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DOI: 10.1148/radiol.2443061756
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(Radiology 2007;244:807-814.)
© RSNA, 2007

Innovations

PET/MR Images Acquired with a Compact MR-compatible PET Detector in a 7-T Magnet1

Martin S. Judenhofer, BS, Ciprian Catana, PhD, Brian K. Swann, MS, Stefan B. Siegel, PhD, Wulf-Ingo Jung, PhD, Robert E. Nutt, MS, Simon R. Cherry, PhD, Claus D. Claussen, MD, and Bernd J. Pichler, PhD

1 From the Laboratory for Preclinical Imaging and Imaging Technology, Clinic of Radiology, University of Tübingen, Röntgenweg 13, 72076 Tübingen, Germany (M.S.J., C.D.C., B.J.P.); Department of Biomedical Engineering, University of California, Davis, Calif (C.C., S.R.C.); Siemens Preclinical Solutions, Knoxville, Tenn (B.K.S., S.B.S., R.E.N.); and Bruker BioSpin MRI, Ettlingen, Germany (W.I.J.). Received October 10, 2006; revision requested December 7; revision received December 20; accepted January 23, 2007; final version accepted March 1. Supported by NIH grant EB004483-01. Address corres-pondence to B.J.P. (e-mail: bernd.pichler{at}med.uni-tuebingen.de).


   ABSTRACT
TOP
 ABSTRACT
INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Purpose: To prospectively use compact avalanche photodiodes instead of photomultiplier tubes to integrate a positron emission tomographic (PET) detector and a 7-T magnetic resonance (MR) imager.

Materials and Methods: All animal experiments were performed in accordance with the University of Tübingen guidelines and the German law for the protection of animals. A compact lutetium oxyorthosilicate–avalanche photodiode PET detector was built and optimized to operate within a 7-T MR imager. The detector performance was investigated both outside and inside the magnet, and MR image quality was evaluated with and without the PET detector. Two PET detectors were set up opposite each other and operated in coincidence to acquire PET images in the step-and-shoot mode in a mouse head specimen after injection of fluorine 18 fluorodeoxyglucose.

Results: The performance of the PET detector when operated inside the magnet during MR image acquisition showed little degradation in energy resolution (increase from 14.6% to 15.9%). The PET detector did not influence MR imaging. The fused PET and MR images showed an anatomic match and no degradation of image quality.

Conclusion: Simultaneous PET and MR imaging with a 7-T system was deemed feasible.

© RSNA, 2007

Editor's note: In January 2006 (From the Editor), I announced a new section in Radiology—Innovations. Under this banner, we will publish original research that may possibly have far-reaching implications. Authors interested in having their manuscripts considered for Innovations should first read the Editoral to learn of more specifics. Regarding the manuscript by Judenhofer et al, positive comments we received concerning its publication in Innovations included: "Just as PET/CT has demonstrated the great power of multimodality imaging, so this work is an important step toward use of MR imaging in combined PET/MR imaging machines, thus harnessing the important and different information provided by MR imaging."

—Anthony V. Proto, MD, Editor


   INTRODUCTION
TOP
 ABSTRACT
 INTRODUCTION
MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Currently, detector research is focused on multimodality in vivo imaging in an effort to combine functional and morphologic information for clinical diagnosis and preclinical research with laboratory animals. Combined positron emission tomography (PET) and computed tomography (CT) (1) have already been shown to have great value, especially in tumor diagnosis and staging (1–4). However, in contrast to CT, magnetic resonance (MR) imaging does not expose the patient to additional radiation; furthermore, MR imaging can yield images with high soft-tissue contrast without the use of contrast agents. The ionizing radiation dose and the use of large amounts of contrast agents can alter the biologic process of interest and should therefore be minimized. If stand-alone PET and MR imaging systems are used and image data are fused manually, movement from one imaging device to another or long examination times often make coregistration impossible, especially in small regions such as lymph nodes. Thus, current developments are fostering the combination of PET and MR imaging for simultaneous data acquisition.

Attempts to combine PET and MR imaging go back approximately 10 years (5–9). A limiting feature of early combined PET and MR imaging designs was the bulky photomultiplier-based PET detectors that were sensitive to magnetic fields (10). Originally, the focus was on an optical fiber–based system that channeled the scintillation light produced in the PET detectors to photomultiplier tubes positioned outside the magnet and in the fringe magnetic field (5,7–9). The drawback of this concept was that the PET signal quality suffered from light loss caused by the light transmission via optical fibers over several meters, resulting in reduced timing resolution, energy resolution, and crystal decoding accuracy. However, current approaches that combine PET and MR imaging are based on avalanche photodiode (APD) technology (11–14) or complex MR imaging modifications, such as (a) field-cycled MR imaging, in which the PET detector is used to acquire data when the magnetic field is turned off (15), or (b) split magnets (16). APDs are light detectors that have been successfully used with PET technology (11,17). Catana et al (18) used APDs combined with short fibers to create a small-animal PET/MR system—and thus ensure an MR field of view without any metal or electronic parts—by placing the APDs so that they were offset in the transverse direction. Since only short fibers were needed, the performance of the PET detector was only slightly degraded.

Our goal was to develop and optimize a compact full PET detector ring that could be used while a subject was in the MR imager, thus making optical fiber coupling redundant. Our intent was to enable both the PET scanner and the MR imager to operate at their full performance potential without influencing each other. In addition, we intended the design of the system to allow simultaneous PET/MR imaging.

Initial measurements obtained with PET technology based on lutetium oxyorthosilicate (LSO) crystal blocks coupled to 3 x 3 APD arrays showed the feasibility of using an APD-based PET detector inside a high-field-strength MR system (18,19). However, preliminary results (19) showed there was substantial interference between the two imaging systems and degraded system performance. Thus, the purpose of our study was to prospectively use compact APDs instead of photomultiplier tubes to integrate a PET detector system within a 7-T MR imager.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Design of the MR-compatible PET Scanner
The PET system was designed to be installed within either a 7-T BioSpec 70/30 Ultra Shielded Refrigerated MR imager (Bruker BioSpin MRI, Ettlingen, Germany) or a 7-T ClinScan MR system (Bruker BioSpin MRI). In this study, we used the BioSpec 70/30 USR MR imager, which operated at a 300-MHz resonance frequency. For all MR data acquisition and analysis, the ParaVision (Bruker BioSpin MRI) software platform was used. The full-ring PET scanner was under construction at the time of this writing. (It has since been completed and is being tested.) It consisted of 10 block detectors arranged in a ring with an inner diameter of 60 mm. The crystal blocks had a 19.0 x 19.0 x 4.5-mm volume and formed a 19-mm transverse field of view and an approximately 45-mm transaxial field of view. This field of view was sufficient for imaging the brain, heart, or abdominal region of a mouse. The outer diameter of the PET ring was only 120 mm and fit into the small gradient set (B-GA 12; Bruker BioSpin MRI) of the MR imager. The microimaging radiofrequency (RF) coil (Bruker BioSpin MRI) had a 60-mm outer diameter and a 36-mm inner diameter and fit inside the PET scanner (Fig 1).


Figure 1
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  		Figure 1: Schematic drawings of the front (left image) and side (right image) of the magnet with the gradient set, PET system, and RF probe in place.

 
PET Detector Design
Each PET detector consisted of a 19 x 19-mm crystal block (Siemens Preclinical Solutions, Knoxville, Tenn) that comprised 12 x 12 individual 1.5 x 1.5 x 4.5-mm crystals separated with highly reflective foil (17). The crystal block was coupled with a 3-mm-thick light guide to a monolithic 3 x 3 APD array (APD 5054; Hamamatsu, Hamamatsu City, Japan) (Fig 2) where the individual APDs had an active 5 x 5-mm surface (17). The APDs were operated at a negative bias voltage below the breakdown voltage of approximately –405 V. On the basis of the findings of a previous study (19) performed with large hybrid amplifiers, the nine APD signals were fed into a highly integrated nine-channel charge-sensitive preamplifier (Siemens Preclinical Solutions) (17). The output signals of the preamplifier were buffered and were led outside the magnet by 6-m fully shielded nonmagnetic coaxial cables (Leoni, Nürnberg, Germany). All electronic parts were selected to be nonmagnetic and were mounted on a custom-made six-layer printed circuit board that had a flexible connection between the APD and the preamplifier to allow height adjustment of the LSO-APD detector in the radial position of the gantry (Fig 2). The printed circuit board was optimized for operation in a high magnetic field to keep eddy currents in the power and ground planes low.


Figure 2
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  		Figure 2: Top: Crystal block with light guide (a) and APD (b) mounted on the custom-made flex-rigid printed circuit board containing the preamplifier (d), buffers (e), and connectors (f). The flexible APD receptacle (c) enables adaptation for accurate radial positioning of the PET detector. Middle: Printed circuit board (28 x 192 mm) and crystal mounted into detector housing. Bottom: Fully enclosed PET detector.

 
While the LSO-APD detector, preamplifier, and buffers (Fig 2) resided inside the magnet, the nine analog signals were processed outside the 0.0005-T line with a custom-made nine-to-four channel analog multiplexer that provided event position and energy information from the 12 x 12 crystal block. For the test setup used in this study, the analog signals were postprocessed with standard nuclear instrument modules, and analog-to-digital conversion was performed with a PD2-MFS-8-2M/14 data acquisition board (United Electronic Industries) mounted on a standard personal computer (20). For the final detector ring, dedicated PET electronics (Siemens Preclinical Solutions) were used to digitize the signals and perform coincidence processing for all 10 detectors. On the basis of the problems seen in preliminary studies (19), we used double-sided printed circuit board material coated with 10-µm-thick copper for electromagnetic shielding to protect the PET front-end electronics from distortions induced by MR sequences. The copper layer was only 10 µm thick to avoid MR image artifacts caused by eddy currents induced by the material. The copper material was water jet cut into pieces and soldered together at the edges to form a solid box (Fig 2).

Performance Test of the LSO-APD PET Detector Outside the Magnetic Field
As described, several important optimizations of the front-end electronics were performed on the basis of experience reported in a previous work (19) to minimize interference between the PET and MR systems. To ensure proper performance of the new front-end board layout, all nine channels of the preamplifier were tested by simulating a 100-pF APD capacitance on the amplifier input. The results were compared with previously published performance parameters of the amplifier (17). Three authors (M.S.J., B.K.S., and S.B.S.) working together obtained measurements and evaluated data.

A completely assembled and shielded PET detector was tested outside the magnetic field. The crystal block and light guide were glued with UV-curable optical glue (OP 20; Dymax), whereas the light guide was coupled to the APD array with optical grease (Bicron BC 630; Saint-Gobain Ceramics & Plastics) so that the same light guide and LSO block was used for all the boards tested. Position profiles that showed the individual crystals of the LSO-APD block detector were acquired by exposing the crystal block to a 100-kBq germanium 68 point source. The APD array was biased at –380 V. All measurements described herein were obtained at room temperature without additional cooling of the PET detector. For data acquisition, the four signals from the multiplexer were shaped with a semi-Gaussian filter (300-nsec shaping time) and subsequently fed to the previously mentioned data acquisition board for digitization and signal processing (20).

To generate an analog-to-digital conversion timing signal for the data acquisition board at the peak of the analog pulse, the four signals from the multiplexer were split before shaping, and one signal path was processed with a custom-made fast summing circuit and then fed to a constant fraction discriminator (CFD 103; Paul Scherrer Institute, Zurich, Switzerland) and gate and delay generator (DT 102; PSI) to generate an accurate timing signal. Position profiles and the mean peak-to-valley ratio for three center crystals and four edge crystals (two on each side) were calculated. The single crystal energy spectra of the PET detector and the energy resolution of center and corner crystals were calculated with custom software (20). The quantitative values are reported as mean values ± one standard deviation. One author (M.S.J.) acquired and evaluated data.

Performance Test of the LSO-APD PET Detector Inside the 7-T MR Imager
Comparable tests with an individual PET detector, as described previously, were performed inside the 7-T imager and during the application of pulse sequences to determine whether the PET detector (electronics and shielding) and the MR imager (gradients and RF signals) interfere with each other and whether there is a loss of PET or MR image quality. One author (M.S.J.) obtained measurements outside the magnet and evaluated data.

For MR measurements, an RF coil with a 72-mm inner diameter was used with the 200-mm gradient set (B-GA 20; Bruker BioSpin MRI). A polymethylmethacrylate cylinder (outer diameter, 28 mm; length, 100 mm) filled with silicone oil (M10; Roth, Karlsruhe, Germany) was used as a homogeneous phantom and placed inside the RF coil. A standard fast low-angle shot imaging sequence (repetition time msec/echo time msec, 400/6; flip angle, 30°; 256 x 256 pixels) was used to acquire transverse images of the phantom. The PET detector was positioned with its center field of view aligned with the magnet isocenter and radially at the outer edge of the RF coil in the same way the PET detectors were arranged in the final setup (Fig 1). MR images were acquired with and without the PET detector inside the magnetic field.

As a measure of image quality, the ratio of the signal intensity of the phantom to the signal intensities of the background outside the phantom and the signal-to-noise ratio were determined. Signal intensity was measured in five concentric 20-mm regions of interest placed on five adjacent transverse image sections, and a mean value was calculated. To calculate the signal-to-background and signal-to-noise ratios, four regions of interest—each with a diameter of 10 mm—were placed at the corners outside the phantom on the same transverse image sections where the regions of interest had been placed. The mean signal intensity of these background regions of interest was used as the background, and their standard deviation was used as the noise. The PET detector was switched on, and data were acquired during the MR acquisition. The performance of the PET detector inside the MR imager and while MR sequences were performed was assessed by acquiring position profiles and energy spectra and subsequently comparing them with measurements obtained outside the MR imager.

Coincidence Detector Setup for Simultaneous PET/MR
Since the previous test results demonstrated good performance of the optimized LSO-APD PET detector, even when operated simultaneously with the MR imager, two PET detector modules were mounted with a 69-mm interval between the crystal block surfaces on a gantry made of polymethylmethacrylate (Fig 3) and set up in coincidence to allow acquisition of PET data. The gantry had an inner diameter of 60 mm to hold the microimaging RF probe and an outer diameter of 200 mm to fit inside the 200-mm gradient set of the MR imager. Signal processing was similar to that with the single detector setup. Coincidences (12-nsec time window) were generated by feeding the fast summed analog output signal of each detector block into a constant fraction discriminator and feeding both constant fraction discriminator outputs into a coincidence unit (LRS 466; LeCroy, Chestnut Ridge, NY) that triggered the eight-channel data acquisition board. One author (M.S.J.) prepared the setup.


Figure 3
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  		Figure 3: Left: This prototype gantry, which fits inside the gradient set and holds two PET detectors and the RF coil, was used to acquire the first simultaneous PET/MR images inside a 7-T magnet. Right: Frontal view of the PET gantry, with the RF coil placed inside the imager.

 
PET imaging was performed by rotating the object with a step-and-shoot acquisition with 12 projections over 180°. After acquisition, all data were sorted into a sinogram, and each projection was normalized and corrected for decay by scaling the respective sinogram values. Normalization data were obtained by acquiring coincidence data from a flood source phantom with inner dimensions of 22 x 22 x 1 mm and filled with 20 MBq of fluorine 18 (18F), which was placed between the detectors. Normalized PET emission sinogram data were reconstructed without further corrections by using standard filtered back projection (cutoff frequency of 0.5) into a 128 x 128 matrix. After reconstruction, the images were smoothed with a 2.5-mm full width half maximum Gaussian filter. Two authors (M.S.J. and S.B.S.) working together performed imaging tasks.

Simultaneous 18F Fluorodeoxyglucose PET and MR Imaging of a Mouse Head
All animal experiments were performed in accordance with University of Tübingen guidelines for the use of living and dead animals in scientific studies and German laws for the protection of animals. One female C57 BL/6 mouse was intravenously injected with a 200-MBq dose of 18F fluorodeoxyglucose (FDG) and sacrificed 45 minutes after radiotracer uptake. The simultaneous PET/MR examinations were performed outside the University of Tübingen at Bruker BioSpin MRI; therefore, federal regulations limited the maximum amount of radioactivity that could be handled at this site. Thus, only the mouse head was transported to the company's site. The head had a remaining activity of about 8 MBq at the start of the examination (4 hours after injection). The head was placed in the 19-mm transverse field of view of the PET scanner. A total of 12 projections (6-minute duration for each) were acquired. During acquisition of each projection, coronal MR images were acquired with a fast low-angle shot sequence (394/5.9, 40° flip angle, six signals acquired, 1-mm section thickness, 256 x 256 pixels, 6-minute acquisition time). The PET data were reconstructed as described. Coronal views of the PET data were fused with the MR images by using MiraView software (Siemens Preclinical Solutions). Matching of PET data to MR data was evaluated visually. Three authors (M.S.J., W.I.J., and B.J.P.) working in consensus evaluated MR images. Two authors (M.S.J. and B.J.P.) working in consensus performed PET image acquisition, fusion, and evaluation.


   RESULTS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 
Performance of the LSO-APD PET Detector Outside the Magnetic Field
The measured preamplifier performance in combination with the front-end board, which was modified for use in an MR imager, was maintained and comparable with the results reported previously (17). All crystals in the position profile acquired from the LSO-APD PET detector outside the MR imager could be separated, and the mean peak-to-valley ratio measured from a profile through a center crystal row was 8.8 ± 2.9 for the center crystals and 2.7 ± 1.3 for the crystals at the edges (Fig 4). The energy spectra of one center crystal and one corner crystal clearly showed the 511-keV peak was well separated from the Compton edge. The mean energy resolutions of four center crystals were 14.6% ± 0.32 and 19.9% ± 3.34 for the crystals in the corners of the block (Table). The lower 511-keV photopeak position of the corner crystals indicated a 21.5% light loss compared with the center crystals.


Figure 4A
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  		Figure 4a: (a) Position profiles of the LSO-APD PET detector (12 x 12 crystals) acquired inside and outside the imager show that all 144 crystals can be identified. (b) Plot of energy histograms of center and corner crystals acquired inside and outside the imager. (c) Profiles through the position profile acquired inside and outside the imager show the peak-to-valley ratios.

 

Figure 4B
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  		Figure 4b: (a) Position profiles of the LSO-APD PET detector (12 x 12 crystals) acquired inside and outside the imager show that all 144 crystals can be identified. (b) Plot of energy histograms of center and corner crystals acquired inside and outside the imager. (c) Profiles through the position profile acquired inside and outside the imager show the peak-to-valley ratios.

 

Figure 4C
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  		Figure 4c: (a) Position profiles of the LSO-APD PET detector (12 x 12 crystals) acquired inside and outside the imager show that all 144 crystals can be identified. (b) Plot of energy histograms of center and corner crystals acquired inside and outside the imager. (c) Profiles through the position profile acquired inside and outside the imager show the peak-to-valley ratios.

 

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  		Energy Resolution and Peak Positions Measured with the LSO-APD Detector inside and outside the MR Imager

 
Performance of the LSO-APD PET Detector Inside the 7-T MR Imager
The MR images acquired with and without the operating PET detector mounted to the RF coil showed no considerable degradation in image quality, especially toward the bottom of the phantom where the PET detector was located (Fig 5). The notch on the top of the images of the phantom was caused by an air bubble inside the phantom. The signal-to-noise ratio of the phantom imaged without the detector was 175, and the signal-to-background ratio was 92. When the PET detector was placed in the MR imager, the signal-to-noise ratio was 177 and the signal-to-background ratio was 93.


Figure 5
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  		Figure 5: Transverse unenhanced fast low-angle shot MR images (400/6, 30° flip angle, 256 x 256 pixels) of a silicone oil phantom acquired with and without the PET detector mounted to the RF coil. The rectangle below the right image indicates the position of the PET detector.

 
Quantitative analysis of the MR images enabled us to confirm that MR image quality was maintained when the PET detector was used inside the MR imager. The position profile acquired with the PET detector when it was located inside the MR imager and while a fast low-angle shot imaging sequence was being performed (Fig 4) showed only minor changes compared with the position profile acquired outside the MR imager (Fig 4). The mean peak-to-valley ratios decreased to 5.7 ± 1.9 for the center crystals and to 2.4 ± 1.3 for the edge crystals. The mean 511-keV photopeak positions for corner and center crystals decreased by 6% compared with measurements obtained outside the magnet (Table). The energy resolution of the center crystals (full width half maximum at 511 keV) increased from 14.6% ± 0.3 to 15.9% ± 0.7 when it was measured inside the magnet while MR imaging sequences were performed.

Simultaneous FDG PET and MR Imaging of a Mouse Head
No degradation of image quality was observed on the simultaneously acquired images of the mouse head (Fig 6). The PET/MR images showed an anatomic match and FDG uptake, especially in the cortex and Harderian glands of the mouse (Fig 6).


Figure 6
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  		Figure 6: Simultaneously acquired PET (filtered back projection, 2.5-mm Gaussian postsmoothing filter) and coronal unenhanced fast low-angle shot MR (394/5.9, 40° flip angle, six signals acquired, 1-mm section thickness, 256 x 256 pixels) images of a mouse head injected with FDG. The fused PET/MR images show good alignment of images acquired with the two imaging modalities. The increased uptake of the PET images correlates with the location of the Harderian glands behind the eyes in the MR images.

 

   DISCUSSION
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
ADVANCES IN KNOWLEDGE
 References
 
Our test results for the preamplifier in combination with the dedicated multilayer board, which had been modified for MR applications, showed superior performance compared with those for the hybrid amplifier used in the prototype setup (19). Testing of the fully assembled LSO-APD block detector revealed that all crystals can be resolved in the acquired position profiles. The mean energy resolution of the block detector was 14.6% for the center crystals and 19.9% for the corner crystals. These results are comparable to those for other APD-based block detectors (21) and are even comparable to results obtained with state-of-the-art photo multiplier tube block readout schemes (21–23).

When the modified PET detector module was used inside the 7-T imager and while MR images were acquired, the energy resolution increased from 14.6% to 15.9% for the center crystals and from 19.9% to 21.9% for the corner crystals. This is an absolute change of only 2.0% in energy resolution, and it is probably a result of a slight increase in temperature inside the MR system. Earlier test results have shown that the gain and noise of an LSO-APD detector vary with the temperature changes by approximately 3% per Kelvin. The MR phantom images showed no visible interference when imaging was performed while an operating PET detector was located around the RF coil; this finding was in sharp contrast to the results presented by Pichler et al (19). The signal-to-noise and signal-to-background ratios remained the same when MR imaging was performed with or without any PET detector material inside the MR imager. This finding confirmed that the printed circuit board material covered with a thin copper layer on both sides was a good choice for PET detector shielding and that the use of this material drastically reduced eddy currents compared with the use of other shielding materials tested (19).

The two coincident PET detectors performed well when they were used inside the 7-T system. The first simultaneous PET/MR images of an FDG-enhanced mouse head revealed no degradation of image quality for either MR imaging or PET. To our knowledge, this is the first example of simultaneous PET/MR imaging of a biologic specimen with a combined PET/MR imager in which the entire PET detector resided in the active imaging region of the MR imager. The fused images showed the expected FDG uptake in the brain and Harderian glands of the mouse, which matched the anatomic landmarks. Since the PET field of view was physically aligned with the MR imaging field of view, fusing PET and MR images was not a problem. Compared with other approaches involving the use of PET/MR systems (5,7–9,24), the LSO-APD detector used in our study provided multiple PET sections that covered a transverse field of view of 19 mm and was easily extended owing to the compact and modular nature of the PET detector block design. In addition, the fully integrated detector design makes light fibers redundant and provides better energy resolution (8), although further studies with a full PET detector ring in the magnet are needed to show whether MR image quality is maintained and spectroscopy is feasible.

Our study had limitations. First, our use of only two PET detectors in coincidence prevented us from acquiring dynamic data. However, we wanted to determine whether combined PET and MR acquisition was feasible with our detector design. Second, potential artifacts from a full-ring PET system need to be evaluated further from the MR perspective. Third, the capability to perform more demanding MR sequences, such as echo-planar imaging, needs to be assessed in the presence of the full system.

In conclusion, our results confirm that simultaneous PET and high-field-strength MR imaging with LSO-APD–based PET detectors is feasible without sacrificing the quality of images obtained with either system. The next step will be to focus on the evaluation of the full-ring PET insert. While this work was concentrated on small-animal PET/MR imaging, our results can be transferred to a clinical system, where MR imaging is usually performed at a much lower magnetic field strength. The combination of PET and MR imaging can open new opportunities in preclinical research and clinical diagnosis. When nuclear MR spectroscopy is used with MR imaging and PET, trimodal imaging might be possible and thereby add even more information in biomedical research studies.


   ADVANCES IN KNOWLEDGE
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
References
 


   ACKNOWLEDGMENTS
 
We thank Yuan-Chuan Tai, PhD, (Washington University) for providing the light guides used in this work. Furthermore, we thank Sascha Köhler, PhD, and Wolfgang Kreibich (Bruker BioSpin MRI) for their assistance with the MR measurements. Help from Gerald Reischl, PhD, and Klaus-Dieter Keller, PhD, (University of Tübingen) in providing the FDG was greatly appreciated. Finally, the support of Ralf Ladebeck; Robert Krieg, PhD; Ronald Grazioso; Nan Zhang, PhD; and Matthias Schmand, PhD, (Siemens Preclinical Solutions) was valuable to this project.


   FOOTNOTES
 

Abbreviations: APD = avalanche photodiode • FDG = fluorine 18 fluorodeoxyglucose • LSO = lutetium oxyorthosilicate • RF = radiofrequency

Author contributions: Guarantors of integrity of entire study, M.S.J., B.J.P.; study concepts/study design or data acquisition or data analysis/interpretation, all authors; manuscript drafting or manuscript revision for important intellectual content, all authors; manuscript final version approval, all authors; literature research, M.S.J., C.C., B.K.S., S.R.C., B.J.P.; experimental studies, M.S.J., C.C., B.K.S., S.B.S., W.I.J., R.E.N., S.R.C., B.J.P.; statistical analysis, M.S.J., C.C., B.J.P.; and manuscript editing, M.S.J., C.C., S.B.S., R.E.N., S.R.C., C.D.C., B.J.P.

See also the editorial by Zaidi in this issue.


   References
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
 DISCUSSION
 ADVANCES IN KNOWLEDGE
 References
 

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  4. Reinhardt MJ, Joe AY, Jaeger U, et al. Diagnostic performance of whole body dual modality 18F-FDG PET/CT imaging for N- and M-staging of malignant melanoma: experience with 250 consecutive patients. J Clin Oncol 2006;24:1178–1187.[Abstract/Free Full Text]
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