SIPHRA is an integrated circuit (IC) for the readout of photon detectors, such as photomultiplier tubes (PMTs), silicon photomultipliers (SiPMs), and multi-pixel photon counters (MPPCs). The IC has 16 input channels and one summing channel. Each channel can be used for pulse height spectroscopy and timing. The summing channel is important for the readout of detector arrays with monolithic scintillators. The programmable shaping time of 200 ns, 400 ns, 800 ns, or 1600 ns allows for pulse-height spectroscopy using scintillators with different light emission properties. The current mode input stage (CMIS) is designed for large negative charge (-16 nC, -8 nC, -4 nC, and -0.4 nC), depending on the programmable attenuation, and it accommodates large capacitive load (several nF) and large leakage current (up to -100 μA from dark counts). Alternatively, the CMIS can be by-passed to allow for positive charge depending on programmable gain (+40 pC, +4 pC, +0.4 pC). The IC contains one 12-bit analog-to-digital converter (ADC) that allows for digitization of the pulseheights from all channels, including the summing channel at a sampling rate of 50 ksps. Every channel output is available for external use and provides either the analog or a digital trigger/timing pulse with fixed width or time-over-threshold. The programmable channel output facilitates many applications, such as external waveform sampling and digitization, pulse height and time spectroscopy, pulse counting, and triggering. The IC operates at 3.3-V supply voltage and dissipates about 15 mW without CMIS and 30 mW with CMIS active. To save power, any channel or function can be powered down. The ASIC has a serial peripheral interface (SPI) for programming its register settings and for slow ADC data readout; faster readout with up to 1 Mbit/s is possible via a serial data transmission line.

The ESA Next Generation Radiation Monitor (NGRM) will be the successor of the ESA Standard Radiation Environment Monitor (SREM) that is still measuring near-Earth and interplanetary space radiation environment onboard 6 different spacecrafts. NGRM will measure protons from 2 MeV up to 200 MeV, electrons from 100keV up to 7MeV, as well as LET spectrum of ions. Compared to SREM, NGRM will provide a much better energy resolution, will be smaller (<;1L), lighter (<;1kg) and consume less energy (<;1W). In this paper we describe the detection concept of the NGRM and present the detailed Monte Carlo analysis of the performance of the NGRM detector system.

We describe an MR-compatible SPECT camera for small animals. The SPECT camera system can be inserted into the bore of a state-of-the-art MRI system and allows researchers to acquire tomographic images from a mouse in-vivo with the MRI and the SPECT acquiring simultaneously. The SPECT system provides functional information, while MRI provides anatomical information. Until today it was impossible to operate conventional SPECT inside the MRI because of mutual interference. The new SPECT technology is based on semiconductor radiation sensors (CZT, ASICs), and it fits into conventional high field MRI systems with a minimum 12-cm bore size. The SPECT camera has an MR-compatible multi-pinhole collimator for mice with a Ø25-mm field-of-view. For the work reported here we assembled a prototype SPECT camera system and acquired SPECT and MRI data from radioactive sources and resolution phantoms using the camera outside and inside the MRI.

The goal of this paper was to investigate the benefits that could be realistically achieved on a microCT imaging system with an energy-resolved photon-counting x-ray detector. To this end, we built and evaluated a prototype microCT system based on such a detector. The detector is based on cadmium telluride (CdTe) radiation sensors and application-specific integrated circuit (ASIC) readouts. Each detector pixel can simultaneously count x-ray photons above six energy thresholds, providing the capability for energy-selective x-ray imaging. We tested the spectroscopic performance of the system using polychromatic x-ray radiation and various filtering materials with K-absorption edges. Tomographic images were then acquired of a cylindrical PMMA phantom containing holes filled with various materials. Results were also compared with those acquired using an intensity-integrating x-ray detector and single-energy (i.e. non-energy-selective) CT. This paper describes the functionality and performance of the system, and presents preliminary spectroscopic and tomographic results. The spectroscopic experiments showed that the energy-resolved photon-counting detector was capable of measuring energy spectra from polychromatic sources like a standard x-ray tube, and resolving absorption edges present in the energy range used for imaging. However, the spectral quality was degraded by spectral distortions resulting from degrading factors, including finite energy resolution and charge sharing. We developed a simple charge-sharing model to reproduce these distortions. The tomographic experiments showed that the availability of multiple energy thresholds in the photon-counting detector allowed us to simultaneously measure target-to-background contrasts in different energy ranges. Compared with single-energy CT with an integrating detector, this feature was especially useful to improve differentiation of materials with different attenuation coefficient energy dependences.

The overall aim of this work was to evaluate the potential for improving in vivo small animal microCT through the use of an energy resolved photon-counting detector. To this end, we developed and evaluated a prototype microCT system based on a second-generation photon-counting x-ray detector which simultaneously counted photons with energies above six energy thresholds. First, we developed a threshold tuning procedure to reduce the dependence of detector uniformity and to reduce ring artifacts. Next, we evaluated the system in terms of the contrast-to-noise ratio in different energy windows for different target materials. These differences provided the possibility to weight the data acquired in different windows in order to optimize the contrast-to-noise ratio. We also explored the ability of the system to use data from different energy windows to aid in distinguishing various materials. We found that the energy discrimination capability provided the possibility for improved contrast-to-noise ratios and allowed separation of more than two materials, e.g., bone, soft-tissue and one or more contrast materials having K-absorption edges in the energy ranges of interest.

We describe a single photon emission computed tomograph (SPECT) which can be operated inside state-of-theart magnetic resonance imaging (MRI) systems. The combined SPECT/MRI system allows one to acquire simultaneously the data from both modalities and co-register the images in space and time. Unlike conventional SPECT systems, which use scintillators and photomultiplier tubes, the new SPECT is based on the semiconductor cadmium zinc telluride (CZT) and application specific integrated circuits (ASICs) - two technologies which are MR-compatible and almost insensitive to magnetic fields. The γ-radiation imaging system has a 3-cm diameter fieldof- view and allows one to acquire regional tomographic images in mice without moving the camera.

Purpose: We had previously shown the MR compatibility of CZT based single photon detectors and associated nuclear data acquisition electronics [1]. The purpose of the present work was to study the feasibility a combined MRSPECT system operating inside a 4T MR system for dual-modality MR-SPECT imaging. The current investigation included a study of the effect of SPECT detector/collimator system on the MR images (MRI) as well as the effect of MR on SPECT imaging and the feasibility of simultaneous MR – SPECT imaging. Methods: In order to simulate an experimental SPECT system that would work inside an MR scanner we constructed a miniversion of such a system using a 1x1 inch (16x16) pixellated CZT detector [1] in conjunction with the same size parallel –hole collimator made of lead (Pb). The collimator/detector assembly was inserted into a custom dual modality RF coil [2] for simultaneous imaging. The nuclear electronics were kept outside the scanner room while all the wires were passed through the penetration panel with appropriate filtering. The detector assembly was rotated around the object to acquire 30 equally spaced angular views around 360-deg for SPECT. From each nuclear projection image 3 rows were selected to make a 4.76mm slice in the reconstructed image. The data acquisition time for each scintigraphic projection image was 1.5 min. The system is shown in Fig.1. The MR pulse sequence parameters were: 2DT1 Weighted Spin Echo, TR/TE=500/20ms, matrix = 128x128, FOV = 40x40mm, thickness=4.0mm, NEX=2, BW = 33.3 kHz. The number of slices was 7 with no gap. Parallel beam filtered backprojection algorithm was used for the SPECT reconstruction that was interpolated to the same matrix size as the MRI viz. 128x128. Each projection image was normalized by a flood image obtained at the same view to take into account of the Lorentz effect due to the magnetic field and to correct for detector non-uniformity. A center of rotation correction due to the Lorentz effect was also applied to each projection after the flood field correction. In order to test the effect of the operating MR system on the pulse height spectrum obtained with the CZT detector we measured the energy spectrum (PHS) of Tc-99m inside the magnet with and without the MR system being operational. We had previously shown that the static magnetic field had no effect on the PHS [1]. Two multimodality phantoms made of Plexiglas were filled with a paramagnetic solution (Gd-DTPA) and 1 mCi of Tc-99m for sequential or simultaneous imaging. The first phantom was a hollow cylinder with an inner diameter of 22.5mm and a length of 82mm (uniform phantom). The second phantom was an identical cylinder that had 3 Plexiglas rods with 4.8mm diameter each separated from each other by 9.6mm center-to-center for resolution measurement in the reconstructed images (resolution phantom). Results: Figure 2 compares the PHS of Tc-99m obtained inside the 4T magnet with and without the MR system performing scanning. It is seen that the two spectra are identical and thus an operating MR system has no effect on SPECT data acquisition. This implies that it should be possible to acquire simultaneous MR and SPECT images with such a system. Figure 4 shows the MR and SPECT images of the uniformity phantom: A) MRI while SPECT off; B) MRI with SPECT on; C) SPECT while MRI off; and D) SPECT while MRI on. A close inspection of Figures 4.A and 4.B and Figures 4.C and 4.D shows no observable degradation from simultaneous operation of MR and SPECT data acquisition. Figure 5 shows the MR and SPECT images of the resolution phantom: A) MRI while SPECT off; B) MRI with SPECT on; C) SPECT while MRI off; and D) SPECT while MRI on. Again we do not observe any interference of MR on SPECT and vice versa. Discussion: The work presented here demonstrates experimentally the feasibility of constructing an MR compatible SPECT scanner that could be used for acquiring spatially and temporally co-registered MR and SPECT images that would be of great significance in molecular imaging [3]. The energy spectrum shown in Fig. 3 clearly demonstrates that one collect nuclear imaging data while the MR scanner is acquiring MRI. Figures 4 and 5 show that the spatial resolution in the SPECT images is poor compared to the MRI in the current setup. This is due to the fact that a parallel-hole collimator was used in front of the SPECT detector. However, as we have shown previously [4] it is possible to use the high-resolution MR images as a priori information in the SPECT reconstruction for improving the spatial resolution. Additionally the use of multi-pinhole collimators instead of the parallel-hole collimator would also help improving the SPECT spatial resolution [5].

Purpose: Multi-modality medical imaging was pioneered with PET/CT and, more recently, SPECT/CT. In this work we report on the development of a dual-modality, SPECT/MRI pre-clinical research tool and present in vivo dual-modality SPECT/MRI images of a laboratory animal. Previous reports have shown the ability of semiconductor cadmium-zinc-telluride (CdZnTe, or CZT) to serve as a high-performance, pixellated nuclear imaging device under strong magnetic fields [1]. A stationary tomographic SPECT system capable of obtaining 24 views simultaneously has been designed [2], fabricated, and tested [3]. Reconstruction algorithms for the 24-view SPECT data are under development [4]. In this work we describe the full system design, engineering aspects of materials and EMI shielding techniques, calibration and correction methods, and test results of signal cross-talk and simultaneous phantom and animal imaging.

Purpose: Most of the nuclear imaging systems, especially those that are based on single photon emission detection, require the use of a collimatorusually made out of lead (Pb) or some other high Z material such as tungsten (W). Even some PET systems employ some sort of slat-septa collimation, although the important part of PET image formation is done electronically via coincidence detection. Multi-modality imaging combines two or more complementary imaging systems to provide spatially and temporally co-registered multidimensional images by the chosen modalities. The purpose of this study was to investigate the affect of Pb collimators for a hybrid system that combines single photon scintigraphic imaging with MRI. We have tested a parallel-hole collimator in combination with a shielding plate both made of lead inside a 4T MR system to evaluate their effect on the MR image quality. Although the current common belief for the choice of such MR-compatible shielding and collimator materials is to use a high-Z, non-electrically conductive material, we tested a Pb parallel-hole collimator since it is a cost-effective and proven material in nuclear imaging. Methods: The effect of Pb collimator on MR images was studied by designing and constructing a true-multimodality RF coil specifically developed for this purpose. The coil design is based on a birdcage configuration where the copper rungs of the birdcage are displaced to allow for the integration of a nuclear collimator right into the RF coil [1]. A picture of such a coil designed for a parallelhole collimator is shown in Fig. 1. This is an 8-section birdcage RF coil where the separation between the two rungs are opened up to allow a 2.54 mm x 2.54 mm parallel-hole collimator to be inserted through. In other types of applications where pinhole collimators may be used, a hole of appropriate size may be cut out for the pinhole. This specific RF coil was designed to work with a single CZT detector (16x16 array – 1.6mm pixel size) and had an inside diameter of 30 mm and a wall thickness of 7 mm. It should be emphasized that this fully integrated design allows for bringing the collimator as close to the object as desired to improve spatial resolution without worrying about RF coil attenuation of gamma rays. A cylindrical phantom with a 24mm diameter was filled with CuSO4 and placed inside the RF coil with and without the Pb collimator. The pulse sequence parameters were: T1 Weighted Spin Echo, TR/TE=500/20ms, matrix = 128*128, FOV = 40*40mm, thickness=4.0mm, NEX=2, BW = 33.3 kHz. The center frequency of the RF coil and its Q-factor were measured under different loading conditions. Additionally MR images were acquired under different conditions to evaluate the MR performance in terms of SNR as well as signal uniformity changes resulting from the presence of Pb collimator. MR images without the Pb collimator were acquired as a reference set then the Pb collimator was inserted into the RF coil as shown in Fig. 2. In order to improve the signal uniformity in the images we also placed small Pb sheets on the other 3 sides to balance the Pb collimator as shown in Fig. 3 and repeated the measurements. Results: The center frequency of the RF coil and its Q-value were measured without and with the collimator/detector assembly inserted into the RF coil. The results are given in Table 1. The first column in Table 1 corresponds to the 4 cases shown in Fig. 4. We also analyzed the MR images obtained under these conditions to assess the signal nonuniformity arising from the presence of the Pb collimator and additional shielding near the phantom. MR images corresponding to the 4 cases shown in Table 1 are shown in Fig. 4 with a horizontal profile drawn to compute the signal drop off due to the Pb collimator that was on the left-hand side of Figs. 4B and 4D. The signal non-uniformity was calculated as the % difference in the signal between the sides close to the collimator vs. farthest from it. Table 2 summarizes the signal non-uniformity and the image SNR for the four cases presented here. Discussion: As expected, the introduction of a Pb collimator near or inside the RF coil causes a shift in the center frequency and a drop in the Q-factor. Consequently this results in various degrees of SNR changes as a Pb collimator and additional Pb shield are introduced. The SNR reduction is due to the RF absorption by the collimator and other surrounding metal. It should be noted that the addition of Pb shields as shown in Fig. 3 improves the image signal non-uniformity considerably even in the presence of a Pb collimator 5mm away from the object. Although the introduction of the Pb shields to improve the uniformity results in the degradation of the image SNR (see Table 2) since the time it takes to acquire nuclear images is much longer than for MRI data collection one can easily increase the number of averages in the MRI to overcome this problem. In conclusion our results demonstrate that lead (Pb) collimators could be used in MR -nuclear dual modality imaging systems with appropriate precautions such as balancing the collimator material with other shielding made of the same material as the collimator.

In future HEP accelerators, such as the LHC (CERN), detectors and electronics in the vertex region of the experiments will su er from extreme radiation. Thus radiation hardness is required for both detectors and electronics to survive in this harsh environment. CVD diamond, which is investigated by the RD42 Collaboration at CERN, can meet these requirements. Samples of up to 2 4 cm2 have been grown and re ned for better charge collection properties, which are measured with a source or in a testbeam. A large number of diamond samples has been irradiated with hadrons to uences of up to 5 1015 cm 2 to study the e ects of radiation. Both strip and pixel detectors were prepared in various geometries. Samples with strip metallization have been tested with both slow and fast readout electronics, and the rst diamond pixel detector proved fully functional with LHC electronics.

We present recent results on the use of Chemical Vapor Deposition (CVD) diamond microstrip detectors for charged particle tracking. A series of detectors was fabricated using 1 1 cm 2 diamonds. Good signal-to-noise ratios were observed using both slow and fast readout electronics. For slow readout electronics, 2 s shaping time, the most probable signal-to-noise ratio was 50 to 1. For fast readout electronics, 25 ns peaking time, the most probable signal-to-noise ratio was 7 to 1. Using the rst 2 4 cm 2 diamond from a production CVD reactor with slow readout electronics, the most probable signal-to-noise ratio was 23 to 1. The spatial resolution achieved for the detectors was consistent with the digital resolution expected from the detector pitch.

The RD42 Collaboration R. Wedenig9,—, W. Adam, C. Bauer, E. Berdermann, P. Bergonzo, F. Bogani, E. Borchi, A. Brambilla, M. Bruzzi, C. Colledani, J. Conway, W. Dabrowski, P. Delpierre, A. Deneuville, W. Dulinski, B. van Eijk, A. Fallou, F. Fizzotti, F. Foulon, M. Friedl, K.K. Gan, E. Gheeraert, E. Grigoriev, G. Hallewell, R. Hall-Wilton, S. Han, F. Hartjes, J. Hrubec, D. Husson, H. Kagan, D. Kania, J. Kaplon, C. Karl, R. Kass, K.T. Knöpfle, M. Krammer, A. Logiudice, R. Lu, P.F. Manfredi, C. Manfredotti, R.D. Marshall, D. Meier, M. Mishina, A. Oh, L.S. Pan, V.G. Palmieri, M. Pernicka, A. Peitz, S. Pirollo, P. Polesello, K. Pretzl, M. Procario, V. Re, J.L. Riester, S. Roe, D. Roff, A. Rudge, O. Runolfsson, J. Russ, S. Schnetzer, S. Sciortino, V. Speziali, H. Stelzer, R. Stone, B. Suter, R.J. Tapper, R. Tesarek, M. Trawick, W. Trischuk, E. Vittone, A. Wagner, A.M. Walsh, P. Weilhammer, C. White, W. Zeuner, H. Ziock, M. Zoeller

The value of pre-clinical PET and MRI is established, and SPECT also is now routinely used in drug development, disease treatment, and molecular biology research. The aim of this research is to prepare and test imaging prototypes of both SPECT and PET with similar form-factors for use within the core of a pre-clinical MRI. To test the imaging properties of the SPECT detector in a magnetic field a prototype based on CZT semiconductor detectors has been fabricated. This detector has a field-of-view of 2.54 times 12.7 cm2 and is fitted with a parallel hole collimator for mouse imaging. The prototype is designed to obtain a tomographic dataset by manual positioning within a 90-degree projection range. Shielding, support, positioning, and cooling materials were selected to minimally influence the magnetic field homogeneity. Fabrication of a stationary, polygonal ring of CZT has commenced. An APD-based, light-sharing detector has been evaluated for use in PET. The detector configuration is a 2 times 2 array of 10 mm times 10 mm Hamamatsu S8664-1010 APDs used to read out multi-crystal LYSO blocks. The design of the detector and front-end electronics is optimized for spectroscopic and timing performance, minimization of power dissipation, and low electromagnetic interference. Current plans are for assembling five rings, each with 17 modules, to give a 13.0 cm axial FOV with a 14 cm inner diameter. The results on the tests and performance of both types of modules are reported.

Silicon photomultipliers are of great interest for scintillation light detection in gamma spectroscopy, due to their large gain and high speed. In this article we report about the use of a silicon photomultiplier for radiation detection and gamma spectroscopy. It is our goal to understand silicon photomultipliers to the extent that we could use them in imaging instrumentation for applications in nuclear medicine. We measure gain of up to 106, and pixel dark count rate of a few MHz related to the leakage current of about 1 muA. Using an LED pulser, we measure the signal risetime of less than a few nano-seconds, and 300-ps rms jitter relative to a photomultiplier tube. We use CsI:Tl and LSO:Ce scintillators and measure energy spectra from radionuclides. In CsI:Tl we measure an energy resolution of 32% FWHM at 122 keV from Co-57, and extrapolate a resolution of 30% FWHM at 140 keV. We observe reasonably good energy spectra from LSO:Ce scintillators. Similar energy spectra were obtained with a photomultiplier tube as a reference

Compton collimated imaging may improve the detection of gamma rays emitted by radioisotopes used in single photon emission computed tomography (SPECT). We present a crude prototype consisting of a single 500μm thick, 256 pad silicon detector with pad size of 1.4×1.4mm2, combined with a 15×15×1cm3 NaI scintillator crystal coupled to a set of 20 photo multipliers. Emphasis is placed on the performance of the silicon detector and the associated read-out electronics, which has so far proved to be the most challenging part of the set-up. Results were obtained using the VATAGP3, 128 channel low-noise self-triggering ASIC as the silicon detector's front-end. The noise distribution (σ) of the spectroscopic outputs gave an equivalent noise charge (ENC) with a mean value of 〈σ〉=137e with a spread of 10e, corresponding to an energy resolution of 1.15keV FWHM for the scattered electron energy. Threshold settings above 8.2keV were required for stable operation of the trigger. Coincident Compton scatter events in both modules were observed for photons emitted by 57Co source with principal gamma ray energies of 122 and 136keV.

A new low-noise multichannel application specific integrated circuit (ASIC) has been designed for use in nuclear medical imaging single photon emission computed tomography (SPECT), and prototypes have been produced and tested. The ASIC has 128 charge sensitive amplifiers (CSAs) followed by shapers, threshold discriminators, and sample-and-hold. The ASIC delivers analog signal amplitude and digital address from channels with signals above threshold, and optionally allows one to read signals from any channel using a random-access readout. The total power consumption is 358 mW. The ASIC parameters have been specified and measured as follows: each amplifier can measure charge within a linear dynamic range from -40000 e to +40000 e. The preamplifier noise is 56 e+11.6 e/pF at 10 /spl mu/s shaping time. The ASIC is ideally suited for readout of radiation sensors with a few pF capacitance and less than 100 pA dc input current. The ASIC has been tested with silicon diode pad sensors: the diode had a capacitance of /spl ap/1 pF and a leakage current of 20 pA. We measure an equivalent noise charge of 85 e, and achieve a triggering threshold of 800 e. Energy spectra from radioactive isotopes have been measured with energy resolution of 823 eV FWHM for 5.9 keV X-rays.

Chemical Vapor Deposition (CVD) diamond has been discussed extensively as an alternative sensor material for use very close to the interaction region of the LHC and other machines where extreme radiation conditions exist. During the last seven years the RD42 collaboration has developed diamond detectors and tested them with LHC electronics towards the end of creating a device usable by experiments. The most recent results of this work are presented. Recently, a new form of CVD diamond has been developed: single crystal CVD diamond which resolves many of the issues associated with poly-crystalline CVD material. The first tests of this material are also presented.