Performance of the Tachyon Time-of-Flight PET Camera.

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 mm2 side of 6.15 × 6.15 × 25 mm3 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.

Optimization of a LSO-Based Detector Module for Time-of-Flight PET.

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.

Improving timing performance of double-ended readout in TOF-PET detectors.

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.

OpenPET: A Flexible Electronics System for Radiotracer Imaging.

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.

A front-end readout Detector Board for the OpenPET electronics system.

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.

MODELING TIME DISPERSION DUE TO OPTICAL PATH LENGTH DIFFERENCES IN SCINTILLATION DETECTORS.

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.

A New Time Calibration Method for Switched-capacitor-array-based Waveform Samplers.

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.

High-performance electronics for time-of-flight PET systems.

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 LaBr3 crystals respectively.

Real-time quantitative ex vivo direct autoradiography with 10 µm pixel resolution.

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.

A Design of a PET Detector Using Micro-Channel Plate Photomultipliers with Transmission-Line Readout.

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.

A Multi-Threshold Sampling Method for TOF PET Signal Processing.

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.

Measurement of the asymmetry in the decay Omega+-->LamdaKappa+-->rhopi+Kappa+.

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.

Search for the lepton-number-violating decay Xi(-)-->pmu(-)mu(-).

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.

Evidence for the decay sigma+ --> pmu+ mu-.

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).

Search for DeltaS = 2 nonleptonic hyperon decays.

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.

Observation of the decay K- --> pi(-)mu(+)mu(-) and measurements of the branching ratios for K+/- --> pi(+/-)mu(+)mu(-).

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.

Search for CP violation in charged-Xi and Lambda hyperon decays.

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).