Positron emission tomography (PET) and magnetic resonance imaging (MRI) are widely used in vivo imaging technologies with both clinical and biomedical research applications. The strengths of MRI include high-resolution, high-contrast morphologic imaging of soft tissues; the ability to image physiologic parameters such as diffusion and changes in oxygenation level resulting from neuronal stimulation; and the measurement of metabolites using chemical shift imaging. PET images the distribution of biologically targeted radiotracers with high sensitivity, but images generally lack anatomic context and are of lower spatial resolution. Integration of these technologies permits the acquisition of temporally correlated data showing the distribution of PET radiotracers and MRI contrast agents or MR-detectable metabolites, with registration to the underlying anatomy. An MRI-compatible PET scanner has been built for biomedical research applications that allows data from both modalities to be acquired simultaneously. Experiments demonstrate no effect of the MRI system on the spatial resolution of the PET system and <10% reduction in the fraction of radioactive decay events detected by the PET scanner inside the MRI. The signal-to-noise ratio and uniformity of the MR images, with the exception of one particular pulse sequence, were little affected by the presence of the PET scanner. In vivo simultaneous PET and MRI studies were performed in mice. Proof-of-principle in vivo MR spectroscopy and functional MRI experiments were also demonstrated with the combined scanner. molecular imaging ͉ small animal imaging ͉ multimodality imaging P ositron emission tomography (PET) noninvasively images the distribution in vivo of biomolecules (small molecules, peptides, antibodies, and nanoparticles) labeled with radionuclides that undergo positron decay and produce back-to-back 511-keV annihilation photons (1). Because of the high sensitivity of radioactive assays, PET can measure picomolar concentrations of labeled biomolecules. A wide variety of molecular targets and pathways have been imaged by using PET radiotracers (2, 3), with the avid accumulation of the radiotracer [ 18 F]-2-fluoro-2-deoxy-D-gluocse (FDG) in malignant tumors being just one example that has widespread applications in the clinic and in the study of therapeutic strategies for tumor treatment in animal models. However, the spatial resolution of PET is limited by physical factors associated with positron physics and by the difficulty of acquiring sufficient counting statistics. Furthermore, PET images often lack definitive anatomic information, making interpretation of the precise location of radiotracer accumulation difficult.Magnetic resonance imaging (MRI) can provide high-spatialresolution anatomic images with exquisite soft-tissue contrast by exploiting the differences in relaxation times of protons in different biochemical environments (4, 5). The combination of high spatial resolution and contrast allows the anatomic consequences (e.g., tumor growth, brain atroph...

Detectors with depth-encoding allow a PET scanner to simultaneously achieve high sensitivity and high spatial resolution. Methods: A prototype PET scanner, consisting of depth-encoding detectors constructed by dual-ended readout of lutetium oxyorthosilicate (LSO) arrays with 2 position-sensitive avalanche photodiodes (PSAPDs), was developed. The scanner comprised 2 detector plates, each with 4 detector modules, and the LSO arrays consisted of 7 · 7 elements, with a crystal size of 0.9225 · 0.9225 · 20 mm and a pitch of 1.0 mm. The active area of the PSAPDs was 8 · 8 mm. The performance of individual detector modules was characterized. A line-source phantom and a hot-rod phantom were imaged on the prototype scanner in 2 different scanner configurations. The images were reconstructed using 20, 10, 5, 2, and 1 depth-of-interaction (DOI) bins to demonstrate the effects of DOI resolution on reconstructed image resolution and visual image quality. Results: The flood histograms measured from the sum of both PSAPD signals were only weakly depth-dependent, and excellent crystal identification was obtained at all depths. The flood histograms improved as the detector temperature decreased. DOI resolution and energy resolution improved significantly as the temperature decreased from 20°C to 10°C but improved only slightly with a subsequent temperature decrease to 0°C. A full width at half maximum (FWHM) DOI resolution of 2 mm and an FWHM energy resolution of 15% were obtained at a temperature of 10°C. Phantom studies showed that DOI measurements significantly improved the reconstructed image resolution. In the first scanner configuration (parallel detector planes), the image resolution at the center of the field of view was 0.9-mm FWHM with 20 DOI bins and 1.6-mm FWHM with 1 DOI bin. In the second scanner configuration (detector planes at a 40°angle), the image resolution at the center of the field of view was 1.0-mm FWHM with 20 DOI bins and was not measurable when using only 1 bin. Conclusion: PET scanners based on this detector design offer the prospect of high and uniform spatial resolution (crystal size, ;1 mm; DOI resolution, ;2 mm), high sensitivity (20-mm-thick detectors), and compact size (DOI encoding permits detectors to be tightly packed around the subject and minimizes number of detectors needed). Over the past decade, many small-animal PET scanners have been developed (1-12), and this technology has played a very important role in the rapidly growing field of molecular imaging. High sensitivity is needed to increase signal-to-noise ratio of the images to reliably detect lower levels of radiotracer uptake and to reduce the injected dose (reducing radiation dose to the subject) (13) and scan time (increasing temporal resolution for dynamic studies). High spatial resolution is required to detect small structures and lesions and to improve quantification by reducing the partialvolume effect. A compromise between sensitivity and spatial resolution due to depth-of-interaction (DOI) effects always exists in small-ani...

A prototype small-animal PET scanner was developed based on depth-encoding detectors using dual-ended readout of very small scintillator elements to produce high and uniform spatial resolution suitable for imaging the mouse brain. Methods The scanner consists of 16 tapered dual-ended readout detectors arranged in a ring of diameter 61 mm. The axial field of view is 7 mm and the transaxial field of view is 30 mm. The scintillator arrays consist of 14×14 lutetium oxyorthosilicate (LSO) elements, with a crystal size of 0.43×0.43 mm2 at the front end and 0.80×0.43 mm2 at the back end, and the crystal elements are 13 mm long. The arrays are read out by 8×8 mm2 and a 13×8 mm2 position-sensitive avalanche photodiodes (PSAPDs) placed at opposite ends of the array. Standard nuclear instrumentation module (NIM) electronics and a custom designed multiplexer are used for signal processing. Results The detector performance was measured and all except the very edge crystals could be clearly resolved. The average detector intrinsic spatial resolution in the axial direction was 0.61 mm. A depth of interaction resolution of 1.7 mm was achieved. The sensitivity of the scanner at center of the field of view was 1.02% for a lower energy threshold of 150 keV and 0.68% for a lower energy threshold of 250 keV. The spatial resolution within a field of view that can accommodate the entire mouse brain was ~0.6 mm using a 3D Maximum Likelihood-Expectation Maximization (ML-EM) reconstruction algorithm. Images of a micro hot-rod phantom showed that rods with diameter down to 0.5 mm could be resolved. First in vivo studies were obtained using 18F-fluoride and confirmed that 0.6 mm resolution can be achieved in the mouse head in vivo. Brain imaging studies with 18F-fluorodeoxyglucose were also acquired. Conclusion A prototype PET scanner achieving a spatial resolution approaching the physical limits for a small-bore PET scanner set by positron range and acolinearity was developed. Future plans are to add more detector rings to extend the axial field of view of the scanner and increase sensitivity.

The impinging position of a 511 keV photon onto a continuous scintillator can be obtained from the light distribution measured by a pixelated photodetector such as avalanche photodiode (APD) arrays. This information is extracted using neural networks trained for events with a particular incidence angle. Using a 20 10 10 mm block of lutetium oxyorthosilicate mounted onto a S8550 Hamamatsu APD matrix we achieved an intrinsic resolution of 1.9 mm full-width at half-maximum (FWHM) for perpendicular incident photons and 2.6 mm FWHM at a 40 incidence angle. A possible implementation for tomographic imaging is presented. Index Terms-Depth of interaction (DOI), neural network (NN), positron emission tomorgraphy (PET).

We have constructed a dedicated breast PET/CT scanner capable of high-resolution functional and anatomic imaging. Here, we present an initial characterization of scanner performance during patient imaging. Methods: The system consisted of a lutetium oxyorthosilicate-based dual-planar head PET camera (crystal size, 3 · 3 · 20 mm) and 768-slice cone-beam CT. The position of the PET heads (separation and height) could be adjusted for varying breast dimensions. For scanning, the patient lay prone on a specialized bed and inserted a single pendent breast through an aperture in the table top. Compression of the breast as used in mammography is not required. PET and CT systems rotate in the coronal plane underneath the patient sequentially to collect fully tomographic datasets. PET images were reconstructed with the fully 3-dimensional maximum a posteriori method, and CT images were reconstructed with the Feldkamp algorithm, then spatially registered and fused for display. Phantom scans were obtained to assess the registration accuracy between PET and CT images and the influence of PET electronics and activity on CT image quality. We imaged 4 women with mammographic findings highly suggestive of breast cancer (breast imaging reporting and data system, category 5) in an ongoing clinical trial. Patients were injected with 18 F-FDG and imaged for 12.5 min per breast. From patient data, noise-equivalent counting rates and the singles-to-trues ratio (a surrogate for the randoms fraction) were calculated. Results: The average registration error between PET and CT images was 0.18 mm. PET electronics and activity did not significantly affect CT image quality. For the patient trial, biopsy-confirmed cancers were visualized on dedicated breast PET/CT on all patient scans, including the detection of ductal carcinoma in situ in 1 case. The singles-to-trues ratio was found to be inversely correlated with breast volume in the field of view, suggesting that larger breasts trend toward increased noiseequivalent counting rates for all other things equal. Conclusion: Scanning of the uncompressed breast with dedicated breast PET/CT can accurately visualize suspected lesions in 3 dimensions. Whol e-body (WB) 18 F-FDG PET has clinical utility in breast cancer staging, restaging, and therapy response assessment. A study by Rousseau et al. (1) found that WB PET could identify tumors with pathologic response after a single course of neoadjuvant chemotherapy (sensitivity, 61%; specificity, 96%), whereas mammography had limited accuracy (sensitivity, 31%; specificity, 56%), even after 6 courses of treatment. WB PET has been shown to have a high accuracy for detecting distant metastasis. Mahner et al.(2) measured a sensitivity and specificity for metastatic disease of 87% and 83%, respectively, for WB PET, versus 43% and 98%, respectively, for combined results from chest radiography, abdominal ultrasound, and bone scintigraphy.The combination of WB PET with CT in a single platform (PET/CT) has been shown to have increased utility over either PET...

Small animal PET scanners may be improved by increasing the sensitivity, improving the spatial resolution and improving the uniformity of the spatial resolution across the field of view. This may be achieved by using PET detectors based on crystal elements that are thin in the axial and transaxial directions and long in the radial direction, and by employing depth of interaction (DOI) encoding to minimize the parallax error. With DOI detectors, the diameter of the ring of the PET scanner may also be decreased. This minimizes the number of detectors required to achieve the same solid angle coverage as a scanner with a larger ring diameter and minimizes errors due to non-collinearity of the annihilation photons. In this study, we characterize prototype PET detectors that are finely pixelated with individual LSO crystal element sizes of 0.5 mm × 0.5 mm × 20 mm and 0.7 mm × 0.7 mm × 20 mm, read out at both ends by position sensitive avalanche photodiodes (PSAPDs). Both a specular reflector and a diffuse reflector were evaluated. The detectors were characterized based on the ability to clearly resolve the individual crystal elements, the DOI resolution and the energy resolution. Our results indicate that a scanner based on any of the four detector designs would offer improved spatial resolution and more uniform spatial resolution compared to present day small animal PET scanners. The greatest improvements to spatial resolution will be achieved when the detectors employing the 0.5 mm × 0.5 mm × 20 mm crystals are used. Monte Carlo simulations were performed to demonstrate that 2 mm DOI resolution is adequate to ensure uniform spatial resolution for a small animal PET scanner geometry using these detectors. The sensitivity of such a scanner was also simulated using Monte Carlo simulations and was shown to be greater than 10 % for a four ring scanner with an inner diameter of 6 cm, employing 20 detectors per scanner ring.

Many laboratories are developing depth-encoding detectors to improve the trade-off between spatial resolution and sensitivity in positron emission tomography (PET) scanners. One challenge in implementing these detectors is the need to calibrate the depth of interaction (DOI) response for the large numbers of detector elements in a scanner. In this work, we evaluate two different methods, a linear detector calibration and a linear crystal calibration, for determining DOI calibration parameters. Both methods can use measurements from any source distribution and location, or even the intrinsic LSO background activity, and are therefore well-suited for use in a depth-encoding PET scanner. The methods were evaluated by measuring detector and crystal DOI responses for all eight detectors in a prototype depth-encoding PET scanner. The detectors utilize dual-ended read out of lutetium oxyorthosilicate (LSO) scintillator arrays with position-sensitive avalanche photodiodes (PSAPDs). The LSO arrays have 7×7 elements, with a crystal size of 0.92×0.92×20 mm 3 and pitch of 1.0 mm. The arrays are read out by two 8×8 mm 2 area PSAPDs placed at opposite ends of the arrays. DOI is measured by the ratio of the amplitude of the total energy signals measured by the two PSAPDs. Small variations were observed in the DOI responses of different crystals within an array as well as DOI responses for different arrays. A slightly nonlinear dependence of the DOI ratio on depth was observed and the nonlinearity was larger for the corner and edge crystals. The DOI calibration parameters were obtained from the DOI responses measured in singles mode. The average error between the calibrated DOI and the known DOI was 0.8 mm if a linear detector DOI calibration was used, and 0.5 mm if a linear crystal DOI calibration was used. A line source phantom and a hot rod phantom were scanned on the prototype PET scanner. DOI measurement significantly improved the image spatial resolution no matter which DOI calibration method was used. A linear crystal DOI calibration provided slightly better image spatial resolution compared with a linear detector DOI calibration.

A dedicated breast PET/CT system has been constructed at our institution, with the goal of having increased spatial resolution and sensitivity compared to whole-body systems. The purpose of this work is to describe the design and the performance characteristics of the PET component of this device. Average spatial resolution of a line source in warm background using maximum a posteriori (MAP) reconstruction was 2.5 mm, while average spatial resolution of a phantom containing point sources using filtered back projection (FBP) was 3.27 mm. A sensitivity profile was computed with a point source translated across the axial field of view (FOV) and the peak sensitivity of 1.64% was measured at the center of FOV. Average energy resolution determined on a per-crystal basis was 25%. Characteristic dead time for the front-end electronics and data acquisition (DAQ) was determined to be 145 ns and 3.6 µs, respectively. With no activity outside the FOV, a peak noise-equivalent count rate of 18.6 kcps was achieved at 318 µCi (11.766 MBq) in a cylindrical phantom of diameter 75 mm. After the effects of exposing PET detectors to x-ray flux were evaluated and ameliorated, the combined PET/CT scan was performed. The percentage standard deviations of uniformity along axial and transaixal directions were 3.7% and 2.8%, respectively. The impact of the increased reconstructed spatial resolution compared to typical whole-body PET scanners is currently being assessed in a clinical trial.