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We present the performance of a digital phoswich positron emission tomography (PET) detector, composed by layers of pixilated scintillator arrays, read out by solid state light detectors and an application specific integrated circuit (ASIC). We investigated the use of integrated charge from the scintillation pulses along with time-over-threshold (ToT) to determine the layer of interaction (DOI) in the scintillator. Simulations were performed to assess the effectiveness of the ToT measurements for separating the scintillator events and identifying cross-layer-crystal-scatter (CLCS) events. These simulations indicate that ToT and charge integration from such a detector provide sufficient information to determine the layer of interaction. To demonstrate this in practice, we used a pair of prototype LYSO/BGO detectors. One detector consisted of a 19 × 19 array of 7 mm long LYSO crystals (1.36 mm pitch) coupled to a 16 × 16 array of 8 mm long BGO crystals (1.63 mm pitch). The other detector was similar except the LYSO crystal pitch was 1.63 mm. These detectors were coupled to an 8 × 8 multi-pixel photon counter mounted on a PETsys TOFPET2 ASIC. This high performance ASIC provided digital readout of the integrated charge and ToT from these detectors. We present a method to separate the events from the two scintillator layers using the ToT, and also investigate the performance of this detector. All the crystals within the proposed detector were clearly resolved, and the peak to valley ratio was 11.8 ± 4.0 and 10.1 ± 2.9 for the LYSO and BGO flood images. The measured energy resolution was 9.9% ± 1.3% and 28.5% ± 5.0% respectively for the LYSO and BGO crystals in the phoswich layers. The timing resolution between the LYSO–LYSO, LYSO–BGO and BGO–BGO coincidences was 468 ps, 1.33 ns and 2.14 ns respectively. Results show ToT can be used to identify the crystal layer where events occurred and also identify and reject the majority of CLCS events between layers.

IntroductionRadio thin layer chromatography (radio-TLC) is commonly used to analyze purity of radiopharmaceuticals or to determine the reaction conversion when optimizing radiosynthesis processes. In applications where there are few radioactive species, radio-TLC is preferred over radio-high-performance liquid chromatography due to its simplicity and relatively quick analysis time. However, with current radio-TLC methods, it remains cumbersome to analyze a large number of samples during reaction optimization. In a couple of studies, Cerenkov luminescence imaging (CLI) has been used for reading radio-TLC plates spotted with a variety of isotopes. We show that this approach can be extended to develop a high-throughput approach for radio-TLC analysis of many samples.MethodsThe high-throughput radio-TLC analysis was carried out by performing parallel development of multiple radioactive samples spotted on a single TLC plate, followed by simultaneous readout of the separated samples using Cerenkov imaging. Using custom-written MATLAB software, images were processed and regions of interest (ROIs) were drawn to enclose the radioactive regions/spots. For each sample, the proportion of integrated signal in each ROI was computed. Various crude samples of [18F]fallypride, [18F]FET and [177Lu]Lu-PSMA-617 were prepared for demonstration of this new method.ResultsBenefiting from a parallel developing process and high resolution of CLI-based readout, total analysis time for eight [18F]fallypride samples was 7.5min (2.5min for parallel developing, 5min for parallel readout), which was significantly shorter than the 48min needed using conventional approaches (24min for sequential developing, 24min for sequential readout on a radio-TLC scanner). The greater separation resolution of CLI enabled the discovery of a low-abundance side product from a crude [18F]FET sample that was not discernable using the radio-TLC scanner. Using the CLI-based readout method, we also observed that high labeling efficiency (99%) of [177Lu]Lu-PSMA-617 can be achieved in just 10min, rather than the typical 30min timeframe used.ConclusionsCerenkov imaging in combination with parallel developing of multiple samples on a single TLC plate proved to be a practical method for rapid, high-throughput radio-TLC analysis.

Pulse pileup events degrade the signal-to-noise ratio (SNR) of nuclear medicine data. When such events occur in multiplexed detectors, they cause spatial misposition, energy spectrum distortion and degraded timing resolution, which leads to image artifacts. Pulse pileup is pronounced in PETbox4, a bench top PET scanner dedicated to high sensitivity and high resolution imaging of mice. In that system, the combination of high absolute sensitivity, long scintillator decay time (BGO) and highly multiplexed electronics lead to a significant fraction of pulse pileup, reached at lower total activity than for comparable instruments. In this manuscript, a new pulse pileup rejection method named position shift rejection (PSR) is introduced. The performance of PSR is compared with a conventional leading edge rejection (LER) method and with no pileup rejection implemented (NoPR). A comprehensive digital pulse library was developed for objective evaluation and optimization of the PSR and LER, in which pulse waveforms were directly recorded from real measurements exactly representing the signals to be processed. Physical measurements including singles event acquisition, peak system sensitivity and NEMA NU-4 image quality phantom were also performed in the PETbox4 system to validate and compare the different pulse pile-up rejection methods. The evaluation of both physical measurements and model pulse trains demonstrated that the new PSR performs more accurate pileup event identification and avoids erroneous rejection of valid events. For the PETbox4 system, this improvement leads to a significant recovery of sensitivity at low count rates, amounting to about 1/4th of the expected true coincidence events, compared to the LER method. Furthermore, with the implementation of PSR, optimal image quality can be achieved near the peak noise equivalent count rate (NECR).

HiPET is a recently developed prototype preclinical PET scanner dedicated to high sensitivity and high resolution molecular imaging. The HiPET system employs a phoswich depth of interaction (DOI) detector design, which also allows identification of the large majority of the cross layer crystal scatter (CLCS) events. This work evaluates its performance characteristics following the National Electrical Manufacturers Association (NEMA) NU4-2008 protocol. The HiPET consists of twenty flat panel type detectors arranged in two rings. The inner diameter is 160 mm and the axial field of view (FOV) is 104 mm. Each detector is comprised of two layers of phoswich scintillator crystal arrays, a tapered, pixelated glass lightguide and a multi anode photomultiplier tube (MAPMT). The front (gamma ray entrance) layer is a 48 × 48 pixelated cerium doped lutetium yttrium orthosilicate (LYSO) scintillator array with individual crystals measuring 1.01 × 1.01 × 6.1 mm. The back (towards the PMT) layer is a 32 × 32 pixelated bismuth germanate (BGO) scintillator array with individual crystals measuring 1.55 × 1.55 × 8.9 mm. For energy windows of 250-650 keV and 350-650 keV, the peak absolute sensitivity at the center of the FOV was 13.5% and 10.4% including CLCS events, and 11.8% and 8.9% excluding CLCS events, respectively. The average detector energy resolution derived by averaging the individual crystal spectra was 11.7% ± 1.4% for LYSO and 17.0% ± 1.4% for BGO. The 3D ordered-subsets expectation maximization (OSEM) reconstructed image of a point source in air, ranged from 0.73 mm to 1.19 mm, with an average value of 0.93 ± 0.09 mm at all measured locations. The peak noise equivalent count rate (NECR) and scatter fraction were 179 kcps at 12.4 MBq and 6.9% for the mouse-sized phantom, and 63 kcps at 11.3 MBq and 18.3% for the rat-sized phantom. For the NEMA image quality phantom, the uniformity was 5.8%, and the spillover ratios measured in the water- and air-filled cold region chambers were 0.047 and 0.044, respectively. The recovery coefficients (RC) ranged from 0.31 to 0.92. These results and in vivo evaluation demonstrate that the HiPET can achieve high quality molecular imaging for biomedical applications.

G8 is a benchtop integrated PET/CT scanner dedicated to high-sensitivity and high-resolution imaging of mice. This work characterizes its National Electrical Manufacturers Association NU 4-2008 performance where applicable and also assesses the basic imaging performance of the CT subsystem. Methods: The PET subsystem in G8 consists of 4 flat-panel detectors arranged in a boxlike geometry. Each panel consists of 2 modules of a 26 × 26 pixelated bismuth germanate scintillator array with individual crystals measuring 1.75 × 1.75 × 7.2 mm. The crystal arrays are coupled to multichannel photomultiplier tubes via a tapered, pixelated glass lightguide. A cone-beam CT scanner consisting of a MicroFocus x-ray source and a complementary metal oxide semiconductor detector provides anatomic information. Sensitivity, spatial resolution, energy resolution, scatter fraction, count-rate performance, and the capability of performing phantom and mouse imaging were evaluated for the PET subsystem. Noise, dose level, contrast, and resolution were evaluated for the CT subsystem. Results: With an energy window of 350–650 keV, the peak sensitivity was 9.0% near the center of the field of view. The crystal energy resolution ranged from 15.0% to 69.6% in full width at half maximum (FWHM), with a mean of 19.3% ± 3.7%. The average intrinsic spatial resolution was 1.30 and 1.38 mm FWHM in the transverse and axial directions, respectively. The maximum-likelihood expectation maximization reconstructed image of a point source in air averaged 0.81 ± 0.11 mm FWHM. The peak noise-equivalent count rate for the mouse-sized phantom was 44 kcps for a total activity of 2.9 MBq (78 μCi), and the scatter fraction was 11%. For the CT subsystem, the value of the modulation transfer function at 10% was 2.05 cycles/mm. Conclusion: The overall performance demonstrates that the G8 can produce high-quality images for molecular imaging–based biomedical research.

Small animal positron emission tomography (PET) systems are often designed by employing close geometry configurations. Due to the different characteristics caused by geometrical factors, these tomographs require data acquisition protocols that differ from those optimized for conventional large diameter ring systems. In this work we optimized the energy window for data acquisitions with PETbox4, a 50 mm detector separation (box-like geometry) pre-clinical PET scanner, using the Geant4 Application for Tomographic Emission (GATE). The fractions of different types of events were estimated using a voxelized phantom including a mouse as well as its supporting chamber, mimicking a realistic mouse imaging environment. Separate code was developed to extract additional information about the gamma interactions for more accurate event type classification. Three types of detector backscatter events were identified in addition to the trues, phantom scatters and randoms. The energy window was optimized based on the noise equivalent count rate (NECR) and scatter fraction (SF) with lower-level discriminators (LLD) corresponding to energies from 150 keV to 450 keV. The results were validated based on the calculated image uniformity, spillover ratio (SOR) and recovery coefficient (RC) from physical measurements using the National Electrical Manufacturers Association (NEMA) NU-4 image quality phantom. These results indicate that when PETbox4 is operated with a more narrow energy window (350-650 keV), detector backscatter rejection is unnecessary. For the NEMA NU-4 image quality phantom, the SOR for the water chamber decreases by about 45% from 15.1% to 8.3%, and the SOR for the air chamber decreases by 31% from 12.0% to 8.3% at the LLDs of 150 and 350 keV, without obvious change in uniformity, further supporting the simulation based optimization. The optimization described in this work is not limited to PETbox4, but also applicable or helpful to other small inner diameter geometry scanners.

A beta camera has been developed that allows planar imaging of the spatial and temporal distribution of beta particles using a 14 × 14 mm^2 position sensitive avalanche photodiode (PSAPD). This camera system, which we call Betabox, can be directly coupled to microfluidic chips designed for cell incubation or other biological applications. Betabox allows for imaging the cellular uptake of molecular imaging probes labeled with charged particle emitters such as ^(18)F inside these chips. In this work, we investigate the quantitative imaging capabilities of Betabox for ^(18)F beta particles, in terms of background rate, efficiency, spatial resolution, and count rate. Measurements of background and spatial resolution are considered both at room temperature (21 °C ± 1 °C) and at an elevated operating temperature (37 °C ± 1 °C), as is often required for biological assays. The background rate measured with a 4 keV energy cutoff is below 2 cph mm^(−2) at both 21 and 37 °C. The absolute efficiency of Betabox for the detection of ^(18)F positron sources in contact with a PSAPD with the surface passivated from ambient light and damage is 46% ± 1%. The lower detection limit is estimated using the Rose Criterion to be 0.2 cps mm^(−2) for 1 min acquisitions and a 62 × 62 µm^2 pixel size. The upper detection limit is approximately 21 000 cps. The spatial resolution at both 21 and 37 °C ranges from 0.4 mm FWHM at the center of the field of view (FOV), and degrades to 1 mm at a distance of 5 mm away from center yielding a useful FOV of approximately 10 × 10 mm^2. We also investigate the effects on spatial resolution and sensitivity that result from the use of a polymer based microfluidic chip. For these studies we place varying layers of low-density polyethylene (LDPE) between the detector and the source and find that the spatial resolution degrades by ~180 µm for every 100 µm of LDPE film. Sensitivity is reduced by half with the inclusion of ~200 µm of additional LDPE film. Lastly, we demonstrate the practical utilization of Betabox, with an imaging test of its linearity, when coupled to a polydimethylsiloxane microfluidic chip designed for cell based assays.

Semiconductor radiation detectors offer direct absorption of a high-energy photon by the detector material, and promise the lowest energy resolution for x-ray and gamma-ray spectroscopy. The enhanced energy resolution for x-rays or gamma-rays could be very attractive for monitoring hazardous industrial waste, sensing radiological weapons or nuclear threats, and medical imaging. High stopping power (photoelectric absorption efficiency, ∝ Z4–5) semiconductor radiation detectors that operate at room temperature with a performance approaching the energy resolution of cooled high purity germanium detectors (HPGe) is desirable, but have yet been realized. Large bandgap materials with lower noise tend to have low Z and thus poor efficiency for stopping gamma radiation. On the other hand, high Z materials, which are efficient at stopping gamma radiation, present low bandgaps resulting in large dark currents and require cooling to achieve good energy resolution. CdTe is by far the most promising material that has both large bandgap and high Z, but it also suffers from hole tailing effect for photon energy >100 keV due to the poor hole transport [1, 2].

PETbox is a low-cost benchtop PET scanner dedicated to high throughput preclinical imaging that is currently under development at our institute. This paper presents the design and characterization of the detectors that are used in the PETbox system. In this work, bismuth germanate scintillator was used for the detector, taking advantage of its high stopping power, high photoelectric event fraction, lack of intrinsic background radiation and low cost. The detector block was segmented into a pixelated array consisting of 20 × 44 elements, with a crystal pitch of 2.2 mm and a crystal cross section of 2 mm × 2 mm. The effective area of the array was 44 mm × 96.8 mm. The array was coupled to two Hamamatsu H8500 position sensitive photomultiplier tubes, forming a flat-panel type detector head with a sensitive area large enough to cover the whole body of a typical laboratory mouse. Two such detector heads were constructed and their performance was characterized. For one detector head, the energy resolution ranged from 16.1% to 38.5% full width at half maximum (FWHM), with a mean of 20.1%; for the other detector head, the energy resolution ranged from 15.5% to 42.7% FWHM, with a mean of 19.6 %. The intrinsic spatial resolution was measured to range from 1.55 mm to 2.39 mm FWHM along the detector short axis and from 1.48 mm to 2.33 mm FWHM along the detector long axis, with an average of 1.78 mm. Coincidence timing resolution for the detector pair was measured to be 4.1 ns FWHM. These measurement results show that the detectors are suitable for our specific application.

It has been observed that microfluidic chips used for synthesizing (18)F-labeled compounds demonstrate visible light emission without nearby scintillators or fluorescent materials. The origin of the light was investigated and found to be consistent with the emission characteristics from Cerenkov radiation. Since (18)F decays through the emission of high-energy positrons, the energy threshold for beta particles, i.e. electrons or positrons, to generate Cerenkov radiation was calculated for water and polydimethylsiloxane (PDMS), the most commonly used polymer-based material for microfluidic chips. Beta particles emitted from (18)F have a continuous energy spectrum, with a maximum energy that exceeds this energy threshold for both water and PDMS. In addition, the spectral characteristics of the emitted light from (18)F in distilled water were also measured, yielding a broad distribution from 300 nm to 700 nm, with higher intensity at shorter wavelengths. A photograph of the (18)F solution showed a bluish-white light emitted from the solution, further suggesting Cerenkov radiation. In this study, the feasibility of using this Cerenkov light emission as a method for quantitative measurements of the radioactivity within the microfluidic chip in situ was evaluated. A detector previously developed for imaging microfluidic platforms was used. The detector consisted of a charge-coupled device (CCD) optically coupled to a lens. The system spatial resolution, minimum detectable activity and dynamic range were evaluated. In addition, the calibration of a Cerenkov signal versus activity concentration in the microfluidic chip was determined. This novel method of Cerenkov radiation measurements will provide researchers with a simple yet robust quantitative imaging tool for microfluidic applications utilizing beta particles.

A phoswich depth of interaction (DOI) detector composed by two layers of scintillator arrays (LYSO and BGO) has been proposed. To use this design for building practical systems, a tapered multiple-element glass lightguide is necessary to couple the scintillator arrays to a PSPMT. The complete individual detector module offers an overall dimension of 52×52 mm2 that matches the external dimensions of the PSPMT package, allowing continuous positioning of the scintillator arrays for creating flat panel detectors without introducing gaps between detector modules. In this work, we define the elements adjacent to the edge in an array as “edge+” elements. Both of the layers of the phoswich detector and the lightguide have such edge+ elements. The lightguide was modified according to the photodetector geometry, aiming at improving edge and corner crystal identification and energy resolution: In an assembled detector, the optimized edge+ lightguide elements superimpose on top of both edge and edge+ anodes of the PSPMT. In that case, the edge+ lightguide elements distribute light onto both the edge and edge+ anodes, while edge lightguide elements primarily distribute light onto the edge anodes. This different light sharing scheme improves the capability of identifying edge+ crystals from edge crystals. In addition, the edge lightguide elements were enlarged to improve the light collection and corresponding energy resolution for the edge crystals. Flood images were acquired and energy resolution was calculated for individual crystals. All the edge and edge+ crystals were clearly resolved. Energy resolution was improved from 24.4 ± 4.4% to 20.0 ± 1.7% for edge crystals. The new proposed lightguide successfully improved edge crystal identification and energy resolution in the proposed pixelated scintillator detector.

A new phoswich detector is being developed at the Crump Institute, aiming to provide improvements in sensitivity, and spatial resolution for PET. The detector configuration is comprised of two layers of pixelated scintillator crystal arrays, a glass lightguide and a light detector. The annihilation photon entrance (top) layer is a $48\times48$ array of $1.01 \times 1.01 \times7 ~\hbox{mm}^{3}$ LYSO crystals. The bottom layer is a $32 \times 32$ array of $1.55 \times1.55 \times 9~\hbox{mm}^{3}$ BGO crystals. A tapered, multiple-element glass lightguide is used to couple the exit end of the BGO crystal array ( $52 \times 52~\hbox{mm}^{2}$ ) to the photosensitive area of the Position Sensitive Photomultiplier Tube ( $46 \times 46~\hbox{mm}^{2}$ ), allowing the creation of flat panel detectors without gaps between the detector modules. Both simulations and measurements were performed to evaluate the characteristics and benefits of the proposed design. The GATE Monte Carlo simulation indicated that the total fraction of the cross layer crystal scatter (CLCS) events in singles detection mode for this detector geometry is 13.2%. The large majority of these CLCS events (10.1% out of 13.2%) deposit most of their energy in a scintillator layer other than the layer of first interaction. Identification of those CLCS events for rejection or correction may lead to improvements in data quality and imaging performance. Physical measurements with the prototype detector showed that the LYSO, BGO and CLCS events were successfully identified using the delayed charge integration (DCI) technique, with more than 95% of the LYSO and BGO crystal elements clearly resolved. The measured peak-to-valley ratios (PVR) in the flood histograms were 3.5 for LYSO and 2.0 for BGO. For LYSO, the energy resolution ranged from 9.7% to 37.0% full width at half maximum (FWHM), with a mean of $13.4 \pm 4.8\%$ . For BGO the energy resolution ranged from 16.0% to 33.9% FWHM, with a mean of $18.6\pm 3.2\%$ . In conclusion, these results demonstrate that the proposed detector is feasible and can potentially lead to a high spatial resolution, high sensitivity and DOI PET system.

The Optical Positron Emission Tomography (OPET) system will provide the capabilities of bioluminescence and PET modalities in one combined imaging system. Bioluminescence light emitted in vivo has a spectrum ranging from 550–700 nm. The optical photon signals detected are approximately two orders lower in amplitude and one order shorter in duration compared to scintillation light pulses from a 511 keV γ-event. In this study the performance characteristics of the OPET detectors for bioluminescence imaging were evaluated. The detection uniformity was measured for wavelengths between 400–700 nm and normalization maps were used to correct the acquired images. The spatial resolution along the transaxial and axial directions was 2.30 and 2.38 mm FWHM, respectively. The average dark counts for the detector was 76±30 cps at an ambient temperature of 20°.

The X-ray spectrum from a small-animal computed tomography (CT) device directly influences the dose delivered to the animal as well as the image quality. For any given imaging task, the spectrum needs to be optimized. The aim of this study is to present a model of a small animal microCT X-ray source and to calculate practical spectra for different anode and filter combinations by using the Monte Carlo method. The GATE simulation package was used for this work. The X-ray source of the MicroCAT II scanner was modeled as a tungsten enclosure containing the X-ray tube with dielectric oil and a thin carbon exit window. The anode was a tungsten or molybdenum cylinder block tilted with respect to the electron-beam axis. Beam filtration was achieved through a disk of variable thickness and composition positioned at the exit window. Photon detection was performed with a thick (1 cm) cylindrical tungsten plate (20 cm diameter). Simulated spectra were compared with spectra measured with a CdTe detector for three beams: 50, 70 and 80 kVp with 2 mm of aluminum filtration. There was a reasonable agreement between measurements and calculations. Identified features of the measured spectra were: K absorption edges of Cd and Te (from the detector itself) and characteristic X-rays from Te (Kα and Kβ) and possibly Sb (Kα). Discrepancies between simulations and measurements can be attributed to elements that have been left out of simulations, imperfect detector alignment, and uncertainty in elemental composition of simulated materials. The Monte Carlo method is a convenient and fast tool allowing the calculation of spectra possessing the essential features of real spectra. These and other spectra will be used in future work to study the influence of X-ray spectrum on dose and image quality for various X-ray CT system configurations.

The design of a dual modality imaging system for small animal optical and positron emission tomography imaging (OPET) is underway. Its detector must be capable of imaging high energy γ-rays from PET while also resolving optical wavelength photons from bioluminescence. GSO, high purity GSO, BGO, LSO, LYSO, and LaBr scintillators were investigated for their use in the OPET detector. Of specific interest were scintillators with low afterglow, since afterglow photons in the decay of the larger γ-ray events are indistinguishable from the photons generated by bioluminescence. Samples from these crystals were coupled to a photomultiplier tube (PMT) and produced scintillation light from γ-ray events originating from a positron source. The PMT output was directed to a special signal processing circuit that allowed measurement of single photons at different times in the decay of the scintillation. GSO and BGO exhibited optimal performance for use in the OPET system due to their low afterglow. LSO, LYSO, and LaBr were determined unsuitable for use with the current OPET design due to their significant afterglow components. The effect of the afterglow of GSO on the detection of the bioluminescence signal-to-noise ratio (SNR) was evaluated for the OPET system.

Inevitable discrepancies between the mouse tissue optical properties assumed by an experimenter and the actual physiological values may affect the tomographic localization of bioluminescent sources. In a previous work, the simplifying assumption of optically homogeneous tissues led to inaccurate localization of deep sources. Improved results may be obtained if a mouse anatomical map is provided by a high-resolution imaging modality and optical properties are assigned to segmented tissues. In this work, the feasibility of this approach was explored by simulating the effect of different magnitude optical property errors on the image formation process of a combined optical–PET system. Some comparisons were made with corresponding simulations using higher spatial resolution data that are typically attainable by CCD cameras. In addition, simulation results provided insights on some of the experimental conditions that could lead to poor localization of bioluminescent sources. They also provided a rough guide on how accurately tissue optical properties need to be known in order to achieve correct localization of point sources with increasing tissue depth under low background noise conditions.

The feasibility and limits in performing tomographic bioluminescence imaging with a combined optical-PET (OPET) system were explored by simulating its image formation process. A micro-MRI based virtual mouse phantom was assigned appropriate tissue optical properties to each of its segmented internal organs at wavelengths spanning the emission spectrum of the firefly luciferase at 37 degrees C. The TOAST finite-element code was employed to simulate the diffuse transport of photons emitted from bioluminescence sources in the mouse. OPET measurements were simulated for single-point, two-point and distributed bioluminescence sources located in different organs such as the liver, the kidneys and the gut. An expectation maximization code was employed to recover the intensity and location of these simulated sources. It was found that spectrally resolved measurements were necessary in order to perform tomographic bioluminescence imaging. The true location of emission sources could be recovered if the mouse background optical properties were known a priori. The assumption of a homogeneous optical property background proved inadequate for describing photon transport in optically heterogeneous tissues and led to inaccurate source localization in the reconstructed images. The simulation results pointed out specific methodological challenges that need to be addressed before a practical implementation of OPET-based bioluminescence tomography is achieved.

A combined optical positron emission tomography (OPET) system is capable of both optical and PET imaging in the same setting, and it can provide information/interpretation not possible in single-mode imaging. The scintillator array here serves the dual function of coupling the optical signal from bioluminescence/fluorescence to the photodetector and also of channeling optical scintillations from the gamma rays. We report simulation results of the PET part of OPET using GATE, a Geant4 simulation package. The purpose of this investigation is the definition of the geometric parameters of the OPET tomograph. OPET is composed of six detector blocks arranged in a hexagonal ring-shaped pattern with an inner radius of 15.6 mm. Each detector consists of a two-dimensional array of 8 /spl times/ 8 scintillator crystals each measuring 2 /spl times/ 2 /spl times/ 10 mm/sup 3/. Monte Carlo simulations were performed using the GATE software to measure absolute sensitivity, depth of interaction, and spatial resolution for two ring configurations, with and without gantry rotations, two crystal materials, and several crystal lengths. Images were reconstructed with filtered backprojection after angular interleaving and transverse one-dimensional interpolation of the sinogram. We report absolute sensitivities nearly seven times that of the prototype microPET at the center of field of view and 2.0 mm tangential and 2.3 mm radial resolutions with gantry rotations up to an 8.0 mm radial offset. These performance parameters indicate that the imaging spatial resolution and sensitivity of the OPET system will be suitable for high-resolution and high-sensitivity small-animal PET imaging.

MicroPET II is a second-generation animal PET scanner designed for high-resolution imaging of small laboratory rodents. The system consists of 90 scintillation detector modules arranged in three contiguous axial rings with a ring diameter of 16.0 cm and an axial length of 4.9 cm. Each detector module consists of a 14 x 14 array of lutetium oxyorthosilicate (LSO) crystals coupled to a multi-channel photomultiplier tube (MC-PMT) through a coherent optical fibre bundle. Each LSO crystal element measures 0.975 mm x 0.975 mm in cross section by 12.5 mm in length. A barium sulphate reflector material was used between LSO elements leading to a detector pitch of 1.15 mm in both axial and transverse directions. Fused optical fibre bundles were made from 90 microm diameter glass fibres with a numerical aperture of 0.56. Interstitial extramural absorber was added between the fibres to reduce optical cross talk. A charge-division readout circuit was implemented on printed circuit boards to decode the 196 crystals in each array from the outputs of the 64 anode signals of the MC-PMT. Electronics from Concorde Microsystems Inc. (Knoxville, TN) were used for signal amplification, digitization, event qualification, coincidence processing and data capture. Coincidence data were passed to a host PC that recorded events in list mode. Following acquisition, data were sorted into sinograms and reconstructed using Fourier rebinning and filtered hackprojection algorithms. Basic evaluation of the system has been completed. The absolute sensitivity of the microPET II scanner was 2.26% at the centre of the field of view (CFOV) for an energy window of 250-750 keV and a timing window of 10 ns. The intrinsic spatial resolution of the detectors in the system averaged 1.21 mm full width at half maximum (FWHM) when measured with a 22Na point source 0.5 mm in diameter. Reconstructed image resolution ranged from 0.83 mm FWHM at the CFOV to 1.47 mm FWHM in the radial direction, 1.17 mm FWHM in the tangential direction and 1.42 mm FWHM in the axial direction at 1 cm offset from the CFOV. These values represent highly significant improvements over our earlier microPET scanner (approximately fourfold sensitivity increase and 25-35% improvement in linear spatial resolution under equivalent operating conditions) and are expected to be further improved when the system is fully optimized.

We report our investigation of a positron emission tomography (PET) detector with 1 m spatial resolution. The prototype detector consists of a 9×9 array of 1×1×10 mm 3 lutetium oxyorthosilicate (LSO) scintillator crystals coupled to Hamamatsu R5900-M64 or R5900-C12 position sensitive PMT by either optical fibers or an optical fiber bundle. With a 511 eV gamma source, the intrinsic spatial resolution of this detector was measured to be 0.92 mm. All crystals were well resolved in the flood source histogram. The measured energy and coincidence timing resolutions were around 26% and 4 ns, respectively, demonstrating that sufficient light can be extracted from these small crystals for PET applications.

We are currently developing a small animal positron emission tomography (PET) scanner with a design goal of 1 microlitre (1 mm3) image resolution. The detectors consist of a 12 x 12 array of 1 x 1 x 10 mm lutetium oxyorthosilicate (LSO) scintillator crystals coupled to a 64-channel photomultiplier tube (PMT) via 5 cm long optical fibre bundles. The optical fibre connection allows a high detector packing fraction despite the dead space surrounding the active region of the PMT. Optical fibre bundles made from different types of glass were tested for light transmission, and also their effects on crystal identification and energy resolution, and compared to direct coupling of the LSO arrays to the PMTs. We also investigated the effects of extramural absorber (EMA) in the fibre bundles. Based on these results, fibre bundles manufactured from F2 glass were selected. We built three pairs of prototype detectors (directly coupled LSO array, fibre bundle without EMA and fibre bundle with EMA) and measured flood histograms, energy resolution, intrinsic spatial resolution and timing resolution. The results demonstrated an intrinsic spatial resolution (FWHM) of 1.12 mm (directly coupled), 1.23 mm (fibre bundle without EMA coupling) and 1.27 mm (fibre bundle with EMA coupling) using an approximately 500 microm diameter Na-22 point source. Using a 330 microm outer diameter steel needle line source filled with F-18, spatial resolution for the detector with the EMA optical fibre bundle improved to 1.05 mm. The respective timing and energy FWHM values were 1.96 ns, 21% (directly coupled), 2.20 ns, 23% (fibre bundle without EMA) and 2.99 ns, 30% (fibre bundle with EMA). The peak-to-valley ratio in the flood histograms was better with EMA (5:1) compared to the optical fibre bundle without EMA (2.5:1), due to the decreased optical cross-talk. In comparison to the detectors used in our current generation microPET scanner, these detectors substantially improve on the spatial resolution, preserve the timing resolution and provide adequate energy resolution for a modern high-resolution animal PET tomograph.

Surface treatment of scintillators is critical for light collection from narrow rectangular crystals. The authors investigated chemical polishing which is less labor intensive and expensive than hand and machine polishing. They used phosphoric acid to chemically polish LSO crystals at 110/spl deg/C, 150/spl deg/C and 190/spl deg/C, and compared these with unpolished and mechanically polished crystals. Groups of five 2/spl times/2/spl times/10 mm crystals were etched for different times at each temperature. Weight loss, light output and energy resolution were measured as a function of treatment time and temperature. The authors found that chemical polishing can increase light output by 250% relative to unpolished crystals and by 16% relative to mechanically polished crystals. The energy resolution was relatively independent of surface treatment, with values of between 14 and 16%. The rate of loss of LSO was 0.02%/min at 110/spl deg/C, 0.1%/min at 150/spl deg/C and 0.36%/min at 190/spl deg/C. Maximum light output occurred when 5-10% of the LSO was removed. The crystals were also imaged using scanning electron microscopy and the surface roughness quantitatively assessed by using a profilometer. These measurements helped to clarify the effect of acid polishing on the surface of the scintillator. In summary, chemical polishing appears to be a convenient and effective method for improving light output from small LSO crystals.

A method is described to improve the spatial sampling of microPET, a high resolution PET scanner designed for imaging small laboratory animals. The high intrinsic resolution of the microPET detector (1.58 mm FWHM), in combination with the stationary ring geometry of the tomograph, generate an imaging system which is inherently spatially undersampled. As a result the imaging resolution measured with a three-dimensional (3-D) filtered backprojection (FBP) algorithm for a point source at the center of the field of view (CFOV) is only 1.8 mm FWHM and has large fluctuations at positions near the CFOV. A small wobble motion was introduced via a circular motion of the scanner bed in the transverse plane, with a wobble radius of 300 /spl mu/m. The data was acquired with a step-and-shoot method by dividing the wobble circle into a number of equidistantly sampled intervals. The separate sinograms were interpolated to a finely resampled sinogram, which was reconstructed with the 3-D filtered backprojection algorithm. The resulting images demonstrated full recovery of the intrinsic detector resolution and elimination of the local nonuniformities of the point spread function (PSF) at the CFOV, with three wobble samples. The resulting average resolution improvement fur the central 5 cm of the FOV was approximately 13% in the radial and 19% in the tangential direction, with an associated 50% penalty in the reconstructed image noise.

A Bayesian method is described for reconstruction of high-resolution 3D images from the microPET small-animal scanner. Resolution recovery is achieved by explicitly modelling the depth dependent geometric sensitivity for each voxel in combination with an accurate detector response model that includes factors due to photon pair non-collinearity and inter-crystal scatter and penetration. To reduce storage and computational costs we use a factored matrix in which the detector response is modelled using a sinogram blurring kernel. Maximum a posteriori (MAP) images are reconstructed using this model in combination with a Poisson likelihood function and a Gibbs prior on the image. Reconstructions obtained from point source data using the accurate system model demonstrate a potential for near-isotropic FWHM resolution of approximately 1.2 mm at the center of the field of view compared with approximately 2 mm when using an analytic 3D reprojection (3DRP) method with a ramp filter. These results also show the ability of the accurate system model to compensate for resolution loss due to crystal penetration producing nearly constant radial FWHM resolution of 1 mm out to a 4 mm radius. Studies with a point source in a uniform cylinder indicate that as the resolution of the image is reduced to control noise propagation the resolution obtained using the accurate system model is superior to that obtained using 3DRP at matched background noise levels. Additional studies using pie phantoms with hot and cold cylinders of diameter 1-2.5 mm and 18FDG animal studies appear to confirm this observation.

MicroPET is a high resolution positron emission tomography (PET) scanner designed for imaging small laboratory animals. It consists of a ring of 30 position-sensitive scintillation detectors, each with an 8/spl times/8 array of small lutetium oxyorthosilicate (LSO) crystals coupled via optical fibers to a multi-channel photomultiplier tube. The detectors have an intrinsic resolution averaging 1.68 mm, an energy resolution between 15 and 25% and 2.4 ns timing resolution at 511 keV. The detector ring diameter of microPET is 17.2 cm with an imaging field of view of 112 mm transaxially by 18 mm axially. The scanner has no septa and operates exclusively in 3D mode. Reconstructed image resolution 1 cm from the center of the scanner is 2.0 mm and virtually isotropic, yielding a volume resolution of 8 mm/sup 3/. For comparison, the volume resolution of state-of-the-art clinical PET systems is in the range of 50-75 mm/sup 3/. Initial images of phantoms have been acquired and are reported. A computer controlled bed is under construction and will incorporate a small wobble motion to improve spatial sampling. This is projected to further enhance spatial resolution. MicroPET is the first PET scanner to incorporate the new scintillator LSO and to our knowledge is the highest resolution multi-ring PET scanner currently in existence.

To date most examinations of silicon photomultipliers (SiPM) for use in positron emission tomography have involved studies of arrays or single crystals of Lutetium based scintillators. In this work, we describe measurements bearing on the suitability of SiPMs for use with arrays of pixelated BGO scintillators in preclinical tomographs. Most preclinical scanners rely on photomultiplier tubes (PMTs) coupled to scintillators through a light guide to provide spreading of light over a number of sensor pixels. Because SiPMs have similar gains to PMTs and similar or higher photon detection probability, they could prove an attractive, compact alternative to PMTs for use in preclinical scanners. However, because of the relatively high noise of SiPMs compared to PMTs, combined with the low light output and long decay time of BGO, whether SiPMs can provide an adequate alternative for PMTs in such scanners is an open question. We have examined three commercially available SiPM arrays coupled to a BGO scintillator array. These SiPM sensors are: the Philips digital photon counting SiPM (PDPC), the SensL Matrix9 SiPM array and the Hamamatsu S11828 16 channel Multipixel Photon Counter (MPPC). In particular, we examined the triggering and noise of these devices with respect to the energy and spatial resolution obtained with the BGO array. While we were able to identify the crystals in the BGO array with each sensor, the high noise of these devices when operating at room temperature has a large impact on the energy and spatial resolution. Our results indicate better spatial and energy resolution were obtained with the digital SiPM, which allows suppression of noisy microcells.

PETbox4 is a new, fully tomographic bench top PET scanner dedicated to high sensitivity and high resolution imaging of mice. This manuscript characterizes the performance of the prototype system using the National Electrical Manufacturers Association (NEMA) NU 4–2008 standards, including studies of sensitivity, spatial resolution, energy resolution, scatter fraction, count-rate performance and image quality. The PETbox4 performance is also compared with the performance of PETbox, a previous generation limited angle tomography system. PETbox4 consists of four opposing flat-panel type detectors arranged in a box like geometry. Each panel is made by a 24 × 50 pixelated array of 1.82 × 1.82 × 7mm bismuth germanate (BGO) scintillation crystals with a crystal pitch of 1.90mm. Each of these scintillation arrays is coupled to two Hamamatsu H8500 photomultiplier tubes via a glass light guide. Volumetric images for a 45 × 45 × 95 mm field of view (FOV) are reconstructed with a maximum likelihood expectation maximization (ML-EM) algorithm incorporating a system model based on a parameterized detector response. With an energy window of 150–650 keV, the peak absolute sensitivity is approximately 18% at the center of FOV. The measured crystal energy resolution ranges from 13.5% to 48.3% full width at half maximum (FWHM), with a mean of 18.0%. The intrinsic detector spatial resolution is 1.5mm FWHM in both transverse and axial directions. The reconstructed image spatial resolution for different locations in the FOV ranges from 1.32mm to 1.93 mm, with an average of 1.46 mm. The peak noise equivalent count rate for the mouse-sized phantom is 35 kcps for a total activity of 1.5 MBq (40 µCi) and the scatter fraction is 28%. The standard deviation in the uniform region of the image quality phantom is 5.7%. The recovery coefficients range from 0.10 to 0.93. In comparison to the first generation two panel PETbox system, PETbox4 achieves substantial improvements on sensitivity and spatial resolution. The overall performance demonstrates that the PETbox4 scanner is suitable for producing high quality images for molecular imaging based biomedical research.

Measured attenuation correction in PET is routinely performed using transmission scans. Acquisition time and noise considerations necessitate low-pass filtering of the transmission data before generating the attenuation correction matrix. This smoothing operation reduces noise propagation from transmission into emission data but also introduces image artifacts that are mostly pronounced around areas of strongly varying attenuation coefficients. The source of these artifacts, which lies in the mismatch of the spatial resolutions of emission and transmission data, was investigated in this study. The effect of different transmission and emission sinogram smoothing protocols on the emission images was also investigated. A method is proposed that addresses the problem, in conjunction with the filtering step during reconstruction. Instead of the standard in-plane low-pass filtering of the emission data during reconstruction, emission and transmission sinograms can be volumetrically filtered to the desired image resolution prior to attenuation correction while reconstruction is performed with no smoothing (Ramp filter). This operation reduces or eliminates resolution mismatch and consequent image artifacts, especially in cardiac studies. The proposed method improves the accuracy of the activity distribution in emission images with minimal computational and signal-to-noise ratio (SNR) cost.

We extract the carrier mobility-lifetime products for epitaxially grown GaSb and demonstrate the spectral response to gamma rays of a GaSb p–i–n photodiode with a 2-µm-thick absorption region. Under exposure from 55Fe and 241Am radioactive sources at 140 K, the photodiode exhibits full width at half maximum energy resolutions of 1.238 ± 0.028 and 1.789 ± 0.057 keV at 5.89 and 59.5 keV, respectively. We observe good linearity of the GaSb photodiode across a range of photon energies. The electronic noise and charge trapping noise are measured and shown to be the main components limiting the measured energy resolutions.

PETBox4 is a preclinical system dedicated to imaging mice. This system is composed by four detector panels, each made by a 24 × 50 array of 1.825 × 1.825 × 7 mm BGO crystals. The face to face crystal separation of the detectors is 5 cm, generating a 4.5 × 4.5 × 9.4 cm field of view (FOV). The result is a tomograph with a large detection solid angle, which in combination with a wide open energy window achieves high peak detection efficiency (~17%). Despite the small size of the typical imaged subjects, these design features also increase the fraction of accepted crystal and object scattered events, which reduce the overall image signal to noise ratio. In this work, we investigated the system acquisition configuration settings necessary to optimize the NEC (Noise equivalent Counts) and image quality. A Monte Carlo simulation software package was used (GATE) to investigate the different types of events detected as a function of energy window settings and multiple event acceptance policy. This was done for a realistic distribution of activity and attenuation coefficients in the PETBox4 FOV, based on emission data from an in-vivo preclinical PET image from an average sized mouse (30g). Based on the simulations, the NEC rate was optimized for a dual energy window that accepts both low (50-175 keY) and high (410-650 keY) energy events. This indicates that low energy events that are composed mostly from single crystal scatter contribute useful image information, while events in the middle of the energy spectrum (175keV-410keV), tend to include large fractions of crystal backscatter as well as object scatter and do not contribute significantly in data signal to noise ratio. As a result of the same simulations, a new policy for the acceptance of multiple events was introduced, implementing a "KiIlAIl" multiple policy, further improving the NEe. Overall, these two optimization parameters improved the NEC rate by 56%, providing a significant advantage in image signal to noise ratio.

OPET is an imaging system that detects both highenergy γ-rays and optical wavelength photons. This system is capable of non-invasively and repeatedly imaging small animal models in-vivo for the presence of γ-rays from PET tracers and optical signals from bioluminescence. OPET consists of six detector modules arranged to form a ring with a bore diameter of 3.5cm, where each module is made up of a 64-channel multichannel photomultiplier tube (PMT) coupled to an 8×8 BGO crystal array. The crystal lengths and widths are 2.15mm x 2.15mm respectively. The crystal heights vary along the face of the PMT such that the array presents a smooth curving profile. The front surface of the crystals is left open to allow optical wavelength photons to enter the detector module. A charge-division readout circuit is implemented to decode the signals from the 64 outputs of each PMT into four signals from which position information is obtained. These signals are amplified, shaped, and digitized, and finally processed using a field programmable gate array (FPGA). Flood images are used to obtain position look-up tables as well as pulse height spectra for each crystal in both PET and Optical mode operation. Basic system performance parameters have been measured for both operation modes of OPET including: sensitivity and count-rate performance. In PET mode the system has a sensitivity of 3.5% at the center of the field of view and an energy resolution around 36%. The sensitivity for optical wavelength photons was measured for each detector using a calibrated light source and varied from 0.6% to 2.0%. The maximum count rate in Optical mode is 106 cps per detector. The first dual-mode data acquisition of OPET was performed using a two chamber phantom. In one chamber was a set of six 0.8mm diameter rods filled with FDG, while the other chamber was filled with an enzyme and substrate used in bioluminescence studies of small animals. Images from both modes of operation are be presented.

Purpose PETbox is a low cost bench top preclinical PET scanner dedicated to pharmacokinetic and pharmacodynamic mouse studies. A prototype system was developed at our institute, and this manuscript characterizes the performance of the prototype system. Procedures The PETbox detector consists of a 20 × 44 bismuth germanate crystal array with a thickness of 5 mm and cross-section size of 2.05 × 2.05 mm. Two such detectors are placed facing each other at a spacing of 5 cm, forming a dual-head geometry optimized for imaging mice. The detectors are kept stationary during the scan, making PETbox a limited angle tomography system. 3D images are reconstructed using a maximum likelihood and expectation maximization (ML–EM) method. The performance of the prototype system was characterized based on a modified set of the NEMA NU 4-2008 standards. Results In-plane image spatial resolution was measured to be an average of 1.53 mm full width at half maximum for coronal images and 2.65 mm for the anterior–posterior direction. The volumetric reconstructed resolution was below 8 mm3 at most locations in the field of view (FOV). The sensitivity, scatter fraction, and noise equivalent count rate (NECR) were measured for different energy windows. With an energy window of 150 - 650 keV and a timing window of 20 ns optimized for mouse imaging, the peak absolute sensitivity was 3.99% at the center of FOV and a peak NECR of 20 kcps was achieved for a total activity of 3.2 MBq (86.8 μCi). Phantom and in vivo imaging studies were performed and demonstrated the utility of the system at low activity levels. The quantitation capabilities of the system were also characterized showing that despite the limited angle tomography, reasonably good quantification accuracy was achieved over a large dynamic range of activity levels. Conclusions The presented results demonstrate the potential of this new tomograph for small animal imaging.

The increased emphasis on mammography acceptance testing and quality control make the star pattern test an essential component of the procedure for diagnosing poor resolution due to large focal spots. In this paper procedures for calculating magnification and the correct angle of the star pattern to use are described. The preferred placement of the star pattern in the field and proper exposure technique is also discussed.

Bioluminescence tomography (BLT) is an inverse source problem that localizes and quantifies bioluminescent probe distribution in 3-D. The generic BLT model is ill-posed, leading to nonunique solutions and aberrant reconstruction in the presence of measurement noise and optical parameter mismatches. In this paper, we introduce the knowledge of the number of bioluminescence sources to stabilize the BLT problem. Based on this regularized BLT model, we develop a differential evolution-based reconstruction algorithm to determine the source locations and strengths accurately and reliably. Then, we evaluate this novel approach in numerical, phantom, and mouse studies.

Bioluminescence imaging has been extensively applied to in vivo small animal imaging. Quantitative three-dimensional bioluminescent source information obtained by using bioluminescence tomography can directly and much more accurately reflect biological changes as opposed to planar bioluminescence imaging. Preliminary simulated and experimental reconstruction results demonstrate the feasibility and promise of bioluminescence tomography. However, the use of multiple approximations, particularly the diffusion approximation theory, affects the quality of in vivo small animal-based image reconstructions. In the development of new reconstruction algorithms, high-order approximation models of the radiative transfer equation and spectrally resolved data introduce new challenges to the reconstruction algorithm and speed. In this paper, an SP(3)-based (the third-order simplified spherical harmonics approximation) spectrally resolved reconstruction algorithm is proposed. The simple linear relationship between the unknown source distribution and the spectrally resolved data is established in this algorithm. A parallel version of this algorithm is realized, making BLT reconstruction feasible for the whole body of small animals especially for fine spatial domain discretization. In simulation validations, the proposed algorithm shows improved reconstruction quality compared with diffusion approximation-based methods when high absorption, superficial sources and detection modes are considered. In addition, comparisons between fine and coarse mesh-based BLT reconstructions show the effects of numerical errors in reconstruction image quality. Finally, BLT reconstructions using in vivo mouse experiments further demonstrate the potential and effectiveness of the SP(3)-based reconstruction algorithm.

PETbox, an integrated low-cost benchtop preclinical PET scanner dedicated to high throughput quantitative mouse studies is currently under development at the Crump Institute for Molecular Imaging, UCLA. The system employs two opposing flat panel BGO detector blocks on a static gantry and produces a limited number of projection angles around a central antero-posterior (AP) view. Iterative reconstruction based on a Maximum Likelihood and Expectation Maximization (ML-EM) algorithm was developed with incorporation of a probability matrix, which takes into account the detection probability and inter-crystal scattering. Point sources at different positions were simulated with the GATE software package and showed a peak sensitivity of 3.87% at a 150-650 keV energy window. When applying a sensitivity correction map, the image values of the reconstructed point sources at different positions are within a 5.3% standard deviation, indicating good quantification accuracy. The spatial resolution of point sources is largely uniform along the coronal direction across the field of view (FOV) with worse resolution along the direction orthogonal to the detector blocks, which is mainly due to the reduced spatial sampling in that direction. A voxelized digital mouse (Moby phantom) was also simulated and then iteratively reconstructed. Attenuation and non-uniform sensitivity corrections were applied. Quantification accuracy was evaluated with the reconstructed Moby data and two analytically simulated data sets. To a first approximation, reasonably good quantification accuracy was obtained for most important organs with the PETbox reconstructed image. This evaluation indicates that the integrated detector system design, reconstruction algorithm and attenuation correction work well, which is very important for quantitative pharmacokinetic and pharmacodynamic studies. In-vivo mouse studies were also acquired and reconstructed. The obtained images proved the capability of PETbox as a dedicated mouse scanner for both dynamic and static acquisitions.

The Inveon dedicated PET (DPET) scanner is the latest generation of preclinical PET systems devoted to high-resolution and high-sensitivity murine model imaging. In this study, we report on its performance based on the National Electrical Manufacturers Association (NEMA) NU-4 standards.The Inveon DPET consists of 64 lutetium oxyorthosilicate block detectors arranged in 4 contiguous rings, with a 16.1-cm ring diameter and a 12.7-cm axial length. Each detector block consists of a 20 x 20 lutetium oxyorthosilicate crystal array of 1.51 x 1.51 x 10.0 mm elements. The scintillation light is transmitted to position-sensitive photomultiplier tubes via optical light guides. Energy resolution, spatial resolution, sensitivity, scatter fraction, and counting-rate performance were evaluated. The NEMA NU-4 image-quality phantom and a healthy mouse injected with (18)F-FDG and (18)F(-) were scanned to evaluate the imaging capability of the Inveon DPET.The energy resolution at 511 keV was 14.6% on average for the entire system. In-plane radial and tangential resolutions reconstructed with Fourier rebinning and filtered backprojection algorithms were below 1.8-mm full width at half maximum (FWHM) at the center of the field of view. The radial and tangential resolution remained under 2.0 mm, and the axial resolution remained under 2.5-mm FWHM within the central 4-cm diameter of the field of view. The absolute sensitivity of the system was 9.3% for an energy window of 250-625 keV and a timing window of 3.432 ns. At a 350- to 625-keV energy window and a 3.432-ns timing window, the peak noise equivalent counting rate was 1,670 kcps at 130 MBq for the mouse-sized phantom and 590 kcps at 110 MBq for the rat-sized phantom. The scatter fractions at the same acquisition settings were 7.8% and 17.2% for the mouse- and rat-sized phantoms, respectively. The mouse image-quality phantom results demonstrate that for typical mouse acquisitions, the image quality correlates well with the measured performance parameters in terms of image uniformity, recovery coefficients, attenuation, and scatter corrections.The Inveon system, compared with previous generations of preclinical PET systems from the same manufacturer, shows significantly improved energy resolution, sensitivity, axial coverage, and counting-rate capabilities. The performance of the Inveon is suitable for successful murine model imaging experiments.

Small animal PET imaging imposes high performance requirements on image resolution and system sensitivity. Scanners with larger solid angle, achieved by using smaller crystal ring diameter and longer axial field-of-view (FOV), have higher absolute sensitivity. The Inveon dedicated PET (DPET) system, the latest generation of commercial tomographs from Siemens Preclinical Solutions, Inc., is such a high sensitivity scanner. Its crystal ring diameter and axial extent is 16.1 cm and 12.7 cm respectively, which give a solid angle coverage of 62%. However, this geometry also accentuates inter-crystal penetration, especially in the axial direction and causes axial blurring. Axial and radial blurring recovery is crucial for high resolution small animal PET imaging. System response modeling in combination with iterative reconstruction algorithms like Maximum a posteriori reconstruction (MAP) can be used to recover the resolution loss. Blurring in both radial and axial directions were simulated in GATE with a planar strip of 18F source placed inside the Inveon scanner. Detector response for each possible line of response (LOR) was calculated and can be incorporated into iterative reconstruction algorithms. As the entrance ring difference (δ) increased, the recorded coincidences tended to shift to a larger ring difference. As the entrance radial offset (u) increased, the recorded coincidences tended to shift to a smaller radial offset. The blurring effect got larger as δ and/or u increased. A real 18F plane source printed on carbon paper was also imaged and compared with the simulation data as a validation. The coincidence counts recorded for each ring difference showed a very good agreement between simulation and experiment except for very large oblique angles. This discrepancy should be improved with the inclusion of the scanner end shields in future simulations.

The development of a prototype dual-modality optical and PET (OPET) small animal imaging tomograph is underway in the Crump Institute for Molecular Imaging at the University of California Los Angeles. OPET consists of a single ring of six detector modules with a diameter of 3.5 cm. Each detector has an 8times8 array of optically isolated BGO scintillators which are coupled to multichannel photomultiplier tubes and open on the front end. The system operates in either PET or optical mode and reconstructs the data sets as 3D tomograms. The detectors are capable of detecting both annihilation events (511 keV) from PET tracers as well as single photon events (SPEs) (2-3 eV) from bioluminescence. Detector channels are readout using a custom multiplex readout scheme and then filtered in analog circuitry using either a gamma-ray or SPE specific filter. Shaped pulses are sent to a digital signal processing (DSP) unit for event processing. The DSP unit has 100 MHz analog-to-digital converters on the front-end which send digitized samples to field programmable gate arrays which are programmed via user configurable algorithms to process PET coincidence events or bioluminescence SPEs. Information determined using DSP includes: event timing, energy determination-discrimination, position determination-lookup, and coincidence processing. Coincidence or SPE events are recorded to an external disk and minimal post processing is required prior to image reconstruction. Initial imaging results from a phantom filled with 18FDG solution and an optical pattern placed on the front end of a detector module in the vicinity of a SPE source are shown.

This work proposes a novel method for quantitative imaging of radioactivity on microfluidic chips by using visible light emission from Cerenkov radiation. Cerenkov radiation is generated when charged particles travel through an optically transparent material with a velocity greater than that of light in that material. It has been observed at UCLA that microfluidic chips used for 18F-related radio-synthesis studies have shown unidentified visible light emissions. In this study, the origin of the light was investigated and its feasibility as a quantitative imaging source was tested.

Dual-energy x-ray computed tomography (DECT) has proved its potential in eliminating beam-hardening artifacts from reconstructed images. The method typically involves two successive scans with different x-ray tube voltage settings. This work proposes a new approach to dual-energy through x-ray beam filtration that requires only one scan and a single tube voltage setting.

A BGO based small animal PET scanner dedicated for imaging small rodent was simulated by GATE. The virtual PET scanner had the same ring diameter, axial field of view (FOV) and crystal arrangement as the LSO based Siemens Inveon PET system, but was simulated with varied crystal lengths. The simulated system sensitivity was 11.6%, 19.3% and 25.5% for 10, 15 and 20 mm BGO at an energy window of 250-750 keV. The spatial resolution was measured at radial offsets of 0, 15 and 28 mm from the center of the FOV for the three crystal thicknesses. The FWHM in the radial and tangential directions was below 2.5 mm and 1.8 mm respectively for all three crystal thicknesses, up to a 30 mm diameter FOV. Scatter fraction and count rate performance were measured using a line source inserted in a water cylinder 70 mm long and 25 mm diameter for the 20 mm BGO system. The maximum NECR was 0.99 Mcps at 24 MBq and the phantom scatter fraction was 4.5% with an energy window of 250-750 keV and a timing window of 12 ns. The BGO based PET scanner was compared with the Inveon and the microPET Focus 220 systems. With the same crystal thickness, the BGO scanner had higher system sensitivity than the LSO and further improvement in sensitivity can be achieved by using thicker crystals without sacrificing much spatial resolution. Both radial and tangential resolutions were comparable to the LSO based systems. At the evaluated energy window of 250-750 keV, the phantom scatter fraction was similar to the Inveon system, while the crystal scatter fraction was about 10% lower. The maximum NECR was lower than Inveon and was achieved at a lower activity level. Simulation of the BGO PET scanner proved the design concept of a high sensitivity small animal PET scanner, with comparable spatial resolution, similar phantom scatter fraction, and acceptable count rate performance.

Microfluidic chips are an emerging technology that facilitates the study of molecular processes in nano-liter levels in a finely controlled manner. A prototype imaging system capable of detecting and quantifying very small amounts of beta particles in a microfluidic chip, utilizing a scintillator, an optical lens, and a charge coupled device (CCD), was developed and proof of concept was previously demonstrated [1]. In this study, we optimized the system design parameters, such as CCD binning and types of scintillators, and evaluated the system performance for spatial resolution and minimum detectable activity. Pixel binning of the CCD during readout process improved the signal-to-noise ratio with no spatial resolution degradation in beta particle imaging of 18F, up to a binning of 3times3 in our study, which was equivalent to 44 mum times 44 mum pixels in an object plane with a magnification of 0.5. The full width at half maximum (FWHM) of line sources with a finite width of 115 mum were measured to be 493 mum with a plastic scintillator and 289 mum with a CsI(Tl) scintillator. The minimum detectable activities were measured to be 360 pCi/mm2 with a plastic scintillator and 40 pCi/mm2 with a CsI(Tl) scintillator for a 5-min acquisition.

The development of a prototype dual-modality optical and PET (OPET) small animal imaging tomograph is underway in the Crump Institute for Molecular Imaging at the University of California Los Angeles. OPET consists of a single ring of six detector modules with a diameter of 3.5 cm. Each detector has an 8 times 8 array of optically isolated BGO scintillators which are coupled to multichannel photomultiplier tubes and open on the front end. The system operates in either PET or optical mode and reconstructs the data sets as 3D tomograms. The detectors are capable of detecting both annihilation events (511 keV) from PET tracers as well as single photon events (SPEs) (2-3 eV) from bioluminescence. Detector channels are readout using a custom multiplex readout scheme and then filtered in analog circuitry using either a gamma-ray or SPE specific filter. Shaped pulses are sent to a digital signal processing (DSP) unit for event processing. The DSP unit has 100 MHz analog-to-digital converters on the front-end which send digitized samples to field programmable gate arrays which are programmed via user configurable algorithms to process PET coincidence events or bioluminescence SPEs. Information determined using DSP includes: event timing, energy determination-discrimination, position determination-lookup, and coincidence processing. Coincidence or SPE events are recorded to an external disk and minimal post processing is required prior to image reconstruction. Initial results from a two-detector version of OPET are shown for a PET phantom image reconstruction as well as for an optical mode acquisition. Projection images of an optical pattern placed on the front end of the detector module in the vicinity of a SPE source and positron emitting source are shown. Current work includes assembly of the full ring system and tomographic reconstruction of an optical source.

Novel molecular imaging applications increasingly involve studies where very low amount of activity is present. This operation point can be challenging in terms of image quality especially when the intrinsic detector activity from scintillators such as LSO is considered. LSO crystals contain 176Lu which emits beta- particles followed by gamma photons, resulting in the detection of true and random events. This background activity has been shown to contribute a significant percentage of the total detected true events, when a very weak activity distribution is imaged. This can affect the weakest signal that can be detected by an LSO PET scanner which determines its "detection limit" and is evaluated by the parameter of minimum detectable activity or MDA. A series of acquisitions was performed in order to study the effect of the energy window to the intrinsic true and randoms rate. The experiments were also simulated with GATE, and the results were validated by comparing them with the experiment. Four square regions each with a unique signal to background activity concentration ratio were used. The background activity level was held constant between the different regions, while the activity of the point sources varied in 4 selected levels. The signal to background ratio was calculated separately for each region. The energy spectrum of the intrinsic background activity and its contribution to the total energy spectrum both for singles and coincidences was estimated through the GATE simulation. We histogrammed both the measured and simulated data on various time frames, which we reconstructed using the filtered backprojection algorithm. Every image was quantified based on the Currie equation in order to associate an MDA value for each of the 4 point sources as a function of the frame length. In the case of the microPET Focus 220 a total amount of 4nCi/mm3 can be reliably detected for frame lengths longer than 5 min and at regions where the signal to background activity concentration ratio is higher than 4. In the case of higher contrast regions detection can be achieved even for frame lengths down to 1 min.

Purpose: To examine the effect of different scintillator surface treatments (BGO crystals) on the fraction of scintillation photons that exit the crystal and reach the photodetector (SiPM). Methods: Positron Emission Tomography is based on the detection of light that exits scintillator crystals, after annihilation photons deposit energy inside these crystals. A considerable fraction of the scintillation light gets trapped or absorbed after going through multiple internal reflections on the interfaces surrounding the crystals. BGO scintillator crystals generate considerably less scintillation light than crystals made of LSO and its variants. Therefore, it is crucial that the small amount of light produced by BGO exits towards the light detector. The surface treatment of scintillator crystals is among the factors affecting the ability of scintillation light to reach the detectors. In this study, we analyze the effect of different crystal surface treatments on the fraction of scintillation light that is detected by the solid state photodetector (SiPM), once energy is deposited inside a BGO crystal. Simulations were performed by a Monte Carlo based software named GATE, and validated by measurements from individual BGO crystals coupled to Philips digital-SiPM sensor (DPC-3200). Results: The results showed an increment in light collection of about 4 percent when only the exit face of the BGO crystal, is unpolished; compared to when all the faces are polished. However, leaving several faces unpolished caused a reduction of at least 10 percent of light output when the interaction occurs as far from the exit face of the crystal as possible compared to when it occurs very close to the exit face. Conclusion: This work demonstrates the advantages on light collection from leaving unpolished the exit face of BGO crystals. The configuration with best light output will be used to obtain flood images from BGO crystal arrays coupled to SiPM sensors.

The dual modality imaging system for small animal optical and positron emission tomography imaging (OPET) is currently being developed at the Crump Institute for Molecular Imaging. The OPET system is capable of detecting and simultaneously imaging both high-energy gamma-rays from PET tracers and single photons in the optical wavelength from bioluminescence. The PET performance characteristics for a two-detector prototype of the OPET system are described here. Each detector consists of an 8 /spl times/ 8 array of GSO crystals cut to different lengths providing a detector surface with a circular profile and a 3.4 cm ring diameter. The crystals were inserted into a laser-cut grid assembly constructed from 0.067 mm thick reflector film. Crystals were coupled to a 64-channel photomultiplier tube (PMT) and a charge-division readout circuit was implemented to decode the signals from the 64 outputs of the PMT into event location data. Flood images were taken to obtain look-up tables as well as energy spectra for each crystal. The timing resolution was determined to be 6.6 nsec. A 0.3 mm Na-22 point source was stepped across the field of view (FOV) and the intrinsic detector spatial resolution was 1.23 mm FWHM. The estimated peak absolute system sensitivity at the center of FOV for the two detector system was 0.33%. Data from a line phantom was acquired and reconstructed with an expectation maximization 3D reconstruction algorithm.

Due to its long radiological half life, 124I can be used for investigation of lengthy biological processes. However, it is not an ideal PET isotope because it has poor positron yield and half of the positrons are emitted in cascade with 603 keV gammas. This decay spectrum results in coincidences between gammas and annihilation photons (dirty coincidences) leading to both inaccurate quantitation and overall reduction in the image quality. It has been shown that noise equivalent counts (NEC) have a direct link to signal to noise ratio of the reconstructed images. The purpose of this work was to investigate the effect of varying the energy window of the digitizer on the NEC measured from the non-pure positron emitter 124I and hence, on the image quality. Simulations were performed using the GATE software for a 1.32 cm radius cylinder filled with 300 muCi of 124I placed at the center of field of view. We report that by windowing in on the upper energy window of the data acquisition from a symmetric 50% (255-766 keV) energy window, to an asymmetric 255-600 keV one, it is possible to preferentially reduce the dirty coincidences with minor sensitivity changes. These findings will be verified with experimental data, and will lead to optimizing the data acquisition chain for 124I imaging performance

A new device is being developed that will allow imaging of the distribution of charged particles in microfluidic circuits, using a scintillator, a lens and a charge coupled device (CCD). Microfluidic chips are an emerging technology that will facilitate the study of molecular processes in pico-liter levels in a finely controlled manner. The ability to quantify low amounts of radioactivity in a microfluidic chip will provide researchers with a platform to investigate molecular processes in a controlled in-vitro environment. The new detector system consists of a plastic scintillator, a C-mount compact lens, and a Sony ICX285AL interline-transfer CCD cooled with a Peltier device. A microfluidic chip filled with fluoro-deoxy-glucose (FDG) solution was coupled to a plastic scintillator and the set up was placed inside a light-tight box for imaging. Preliminary studies were performed to test the feasibility of using the scintillator-based CCD detector system for this application. The CCD performance parameters were characterized by the photon transfer curve. The camera gain constant and read-out noise were measured to be 0.746 e-/ADU (analog-to-digital unit) and 8 e-, respectively. The dark current was also investigated with different temperatures and binning factors. The spatial resolution was measured and line pairs of FDG in a microfluidic chip were discernable down to a 0.5 mm separation. The system was able to quantify the activity level reliably down to 5 nCi/mm2.