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Nuclear Inst. and Methods in Physics Research, A 959 (2020) 163575
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Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
Design, evaluation and initial imaging results of a PET insert based on strip-line readout for simultaneous PET/MRI H. Kim a ,∗, Y. Hua b , H.-T. Chen a , H.-M. Tsai a , C.-T. Chen a , G. Karczmar a , X. Fan a , D. Xi b , Q. Xie b , C.-Y. Chou c , C.-M. Kao a a
Department of Radiology, University of Chicago, Chicago, IL 60637, USA Biomedical Engineering Department, Huazhong University of Science and Technology, Wuhan, China c Department of Bio-Industrial Mechatronics Engineering, National Taiwan University, Taipei, Taiwan b
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Keywords: Positron Emission Tomography Magnetic Resonance Imaging Strip-line readout Sampling
ABSTRACT We present the development of a PET insert system for potential simultaneous PET/MR imaging using a 9.4 T small animal MRI scanner to test our system. The detectors of the system adopt a strip-line based multiplexing readout method for SiPM signals. In this readout, multiple SiPM outputs in a row share a common strip-line. The position information about a hit SiPM is encoded in the propagation time difference of the signals arriving at the two ends of the strip-line. The use of strip-lines allows us to place the data acquisition electronics remotely from the detector module to greatly simplify the design of the detector module and minimize the mutual electromagnetic interference. The prototype is comprised of 14 detector modules, each of which consists of an 8x4 LYSO scintillator array (each LYSO crystal is 3x3x10 mm3 ) coupled to two units of Hamamatsu MPPC arrays (4x4, 3.2 mm pitch) that are mounted on a strip-line board. On the strip-line board, outputs of the 32 SiPMs are routed to 2 strip-lines so that 16 SiPM signals share a strip-line. The detector modules are installed inside a plastic cylindrical supporting structure with an inner and outer diameter of 60 mm and 115 mm, respectively, to fit inside a Bruker BioSpec 9.4 T MR scanner. The axial field of view of the prototype is 25.4 mm. The strip-lines were extended by using 5-meter cables to a sampling data acquisition (DAQ) board placed outside the magnet. The detectors were not shielded in the interest of investigating how they may affect and be affected by the MRI. Experimental tests were conducted to evaluate detection performance, and phantom and animal imaging were carried out to assess the spatial resolution and the MR compatibility of the PET insert. Initial results are encouraging and demonstrate that the prototype insert PET can potentially be used for PET/MR imaging if appropriate shielding will be implemented for minimizing the mutual interference between the PET and MRI systems.
1. Introduction Simultaneous PET/MR imaging combines the complementary merits of PET and MRI; the former can provide functional imaging with high sensitivity and specificity [1,2], and the latter can produce anatomical imaging at high spatial resolution [3]. Compared to CT, MRI has higher soft-tissue contrast, no radiation dose, and diverse contrast mechanisms through various pulse sequences. Therefore, PET/MRI has been considered as an alternative to PET/CT in many clinical and pre-clinical imaging applications. MRI uses strong magnetic fields (3–15 T), fast switching gradient fields, and radio-frequency (RF) signals for image formation. These fields and RF signals are known to affect the operations of a PET detector and its electronics. On the other hand, the existence of PET inside the MR scanner affects the magnetic field homogeneity and
can produce RF signals that interfere with MRI. Therefore, building a PET/MRI system requires MR-compatible detector materials and system design to maintain proper functions of both modalities. Various PET/MRI systems have been developed based on different detector technologies to overcome these challenges. Earlier developments used long optical fibers to separate the scintillators, which are inside the magnet, and photo-sensors [4–6] such as the photo-multiplier tube (PMT), which are outside the magnet, to eliminate the strong magnetic effects and constraints of space limitation of MR. The long optical fibers, however, induced significant loss of the scintillation lights. With the advent of compact and magnetic field in-sensitive photosensors such as the avalanche photo-diode and silicon photo-multiplier (SiPM) [7,8], it became possible to integrate the photo-sensors with the detector modules that are inserted inside the magnet. This design approach has been adopted by many research groups [9–15]. At present,
∗ Corresponding author. E-mail address: [email protected] (H. Kim).
https://doi.org/10.1016/j.nima.2020.163575 Received 15 April 2019; Received in revised form 15 January 2020; Accepted 31 January 2020 Available online 5 February 2020 0168-9002/© 2020 Elsevier B.V. All rights reserved.
H. Kim, Y. Hua, H.-T. Chen et al.
Nuclear Inst. and Methods in Physics Research, A 959 (2020) 163575
SiPM has emerged as the photo-sensor of choice for PET, particularly for PET/MRI, because of its in-sensitivity to magnetic fields, compact size, and PMT-compatible signal amplification. Excellent reviews of the current development status can be found in the literature [16–18]. Commercial SiPMs are available in 1 × 1 mm2 to 6 × 6 mm2 pixel sizes. Therefore, a practical PET system that typically employs a large number of SiPMs requires an efficient signal multiplexing scheme for readout [19–21]. Most proposed multiplexing readouts, however, significantly compromise the timing characteristics of detector. We have proposed and successfully demonstrated a multiplexing readout method for SiPM-based time-of-flight (TOF) PET [22]. In this method, multiple SiPMs share a single strip-line that is read at the two ends. The position of the hit SiPM on the strip-line is decoded from the time difference between the SiPM output signals arriving at the two ends. Our recent study shows that a single strip-line can connect 32 SiPMs without compromising the detector performance, achieving ∼270 ps coincidence time resolution [23]. With strip-line readout, data acquisition (DAQ) electronics can be separated from the detector module and located remotely via long cables that extend the strip-lines. As a result, the mechanic, electronic and RF shielding designs of the detector may be greatly simplified; the detector can be kept highly compact with minimal materials, which also help to minimize the impact on the magnetic field uniformity. Therefore, we believe PET detectors based on strip-line readout are favorable for developing PET/MRI systems. PET insert systems that adopted long signal transmission cables for coping the aforementioned challenges in developing PET inserts, which was similar to our approach, have been reported in [10,15]. Previously, we have developed detector modules with strip-line readout and demonstrated that the technology is feasible for PET/MRI [24]. Here, we prototyped a PET insert system, adopting PET detectors with strip-line readouts, in combination with the use of a compact 72-channel electronics board that provides an FPGA based sampling DAQ that we previously proposed [25]. While our detectors are TOF capable and have limited resolution for small animal imaging, we tested the technology on a 9.4 T preclinical MRI scanner due to accessibility of the scanner and the more challenging electromagnetic environment and space condition it presented. Experimental tests were conducted to measure the performance of the prototypical PET insert inside and outside the MR scanner. A resolution phantom and uniform phantom were used to assess the MR compatibility of the PET insert. In this paper, we report the design and evaluation of the PET insert, and initial imaging results using the PET insert.
2.1. Detector module
2. Materials and methods
The PET insert was constructed by assembling 14 detector modules in a plastic supporting structure that was designed to fit inside a small animal MR scanner, Bruker 9.4 T BioSpec 94/30 USR (Bruker Biospin, Billerica, MA, US). The inner (ID) and outer diameter (OD) of the PET insert are 6 cm and 12 cm, respectively; they were determined from the MR scanner’s bore size and a RF coil used in this study. The 14 DMs were arranged with the support structure so that the distance between two oppositely facing DMs was 62 mm, and the gap between two adjacent DMs was 1.5 mm. The axial field of the PET insert is 25.4 mm. Figs. 2(a) and 2(b) show the fully assembled PET insert. 5 m long mini coaxial cables (RG-316) were used to transmit the outputs to the DAQ board, which was located outside the MRI room to minimize interference with MR operations. A total of 56 MMCX cables, 4 for each detector module, were used. The bias voltages to the 14 DMs were adjusted in a custom-made voltage distributor that employed LM337 voltage regulators and two units of Lambda GEN750W power supplies. A total of 14 6 m-length cables (24 gauge four conductor stranded) were used to connect the DMs and the power distributor, which was also placed outside the MRI room. All the signal and power cables passed through a penetration panel to be connected to the equipment located outside the MRI room. As shown in Fig. 2, no detector shielding is used in current implementation in the interest of simplifying the design of the PET insert.
The detector module (DM) consists of LYSO scintillator arrays, Hamamatsu Multi-Pixel photon counter (MPPC) arrays, and a stripline board (SLB) as shown in Fig. 1. Two 4 × 4 LYSO arrays (each LYSO crystal is 3 × 3 × 10 mm3 and the crystal pitch is 3.2 mm) are packed together to form a 4 × 8 LYSO array, shown in Fig. 1(b). The LYSO crystals (Shanghai EBO Optoelectronics Technology, China) are coupled one-to-one to SiPM pixels by using optical glue (Cargille Meltmount, refractive index 1.582). Enhanced specular reflectors (ESR, 3M™) are used between LYSOs in the arrays for optical isolation. Hamamatsu MPPC arrays, containing 4 × 4 pixels (each MPPC pixel has a detection surface area of 3 × 3 mm2 with a 3.2 mm pitch that matches the LYSO crystal pitch), were selected for the photo-sensor because of their superior gain uniformity. A SiPM array board shown in Fig. 1(c) was manufactured to house two MPPC arrays, provides bias voltage to SiPMs and routes the outputs of the 32 SiPMs to the SLB via connectors. Due to the rapid development cycle of MPPC, two different models of the MPPC array were used in the PET insert, including 8 units of S12642-0404-PA and 20 units of S13361-3050NE [28]. The main difference between these models is their breakdown voltages, which are measured to be about 65 V and 53 V for S12642 and S13361, respectively. SiPMs in a detector module use the same bias voltage. For each detector module, the bias voltage was adjusted to be about 2.5 V above the module’s breakdown voltage so that all modules produce similar pulse heights for 511 keV photons. The detector was operated at room temperature maintained at 18 ◦ C to avoid SiPM gain variation due to temperature change. The SLB is an 18 × 110 mm2 printed circuit board (PCB) containing two strip-lines, one laid out on the top side and another at the bottom side. Each strip-line receives the outputs of 16 SiPMs so that a total of 32 SiPM outputs of two MPPC arrays are routed into two strip-lines. Fig. 1(d) depicts a stripline layout across 16 SiPM outputs. The strip-line was configured to have 50 ohm characteristic impedance, and, as shown in Fig. 1(e), the trace length between adjacent SiPM outputs was set to 25 mm in order to increase the signal propagation time. Because position decoding on strip-line is based on the measured time difference, this alleviated the requirement on the time accuracy for the DAQ. Fig. 1(f) shows the connection scheme of the SiPM output (negative polarity) to the stripline. Schottky diodes were used for decoupling the capacitances of the SiPMs [29] that shared a strip-line. The ends of each strip-line were connected to micro-miniature coaxial (MMCX) connectors, mounted at the edge of SLB, for transferring the output signals to DAQ. 2.2. PET insert assembly
The prototypical PET insert employs 14 detector modules. Each module has 2 strip-lines and therefore 4 outputs, yielding a total of 56 outputs for the PET insert. For DAQ, a DRS4 waveform sampler [26] was used in our previous evaluation studies of detector modules. In this study, we used the DSR4 sampler again, specifically the 4-channel DRS4 evaluation board, for the purpose of validating and evaluating individual detector modules of the PET insert. For practical use, however, it is beneficial to simplify the DAQ electronics. Hence, for the developed PET insert, we used another sampling DAQ technology that we previously proposed and developed for scintillation detectors. This technology, called multi-voltage threshold (MVT) [25], samples a waveform with respect to up to 4 user-defined voltage thresholds and can be implemented by using FPGAs. As a result, multi-channel DAQ boards capable of supporting TOF detection can be readily developed and modified. In this work, we employed a single 10 × 18 cm2 72channel MVT DAQ board [27] to handle the 56 outputs of the PET insert. In Sections 2.1–2.3, the implementation of the strip-line readout in the detector module, overall design features of the PET insert, and DAQ system are described in detail. In Sections 2.4 and 2.5, we describe the experimental setups and processing the acquired data, respectively. 2
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Nuclear Inst. and Methods in Physics Research, A 959 (2020) 163575
Fig. 1. Components of a DM of the PET insert. (a) A strip-line board (SLB). (b) LYSO arrays coupled to (c) a SiPM array board. (d) A strip-line layout connecting 16 SiPMs output. (e) A close-up of the strip-line on SLB. (f) A schematic of a SiPM output coupling to a strip-line through diodes.
Fig. 2. (a) A cross-sectional view and (b) a side view of the PET insert after all the detector modules were assembled into the supporting structure. The PET insert assembly is not shielded for RF signals.
2.3. Multi-voltage threshold waveform sampling
2.4. Experimental setup
Digitization of the signal outputs of the PET insert was accomplished in a board implementing MVT waveform sampling technology [25] that we have previously developed. In MVT, the input signal is passed to four comparators with user-defined reference voltages. The leading and trailing transition times of the outputs of the comparators are recorded by time-to-digital converters (TDC). Subsequently, the waveform of the input signal is reconstructed from time stamps and the known reference voltages used. An example of the MVT operation is shown in Fig. 3(a). In this case, the blue circles show the samples obtained by using 4 reference voltages, including 50, 150, 300 and 600 mV, and the red solid curve shows the reconstructed waveform obtained by linearly interpolating the samples. The pulse-height of the reconstructed waveform was estimated by integrating the area under the solid curve. Waveform data acquired by a DRS4 sampler were used to check the accuracy of the energy estimation. The area integration method results in ∼10.6% energy resolution at 511 keV photo-peak. On the other hand, ∼8.5% of energy resolution was measured by using the maximum amplitude of the waveforms. For comparison, the dashed curve shows the reconstructed waveform obtained by using a bi-exponential fitting method we previously reported [25]. Previously we have demonstrated that both the comparators and the TDCs can be implemented by using FPAGs that are widely used in modern electronics. Fig. 3(b) shows the FPGA (Altera EP4C115F29C7) based MVT board used in this study (manufactured by Raycan Technology Co., China). This board provides 72 input channels, each of which has 4 comparators. The threshold voltages (0–1 V) of the comparators were programmably configured by using digital-to-analog converters (DAC) (DAC7668, Texas Instruments). The bin width of the TDCs implemented on the board was calibrated to be 95 ps. Because the MVT uses flat flexible cables (FFC) for inputs, an adapter board was built to convert the MMCX connectors from the PET insert to the FFC connectors. The time stamps produced by the MVT were packed into user datagram protocol (UDP) packets and transferred to a computer through a Gigabit Ethernet interface.
The DMs were individually tested on a bench top setup before they were assembled into the PET insert. As already mentioned, a DRS4 waveform sampler that can provide 5 GHz sampling was used in order to obtain accurate performance assessment for the DMs. Positiondecoding accuracy, energy resolution, and coincidence time resolution were characterized. The breakdown voltage of each DM was determined from pulse-height measurements by varying the bias voltage. For measuring the coincidence time resolution, a reference detector, consisting of a Hamamatsu MPPC S10931-050P (3 × 3 mm2 active area) coupled to a 3 × 3 × 10 mm3 LYSO crystal, was used. Measurements were performed by placing a 68 Ge line source (sealed in a cylindrical stainless tube with 19 cm length and 3.2 mm (1.9 mm) of OD (ID)) or a 22 Na point source (∼0.5 mm diameter sphere sealed in a 2.5 cm diameter plastic disk) between the reference detector and the DM under evaluation. After the PET insert was fully assembled, the above performance measurements for the DMs were repeated by using the 72-channel MVT board. A custom-made resolution phantom filled with 18 F solution was also used to assess the spatial resolution of the PET insert. Fig. 4 shows the 68 Ge line source and the resolution phantom. For 68 Ge source imaging, the line source was moved along the radial direction in 5 mm step from the center of the FOV. The resolution phantom has six groups of holes of different diameters, including 1.0, 1.4, 1.8, 2.2, 2.6, and 3.0 mm. The distance between holes within each group is twice the hole diameter, and the depth of the holes is 30 mm. For experiments with MR, the PET insert was placed inside a Bruker 9.4 T BioSpec MR scanner (116 mm bore diameter). An actively shielded gradient coils (maximum strength 230 mT/m) was used. MR imaging was performed by using a 35 mm ID (59 mm OD) quadrature mouse volume coil (RAPID MR International, Columbus, OH, USA). Fig. 5 shows the arrangement of the PET insert, the RF coil, and the gradient coil inside the MR bore. Two MR pulse sequences were used for MR data acquisition: FLASH (fast low angle shot) T1-weighted gradient echo with TR/TE = 350/3 ms and flip angle = 30, and RARE (rapid acquisition with relaxation enhancement) T2-weighted fast spin echo with TR/TE = 4000/25 ms and RARE factor = 8. 3
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Nuclear Inst. and Methods in Physics Research, A 959 (2020) 163575
Fig. 3. (a) An example of waveforms reconstructed from the MVT stamps. (b) A 72-channel MVT board and an adapter board for MMCX to FFC conversion.
(see Fig. 3(a)) to reconstruct the waveform, and then the pulse height for the hit is obtained by integrating the area under the reconstructed waveform. The pulse height of the single event is taken as the average of the pulse heights of the two hits, calibrated to keV by using a 22 Na source. • Coincidence filtering: An energy window of 410–610 keV is used to select photo-peak singles events. For coincidence detection, the event time of a single is taken as the average of the first time stamps of the two hits corresponding to the single (the time stamps of the leading transitions of the lowest threshold, (T𝐿1 + T𝑅1 )/2). Based on this event time, two singles that are within a 2 ns time window are registered as a coincidence event. • Image reconstruction: A line-of-response (LOR) based maximum likelihood expectation maximization (MLEM) algorithm [30] is employed to reconstruct the resulting coincidence data. 30 iterations are used in the image reconstruction. The system response matrix is calculated by using Siddon’s ray tracing algorithm [31]. Currently, the algorithm does not implement corrections for attenuation and scatter. Therefore, the images shown are the baseline results that can be produced by the PET insert.
Fig. 4. (a) A 68 Ge line source in a positioning holder. (b) A resolution phantom has holes with diameter ranging from 1.0 mm to 3.0 mm. The holes are filled with 18 F solution for imaging.
2.5. PET data processing The raw UDP packets generated by the MVT board were saved on the hard drives of the DAQ computer for off-line analysis. The basic element of the UDP packet was one MVT hit, which contained the time stamps of leading and trailing transitions (8 time stamps total) and the input channel number. Fig. 6 shows a simplified diagram of two MVT hits on a strip-line to produce a single event. T𝐿(𝑅)𝑖 in the figure represents the MVT time stamp of the 𝑖th voltage threshold on the two ends of a strip-line. The data analysis flow is described below.
3. Results 3.1. Detector performance 3.1.1. Position decoding by differential time on strip-line Fig. 7(a) shows a pulse-height spectrum (PHS) using singles collected from a selected strip-line. The gain variation of the LYSO/SiPM pixels sharing the strip-line broadens the photo-peak and even creates a small side peak. Fig. 7(b) shows the differential-time (𝛥T) histogram on a strip-line that illustrates the accuracy in position decoding. This particular 𝛥T histogram was obtained by using mainly events in the photo-peak identified on PHS in Fig. 7(a). The 𝛥T histogram clearly shows 16 peaks that reflect the positions of the 16 LYSO/SiPM pixels on the strip-line. A Gaussian fit was applied to each of the 16 peaks, and 𝛥T ranges corresponding to individual LYSO/SiPM pixels were determined by using the midpoints between the center positions of the resulting Gaussians. The average peak-to-valley ratio (P/V), which provides a figure of merit for the accuracy of position encoding, was about 13:1 for the entire PET insert. From the 𝛥T histogram, the time difference
• Single gamma event tagging: Following the 511 keV gamma interaction within the detector, two signals are formed on a stripline and propagate toward the ends of the strip-line, as depicted in Fig. 6. Therefore, two MVT hits appearing at the two ends of a strip-line within a 7 ns time window are tagged as a single event (∣T𝐿1 −T𝑅1 ∣ < 7 ns). In this step, the time difference is determined based on the leading transition time of the lowest threshold. • Position and pulse-height of a single event: The position of a single hit along the strip-line is determined from the time difference of the two registered MVT hits produced by the above step. In this case, however, all four time stamps of the leading transitions of each MVT hit are averaged to obtain the hit time (𝛥T = 𝛴T𝐿𝑖 −𝛴T𝑅𝑖 ). For each hit, its samples are linearly interpolated
Fig. 5. (a) The PET insert and the MVT board are connected by 5 m long MMCX cables. (b) The PET insert is installed into a Bruker 9.4 T small animal imaging MR scanner. (c) A cross-sectional view shows the arrangement of the PET insert, the RF coil, and the gradient coil inside the MRI scanner.
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some input channels fail to produce stable results due to noise. As a result, to ensure all channels function properly, we cannot decrease the lowest threshold below 50 mV with our present MVT board, yielding a system CRT of about 500 ps. 3.1.4. Count rate performance Count rate performance of the MVT board was measured by using the resolution phantom filled with 18 F solution with an initial activity of 265 μCi (most PET mouse imaging uses lower activities). Fig. 11(a) shows the MVT hit rate as a function of radioactivity as it decays to the background level over 12 h. The slight change in the count rate slope around 100 μCi may be due to the data writing load in the DAQ computer. As shown in Fig. 11(a), the count rate was fitted well by a 2nd order polynomial function 𝑦 = p2 𝑥2 + p1 𝑥 + p0 with p2 = −1.292 × 10−5 , p1 = 0.01428, and p0 = 0.04505. The sensitivity of the PET insert was measured by using 22 Na point source with 15 μCi radioactivity. The 22 Na source was moved along the central axis of the PET insert with a 2 mm step. Sensitivity was calculated by dividing the measured coincidence rates to the known activity of the source. Fig. 11(b) shows the sensitivity result (red curve), which shows a peak of about 1.1% at the axial center and steadily decreases toward the ends of the axial FOV. For comparison, the axial sensitivity profile calculated for the PET insert by using Geant4 simulation is also shown (black curve). Note that the simulation result takes into account mainly the geometric factor and uses two photons being emitted at 180 degrees to each other.
Fig. 6. Diagram of two MVT hits from a strip-line. The position, time and pulse-height of a single event are determined from the time stamps of the two MVT hits (see text for details).
between two adjacent LYSO/SiPM pixels was measured to be about 340 ps. This is consistent with the signal propagation delay on 25 mm trace, assuming a signal propagation speed 0.5 𝑐 on the strip-line (𝑐 is the speed of light in a vacuum). 3.1.2. Pulse-height spectra The PHS of individual LYSO/SiPM was obtained after events were assigned to crystals based on the 𝛥T measurement as described above. Fig. 8 shows the resulting pixel-level PHS obtained for a selected stripline. Fig. 9 shows the pulse-height corresponding to 511 keV and the energy resolution (FWHM) at 511 keV for all 32 pixels of a selected detector module, derived from pixel-level PHS as illustrated in Fig. 8. The gain variations (RMS) within detector modules are found to be 5%– 10%, and the pixel-level energy resolution at 511 keV is measured to be 9%–13% for the PET insert, which contains 448 pixels in total. The correction for possible SiPM saturation was not applied in the energy resolution.
3.2. Initial phantom imaging Figs. 12(a) shows an image of two 22 Na point sources placed on the horizontal axis. The two point sources, located at + 2.5 mm and −2.5 mm, were separated by 5.0 mm. The intensity of the image on the axis is shown on Fig. 12(d). 24 M coincidence events were collected and reconstructed by using the MLEM algorithm as described in Section 2.5. The average width of the intensity profiles was measured to be 1.5 mm in FWHM by applying a Gaussian fit. The transverse image of the 68 Ge line source and a line profile of the image along the radial direction are shown in Figs. 12(b) and 12(e). A total of 34 M coincidence events were used for reconstruction. The 68 Ge source position and width in the reconstructed images are obtained by Gaussian fit and summarized in Table 1. The positions of the 68 Ge source measured from the image match well with the actual positions. The width of the source image increased as the source moves away from the center due to the parallax error effect. At the trans-axial center, the 2.1 mm FWHM obtained for the 68 Ge line source was greater than that obtained for the 22 Na source. This may be attributed to the size of the line source: from an X-ray image of the source, the ID of the 68 Ge source container was measured to be 1.9 mm. Figs. 12(c), (f) and (g) show the transverse image of the resolution phantom and line profiles through the centers of the 1.8 mm and 2.6 mm diameter holes. Some of the 1.8 mm diameter
3.1.3. Coincidence time resolution For selecting coincidence events, a 410–610 keV energy window was applied. Fig. 10(a) shows a histogram of the coincidence differential time measured between two selected detectors, showing a coincidence resolving time (CRT) of approximately 495 ps. Fig. 10(b) shows the histogram of the 128 CRTs measured for all possible pairs in the selected two DMs. This histogram shows a distribution peaks at 500 ps and ranges from 380 ps to 620 ps. For the same 128 coincidence pairs, the CRT is also measured by varying the lowest threshold voltage used for the MVT from 30 mV to 90 mV in a 10 mV step. As described in Section 2.4, for coincidence filtering the event time was determined by the lowest threshold, which was set to 50 mV by default. The experiment producing Fig. 10(b) was repeated when varying the lowest threshold to obtain the CRT distribution at each lowest threshold setting. Fig. 10(c) shows the dependence of the resulting CRT on the MVT lowest threshold. Evidently, it shows that the CRT can be improved by lowering the lowest threshold. However, as the threshold is lowered
Fig. 7. (a) A pulse-height histogram obtained for a selected strip-line (sum of 16 uncalibrated PHS of the LYSO/SiPM pixels on the strip-line). (b) A histogram of differential time (𝛥T) on a strip-line. The position of each detector pixel is determined by 𝛥T.
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Fig. 8. Pulse-height spectra of 16 LYSO/SiPM detector pixels on a strip-line, after events are assigned to individual pixels using the method described in Section 3.1.1.
Fig. 9. (a) Pulse-heights corresponding 511 keV photo-peak for all detector pixels in a DM. (b) Energy resolutions (FWHM) at the photo-peak for the same detector pixels in (a).
Fig. 10. (a) Coincidence differential time between two detector pixels, showing a 495 ps CRT. (b) A histogram of CRTs measured for 128 coincidence pairs. (c) The measured CRT as a function of the lowest MVT threshold voltage.
Fig. 11. (a) Measured count rate of the MVT board as a function of radioactivity in μCi, fitted by a 2nd order polynomial. (b) The sensitivity of the PET insert along the axial direction.
holes, especially those in the central region where parallax errors are not strong, can be resolved in the image and profiles. This suggests that 1.8 mm spatial resolution may be achievable in the central region of the
insert. 68 M coincidence events were used to generate the resolution phantom image. It is worth noting that these imaging results were obtained without corrections for attenuation and scatter; therefore, 6
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Fig. 12. Phantom images acquired outside the MRI scanner and line profiles: (a) and (d) for the source placed at various radial positions, (c), (f) and (g) for the resolution phantom. Table 1 The positions and widths of the shown in Fig. 12(e).
68
Position in image (mm)
FWHM in image (mm)
0 5 10 15 20 25
0.0 5.1 9.9 15.0 20.0 25.2
2.1 2.1 2.1 2.3 3.1 3.3
Na point source at two positions 5.0 mm apart, (b) and (e) for the
68
Ge line
data acquired by the DRS4 sampler, which has ∼5 ps FWHM electronic time resolution. The leading edge discriminator with 3–4 mV threshold was applied on the rising part of waveforms for the CRT measurement. At the same bias voltage, the CRT measured inside the MR scanner was larger than that measured outside. However, a comparable CRT can be obtained by increasing the SiPM bias voltage by ∼0.5 V when running the PET insert inside the MRI scanner. The SiPM bias voltage increase may cause the dark counts to increase. However, the noise level of the DMs was observed to be the same (∼0.5 mV RMS) after the bias voltage increased. This indicates that the effect due to SiPM dark counts increase is negligible with 0.5 V increase in the bias voltage. We are currently conducting more studies to understand the mechanism responsible for pulse-height reduction inside the magnetic field. The resolution phantom images were obtained by acquiring data after placing the PET insert inside the MR scanner, and were compared to the image obtained when it was outside. Two data sets were acquired to examine the effects of MR pulsing: one was collected without any MR pulsing, and the other was collected during MRI data acquisition that used a gradient echo pulse sequence with TR/TE = 350/3 ms. Figs. 14(a) and 14(b) show the images obtained from the two data sets, each containing 30 M coincidence events. Sample line profiles of the images are shown in Figs. 14(c) and 14(d). We do not observe significant differences between these two images and intensity profiles, suggesting that the effects due to the particular MR pulsing examined is insignificant. Also, the 1.8 mm diameter holes are resolved, consistent with the expected intrinsic resolution of 1.6 mm of the PET insert. The experiments shown in Figs. 12 and 14 were not conducted on the same day due to the difficulties in setting up the experiment in the MRI room. However, the results in Fig. 14 do show the similar patterns with those in Figs. 12(c) and 12(f). We do observe elevated activities in holes at the upper right region of the phantom. We suspect that this is due to filling artifacts of the 18 F solution.
Ge line sources derived by Gaussian fit to the profiles
Source position (mm)
22
they represent the base-line image quality we can expect to obtain with the PET insert. 3.3. PET/MR compatibility 3.3.1. Effects on PET due to MR When the PET insert was placed inside the MR scanner, the PET signal pulse height decreased. While the MVT board is a practical solution, it provides limited information about the PET signals. Therefore, for studying more precisely the influence of the MR magnetic fields on the performance of the PET insert, we used a 4-channel DRS4 evaluation board to capture a full PET signal waveform at 2.5 GS/s sampling rate for a duration of 400 ns. Fig. 13(a) shows the averaged PET signal waveforms when the PET insert is placed inside and outside the MRI scanner. Only 511 keV events were used to obtain the average. There are no noticeable changes in the waveform shape other than reduction in amplitude. The pulse-height of a strip-line was also measured by placing the PET insert at different axial locations away from the magnet center to examine the effect of the magnetic field strength. Fig. 13(b) shows the relative pulse-height of a strip-line with respect to that measured outside the MRI scanner. The result shows that pulse-height reduction is affected by magnetic field strength: ∼15% of pulse-height reduction was observed close to the magnet center and the reduction started to decrease at 60 cm from the center and became negligible at 100 cm. CRT of the PET insert was also found to degrade inside the MRI scanner, which we postulate is due to the reduced pulse height and therefore may be compensated by increasing the signal gain. Fig. 13(c) shows the CRT measured from waveform
3.3.2. Effects on MR due to PET An MR image of the resolution phantom filled with 18 F solution acquired during simultaneous PET/MR operation is shown in Fig. 15(a). No image artifacts are observed. The potential effects on MR imaging due to the PET insert were studied more carefully by using an uniform cylinder (OD = 2.5 cm, length = 12 cm) filled with water. Two different MR pulse sequences, including gradient echo and spin 7
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Fig. 13. (a) Signal waveforms of the PET insert placed inside and outside the MRI scanner. (b) The relative pulse-height measured at different location along the MR bore. (c) Coincidence resolving time as a function of bias voltage over the breakdown.
Fig. 14. Resolution phantom images acquired inside the MRI scanner and line profiles: (a) and (c) when only the PET insert is running, and (b) and (d) while both the PET insert and the MR scanner are running.
in the middle column. The PET image was reconstructed by using 1.2 M coincidence events that are collected over 15 min. Qualitatively, there is a good match between the PET and MR images.
echo, were employed for comparison. Figs. 15(b) and 15(c) show the images obtained with the spin echo sequence. Figs. 15(d) and 15(e) show axial total-intensity profiles (by summing pixel intensities in an image slice) obtained before and after placing the PET insert inside the MR scanner. The signal cables of the PET insert are on the left-hand side of the profiles shown, and the power cables are on the right-hand side. When the PET insert was placed inside the MR scanner, while the profiles remain approximately the same on the signal-cable side, their amplitude decreases on the power-cable side. This reduction is larger in the case of the gradient echo sequence. The profiles of the signal-tonoise ratio (S/N) and uniformity calculated as described in [32] also have similar degradation on the power-cable side (results not shown). We suspect the reduction is caused by the biasing cables and connectors on the PET insert, which appear to affect gradient field switching.
4. Discussion When the PET insert is present, we do observe some effects to MRI, especially when using gradient echo pulse sequences. We are investigating which components of the PET insert that are responsible. In particular, we have placed the PET insert inside the MR scanner with or without the biasing cables. By comparing the resulting MR image profiles, we have strong evidence that suggests the effects are due to the cables and connectors. We postulate that the gradient field switching induces eddy currents in the cables and connectors to affect the magnetic field uniformity. Moreover, unlike the signal cables, the biasing cable is close to the detector center; this proximity might also cause asymmetric effects on MR imaging. Currently, we use separate cables for signal transmission and biasing so that the prototypical PET insert is easier to construct and trouble shoot. We may, however, be able to avoid biasing cables and connectors in our next design by making all cable connections on the same side of the detector module. Also, we may employ flat flexible cables for the signal and biasing cables so the PET insert is more practical to use. Currently, our PET insert is completely unshielded in order to evaluate the intrinsic PET/MR mutual interference of the proposed technology. The aforementioned
3.4. Initial simultaneous PET/MR mouse imaging A PET/MR image of a mouse brain is acquired by running the PET insert and MR scanner simultaneously. An initial activity of 180 μCi 18 F-fluorodeoxyglucose (FDG) was administrated to a healthy 30 gram male mouse through intraperitoneal (IP) injection, and data acquisition started 80 min post FDG injection. A spin echo pulse sequence with TR/TE = 4000/25 ms was used for MR imaging. Fig. 16 shows the resulting images of the mouse brain (MR images shown in gray-scale and PET images in color), and the co-registered PET/MR images are shown 8
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Fig. 15. (a) A transverse view of the MR image of the resolution phantom. (b) A transverse and (c) sagittal view of the MR image of the uniform phantom filled with water. Signal profiles in MR images along the axial slice are shown for (d) gradient echo sequence and (e) spin echo sequence.
Fig. 16. PET/MR combined images of a mouse brain. The left-hand side shows the trans-axial and sagittal views of the image, and the right-hand side shows the coronal views at the two different slices. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
effects may be addressed by applying adequate but minimal detector shielding utilizing insights gained in this study. Due to the use of 3.2 mm detector pixels, our PET insert currently has limited resolution for small animal imaging. To improve the spatial resolution, we can decrease the detector pixel size to 2 mm or smaller, to achieve an image resolution of about 1 mm or better. Also, in this work we have focused our attention on the instrumentation aspect of the insert. Therefore, we have only developed a most basic MLEM algorithm that is based on Siddon raytracing and incorporates prereconstruction corrections of detector-efficiency non-uniformity and random events. Filtered back projection algorithm will be implemented in image reconstruction as a basis for measuring the image spatial resolution in accordance to NEMA standards. It is well known that the image reconstruction process can critically impact both the resolution and statistical properties of the resulting images. In future work, we will employ more advanced image reconstruction algorithms that incorporate accurate modeling of the imaging physics. While TOF is not important for small animal imaging, we seek to develop PET insert for clinical PET/MR imaging; and, therefore, in this work we are also interested in examining the TOF capability with the proposed technology. In this study, by use of MVT sampling DAQ, we obtain a CRT of about 500 ps. However, Fig. 13(c) shows that the PET insert can achieve ∼270 ps when using the DRS4 waveform sampler. Previously, Zeng et al. [33] reported a ∼300 ps CRT for SiPM
based detector when using the MVT DAQ. Therefore, we believe that it is possible to improve the CRT of the PET insert toward 300 ps when optimizing the MVT DAQ board for our detector modules. To further improve the CRT, we can modify the design to reduce the baseline noise level of the MVT board and increase the preamplifier (preamp) gain at the signal input. Currently, the preamp gain is set to 6 by the manufacturer, not optimized for our particular detector modules with MPPCs with strip-line readout. In this study, we calculate the CRT based on the time stamps obtained at each of the four MVT thresholds (50, 150, 300 and 600 mV) and observe that the best result is obtained with 50 mV. However, better results might be possible by using some extrapolation methods that use multiple time stamps. The current thresholds setting is only empirical. Systematic investigation seeking to optimize the MVT thresholds is underway. 5. Summary We have prototyped a PET insert for simultaneous PET/MRI. The PET insert consists of 14 detector modules arranged in a cylindrical shape to fit within a 9.4 T small animal MRI scanner. The prototypical PET insert adopts a strip-line signal multiplexing method and multi-voltage threshold waveform sampling previously proposed and demonstrated by us. These methods allow us to significantly reduce the readout channel number and electronics complexity while still 9
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maintaining high detection performance. Performance of the detector modules and the PET insert, and the mutual effects between the PET insert and MR scanner are experimentally evaluated. Initial imaging studies of resolution phantom and a healthy mouse are conducted. Our initial results are encouraging, but there are several issues that require further detailed investigations in future studies, which include minimizing the mutual interference between the PET and MRI systems with an effective shielding.
[8] D. Renker, Geiger-mode avalanche photodiodes, history, properties and problems, Nucl. Instrum. Methods Phys. Res. A 567 (2006) 48–56. [9] B. Pichler, et al., Performance test of an LSO-APD detector in a 7 T MRI scanner for simultaneous PET/MRI, J. Nucl. Med. 47 (2006) 639–647. [10] J. Kang, et al., A small animal PET based on GAPDs and charge signal transmission approach for hybrid PET-MR imaging, J. Instrum. 6 (2011) P08012. [11] H. Yoon, et al., Initial results of simultaneous PET/MRI experiments with an MRI-compatible silicon photomultiplier PET scanner, J. Nucl. Med. 53 (4) (2012) 608–614. [12] J. Wehner, et al., PET/MRI insert using digital SiPMs: investigation of MR-compatibility, Nucl. Instrum. Methods Phys. Res. A 734 (2014) 116–121. [13] A.L. Goertzen, et al., First results from a high-resolution small animal SiPM PET insert for PET/MR imaging at 7 T, IEEE Trans. Nucl. Sci. 63 (2016) 2424–2433. [14] G. Ko, et al., Evaluation of a silicon photomultiplier PET insert for simultaneous PET and MR imaging, Med. Phys. 43 (2016) 72–83. [15] N. Omidvari, et al., PET performance evaluation of MADPET4: a small animal PET insert for a 7T MRI scanner, Phys. Med. Biol. 62 (2017) 8671–8692. [16] B. Pichler, H.F. Wehrl, A. Kolb, M.S. Judenhofer, PET/MRI: The next generation of multimodality imaging, Semin. Nucl. Med. 38 (3) (2008) 199–208. [17] H. Zaidi, A.D. Guerra, An outlook on future design of hybrid PET/MRI systems, Med. Phys. 38 (10) (2011) 5667–5689. [18] S. Vandenberghe, P.K. Marsden, PET-MRI: a review of challenges and solutions in the development of integrated multimodality imaging, Phys. Med. Biol. 60 (2015) R115–R154. [19] Y. Wang, et al., Design and performance evaluation of a compact, large-area PET detector module based on silicon photomultiplier, Nucl. Instrum. Methods Phys. Res. A 670 (2012) 49–54. [20] A.L. Goertzen, et al., Design and performance of a resistor multiplexing readout circuit for a SiPM detector, IEEE Trans. Nucl. Sci. 60 (2013) 1541–1549. [21] H. Choi, Y. Choi, D.J. Kwak, J. Lee, Prototype time-of-flight PET utilizing capacitive multiplexing readout, Nucl. Instrum. Methods Phys. Res. A 921 (2019) 43–49. [22] H. Kim, et al., A silicon photo-multiplier signal readout using strip-line and waveform sampling for position emission tomography, Nucl. Instrum. Methods Phys. Res. A 830 (2016) 119–129. [23] H. Kim, et al., A systematic study on the strip-line readout method for SiPMbased TOF PET, in: IEEE NSS/MIC Conference Record, 2017, http://dx.doi.org/ 10.1109/NSSMIC.2017.8533065. [24] H. Kim, et al., A feasibility study of a PET/MRI insert detector using strip-line and waveform sampling data acquisitioin, Nucl. Instrum. Methods Phys. Res. A 784 (2015) 557–564. [25] D. Xi, et al., FPGA-only MVT digitizer for TOF PET, IEEE Trans. Nucl. Sci. 60 (2013) 3253–3261. [26] S. Ritt, R. Dinapoli, U. Hartmann, Application of the DRS chip for fast waveform digitizing, Nucl. Instrum. Methods Phys. Res. A 623 (2010) 486–488. [27] X. Mei, et al., A 72-channel FPGA-only MVT digitizer board and a micro-system for coincidence detection/imaging, in: IEEE NSS/MIC Conference Record, 2014, http://dx.doi.org/10.1109/NSSMIC.2014.7431224. [28] S13361-3050NE-04 Specifications, https://www.hamamatsu.com/jp/en/product/ type/S13361-3050NE-04/index.html. [29] SensL application note on readout methods for arrays of SiPM http://sensl.com/ downloads/ds/TN-Readout_Methods_for_Arrays_of_SiPM.pdf. [30] L. Shepp, Y. Vardi, Maximum likelihood reconstruction for emission tomography, IEEE Trans. Med. Imaging 1 (2) (1982) 113–122. [31] R. Siddon, Fast calculation of the exact radiological path for a three-dimensional CT array, Med. Phys. 12 (2) (1985) 252–255. [32] H.F. Wehrl, Assessment of MR compatibility of a PET insert developed for simultaneous multiparametric PET/MR imaging on an animal system operating at 7 T, Magn. Reson. Med. 65 (2011) 269–279. [33] C. Zeng, et al., Timing optimization for digital PET detector module based on FPGA-only MVT digitizers, in: IEEE NSS/MIC Conference Record, 2014, http://dx.doi.org/10.1109/NSSMIC.2014.7430851.
Declaration of competing interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. CRediT authorship contribution statement H. Kim: Conceptualization, Software, Formal analysis, Investigation, Writing - original draft. Y. Hua: Resources, Investigation. H.-T. Chen: Software. H.-M. Tsai: Software, Visualization. C.-T. Chen: Resources, Funding acquisition. G. Karczmar: Resources, Funding acquisition. X. Fan: Resources, Investigation. D. Xi: Resources, Investigation. Q. Xie: Resources, Funding acquisition. C.-Y. Chou: Resources, Funding acquisition. C.-M. Kao: Conceptualization, Writing - review & editing, Resources, Funding acquisition. Acknowledgments This work was supported in part by the University of Chicago Medicine Comprehensive Cancer Center pilot award, United States; Ministry of Science and Technology of Taiwan #MOST 107-2321-B055-004, ; the National Natural Science Foundation of China (NSFC) Grant #61425001, #61671215, #61604059, the National Key Scientific Instrument and Equipment Development Project of China #2013YQ030923, and the National Key Research and Development Program #2016YFF0101501. The Hamamatsu Corporation generously provided MPPC arrays used in this study. The authors thank Inna Gertsenshteyn for proof-reading the manuscript. References [1] T. Jones, Molecular imaging with PET – the future challenges, Br. J. Radiol. 75 (2002) S6–S15. [2] S.M. Ametamey, M. Honer, P.A. Schubiger, Molecular imaging with PET, Chem. Rev. 108 (2008) 1501–1516. [3] Y. Cossuin, A. Hocq, P. Gillis, Q. Vuong, Physics of magnetic resonance imaging: from spin to pixel, J. Phys. D: Appl. Phys. 43 (2010) 213001. [4] Y. Shao, et al., Simultaneous PET and MR imaging, Phys. Med. Biol. 42 (1997) 1965–1970. [5] R. Raylman, et al., Simultaneous MRI and PET imaging of a rat brain, Phys. Med. Biol. 51 (2006) 6371–6379. [6] S. Yamamoto, Design and performance from an integrated PET/MRI system for small animals, Ann. Nucl. Med. 24 (2010) 89–98. [7] P. Buzhan, et al., Silicon photomultiplier and its possible applications, Nucl. Instrum. Methods Phys. Res. A 504 (2003) 48–52.
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Contents lists available at ScienceDirect
Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
Design, evaluation and initial imaging results of a PET insert based on strip-line readout for simultaneous PET/MRI H. Kim a ,∗, Y. Hua b , H.-T. Chen a , H.-M. Tsai a , C.-T. Chen a , G. Karczmar a , X. Fan a , D. Xi b , Q. Xie b , C.-Y. Chou c , C.-M. Kao a a
Department of Radiology, University of Chicago, Chicago, IL 60637, USA Biomedical Engineering Department, Huazhong University of Science and Technology, Wuhan, China c Department of Bio-Industrial Mechatronics Engineering, National Taiwan University, Taipei, Taiwan b
ARTICLE
INFO
Keywords: Positron Emission Tomography Magnetic Resonance Imaging Strip-line readout Sampling
ABSTRACT We present the development of a PET insert system for potential simultaneous PET/MR imaging using a 9.4 T small animal MRI scanner to test our system. The detectors of the system adopt a strip-line based multiplexing readout method for SiPM signals. In this readout, multiple SiPM outputs in a row share a common strip-line. The position information about a hit SiPM is encoded in the propagation time difference of the signals arriving at the two ends of the strip-line. The use of strip-lines allows us to place the data acquisition electronics remotely from the detector module to greatly simplify the design of the detector module and minimize the mutual electromagnetic interference. The prototype is comprised of 14 detector modules, each of which consists of an 8x4 LYSO scintillator array (each LYSO crystal is 3x3x10 mm3 ) coupled to two units of Hamamatsu MPPC arrays (4x4, 3.2 mm pitch) that are mounted on a strip-line board. On the strip-line board, outputs of the 32 SiPMs are routed to 2 strip-lines so that 16 SiPM signals share a strip-line. The detector modules are installed inside a plastic cylindrical supporting structure with an inner and outer diameter of 60 mm and 115 mm, respectively, to fit inside a Bruker BioSpec 9.4 T MR scanner. The axial field of view of the prototype is 25.4 mm. The strip-lines were extended by using 5-meter cables to a sampling data acquisition (DAQ) board placed outside the magnet. The detectors were not shielded in the interest of investigating how they may affect and be affected by the MRI. Experimental tests were conducted to evaluate detection performance, and phantom and animal imaging were carried out to assess the spatial resolution and the MR compatibility of the PET insert. Initial results are encouraging and demonstrate that the prototype insert PET can potentially be used for PET/MR imaging if appropriate shielding will be implemented for minimizing the mutual interference between the PET and MRI systems.
1. Introduction Simultaneous PET/MR imaging combines the complementary merits of PET and MRI; the former can provide functional imaging with high sensitivity and specificity [1,2], and the latter can produce anatomical imaging at high spatial resolution [3]. Compared to CT, MRI has higher soft-tissue contrast, no radiation dose, and diverse contrast mechanisms through various pulse sequences. Therefore, PET/MRI has been considered as an alternative to PET/CT in many clinical and pre-clinical imaging applications. MRI uses strong magnetic fields (3–15 T), fast switching gradient fields, and radio-frequency (RF) signals for image formation. These fields and RF signals are known to affect the operations of a PET detector and its electronics. On the other hand, the existence of PET inside the MR scanner affects the magnetic field homogeneity and
can produce RF signals that interfere with MRI. Therefore, building a PET/MRI system requires MR-compatible detector materials and system design to maintain proper functions of both modalities. Various PET/MRI systems have been developed based on different detector technologies to overcome these challenges. Earlier developments used long optical fibers to separate the scintillators, which are inside the magnet, and photo-sensors [4–6] such as the photo-multiplier tube (PMT), which are outside the magnet, to eliminate the strong magnetic effects and constraints of space limitation of MR. The long optical fibers, however, induced significant loss of the scintillation lights. With the advent of compact and magnetic field in-sensitive photosensors such as the avalanche photo-diode and silicon photo-multiplier (SiPM) [7,8], it became possible to integrate the photo-sensors with the detector modules that are inserted inside the magnet. This design approach has been adopted by many research groups [9–15]. At present,
∗ Corresponding author. E-mail address: [email protected] (H. Kim).
https://doi.org/10.1016/j.nima.2020.163575 Received 15 April 2019; Received in revised form 15 January 2020; Accepted 31 January 2020 Available online 5 February 2020 0168-9002/© 2020 Elsevier B.V. All rights reserved.
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SiPM has emerged as the photo-sensor of choice for PET, particularly for PET/MRI, because of its in-sensitivity to magnetic fields, compact size, and PMT-compatible signal amplification. Excellent reviews of the current development status can be found in the literature [16–18]. Commercial SiPMs are available in 1 × 1 mm2 to 6 × 6 mm2 pixel sizes. Therefore, a practical PET system that typically employs a large number of SiPMs requires an efficient signal multiplexing scheme for readout [19–21]. Most proposed multiplexing readouts, however, significantly compromise the timing characteristics of detector. We have proposed and successfully demonstrated a multiplexing readout method for SiPM-based time-of-flight (TOF) PET [22]. In this method, multiple SiPMs share a single strip-line that is read at the two ends. The position of the hit SiPM on the strip-line is decoded from the time difference between the SiPM output signals arriving at the two ends. Our recent study shows that a single strip-line can connect 32 SiPMs without compromising the detector performance, achieving ∼270 ps coincidence time resolution [23]. With strip-line readout, data acquisition (DAQ) electronics can be separated from the detector module and located remotely via long cables that extend the strip-lines. As a result, the mechanic, electronic and RF shielding designs of the detector may be greatly simplified; the detector can be kept highly compact with minimal materials, which also help to minimize the impact on the magnetic field uniformity. Therefore, we believe PET detectors based on strip-line readout are favorable for developing PET/MRI systems. PET insert systems that adopted long signal transmission cables for coping the aforementioned challenges in developing PET inserts, which was similar to our approach, have been reported in [10,15]. Previously, we have developed detector modules with strip-line readout and demonstrated that the technology is feasible for PET/MRI [24]. Here, we prototyped a PET insert system, adopting PET detectors with strip-line readouts, in combination with the use of a compact 72-channel electronics board that provides an FPGA based sampling DAQ that we previously proposed [25]. While our detectors are TOF capable and have limited resolution for small animal imaging, we tested the technology on a 9.4 T preclinical MRI scanner due to accessibility of the scanner and the more challenging electromagnetic environment and space condition it presented. Experimental tests were conducted to measure the performance of the prototypical PET insert inside and outside the MR scanner. A resolution phantom and uniform phantom were used to assess the MR compatibility of the PET insert. In this paper, we report the design and evaluation of the PET insert, and initial imaging results using the PET insert.
2.1. Detector module
2. Materials and methods
The PET insert was constructed by assembling 14 detector modules in a plastic supporting structure that was designed to fit inside a small animal MR scanner, Bruker 9.4 T BioSpec 94/30 USR (Bruker Biospin, Billerica, MA, US). The inner (ID) and outer diameter (OD) of the PET insert are 6 cm and 12 cm, respectively; they were determined from the MR scanner’s bore size and a RF coil used in this study. The 14 DMs were arranged with the support structure so that the distance between two oppositely facing DMs was 62 mm, and the gap between two adjacent DMs was 1.5 mm. The axial field of the PET insert is 25.4 mm. Figs. 2(a) and 2(b) show the fully assembled PET insert. 5 m long mini coaxial cables (RG-316) were used to transmit the outputs to the DAQ board, which was located outside the MRI room to minimize interference with MR operations. A total of 56 MMCX cables, 4 for each detector module, were used. The bias voltages to the 14 DMs were adjusted in a custom-made voltage distributor that employed LM337 voltage regulators and two units of Lambda GEN750W power supplies. A total of 14 6 m-length cables (24 gauge four conductor stranded) were used to connect the DMs and the power distributor, which was also placed outside the MRI room. All the signal and power cables passed through a penetration panel to be connected to the equipment located outside the MRI room. As shown in Fig. 2, no detector shielding is used in current implementation in the interest of simplifying the design of the PET insert.
The detector module (DM) consists of LYSO scintillator arrays, Hamamatsu Multi-Pixel photon counter (MPPC) arrays, and a stripline board (SLB) as shown in Fig. 1. Two 4 × 4 LYSO arrays (each LYSO crystal is 3 × 3 × 10 mm3 and the crystal pitch is 3.2 mm) are packed together to form a 4 × 8 LYSO array, shown in Fig. 1(b). The LYSO crystals (Shanghai EBO Optoelectronics Technology, China) are coupled one-to-one to SiPM pixels by using optical glue (Cargille Meltmount, refractive index 1.582). Enhanced specular reflectors (ESR, 3M™) are used between LYSOs in the arrays for optical isolation. Hamamatsu MPPC arrays, containing 4 × 4 pixels (each MPPC pixel has a detection surface area of 3 × 3 mm2 with a 3.2 mm pitch that matches the LYSO crystal pitch), were selected for the photo-sensor because of their superior gain uniformity. A SiPM array board shown in Fig. 1(c) was manufactured to house two MPPC arrays, provides bias voltage to SiPMs and routes the outputs of the 32 SiPMs to the SLB via connectors. Due to the rapid development cycle of MPPC, two different models of the MPPC array were used in the PET insert, including 8 units of S12642-0404-PA and 20 units of S13361-3050NE [28]. The main difference between these models is their breakdown voltages, which are measured to be about 65 V and 53 V for S12642 and S13361, respectively. SiPMs in a detector module use the same bias voltage. For each detector module, the bias voltage was adjusted to be about 2.5 V above the module’s breakdown voltage so that all modules produce similar pulse heights for 511 keV photons. The detector was operated at room temperature maintained at 18 ◦ C to avoid SiPM gain variation due to temperature change. The SLB is an 18 × 110 mm2 printed circuit board (PCB) containing two strip-lines, one laid out on the top side and another at the bottom side. Each strip-line receives the outputs of 16 SiPMs so that a total of 32 SiPM outputs of two MPPC arrays are routed into two strip-lines. Fig. 1(d) depicts a stripline layout across 16 SiPM outputs. The strip-line was configured to have 50 ohm characteristic impedance, and, as shown in Fig. 1(e), the trace length between adjacent SiPM outputs was set to 25 mm in order to increase the signal propagation time. Because position decoding on strip-line is based on the measured time difference, this alleviated the requirement on the time accuracy for the DAQ. Fig. 1(f) shows the connection scheme of the SiPM output (negative polarity) to the stripline. Schottky diodes were used for decoupling the capacitances of the SiPMs [29] that shared a strip-line. The ends of each strip-line were connected to micro-miniature coaxial (MMCX) connectors, mounted at the edge of SLB, for transferring the output signals to DAQ. 2.2. PET insert assembly
The prototypical PET insert employs 14 detector modules. Each module has 2 strip-lines and therefore 4 outputs, yielding a total of 56 outputs for the PET insert. For DAQ, a DRS4 waveform sampler [26] was used in our previous evaluation studies of detector modules. In this study, we used the DSR4 sampler again, specifically the 4-channel DRS4 evaluation board, for the purpose of validating and evaluating individual detector modules of the PET insert. For practical use, however, it is beneficial to simplify the DAQ electronics. Hence, for the developed PET insert, we used another sampling DAQ technology that we previously proposed and developed for scintillation detectors. This technology, called multi-voltage threshold (MVT) [25], samples a waveform with respect to up to 4 user-defined voltage thresholds and can be implemented by using FPGAs. As a result, multi-channel DAQ boards capable of supporting TOF detection can be readily developed and modified. In this work, we employed a single 10 × 18 cm2 72channel MVT DAQ board [27] to handle the 56 outputs of the PET insert. In Sections 2.1–2.3, the implementation of the strip-line readout in the detector module, overall design features of the PET insert, and DAQ system are described in detail. In Sections 2.4 and 2.5, we describe the experimental setups and processing the acquired data, respectively. 2
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Fig. 1. Components of a DM of the PET insert. (a) A strip-line board (SLB). (b) LYSO arrays coupled to (c) a SiPM array board. (d) A strip-line layout connecting 16 SiPMs output. (e) A close-up of the strip-line on SLB. (f) A schematic of a SiPM output coupling to a strip-line through diodes.
Fig. 2. (a) A cross-sectional view and (b) a side view of the PET insert after all the detector modules were assembled into the supporting structure. The PET insert assembly is not shielded for RF signals.
2.3. Multi-voltage threshold waveform sampling
2.4. Experimental setup
Digitization of the signal outputs of the PET insert was accomplished in a board implementing MVT waveform sampling technology [25] that we have previously developed. In MVT, the input signal is passed to four comparators with user-defined reference voltages. The leading and trailing transition times of the outputs of the comparators are recorded by time-to-digital converters (TDC). Subsequently, the waveform of the input signal is reconstructed from time stamps and the known reference voltages used. An example of the MVT operation is shown in Fig. 3(a). In this case, the blue circles show the samples obtained by using 4 reference voltages, including 50, 150, 300 and 600 mV, and the red solid curve shows the reconstructed waveform obtained by linearly interpolating the samples. The pulse-height of the reconstructed waveform was estimated by integrating the area under the solid curve. Waveform data acquired by a DRS4 sampler were used to check the accuracy of the energy estimation. The area integration method results in ∼10.6% energy resolution at 511 keV photo-peak. On the other hand, ∼8.5% of energy resolution was measured by using the maximum amplitude of the waveforms. For comparison, the dashed curve shows the reconstructed waveform obtained by using a bi-exponential fitting method we previously reported [25]. Previously we have demonstrated that both the comparators and the TDCs can be implemented by using FPAGs that are widely used in modern electronics. Fig. 3(b) shows the FPGA (Altera EP4C115F29C7) based MVT board used in this study (manufactured by Raycan Technology Co., China). This board provides 72 input channels, each of which has 4 comparators. The threshold voltages (0–1 V) of the comparators were programmably configured by using digital-to-analog converters (DAC) (DAC7668, Texas Instruments). The bin width of the TDCs implemented on the board was calibrated to be 95 ps. Because the MVT uses flat flexible cables (FFC) for inputs, an adapter board was built to convert the MMCX connectors from the PET insert to the FFC connectors. The time stamps produced by the MVT were packed into user datagram protocol (UDP) packets and transferred to a computer through a Gigabit Ethernet interface.
The DMs were individually tested on a bench top setup before they were assembled into the PET insert. As already mentioned, a DRS4 waveform sampler that can provide 5 GHz sampling was used in order to obtain accurate performance assessment for the DMs. Positiondecoding accuracy, energy resolution, and coincidence time resolution were characterized. The breakdown voltage of each DM was determined from pulse-height measurements by varying the bias voltage. For measuring the coincidence time resolution, a reference detector, consisting of a Hamamatsu MPPC S10931-050P (3 × 3 mm2 active area) coupled to a 3 × 3 × 10 mm3 LYSO crystal, was used. Measurements were performed by placing a 68 Ge line source (sealed in a cylindrical stainless tube with 19 cm length and 3.2 mm (1.9 mm) of OD (ID)) or a 22 Na point source (∼0.5 mm diameter sphere sealed in a 2.5 cm diameter plastic disk) between the reference detector and the DM under evaluation. After the PET insert was fully assembled, the above performance measurements for the DMs were repeated by using the 72-channel MVT board. A custom-made resolution phantom filled with 18 F solution was also used to assess the spatial resolution of the PET insert. Fig. 4 shows the 68 Ge line source and the resolution phantom. For 68 Ge source imaging, the line source was moved along the radial direction in 5 mm step from the center of the FOV. The resolution phantom has six groups of holes of different diameters, including 1.0, 1.4, 1.8, 2.2, 2.6, and 3.0 mm. The distance between holes within each group is twice the hole diameter, and the depth of the holes is 30 mm. For experiments with MR, the PET insert was placed inside a Bruker 9.4 T BioSpec MR scanner (116 mm bore diameter). An actively shielded gradient coils (maximum strength 230 mT/m) was used. MR imaging was performed by using a 35 mm ID (59 mm OD) quadrature mouse volume coil (RAPID MR International, Columbus, OH, USA). Fig. 5 shows the arrangement of the PET insert, the RF coil, and the gradient coil inside the MR bore. Two MR pulse sequences were used for MR data acquisition: FLASH (fast low angle shot) T1-weighted gradient echo with TR/TE = 350/3 ms and flip angle = 30, and RARE (rapid acquisition with relaxation enhancement) T2-weighted fast spin echo with TR/TE = 4000/25 ms and RARE factor = 8. 3
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Fig. 3. (a) An example of waveforms reconstructed from the MVT stamps. (b) A 72-channel MVT board and an adapter board for MMCX to FFC conversion.
(see Fig. 3(a)) to reconstruct the waveform, and then the pulse height for the hit is obtained by integrating the area under the reconstructed waveform. The pulse height of the single event is taken as the average of the pulse heights of the two hits, calibrated to keV by using a 22 Na source. • Coincidence filtering: An energy window of 410–610 keV is used to select photo-peak singles events. For coincidence detection, the event time of a single is taken as the average of the first time stamps of the two hits corresponding to the single (the time stamps of the leading transitions of the lowest threshold, (T𝐿1 + T𝑅1 )/2). Based on this event time, two singles that are within a 2 ns time window are registered as a coincidence event. • Image reconstruction: A line-of-response (LOR) based maximum likelihood expectation maximization (MLEM) algorithm [30] is employed to reconstruct the resulting coincidence data. 30 iterations are used in the image reconstruction. The system response matrix is calculated by using Siddon’s ray tracing algorithm [31]. Currently, the algorithm does not implement corrections for attenuation and scatter. Therefore, the images shown are the baseline results that can be produced by the PET insert.
Fig. 4. (a) A 68 Ge line source in a positioning holder. (b) A resolution phantom has holes with diameter ranging from 1.0 mm to 3.0 mm. The holes are filled with 18 F solution for imaging.
2.5. PET data processing The raw UDP packets generated by the MVT board were saved on the hard drives of the DAQ computer for off-line analysis. The basic element of the UDP packet was one MVT hit, which contained the time stamps of leading and trailing transitions (8 time stamps total) and the input channel number. Fig. 6 shows a simplified diagram of two MVT hits on a strip-line to produce a single event. T𝐿(𝑅)𝑖 in the figure represents the MVT time stamp of the 𝑖th voltage threshold on the two ends of a strip-line. The data analysis flow is described below.
3. Results 3.1. Detector performance 3.1.1. Position decoding by differential time on strip-line Fig. 7(a) shows a pulse-height spectrum (PHS) using singles collected from a selected strip-line. The gain variation of the LYSO/SiPM pixels sharing the strip-line broadens the photo-peak and even creates a small side peak. Fig. 7(b) shows the differential-time (𝛥T) histogram on a strip-line that illustrates the accuracy in position decoding. This particular 𝛥T histogram was obtained by using mainly events in the photo-peak identified on PHS in Fig. 7(a). The 𝛥T histogram clearly shows 16 peaks that reflect the positions of the 16 LYSO/SiPM pixels on the strip-line. A Gaussian fit was applied to each of the 16 peaks, and 𝛥T ranges corresponding to individual LYSO/SiPM pixels were determined by using the midpoints between the center positions of the resulting Gaussians. The average peak-to-valley ratio (P/V), which provides a figure of merit for the accuracy of position encoding, was about 13:1 for the entire PET insert. From the 𝛥T histogram, the time difference
• Single gamma event tagging: Following the 511 keV gamma interaction within the detector, two signals are formed on a stripline and propagate toward the ends of the strip-line, as depicted in Fig. 6. Therefore, two MVT hits appearing at the two ends of a strip-line within a 7 ns time window are tagged as a single event (∣T𝐿1 −T𝑅1 ∣ < 7 ns). In this step, the time difference is determined based on the leading transition time of the lowest threshold. • Position and pulse-height of a single event: The position of a single hit along the strip-line is determined from the time difference of the two registered MVT hits produced by the above step. In this case, however, all four time stamps of the leading transitions of each MVT hit are averaged to obtain the hit time (𝛥T = 𝛴T𝐿𝑖 −𝛴T𝑅𝑖 ). For each hit, its samples are linearly interpolated
Fig. 5. (a) The PET insert and the MVT board are connected by 5 m long MMCX cables. (b) The PET insert is installed into a Bruker 9.4 T small animal imaging MR scanner. (c) A cross-sectional view shows the arrangement of the PET insert, the RF coil, and the gradient coil inside the MRI scanner.
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some input channels fail to produce stable results due to noise. As a result, to ensure all channels function properly, we cannot decrease the lowest threshold below 50 mV with our present MVT board, yielding a system CRT of about 500 ps. 3.1.4. Count rate performance Count rate performance of the MVT board was measured by using the resolution phantom filled with 18 F solution with an initial activity of 265 μCi (most PET mouse imaging uses lower activities). Fig. 11(a) shows the MVT hit rate as a function of radioactivity as it decays to the background level over 12 h. The slight change in the count rate slope around 100 μCi may be due to the data writing load in the DAQ computer. As shown in Fig. 11(a), the count rate was fitted well by a 2nd order polynomial function 𝑦 = p2 𝑥2 + p1 𝑥 + p0 with p2 = −1.292 × 10−5 , p1 = 0.01428, and p0 = 0.04505. The sensitivity of the PET insert was measured by using 22 Na point source with 15 μCi radioactivity. The 22 Na source was moved along the central axis of the PET insert with a 2 mm step. Sensitivity was calculated by dividing the measured coincidence rates to the known activity of the source. Fig. 11(b) shows the sensitivity result (red curve), which shows a peak of about 1.1% at the axial center and steadily decreases toward the ends of the axial FOV. For comparison, the axial sensitivity profile calculated for the PET insert by using Geant4 simulation is also shown (black curve). Note that the simulation result takes into account mainly the geometric factor and uses two photons being emitted at 180 degrees to each other.
Fig. 6. Diagram of two MVT hits from a strip-line. The position, time and pulse-height of a single event are determined from the time stamps of the two MVT hits (see text for details).
between two adjacent LYSO/SiPM pixels was measured to be about 340 ps. This is consistent with the signal propagation delay on 25 mm trace, assuming a signal propagation speed 0.5 𝑐 on the strip-line (𝑐 is the speed of light in a vacuum). 3.1.2. Pulse-height spectra The PHS of individual LYSO/SiPM was obtained after events were assigned to crystals based on the 𝛥T measurement as described above. Fig. 8 shows the resulting pixel-level PHS obtained for a selected stripline. Fig. 9 shows the pulse-height corresponding to 511 keV and the energy resolution (FWHM) at 511 keV for all 32 pixels of a selected detector module, derived from pixel-level PHS as illustrated in Fig. 8. The gain variations (RMS) within detector modules are found to be 5%– 10%, and the pixel-level energy resolution at 511 keV is measured to be 9%–13% for the PET insert, which contains 448 pixels in total. The correction for possible SiPM saturation was not applied in the energy resolution.
3.2. Initial phantom imaging Figs. 12(a) shows an image of two 22 Na point sources placed on the horizontal axis. The two point sources, located at + 2.5 mm and −2.5 mm, were separated by 5.0 mm. The intensity of the image on the axis is shown on Fig. 12(d). 24 M coincidence events were collected and reconstructed by using the MLEM algorithm as described in Section 2.5. The average width of the intensity profiles was measured to be 1.5 mm in FWHM by applying a Gaussian fit. The transverse image of the 68 Ge line source and a line profile of the image along the radial direction are shown in Figs. 12(b) and 12(e). A total of 34 M coincidence events were used for reconstruction. The 68 Ge source position and width in the reconstructed images are obtained by Gaussian fit and summarized in Table 1. The positions of the 68 Ge source measured from the image match well with the actual positions. The width of the source image increased as the source moves away from the center due to the parallax error effect. At the trans-axial center, the 2.1 mm FWHM obtained for the 68 Ge line source was greater than that obtained for the 22 Na source. This may be attributed to the size of the line source: from an X-ray image of the source, the ID of the 68 Ge source container was measured to be 1.9 mm. Figs. 12(c), (f) and (g) show the transverse image of the resolution phantom and line profiles through the centers of the 1.8 mm and 2.6 mm diameter holes. Some of the 1.8 mm diameter
3.1.3. Coincidence time resolution For selecting coincidence events, a 410–610 keV energy window was applied. Fig. 10(a) shows a histogram of the coincidence differential time measured between two selected detectors, showing a coincidence resolving time (CRT) of approximately 495 ps. Fig. 10(b) shows the histogram of the 128 CRTs measured for all possible pairs in the selected two DMs. This histogram shows a distribution peaks at 500 ps and ranges from 380 ps to 620 ps. For the same 128 coincidence pairs, the CRT is also measured by varying the lowest threshold voltage used for the MVT from 30 mV to 90 mV in a 10 mV step. As described in Section 2.4, for coincidence filtering the event time was determined by the lowest threshold, which was set to 50 mV by default. The experiment producing Fig. 10(b) was repeated when varying the lowest threshold to obtain the CRT distribution at each lowest threshold setting. Fig. 10(c) shows the dependence of the resulting CRT on the MVT lowest threshold. Evidently, it shows that the CRT can be improved by lowering the lowest threshold. However, as the threshold is lowered
Fig. 7. (a) A pulse-height histogram obtained for a selected strip-line (sum of 16 uncalibrated PHS of the LYSO/SiPM pixels on the strip-line). (b) A histogram of differential time (𝛥T) on a strip-line. The position of each detector pixel is determined by 𝛥T.
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Fig. 8. Pulse-height spectra of 16 LYSO/SiPM detector pixels on a strip-line, after events are assigned to individual pixels using the method described in Section 3.1.1.
Fig. 9. (a) Pulse-heights corresponding 511 keV photo-peak for all detector pixels in a DM. (b) Energy resolutions (FWHM) at the photo-peak for the same detector pixels in (a).
Fig. 10. (a) Coincidence differential time between two detector pixels, showing a 495 ps CRT. (b) A histogram of CRTs measured for 128 coincidence pairs. (c) The measured CRT as a function of the lowest MVT threshold voltage.
Fig. 11. (a) Measured count rate of the MVT board as a function of radioactivity in μCi, fitted by a 2nd order polynomial. (b) The sensitivity of the PET insert along the axial direction.
holes, especially those in the central region where parallax errors are not strong, can be resolved in the image and profiles. This suggests that 1.8 mm spatial resolution may be achievable in the central region of the
insert. 68 M coincidence events were used to generate the resolution phantom image. It is worth noting that these imaging results were obtained without corrections for attenuation and scatter; therefore, 6
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Fig. 12. Phantom images acquired outside the MRI scanner and line profiles: (a) and (d) for the source placed at various radial positions, (c), (f) and (g) for the resolution phantom. Table 1 The positions and widths of the shown in Fig. 12(e).
68
Position in image (mm)
FWHM in image (mm)
0 5 10 15 20 25
0.0 5.1 9.9 15.0 20.0 25.2
2.1 2.1 2.1 2.3 3.1 3.3
Na point source at two positions 5.0 mm apart, (b) and (e) for the
68
Ge line
data acquired by the DRS4 sampler, which has ∼5 ps FWHM electronic time resolution. The leading edge discriminator with 3–4 mV threshold was applied on the rising part of waveforms for the CRT measurement. At the same bias voltage, the CRT measured inside the MR scanner was larger than that measured outside. However, a comparable CRT can be obtained by increasing the SiPM bias voltage by ∼0.5 V when running the PET insert inside the MRI scanner. The SiPM bias voltage increase may cause the dark counts to increase. However, the noise level of the DMs was observed to be the same (∼0.5 mV RMS) after the bias voltage increased. This indicates that the effect due to SiPM dark counts increase is negligible with 0.5 V increase in the bias voltage. We are currently conducting more studies to understand the mechanism responsible for pulse-height reduction inside the magnetic field. The resolution phantom images were obtained by acquiring data after placing the PET insert inside the MR scanner, and were compared to the image obtained when it was outside. Two data sets were acquired to examine the effects of MR pulsing: one was collected without any MR pulsing, and the other was collected during MRI data acquisition that used a gradient echo pulse sequence with TR/TE = 350/3 ms. Figs. 14(a) and 14(b) show the images obtained from the two data sets, each containing 30 M coincidence events. Sample line profiles of the images are shown in Figs. 14(c) and 14(d). We do not observe significant differences between these two images and intensity profiles, suggesting that the effects due to the particular MR pulsing examined is insignificant. Also, the 1.8 mm diameter holes are resolved, consistent with the expected intrinsic resolution of 1.6 mm of the PET insert. The experiments shown in Figs. 12 and 14 were not conducted on the same day due to the difficulties in setting up the experiment in the MRI room. However, the results in Fig. 14 do show the similar patterns with those in Figs. 12(c) and 12(f). We do observe elevated activities in holes at the upper right region of the phantom. We suspect that this is due to filling artifacts of the 18 F solution.
Ge line sources derived by Gaussian fit to the profiles
Source position (mm)
22
they represent the base-line image quality we can expect to obtain with the PET insert. 3.3. PET/MR compatibility 3.3.1. Effects on PET due to MR When the PET insert was placed inside the MR scanner, the PET signal pulse height decreased. While the MVT board is a practical solution, it provides limited information about the PET signals. Therefore, for studying more precisely the influence of the MR magnetic fields on the performance of the PET insert, we used a 4-channel DRS4 evaluation board to capture a full PET signal waveform at 2.5 GS/s sampling rate for a duration of 400 ns. Fig. 13(a) shows the averaged PET signal waveforms when the PET insert is placed inside and outside the MRI scanner. Only 511 keV events were used to obtain the average. There are no noticeable changes in the waveform shape other than reduction in amplitude. The pulse-height of a strip-line was also measured by placing the PET insert at different axial locations away from the magnet center to examine the effect of the magnetic field strength. Fig. 13(b) shows the relative pulse-height of a strip-line with respect to that measured outside the MRI scanner. The result shows that pulse-height reduction is affected by magnetic field strength: ∼15% of pulse-height reduction was observed close to the magnet center and the reduction started to decrease at 60 cm from the center and became negligible at 100 cm. CRT of the PET insert was also found to degrade inside the MRI scanner, which we postulate is due to the reduced pulse height and therefore may be compensated by increasing the signal gain. Fig. 13(c) shows the CRT measured from waveform
3.3.2. Effects on MR due to PET An MR image of the resolution phantom filled with 18 F solution acquired during simultaneous PET/MR operation is shown in Fig. 15(a). No image artifacts are observed. The potential effects on MR imaging due to the PET insert were studied more carefully by using an uniform cylinder (OD = 2.5 cm, length = 12 cm) filled with water. Two different MR pulse sequences, including gradient echo and spin 7
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Fig. 13. (a) Signal waveforms of the PET insert placed inside and outside the MRI scanner. (b) The relative pulse-height measured at different location along the MR bore. (c) Coincidence resolving time as a function of bias voltage over the breakdown.
Fig. 14. Resolution phantom images acquired inside the MRI scanner and line profiles: (a) and (c) when only the PET insert is running, and (b) and (d) while both the PET insert and the MR scanner are running.
in the middle column. The PET image was reconstructed by using 1.2 M coincidence events that are collected over 15 min. Qualitatively, there is a good match between the PET and MR images.
echo, were employed for comparison. Figs. 15(b) and 15(c) show the images obtained with the spin echo sequence. Figs. 15(d) and 15(e) show axial total-intensity profiles (by summing pixel intensities in an image slice) obtained before and after placing the PET insert inside the MR scanner. The signal cables of the PET insert are on the left-hand side of the profiles shown, and the power cables are on the right-hand side. When the PET insert was placed inside the MR scanner, while the profiles remain approximately the same on the signal-cable side, their amplitude decreases on the power-cable side. This reduction is larger in the case of the gradient echo sequence. The profiles of the signal-tonoise ratio (S/N) and uniformity calculated as described in [32] also have similar degradation on the power-cable side (results not shown). We suspect the reduction is caused by the biasing cables and connectors on the PET insert, which appear to affect gradient field switching.
4. Discussion When the PET insert is present, we do observe some effects to MRI, especially when using gradient echo pulse sequences. We are investigating which components of the PET insert that are responsible. In particular, we have placed the PET insert inside the MR scanner with or without the biasing cables. By comparing the resulting MR image profiles, we have strong evidence that suggests the effects are due to the cables and connectors. We postulate that the gradient field switching induces eddy currents in the cables and connectors to affect the magnetic field uniformity. Moreover, unlike the signal cables, the biasing cable is close to the detector center; this proximity might also cause asymmetric effects on MR imaging. Currently, we use separate cables for signal transmission and biasing so that the prototypical PET insert is easier to construct and trouble shoot. We may, however, be able to avoid biasing cables and connectors in our next design by making all cable connections on the same side of the detector module. Also, we may employ flat flexible cables for the signal and biasing cables so the PET insert is more practical to use. Currently, our PET insert is completely unshielded in order to evaluate the intrinsic PET/MR mutual interference of the proposed technology. The aforementioned
3.4. Initial simultaneous PET/MR mouse imaging A PET/MR image of a mouse brain is acquired by running the PET insert and MR scanner simultaneously. An initial activity of 180 μCi 18 F-fluorodeoxyglucose (FDG) was administrated to a healthy 30 gram male mouse through intraperitoneal (IP) injection, and data acquisition started 80 min post FDG injection. A spin echo pulse sequence with TR/TE = 4000/25 ms was used for MR imaging. Fig. 16 shows the resulting images of the mouse brain (MR images shown in gray-scale and PET images in color), and the co-registered PET/MR images are shown 8
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Fig. 15. (a) A transverse view of the MR image of the resolution phantom. (b) A transverse and (c) sagittal view of the MR image of the uniform phantom filled with water. Signal profiles in MR images along the axial slice are shown for (d) gradient echo sequence and (e) spin echo sequence.
Fig. 16. PET/MR combined images of a mouse brain. The left-hand side shows the trans-axial and sagittal views of the image, and the right-hand side shows the coronal views at the two different slices. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
effects may be addressed by applying adequate but minimal detector shielding utilizing insights gained in this study. Due to the use of 3.2 mm detector pixels, our PET insert currently has limited resolution for small animal imaging. To improve the spatial resolution, we can decrease the detector pixel size to 2 mm or smaller, to achieve an image resolution of about 1 mm or better. Also, in this work we have focused our attention on the instrumentation aspect of the insert. Therefore, we have only developed a most basic MLEM algorithm that is based on Siddon raytracing and incorporates prereconstruction corrections of detector-efficiency non-uniformity and random events. Filtered back projection algorithm will be implemented in image reconstruction as a basis for measuring the image spatial resolution in accordance to NEMA standards. It is well known that the image reconstruction process can critically impact both the resolution and statistical properties of the resulting images. In future work, we will employ more advanced image reconstruction algorithms that incorporate accurate modeling of the imaging physics. While TOF is not important for small animal imaging, we seek to develop PET insert for clinical PET/MR imaging; and, therefore, in this work we are also interested in examining the TOF capability with the proposed technology. In this study, by use of MVT sampling DAQ, we obtain a CRT of about 500 ps. However, Fig. 13(c) shows that the PET insert can achieve ∼270 ps when using the DRS4 waveform sampler. Previously, Zeng et al. [33] reported a ∼300 ps CRT for SiPM
based detector when using the MVT DAQ. Therefore, we believe that it is possible to improve the CRT of the PET insert toward 300 ps when optimizing the MVT DAQ board for our detector modules. To further improve the CRT, we can modify the design to reduce the baseline noise level of the MVT board and increase the preamplifier (preamp) gain at the signal input. Currently, the preamp gain is set to 6 by the manufacturer, not optimized for our particular detector modules with MPPCs with strip-line readout. In this study, we calculate the CRT based on the time stamps obtained at each of the four MVT thresholds (50, 150, 300 and 600 mV) and observe that the best result is obtained with 50 mV. However, better results might be possible by using some extrapolation methods that use multiple time stamps. The current thresholds setting is only empirical. Systematic investigation seeking to optimize the MVT thresholds is underway. 5. Summary We have prototyped a PET insert for simultaneous PET/MRI. The PET insert consists of 14 detector modules arranged in a cylindrical shape to fit within a 9.4 T small animal MRI scanner. The prototypical PET insert adopts a strip-line signal multiplexing method and multi-voltage threshold waveform sampling previously proposed and demonstrated by us. These methods allow us to significantly reduce the readout channel number and electronics complexity while still 9
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maintaining high detection performance. Performance of the detector modules and the PET insert, and the mutual effects between the PET insert and MR scanner are experimentally evaluated. Initial imaging studies of resolution phantom and a healthy mouse are conducted. Our initial results are encouraging, but there are several issues that require further detailed investigations in future studies, which include minimizing the mutual interference between the PET and MRI systems with an effective shielding.
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Declaration of competing interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. CRediT authorship contribution statement H. Kim: Conceptualization, Software, Formal analysis, Investigation, Writing - original draft. Y. Hua: Resources, Investigation. H.-T. Chen: Software. H.-M. Tsai: Software, Visualization. C.-T. Chen: Resources, Funding acquisition. G. Karczmar: Resources, Funding acquisition. X. Fan: Resources, Investigation. D. Xi: Resources, Investigation. Q. Xie: Resources, Funding acquisition. C.-Y. Chou: Resources, Funding acquisition. C.-M. Kao: Conceptualization, Writing - review & editing, Resources, Funding acquisition. Acknowledgments This work was supported in part by the University of Chicago Medicine Comprehensive Cancer Center pilot award, United States; Ministry of Science and Technology of Taiwan #MOST 107-2321-B055-004, ; the National Natural Science Foundation of China (NSFC) Grant #61425001, #61671215, #61604059, the National Key Scientific Instrument and Equipment Development Project of China #2013YQ030923, and the National Key Research and Development Program #2016YFF0101501. The Hamamatsu Corporation generously provided MPPC arrays used in this study. The authors thank Inna Gertsenshteyn for proof-reading the manuscript. References [1] T. Jones, Molecular imaging with PET – the future challenges, Br. J. Radiol. 75 (2002) S6–S15. [2] S.M. Ametamey, M. Honer, P.A. Schubiger, Molecular imaging with PET, Chem. Rev. 108 (2008) 1501–1516. [3] Y. Cossuin, A. Hocq, P. Gillis, Q. Vuong, Physics of magnetic resonance imaging: from spin to pixel, J. Phys. D: Appl. Phys. 43 (2010) 213001. [4] Y. Shao, et al., Simultaneous PET and MR imaging, Phys. Med. Biol. 42 (1997) 1965–1970. [5] R. Raylman, et al., Simultaneous MRI and PET imaging of a rat brain, Phys. Med. Biol. 51 (2006) 6371–6379. [6] S. Yamamoto, Design and performance from an integrated PET/MRI system for small animals, Ann. Nucl. Med. 24 (2010) 89–98. [7] P. Buzhan, et al., Silicon photomultiplier and its possible applications, Nucl. Instrum. Methods Phys. Res. A 504 (2003) 48–52.
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