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A new small animal PET based on SiPM and monolithic LYSO crystals has been developed. Eight detector modules form the PET ring, each mounting an array of ${\text{12}} \times {\text{12}}$ SiPMs coupled to a readout providing the summed signals of the pixels on each of the 12 rows and 12 columns of the SiPM array. This design makes it possible to accurately determine the centroid of the scintillation light distribution with about 1.6 mm full width at half maximum (FWHM) resolution without correction for the 1 mm source size, and the photon depth of interaction (DOI) with nearly 2 mm FWHM. This single ring PET system has a homogeneous spatial resolution across the entire 80 mm transaxial field of view (FOV) of about 1 mm FWHM. The noise equivalent count rate (NECR) peak is estimated to occur at around 39.2 MBq with a rate of approximately 82.7 kcps for the mouse-like phantom and 22 kcps at 48.1 MBq for the rat-like phantom. Following the NEMA protocol, the peak absolute sensitivity in the center of the FOV is 2.8% for a 30% peak energy window. A pilot test injecting NaF to a mouse of 20 grams is also presented. Finally, the PET ring has been tested in front of a high field 15.2 T Magnetic Resonance (MR). No significant variation on energy and spatial resolution across the FOV has been observed due to the presence of the magnetic field.

The main aim of this work is to provide a method to retrieve the intrinsic resolution of detector blocks based on monolithic crystals in a fully assembled scanner. This method suggests a software collimation to the original data. The results are compared with the traditional approach of separating two detector blocks far enough, resulting in geometrical collimation.An empirical equation has been deduced to fit the experimental data in which the detector intrinsic resolution follows a Gaussian distribution and the contribution of the source, given the small size of 0.25 mm in diameter, follows a Lorentzian profile. The experiments resulted in an average detector intrinsic spatial resolution of 0.6 mm FWHM, with a standard deviation error of 0.1 mm. These tests show a method to determine the intrinsic resolution of monolithic-based detector blocks, once assembled in the PET system, with high accuracy.

Purpose: To explore the potential of grating-based x-ray phase-contrast imaging for clinical applications, a first compact gantry system was developed. It is designed such that it can be implemented into anin-vivo small-animal phase-contrast computed tomography (PC-CT) scanner. The purpose of the present study is to assess the accuracy and quantitativeness of the described gantry in both absorption and phase-contrast. Methods: A phantom, containing six chemically well-defined liquids, was constructed. A tomography scan with cone-beam reconstruction of this phantom was performed yielding the spatial distribution of the linear attenuation coefficient μ and decrement δ of the complex refractive index. Theoretical values of μ and δ were calculated for each liquid from tabulated data and compared with the experimentally measured values. Additionally, a color-fused image representation is proposed to display the complementary absorption and phase-contrast information in a single image. Results: Experimental and calculated data of the phantom agree well confirming the quantitativeness and accuracy of the reconstructed spatial distributions of μ and δ. The proposed color-fused image representation, which combines the complementary absorption and phase information, considerably helps in distinguishing the individual substances. Conclusions: The concept of grating-based phase-contrast computed tomography(CT) can be implemented into a compact, cone-beam geometry gantry setup. The authors believe that this work represents an important milestone in translating phase-contrast x-ray imaging from previous proof-of-principle experiments to first preclinical biomedical imaging applications on small-animal models.

We have demonstrated that structures down to 150 nm can be visualized in X-ray projection images using nanofocus X-ray sources. Due to their unlimited depth of focus, they do not possess a limit on the specimen size. This is essential for three-dimensional tomographic imaging of samples with a diameter larger than a few microns. Further simulation studies have shown that optimization of the detector response curve and switching from a reflective X-ray target to a transmission target should allow us to reach sub-100-nm resolutions.

PURPOSE: Previously, we demonstrated the potential of positron emission tomography detectors consisting of monolithic scintillation crystals read out by arrays of solid-state light sensors. We reported detector spatial resolutions of 1.1-1.3 mm full width at half maximum (FWHM) with no degradation for angles of incidence up to 30 degrees, energy resolutions of approximately 11% FWHM, and timing resolutions of approximately 2 ns FWHM, using monolithic LYSO:Ce3+ crystals coupled to avalanche photodiode (APD) arrays. Here we develop, validate, and demonstrate a simple model of the detector point spread function (PSF) of such monolithic scintillator detectors. METHODS: A PSF model was developed that essentially consists of two convolved components, one accounting for the spatial distribution of the energy deposited by annihilation photons within the crystal, and the other for the influences of statistical signal fluctuations and electronic noise. The model was validated through comparison with spatial resolution measurements on a detector consisting of an LYSO:Ce3+ crystal read out by two APD arrays. RESULTS: The model is shown to describe the measured detector spatial response well at the noise levels found in the experiments. In addition, it is demonstrated how the model can be used to correct the measured spatial response for the influence of the finite diameter of the annihilation photon beam used in the experiments, thus obtaining an estimate of the intrinsic detector PSF. CONCLUSIONS: Despite its simplicity, the proposed model is an accurate tool for analyzing the detector PSF of monolithic scintillator detectors and can be used to estimate the intrinsic detector PSF from the measured one.

Item does not contain fulltext Much research is being conducted on position-sensitive scintillation detectors for medical imaging, particularly for emission tomography. Monte Carlo simulations play an essential role in many of these research activities. As the scintillation process, the transport of scintillation photons through the crystal(s), and the conversion of these photons into electronic signals each have a major influence on the detector performance; all of these processes may need to be incorporated in the model to obtain accurate results. In this work the optical and scintillation models of the GEANT4 simulation toolkit are validated by comparing simulations and measurements on monolithic scintillator detectors for high-resolution positron emission tomography (PET). We have furthermore made the GEANT4 optical models available within the user-friendly GATE simulation platform (as of version 3.0). It is shown how the necessary optical input parameters can be determined with sufficient accuracy. The results show that the optical physics models of GATE/GEANT4 enable accurate prediction of the spatial and energy resolution of monolithic scintillator PET detectors.

In the field of biomedical X-ray imaging, novel techniques, such as phase-contrast and dark-field imaging, have the potential to enhance the contrast and provide complementary structural information about a specimen. In this paper, a first prototype of a preclinical X-ray phase-contrast CT scanner based on a Talbot-Lau interferometer is characterized. We present a study of the contrast-to-noise ratios for attenuation and phase-contrast images acquired with the prototype scanner. The shown results are based on a series of projection images and tomographic data sets of a plastic phantom in phase and attenuation-contrast recorded with varying acquisition settings. Subsequently, the signal and noise distribution of different regions in the phantom were determined. We present a novel method for estimation of contrast-to-noise ratios for projection images based on the cylindrical geometry of the phantom. Analytical functions, representing the expected signal in phase and attenuation-contrast for a circular object, are fitted to individual line profiles of the projection data. The free parameter of the fit function is used to estimate the contrast and the goodness of the fit is determined to assess the noise in the respective signal. The results depict the dependence of the contrast-to-noise ratios on the applied source voltages, the number of steps of the phase stepping routine, and the exposure times for an individual step. Moreover, the influence of the number of projection angles on the image quality of CT slices is investigated. Finally, the implications for future imaging purposes with the scanner are discussed.

Monolithic scintillator blocks provide an efficient way to significantly increase the sensitivity of high-resolution positron emission tomography (PET) systems. Using statistical or neural network-based positioning schemes, the measured intrinsic detector resolution for perpendicular incident photons is 1.7 mm full-width at half-maximum (FWHM) for 10 mm thick LSO blocks and 2.0 mm FWHM for 20 mm thick LSO blocks. Because the positioning algorithms determine the incidence position as opposed to the combination of interaction position and interaction depth, a very good parallax correction can be achieved. Energy and detector time resolution were 11.5% FWHM and 1.6 ns FWHM, respectively.

The absence of very small crystal pixels in monolithic scintillation detectors has a number of potential advantages such as higher sensitivity, better energy resolution and continuous coordinates. In such detectors, the photon incidence position on the detector surface is derived from the measured scintillation light distribution using artificial neural networks (NNs). To this end, each detector module has to be position-calibrated by training the NNs. An automated procedure to simultaneously obtain the calibration data to train NNs for all detector modules in a fully assembled PET system has been developed and evaluated on a simulator set-up.

We have compared the performance of PET detector modules using continuous scintillator blocks with modules based on a large number of small crystal pixels. We show that a good spatial resolution can be obtained in undivided scintillators. One of the main advantages of the undivided scintillators is the excellent depth of interaction identification capability. There is almost no degradation of the spatial resolution upto incidence angles of 40° away from normal incidence.

The Crystal Clear Collaboration has designed and is building a high-resolution small animal PET scanner. The design is based on the use of the Hamamatsu R7600-M64 multi-anode photomultiplier tube and a LSO/LuYAP phoswich matrix with one to one coupling between the crystals and the photo-detector. The complete system will have 80 PM tubes in four rings with an inner diameter of 137 mm and an axial field of view of 110 mm. The PM pulses are digitized by free-running ADCs and digital data processing determines the gamma energy, the phoswich layer and even the pulse arrival time. Single gamma interactions are recorded and coincidences are found by software. The gantry allows rotation of the detector modules around the field of view. Simulations, and measurements a 2×4 module test set-up predict a spatial resolution of 1.5 mm in the centre of the field of view and a sensitivity of 5.9% for a point source in the centre of the field of view.

LuPO/sub 4/:Nd and YPO/sub 4/:Nd represent new and very fast potential scintillator materials. For most types of ionizing excitation, the luminescence of these materials is dominated by an emission band whose maximum intensity occurs at about 192 nm. The origin of this band lies in the fast 5d-4f transitions of the Nd/sup 3+/ ions. An additional contribution due to host-defect emission varies for different samples. These defects appear to be responsible for the limited light yield of LuPO/sub 4/:Nd.

Detailed results are given for the study of the spatial resolution obtained with different configurations of lutetium oxyorthosilicate (LSO) scintillation crystals read out by a Hamamatsu 2 /spl times/ 8 avalanche photodiode (APD) array. Block detectors as well as small-sized individual LSO crystals have been coupled to the APD matrix. Both one-to-one coupling and light-sharing schemes were investigated. The spatial resolution is derived for each configuration. Two different configurations with two layers of LSO crystals shifted by half a crystal size with respect to each other are used for investigations on the depth-of-interaction information. Prior to the spatial resolution measurements, a comparison of the light output for various crystal surface treatments and different optical coupling materials has been made.

The grating based approach to phase contrast imaging is rather inefficient in the use of the available x-ray flux due to the presence of two absorption gratings and it requires longer scan times compared to conventional CT because multiple images are needed at each projection angle. To avoid these drawbacks, a proof-of-principle experiment was developed to obtain absorption, phase contrast (DPC) and dark field images (DCI) in a single exposure using only a non-absorbing phase grating, a micro-focus source in cone-beam geometry and a highresolution x-ray detector.

A 3D small animal PET scanner, designed and built at the University of Brussel, has been operational since the beginning of 1996. The scanner has a transaxial field of view (FOV) of 110 mm and an axial FOV of 52 mm. The absolute sensitivity is around 35,000 coincidences s−1 MBq−1 for a point source at the center of the scanner and the time resolution is 29 ns FWHM. To measure the achievable spatial resolution, a thin 22Na source was scanned at various distances from the scanner axis. The resolution in a reconstructed image for a source close to the center is 3.0 mm FWHM. This figure can be improved at the expense of the sensitivity by lowering the voltage on the anodes in the wire chamber. Finally, to assess the overall image quality, scans were made of a cylindrical phantom with holes of varying diameter.

We have developed a compact grating-based in-vivo phase-contrast micro-CT system with a rotating gantry. The 50 W microfocus x-ray source operates with 20 to 50 kV peak energy. The length of the rotating interferometer is around 47 cm. Pixel size in the object is 30 micron; the field of view is approx. 35 mm in diameter, suited to image a mouse. The interferometer consists of three gratings: an absorption grating close to the x-ray source, a phase grating to introduce a π/2 phase shift and an absorption analyzer grating positioned at the first fractional Talbot distance. Numerous drives and actuators are used to provide angular and linear grating alignment, phase stepping and object/gantry precision positioning. Phantom studies were conducted to investigate performance, accuracy and stability of the scanner. In particular, the influences of gantry rotation and of temperature fluctuations on the interferometric image acquisition were characterized. Also dose measurements were performed. The first imaging results obtained with the system show the complementary nature of phase-contrast micro-CT images with respect to absorption-based micro-CT. Future improvements, necessary to optimize the scanner for in-vivo small-animal CT scanning on a regular and easy-to-use basis, are also discussed.

To explore the future clinical potential of improved soft-tissue visibility with grating-based X-ray phase contrast (PC), we have developed a first preclinical computed tomography (CT) scanner featuring a rotating gantry. The main challenge in the transition from previous bench-top systems to a preclinical scanner are phase artifacts that are caused by minimal changes in the grating alignment during gantry rotation. In this paper, we present the first experimental results from the system together with an adaptive phase recovery method that corrects for these phase artifacts. Using this method, we show that the scanner can recover quantitatively accurate Hounsfield units in attenuation and phase. Moreover, we present a first tomography scan of biological tissue with complementary information in attenuation and phase contrast. The present study hence demonstrates the feasibility of grating-based phase contrast with a rotating gantry for the first time and paves the way for future in vivo studies on small animal disease models (in the mid-term future) and human diagnostics applications (in the long-term future).

After successful demonstrations of soft-tissue phase-contrast imaging with grating interferometers at synchrotron radiation sources and at laboratory based x-ray tubes, a first preclinical CT scanner with grating based phase contrast imaging modality has been constructed. The rotating gantry is equipped with a three-grating interferometer, a 50 watt tungsten anode source and a Hamamatsu flat panel detector. The total length of the interferometer is 45 cm, and the bed of the scanner is optimized for mice, with a scanning diameter of 35 mm. From one single scan both phase-contrast and standard attenuation based tomography can be attained, providing an overall gain in image contrast.

Item does not contain fulltext Cramér-Rao theory can be used to derive the lower bound on the spatial resolution achievable with position-sensitive scintillation detectors as a function of the detector geometry and the pertinent physical properties of the scintillator, the photosensor and the readout electronics. Knowledge of the Cramér-Rao lower bound (CRLB) can for example be used to optimize the detector design and to test the performance of the method used to derive position information from the detector signals. Here, this approach is demonstrated for monolithic scintillator detectors for positron emission tomography. Two detector geometries are investigated: a 20 × 10 × 10 mm(3) and a 20 × 10 × 20 mm(3) monolithic LYSO:Ce(3+) crystal read out by one or two Hamamatsu S8550SPL avalanche photodiode (APD) arrays, respectively. The results indicate that in these detectors the CRLB is primarily determined by the APD excess noise factor and the number of scintillation photons detected. Furthermore, it is shown that the use of a k-nearest neighbor (k-NN) algorithm for position estimation allows the experimentally obtained spatial resolution to closely approach the CRLB. The approach outlined in this work can in principle be applied to any scintillation detector in which position information is encoded in the distribution of the scintillation light over multiple photosensor elements.

This paper presents a study of possible models to describe the relation between the scintillation light point-of-origin and the measured photo detector pixel signals in monolithic scintillation crystals. From these models the X, Y and depth of interaction (DOI) coordinates can be estimated simultaneously by nonlinear least-square fitting. The method depends only on the information embedded in the signals of individual events, and therefore does not need any prior position training or calibration. Three possible distributions of the light sources were evaluated: an exact solid-angle-based distribution, an approximate solid-angle distribution and an extended approximate solid-angle-based distribution which includes internal reflection at side and bottom surfaces. The performance of the general model using these three distributions was studied using Monte Carlo simulated data of a 20 × 20 × 10 mm lutetium oxyorthosilicate (Lu 2 SiO 5 or LSO) block read out by 2 Hamamatsu S8550 avalanche photo diode arrays. The approximate solid-angle-based model had the best compromise between resolution and simplicity. This model was also evaluated using experimental data by positioning a narrow 1.2 mm full width at half maximum (FWHM) beam of 511 keV photons at known positions on the 20 × 20 × 10 mm LSO block. An average intrinsic resolution in the X-direction of 1.4 mm FWHM was obtained for positions covering the complete block. The intrinsic DOI resolution was estimated at 2.6 mm FWHM.

Proper selection of modern key components allows eliminating most artifacts in micro-CT and nano-CT systems already during data acquisition. X-ray cameras with direct photon detection allow avoiding ring artifacts. Newly developed fully depleted CCD sensors show an energy response similar to traditional cameras with a thin scintillator, but without any geometrical distortions and flashes from x-ray photons penetrating through the fiber optics. Air-bearing rotation stages and piezo-positioning minimizes mechanical inaccuracies in acquiring angular projections. Beam hardening can be eliminated by energy-selective photon counting imaging.

An integrated microCT/microXRF system has been designed and built at SkyScan. The two sub-systems are aligned. The microCT provides 3D morphological information of the sample, which can be also used for attenuation correction during microXRF reconstruction. The microXRF, based on a pin-hole collimator and a photon-counting energy-sensitive 2Ddetector, obtains 2D projections of 3D chemical composition inside the sample with 50-70 microns spatial resolution. The reconstruction of 3D microXRF scans is challenging because of very low photon counting statistics due to limited power of laboratory x-ray sources and the strong self-absorption of the low-energy fluorescence photons. We have developed a maximum-likelihood expectation-maximization (ML-EM) algorithm based on Poisson model. This algorithm has proven to be rather robust and good reconstructions have been obtained with sample scans. Regularization is necessary to achieve stable reconstruction. One method is to apply smoothing between iterations. Two different smoothing kernels have been evaluated: 3D symmetric Gaussian kernel and minimization of total variation. For further improvement, a multi-ray resolution recovery technique has been evaluated. The self-absorption is currently compensated by a simplified method: the correction coefficients are pre-calculated and obtained by forward-projecting the attenuation map for both the primary X-rays and the fluorescence photons. The attenuation maps at the energy of fluorescence photons are approximated from the CT image.

X-ray fluorescence (XRF) allows imaging of the chemical composition of a specimen. We developed a 2 nd generation prototype laboratory system that can produce 3D chemical maps using microXRF as well as volumetric microCT images. The latter can be used to overlay morphological information on top of the XRF image for co-registration. It is also employed for attenuation correction during the tomographic reconstruction of the XRF images. The new system has various hardware and software changes to improve the performance, stability and flexibility. A deep depleted CCD was employed to improve the detection efficiency for high-energy fluorescence X rays. The use of a deep depleted CCD requires signal-clustering techniques to correct for charge diffusion in the CCD to obtain the correct energy of the fluorescence x rays. Furthermore, energy drift correction techniques were put in place to ensure stability of energy measurement during very long scan times. To minimize the contribution of the long CCD readout times to the total scan time, the exposure frames are dynamically adjust during the scan to the maximum time allowed for operation under photon counting mode. The XRF component has a spatial resolution of 70 μm and an energy resolution of 180 eV at 6.4 keV.

Most X-ray systems are limited in spatial resolution by the x-ray source performance. In laboratory sources, x-rays are generated by the interaction of an electron beam with a metal target. Bulk target sources produce a spot size in the micron range. Thin layer targets allow a spot size improvement down to hundreds of nanometers, but with a significant flux reduction. Until now a spatial resolution under 100 nm could only be achieved by imaging with Fresnel zone plates with limited depth of focus, typically - several microns. This is acceptable for imaging of flat objects, but it creates a problem for tomography, which requires all parts of a bulk object to be in focus. To overcome the limitations, we invented an x-ray source with a new type of target. Because x-ray cameras can only collect photons from a small angle, the new emitter is physically shaped in such way that the camera can see it as a small dot, but it has a big length along the direction perpendicular to the camera creating a significant flux without compromising the resolution. Evaluation shows that structures down to 50 nm can be distinguished while maintaining a significant x-ray flux and infinite depth of focus required for nano-tomographical reconstruction.

The monolithic scintillator block approach for gamma detection in the Positron Emission Tomography (PET) avoids estimating Depth of Interaction (DOI), reduces dead zones in detector and diminishes costs of detector production. Neural Networks (NN) are very efficient to determine the entrance point of a gamma incident on a scintillator block. This paper presents results on the robustness of the spatial resolution as a function of the random fraction in the data, temperature and HV fluctuations. This is important when implementing the method in a real scanner. Measurements were done with two Hamamatsu S8550 APD arrays, glued on a 20×20×10 mm3 monolithic LSO crystal block.

Vsing a nano-focus x-ray source we have been able to visualize structures down to 150 nm in x-ray projection images. Because of their unlimited depth of focus, these sources don't restrict the specimen size in case of 3D tomographic imaging. Simulation studies have shown that optimization of the detector response curve and switching from a reflective x-ray target to a transmission target might allow us to reach sub-100 nm resolutions.

We developed positron emission tomography (PET) detectors based on monolithic scintillation crystals and position-sensitive light sensors. Intrinsic depth-of-interaction (DOI) correction is achieved by deriving the entry points of annihilation photons on the front surface of the crystal from the light sensor signals. Here we characterize the next generation of these detectors, consisting of a 20 mm thick rectangular or trapezoidal LYSO:Ce crystal read out on the front and the back (double-sided readout, DSR) by Hamamatsu S8550SPL avalanche photodiode (APD) arrays optimized for DSR. The full width at half maximum (FWHM) of the detector point-spread function (PSF) obtained with a rectangular crystal at normal incidence equals approximately 1.05 mm at the detector centre, after correction for the approximately 0.9 mm diameter test beam of annihilation photons. Resolution losses of several tenths of a mm occur near the crystal edges. Furthermore, trapezoidal crystals perform almost equally well as rectangular ones, while improving system sensitivity. Due to the highly accurate DOI correction of all detectors, the spatial resolution remains essentially constant for angles of incidence of up to at least 30 degrees . Energy resolutions of approximately 11% FWHM are measured, with a fraction of events of up to 75% in the full-energy peak. The coincidence timing resolution is estimated to be 2.8 ns FWHM. The good spatial, energy and timing resolutions, together with the excellent DOI correction and high detection efficiency of our detectors, are expected to facilitate high and uniform PET system resolution.

This work is part of the design study for a medical positron emission tomography (PET) system based on the use of BaF 2 scintillator and photosensitive wire chambers with TMAE. The number of photoelectrons per MEV (phe/MeV) is measured in the wire chamber and on a photomultiplier for crystals from five different suppliers. For the best crystals 26.7±2.0 phe/MeV in TMAE are observed, corresponding to an average quantum efficiency of TMAE for the fast component of BaF 2 of (7.6±0.7)%. The performance of a PET scanner based on this principle was evaluated with a detailed Monte Carlo simulation. It is shown that the reduced energy resolution with wire chambers has only a marginal effect on the scatter rejection in the scanner. Finally LaF 3 :Nd 3+ as an alternative to BaF 2 was tested and found to give a disappointing photoelectron yield.

The optimal pixel size of SiPMT arrays to read out monolithic 20x20x10 mm LSO scintillator block detectors has been studied. The influence of different SiPMT parameters on the achievable spatial resolution was compared. It was concluded that a SiPMT array with 3*3 mm pixel size, pitch 3.4 mm pixel, 30% detection efficiency, 100 ìm *100 ìm micro-cell size and crosstalk smaller than 20% is a good and realistic design that achieves better spatial and energy resolution than our current monolithic scintillator block readout based on Hamamatsu S8550 APD arrays.

We are presently developing a novel PET scanner for human brain functional imaging based on monolithic scintillator crystals read by APD matrices, capable of being inserted into an MRI system. In this work we report on the detailed study that has been made of the design of the detector blocks, aiming at defining the most suitable geometrical and readout configuration for optimizing the overall performance of the entire scanner. Both parallel and trapezoidal geometries have been simulated, featuring two layers of active scintillator material with different or similar thickness and APD readout on the front or back side. Results of this study indicate that a trapezoidal geometry with equal thickness of both layers is the best solution for improving the expected scanner performance.

The performance of a small-animal positron emission tomography (PET) scanner based on monolithic scintillation detectors read-out by avalanche photo-diode arrays has been investigated by simulation. By minimizing dead space, both within and between the modules, these detectors offer increased detection efficiency compared to pixellated detectors. The spatial resolution of the scanner was investigated in 2-D by simulating a point source at various radial distances from the center. To model the detector response, measured detector line-spread functions were used. An optimum value of approximately 1 mm FWHM was found at 10 mm radial distance from the scanner central axis. Point-source sensitivity profiles in the radial and axial directions were simulated at 1 MBq activity using the Monte-Carlo code GATE. They indicated that monolithic designs increase the sensitivity roughly by a factor of two compared to pixellated designs. NECR curves simulated for these scanner designs show no significant degradation of the performance for activities up to 40 MBq.

Monolithic scintillation detectors are very promising for high resolution and high sensitivity positron emission tomography. These detectors consist of a few cubic centimeters of scintillating material coupled to one or more position-sensitive APD arrays. The entry point of an impinging annihilation photon is estimated from the light distribution on the APD pixels. In this paper, a model will be derived that predicts the line spread function of these detectors. This model includes the influences of the finite width of the measurement beam, scatter of radiation inside the detector, light yield and intrinsic energy resolution of the scintillator, quantum efficiency, gain and excess noise factor of the avalanche photo-diode arrays, and electronic noise. With this model a better understanding of the behavior of the detector is possible, which can make optimization more efficient.

Recent years have seen an increased interest in high resolution, high sensitivity detectors for small animal positron emission tomography (PET). Detectors based on large (e.g. several cm/sup 3/ of LSO) monolithic scintillators optically coupled to one or more avalanche photodiode (APD) array(s) minimize loss of sensitivity due to dead space. In such detectors, the primary event position and depth of interaction (DOI) correction are derived from the measured distribution of scintillation light over the APD array(s). The performance of detectors consisting of a monolithic LYSO scintillator read out by two Hamamatsu S8550 APD arrays is investigated. Monte Carlo simulations indicate that the scanner sensitivity and uniformity are improved when trapezoidal crystals are used instead of rectangular ones. The performance of both types of detectors was studied. Experiments show that the influence of the angle of incidence on the spatial resolution of both detector types is small, ranging from /spl sim/2.0 mm FWHM at 0/spl deg/ to /spl sim/2.4 mm FWHM at 30/spl deg/. An energy resolution of 11.5% at 511 keV and a time resolution of 2.4 ns was measured. Monte Carlo simulations predict a sensitivity of /spl sim/21% for a point source at the center of the field of view for a four-ring scanner based on trapezoidal monolithic detectors.

In this work results of ongoing research are presented for the LYSO/LuYAP phoswich solution adapted in the ClearPETtrade scanner with depth-of-interaction capability. The use of artificial neural network has been found as a very efficient phoswich crystal layer identification method that works for scintillators used at present, as well as on prospective materials. The LYSO/LuAP and LYSO/GSO phoswiches have been identified as a promising candidates for future PET scanners

Minimizing dead space is one way to increase the detection efficiency of small-animal PET scanners. By using monolithic scintillator crystals (e.g., 20 mm/spl times/10 mm/spl times/10 mm LSO), loss of efficiency due to inter-crystal reflective material is minimized. Readout of such crystals can be performed by means of one or more avalanche photo-diode (APD) arrays optically coupled to the crystal. The entry point of a gamma photon on the crystal surface can be estimated from the measured distribution of the scintillation light over the APD array(s). By estimating the entry point, correction for the depth-of-interaction (DOI) is automatically provided. We are studying the feasibility of such detector modules. To this end, a 64-channel test setup has been developed. Experiments to determine the effect on the spatial resolution of crystal surface finish and detector geometry have been carried out. The first results of these experiments are presented and compared to simulation results. The crystal surface finish has only a small influence on the spatial resolution. The spatial resolution of 20 mm/spl times/10 mm/spl times/10 mm detectors is significantly better when read out on the front side than when read out on the back side. With a 20 mm/spl times/10 mm/spl times/20 mm crystal coupled to two APD arrays, a very small resolution degradation of only /spl sim/0.2 mm is observed for an incidence angle of 30/spl deg/ compared to normal incidence.

This paper describes the LSO/LuYAP phoswich detector head developed for the ClearPET small animal PET scanner demonstrator that is under construction in Lausanne within the Crystal Clear Collaboration. The detector head consists of a dual layer of 8/spl times/8 LSO and LuYAP crystal arrays coupled to a multi-anode photomultiplier tube (Hamamatsu R7600-M64). Equalistion of the LSO/LuYAP light collection is obtained through partial attenuation of the LSO scintillation light using a thin aluminum deposit of 20-35 nm on LSO and appropriate temperature regulation of the phoswich head between 30/spl deg/C to 60/spl deg/C. At 511keV, typical FWHM energy resolutions of the pixels of a phoswich head amounts to (28/spl plusmn/2)% for LSO and (25/spl plusmn/2)% for LuYAP. The LSO versus LuYAP crystal identification efficiency is better than 98%. Six detector modules have been mounted on a rotating gantry. Axial and tangential spatial resolutions were measured up to 4 cm from the scanner axis and compared to Monte Carlo simulations using GATE. FWHM spatial resolution ranges from 1.3 mm on axis to 2.6 mm at 4 cm from the axis.

The Crystal Clear Collaboration has designed and built a family of high resolution small animal PET scanners. These were designed to be used in research laboratories and provide maximum modularity and flexibility. The source code of the data acquisition and reconstruction software is freely available to the users. The design is based on the use of the Hamamatsu R7600-M64 multi-anode photomultiplier tube and an LSO/LuYAP phoswich matrix with one-to-one coupling between the crystals and the photo-detector. A complete system has 80 PMT tubes in four rings with a minimum inner diameter of 137 mm and an axial field of view of 110 mm. The detectors are rotating over 360 degrees so that partially filled ring geometries can be used. This greatly simplifies the combination of PET with other imaging modalities. Single gamma interactions are recorded in list mode format and coincidences are found by software.

Monolithic scintillator detectors, consisting of several cm3 of scintillating material coupled to one or more Hamamatsu S8550 avalanche photodiode (APD) arrays, have been proposed as detectors for high resolution positron emission tomography (PET). These detectors eliminate the dead space caused by reflective material between crystal pixels, and provide good depth of interaction (DOI) information. The entry point of impinging annihilation photons on the front surface of the detector is derived from the scintillation light distribution on the APD array(s). In this work, the various contributions to the energy and timing resolutions were analysed, as well as the influence of several contributions to the spatial resolution. Energy resolutions of around 11.5% FWHM at 511 keV were measured, the major part of which was found to be due to the intrinsic photon variance of the scintillator. A timing resolution of 1.6 ns FWHM was measured against a BaF2-PMT detector. The timing resolution was found to be strongly gain dependent. Spatial resolutions of around 1.75 mm FWHM were measured, fairly independent of the APD gain. This figure still needs to be corrected for the ~0.9 mm FWHM measurement beam.

PET-detectors based on monolithic scintillator blocks have the potential to significantly improve the sensitivity. The impinging photon position of the 511 keV photons on the detector surface is derived from the scintillation light distribution measured by an APD array. To extract this position information, we used Neural Networks (NN). To this end, each detector module has to be position-calibrated by training neural networks. The neural networks can immediately yield a parallax insensitive incidence position of the photon when it was trained with sample data of photons impinging under a similar angle. Different procedures to obtain the calibration data in an automated way on a fully assembled PET system have been developed and validated on a simulator set-up.

Contents. Preface S. Tavernier.-A Look at Medical Imaging Trends through the Eyes of a Medical Doctor S.S. Makeyev.- Introduction.-Historical Aspect of Nuclear Medicine.-Nowadays in Nuclear Medicine.-Perspectives of Nuclear Medicine Imaging.- New Trends in X-Ray CT Imaging R. Deych and E. Dolazza.- Present Status of X-Ray CT.-Detector Instrumentation in Medical CT.- Scintillator.-Photodetectors.-Future Evolution of Data Measurement Systems.- The Evolution of Spect- from Anger to Today and Beyond W.W. Moses, A. Gektin et al.- Introduction.-General Considerations.-SPECT.- The Anger Camera.-Optimizing Positioning in Anger Cameras.- Collimators.-Scintillators for Spect.- Recently Developed Scintillator Materials.- Conclusion.- New Trends in PET Detector Developments P.Lecoq.- Introduction.-PET Based Molecular Imaging.-Improving Sensitivity.- Improving Spatial and Temporal Resolution.-Multimodaility and Multifunctionality.-New Conversion Materials.- New Photodetectors.-Highly Integrated and Low Noise Electronics.-Intelligent and Triggerable Data Acquisition Systems.-Simulation Software.-New Reconstruction and Visualisation Algorithms.-Conclusion.-Semiconductor Detectors in Radiation Medicine: Radiotherapy and related Applcations A.B. Rosenfeld.- Introduction.-Integral Semiconductor Dosimetry in Radiation Therapy.-Mosfet Detectors.-Semiconductor Radiation Detectors in Hadron Therapy.- Semiconductor Radiation Detectors for Microdosimetry in Radiation Therapy.-Application of Scintillator Based Detector in Radiation Therapy.-Conclusion.-First Results with the ClearPET small Animal PET Scanners S. Tavernier et al.- Introduction.-Description of the ClearPET Scanners.-Measured Performance and Comparison with Monte Carlo Simulations.- Image Reconstruction.-Conclusions.-Investigation of Crystal Identification Methods for ClearPETTM Phoswich Detector D. Wisniewski et al.- Introduction.-Measurement Setup.-Crystal Identification Methods.- Experimental Results.- Conclusions.- Directions in Scintillation Materials Research P. Dorenbos.- Introduction.-Historic Developments.- Fundamental Limits.- Directions in Scintillation Materials Research.-Summary and Conclusions.-Scintillation Detectors for Medical and Biology Applications: Materials, Design and Light Collection Conditions M. Globus, B. Grinyov.- Introduction.-2. Some Features and Regularities of Light Collection in Scintillators.- Medical Diagnostics Instrumentation.- Thin Scintillation Films for Biological Microtomography. Conclusions.- Current and Future Use of LSO: CE Scintillators in PET C.L. Melcher et al.- Introduction.-Physical Properties.- Scintillation Properties.-Crystal Growth.-Detector Design.- Future Uses of LSO: CE in PET.-Conclusion.-Inorganic Scintillators in Positron Emission Tomography C.W.E. van Eijk.- Introduction.-Inorganic Scintillators.- Position Resolution and Depth of Interaction.-Coincidence-Time Resolution, Random Coincidences, Time of Flight and Dead Time.-Conclusion.-Crystal Fibers and thin Films for Imaging Applications C. Pedrini and C. Dujardin.-. Introduction.-Single Crystal Fibers.- Scintillating Thin Films Deposited on Substrate.- Scintillation thin Layers created by Irradiation.-Conclusions. Non-Proportionality and Energy Resolution of Scintillation Detectors M. Moszynski.-Introduction.-Outline of the Problem.Study of Energy Resolution and Non-Proportionality.- Discussion and Conclusions.

We are investigating the possibility of using monolithic scintillator blocks as detectors for small animal positron emission tomography (PET). These detectors consist of several cm/sup 3/ of scintillating material, read out by one or more avalanche photo-diode (APD) arrays. The entry point of an incoming gamma photon is estimated from the distribution of the scintillation light over the APD pixels. To optimize the detector design, the influence of different design parameters is investigated using Geant4 simulations. To make it possible to study the influence of individual design parameters on the intrinsic spatial resolution of the detector, the use of a performance measure is proposed that is independent of the algorithm used to estimate the entry point, namely the Cramer-Rao lower bound on the estimation of the coordinates of a point source of light inside the crystal. To illustrate the use of this method, the influence of optical transport inside the detector is investigated for different detector designs, surface finishes and APD pixel sizes. A comparison with resolutions obtained from simulations involving beams of 511 keV annihilation photons indicates that this approach gives valid results.

The paper describes the LSO/LuYAP phoswich detector head developed for the ClearPET small animal PET scanner demonstrator that is under construction within the Crystal Clear Collaboration. The detector head consists of a dual layer of 8/spl times/8 LSO and LuYAP crystal arrays coupled to a multi-channel photomultiplier tube (Hamamatsu R7600-M64). Equalisation of the LSO/LuYAP light collection is obtained through partial attenuation of the LSO scintillation light using a thin aluminium deposit of 20-35 nm on LSO and appropriate temperature regulation of the phoswich head between 30 to 60 /spl deg/C. At 511 keV, typical FWHM energy resolutions of the pixels of a phoswich head amounts to (28/spl plusmn/2)% for LSO and (25/spl plusmn/2)% for LuYAP. The LSO versus LuYAP crystal identification efficiency is better than 98%. Six detector modules have been mounted on a rotating gantry. Axial and tangential spatial resolutions were measured up to 4 cm from the scanner axis and compared to Monte Carlo simulations using GATE. FWHM spatial resolution ranges from 1.3 mm on axis to 2.6 mm at 4 cm from the axis.

Neutrinos, Imagerie; Second generation high-performance PET scanners, called ClearPET$^{TM}$, have been developed by working groups of the Crystal Clear Collaboration (CCC). High sensitivity and high spatial resolution for the ClearPET camera is achieved by using a phoswich arrangement combining two different types of lutetium-based scintillator materials: LSO from CTI and LuYAP:Ce from the CCC (ISTC project). In a first ClearPET prototype, phoswich arrangements of 8$\times$8 crystals of 2$\times$2$\times$10 mm$^3$ are coupled to multi-channel photomultiplier tubes (Hamamatsu R7600). A unit of four PMTs arranged in-line represents one of 20 sectors of the ring design. The opening diameter of the ring is 120 mm, the axial detector length is 110 mm. The PMT pulses are digitized by free-running ADCs and digital data processing determines the gamma energy, the phoswich layer and even the exact pulse starting time, which is subsequently used for coincidence detection. The gantry allows rotation of the detector modules around the field of view. Preliminary data shows a correct identification of the crystal layer about (98$\pm$1)%. Typically the energy resolution is (23.3$\pm$0.5)% for the luyap layer and (15.4$\pm$0.4)% for the lso layer. Early studies showed the timing resolution of 2 ns FWHM and 4.8 ns FWTM. the intrinsic spatial resolution ranges from 1.37 mm to 1.61 mm full-width of half-maximum (FWHM) with a mean of 1.48 mm FWHM. further improvements in image and energy resolution are expected when the system geometry is fully modeled.

A 2nd generation high performance small animal PET scanner, called ClearPET/spl trade/, has been designed and a first prototype is built by working groups of the Crystal Clear Collaboration (CCC). In order to achieve high sensitivity and maintain good uniform spatial resolution over the field of view in high resolution PET systems, it is necessary to extract the depth of interaction (DOI) information and correct for spatial degradation. The design of the first ClearPET/spl trade/ Demonstrator based on the use of the multi-anode photomultiplier tube (Hamamatsu R7600-M64) and a LSO/LuYAP phoswich matrix. The two crystal layers of 8*8 crystals (2*2*10 mm/sup 3/) are stacked on each other and mounted without light guide as one to one on the PMT. A unit of four PMTs arranged in-line represents one of 20 sectors of the ring design. The opening diameter of the crystal ring is 137 mm, the axial detector length is 110 mm. The PMT pulses are digitized by free-running ADCs and digital data processing determines the gamma energy, the phoswich layer and even the pulse arrival time. Single gamma interactions are recorded and coincidences are found by software. The gantry allows rotation of the detector modules around the field of view. The measurements have been done using the first LSO/LuYAP detector cassettes.

A 3D small animal PET scanner, designed and built at the University of Brussels, has been operational since the beginning of 1996. The scanner has a transaxial field of view (FOV) of 110 mm and an axial FOV of 51 mm. The absolute sensitivity is around 35.000 coincidences/s/MBq for a point source at the center of the scanner and the time resolution is 29 ns FWHM. To measure the achievable spatial resolution, a thin /sup 22/Na source was scanned at various distances from the scanner axis. The resolution in a reconstructed image for a source close to the center is 3.0 mm FWHM. Finally, to assess the overall image quality, scans were made of a cylindrical phantom with holes of varying diameter.

We have obtained optically clear samples of LuAP:Ce from the Institute of Physical Research, Ashtarak, Armenia, measuring up to 5/spl times/5/spl times/40 mm/sup 3/. The photo electron yield for a number of these samples was measured with a Philips XP2020Q photomultiplier tube, and 556 photo electrons per MeV was obtained for the best samples. This is comparable to what is obtained from BGO crystals under the same condition, but a factor of 5 less than what is obtained with an LSO sample of similar dimensions. The light in LuAP:Ce increases with cerium concentration between 0.13 and 0.47%.

The authors have designed and built a small high resolution PET scanner using BaF/sub 2/ scintillation crystals and photosensitive wire chambers using TMAE as photosensitive agent. The very first measurements performed on the detector are presented. >

The LuAlO/sub 3/:Ce/sup 3+/ system was shown to be a promising scintillator for medical imaging devices. Recently, efforts were focused on the improvement of its scintillating properties. Several large size crystals with various cerium concentrations were grown. Emission spectra, fluorescence decay times and light yields, both under /spl gamma/-ray and X-ray excitation, were measured under various experimental conditions. A reabsorption process is shown to take place in this material. This process is responsible for the observed decrease of the light yield when increasing the size of the sample.

The possibility of extracting depth-of-interaction (DOI) has been studied for block detector configurations, consisting of either a block of LSO or a saw-shaped crystal coupled to a Hamamatsu Si-avalanche photodiode (APD) array 1370. The results of this study, as well as previous results on DOI, obtained with individual crystal configurations consisting of two layers of lutetium oxyorthosilicate (LSO) crystals shifted over half a crystal width with respect to each other, are applied in corrections on the spatial resolution obtained with these Positron Emission Tomography (PET) detector modules. Further improvement in the composition of crystal matrices has been realized. The energy resolution and the light yield for crystals in such a crystal matrix have been evaluated.