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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 99mTc (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 μm2 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.

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.

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.

Simultaneous measurements of multiple physiological parameters are essential for the study of brain disease mechanisms and the development of suitable therapies to treat them. In this study, we developed a measurement system for simultaneous optical imaging and PET for awake mice. The key elements of this system are the OpenPET, optical imaging and fixation apparatus for an awake mouse. The OpenPET is our original open-type PET geometry, which can be used in combination with another device because of the easily accessible open space of the former. A small prototype of the axial shift single-ring OpenPET was used. The objective lens for optical imaging with a mounted charge-coupled device camera was placed inside the open space of the AS-SROP. Our original fixation apparatus to hold an awake mouse was also applied. As a first application of this system, simultaneous measurements of cerebral blood flow (CBF) by laser speckle imaging (LSI) and [(11)C]raclopride-PET were performed under control and 5% CO2 inhalation (hypercapnia) conditions. Our system successfully obtained the CBF and [(11)C]raclopride radioactivity concentration simultaneously. Accumulation of [(11)C]raclopride was observed in the striatum where the density of dopamine D2 receptors is high. LSI measurements could be stably performed for more than 60 minutes. Increased CBF induced by hypercapnia was observed while CBF under the control condition was stable. We concluded that our imaging system should be useful for investigating the mechanisms of brain diseases in awake animal models.

We previously have developed the photoresponsive tetra-gel and nanoparticles for controlling the function of the encapsulated substance by UV irradiation. However, the penetration ability of the UV is not high enough. Here, we developed a radiation-responsive tetra-gel and nanoparticle based on γ-ray-responsive X-shaped polyethylene glycol (PEG) linker with a disulfide bond. The nanoparticle could retain small molecules and biomacromolecules. γ-Rays were used as a trigger signal because of their higher penetrating ability. This allowed a spatiotemporal release and control of the encapsulated substances from the nanoparticle in the deeper region, which is impossible by using light exposure (ultraviolet, visible, and near-infrared).

Radioactive ion (RI) beams combined with in-beam positron emission tomography enable accurate in situ beam range verification in heavy ion therapy. However, the energy spread of the radioactive beams generated as secondary beams is wider than that of conventional stable heavy ion beams which causes Bragg peak region and distal falloff region broadening. Therefore, the energy spread of the RI beams should be measured carefully for their quality control. Here, we proposed an optical imaging technique for the energy spread estimation of radioactive oxygen ion beams. A polymethyl methacrylate phantom (10.0 × 10.0 × 9.9 cm3) was irradiated with an 15O beam (mean energy = 247.7 MeV u−1, standard deviation = 6.8 MeV u−1) in the Heavy Ion Medical Accelerator in Chiba. Three different momentum acceptances of 1%, 2% and 4% were used to get energy spreads of 1.9 MeV u−1, 3.4 MeV u−1 and 5.5 MeV u−1, respectively. The in-beam luminescence light and offline beam Cerenkov light images were acquired with an optical system consisting of a lens and a cooled charge-coupled device camera. To estimate the energy spread of the 15O ion beams, we proposed three optical parameters: (1) distal-50% falloff length of the prompt luminescence signals; (2) full-width at half maximum of the Cerenkov light signals in the beam direction; and (3) positional difference between the peaks of the Cerenkov light and the luminescence signals. These parameters estimated the energy spread with the respective mean squared errors of 2.52 × 10−3 MeV u−1, 5.91 × 10−3 MeV u−1, and 0.182 MeV u−1. The distal-50% falloff length of the luminescence signals provided the lowest mean squared error among the optical parameters. From the findings, we concluded optical imaging using luminescence and Cerenkov light signals offers an accurate energy spread estimation of 15O ion beams. In the future, the proposed optical parameters will be used for energy spread estimation of other RI beams as well as stable ion beams.

Positronium (Ps), a hydrogen-like atom consisting of a positron and an electron, is efficiently formed in the human body during positron emission tomography (PET) examination, and its decay rate into gamma-ray photons is significantly influenced by the chemical environment, especially the dissolved oxygen concentration (pO2) due to the unpaired electrons. However, the functionality of PET has been underestimated by neglecting the specific information provided by Ps. By comparing the decay rates in O2-, N2-, and air-saturated waters, here we show that Ps probes the absolute value of pO2 with a good linearity and a resolution better than 10 mmHg. This is a sufficient sensitivity for discriminating a hypoxic region in a tumor at approximately 6 mmHg from healthy tissues at approximately 40 mmHg. This method depends only on the fundamental properties of Ps and is independent of specific radiopharmaceuticals. The applications of Ps spin states and reactions will greatly enhance PET functionalities in the next decade.

Positron emission tomography (PET) is a practical tool for range verification of hadron therapy. As well, the quantitative washout of the positron emitters has a potential usefulness as a diagnostic index, but the modelling for this has not been established. In this study, we measured washout rates of rabbit brain and performed kinetic analysis to explore the washout mechanism. Six rabbit brains were irradiated by 11C and 15O ion beams, and dynamic PET scan was performed using our original depth of interest (DOI)-PET prototype. The washout rate was obtained based on the two-compartment model, where efflux from tissue to blood (k2), influx (k3) and efflux (k4) from the first to second compartments in tissue were evaluated. The observed k2, k3 and k4 of 11C were 0.086, 0.137 and 0.007 min−1, and those of 15O were 0.502, 0.360 and 0.007 min−1, respectively. It was suggested permeability of a molecule containing 11C atoms might be regulated by a transporter. The k2 of 15O was comparable with 15O-water. This study provides basic data for modelling of the washout effect.

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.

In heavy-ion therapy, the stopping position of primary ions in tumours needs to be monitored for effective treatment and to prevent overdose exposure to normal tissues. Positron-emitting ion beams, such as 11C and 15O, have been suggested for range verification in heavy-ion therapy using in-beam positron emission tomography (PET) imaging, which offers the capability of visualizing the ion stopping position with a high signal-to-noise ratio. We have previously demonstrated the feasibility of in-beam PET imaging for the range verification of 11C and 15O ion beams and observed a slight shift between the beam stopping position and the dose peak position in simulations, depending on the initial beam energy spread. In this study, we focused on the experimental confirmation of the shift between the Bragg peak position and the position of the maximum detected positron-emitting fragments via a PET system for positron-emitting ion beams of 11C (210 MeV u-1) and 15O (312 MeV u-1) with momentum acceptances of 5% and 0.5%. For this purpose, we measured the depth doses and performed in-beam PET imaging using a polymethyl methacrylate (PMMA) phantom for both beams with different momentum acceptances. The shifts between the Bragg peak position and the PET peak position in an irradiated PMMA phantom for the 15O ion beams were 1.8 mm and 0.3 mm for momentum acceptances of 5% and 0.5%, respectively. The shifts between the positions of two peaks for the 11C ion beam were 2.1 mm and 0.1 mm for momentum acceptances of 5% and 0.5%, respectively. We observed larger shifts between the Bragg peak and the PET peak positions for a momentum acceptance of 5% for both beams, which is consistent with the simulation results reported in our previous study. The biological doses were also estimated from the calculated relative biological effectiveness (RBE) values using a modified microdosimetric kinetic model (mMKM) and Monte Carlo simulation. Beams with a momentum acceptance of 5% should be used with caution for therapeutic applications to avoid extra dose to normal tissues beyond the tumour when the dose distal fall-off is located beyond the treatment volume.

Positron emission tomography (PET) has been used for dose verification in charged particle therapy. The causes of washout of positron emitters by physiological functions should be clarified for accurate dose verification. In this study, we visualized the distribution of irradiated radioactive beams, 11C and 15O beams, in the rabbit whole-body using our original depth-of-interaction (DOI)-PET prototype to add basic data for biological washout effect correction. Time activity curves of the irradiated field and organs were measured immediately after the irradiations. All data were corrected for physical decay before further analysis. We also collected expired gas of the rabbit during beam irradiation and the energy spectrum was measured with a germanium detector. Irradiated radioactive beams into the brain were distributed to the whole body due to the biological washout process, and the implanted 11C and 15O ions were concentrated in the regions which had high blood volume. The 11C-labelled 11CO2 was detected in expired gas under the 11C beam irradiation, while no significant signal was detected under the 15O beam irradiation as a form of C15O2. Results suggested that the implanted 11C ions form molecules that diffuse out to the whole body by undergoing perfusion, then, they are incorporated into the blood-gas exchange in the respiratory system. This study provides basic data for modelling of the biological washout effect.

Research on the distribution and dynamics of photoassimilates in plants, especially those in fruits, is important for increasing food production. Positron emission tomography (PET) and carbon-11 ( 11C) can be used for obtaining 3-D images of photoassimilates. For plant experiments, however, it is important to adjust the system to the plant’s growth environment. A small OpenPET prototype can visualize an open space between two detector rings, and thus facilitate adjustment of the environment around a target plant during fruit PET imaging. In this work, we improve the OpenPET for plant studies and discuss its applications in the analysis of the dynamics and distributions of photoassimilates in plants with fruits. 11C-labeled photoassimilate translocation into the fruit of a strawberry plant is imaged using 11C-labeled CO2 and the OpenPET. The OpenPET has been found to be promising for researching the physiological functions of the photoassimilates in plants with fruits.

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-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 × 12 array of multi pixel photon counters (MPPCs) is connected to a lutetium fine silicate (LFS) crystal array of 4.1 × 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 × 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.

Brain PET imaging has important roles in neurology, neuro-oncology and molecular imaging research. We have developed a helmet-type PET prototype and have shown that the proposed hemispherical geometry had high potential for realizing high-sensitivity and low-cost brain imaging. However, there is no standard performance evaluation method for helmet-type PET, which would be a bottleneck to its commercialization. Therefore, we investigated appropriate performance evaluation methods for a helmet-type PET based on the NEMA NU 2-2018 standards. For those measurement methods that are not applicable to the helmet-type PET, we changed them while keeping the basic concept of the original NEMA standards. We measured spatial resolution, sensitivity, scatter fraction, count rate characteristics, accuracy of corrections for count losses and randoms, and image quality. We partially changed the measurement methods by making brain-size phantoms and by optimizing the length or the position of radioactive sources. The spatial resolution was 2.8 mm at 1-cm offset position by the filtered back-projection method. Sensitivities measured by the NEMA original setup and the proposed setup were 13.4 and 57.1 kcps/MBq. The respective values measured with our developed brain-size scatter phantom and with the conventional whole-body-size scatter phantom were: scatter fractions of 35% and 35%; peak NECRs of 25.1 kcps at 3.2 kBq/ml and 19.8 kcps at 2.6 kBq/ml; and maximum absolute biases of 5.5% and 16.0%. The image quality was evaluated with the developed brain-size phantom, and good image quality was obtained. The helmet-type PET prototype showed high-sensitivity even with the small number of 54 detectors. The spatial resolution was better than 4.0 mm over the field-of-view. In conclusion, we proposed the performance evaluation methods for a brain-dedicated PET system with a hemispherical geometry. The proposed method could facilitate evaluation of performance characteristics of brain-dedicated PET scanners and optimization of its scanning and reconstruction parameters.

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.

Single photon emission computed tomography (SPECT) has been widely used for preclinical research. High resolution small animal SPECT imaging can be achieved by employing a pinhole or multi-pinhole collimators. However, the useful FOV is narrow compared to that of a parallel-hole collimator. The aim of this study is to optimize the hole diameter of a parallel square-hole collimator for high resolution whole-body small animal SPECT imaging using GATE and STIR software. The small animal SPECT consists of a tungsten parallel square-hole collimator (L= 30 mm), a monolithic GAGG crystal (49 × 49 × 5 mm3), and a PMT (H9500). The SPECT images of an Ultra-Micro hot phantom and NEMA NU-4 phantom were acquired with different square-hole sizes of 0.2, 0.4, 0.6, 0.8, 1.0, and 1.2 mm. The collimator length and septal thickness were kept as 30 mm and 0.1 mm, respectively. The projection image (128 × 128 matrix, pixel size of 0.3828 × 0.3828 mm2) was reconstructed with FBP, and OSMAPOSL algorithms provided by STIR software. With the collimator hole size of 0.4 mm, the spatial resolution of 1.11 mm and sensitivity of 0.001% could be achieved. The contrast recovery coefficient could be improved by using the 0.4 mm hole size and the hot rods (D= 1.35 mm) could be identified clearly. The important message from this work is that the small animal SPECT system can be optimized with the open-source software packages of GATE and STIR. In the future, a prototype small animal SPECT will be developed based on the GATE simulation results.

A modular radiofrequency (RF) shielded PET insert, electrically floating from the RF ground, enables the RF field from outside to penetrate through the inter-modular gaps to the imaging region (ROI) with less attenuation compared to the electrically grounded PET insert. In this study, we developed an electrically floating PET insert by implementing conventional electric data transmission and power source cables and, performed a comparison with the same PET ring as a grounded PET insert inside a 3 T MRI system. An existing electrical data transmission capable RF shielded PET detector ring is modified to perform the study. The PET inserts, with the inner diameter of 240-mm, had eight RF shielded PET detector modules with intermodular gap of 5 mm. Here the term ground PET is used to mean that both the RF shields of cables and the PET detector modules are together connected to the RF ground. On the other hand, the term floating PET is used to mean that the RF shield boxes of PET detector modules kept electrically isolated from the cable shield that remained connected to RF ground. The MRI system built-in body RF coil (700-mm in diameter) is used both as RF transmitter and receiver. A homogeneous phantom study revealed an approximate 10% and 20% reduction in mean RF field attenuation for the front 125 mm axial-FOV of respectively floating and ground PET insert. Compared to the front of PET insert, an approximate 20% and 40% field attenuation was seen for an imaging region close to the PET insert cable-end. However, RF field homogeneity for both the front and center of the PET insert was comparable to the MRI-only values.

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.

Nuclear medicine imaging is an important tool for cancer diagnosis, brain research, molecular imaging research and so on. Therefore, various imaging techniques and methods are being developed and investigated in nuclear medicine physics. In this report, we introduce state-of-the-art techniques, such as Compton camera imaging, time-of-flight positron emission tomography, semiconductor detectors for medical applications, image reconstruction and deep learning, which were reported in the 2017 IEEE Nuclear Science SymposiumMedical Imaging Conference.

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 world's first open type PET, named OpenPET, is being developed at the National Institute Radiological Sciences. We are aiming to employ the OpenPET in radiotherapy and to track the respiratory motion of a tumor in the thoracoabdominal region of a patient. By using PET images, we expect that the tumor itself can be directly visualized without using radio-opaque markers, but this goal has not been achieved yet. Our demonstration results using a small prototype OpenPET system showed that the system can output reconstructed images at about 2 frames per second with about a 2 s delay; this delay is mainly due to the reconstruction calculation time. In this study, we developed a time delay correction method for tumor tracking for the OpenPET. Since it is difficult to correct the time delay using only the tumor location 2 s before, we introduced another sensor to acquire the respiration phase for correction. In the proposed method, the relationship between tumor motion and the additional sensor output signal was calculated by support vector regression (SVR) and the time delay was corrected by using the regression line which represents the relationship. We simulated this time delay correction method with computer-generated PET images which had practical respiration motions obtained from clinical MRI. As a result, we could track the tumor within almost 1.5 mm mean error when we assumed the use of a belt type respiratory motion sensor. We also found from the simulation that disturbance of the relationship can be detected by gates. Thus we could avoid the tracking errors which are caused by disturbance of the correlation.

PET is recognized as a successful method to pursue cancer diagnosis and molecular imaging. However, in order to meet emerging demands for widened imaging applications such as textbfin-situ real-time single-cell tracking, we need to break through the principle of PET itself. In this paper, therefore, we propose a new concept of whole gamma imaging (WGI), which is a novel combination of PET and Compton imaging. An additional detector ring, which is used as the scatterer, is inserted in a conventional PET ring so that single gamma rays can be detected by the Compton imaging method. Therefore, in addition to Compton imaging (single-gamma mode), missing pairs of annihilation photons in PET, at least one of which is undetected, can be used for imaging (PET mode). Further large sensitivity gain can be expected for triple gamma emitters such as $^{\mathbf {44}}$textbfSc, that emits a pair of 511 keV photons and a 1157 keV gamma ray almost at the same time (triple-gamma mode). In principle, only a single decay would be enough to localize the source position: (1) the coincidence detection of a pair of 511keV photons locates the source position along a line-of-response (LOR), and (2) the source position is identified as one of two intersection points of the LOR with a Compton cone after measuring the 1157 keV gamma ray. Using GEANT4, we simulated an “insert geometry”, in which a scatter ring (24 x 24 array of 1 x 1 x 6 mm$^{\mathbf {3}}$ GAGG crystals, 20 cm diameter and 5 cm long,) was inserted into a PET ring (16 × 16 × 4-DOI array of 2.9 × 2.9 × 7.5 mm$^{\mathbf {3}}$ GSOZ crystals, 66 cm diameter and 22 cm long). In the single-gamma mode, spatial resolution for the 511keV source obtained by 3D OSEM was 6.2 mm FWHM (center)-3.0 mm FWHM (8 cm off-center). In the triple-gamma mode, the position distribution of a $^{\mathbf {22}}$Na point source projected on a line-of-response was 7.3 mm FWHM at the 5 cm off-center position without applying any image reconstruction. From the simulation results, we were able to develop the first prototype of the WGI system.

Studies on cell regulation are attracting worldwide attention in order to realize regenerative medicine. Therefore, a nuclear medicine imaging method, which can use tracers having substantially the same composition as a target biomolecule, is required. Autoradiography is a high-resolution nuclear medicine imaging method. However, this method does not have a dynamical imaging capability due to the principle on which it is based. In this research, we develop a nuclear medicine imaging system for dynamic cell observation. Specifically, in order to prevent broadening of the scintillation position because of scintillation light spreading in the scintillator, β-rays are detected by a thin scintillator plate. A scientific CMOS camera with low readout noise and high resolution was used to detect scintillation light. The scintillator plate was a CsI crystal connected to an optical fiber array. The scintillation light generated from the scintillator plate was extracted through optical fibers. The imaging lenses adopted in this research were a high-speed imaging and a conversion lens.Imaging resolution was adjusted to 6.5 × 6.5 mm2 / pixel and exposure time was 10 s. In order to evaluate the imaging performance of the proposed system, we measured samples containing a radiopharmaceutical with a prototype system. In this measurement, two different activity 18F-solutions were used as the imaging target. They were fixed with a uniform thickness of 0.4 mm. In the imaging results, we measured different image values depending on their activities. On the other hand, in the region of no 18F-solutions, the image value was almost constant. The output had sufficient linearity for images with activity exceeding 0.05 Bq / pixel. Imaging results showed that our system has sufficient sensitivity for imaging uptake of 18F by single cells.

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.

In some recently developed brain-dedicated PET systems, a seated position is selected for the patient. The PET scanners with the seated measurement style have a merit of allowing system downsizing. However, the effect of the seated position for head motion suppression compared with the conventional supine position is not clear. Although motion correction methods have been developed, they are not in practical use at a routine level. Even when using the motion correction method, it is desirable to make the head motion as small as possible. In this study, we developed a contactless head motion tracking system, and we conducted a volunteer study of the head motion tracking to assess the optimum measurement position for the brain PET. We developed the tracking system using the Microsoft Kinect sensor as a range sensor. As a result of the accuracy evaluation, translation accuracy of about 1 mm and rotation accuracy of less than 1 deg were achieved. In the volunteer study, we measured the head motion of volunteers with supine, normal sitting, and reclining. Additionally, we measured these positions with and without head fixation. As a result, the normal sitting position without head fixation had the largest head motion, and the reclining position had smaller motion than the supine position. The head fixation was effective for motion suppression in all cases. The lowest head motion was achieved with the reclining and supine positions (average displacement in 60 s was about 0.5 mm). In order to assess the effect of the head motion, we added the motion information obtained with the volunteer study to brain phantom data measured with our helmet-chin PET prototype. In the case of the reclining with head fixation, we could obtain reconstruction images of equal quality to the case without motion even though we did not apply any motion correction. We concluded that the reclining position had a high head motion suppression effect equal to or better than that of the supine position. （ M-15-058）, IEEE NSS-MIC2017

Dose monitoring and range verification are important tools in carbon ion therapy. For their implementation, positron emission tomography (PET) can be used to image the $\beta^{+}$-activation of tissue during treatment. Predictions of these $\beta^{+}$-activity distributions are usually obtained from Monte Carlo simulations, which demands high computational time and thus limits the applicability of this technique in clinical scenario. Nevertheless, it is desirable to explore faster approaches able to give such a prediction, since only its comparison with the measured distributions allows a definite assessment of potential range deviations from the planned treatment. For the first time, we present an approach to perform deconvolution from PET data in carbon ion therapy and reconstruct the dose. A filtering method is used to predict positron emitter profiles from dose profiles in short time. In order to reverse the convolution and estimate a dose distribution from a positron emitter distribution, we apply an evolutionary algorithm. Filters are obtained from either a library or are created in advance for a specific problem, assuming that a prediction of the positron emitter distribution is available. To perform the latter method and find the best filter for a specific problem, we use another evolutionary algorithm, hence optimizing the filter on-the-fly for the given treatment scheme. The application of our method is shown for dose and positron emitter distributions in homogeneous phantoms using simulated and newly measured online PET data. Carbon ion ranges can be predicted within 2 mm and the shape of the dose distribution is reconstructed with an overall promising agreement.

We have developed an oval shape RF-penetrable PET insert in a motivation for PET/MR body imaging with an existing 3T clinical MRI system. The MRI built-in body RF coil was used both as transmitter and receiver. We implemented 12 RF-shielded PET modules on a 440-mm major-axis and 350-mm minor-axis oval frame in which 8 PET modules contained PET detectors and electronics. The RF shielded PET detectormodules with electric data transmission cable were electrically floated from the MRI ground to use the MRI built in RF coil as transmitter and receiver. The system was positioned on the MRI patient bed. Experimental results showed about 6 percentagepoint RF field homogeneity reduction, about 50{\% SNR reduction for spin-echo sequence compare to MRI-only data for the 150mm diameter X 120 mm length cylindrical imaging region. Also, about 1.25 times increase in maximum local SAR was seen for the central 130 mm diameter transverse plane. The system successfully cleared a major step towards developing large scale oval PET insert for body imaging in an existing MRI system.

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.

The dominant factor limiting the intrinsic spatial resolution of a positron emission tomography (PET) system is the size of the crystal elements in the detector. To increase sensitivity and achieve high spatial resolution, it is essential to use advanced depth-of-interaction (DOI) detectors and arrange them close to the subject. The DOI detectors help maintain high spatial resolution by mitigating the parallax error caused by the thickness of the scintillator near the peripheral regions of the field-of-view. As an optimal geometry for a brain PET scanner, with high sensitivity and spatial resolution, we proposed and developed the helmet–chin PET scanner using 54 four-layered DOI detectors consisting of a 16 × 16 × 4 array of GSOZ scintillator crystals with dimensions of 2.8 × 2.8 × 7.5 mm3. All the detectors used in the helmet–chin PET scanner had the same spatial resolution. In this study, we conducted a feasibility study of a new add-on detector arrangement for the helmet PET scanner by replacing the chin detector with a segmented crystal cube, having high spatial resolution in all directions, which can be placed inside the mouth. The crystal cube (which we have named the mouth-insert detector) has an array of 20 × 20 × 20 LYSO crystal segments with dimensions of 1 × 1 × 1 mm3. Thus, the scanner is formed by the combination of the helmet and mouth-insert detectors, and is referred to as the helmet–mouth-insert PET scanner. The results show that the helmet–mouth-insert PET scanner has comparable sensitivity and improved spatial resolution near the center of the hemisphere, compared to the helmet–chin PET scanner.

Positron emission tomography (PET) plays important roles in cancer diagnosis, neuroimaging and molecular imaging research; but potential points remain for which big improvements could be made, including spatial resolution, sensitivity and manufacturing costs. Therefore, research on next generation PET technologies remains a hot topic worldwide. Depth-of-interaction (DOI) measurement in the radiation sensor will be a key technology to get any significant improvement in sensitivity while maintaining high spatial resolution. In this talk, we review PET innovations being made at the National Institute of Radiological Sciences (NIRS).

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.

Positron annihilation spectroscopy (PAS) spectra of biological tissue in 11C irradiation are reported and spatial resolution coefficient of positron emission tomography (PET) obtained from the PAS spectrum is discussed for 11C irradiation. A PAS spectrum of the biological tissue with water is the same as that of the water pool phantom in 11C irradiation. However, a PAS spectrum of the biological tissue with less water differs from that of the water pool phantom. The PET spatial resolution coefficient depends on the kind of biological tissue. However, the PET spatial resolution coefficient, 0.00243±0.00014, can be used as a common value of maximum limit.

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 are developing an open-type PET ‘OpenPET’ geometry. One possible geometry is a dual-ring OpenPET, which consists of two detector rings separated by a gap for entrance of a radiotherapy beam or for inserting other modalities. In our previous simulations and experiments the OpenPET imaging geometry was shown to be feasible by applying iterative reconstruction methods. However, the gap violates Orlov’s completeness condition for accurate tomographic reconstruction. In this study, we propose a solution for the incompleteness problem by adding bridge detectors to fill in parts of the gaps of the OpenPET geometry; we call this bridged OpenPET. Although this geometry was considered previously, its analytical property was not discussed. Therefore, we applied the direct Fourier method as an analytical reconstruction method to the bridged OpenPET, dual-ring OpenPET and conventional cylindrical PET for comparison. Numerical simulations showed that the additional bridge detectors compensate for the incompleteness of the OpenPET by covering one direction perpendicular to the transaxial slices of the imaging subjects.

For clinical studies, the effects of the intrinsic radioactivity of lutetium-based scintillators such as LSO used in PET imaging can be ignored within a narrow energy window. However, the intrinsic radioactivity becomes problematic when used in low-count-rate situations such as gene expression imaging or in-beam PET imaging. Time-of-flight (TOF) measurement capability promises not only to improve PET image quality, but also to reduce intrinsic random coincidences. On the other hand, we have developed a new reduction method for intrinsic random coincidence events based on multiple-coincidence information. Without the energy window, an intrinsic random coincidence is detected simultaneously with an intrinsic true coincidence as a multiple coincidence. The multiple-coincidence events can serve as a guide to identification of the intrinsic coincidences. After rejection of multiple-coincidence events detected with a wide energy window, data obtained included a few intrinsic random and many intrinsic true coincidence events. We analyzed the effect of intrinsic radioactivity and used Monte Carlo simulation to test both the TOF-based method and the developed multiple-coincidence-based (MC-based) method for a whole-body LSO-PET scanner. Using the TOF- and MC-based reduction methods separately, we could reduce the intrinsic random coincidence rates by 77 and 30 %, respectively. Also, the intrinsic random coincidence rate could be reduced by 84 % when the TOF+MC reduction methods were applied. The developed MC-based method showed reduced number of the intrinsic random coincidence events, but the reduction performance was limited compared to that of the TOF-based reduction method.

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.

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

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.