolution in the triple-gamma imaging.

In carbon ion therapy, visualization of the range of incident particles in a patient body is important for treatment verification. In-beam positron emission tomography (PET) imaging in ion therapy is one of the method for verification of treatment due to the high quality of PET images. We have already shown the feasibility of radioactive beams of 15O and 11C using in-beam PET imaging using our OpenPET system. Recently, we have developed a whole gamma imaging (WGI) system which can work as PET, single gamma-ray and triple gamma-ray imaging simultaneously. The WGI system has high potential to detect location of 10C, which emits positron with a simultaneous gamma-ray of 718 keV, and activity of other produced positron emitting nuclei within patient body during ion therapy. In this work, we focus on investigation of performance of WGI system for 10C irradiation. First, the performance of scatterer detector of the WGI system regarding the radiation hardness was studied for 10C irradiation, then the performance of WGI was studied by simulation using the Geant4 code. The scatterer detector consisted of a segmented GAGG crystal and a multi-pixel photon counter (MPPC) and its performance was studied as a PMMA phantom was irradiated with 5 and 400 spills of 10C with energy of 350 MeV/u. No damage to the MPPC (scatterer detector) was observed even though after irradiation of 400 spills (~1.9×105 particle per spill). Sensitivity values of WGI system by simulation of a 10C nuclide at the center of field of view (FOV) for PET mode imaging, single gamma-ray of 718 keV imaging and triple gamma-ray imaging were 7.85%, 0.28% and 0.012% respectively. The performance of WGI system is going to be evaluated for 10C irradiation in the near future.

PET detectors with both time-of-flight (TOF) and depth-of-interaction (DOI) capabilities are ideal, but DOI-TOF PET detectors have not been studied well. In this paper, we develop a 4-layer DOI TOF detector. The key item is a recently commercialized module of multi pixel photon counters (MPPCs) (C13500 series, Hamamatsu Photonics K. K., Japan). The MPPC array is the through silicon via (TSV) type with an effective area of 6 x 6 mm3 (0.075 mm pitch sub pixels), which is arranged as an 8 x 8 array. At first, a pair of lutetium fine silicate (LFS) crystals each sized in 4 x 4 x 10 mm3 was placed on the center of each one MPPC channel. The CRT value was 252 ps. Afterward one of those detectors was used as the reference detector. Then cerium doped lutetium yttrium orthosilicate (LYSO) crystals sized in 2.8 x 2.8 x 5 mm3 were arranged in 8 x 8 x 4-layer based on our special reflector arrangement. The 4-layer LYSO array were coupled to the MPPC modules via a light guide. In the position map of the 4-layer LYSO crystal block, most of crystals could be identified. The averaged CRT value for crystals of the central part in the position map for the reference detector was 486 ± 106 ps. In timing spectrum of the third layer and the fourth layer, other components were observed. This was presumed to be due to contamination caused by incomplete crystal identification. Individual specificity for each MPPC channel was required to be calibrated for complete crystal identification. In conclusion, we developed a 4-layer DOI TOF detector and obtained around 500 ps averaged CRT value for the reference detector.

We have developed four-layered depth-of-interaction (DOI) detectors based on light sharing. Reflectors, which are inserted in every two lines of crystal segments and shifted differently depending on each layer, project 3-D crystal positions to a 2-D position histogram without any overlapping after applying the Anger-type calculation. The DOI measurement itself has the potential to improve time resolution because the depth-dependent timing delay can be corrected. However, light sharing tends to increase variance of light paths inside the crystal block, thus resulting in worsened time resolution. Although we have reported advantages of our DOI detectors in terms of position and energy resolutions, we had not evaluated their potential for time resolution. In this paper, therefore, we measured timing performance with the help of a digital photon counter (DPC), which offers precise control of event triggering. There are several studies that have reported one-to-one coupling of the scintillator to the DPC pixel, but DPCs have not been studied well for light-sharing detectors. Therefore, in this work, we optimized measurement parameters of the DPCs for our four-layered DOI detector. The DOI detector consists of 256 LGSO crystals which are arranged in four layers of ${8} \times {8}$ arrays, coupled to the DPC array. Each crystal element is $2.9\times 2.9 \times 5~\hbox{mm}^{3}$ . Each die of the DPC array provides an individual timestamp. Crystal identification performance largely depended on the dark count rate of the DPC array, which can be reduced by means of cell inhibition. We measured several conditions of the inhibition rate of microcells and temperature. For increased inhibition rate, we observed degraded time resolution, although positioning performance and energy resolution were improved. Regarding the temperature dependency within 10 to $ - 7 \ ^\circ \hbox {C}$ , we found that time resolution was insensitive. At 10 $\ ^\circ \hbox {C}$ and 20% inhibition rate, average time resolution over all crystals was $267 \pm 32~\hbox{ps}$ (full width half maximum). Better positioning performance and energy resolution were obtained for colder temperatures.

The purpose of this study is to perform a first characterization and proof of principle investigation of a Compton camera setup composed by a scatterer component consisting of a pixelated GAGG crystal read out by a SiPM multi-pixel photon counter (MPPC) and an absorber component consisting of a monolithic LaBr 3 (Ce) scintillator read out by a 256-fold multianode photomultiplier (PMT). The rationale of the study is to develop a Compton camera system as a future ion beam range verification device during particle therapy, via prompt gamma imaging. The properties to be investigated are the reconstruction efficiency and accuracy achievable with this system for detecting prompt-$\gamma$ rays. The Compton camera system described has been tested with a laboratory radioactive Cesium137 source, in a certain geometrical configuration. The readout system is based on individual spectroscopy (NIM+VME) electronic modules, digitizing energy and time signals. The data have been analyzed to produce an input for the image reconstruction, performed using the MEGAlib toolkit software.

We have developed the brain-dedicated compact PET scanner which has a hemispherical helmet detector unit and an add-on detector unit located at the neck position (helmet-neck PET). To realize clinical applications of the helmet-neck PET scanner, we need to evaluate the performance regarding their absolute quantitation, image contrast, noise and uniformity using appropriate phantoms. However, existing standard phantoms are not applicable for the helmet-neck PET scanner because of its hemisphere geometry. Therefore, in this study, we developed two new phantoms, a small sphere contrast phantom and a 3dimensional hemispherical Hoffman brain phantom, which model a brain tumor and static regional cerebral blood flow, respectively. These phantoms have an adaptive structure for use with the PET scanner with hemispherical detector arrangement. In addition, we tested the developed contrast phantom using the helmet-neck PET prototype. The PET image obtained by the helmet-neck PET prototype showed high image contrast and acceptable noise level. These phantoms are useful for evaluating the imaging performance of compact helmet-shape scanners.

We are developing a compact and high-sensitivity brain-dedicated PET scanner, which consists of a hemispherically arranged detector unit and an add-on detector unit. Following our first idea of using a chin position as the location of the add-on detector (helmet-chin PET), a neck position has been selected as the location (helmet-neck PET). We have developed prototypes of both geometries using the 4-layered depth-ofinteraction (DOI) detector (2.8 mm sized GSOZ scintillators and a 64 ch PMT) which does not have time-of-flight (TOF) capability. On the other hand, a new TOF-PET detector module which consists of 4.1 mm sized LFS scintillators and multi pixel photon counters (MPPCs) is now commercially available (C135004075LC-12, Hamamatsu Photonics K.K.). Using one-to-one crystal-photodetector coupling, it is possible to achieve a 280 ps coincidence resolving time (CRT), although the Anger-type calculation, which is essential for the 4-layered DOI detector, cannot be applied. In this paper, therefore, we investigated the suitability of this TOF-PET detector module for the next helmetneck PET. At first, the helmet-neck PET with 20 mm crystal length, which is a standard parameter of the module, was modeled by GEANT4, and performance results were compared with those of the current helmet-neck PET prototype. We observed a compromised image quality; the non-DOI capability counteracted the TOF gain. On the other hand, the helmet-neck PET with shortened crystal length (10 mm) outperformed the current helmet-neck PET prototype; the shortened crystal length reduced the parallax error, and TOF information compensated for the loss of sensitivity. Based on the GEANT4 simulation results, we fabricated a one-pair prototype of the TOF-PET module with 10 mm length, which yielded 11.6 % energy resolution at 511 keV and 281.8\pm 9.6 ps CRT at FWHM. In conclusion, we confirmed suitability of the 280 ps-CRT non-DOI detector for the next helmet-neck PET.

In a conventional PET scanner, coincidence events are measured with a limited energy window for detection of photoelectric events in order to reject Compton scatter events that occur in a patient, but Compton scatter events caused in detector crystals are also rejected. Scatter events within the patient causes scatter coincidences, but inter crystal scattering (ICS) events have useful information for determining an activity distribution. Some researchers have reported the feasibility of PET scanners based on a Compton camera for tracing ICS into the detector. However, these scanners require expensive semiconductor detectors for high-energy resolution. In the Anger-type block detector, single photons interacting with multiple detectors can be obtained for each interacting position and complete information can be gotten just as for photoelectric events in the single detector. ICS events in the single detector have been used to get coincidence, but single photons interacting with multiple detectors have not been used to get coincidence. In this work, we evaluated effect of sensitivity improvement using Compton kinetics in several types of DOI-PET scanners. The proposed method promises to improve the sensitivity using coincidence events of single photons interacting with multiple detectors, which are identified as the first interaction (FI). FI estimation accuracy can be improved to determine FI validity from the correlation between Compton scatter angles calculated on the coincidence line-of-response. We simulated an animal PET scanner consisting of 42 detectors. Each detector block consists of three types of scintillator crystals (LSO, GSO and GAGG). After the simulation, coincidence events are added as information for several depth-of-interaction (DOI) resolutions. From the simulation results, we concluded the proposed method promises to improve the sensitivity considerably when effective atomic number of a scintillator is low. Also, we showed that FI estimate accuracy is improved, as DOI resolution is high.

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.

Parallax error caused by the detector crystal thickness degrades spatial resolution at the peripheral regions of the field-of-view (FOV) of a scanner. To resolve this issue, depth-of-interaction (DOI) measurement is a promising solution to improve the spatial resolution and its uniformity over the entire FOV. Even though DOI detectors have been used in dedicated systems with a small ring diameter such as for the human brain, breast and small animals, the use of DOI detectors for a large bore whole-body PET system has not been demonstrated yet. We have developed a four-layered DOI detector, and its potential for a brain dedicated system has been proven in our previous development. In the present work, we investigated the use of the four-layer DOI detector for a large bore PET system by developing the world's first whole-body prototype. We evaluated its performance characteristics in accordance with the NEMA NU 2 standard. Furthermore, the impact of incorporating DOI information was evaluated with the NEMA NU 4 image quality phantom. Point source images were reconstructed with a filtered back projection (FBP), and an average spatial resolution of 5.2 ± 0.7 mm was obtained. For the FBP image, the four-layer DOI information improved the radial spatial resolution by 48% at the 20 cm offset position. The peak noise-equivalent count rate (NECR) was 22.9 kcps at 7.4 kBq ml-1 and the scatter fraction was 44%. The system sensitivity was 5.9 kcps MBq-1. For the NEMA NU 2 image quality phantom, the 10 mm sphere was clearly visualized without any artifacts. For the NEMA NU 4 image quality phantom, we measured the phantom at 0, 10 and 20 cm offset positions. As a result, we found the image with four-layer DOI could visualize the 2 mm-diameter hot cylinder although it could not be recognized on the image without DOI. The average improvements in the recovery coefficients for the five hot rods (1-5 mm) were 0.3%, 4.4% and 26.3% at the 0, 10 and 20 cm offset positions, respectively (except for the 1 mm-diameter rod at the 20 cm offset position). Although several practical issues (such as adding end-shields) remain to be addressed before the scanner is ready for clinical use, we showed that the four-layer DOI technology provided higher and more uniform spatial resolution over the FOV and improved contrast for small uptake regions located at the peripheral FOV, which could improve detectability of small and distal lesions such as nodal metastases, especially in obese patients.

A strong demand is expected for high-sensitivity, high-resolution and low-cost brain positron emission tomography (PET) imaging for early diagnosis of dementia, as well as for general neuroscience studies. Therefore, we have proposed novel geometries of a hemisphere detector arrangement for high-sensitivity brain imaging, in which an add-on detector at the chin position or neck position helps in sensitivity uniformity improvement. In this study, we developed the first prototype system for proof-of-concept using four-layer depth-of-interaction detectors, each of which consisted of 16 × 16 × 4 Zr-doped GSO crystals with dimensions of 2.8 × 2.8 × 7.5 mm3 and a high-sensitivity 64-channel flat-panel photomultiplier tube. We used 47 detectors to form a hemisphere detector with a hemisphere shape of 25 cm inner diameter and 50 cm outer diameter, and we used seven detectors for each of the add-on detectors. The total detector number of 54 was about one-fourth that of a typical whole-body PET scanner. The hemisphere detector for the prototype system was realized by multiple rings having different numbers of detectors and a cross-shaped top detector unit covering the top. Performance evaluation showed uniform spatial resolutions of 3-4 mm by the filtered back-projection method. Imaging tests of a hot-rod phantom done with an iterative method were able to resolve 2.2 mm rods. Peak sensitivity was measured as more than 10% at a region near the top of the head, which was achieved with the help of the top detector unit. In addition, using the prototype system, we performed the first FDG clinical test with a healthy volunteer. The results showed that the proposed geometries had high potential for realizing high-sensitivity, high-resolution, and low-cost brain PET imaging. As for the add-on detector position, it was shown that the neck position resulted in higher sensitivity and wider field of view (FOV) than the chin position because the add-on detector at the neck position can be placed continuously to the hemisphere detector and close to the FOV.

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.

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.

We developed a modified, more practical four-layered depth-of-interaction (DOI) detector based on the light sharing method. Reflectors, which are inserted in every two lines of crystal segments and shifted differently depending on each layer, project 3-D crystal positions onto a 2-D position histogram without any overlapping after applying an Anger-type calculation. The best crystal separation we have ever made based on this method was the 4-layered 32 × 32 array of LYSO crystals sized at 1.45 × 1.45 × 5 mm3. However, assembling crystals of a tiny size tends to cost a lot, and fine tuning of the light guide and the front-end circuit is required to have fine crystal identification from photo sensor signals of coarser pixel pitch. In this paper, therefore, we proposed a more practical 4-layered DOI detector. The key idea is that the crystals in the top layer, which have the highest detection efficiency, mostly contribute to PET spatial resolution. We applied two new ideas: (1) use of 1/4 size crystals only for the 1st (top) layer and (2) inserting a thin light guide between the 1st and the 2nd layers of crystal array. In the developed prototype detector, the 1st layer used 32 × 32 LYSO crystals of quarter size (1.4 × 1.4 × 5.0 mm3) compared with the other layers (16 × 16 arrays of crystals of 2.8 × 2.8 × 5.0 mm3). For better crystal identification of small crystals in the 1st layer, we optimized the optical condition between crystals such as use of an optical glue or air. Also, a thin light guide was inserted between the 1st and the 2nd layers for improvement of crystal identification of the 1st layer. With the appropriate insertion of the light guide, all crystals of the 1st layer were identified as well as the crystals in the other layers.

High sensitivity, high spatial resolution and low cost dedicated brain PET scanners are required for early diagnosis of Alzheimer's disease and brain function studies. As an optimal geometry, we proposed and developed the first prototype of the helmet-chin PET scanner. The scanner was constructed from 4-layer DOI detectors constructed from GSO crystals which were originally developed for our OpenPET. The helmet part of the helmet-chin PET consisted of three ring detectors with different radii arranged on a surface of a hemisphere and a top cover. In this study, for our next development, we optimized the size of the detectors to be arranged on the helmet, and compared two types of geometrical arrangements of the detectors on the hemisphere: a spherical arrangement in which the center of each detector faces toward the center of the hemisphere, and a multi-ring arrangement which has a similar detector arrangement to that of the first prototype. Geant4 simulation toolkit was used to model the scanners. The simulated scanners were constructed from LYSO crystals with a size of 1 × 1 × 5 mm3 (transaxial × axial × DOI). A dead-space of 2 mm was assumed in the axial and transaxial directions of the detector such as for wrapping. While fixing the number of the DOI layers to 4, the size of the detectors was varied by changing the number of crystals in the transaxial and axial directions. The results showed that, for the helmet detector, a detector with a size of 42 × 42 × 20 mm3 had the highest sensitivity evaluated using a hemispherical phantom. Then, the helmet-chin PET was modeled based on a detector whose size is 42 × 42 × 20 mm3 with the two geometrical arrangements and their sensitivities were compared. The respective sensitivities for hemispherical and multi-ring arrangements were 3.4% and 2.9% for a hemispherical phantom.

PET is the most sensitive imaging technique in the field of nuclear medicine. However, the single count rate is typically 10 times higher than the coincidence count rate. This means that there are a lot of pairs of undetected annihilation photons. For single gamma imaging, Compton imaging is known as the only method to obtain directional information of an incoming gamma ray without using collimators. Compton kinematics collimation suffers from poor angular resolution due to Doppler broadening, energy blurring, and position blurring. Therefore, Compton imaging is still challenging compared to PET imaging in general. However, non-coincidence single photons have the potential to contribute to improved imaging performance. Previously, we have proposed a system which combines PET and Compton imaging and we showed high sensitivity of this system using Compton kinematics collimation events. In this paper, we evaluate sensitivity and spatial resolution for feasibility of this system. This system modeled the PET ring and the inserted Compton scatter ring. The PET ring consisted of 48 DOI detectors arranged in five rings with a ring diameter of 80 cm. The PET ring was made of GAGG with a size of 2.9 × 2.9 × 5 mm3 including 4-layered DOI capability. Ring diameters of the scatter ring were selected as 30, 40, 50 and 60 cm. Detector material of the scatter ring was Si with a pixel size of 0.9 × 0.9 × 0.6 mm3. From point source measurements, images of PET and Compton events were individually reconstructed by the 3D-LM-OSEM algorithm. From point source measurements at the center of the field-of-view, spatial resolution with 30 cm diameter was 3 mm. Also, as the position of the point source was close to the detector, spatial resolution was improved. The proposed system had a potential of high spatial resolution for not only PET imaging but also Compton imaging with the small scatter ring.

We are developing a PET/MRI system based on 4-layered depth-of-interaction (DOI) detectors integrated with a birdcage RF-coil. The PET detectors are placed close to the objective to achieve both high sensitivity and high spatial resolution even at the edge of the field of view (FOV). In a previous study, we reported the first prototype which had a single detector ring to show a proof-of-concept and lower than 1.6 mm spatial resolution was achieved. However, the axial FOV of the prototype was only 1 cm, which could not be extended due to the size of the readout circuit boards. In this paper, we finally constructed a new RF-coil for the second prototype, which is combined with 24 DOI-PET detectors. We evaluated performance in simultaneous measurements. The second prototype consisted of an RF-coil and 24 PET detector units with two PET detectors each. The detector units were shielded with carbon fiber shielding boxes. Each PET detector consisted of a 14 × 14 × 4-layer array of lutetium fine silicate (LFS) crystals (1.9 mm × 1.9 mm × 4.0 mm), an 8 × 8 multi-pixel photon counter module (MPPC, S11206-0808FC(X)) and a readout circuit board with an ASIC. The RF coil was dedicated to a 3T MRI (MAGNETOM Verio, Siemens). We conducted performance tests of the second prototype in simultaneous measurements. As a result, the influence of the MRI measurements on the PET performance was found to be negligible. In addition, secondary magnetic field due to the eddy current effect can be reduced, compared with the previous prototype system.

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 develop an OpenPET system which can provide an accessible open space to the patient during PET scanning. Our first-generation OpenPET geometry which we called dual-ring OpenPET consisted of two separated detector rings and it could extend its axial field of view (FOV) therefore enabling imaging the gap region in addition to the in-ring region. However, applications such as dose verification by in-beam PET measurement during particle therapy and real-time tumor tracking by PET require sensitivity focused onto the gap rather than on the wide FOV. In this paper, we propose a second-generation OpenPET geometry, single-ring OpenPET, which can provide an accessible and observable open space with higher sensitivity and a reduced number of detectors than the earlier one. The proposed geometry has a cylinder shape cut at a slant angle, in which the shape of each cut end becomes an ellipse. We provided a theoretical analysis for sensitivity of the proposed geometry, compared with the dual-ring OpenPET and a geometry where the conventional PET was positioned at a slant angle against the patient bed to form an accessible open space, which we called a slant PET. The central sensitivity depends on the solid angle of these geometries. As a result, we found that the single-ring OpenPET has a sensitivity 1.2 times higher than the dual-ring OpenPET and 1.3 times higher than the slant PET when designed for a 600 mm bed width with 300 mm accessible open space and about 200 detector blocks, each with a front area of 2500 mm². In addition, numerical simulation was carried out to show the imaging property of the proposed geometry realized with the ellipsoidal rings and these results indicate that the depth-of-interaction detector can provide uniform resolution even when the detectors are arranged in an ellipsoidal ring.

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.