Search

Groups such as E6 with extra SU(2)L singlet down quarks give rise to flavor-changing neutral currents (FCNC's) through the mixing of four or more down quarks. These currents may substantially contribute to the processes of Bd0-B̄d0 and Bs0-B̄s0 mixing, within the allowed range of nonzero values for the four quark mixing parameters. These would then introduce new angles and phases into the CP-violating B decay asymmetries. Other, negative FCNC experiments put limits on these angles. Using a maximum likelihood analysis with contributions from both the standard model and the FCNC's, we find large regions for new physics phases contributing to CP-violating B decay asymmetries. From the bound on the mixing elements, the new singlet down quark might have a mass above the 500 GeV range. © 1994 The American Physical Society.

The probability of possible values for the four angles of the CKM matrix are determined by finding the χ2 values for each choice of angles fitting to ten standard experiments. The χ2 isosurface plots are first displayed in the space of three of the CKM angles showing the dependence of the bounds of δ and greatly improving the lower bound on Vcb. The maximum likelihood χ2 contour plot in the ρ-η plane shows the probability of each location for the vertex of the unitarity triangle which controls CP-violating asymmetries in B0 decays. We find for mt=150 GeV that at the 2σ level the signal for asymmetries measuring sin(2β) is at least sin(2β)0.25. At the 1σ level the signal is at least sin(2β)0.40, with the preferred range of sin(2β)=0.6±0.2. © 1994 The American Physical Society.

Objective. Heterostructured scintillators offer a promising solution to balance the sensitivity and timing in TOF-PET detectors. These scintillators utilize alternating layers of materials with complementary properties to optimize performance. However, the layering compromises time resolution due to light transport issues. This study explores double-sided readout-enabling improved light collection and Depth-of-Interaction (DOI) information retrieval-to mitigate this effect and enhance the timing capabilities of heterostructures. Approach. The time resolution and DOI performances of 3 × 3 × 20 mm 3 BGO&EJ232 heterostructures were assessed in a single and double-sided readout (SSR and DSR, respectively) configuration using high-frequency electronics. Main results. Selective analysis of photopeak events yielded a DOI resolution of 6.4 ± 0.04 mm. Notably, the Coincidence Time Resolution (CTR) improved from 262 ± 8 ps (SSR) to 174 ± 6 ps (DSR) when measured in coincidence with a fast reference detector. Additionally, symmetrical configuration of two identical heterostructures in coincidence was tested, yielding in DSR a CTR of 254 ± 8 ps for all photopeak events and 107 ± 5 ps for the fastest events. Significance. By using high-frequency double-sided readout, we could measure DOI resolution and improve the time resolution of heterostructures of up to 40%. The DOI information resulted intrinsically captured in the average between the timestamps of the two SiPMs, without requiring any further correction., (Creative Commons Attribution license.)

We investigate a highly multiplexing readout for depth-of-interaction (DOI) and time-of-flight PET detector consisting of an N×N crystals whose light outputs at the front and back ends are detected by using silicon photomultipliers (SiPM). The front N×N SiPM array is read by using a stripline (SL) configured to support discrimination of the row position of the signal-producing crystal. The back N×N SiPM array is similarly read by an SL for column discrimination. Hence, the detector has only four outputs. We built 4×4 and 8×8 detector modules (DM) by using 3.0×3.0×20 mm 3 lutetium-yttrium oxyorthosilicates. The outputs were sampled and processed offline. For both DMs, crystal discrimination was successful. For the 4×4 DM, we obtained an average energy resolution (ER) of 14.1%, an average DOI resolution of 2.5 mm, a non DOI-corrected coincidence resolving time (CRT), measured in coincidence with a single-pixel reference detector, of about 495 ps. For the 8×8 DM, the average ER, average DOI resolution and average CRT were 16.4%, 2.9 mm, and 641 ps, respectively. We identified the intercrystal scattering as a probable cause for the CRT deterioration when the DM was increased from 4×4 to 8×8.

Objective. Standard signal processing approaches for scintillation detectors in positron emission tomography (PET) derive accurate estimates for 511 keV photon time of interaction and energy imparted to the detection media from aggregate characteristics of electronic pulse shapes. The ultimate realization of a scintillation detector for PET is one that provides a unique timestamp and position for each detected scintillation photon. Detectors with these capabilities enable advanced concepts for three-dimensional (3D) position and time of interaction estimation with methods that exploit the spatiotemporal arrival time kinetics of individual scintillation photons. Approach. In this work, we show that taking into consideration the temporal photon emission density of a scintillator, the channel density of an analog silicon photomultiplier (SiPM) array, and employing fast electronic readout with digital signal processing, a detector that counts and timestamps scintillation photons can be realized. To demonstrate this approach, a prototype detector was constructed, comprising multichannel electronic readout for a bismuth germanate (BGO) scintillator coupled to an SiPM array. Main Results. In proof-of-concept measurements with this detector, we were able to count and provide unique timestamps for 66% of all optical photons, where the remaining 34% (two-or-more-photon pulses) are also independently counted, but each photon bunch shares a common timestamp. We show this detector concept can implement 3D positioning of 511 keV photon interactions and thereby enable corrections for time of interaction estimators. The detector achieved 17.6% energy resolution at 511 keV and 237 ± 10 ps full-width-at-half-maximum coincidence time resolution (CTR) (fast spectral component) versus a reference detector. We outline the methodology, readout, and approach for achieving this detector capability in first-ever, proof-of-concept measurements for scintillation photon counting detector with analog silicon photomultipliers. Significance. The presented detector concept is a promising design for large area, high sensitivity TOF-PET detector modules that can implement advanced event positioning and time of interaction estimators, which could push state-of-the-art performance., (Creative Commons Attribution license.)

State-of-the-art (SoA) electronic readout for silicon photomultiplier (SiPM)-based scintillation detectors that demonstrate experimental limits in achievable coincidence time resolution (CTR) leverage low noise, high frequency signal processing to facilitate a single photon time response that is near the limit of the SiPMs architecture. This readout strategy can optimally exploit fast luminescence and prompt photon populations, and promising measurements show detector concepts employing this readout can greatly advance PET detector CTR, relative to SoA in clinical systems. However, the technique employs power hungry components which make the electronics chain impractical for channel-dense time-of-flight (TOF)-PET detectors. We have developed and tested a low noise and high frequency readout circuit which is performant at low power and consists of discrete elements with small footprints, making it feasible for integration into TOF-PET detector prototypes. A 3 × 3 mm 2 Broadcom SiPM with this readout chain exhibited sub-100 ps single photon time resolution at 10 mW of power consumption, with a relatively minor performance degradation to 120 ± 2 ps FWHM at 5 mW. CTR measurements with 3 × 3 × 20 mm 3 LYSO and fast LGSO scintillators demonstrated 127 ± 3 ps and 113 ± 2 ps FWHM at optimal power operation and 133 ± 2 ps and 121 ± 3 ps CTR at 5 mW. BGO crystals 3 × 3 × 20 mm 3 in size show 271 ± 5 ps FWHM CTR (1174 ± 14 ps full-width-at-tenth-maximum (FWTM)) at optimal power dissipation and 289 ± 8 ps (1296 ± 33 ps FWTM) at 5 mW. The compact and low power readout topology that achieves this performance thereby offers a platform to greatly advance PET system CTR and also opportunities to provide high performance TOF-PET at reduced material cost., (Creative Commons Attribution license.)

The Baram River is one of the largest rivers in Sarawak, where many large industries, such as plywood, sawmills, shipyards, interisland ports, and other wood-based industries are located along the river. Microplastic contamination has become a widespread and growing concern worldwide because of the small sizes of microplastics and their presence in seafood such as fish, squid, scallop, crabs, shrimp, and mussels. In this study, microplastics were found in all sampling stations. Out of the 4017 microplastics found in the water and sediment, microplastics fragment accounted for 67.8% of total microplastics, followed by fiber, film, pellet, and foam. Five microplastic polymer types were detected by ATR-FTIR, including polyethylene (PE), polyester (PET) fibers, silicon polymer, nitrile, and polystyrene (PS). The most common microplastics size range in Baram River was 0.3-1 mm, with blue as the highly abundant color., (Copyright © 2021 Elsevier Ltd. All rights reserved.)

Scintillation crystals of 20mm length or longer are needed for clinical time-of-flight positron emission tomography (TOF-PET) to ensure effective detection efficiency for gamma photons. However, the use of long crystals would deteriorate the key performance of TOF-PET detectors, time and spatial resolution, because of the variations in the travel times of the photons in crystals and the effects of parallax errors. In this work, we studied double-ended readout TOF-PET detectors based on coupling a long scintillation crystal to SiPMs at both ends for correcting the depth-dependent effects to improve the coincidence time resolution (CTR). In particular, we focused our attention to analyze timing performance using different correction methods, including trigger times of the individual photodetectors at both ends of the crystal, the simple average of the trigger times, and the weighted average based on the inverse variances of the depth-dependent corrected trigger times. For a 3 mm × 3 mm × 25mm unpolished lutetium fine silicate (LFS) crystal with double-ended readout and practical head-on irradiation, a CTR of 246ps FWHM can be achieved using depth-dependent timing-correction and weighted average time method compared to 280ps FWHM using the conventional simple average time method and 393ps FWHM using the conventional single-ended readout. The results show that the depth-dependent timing-correction and weighted average time method in double-ended readout can effectively correct for the trigger time variations in TOF-PET detector utilizing long unpolished crystals, resulting in an improvement in the CTR of as much as 37% compared to single-ended readout.

Link efficiency, data integrity, and continuity for high-throughput and real-time systems is crucial. Most of these applications require specialized hardware and operating systems as well as extensive tuning in order to achieve high efficiency. Here, we present an implementation of gigabit Ethernet data streaming which can achieve 99.26% link efficiency while maintaining no packet losses. The design and implementation are built on OpenPET, an opensource data acquisition platform for nuclear medical imaging, where (a) a crate hosting multiple OpenPET detector boards uses a User Datagram Protocol over Internet Protocol (UDP/IP) Ethernet soft-core, that is capable of understanding PAUSE frames, to stream data out to a computer workstation; (b) the receiving computer uses Netmap to allow the processing software (i.e., user space), which is written in Python, to directly receive and manage the network card's ring buffers, bypassing the operating system kernel's networking stack; and, ((c) a multi-threaded application using synchronized queues is implemented in the processing software (Python) to free up the ring buffers as quickly as possible while preserving data integrity and flow continuity.)

OpenPET is an open source, modular, extendible, and high-performance platform suitable for multi-channel data acquisition and analysis. Due to the flexibility of the hardware, firmware, and software architectures, the platform is capable of interfacing with a wide variety of detector modules not only in medical imaging but also in homeland security applications. Analog signals from radiation detectors share similar characteristics - a pulse whose area is proportional to the deposited energy and whose leading edge is used to extract a timing signal. As a result, a generic design method of the platform is adopted for the hardware, firmware, and software architectures and implementations. The analog front-end is hosted on a module called a Detector Board, where each board can filter, combine, timestamp, and process multiple channels independently. The processed data is formatted and sent through a backplane bus to a module called Support Board, where 1 Support Board can host up to eight Detector Board modules. The data in the Support Board, coming from 8 Detector Board modules, can be aggregated or correlated (if needed) depending on the algorithm implemented or runtime mode selected. It is then sent out to a computer workstation for further processing. The number of channels (detector modules), to be processed, mandates the overall OpenPET System Configuration, which is designed to handle up to 1,024 channels using 16-channel Detector Boards in the Standard System Configuration and 16,384 channels using 32-channel Detector Boards in the Large System Configuration.

We present Monte Carlo computational methods for estimating the coincidence resolving time (CRT) of scintillator detector pairs in positron emission tomography (PET) and present results for Lu2SiO5 : Ce (LSO), LaBr3 : Ce, and a hypothetical ultra-fast scintillator with a 1 ns decay time. The calculations were applied to both single-ended and double-ended photodetector readout with constant-fraction triggering. They explicitly include (1) the intrinsic scintillator properties (luminosity, rise time, decay time, and index of refraction), (2) the exponentially distributed depths of interaction, (3) the optical photon transport efficiency, delay, and time dispersion, (4) the photodetector properties (fill factor, quantum efficiency, transit time jitter, and single electron response), and (5) the determination of the constant fraction trigger level that minimizes the CRT. The calculations for single-ended readout include the delayed photons from the opposite reflective surface. The calculations for double-ended readout include (1) the simple average of the two photodetector trigger times, (2) more accurate estimators of the annihilation photon entrance time using the pulse height ratio to estimate the depth of interaction and correct for annihilation photon, optical photon, and trigger delays, and (3) the statistical lower bound for interactions at the center of the crystal. For time-of-flight (TOF) PET we combine stopping power and TOF information in a figure of merit equal to the sensitivity gain relative to whole-body non-TOF PET using LSO. For LSO crystals 3 mm  ×  3 mm  ×  30 mm, a decay time of 37 ns, a total photoelectron count of 4000, and a photodetector with 0.2 ns full-width at half-maximum (fwhm) timing jitter, single-ended readout has a CRT of 0.16 ns fwhm and double-ended readout has a CRT of 0.111 ns fwhm. For LaBr3 : Ce crystals 3 mm  ×  3 mm  ×  30 mm, a rise time of 0.2 ns, a decay time of 18 ns, and a total of 7600 photoelectrons the CRT numbers are 0.14 ns and 0.072 ns fwhm, respectively. For a hypothetical ultra-fast scintillator 3 mm  ×  3 mm  ×  30 mm, a decay time of 1 ns, and a total of 4000 photoelectrons, the CRT numbers are 0.070 and 0.020 ns fwhm, respectively. Over a range of examples, values for double-ended readout are about 10% larger than the statistical lower bound.

We present a 16-channel front-end readout board for the OpenPET electronics system. A major task in developing a nuclear medical imaging system, such as a positron emission computed tomograph (PET) or a single-photon emission computed tomograph (SPECT), is the electronics system. While there are a wide variety of detector and camera design concepts, the relatively simple nature of the acquired data allows for a common set of electronics requirements that can be met by a flexible, scalable, and high-performance OpenPET electronics system. The analog signals from the different types of detectors used in medical imaging share similar characteristics, which allows for a common analog signal processing. The OpenPET electronics processes the analog signals with Detector Boards. Here we report on the development of a 16-channel Detector Board. Each signal is digitized by a continuously sampled analog-to-digital converter (ADC), which is processed by a field programmable gate array (FPGA) to extract pulse height information. A leading edge discriminator creates a timing edge that is "time stamped" by a time-to-digital converter (TDC) implemented inside the FPGA. This digital information from each channel is sent to an FPGA that services 16 analog channels, and then information from multiple channels is processed by this FPGA to perform logic for crystal lookup, DOI calculation, calibration, etc.

We have constructed and characterized a time-of-flight Positron Emission Tomography (TOF PET) camera called the Tachyon. The Tachyon is a single-ring Lutetium Oxyorthosilicate (LSO) based camera designed to obtain significantly better timing resolution than the ~ 550 ps found in present commercial TOF cameras, in order to quantify the benefit of improved TOF resolution for clinically relevant tasks. The Tachyon's detector module is optimized for timing by coupling the 6.15 × 25 mm 2 side of 6.15 × 6.15 × 25 mm 3 LSO scintillator crystals onto a 1-inch diameter Hamamatsu R-9800 PMT with a super-bialkali photocathode. We characterized the camera according to the NEMA NU 2-2012 standard, measuring the energy resolution, timing resolution, spatial resolution, noise equivalent count rates and sensitivity. The Tachyon achieved a coincidence timing resolution of 314 ps +/- ps FWHM over all crystal-crystal combinations. Experiments were performed with the NEMA body phantom to assess the imaging performance improvement over non-TOF PET. The results show that at a matched contrast, incorporating 314 ps TOF reduces the standard deviation of the contrast by a factor of about 2.3.

We have developed a new time calibration method for the DRS4 waveform sampler that enables us to precisely measure the non-uniform sampling interval inherent in the switched-capacitor cells of the DRS4. The method uses the proportionality between the differential amplitude and sampling interval of adjacent switched-capacitor cells responding to a sawtooth-shape pulse. In the experiment, a sawtooth-shape pulse with a 40 ns period generated by a Tektronix AWG7102 is fed to a DRS4 evaluation board for calibrating the sampling intervals of all 1024 cells individually. The electronic time resolution of the DRS4 evaluation board with the new time calibration is measured to be ~2.4 ps RMS by using two simultaneous Gaussian pulses with 2.35 ns full-width at half-maximum and applying a Gaussian fit. The time resolution dependencies on the time difference with the new time calibration are measured and compared to results obtained by another method. The new method could be applicable for other switched-capacitor-array technology-based waveform samplers for precise time calibration.

In this paper we review the primary factors that affect the timing precision of a scintillation detector. Monte Carlo calculations were performed to explore the dependence of the timing precision on the number of photoelectrons, the scintillator decay and rise times, the depth of interaction uncertainty, the time dispersion of the optical photons (modeled as an exponential decay), the photodetector rise time and transit time jitter, the leading-edge trigger level, and electronic noise. The Monte Carlo code was used to estimate the practical limits on the timing precision for an energy deposition of 511 keV in 3 mm × 3 mm × 30 mm Lu2SiO5:Ce and LaBr3:Ce crystals. The calculated timing precisions are consistent with the best experimental literature values. We then calculated the timing precision for 820 cases that sampled scintillator rise times from 0 to 1.0 ns, photon dispersion times from 0 to 0.2 ns, photodetector time jitters from 0 to 0.5 ns fwhm, and A from 10 to 10,000 photoelectrons per ns decay time. Since the timing precision R was found to depend on A(-1/2) more than any other factor, we tabulated the parameter B, where R = BA(-1/2). An empirical analytical formula was found that fit the tabulated values of B with an rms deviation of 2.2% of the value of B. The theoretical lower bound of the timing precision was calculated for the example of 0.5 ns rise time, 0.1 ns photon dispersion, and 0.2 ns fwhm photodetector time jitter. The lower bound was at most 15% lower than leading-edge timing discrimination for A from 10 to 10,000 photoelectrons ns(-1). A timing precision of 8 ps fwhm should be possible for an energy deposition of 511 keV using currently available photodetectors if a theoretically possible scintillator were developed that could produce 10,000 photoelectrons ns(-1).

We characterize the nature of the time dispersion in scintillation detectors caused by path length differences of the scintillation photons as they travel from their generation point to the photodetector. Using Monte Carlo simulation, we find that the initial portion of the distribution (which is the only portion that affects the timing resolution) can usually be modeled by an exponential decay. The peak amplitude and decay time depend both on the geometry of the crystal, the position within the crystal that the scintillation light originates from, and the surface finish. In a rectangular parallelpiped LSO crystal with 3 mm × 3 mm cross section and polished surfaces, the decay time ranges from 10 ps (for interactions 1 mm from the photodetector) up to 80 ps (for interactions 50 mm from the photodetector). Over that same range of distances, the peak amplitude ranges from 100% (defined as the peak amplitude for interactions 1 mm from the photodetector) down to 4% for interactions 50 mm from the photodetector. Higher values for the decay time are obtained for rough surfaces, but the exact value depends on the simulation details. Estimates for the decay time and peak amplitude can be made for different cross section sizes via simple scaling arguments.

We measured the timing resolution of 189 R9800-100 photomultiplier tubes (PMTs), which are a SBA (Super Bialkali, high quantum efficiency) variant of the R9800 high-performance PMT manufactured by Hamamatsu Photonics, and correlated their timing resolutions with various measures of PMT performance, namely Cathode Luminous Sensitivity (CLS), Anode Luminous Sensitivity (ALS), Gain times Collection Efficiency (GCE), Cathode Blue Sensitivity Index (CBSI), Anode Blue Sensitivity Index (ABSI) and dark current. The correlation results show: (1) strong correlations between timing resolution and ALS, ABSI, and GCE; (2) moderate correlations between timing resolution and CBSI; and (3) weak or no correlations between timing resolution and dark current and CLS. The results disclosed that all three measures that include data collected from the anode (ALS, ABSI, and GCE) affect the timing resolution more than either of the two measures that only include photocathode data (CBSI and CLS). We conclude that: (1) the photocathode Quantum Efficiency (QE) and the product of the Gain and the Collection Efficiency (GCE) are the two dominant factors that affect the timing resolution, (2) the GCE variation affects the timing resolution more than the QE variation in the R9800 PMT, and (3) the performance depends on photocathode position.

We have designed and built a high-performance readout electronics system for time-of-flight positron emission tomography (TOF PET) cameras. The electronics architecture is based on the electronics for a commercial whole-body PET camera (Siemens/CPS Cardinal electronics), modified to improve the timing performance. The fundamental contributions in the electronics that can limit the timing resolution include the constant fraction discriminator (CFD), which converts the analog electrical signal from the photo-detector to a digital signal whose leading edge is time-correlated with the input signal, and the time-to-digital converter (TDC), which provides a time stamp for the CFD output. Coincident events are identified by digitally comparing the values of the time stamps. In the Cardinal electronics, the front-end processing electronics are performed by an Analog subsection board, which has two application-specific integrated circuits (ASICs), each servicing a PET block detector module. The ASIC has a built-in CFD and TDC. We found that a significant degradation in the timing resolution comes from the ASIC's CFD and TDC. Therefore, we have designed and built an improved Analog subsection board that replaces the ASIC's CFD and TDC with a high-performance CFD (made with discrete components) and TDC (using the CERN high-performance TDC ASIC). The improved Analog subsection board is used in a custom single-ring LSO-based TOF PET camera. The electronics system achieves a timing resolution of 60 ps FWHM. Prototype TOF detector modules are read out with the electronics system and give coincidence timing resolutions of 259 ps FWHM and 156 ps FWHM for detector modules coupled to LSO and LaBr 3 crystals respectively.

We present a novel structure for the back-side readout silicon photomultipler (SiPM). Current SiPMs are front-illuminated structures with front-side readout, which have relatively small geometric fill factor leading to degradation in their photon detection efficiency (PDE). Back-side readout devices will provide an advantageous solution to achieve high PDE. We designed and investigated a novel structure that would allow back-side readout while creating a region of high electric field optimized for avalanche breakdown. In addition, this structure has relatively high fill factor and also allow direct coupling of individual micro-cell of the SiPM to application-specific integrated circuits. We will discuss the performance that can be attained with this structure through device simulation and the process flow that can be used to fabricate this structure through process simulation.

We present three new autoradiography methods to map positron emission rate of a bio-specimen slice with high resolution. One is based on LBNL scientific charge coupled device (CCD) and the other two are based on conventional CCDs. High conversion efficiency (100k e-h pairs / 0.5 MeV positron) and low dark current (1.75 × 10(-4) e-/pix/sec) can be achieved using the LBNL CCD. The theoretical calculations and preliminary experiments show that an 86 μm spatial resolution can be achieved when imaging a 100 μm thick tissue soaked with (18)F which produce higher energy positron. The main disadvantage of the LBNL CCD we tested is that a very low operating temperature is required to eliminate dark current. This dramatically increases the system cost. In addition, the integration time of the CCD needs to be short enough to avoid overlapping of the positron trajectories. Conventional CCDs have lower conversion efficiency (2k e-h pairs / 0.5 MeV positron) and higher dark current (200 e-/pix/sec), but are more cost-efficient and the requirement for the readout frequency is much lower. The conversion efficiency of the conventional CCD imager can be improved by 17 times by inserting a 100 μm layer of phosphor between the sample and the imager. However, the light emitted from the phosphor screen will be ~100 μm diameter, which severely degrades the spatial resolution. A high readout frequency is also required to avoid the overlapping. The CCD systems designed in this study will be used to map positron emission rate of bio-specimens such as cancerous tissues acquired in regular biopsy procedure. They can also be used to corroborate tracer kinetic modeling at a cellular level.

We report on an investigation of a mulit-anode microchannel plate PMT for time-of-flight PET detector modules. The primary advantages of an MCP lie in its excellent timing properties (fast rise time and low transit time spread), compact size, and reasonably large active area, thus making it a good candidate for TOF applications. In addition, the anode can be segmented into an array of collection electrodes with fine pitch to attain good position sensitivity. In this paper, we investigate using the Photonis Planacon MCP-PMT with a pore size of 10 µm to construct a PET detector module, specifically for time-of-flight applications. We measure the single electron response by exciting the Planacon with pulsed laser diode. We also measure the performance of the Planacon as a PET detector by coupling a 4 mm × 4 mm × 10 mm LSO crystal to individual pixel to study its gain uniformity, energy resolution, and timing resolution. The rise time of the Planacon is 440 ps with pulse duration of about 1 ns. A transit time spread of 120 ps FWHM is achieved. The gain is fairly uniform across the central region of the Planacon, but drops off by as much as a factor of 2.5 around the edges. The energy resolution is fairly uniform across the Planacon with an average value of 18.6±0.7% FWHM. While the average timing resolution of 252±7 ps FWHM is achieved in the central region of the Planacon, it degrades to 280±9 ps FWHM for edge pixels and 316±15 ps FWHM for corner pixels. We compare the results with measurements performed with a fast timing conventional PMT (Hamamatsu R-9800). We find that the R9800, which has significantly higher PDE, has a better timing resolution than the Planacon. Furthermore, we perform detector simulations to calculate the improvement that can be achieved with a higher PDE Planacon. The calculation shows that the Planacon can achieve significantly better timing resolution if it can attain the same PDE as the R-9800, while only a 30% improvement is needed to yield a similar timing resolution as the R-9800.

We have explored methods for optimizing the timing resolution of an LSO-based detector module for a single-ring, "demonstration" time-of-flight PET camera. By maximizing the area that couples the scintillator to the PMT and minimizing the average path length that the scintillation photons travel, a single detector timing resolution of 218 ps fwhm is measured, which is considerably better than the ~385 ps fwhm obtained by commercial LSO or LYSO TOF detector modules. We explored different surface treatments (saw-cut, mechanically polished, and chemically etched) and reflector materials (Teflon tape, ESR, Lumirror, Melinex, white epoxy, and white paint), and found that for our geometry, a chemically etched surface had 5% better timing resolution than the saw-cut or mechanically polished surfaces, and while there was little dependence on the timing resolution between the various reflectors, white paint and white epoxy were a few percent better. Adding co-dopants to LSO shortened the decay time from 40 ns to ~30 ns but maintained the same or higher total light output. This increased the initial photoelectron rate and so improved the timing resolution by 15%. Using photomultiplier tubes with higher quantum efficiency (blue sensitivity index of 13.5 rather than 12) improved the timing resolution by an additional 5%. By choosing the optimum surface treatment (chemically etched), reflector (white paint), LSO composition (co-doped), and PMT (13.5 blue sensitivity index), the coincidence timing resolution of our detector module was reduced from 309 ps to 220 ps fwhm.

A computer simulation study has been conducted to investigate the feasibility of a positron emission tomography (PET) detector design by using micro-channel plate (MCP) photomultiplier tubes (PMT) with transmission-line (TL) read-out and waveform sampling. The detector unit consisted of a 24×24 array of pixelated LSO crystals, each of which was 4×4×25 mm(3) in size, and two 102×102 mm(2) MCP-PMTs coupled to both sides of the scintillator array. The crystal (and TL) pitch was 4.25 mm and reflective medium was inserted between the crystals. The transport of the optical photons inside the scintillator were simulated by using the Geant4 package. The output pulses of the MCP-PMT/TL unit were formed by applying the measured single photo-electron response of the MCP-PMT/TL unit to each individual photon that interacts with the photo-cathode of the MCP-PMT. The waveforms of the pulses at both ends of the TL strips were measured and analyzed to produce energy and timing information for the detected event. An experimental setup was developed by employing a Photonis Planacon MCP-PMT (XP85022) and a prototype TL board for measuring the single photo-electron response of the MCP-PMT/TL. The simulation was validated by comparing the predicted output pulses to measurements obtained with a single MCP-PMT/TL coupled to an LSO crystal exposed to 511 keV gamma rays. The validated simulation was then used to investigate the performance of the proposed new detector design. Our simulation result indicates an energy resolution of ~11% at 511 keV. When using a 400-600 keV energy window, we obtain a coincidence timing resolution of ~323 ps FWHM and a coincidence detection efficiency of ~40% for normally-incident 511keV photons. For the positioning accuracy, it is determined by the pitch of the TLs (and crystals) in the direction normal to the TLs and measured to be ~2.5 mm in the direction parallel to the TLs. The energy and timing obtained at the front- and back-end of the scintillator array also show differences that are correlated with the depth of interaction of the event.

Recent advancements in fast scintillating materials and fast photomultiplier tubes (PMTs) have stimulated renewed interest in time-of-flight (TOF) positron emission tomography (PET). It is well known that the improvement in the timing resolution in PET can significantly reduce the noise variance in the reconstructed image resulting in improved image quality. In order to evaluate the timing performance of scintillation detectors used in TOF PET, we use Monte Carlo analysis to model the physical processes (crystal geometry, crystal surface finish, scintillator rise time, scintillator decay time, photoelectron yield, PMT transit time spread, PMT single-electron response, amplifier response and time pick-off method) that can contribute to the timing resolution of scintillation-detector systems. In the Monte Carlo analysis, the photoelectron emissions are modeled by a rate function, which is used to generate the photoelectron time points. The rate function, which is simulated using Geant4, represents the combined intrinsic light emissions of the scintillator and the subsequent light transport through the crystal. The PMT output signal is determined by the superposition of the PMT single-electron response resulting from the photoelectron emissions. The transit time spread and the single-electron gain variation of the PMT are modeled in the analysis. Three practical time pick-off methods are considered in the analysis. Statistically, the best timing resolution is achieved with the first photoelectron timing. The calculated timing resolution suggests that a leading edge discriminator gives better timing performance than a constant fraction discriminator and produces comparable results when a two-threshold or three-threshold discriminator is used. For a typical PMT, the effect of detector noise on the timing resolution is negligible. The calculated timing resolution is found to improve with increasing mean photoelectron yield, decreasing scintillator decay time and decreasing transit time spread. However, only substantial improvement in the timing resolution is obtained with improved transit time spread if the first photoelectron timing is less than the transit time spread. While the calculated timing performance does not seem to be affected by the pixel size of the crystal, it improves for an etched crystal compared to a polished crystal. In addition, the calculated timing resolution degrades with increasing crystal length. These observations can be explained by studying the initial photoelectron rate. Experimental measurements provide reasonably good agreement with the calculated timing resolution. The Monte Carlo analysis developed in this work will allow us to optimize the scintillation detectors for timing and to understand the physical factors limiting their performance.

We present the design for OpenPET, an electronics readout system designed for prototype radiotracer imaging instruments. The critical requirements are that it has sufficient performance, channel count, channel density, and power consumption to service a complete camera, and yet be simple, flexible, and customizable enough to be used with almost any detector or camera design. An important feature of this system is that each analog input is processed independently. Each input can be configured to accept signals of either polarity as well as either differential or ground referenced signals. Each signal is digitized by a continuously sampled ADC, which is processed by an FPGA to extract pulse height information. A leading edge discriminator creates a timing edge that is "time stamped" by a TDC implemented inside the FPGA. This digital information from each channel is sent to an FPGA that services 16 analog channels, and information from multiple channels is processed by this FPGA to perform logic for crystal lookup, DOI calculation, calibration, etc. As all of this processing is controlled by firmware and software, it can be modified / customized easily. The system is open source, meaning that all technical data (specifications, schematics and board layout files, source code, and instructions) will be publicly available.

As an approach to realizing all-digital data acquisition for positron emission tomography (PET), we have previously proposed and studied a multi-threshold sampling method to generate samples of a PET event waveform with respect to a few user-defined amplitudes. In this sampling scheme, one can extract both the energy and timing information for an event. In this paper, we report our prototype implementation of this sampling method and the performance results obtained with this prototype. The prototype consists of two multi-threshold discriminator boards and a time-to-digital converter (TDC) board. Each of the multi-threshold discriminator boards takes one input and provides up to 8 threshold levels, which can be defined by users, for sampling the input signal. The TDC board employs the CERN HPTDC chip that determines the digitized times of the leading and falling edges of the discriminator output pulses. We connect our prototype electronics to the outputs of two Hamamatsu R9800 photomultiplier tubes (PMTs) that are individually coupled to a 6.25×6.25×25mm(3) LSO crystal. By analyzing waveform samples generated by using four thresholds, we obtain a coincidence timing resolution of about 340 ps and an ∼18% energy resolution at 511 keV. We are also able to estimate the decay-time constant from the resulting samples and obtain a mean value of 44ns with an ∼9 ns FWHM. In comparison, using digitized waveforms obtained at a 20 GSps sampling rate for the same LSO/PMT modules we obtain ∼300 ps coincidence timing resolution, ∼14% energy resolution at 511 keV, and ∼5 ns FWHM for the estimated decay-time constant. Details of the results on the timing and energy resolutions by using the multi-threshold method indicate that it is a promising approach for implementing digital PET data acquisition.

The asymmetry in the rho angular distribution in the sequential decay Omega+-->LamdaKappa+-->rhopi+Kappa+. has been measured to be alphaOmegaalphaLamda=[+1.16+/-0.18(stat)+/-0.17(syst)]x10(-2) using 1.89x10(6) unpolarized Omega+ decays recorded by the HyperCP (E871) experiment at Fermilab. Using the known value of alphaLamda, and assuming that alphaLamda=-alphaLamda, alphaOmega=[-1.81+/-0.28(stat)+/-0.26(syst)]x10(-2). A comparison between this measurement of alphaOmegaalphaLamda and recent measurements of alphaOmegaalphaLamda made by HyperCP shows no evidence of a violation of CP symmetry.

We present a retrospective on the LBNL Positron Emission Mammography (PEM) project, looking back on our design and experiences. The LBNL PEM camera utilizes detector modules that are capable of measuring depth of interaction (DOI) and places them into 4 detector banks in a rectangular geometry. In order to build this camera, we had to develop the DOI detector module, LSO etching, Lumirror-epoxy reflector for the LSO array (to achieve optimal DOI), photodiode array, custom IC, rigid-flex readout board, packaging, DOI calibration and reconstruction algorithms for the rectangular camera geometry. We will discuss the high-lights (good and bad) of these developments.

A sensitive search for the lepton-number-violating decay Xi(-)-->pmu(-)mu(-) has been performed using a sample of approximately 10(9) Xi(-) hyperons produced in 800 GeV/c p-Cu collisions. We obtain B(Xi(-)-->pmu(-)mu(-))<4.0x10(-8) at 90% confidence, improving on the best previous limit by 4 orders of magnitude.

A sensitive search for the rare decays Omega(-)--> Lambdapi(-) and Xi(0)--> ppi(-) has been performed using data from the 1997 run of the HyperCP (Fermilab E871) experiment. Limits on other such processes do not exclude the possibility of observable rates for |DeltaS| = 2 nonleptonic hyperon decays, provided the decays occur through parity-odd operators. We obtain the branching-fraction limits B(Omega(-)-->Lambdapi(-)) < 2.9 x 10(-6) and B(Xi(0)--> ppi(-)) < 8.2 x 10(-6), both at 90% confidence level.

We report the first evidence for the decay Sigma(+)-->pmu(+)mu(-) from data taken by the HyperCP (E871) experiment at Fermilab. Based on three observed events, the branching ratio is B(Sigma(+)-->pmu(+)mu(-))=[8.6(+6.6)(-5.4)(stat)+/-5.5(syst)]x10(-8). The narrow range of dimuon masses may indicate that the decay proceeds via a neutral intermediate state, Sigma(+)-->pP(0),P0-->mu(+)mu(-) with a P0 mass of 214.3+/-0.5 MeV/c(2) and branching ratio B(Sigma(+)-->pP(0),P0-->mu(+)mu(-))=[3.1(+2.4)(-1.9)(stat)+/-1.5(syst)]x10(-8).

We have compared the p and p angular distributions in 117 x 10(6) Xi- -->Lambdapi- -->ppi-pi- and 41 x 10(6) Xi+ -->Lambda pi+ -->p pi+pi+ decays using a subset of the data from the HyperCP experiment (E871) at Fermilab. We find no evidence of CP violation, with the direct-CP-violating parameter AXiLambda identical with (alphaXialphaLambda-alpha Xialpha Lambda)/(alphaXialphaLambda+alphaXialphaLambda)=[0.0+/-5.1(stat)+/-4.4(syst)] x 10(-4).

Using data collected with the HyperCP (E871) spectrometer during the 1997 fixed-target run at Fermilab, we report the first observation of the decay K--->pi(-)mu(+)mu(-) and new measurements of the branching ratios for K+/--->pi(+/-)mu(+)mu(-). By combining the branching ratios for the decays K+-->pi(+)mu(+)mu(-) and K--->pi(-)mu(+)mu(-), we measure Gamma(K+/--->pi(+/-)mu(+)mu(-))/Gamma(K+/--->all) = (9.8+/-1.0+/-0.5)x10(-8). The CP asymmetry between the rates of the two decay modes is [Gamma(K+-->pi(+)mu(+)mu(-))-Gamma(K--->pi(-)mu(+)mu(-))]/[Gamma(K+-->pi(+)mu(+)mu(-))+Gamma(K--->pi(-)mu(+)mu(-))] = -0.02+/-0.11+/-0.04.