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High-resolution imaging of alpha particles is required in the detection of alpha radionuclides in cells or small organs for the development of radio-compounds for targeted alpha-particle therapy or other purposes. We developed an ultrahigh resolution, real time alpha-particle imaging system for observing the trajectories of alpha particles in a scintillator. The developed system is based on a magnifying unit and a cooled electron multiplying charge-coupled device (EM-CCD) camera, combined with a 100-µm-thick Ce-doped Gd3Al2Ga3O12 (GAGG) scintillator plate. Alpha particles from an Am-241 source were irradiated to the GAGG scintillator and imaged with the system. Using our system, we measured the trajectories of the alpha particles having different shapes in real time. In some of these measured trajectories, the line shapes of the alpha particles that flew in the GAGG scintillator were clearly observed. The lateral profiles of the alpha-particle trajectories were imaged with widths of ~ 2 µm. We conclude that the developed imaging system is promising for research on targeted alpha-particle therapy or other alpha particle detections that require high spatial resolution.

It is commonly thought that UV or visible-light luminescence imaging of biological subjects during X-ray irradiation at the energy below 120 keV is impossible because the secondary electrons produced in this energy range do not emit Cerenkov light. Contrary to this consensus, we found UV or visible-light luminescence imaging of the subjects were possible with X-ray irradiations of this energy range. We placed one of the biological subjects in a black box; visible-light luminescence images were measured with a high-sensitivity, cooled charge coupled device (CCD) camera during X-ray irradiation at energy below 120 keV. We also conducted the imaging of air without subjects during irradiation of the same X-ray. The biological subjects emitted visible-light luminescence, and the imaging was possible with the irradiation of the X-ray below 120 keV. The luminescence images were observed in only the X-ray irradiated areas. Also air luminescence images could be obtained and the intensity of the luminescence measured from the images was proportionally increased with the exposure dose. UV or visible-light luminescence imaging of biological subjects was possible during X-ray irradiations lower energy than the Cerenkov-light threshold. The phenomenon was different from general X-ray fluorescence because wavelength of the luminescence is UV or visible-light. The luminescence imaging method is promising for estimating the irradiated area with X-ray, which could be used for interventional radiology (IVR). Also air luminescence imaging would be applied to the exposure dose distribution measurements for X-ray of diagnostic X-ray systems.

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.

The silicon photomultiplier (Si-PM) is a promising photodetector for PET. However, it remains unclear whether Si-PM can be used for a depth-of-interaction (DOI) detector based on the decay time differences of the scintillator where pulse shape analysis is used. For clarification, we tested the Hamamatsu 4 × 4 Si-PM array (S11065-025P) combined with scintillators that used different decay times to develop DOI block detectors using the pulse shape analysis. First, Ce-doped Gd(2)SiO(5) (GSO) scintillators of 0.5 mol% Ce were arranged in a 4 × 4 matrix and were optically coupled to the center of each pixel of the Si-PM array for measurement of the energy resolution as well as its gain variations according to the temperature. Then two types of Ce-doped Lu(1.9)Gd(0.1)Si0(5) (LGSO) scintillators, 0.025 mol% Ce (decay time: ~31 ns) and 0.75 mol% Ce (decay time: ~46 ns), were optically coupled in the DOI direction, arranged in a 11 × 7 matrix, and optically coupled to a Si-PM array for testing of the possibility of a high-resolution DOI detector. The energy resolution of the Si-PM array-based GSO block detector was 18 ± 4.4 % FWHM for a Cs-137 gamma source (662 keV). Less than 1 mm crystals were clearly resolved in the position map of the LGSO DOI block detector. The peak-to-valley ratio (P/V) derived from the pulse shape spectra of the LGSO DOI block detector was 2.2. These results confirmed that Si-PM array-based DOI block detectors are promising for high-resolution small animal PET systems.

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.

Zero-field splittings and the kinetic parameters associated with the triplet spin sublevels of both the triplet ππ* and the triplet ligand-ligand charge-transfer states were measured by the optically-detected magnetic-resonance method for a series of complexes with a d10 closed-shell metal ion of the type, Zn(X-PhS)2(phen), where X = F5, 4-C1, H, 4-CH3, or 4-CH3O. Both the zero-field splittings and the kinetic decay parameters are satisfactorily interpreted in terms of mixing between the lowest phen-localized3ππ* and ligand-ligand charge-transfer states.

Zero-field splittings of the emitting levels of the title compounds were measured by ODMR. For complexes of the type Zn(X-PhS)2(phen) resonances of the3ππ* phosphorescing levels were observed for X=F5, 4-Cl, 4-CH3. For X=4-Cl, resonances were detected for the3LLCT emitting manifold also. For X=4-CH3O, resonances were detected only for the3LLCT emitting states. A model invoking mixing between the two types of states is proposed that satisfactorily accounts for the dependences of the properties of the emitting manifolds on the donating ability of the X substituent.

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.