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In proton therapy, imaging of proton-induced positrons is a useful method to monitor the proton beam distribution after therapy. Usually, a positron emission tomography (PET) system installed in or near the proton beam treatment room is used for this purpose. However, a PET system is sometimes too large and expensive for this purpose. We developed a small field-of-view (FOV) gamma camera for high-energy gamma photons and used it for monitoring the proton-induced positron distribution. The gamma camera used 0.85 mm × 0.85 mm × 10 mm Ce:Gd3Al2Ga3O12 (GAGG) pixels arranged in 20 × 20 matrix to form a scintillator block, which was optically coupled to a 1-inch-square position-sensitive photomultiplier tube (PSPMT). The GAGG detector was encased in a 20-mm thick container and a pinhole collimator was mounted on its front. The gamma camera was set 1.2 m from the 35 cm × 35 cm × 5 cm plastic phantom in the proton therapy treatment room, and proton beams were irradiated to the phantom with two proton energies. The gamma camera had spatial resolution of ~6.7 cm and sensitivity of 3.2 × 10−7 at 1 m from the collimator surface. For both proton energies, positron distribution in the phantom could be imaged by the gamma camera with 10-min acquisition. The lengths of the range of protons measured from the images were almost identical to the simulation results. These results indicate that the developed high-energy gamma camera is useful for imaging positron distributions in proton therapy.

YSO (Ce-doped Y2SiO5) is a promising scintillator for a single-photon imaging system since it has relatively high light output and does not contain any natural radioactivity. Since YSO is not hygroscopic, it may be possible to fabricate a block with small pixels for a high-resolution system. For this purpose, we developed a high-resolution gamma camera system that employs smaller than 1-mm YSO pixels. The gamma camera’s detector used 0.8 × 0.8 × 7-mm YSO pixels. All the surfaces of these YSO pixels were mechanically polished, combined with a 0.1-mm-thick BaSO4 reflector to form a 48 × 48 matrix, and optically coupled to a high quantum efficiency, 2-inch square position sensitive photomultiplier tube (Hamamatsu Photonics H10966 A-100). The YSO block was 43.2 × 43.2 mm. The YSO gamma camera was encased in a 5-mm-thick tungsten container, and a parallel collimator was mounted on its front. The parallel hole collimator was made of a 3-layer (each layer was 5-mm thick) tungsten plate, and each plate had 48 × 48, 0.6-mm holes that were positioned by one-to-one coupling with the YSO pixels. Even with the 0.8-mm YSO pixels, we clearly resolved most of the pixels in a 2-dimensional histogram with a peak-to-valley ratio of 2.9 for the 122-keV gamma photons. The energy resolution was 20.4 % FWHM. The spatial resolutions with a parallel hole collimator 2 mm from the collimator surface were 0.7- and 1.3-mm FWHM for the 122- and ~35-keV gamma photons, respectively. We successfully obtained phantoms and small animal images with our YSO gamma camera system. Our high-resolution system has a potential to be useful for molecular imaging research.

Intra-operative probes have recently become important instruments in nuclear medicine. In such an application, the radiopharmaceutical F-18-fluorodeoxyglucose (FDG) is promising. For the FDG-guided surgery, we developed and tested a positron-imaging detector with background rejection capability.The detector consists of an array of phoswich scintillators, a multi-channel position-sensitive photo-multiplier tube (PSPMT) and an electronic circuit. The scintillators and the PSPMT are encased in a tungsten shield and replaceable collimators are mounted on the top of the detector. Positrons are detected by the plastic scintillators while annihilation photons are detected by the BGOs. By employing a pulse-shape analysis, we can distinguish the true events (positrons) from background gamma events. The dimensions of each plastic scintillator are 2 mm x 2 mm x 3 mm and those of the BGO are 2 mm x 2 mm x 15 mm. These scintillators are optically coupled to each other and combined in an 8 x 8 array, which is optically coupled to a 1-inch square 8 x 8 multi-channel PSPMT via optical fibers. Position determination of the positrons is performed by 64-channel threshold circuits while the pulse shape analysis is applied for the summing signal.The spatial resolution was measured by positioning an F-18 point source onto one pixel of the detector and found than the spillover to the neighbor pixel was less than 20%. The background count rate was less than 2 cps for a 20-cm diameter, 20-cm long cylinder phantom containing 3.7 MBq of F-18.These results indicated that the developed positron-imaging detector will be useful for FDG-guided surgery.

Pinhole SPECT which permits in vivo high resolution 3D imaging of physiological functions in small animals facilitates objective assessment of pharmaceutical development and regenerative therapy in pre-clinical trials. For handiness and mobility, the miniature size of the SPECT system is useful. We developed a small animal SPECT system based on a compact high-resolution gamma camera fitted to a pinhole collimator and an object-rotating unit. This study was aimed at evaluating the basic performance of the detection system and the feasibility of small animal SPECT imaging.The gamma camera consists of a 22 x 22 pixellated scintillator array of 1.8 mm x 1.8 mm x 5 mm NaI(Tl crystals with 0.2-mm gap between the crystals coupled to a 2" flat panel position-sensitive photomultiplier tube (Hamamatsu H8500) with 64 channels. The active imaging region of the camera was 43.8 mm x 43.8 mm. Data acquisition is controlled by a personal computer (Microsoft Windows) through the camera controller. Projection data over 360 degrees for SPECT images are obtained by synchronizing with the rotating unit. The knife-edge pinhole collimators made of tungsten are attached on the camera and have 0.5-mm and 1.0-mm apertures. The basic performance of the detection system was evaluated with 99mTc and 201Tl solutions. Energy resolution, system spatial resolution and linearity of count rate were measured. Rat myocardial perfusion SPECT scans were sequentially performed following intravenous injection of 201TlCl. Projection data were reconstructed using a previously validated pinhole 3D-OSEM method.The energy resolution at 140 keV was 14.8% using a point source. The system spatial resolutions were 2.8-mm FWHM and 2.5-mm FWHM for 99mTc and 201Tl line sources, respectively, at 30-mm source distance (magnification factor of 1.3) using a 1.0-mm pinhole. The linearity between the activity and count rate was good up to 10 kcps. In a rat study, the left ventricular walls were clearly visible in all scans.We developed a compact SPECT system using compact gamma camera for small animals and evaluated basic physical performances. The present system may be of use for quantitation of biological functions such as myocardial blood flow in small animals.

Lutetium oxyorthosilicate (LSO) contains natural radioactivity that emits beta particles and three gamma photons simultaneously. These beta particles and gamma photons increase the single and random rates in a positron emission tomography (PET) system while a beta particle and gamma photon produced in the same decay of Lu-176 and detected by another detector can be beta-gamma coincidence true events. The purpose of this work is to measure the single, random, and true count rates due to the natural radioactivity in LSO and determine the optimum lower energy threshold level for an energy window in an LSO-based clinical PET.First, we measured the energy spectra of these beta particles and gamma photons in LSO using a single crystal to obtain the basic data. Then, we measured single, random, and true count rates of an LSO-based clinical PET from the natural radioactivity as a function of the lower energy threshold.In the PET, single and random count rates due to the natural background activity were gradually decreased as the lower energy threshold level increased. The true count rates due to the beta-gamma coincidence were more than 10 kcps below a lower energy threshold of 250 keV. However, these true count rates due to the natural radioactivity in LSO can be decreased to less than 1 kcps at a lower energy threshold level set at more than 350 keV.With these considerations, in an LSO-based clinical PET, a lower energy threshold level set at above 350 keV is recommended.

We recently developed an integrated positron emission tomography (PET)/magnetic resonance imaging (MRI) (iPET/MRI) scanner for small animals, which had relatively large field-of-view (FOV) covering up to the size of a rat body. The purpose of this study was to report results of simultaneous PET/MRI of a rat body using this scanner with some radiotracers.C-11-methionine (MET), F-18-fluorodeoxyglucose (FDG), or F-18-sodium fluoride (NaF) was injected as a radiotracer for PET portion in addition to gadolinium-ethoxybenzyl-diethylenetriamine penta-acetic acid, a hepatobiliary contrast agent, for MRI portion. Simultaneous PET/MRI was performed in normal rats. PET, MRI, and co-registered fusion images were evaluated regarding image quality and feasibility for rat imaging studies.MET uptake was clearly shown in the liver and pancreas, which was confirmed with magnetic resonance (MR) and fused PET/MR images. PET/MR images depicted intense FDG uptake in the brain, Harderian glands, and myocardium. NaF uptake was observed in all bones and joints within FOV, except in ribs, which was well recognized with the help of MR and fused PET/MR images.This study demonstrated that simultaneous PET/MRI with an integrated dual-modality molecular imaging scanner was a feasible technique for imaging studies targeting on a rat body. However, further developments including attenuation correction methods are required to use this technique routinely in rat imaging studies.

Although simultaneous measurements of PET and magnetic resonance imaging (MRI) can provide interesting results in molecular imaging research, most of the combined systems are huge and animal handling in the system is not easy. To minimize these problems, we developed a compact integrated PET/MRI (iPET/MRI) system for small animals. For the iPET/MRI system, a new MR-compatible PET and a permanent magnet open MRI were designed. In the MRI, a tunnel is opened at the yoke of the magnet. The position-sensitive photo-multiplier tubes (PSPMTs) of the MR-compatible PET are positioned at the back of the yoke where the magnetic field is sufficiently low. The scintillators for the PET system are positioned at the center of the MRI magnets, and the direction of the scintillation photons is changed by slanted light guides, and they are fed to the PSPMTs by 75 cm long optical fiber bundles. The PET detectors employed two types of LGSO crystals (1.9 mm × 2.2 mm × 6 mm and 7 mm) with different decay times (33 and 43 ns) for depth of interaction detection. Sixteen optical fiber-based block detectors are arranged in a 112 mm diameter ring. The transaxial field-of-view (FOV) of the PET system is ~80 mm, and the axial FOV is 21 mm which can be enlarged by the axial motion of the PET detector ring during MRI acquisition. The transaxial and axial resolutions at the center of the PET system was 2.9 and 2.4 mm FWHM, respectively. The absolute sensitivity was 1.5% at the center of the axial FOV. Phantom images revealed no artifact in either the PET or MRI images. We successfully obtained simultaneously measured small animal images using the iPET/MRI system. The open geometry of the developed iPET/MRI facilitates easy accessibility to the subject. The iPET/MRI system appears to be a promising tool for molecular imaging research.

When employing F-18-fluorodeoxyglucose (FDG)-guided surgery to detect positron accumulation in isolated small organs, sampling these organs from opposite directions is a useful way of determining a tumor's position, similar to sampling a small organ with tweezers. The coincidence method is suitable for this purpose because only the positrons between two detectors can be detected. For this purpose, we developed a tweezers-type coincidence imaging detector.The detector employs two depth-of-interaction (DOI) detectors positioned at the tip of the tweezers and images the positron distribution between them using the coincidence method. The DOI detector consists of a 4 x 3 Gd(2)SiO(5):Ce (GSO) array optically coupled to a one-dimensionally arranged quad-photomultiplier tube. These GSOs were arranged to form a DOI detector using the Anger principle. The useful field of view is 20 mm x 15 mm. With these configurations, we could resolve 4 x 3 GSO arrays on a position histogram.Because the imaging detectors were positioned at the tip of the tweezers, one could easily sample the target part manually from opposed sides. A real-time image in coincidence between these two DOI detectors could be obtained. The point spread functions were approximately 3-mm full width at half-maximum (FWHM) parallel to the tweezers and 4-mm FWHM perpendicular to them. The sensitivity was approximately 1% when the two imaging detectors were 10 mm apart.With these results, we conclude that the developed tweezers-type imaging detector has a potential to be a new instrument in nuclear medicine.

For radio-guided surgery on tumors using F-18-FDG, detection of annihilation gamma photons emanating from other parts of the body produces background radiation counts and limits its use in clinical situations. To overcome this limitation, we have developed an intra-operative positron probe with background-rejection capability. The positron probe uses a phoswich detector composed of a plastic scintillator and a bismuth germinate (BGO). A positron from a positron emitter such as F-18 is detected by the plastic scintillator and emits annihilation photons. The BGO detects one of the annihilation photons while a photo-multiplier tube (PMT) detects scintillation photons from both scintillators. The decay time differences of these two scintillators are used to distinguish whether the event is a true event where a positron and a following annihilation photon are detected simultaneously, or a background event. In this configuration, only positrons can be selectively detected, even in an environment of high background gamma photon flux. Spatial resolution was 11-mm full width at half maximum (FWHM) 5 mm from the detector surface. Measured sensitivity for the F-18 point source was 2.6 cps/kBq 5 mm from the detector surface. The background count rate was less than 0.5 cps for a 20-cm diameter cylindrical phantom containing 37 MBq of F-18 solution measured on the phantom surface, while the positron count rate was almost linear over a range of approximately 6 kcps. These results indicate that our developed intra-operative positron probe is valuable for radio-guided surgery on tumors using F-18-FDG in a high flux of background annihilation gamma photons.

In a positron-sensitive probe composed of a plastic scintillator and a bismuth germanate (BGO), scattered annihilation photons in the plastic scintillator become background counts. Although these scattered annihilation photons can be rejected by higher threshold level settings for the scintillation pulse of the plastic scintillator and for that of the BGO, the system sensitivity is reduced. We have theoretically and experimentally optimized the threshold levels for both the plastic scintillator and the BGO. After calculating the energy loss in the plastic scintillator and the BGO for the scattered annihilation photons, we measured the background counts of a positron-sensitive probe by changing these threshold levels. Results revealed that one optimum threshold setting of the positron-sensitive probe was 0.3 of the peak level of the pulse for the plastic scintillator and 0.7 of that for the BGO. With these threshold levels, the background counts could be decreased to less than 0.2% of the true positron counts.

Objective Cerenkov-light imaging provides inherently high resolution because the light is emitted near the positron radionuclide. However, the magnitude for the high spatial resolution of Cerenkov-light imaging is unclear. Its potential molecular imaging applications also remain unclear. We developed an ultrahigh-resolution Cerenkov-light imaging system, measured its spatial resolution, and explored its applications to molecular imaging research. Methods Our Cerenkov-light imaging system consists of a high-sensitivity charged-coupled device camera (Hamamatsu Photonics ORCA2-ER) and a bright lens (Xenon 0.95/25). An extension ring was inserted between them to magnify the subject. A ~100-μm-diameter 22Na point source was made and imaged by the system. For applications of Cerenkov-light imaging, we conducted 18F-FDG administered in vivo, ex vivo whole brain, and sliced brain imaging of rats. Results We obtained spatial resolution of ~220 μm for a 22Na point source with our developed imaging system. The 18F-FDG rat head images showed high light intensity in the eyes for the Cerenkov-light images, although there was no accumulation in these parts in the PET images. The sliced rat brain showed much higher spatial resolution for the Cerenkov-light images compared with CdWO4 scintillator-based autoradiography, although some contrast decrease was observed for them. Conclusion Even though the Cerenkov-light images showed ultrahigh resolution of ~220 μm, their distribution and contrast were sometimes different from the actual positron accumulation in the subjects. Care must be taken when evaluating positron distribution from Cerenkov-light images. However, the ultrahigh resolution of Cerenkov-light imaging will be useful for transparent subjects including phantom studies.