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GAGG phoswich detector aiming for SPECT imaging
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
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Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
A side-by-side LYSO/GAGG phoswich detector aiming for SPECT imaging Qingyang Wei a ,∗, Tianyu Ma b ,∗, Nianming Jiang c , Tianpeng Xu c , Zhenlei Lyu b , Yulin Hu a , Yaqiang Liu b a
Beijing Engineering Research Center of Industrial Spectrum Imaging, School of Automation and Electrical Engineering, University of Science and Technology Beijing, Beijing 100083, China b Department of Engineering Physics, Tsinghua University, Beijing 100084, China c Beijing Novel Medical Equipment Ltd., Beijing 102206, China
ARTICLE Keywords: SPECT Phoswich detector LYSO/GAGG SiPM High performances
INFO
ABSTRACT Single photon emission computed tomography (SPECT) is an effective metabolic and functional imaging technique in both clinical and pre-clinical applications. Gamma-ray detector is a key component of a SPECT system. High performance detectors are required for high-quality SPECT imaging. In this paper, a detector using a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array is developed. The phoswich crystals are discriminated using pulse waveform integration to peak ratio. A 99m Tc and a 57 Co source irradiation experiments were conducted. Results of the experiments demonstrate that the LYSO and GAGG can be clearly discriminated using the simple parameter. 12 × 12 of both LYSO and GAGG crystal arrays with a crystal size of 1.35 × 2.7 × 7 mm3 could be successfully distinguished. The average energy resolutions @140 keV of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% and 21.29%, respectively. The calculation and experiment verifies that the side-by-side LYSO/GAGG phoswich detector achieves relatively high resolution and high sensitivity which is suitable for high performance SPECT systems.
1. Introduction Single photon emission computed tomography (SPECT) is an effective metabolic and functional imaging technique, and is increasingly used as a quantitative imaging tool in recent years. SPECT and SPECT/CT have a wide variety of useful diagnostic applications such as bone scanning for metabolic bone diseases, myocardial perfusion for heart diseases, and neurotransmitter brain imaging for brain diseases, etc [1–3]. They are also widely used in pre-clinical for drug and disease researches [4,5]. Most of the commercial SPECT systems are built with cost-effective, large size monolithic Nal(TI) scintillators decoded by an array of large conventional PMTs using the well-known Anger logic principle [6]. This common design is inexpensive to manufacture, and robust as using few output electronica channels. However, there are some fundamental limitations such as relatively long image acquisition time due to the low sensitivity and count rate limited by signal pile up, and poor intrinsic spatial resolution due to the scintillation photons spread in the large size scintillator. This kind of detector design has an intrinsic spatial resolution limited around 3.5∼4 mm averaged over the useful field of view [7,8]. Nowadays, as new clinical and research applications and are growing, higher quality clinical SPECT images are anticipated. Substantial efforts to improve SPECT image quality have been undertaken by both
the industry and academic researchers. One of the focused research direction is developing high performance detectors for SPECT [9]. Several state-of-art SPECT systems utilize advanced semiconductor technology such as cadmium-zinc-Telluride (CZT) [10,11], in both laboratory prototype and commercially available systems in clinical. CZT is a solid state detector which can achieve very high energy resolution compared to traditional scintillator based detectors. It can also achieve high intrinsic spatial resolution due to the pixelated implement (pixel size is 2.46 × 2.46 × 5 mm3 used in D-SPECT [12]). However, CZT detectors are much more expensive than scintillation detectors. Micro-column scintillators grown on CCD/EMCCD/CMOS sensors is another promising approach being used in pre-clinical SPECT for small animal imaging which can achieve extremely high spatial resolution [13,14]. However, the detector block is small and thin, which leads to low sensitivity. Relatively high resolution and high sensitivity SPECT detector could also be developed by assembling pixelated scintillators with a typical crystal size of 2–3 mm [6,15]. The crystal can be made very small (submillimeter) to furtherly improve the spatial resolution. However, small pixel size has several disadvantages including increasing cost, hard to assemble, requiring small size SiPMs to decode the position, and the small packing fraction leads to low sensitivity. Overall, it is trade-offs between spatial resolution, sensitivity, and cost.
∗ Corresponding author. E-mail addresses: [email protected] (Q. Wei), [email protected] (T. Ma).
https://doi.org/10.1016/j.nima.2019.163242 Received 3 September 2019; Received in revised form 29 November 2019; Accepted 3 December 2019 Available online 6 December 2019 0168-9002/© 2019 Elsevier B.V. All rights reserved.
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Phoswich detector design is a promising technique for improving spatial resolution. It has been widely used in PET detectors to measure the depth of interaction (DOI) information for reducing parallax error, which uses multiple layer scintillators with different decay times and pulse shape discrimination technique [16–18]. Another approach of the phoswich design places scintillator pairs side-by-side in a single layer. The best-known application of this approach is the LabPET system developed by Université de Sherbrooke group, which uses adjacent LYSO/LGSO read out by a single avalanche photodiode (APD) to reduce electronic channels [19]; our group used side-by-side LYSO/YSO read out by four PMTs with quadrant-sharing technique to reduce decoding error [20]; Shao et al. used phoswhich design at the edge of block to solve edge resolving [21]; Han et al. did a simulation study of sideby-side phoswich providing high spatial resolution of < 0.4 mm [22], Jonathan et al. developed a side-by-side fast-LSO/slow-LSO detector readout by position-sensitive photomultiplier (PSPMT) to improve spatial sampling [23]. To the best of our knowledge, all detectors of the second phoswich approach are also designed for PET systems. In this paper, a side-by-side LYSO/GAGG phoswich detector is proposed and realized aiming for SPECT imaging. LYSO and GAGG are the new choice scintillators for SPECT detectors instead of NaI and CsI [24– 30], which have high density, high atomic number (Z-effective), relatively high light output and short decay time. Scintillators with short decay times can reduce signal dead time and pile-up events, thus they can be used for high count rate imaging. Meanwhile, LYSO and GAGG have decay time constants of ∼40 ns and ∼90 ns individually, which can be used for phoswhich detectors [31]. Details of the design and evaluation of a prototype detector for SPECT imaging are demonstrated in this paper.
2.2. Experiment & data processing The detector was covered by a light-tight box and worked at room temperature (∼29 ◦ C) with a SiPM-supplied voltage of 29 V. Two datasets were individually acquired with a 57 Co source (about 1.8 μCi) and a 99m Tc source (about 0.1 mCi) placed on the top of the detector. The samples of the waveforms were transmitted to a PC using the Gigabit Ethernet for further off-line studies. Fig. 3 shows the waveforms of the LYSO/GAGG block in afterglow mode, which are normalized by the peak value of the signals. As shown in the pulse waveforms, increasing the integration time will have better distinguish of LYSO and GAGG, however it might blur the flood map due to the pulse undershoot/overshoot introduced by the capacitors of the SiPM cells [35]. Thus, we used different integration points for energy and integration-to-peak ratio calculation. The energies of E, X and Y signals are calculated by integrating four points before the maximum deviation and 26 points after it. The position of each scintillation event is calculated by the Anger logic. The ratio (r) of waveform integration value (four points before the maximum deviation and 36 points after it) to peak value is used to distinguish the scintillators. Fig. 3(b) shows the distribution of ratio r of the LYSO/GAGG block. There are two peaks in the distribution curve, each one is fitted by a Gaussian function. The fitted results of ratio r of LYSO/GAGG are 14.42 ± 0.73 and 19.84 ± 0.99 respectively, leading to a clear separation of two crystals in the flood map shown in Fig. 3(b). The intersection point of two fitted Gaussian functions is chosen as a threshold value to distinguish LYSO and GAGG crystals which is 16.78 in this platform. The flood map of LYSO is generated using all the events with ratio r belows the threshold (signal decay fast) and the flood map of GAGG is generated using all the events up the threshold (signal decay slow). The distribution curve is higher than the sum curve of two fitted Gaussians around the threshold because the Compton scattering cross-talk events between LYSO and GAGG crystals exist.
2. Materials & methods 2.1. Detector design
3. Results
The proposed design scheme is illustrated in Fig. 1. The crystal block consists of 12 × 12 phoswich element pairs with an overall dimension of 33.7 × 33.7 mm2 . Each pair has a size of 2.75 × 2.75 × 7 mm3 , which is assembled with a LYSO (SIPAT, China, decay time constant < 42 ns) and a GAGG(Ce) (EPIC-Crystal, China, decay time constant 87 ns (90%)/255 ns (10%)) crystal with both sizes of 1.35 × 2.7 × 7 mm3 . There are enhanced spectral reflectors (ESR, 3M, USA) between each pair to control the propagation of scintillation photons. A scintillator block is assembled by coupling the crystals with optical glues (OE6551, Dow Corning, USA). The scintillator block is then coupled to a selfassembled 8 × 8 SiPM array with optical grease (BC-630, Saint-Gobain Crystals, USA) [32]. The 8 × 8 SiPM array has an overall dimension of 33.7 × 33.7 mm2 with 4.2 mm pitch to match the scintillator array. Two SiPM arrays on a same printed circuit board are shown in Fig. 1(c) (Only one array is utilized in the following experiment). The SiPM unit (MicroFB-30035-SMT, SensL, Ireland) with ball grid array (BGA) has an active area of 3.16 × 3.16 mm2 , including 4774 microcells with the size of 35 × 35 μm2 . There is one-millimeter gap between adjacent SiPM pixels in the array. This sparse placement design can reduce the 40% cost of SiPMs compared to the closely arrangement design [33]. An ESR reflector is placed in SiPM gaps to improve scintillation photon collection. Eight-by-eight output channels of the SiPM array are multiplexed by in-chip resistor networks in a front-end ASIC called EXYT [34] on the other side of the printed circuit board. The outputs of the ASIC have three analog signals (E, X, and Y ) for energy and position decoding. Pulses of them are digitized by a 12-bit ADC chip (AD9634, Analog Devices, USA) with a sample rate of 65 MHz, and transmitted to a PC by a Gigabit Ethernet as shown in Fig. 2.
3.1. Position decoding The flood maps of 57 Co (5.4 million events) and 99m Tc (22.6 million events) have high overlap between phoswich pairs as shown in Fig. 4(a) & (d) respectively. While all 12 × 12 crystals for can be distinguished both in LYSO flood map of 99m Tc (r < 16.78) and GAGG flood map of 99m Tc (r ≥ 16.78) as shown in Fig. 4(b) & (c) respectively. LYSO flood map and GAGG flood map of 57 Co are shown in Fig. 4(d) & (e), which are slightly poorer than flood maps of 99m Tc due to the lower gamma photon energy. There are slight overlaps at the edge as the SiPM array has low decoding capability at the edge. There are several low-count crystals in GAGG flood map and high-count crystals in LYSO flood map due to GAGG’s events mis-assigned to LYSO as these crystals have shorter decay time. The quality of the flood maps demonstrate that the intrinsic spatial resolution of the detector approximately equals to the crystal size, which is 2.7 × 1.35 mm2 . 3.2. Photon peaks and energy resolutions Crystal position maps of LYSO and GAGG arrays are generated using a SVD and mean-shift based peak tracking algorithm [36]. The segmentation results of 99m Tc flood maps are shown in Fig. 5(a) & (d) for LYSO and GAGG respectively. The energy spectrum of each crystal is generated by histogramming the energies using the look-up tables. The photon peak and energy resolution of each spectrum is fitted by the Gaussian function. The average peaks of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals of 140 keV are 1719 ± 192 and 1967 ± 191 respectively. The peak of GAGG is 14% higher than the peak of LYSO. The distributions of peak positions of LYSO crystals and GAGG crystals are shown in Fig. 5(b) & (e). The average energy resolutions of 12 × 2
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 1. Details of the detector design. (a) Schematic diagram of the proposed side-by-side phoswich detector configuration consists of the crystal block and SiPM array; (b) Photograph of a side-by-side LYSO/GAGG phoswich array; (c) Photograph of two 8 × 8 SiPM array.
Fig. 2. (a) A phoswich array coupled to a SiPM array; (b) The data acquisition board.
Fig. 3. (a) The normalization waveforms of the LYSO/GAGG block shown in afterglow mode; (b) The distribution of r and the Gaussian fitted results.
12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% ± 4.64% and
The segmentation results of
57 Co
flood maps are shown in Fig. 6(a)
21.29% ± 3.67% respectively. The distributions of energy resolutions
& (d) for LYSO and GAGG respectively. The average peaks of 12 × 12
of LYSO crystals and GAGG crystals are shown in Fig. 5(c) & (f).
LYSO crystals and 12 × 12 GAGG crystals of 122 keV are 1469 ± 177 3
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 4. (a) LYSO/GAGG flood map of 99m Tc without pulse discrimination; (b) LYSO flood map of 99m Tc (r < 16.78); (c) GAGG flood map of flood map of 57 Co without pulse discrimination; (e) LYSO flood map of 57 Co (r < 16.78); (f) GAGG flood map of 57 Co (r ⩾ 16.78).
99m Tc
(r ⩾ 16.78); (d) LYSO/GAGG
Fig. 5. Results of the 99m Tc experiment. Segmentation results of the flood maps (a&d), energy peak positions (b&e) and energy resolutions of 140 keV (d&f) for all crystals. Row one and row two are the results of LYSO and GAGG respectively.
and 1649 ± 169 respectively. The peak of GAGG is 12% higher than the peak of LYSO. The distributions of peak positions of LYSO crystals and GAGG crystals are shown in Fig. 6(b) & (e). The average energy resolutions of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are
30.57% ± 6.18% and 23.10% ± 1.41% respectively. The distributions of energy resolutions of LYSO crystals and GAGG crystals are shown in Fig. 6(c) & (f). 4
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 6. Results of the 57 Co experiment. Segmentation results of the flood maps (a&d), energy peak positions (b&e) and energy resolutions @ 122 keV (d&f) for all crystals. Row one and row two are the results of LYSO and GAGG respectively. Table 1 Stopping power comparison of LYSO, GAGG, and NaI(Tl).
4. Discussions
Scintillator
In this paper, we proposed a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array. The pulse waveforms of LYSO and GAGG can be clearly discriminated using a simple parameter, i.e. the pulse waveform integration to peak ratio which can be easily integrated into an FPGA program. 12 × 12 of both LYSO and GAGG crystal array with a size of 1.35 × 2.7 × 7 mm3 could be successfully distinguished. The average energy resolutions @140 keV of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% and 21.29% respectively, and 30.57% and 23.10% @122 keV respectively. Compared to [28], in which the LYSO and GAGG detector arrays are assembled and investigated individually, our proposed detector has a lower intrinsic spatial resolution. The reason is that we used larger crystal pixel size so that they are decodable with larger SiPM units. Such design limits the intrinsic spatial resolution, but has the advantage of reducing the cost. Besides, there is a trade-off between the detection efficiency and the spatial resolution because thicker crystal will have poorer decoding result due to low light outcome. To achieving high resolution, [28] use 1 mm thick GAGG and 2 mm thick LYSO, which will likely lead to a low sensitivity. The stopping power comparison for 99m Tc (140 keV) and 131 I (364 keV) is listed in Table 1, which are calculated using the XCOM database [37]. Our proposed detector can achieve a sensitivity compatible to a 25.4 mm NaI(Tl) detector, which is suitable for imaging different energy SPECT tracers including 99m Tc and 131 I. The energy resolution of the LYSO/GAGG block presented here is similar to [28], which is poorer than the commonly NaI(Tl) based detectors (typically ≤ 10% @140 keV). The reason is LYSO generates less optical photons (26 000 photons/MeV) than NaI(Tl) does (38 000 photons/MeV); GAGG generates more optical photons (50 000 photons/MeV) while the emission wavelength peak (520 nm) does not match the photon detection efficiency (PDE) peak of the SiPM (420 nm) which is the same as LYSO. Good energy resolution could reduce the scatter events via a narrow energy window. The energy resolution presented here needs to be further improved for clinical use when patient scattering is a concern. However, for small animal SPECT, it is sufficient since the fraction of detected scattered photons is less important [38]. The energy resolution could be further improved by
Density (g/cm3 )
Stopping power (%) 140 keV
364 keV
7 mm LYSO 2 mm LYSOa
7.3
96.7 82.7
45.4 21.6
7 mm GAGG 1 mm GAGGa
6.4
99.8 38.5
57.3 8.3
9.5 mm NaI(Tl)b 25.4 mm NaI(Tl)c
3.6
89.8 99.8
26.3 67.0
a
Used in Ref. [28]. Most commonly used in commercial SPECT systems. c GE StarBright™ SPECT. b
using closely arrangement SiPM array with small gaps, and cooling SiPMs to a low temperature for reducing dark count rate [39]. 176 Lu background radiation of LYSO [40] might introduce noise events. Several studies indicate that the 176 Lu background in LYSO crystal blocks is not expected to significantly deteriorate the performance of SPECT systems [25,30,41]. Moreover, calibration methods could be employed to reduce the background radiation influence [30]. As there is no intrinsic background radiation in GAGG crystals, the LYSO/GAGG phoswich design in this paper reduces half of the 176 Lu background radiation compared to LYSO blocks. 5. Conclusions In this paper, we proposed a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array. The pulse waveforms of LYSO and GAGG can be clearly discriminated. The side-by-side phoswich design improves the spatial resolution in one direction. The proposed design has relatively high resolution and high sensitivity which is suitable for SPECT systems to achieve high imaging performance. 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. 5
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
CRediT authorship contribution statement
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Qingyang Wei: Conceptualization, Methodology, Writing - original draft. Tianyu Ma: Writing - review & editing, Funding acquisition. Nianming Jiang: Resources. Tianpeng Xu: Resources. Zhenlei Lyu: Validation. Yulin Hu: Writing - review & editing. Yaqiang Liu: Funding acquisition. Acknowledgments This work was supported in part by National Natural Science Foundation of China (No. 11975044, No. 11605008 & No. 81727807), Fundamental Research Funds for the Central Universities, China (No. FRFTP-18-035A2 & FRF-TP-19-019A3), Science & Technology on Reliability & Environmental Engineering Laboratory, China (No. 6142004180205). References [1] D.L. Bailey, K.P. Willowson, An evidence-based review of quantitative SPECT imaging and potential clinical applications, J. Nucl. Med. 54 (1) (2013) 83–89. [2] P. Sharma, S. Sharma, S. Ballal, et al., SPECT-CT in routine clinical practice: increase in patient radiation dose compared with SPECT alone, Nucl. Med. Commun. 33 (9) (2012) 926–932. [3] J. Wu, C. Liu, Recent advances in cardiac SPECT instrumentation and imaging methods, Phys. Med. Biol. 64 (6) (2019) 06TR01. [4] S. Neyt, M. Vliegen, B. Verreet, et al., Synthesis, in vitro and in vivo smallanimal SPECT evaluation of novel technetium labeled bile acid analogues to study (altered) hepatic transporter function, Nucl. Med. Biol. 43 (10) (2016) 642–649. [5] B.L. Franc, P.D. Acton, C. Mari, et al., Small-animal SPECT and SPECT/CT: important tools for preclinical investigation, J. Nucl. Med. 49 (10) (2008) 1651–1663. [6] T.E. Peterson, L.R. Furenlid, SPECT detectors: the Anger Camera and beyond, Phys. Med. Biol. 56 (17) (2011) R145. [7] M.T. Madsen, Recent advances in SPECT imaging, J. Nucl. Med. 48 (4) (2007) 661–673. [8] D.A.I. Tiantian, L.I.U. Hui, C.U.I. Junjian, et al., A high-resolution small animal SPECT system developed at Tsinghua, Nucl. Sci. Technol. 22 (6) (2013) 348-344. [9] M. Ljungberg, P.H. Pretorius, Nuclear medicine: Physics and instrumentation special feature review article: SPECT/CT: An update on technological developments and clinical applications, Br. J. Radiol. 91 (1081) (2018). [10] D. Agostini, P.Y. Marie, S. Ben-Haim, et al., Performance of cardiac cadmiumzinc-telluride gamma camera imaging in coronary artery disease: a review from the cardiovascular committee of the European Association of Nuclear Medicine (EANM), Eur. J. Nucl. Med. Mol. Imaging 43 (13) (2016) 2423–2432. [11] F. Nudi, A.E. Iskandrian, O. Schillaci, et al., Diagnostic accuracy of myocardial perfusion imaging with CZT technology: systemic review and meta-analysis of comparison with invasive coronary angiography, JACC: Cardiovasc. Imaging 10 (7) (2017) 787–794. [12] K. Erlandsson, K. Kacperski, D. Van Gramberg, et al., Performance evaluation of D-SPECT: a novel SPECT system for nuclear cardiology, Phys. Med. Biol. 54 (9) (2009) 2635. [13] B.W. Miller, H.H. Barrett, L.R. Furenlid, et al., Recent advances in BazookaSPECT: Real-time data processing and the development of a gamma-ray microscope, Nucl. Instrum. Methods Phys. Res. A 591 (1) (2008) 272–275. [14] V. Nagarkar, V. Gaysinskiy, I. Shestakova, et al., Microcolumnar CsI(Tl) films for small animal SPECT, in: IEEE Nuclear Science Symposium Conference, Vol. 5, 2004, pp. 3334–3337. [15] M. Rozler, H. Liang, H. Sabet, et al., Development of a cost-effective modular pixelated NaI(Tl) detector for clinical SPECT applications, IEEE Trans. Nucl. Sci. 59 (5) (2012) 1831–1840. [16] M. Schm, L. Eriksson, M.E. Casey, et al., Performance results of a new DOI detector block for a high resolution PET-LSO research tomograph HRRT, IEEE Trans. Nucl. Sci. 45 (6) (1998) 3000–3006. [17] J. Seidel, J.J. Vaquero, S. Siegel, et al., Depth identification accuracy of a three layer phoswich PET detector module, IEEE Trans. Nucl. Sci. 46 (3) (1999) 485–490.
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Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
A side-by-side LYSO/GAGG phoswich detector aiming for SPECT imaging Qingyang Wei a ,∗, Tianyu Ma b ,∗, Nianming Jiang c , Tianpeng Xu c , Zhenlei Lyu b , Yulin Hu a , Yaqiang Liu b a
Beijing Engineering Research Center of Industrial Spectrum Imaging, School of Automation and Electrical Engineering, University of Science and Technology Beijing, Beijing 100083, China b Department of Engineering Physics, Tsinghua University, Beijing 100084, China c Beijing Novel Medical Equipment Ltd., Beijing 102206, China
ARTICLE Keywords: SPECT Phoswich detector LYSO/GAGG SiPM High performances
INFO
ABSTRACT Single photon emission computed tomography (SPECT) is an effective metabolic and functional imaging technique in both clinical and pre-clinical applications. Gamma-ray detector is a key component of a SPECT system. High performance detectors are required for high-quality SPECT imaging. In this paper, a detector using a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array is developed. The phoswich crystals are discriminated using pulse waveform integration to peak ratio. A 99m Tc and a 57 Co source irradiation experiments were conducted. Results of the experiments demonstrate that the LYSO and GAGG can be clearly discriminated using the simple parameter. 12 × 12 of both LYSO and GAGG crystal arrays with a crystal size of 1.35 × 2.7 × 7 mm3 could be successfully distinguished. The average energy resolutions @140 keV of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% and 21.29%, respectively. The calculation and experiment verifies that the side-by-side LYSO/GAGG phoswich detector achieves relatively high resolution and high sensitivity which is suitable for high performance SPECT systems.
1. Introduction Single photon emission computed tomography (SPECT) is an effective metabolic and functional imaging technique, and is increasingly used as a quantitative imaging tool in recent years. SPECT and SPECT/CT have a wide variety of useful diagnostic applications such as bone scanning for metabolic bone diseases, myocardial perfusion for heart diseases, and neurotransmitter brain imaging for brain diseases, etc [1–3]. They are also widely used in pre-clinical for drug and disease researches [4,5]. Most of the commercial SPECT systems are built with cost-effective, large size monolithic Nal(TI) scintillators decoded by an array of large conventional PMTs using the well-known Anger logic principle [6]. This common design is inexpensive to manufacture, and robust as using few output electronica channels. However, there are some fundamental limitations such as relatively long image acquisition time due to the low sensitivity and count rate limited by signal pile up, and poor intrinsic spatial resolution due to the scintillation photons spread in the large size scintillator. This kind of detector design has an intrinsic spatial resolution limited around 3.5∼4 mm averaged over the useful field of view [7,8]. Nowadays, as new clinical and research applications and are growing, higher quality clinical SPECT images are anticipated. Substantial efforts to improve SPECT image quality have been undertaken by both
the industry and academic researchers. One of the focused research direction is developing high performance detectors for SPECT [9]. Several state-of-art SPECT systems utilize advanced semiconductor technology such as cadmium-zinc-Telluride (CZT) [10,11], in both laboratory prototype and commercially available systems in clinical. CZT is a solid state detector which can achieve very high energy resolution compared to traditional scintillator based detectors. It can also achieve high intrinsic spatial resolution due to the pixelated implement (pixel size is 2.46 × 2.46 × 5 mm3 used in D-SPECT [12]). However, CZT detectors are much more expensive than scintillation detectors. Micro-column scintillators grown on CCD/EMCCD/CMOS sensors is another promising approach being used in pre-clinical SPECT for small animal imaging which can achieve extremely high spatial resolution [13,14]. However, the detector block is small and thin, which leads to low sensitivity. Relatively high resolution and high sensitivity SPECT detector could also be developed by assembling pixelated scintillators with a typical crystal size of 2–3 mm [6,15]. The crystal can be made very small (submillimeter) to furtherly improve the spatial resolution. However, small pixel size has several disadvantages including increasing cost, hard to assemble, requiring small size SiPMs to decode the position, and the small packing fraction leads to low sensitivity. Overall, it is trade-offs between spatial resolution, sensitivity, and cost.
∗ Corresponding author. E-mail addresses: [email protected] (Q. Wei), [email protected] (T. Ma).
https://doi.org/10.1016/j.nima.2019.163242 Received 3 September 2019; Received in revised form 29 November 2019; Accepted 3 December 2019 Available online 6 December 2019 0168-9002/© 2019 Elsevier B.V. All rights reserved.
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Phoswich detector design is a promising technique for improving spatial resolution. It has been widely used in PET detectors to measure the depth of interaction (DOI) information for reducing parallax error, which uses multiple layer scintillators with different decay times and pulse shape discrimination technique [16–18]. Another approach of the phoswich design places scintillator pairs side-by-side in a single layer. The best-known application of this approach is the LabPET system developed by Université de Sherbrooke group, which uses adjacent LYSO/LGSO read out by a single avalanche photodiode (APD) to reduce electronic channels [19]; our group used side-by-side LYSO/YSO read out by four PMTs with quadrant-sharing technique to reduce decoding error [20]; Shao et al. used phoswhich design at the edge of block to solve edge resolving [21]; Han et al. did a simulation study of sideby-side phoswich providing high spatial resolution of < 0.4 mm [22], Jonathan et al. developed a side-by-side fast-LSO/slow-LSO detector readout by position-sensitive photomultiplier (PSPMT) to improve spatial sampling [23]. To the best of our knowledge, all detectors of the second phoswich approach are also designed for PET systems. In this paper, a side-by-side LYSO/GAGG phoswich detector is proposed and realized aiming for SPECT imaging. LYSO and GAGG are the new choice scintillators for SPECT detectors instead of NaI and CsI [24– 30], which have high density, high atomic number (Z-effective), relatively high light output and short decay time. Scintillators with short decay times can reduce signal dead time and pile-up events, thus they can be used for high count rate imaging. Meanwhile, LYSO and GAGG have decay time constants of ∼40 ns and ∼90 ns individually, which can be used for phoswhich detectors [31]. Details of the design and evaluation of a prototype detector for SPECT imaging are demonstrated in this paper.
2.2. Experiment & data processing The detector was covered by a light-tight box and worked at room temperature (∼29 ◦ C) with a SiPM-supplied voltage of 29 V. Two datasets were individually acquired with a 57 Co source (about 1.8 μCi) and a 99m Tc source (about 0.1 mCi) placed on the top of the detector. The samples of the waveforms were transmitted to a PC using the Gigabit Ethernet for further off-line studies. Fig. 3 shows the waveforms of the LYSO/GAGG block in afterglow mode, which are normalized by the peak value of the signals. As shown in the pulse waveforms, increasing the integration time will have better distinguish of LYSO and GAGG, however it might blur the flood map due to the pulse undershoot/overshoot introduced by the capacitors of the SiPM cells [35]. Thus, we used different integration points for energy and integration-to-peak ratio calculation. The energies of E, X and Y signals are calculated by integrating four points before the maximum deviation and 26 points after it. The position of each scintillation event is calculated by the Anger logic. The ratio (r) of waveform integration value (four points before the maximum deviation and 36 points after it) to peak value is used to distinguish the scintillators. Fig. 3(b) shows the distribution of ratio r of the LYSO/GAGG block. There are two peaks in the distribution curve, each one is fitted by a Gaussian function. The fitted results of ratio r of LYSO/GAGG are 14.42 ± 0.73 and 19.84 ± 0.99 respectively, leading to a clear separation of two crystals in the flood map shown in Fig. 3(b). The intersection point of two fitted Gaussian functions is chosen as a threshold value to distinguish LYSO and GAGG crystals which is 16.78 in this platform. The flood map of LYSO is generated using all the events with ratio r belows the threshold (signal decay fast) and the flood map of GAGG is generated using all the events up the threshold (signal decay slow). The distribution curve is higher than the sum curve of two fitted Gaussians around the threshold because the Compton scattering cross-talk events between LYSO and GAGG crystals exist.
2. Materials & methods 2.1. Detector design
3. Results
The proposed design scheme is illustrated in Fig. 1. The crystal block consists of 12 × 12 phoswich element pairs with an overall dimension of 33.7 × 33.7 mm2 . Each pair has a size of 2.75 × 2.75 × 7 mm3 , which is assembled with a LYSO (SIPAT, China, decay time constant < 42 ns) and a GAGG(Ce) (EPIC-Crystal, China, decay time constant 87 ns (90%)/255 ns (10%)) crystal with both sizes of 1.35 × 2.7 × 7 mm3 . There are enhanced spectral reflectors (ESR, 3M, USA) between each pair to control the propagation of scintillation photons. A scintillator block is assembled by coupling the crystals with optical glues (OE6551, Dow Corning, USA). The scintillator block is then coupled to a selfassembled 8 × 8 SiPM array with optical grease (BC-630, Saint-Gobain Crystals, USA) [32]. The 8 × 8 SiPM array has an overall dimension of 33.7 × 33.7 mm2 with 4.2 mm pitch to match the scintillator array. Two SiPM arrays on a same printed circuit board are shown in Fig. 1(c) (Only one array is utilized in the following experiment). The SiPM unit (MicroFB-30035-SMT, SensL, Ireland) with ball grid array (BGA) has an active area of 3.16 × 3.16 mm2 , including 4774 microcells with the size of 35 × 35 μm2 . There is one-millimeter gap between adjacent SiPM pixels in the array. This sparse placement design can reduce the 40% cost of SiPMs compared to the closely arrangement design [33]. An ESR reflector is placed in SiPM gaps to improve scintillation photon collection. Eight-by-eight output channels of the SiPM array are multiplexed by in-chip resistor networks in a front-end ASIC called EXYT [34] on the other side of the printed circuit board. The outputs of the ASIC have three analog signals (E, X, and Y ) for energy and position decoding. Pulses of them are digitized by a 12-bit ADC chip (AD9634, Analog Devices, USA) with a sample rate of 65 MHz, and transmitted to a PC by a Gigabit Ethernet as shown in Fig. 2.
3.1. Position decoding The flood maps of 57 Co (5.4 million events) and 99m Tc (22.6 million events) have high overlap between phoswich pairs as shown in Fig. 4(a) & (d) respectively. While all 12 × 12 crystals for can be distinguished both in LYSO flood map of 99m Tc (r < 16.78) and GAGG flood map of 99m Tc (r ≥ 16.78) as shown in Fig. 4(b) & (c) respectively. LYSO flood map and GAGG flood map of 57 Co are shown in Fig. 4(d) & (e), which are slightly poorer than flood maps of 99m Tc due to the lower gamma photon energy. There are slight overlaps at the edge as the SiPM array has low decoding capability at the edge. There are several low-count crystals in GAGG flood map and high-count crystals in LYSO flood map due to GAGG’s events mis-assigned to LYSO as these crystals have shorter decay time. The quality of the flood maps demonstrate that the intrinsic spatial resolution of the detector approximately equals to the crystal size, which is 2.7 × 1.35 mm2 . 3.2. Photon peaks and energy resolutions Crystal position maps of LYSO and GAGG arrays are generated using a SVD and mean-shift based peak tracking algorithm [36]. The segmentation results of 99m Tc flood maps are shown in Fig. 5(a) & (d) for LYSO and GAGG respectively. The energy spectrum of each crystal is generated by histogramming the energies using the look-up tables. The photon peak and energy resolution of each spectrum is fitted by the Gaussian function. The average peaks of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals of 140 keV are 1719 ± 192 and 1967 ± 191 respectively. The peak of GAGG is 14% higher than the peak of LYSO. The distributions of peak positions of LYSO crystals and GAGG crystals are shown in Fig. 5(b) & (e). The average energy resolutions of 12 × 2
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 1. Details of the detector design. (a) Schematic diagram of the proposed side-by-side phoswich detector configuration consists of the crystal block and SiPM array; (b) Photograph of a side-by-side LYSO/GAGG phoswich array; (c) Photograph of two 8 × 8 SiPM array.
Fig. 2. (a) A phoswich array coupled to a SiPM array; (b) The data acquisition board.
Fig. 3. (a) The normalization waveforms of the LYSO/GAGG block shown in afterglow mode; (b) The distribution of r and the Gaussian fitted results.
12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% ± 4.64% and
The segmentation results of
57 Co
flood maps are shown in Fig. 6(a)
21.29% ± 3.67% respectively. The distributions of energy resolutions
& (d) for LYSO and GAGG respectively. The average peaks of 12 × 12
of LYSO crystals and GAGG crystals are shown in Fig. 5(c) & (f).
LYSO crystals and 12 × 12 GAGG crystals of 122 keV are 1469 ± 177 3
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 4. (a) LYSO/GAGG flood map of 99m Tc without pulse discrimination; (b) LYSO flood map of 99m Tc (r < 16.78); (c) GAGG flood map of flood map of 57 Co without pulse discrimination; (e) LYSO flood map of 57 Co (r < 16.78); (f) GAGG flood map of 57 Co (r ⩾ 16.78).
99m Tc
(r ⩾ 16.78); (d) LYSO/GAGG
Fig. 5. Results of the 99m Tc experiment. Segmentation results of the flood maps (a&d), energy peak positions (b&e) and energy resolutions of 140 keV (d&f) for all crystals. Row one and row two are the results of LYSO and GAGG respectively.
and 1649 ± 169 respectively. The peak of GAGG is 12% higher than the peak of LYSO. The distributions of peak positions of LYSO crystals and GAGG crystals are shown in Fig. 6(b) & (e). The average energy resolutions of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are
30.57% ± 6.18% and 23.10% ± 1.41% respectively. The distributions of energy resolutions of LYSO crystals and GAGG crystals are shown in Fig. 6(c) & (f). 4
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
Fig. 6. Results of the 57 Co experiment. Segmentation results of the flood maps (a&d), energy peak positions (b&e) and energy resolutions @ 122 keV (d&f) for all crystals. Row one and row two are the results of LYSO and GAGG respectively. Table 1 Stopping power comparison of LYSO, GAGG, and NaI(Tl).
4. Discussions
Scintillator
In this paper, we proposed a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array. The pulse waveforms of LYSO and GAGG can be clearly discriminated using a simple parameter, i.e. the pulse waveform integration to peak ratio which can be easily integrated into an FPGA program. 12 × 12 of both LYSO and GAGG crystal array with a size of 1.35 × 2.7 × 7 mm3 could be successfully distinguished. The average energy resolutions @140 keV of 12 × 12 LYSO crystals and 12 × 12 GAGG crystals are 26.23% and 21.29% respectively, and 30.57% and 23.10% @122 keV respectively. Compared to [28], in which the LYSO and GAGG detector arrays are assembled and investigated individually, our proposed detector has a lower intrinsic spatial resolution. The reason is that we used larger crystal pixel size so that they are decodable with larger SiPM units. Such design limits the intrinsic spatial resolution, but has the advantage of reducing the cost. Besides, there is a trade-off between the detection efficiency and the spatial resolution because thicker crystal will have poorer decoding result due to low light outcome. To achieving high resolution, [28] use 1 mm thick GAGG and 2 mm thick LYSO, which will likely lead to a low sensitivity. The stopping power comparison for 99m Tc (140 keV) and 131 I (364 keV) is listed in Table 1, which are calculated using the XCOM database [37]. Our proposed detector can achieve a sensitivity compatible to a 25.4 mm NaI(Tl) detector, which is suitable for imaging different energy SPECT tracers including 99m Tc and 131 I. The energy resolution of the LYSO/GAGG block presented here is similar to [28], which is poorer than the commonly NaI(Tl) based detectors (typically ≤ 10% @140 keV). The reason is LYSO generates less optical photons (26 000 photons/MeV) than NaI(Tl) does (38 000 photons/MeV); GAGG generates more optical photons (50 000 photons/MeV) while the emission wavelength peak (520 nm) does not match the photon detection efficiency (PDE) peak of the SiPM (420 nm) which is the same as LYSO. Good energy resolution could reduce the scatter events via a narrow energy window. The energy resolution presented here needs to be further improved for clinical use when patient scattering is a concern. However, for small animal SPECT, it is sufficient since the fraction of detected scattered photons is less important [38]. The energy resolution could be further improved by
Density (g/cm3 )
Stopping power (%) 140 keV
364 keV
7 mm LYSO 2 mm LYSOa
7.3
96.7 82.7
45.4 21.6
7 mm GAGG 1 mm GAGGa
6.4
99.8 38.5
57.3 8.3
9.5 mm NaI(Tl)b 25.4 mm NaI(Tl)c
3.6
89.8 99.8
26.3 67.0
a
Used in Ref. [28]. Most commonly used in commercial SPECT systems. c GE StarBright™ SPECT. b
using closely arrangement SiPM array with small gaps, and cooling SiPMs to a low temperature for reducing dark count rate [39]. 176 Lu background radiation of LYSO [40] might introduce noise events. Several studies indicate that the 176 Lu background in LYSO crystal blocks is not expected to significantly deteriorate the performance of SPECT systems [25,30,41]. Moreover, calibration methods could be employed to reduce the background radiation influence [30]. As there is no intrinsic background radiation in GAGG crystals, the LYSO/GAGG phoswich design in this paper reduces half of the 176 Lu background radiation compared to LYSO blocks. 5. Conclusions In this paper, we proposed a side-by-side LYSO/GAGG phoswich array read out by a sparse 8 × 8 SiPM array. The pulse waveforms of LYSO and GAGG can be clearly discriminated. The side-by-side phoswich design improves the spatial resolution in one direction. The proposed design has relatively high resolution and high sensitivity which is suitable for SPECT systems to achieve high imaging performance. 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. 5
Q. Wei, T. Ma, N. Jiang et al.
Nuclear Inst. and Methods in Physics Research, A 953 (2020) 163242
CRediT authorship contribution statement
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Qingyang Wei: Conceptualization, Methodology, Writing - original draft. Tianyu Ma: Writing - review & editing, Funding acquisition. Nianming Jiang: Resources. Tianpeng Xu: Resources. Zhenlei Lyu: Validation. Yulin Hu: Writing - review & editing. Yaqiang Liu: Funding acquisition. Acknowledgments This work was supported in part by National Natural Science Foundation of China (No. 11975044, No. 11605008 & No. 81727807), Fundamental Research Funds for the Central Universities, China (No. FRFTP-18-035A2 & FRF-TP-19-019A3), Science & Technology on Reliability & Environmental Engineering Laboratory, China (No. 6142004180205). References [1] D.L. Bailey, K.P. Willowson, An evidence-based review of quantitative SPECT imaging and potential clinical applications, J. Nucl. Med. 54 (1) (2013) 83–89. [2] P. Sharma, S. Sharma, S. Ballal, et al., SPECT-CT in routine clinical practice: increase in patient radiation dose compared with SPECT alone, Nucl. Med. Commun. 33 (9) (2012) 926–932. [3] J. Wu, C. Liu, Recent advances in cardiac SPECT instrumentation and imaging methods, Phys. Med. Biol. 64 (6) (2019) 06TR01. [4] S. Neyt, M. Vliegen, B. Verreet, et al., Synthesis, in vitro and in vivo smallanimal SPECT evaluation of novel technetium labeled bile acid analogues to study (altered) hepatic transporter function, Nucl. Med. Biol. 43 (10) (2016) 642–649. [5] B.L. Franc, P.D. Acton, C. Mari, et al., Small-animal SPECT and SPECT/CT: important tools for preclinical investigation, J. Nucl. Med. 49 (10) (2008) 1651–1663. [6] T.E. Peterson, L.R. Furenlid, SPECT detectors: the Anger Camera and beyond, Phys. Med. Biol. 56 (17) (2011) R145. [7] M.T. Madsen, Recent advances in SPECT imaging, J. Nucl. Med. 48 (4) (2007) 661–673. [8] D.A.I. Tiantian, L.I.U. Hui, C.U.I. Junjian, et al., A high-resolution small animal SPECT system developed at Tsinghua, Nucl. Sci. Technol. 22 (6) (2013) 348-344. [9] M. Ljungberg, P.H. Pretorius, Nuclear medicine: Physics and instrumentation special feature review article: SPECT/CT: An update on technological developments and clinical applications, Br. J. Radiol. 91 (1081) (2018). [10] D. Agostini, P.Y. Marie, S. Ben-Haim, et al., Performance of cardiac cadmiumzinc-telluride gamma camera imaging in coronary artery disease: a review from the cardiovascular committee of the European Association of Nuclear Medicine (EANM), Eur. J. Nucl. Med. Mol. Imaging 43 (13) (2016) 2423–2432. [11] F. Nudi, A.E. Iskandrian, O. Schillaci, et al., Diagnostic accuracy of myocardial perfusion imaging with CZT technology: systemic review and meta-analysis of comparison with invasive coronary angiography, JACC: Cardiovasc. Imaging 10 (7) (2017) 787–794. [12] K. Erlandsson, K. Kacperski, D. Van Gramberg, et al., Performance evaluation of D-SPECT: a novel SPECT system for nuclear cardiology, Phys. Med. Biol. 54 (9) (2009) 2635. [13] B.W. Miller, H.H. Barrett, L.R. Furenlid, et al., Recent advances in BazookaSPECT: Real-time data processing and the development of a gamma-ray microscope, Nucl. Instrum. Methods Phys. Res. A 591 (1) (2008) 272–275. [14] V. Nagarkar, V. Gaysinskiy, I. Shestakova, et al., Microcolumnar CsI(Tl) films for small animal SPECT, in: IEEE Nuclear Science Symposium Conference, Vol. 5, 2004, pp. 3334–3337. [15] M. Rozler, H. Liang, H. Sabet, et al., Development of a cost-effective modular pixelated NaI(Tl) detector for clinical SPECT applications, IEEE Trans. Nucl. Sci. 59 (5) (2012) 1831–1840. [16] M. Schm, L. Eriksson, M.E. Casey, et al., Performance results of a new DOI detector block for a high resolution PET-LSO research tomograph HRRT, IEEE Trans. Nucl. Sci. 45 (6) (1998) 3000–3006. [17] J. Seidel, J.J. Vaquero, S. Siegel, et al., Depth identification accuracy of a three layer phoswich PET detector module, IEEE Trans. Nucl. Sci. 46 (3) (1999) 485–490.
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