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Development of a circular shape Si-PM-based detector ring for breast-dedicated PET system
Nuclear Inst. and Methods in Physics Research, A 880 (2018) 118–124
Contents lists available at ScienceDirect
Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
Development of a circular shape Si-PM-based detector ring for breast-dedicated PET system Kouhei Nakanishi a, *, Seiichi Yamamoto a, *, Hiroshi Watabe b , Shinji Abe c , Naotoshi Fujita c , Katsuhiko Kato a a b c
Department of Medical Technology, Nagoya University Graduate School of Medicine, Nagoya, Japan Tohoku University CYRIC, Sendai, Japan Department of Radiological Technology, Nagoya University Hospital, Nagoya, Japan
a r t i c l e
i n f o
Keywords: PEM PET Breast Si-PM LGSO Brain
a b s t r a c t In clinical situations, various breast-dedicated positron emission tomography (PET) systems have been used. However, clinical breast-dedicated PET systems have polygonal detector ring. Polygonal detector ring sometimes causes image artifact, so complicated reconstruction algorithm is needed to reduce artifact. Consequently, we developed a circular detector ring for breast-dedicated PET to obtain images without artifact using a simple reconstruction algorithm. We used Lu1.9 Gd0.1 SiO5 (LGSO) scintillator block which was made of 1.5 x 1.9 x 15 mm pixels that were arranged in an 8 x 24 matrix. As photodetectors, we used silicon photomultiplier (Si-PM) arrays whose channel size was 3 x 3 mm. A detector unit was composed of four scintillator blocks, 16 Si-PM arrays and a light guide. The developed detector unit had angled configuration since the light guide was bending. A detector unit had three gaps with an angle of 5.625◦ between scintillator blocks. With these configurations, we could arrange 64 scintillator blocks in nearly circular shape (regular 64-sided polygon) using 16 detector units. The use of the smaller number of detector units could reduce the size of the front–end electronics circuits. The inner diameter of the developed detector ring was 260 mm. This size was similar to those of brain PET systems, so our breast-dedicated PET detector ring can measure not only breast but also brain. Measured radial, tangential and axial spatial resolution of the detector ring reconstructed by the filtered back-projection (FBP) algorithm were 2.1 mm FWHM, 2.0 mm FWHM and 1.7 mm FWHM at center of field of view (FOV), respectively. The sensitivity was 2.0% at center of the axial FOV. With the developed detector ring, we could obtain high resolution image of the breast phantom and the brain phantom. We conclude that our developed Si-PM-based detector ring is promising for a high resolution breast-dedicated PET system that can also be used for brain PET system. © 2017 Elsevier B.V. All rights reserved.
1. Introduction Positron emission tomography (PET) has contributed to diagnoses of various lesions. Especially, in clinical examination of breast cancer for patient with dense breast, PET images are useful because image contrast between breast tissue and breast cancer in PET images is clearer than that in mammographic images. However, with a whole body PET system, there is an attenuation of photons due to body other than breast when breast cancer is measured [1]. Consequently, small PET system dedicated to imaging of breast lesions was proposed [2]. Breastdedicated PET system can realize high sensitivity because photons attenuate due to only breast, as well as detectors can be positioned close to breast. In addition to higher sensitivity, breast-dedicated PET
system can realize higher spatial resolution than that of whole body PET system since deterioration of spatial resolution due to angular deviation is smaller in breast-dedicated PET system. Because of these merits of breast-dedicated PET, various breast-dedicated PET systems have been developed and used in clinical situation [3–7]. For these breast-dedicated PET systems, position sensitive photomultiplier tubes (PSPMTs) are used as photodetectors since PSPMT can provide high gain and the stability is good [3,4]. However, array size of PSPMT is relatively large for detector ring of breast-dedicated PET system, so the detector ring shape reduces the number of polygons. Small numbered polygonal detector ring has large gaps between detector blocks, and a part of projection data are missed due to these gaps. Moreover,
* Corresponding authors.
E-mail addresses: [email protected] (K. Nakanishi), [email protected] (S. Yamamoto). https://doi.org/10.1016/j.nima.2017.10.052 Received 23 August 2017; Received in revised form 13 October 2017; Accepted 18 October 2017 Available online 4 November 2017 0168-9002/© 2017 Elsevier B.V. All rights reserved.
K. Nakanishi et al.
Nuclear Inst. and Methods in Physics Research, A 880 (2018) 118–124
Fig. 1. Schematic diagram of developed detector unit in transverse direction (A), axial direction (B), and detector ring for breast-dedicated PET system with 16 detector units (C).
missing projection data leads to non-uniformity of the coincidence data sampling in field of view (FOV). Non-uniformity of data sampling sometimes artifact in reconstructed image [8,9], so breast-dedicated PET systems which have a small numbered polygonal detector ring need compensation for missing projection data to reduce artifact [10–12]. On the other hand, a silicon photomultiplier (Si-PM) is starting to be used for PET detectors instead of PSPMT [9,13,14]. Si-PM has smaller channel size than a PSPMT in addition to high gain that is similar to a PSPMT. Because of small channels of Si-PM, a detector ring can be made with more number of polygons that are nearly circular shape since more scintillator blocks can be arranged in a detector ring. A circular shape detector ring enables to sample the coincidence data uniformly in FOV, so image artifact is less likely to occur without compensation for missing projection data. However, development of such detector ring for breast-dedicated PET system was not reported yet. Here, we propose our prototype Si-PM based nearly circular-shaped (regular 64sided polygon) detector ring for breast-dedicated PET system.
Fig. 2. Photo of developed detector unit (left) and part of front-end electronics (right).
2. Materials and methods 2.2. Description of developed detector ring for breast-dedicated PET 2.1. Concept of developed detector ring for breast-dedicated PET Fig. 2 shows a photo of the detector unit and a part of the frontend electronics. We used 1.5 × 1.9 × 15 mm Lu1.9 Gd0.1 SiO5 (LGSO) scintillator pixels and arranged them in an 8 × 24 matrix to form scintillator blocks. Four scintillator blocks were optically coupled to Si-PM arrays (Hamamatsu S12642-050) with an angled light guide to form a detector unit. Since the light guide had angles of 5. 625◦ , we could arrange 64 scintillator blocks in a nearly circular shape (a regular 64-sided polygon) using 16 detector units. One reason of the use of the relatively small number of the detector units (in other word; the reason of the use of the relatively larger size vending shaped block detector units) was to reduce the size of the front-end electronics circuits. The other reason was to reduce the edges effect of the Anger type block detectors; the edges parts of the block detectors are difficult to resolve the scintillator pixels and the resolution was worse.
Fig. 1(A) and (B) show schematic diagrams of the transverse direction and the axial direction of our detector unit, respectively. The detector unit was composed of four scintillator blocks, 16 Si-PM arrays and a light guide. The developed detector unit had angled configuration since the light guide was bending. A detector unit had three gaps with an angle of 5.625◦ between scintillator blocks. With these configurations, we could arrange 64 scintillator blocks in a detector ring, i.e. a detector ring could be made with nearly circular shape (regular 64-sided polygon) using 16 detector units. The inner diameter of developed detector ring was 260 mm. This size was similar to those of brain PET systems [15,16], so our breast-dedicated PET detector ring could measure not only breast but also brain. 119
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by 2-dimensional (2D) filtered backprojection (FBP) algorithm or order subset expectation maximization (OSEM) algorithm. The almost same data acquisition system was used for the previously developed brain PET system [16]. 2.3. Performance measurement We measured performances of the developed detector ring with similar methods to the NEMA NU 4-2008 Standards. 2.3.1. Spatial resolution The radial, tangential and axial spatial resolutions of the detector ring were measured using a 0.25 mm diameter 22 Na point source. The lower energy threshold level was set to 350 keV. The lower threshold level was used for all measurements. The measurements were made at the center and moved off-center with 20 mm steps. The accumulated counts of each position were approximately 70k counts. The images of point source were reconstructed with FBP algorithm using with a ramp filter.
Fig. 3. Photo of nearly circular-shaped (regular 64-sided polygon) detector ring.
2.3.2. Sensitivity We measured the sensitivity profile using a 22 Na point source. The position of point source was changed in an axial direction in 5 mm steps.
The signals from the Si-PMs were fed to weight summing boards with small diameter coaxial cables. The weighted sum signals were converted to digital signals by 100 MHz analogue to digital (A–D) converters of the data acquisition system. The digitalized data were integrated for 320 ns, and positions were calculated by Anger principle with field programmable gate array (FPGA). To compensate for variation of gain of Si-PMs due to temperature changes, we included the compensation circuit in the electronics. The compensation circuit controls the bias voltage of the Si-PM arrays. We previously reported configuration and performances of the detector unit and electronics [17]. The detector units had energy resolution of 14.2%FWHM for 511 keV gamma photons and most of the pixels of scintillators were clearly resolved in two-dimensional histogram. Fig. 3 shows a photo of the developed detector ring composed of 16 detector units. We arranged 64 scintillator blocks in a nearly circular shape using 16 detector units. The axial field-of-view (FOV) was 50 mm. Fig. 4(A) shows container and a stand that was used for the positioning and light shielding for the developed detector ring. The stand can adjust the position the detector ring with levers. A photo of the detector ring set inside the container is shown in Fig. 4(B) with front-end electronics circuits. Coincidence events measured with the developed detector ring were fed to a computer in the list mode and sinogram was obtained by sorting the acquired data. The accidental coincidence measured by delayed coincidence was subtracted in the sorting process for the correction of accidental events. The data were reconstructed
2.3.3. Count rate characteristics The count rate performance was measured using a 16 cm diameter, 15 cm long cylindrical phantom which contained 40 MBq 18 F-FDG solution. We calculated the noise equivalent count rate (NECR) from the count data. 2.3.4. Temperature dependent sensitivity change We measured the temperature dependent sensitivity variation of the detector ring using a 22 Na point source with the compensation circuit for the temperature dependence of gain of the Si-PM. The point source was set at the center of FOV of the detector ring. The measurement was carried out by changing the room temperature from 20 ◦ C to 26.5 ◦ C. 2.3.5. Imaging of breast and brain phantoms We carried out imaging of breast phantom and brain phantom which contain 18 F-FDG with developed detector ring. We used the data acquired by measurement of cylindrical phantom for normalization. For imaging of breast phantom, we developed a breast phantom made of acrylic resin. Fig. 5(A) shows a photo of the breast phantom. The diameter of the phantom is 100 mm. In the phantom, there are four spheres which simulate tumor. Fig. 5(B) shows a drawing of the phantom’s transaxial plane which includes spheres. The diameters of the
Fig. 4. Photo of container and stand for detector ring (A) and detector ring set in container and front-end electronics circuits (B).
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Fig. 5. Photo of breast phantom used for imaging (A) and drawing of phantom’s transaxial plane which includes spheres (B).
Fig. 7. Sensitivity profile as function of axial position.
Fig. 6. Radial, tangential and axial spatial resolution as function of distance from center.
3.2. Sensitivity We show the sensitivity profile as a function of the axial position in Fig. 7. The peak of the sensitivity profile was 2.0% at the center of the axial FOV.
spheres were 6 mm, 5 mm, 4 mm and 3 mm. We contained 300 kBq of F18-FDG in the breast phantom. The ratio of radioactivity concentration of background to sphere was 1 to 4. We positioned the phantom at the center of FOV of the detector ring and measurement was made for 20 min. The images were reconstructed using the FBP and the OSEM algorithm (3 iterations, 6 subsets) after single slice rebinning. Assuming uniform attenuation of photons for a 90 mm diameter cylindrical subject, we corrected for photon attenuation. In regard to imaging of brain phantom, we used the 3D Hoffman brain phantom [18]. We contained 20 MBq of18 F-FDG in the phantom. We positioned the phantom at the center of FOV of the detector ring and measurement was made for 40 min. Images were reconstructed with the FBP and the OSEM algorithm (6 iterations, 9 subsets) after single slice rebinning. We assumed a 200 mm diameter cylindrical subject for attenuation correction.
3.3. Count rate characteristics Fig. 8 shows the count rate characteristics of the developed system. The maximum prompt minus delayed count rate was 53-kcps. The peak of NECR was 26-kcps at 20 MBq in the phantom. 3.4. Temperature dependent sensitivity change Fig. 9 shows the count rate variation of the detector ring with a compensation circuit as a function of room temperature. The count rate was changed within ±10% for 6.5◦ . 3.5. Imaging of breast and brain phantoms
3. Results Fig. 10(A) and (B) show transaxial images of the breast phantom reconstructed by the FBP and the OSEM algorithm, respectively. The sphere of 4 mm diameter could be imaged by both reconstruction algorithm which was obvious from the profiles of the images. There was no significant image artifact in the images reconstructed by FBP and OSEM algorithm. Images of the brain phantom reconstructed by the FBP and the OSEM algorithm were shown in Fig. 11(A) and (B), respectively. We could
3.1. Spatial resolution Fig. 6 shows the radial, tangential and axial spatial resolution of the detector ring reconstructed by the filtered back-projection (FBP) algorithm. The radial, tangential and axial spatial resolution were 2.1 mm FWHM, 2.0 mm FWHM and 1.7 mm FWHM at center of field of view (FOV), respectively. 121
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is almost no artifact in images even reconstructed by FBP. Almost no artifact in reconstructed images without software compensation for the gaps between scintillators was a great advantage for the development of a PET system. Other advantage of the developed detector ring is the relatively larger diameter that can enable to measure other part of the body such as head. The larger detector ring diameter can also reduce the pile-up effect when the subject was large and was close to the detectors. Although the diameter of the developed detector ring is larger than those of other breast-dedicated PET systems [3,4] and angular deviation was larger, our detector ring has similar spatial resolution to other high resolution breast-dedicated PET systems [19] because of the smaller channels of Si-PM and scintillator pixels. We did not employ depth of interaction (DOI) detectors, so the spatial resolutions of the developed detector ring were deteriorated due to parallax error at the edges of FOV as shown in Fig. 6. The main reason of avoiding the DOI detectors was the high cost of the DOI detectors; number of the scintillator pixels increases as the number of DOI layers increases and the cost increases almost proportionally with the number of scintillator pixels for high resolution PET systems like breast-dedicated PETs. However, in clinical situations of the breast imaging, the subject size is much smaller than the transaxial FOV and only the central area of the FOV will be used. Therefore, deterioration of the spatial resolution at the edges of FOV may not be a problem in breast imaging. We used through-silicon-via (TSV) type Si-PMs for the detector ring. One of advantages of TSV type Si-PM is that the dark noise of TSV type Si-PM is less than that of other type Si-PMs. Consequently, if axial FOV of our detector ring is increased by increasing number of Si-PMs in axial direction, deterioration of performances such as the accuracy of position calculation by Anger principle due to the increase of the dark noises may be small. The sensitivity of the developed detector ring can be improved by increasing axial FOV. Although the gain of the Si-PM seriously changed without compensation for temperature dependency [20], the sensitivity of the developed detector ring changed within ±10% with compensation circuit. This result indicates the compensation circuit could reduce the sensitivity change of the PET system. However, to achieve the sensitivity changes similar to the PMT based PET system, room temperature or detector temperature needs to be controlled. The developed detector ring is basically compatible with magnetic resonance imaging (MRI) system because Si-PM is insensitivity to the static magnetic field. Thus we may be able to use the Si-PM based breastdedicated PET detector ring in a MRI if the magnetic components are changed to nonmagnetic materials. Since our detector ring is compact
Fig. 8. Prompt, delayed, prompt minus delayed, and NECR of detector ring.
Fig. 9. Sensitivity change with compensation circuit as function of room temperature.
clearly observe the structure of the brain phantom. Image artifacts were not observed in the images reconstructed by both algorithms except for the streaking artifacts in FBP images. 4. Discussion We successfully developed a Si-PM based circular detector ring for breast-dedicated PET system. One of advantages of the detector ring
Fig. 10. Images and profiles of breast phantom reconstructed with FBP (A) and OSEM (B).
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Fig. 11. Images of brain phantom reconstructed with FBP (A) and OSEM (B).
and the outer diameter of the ring is small, developed detector ring can easily be set in narrow space of the MRI gantry. Recently, PET-MRI has been playing an important role for diagnosis [21–23] and our developed detector ring may be used for such a purpose.
resolution reconstructed images of the phantoms without artifact using a simple reconstruction algorithm. We conclude that our developed SiPM-based detector ring is promising for high resolution breast-dedicated PET system that can also be used for brain PET system.
5. Conclusion
Acknowledgment
We developed a Si-PM-based circular-shaped detector ring using 16 detector units. With the developed detector ring, we could acquire high
This work was supported by Health Labor Sciences Research Grant, Japan (201438115A). 123
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References
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Contents lists available at ScienceDirect
Nuclear Inst. and Methods in Physics Research, A journal homepage: www.elsevier.com/locate/nima
Development of a circular shape Si-PM-based detector ring for breast-dedicated PET system Kouhei Nakanishi a, *, Seiichi Yamamoto a, *, Hiroshi Watabe b , Shinji Abe c , Naotoshi Fujita c , Katsuhiko Kato a a b c
Department of Medical Technology, Nagoya University Graduate School of Medicine, Nagoya, Japan Tohoku University CYRIC, Sendai, Japan Department of Radiological Technology, Nagoya University Hospital, Nagoya, Japan
a r t i c l e
i n f o
Keywords: PEM PET Breast Si-PM LGSO Brain
a b s t r a c t In clinical situations, various breast-dedicated positron emission tomography (PET) systems have been used. However, clinical breast-dedicated PET systems have polygonal detector ring. Polygonal detector ring sometimes causes image artifact, so complicated reconstruction algorithm is needed to reduce artifact. Consequently, we developed a circular detector ring for breast-dedicated PET to obtain images without artifact using a simple reconstruction algorithm. We used Lu1.9 Gd0.1 SiO5 (LGSO) scintillator block which was made of 1.5 x 1.9 x 15 mm pixels that were arranged in an 8 x 24 matrix. As photodetectors, we used silicon photomultiplier (Si-PM) arrays whose channel size was 3 x 3 mm. A detector unit was composed of four scintillator blocks, 16 Si-PM arrays and a light guide. The developed detector unit had angled configuration since the light guide was bending. A detector unit had three gaps with an angle of 5.625◦ between scintillator blocks. With these configurations, we could arrange 64 scintillator blocks in nearly circular shape (regular 64-sided polygon) using 16 detector units. The use of the smaller number of detector units could reduce the size of the front–end electronics circuits. The inner diameter of the developed detector ring was 260 mm. This size was similar to those of brain PET systems, so our breast-dedicated PET detector ring can measure not only breast but also brain. Measured radial, tangential and axial spatial resolution of the detector ring reconstructed by the filtered back-projection (FBP) algorithm were 2.1 mm FWHM, 2.0 mm FWHM and 1.7 mm FWHM at center of field of view (FOV), respectively. The sensitivity was 2.0% at center of the axial FOV. With the developed detector ring, we could obtain high resolution image of the breast phantom and the brain phantom. We conclude that our developed Si-PM-based detector ring is promising for a high resolution breast-dedicated PET system that can also be used for brain PET system. © 2017 Elsevier B.V. All rights reserved.
1. Introduction Positron emission tomography (PET) has contributed to diagnoses of various lesions. Especially, in clinical examination of breast cancer for patient with dense breast, PET images are useful because image contrast between breast tissue and breast cancer in PET images is clearer than that in mammographic images. However, with a whole body PET system, there is an attenuation of photons due to body other than breast when breast cancer is measured [1]. Consequently, small PET system dedicated to imaging of breast lesions was proposed [2]. Breastdedicated PET system can realize high sensitivity because photons attenuate due to only breast, as well as detectors can be positioned close to breast. In addition to higher sensitivity, breast-dedicated PET
system can realize higher spatial resolution than that of whole body PET system since deterioration of spatial resolution due to angular deviation is smaller in breast-dedicated PET system. Because of these merits of breast-dedicated PET, various breast-dedicated PET systems have been developed and used in clinical situation [3–7]. For these breast-dedicated PET systems, position sensitive photomultiplier tubes (PSPMTs) are used as photodetectors since PSPMT can provide high gain and the stability is good [3,4]. However, array size of PSPMT is relatively large for detector ring of breast-dedicated PET system, so the detector ring shape reduces the number of polygons. Small numbered polygonal detector ring has large gaps between detector blocks, and a part of projection data are missed due to these gaps. Moreover,
* Corresponding authors.
E-mail addresses: [email protected] (K. Nakanishi), [email protected] (S. Yamamoto). https://doi.org/10.1016/j.nima.2017.10.052 Received 23 August 2017; Received in revised form 13 October 2017; Accepted 18 October 2017 Available online 4 November 2017 0168-9002/© 2017 Elsevier B.V. All rights reserved.
K. Nakanishi et al.
Nuclear Inst. and Methods in Physics Research, A 880 (2018) 118–124
Fig. 1. Schematic diagram of developed detector unit in transverse direction (A), axial direction (B), and detector ring for breast-dedicated PET system with 16 detector units (C).
missing projection data leads to non-uniformity of the coincidence data sampling in field of view (FOV). Non-uniformity of data sampling sometimes artifact in reconstructed image [8,9], so breast-dedicated PET systems which have a small numbered polygonal detector ring need compensation for missing projection data to reduce artifact [10–12]. On the other hand, a silicon photomultiplier (Si-PM) is starting to be used for PET detectors instead of PSPMT [9,13,14]. Si-PM has smaller channel size than a PSPMT in addition to high gain that is similar to a PSPMT. Because of small channels of Si-PM, a detector ring can be made with more number of polygons that are nearly circular shape since more scintillator blocks can be arranged in a detector ring. A circular shape detector ring enables to sample the coincidence data uniformly in FOV, so image artifact is less likely to occur without compensation for missing projection data. However, development of such detector ring for breast-dedicated PET system was not reported yet. Here, we propose our prototype Si-PM based nearly circular-shaped (regular 64sided polygon) detector ring for breast-dedicated PET system.
Fig. 2. Photo of developed detector unit (left) and part of front-end electronics (right).
2. Materials and methods 2.2. Description of developed detector ring for breast-dedicated PET 2.1. Concept of developed detector ring for breast-dedicated PET Fig. 2 shows a photo of the detector unit and a part of the frontend electronics. We used 1.5 × 1.9 × 15 mm Lu1.9 Gd0.1 SiO5 (LGSO) scintillator pixels and arranged them in an 8 × 24 matrix to form scintillator blocks. Four scintillator blocks were optically coupled to Si-PM arrays (Hamamatsu S12642-050) with an angled light guide to form a detector unit. Since the light guide had angles of 5. 625◦ , we could arrange 64 scintillator blocks in a nearly circular shape (a regular 64-sided polygon) using 16 detector units. One reason of the use of the relatively small number of the detector units (in other word; the reason of the use of the relatively larger size vending shaped block detector units) was to reduce the size of the front-end electronics circuits. The other reason was to reduce the edges effect of the Anger type block detectors; the edges parts of the block detectors are difficult to resolve the scintillator pixels and the resolution was worse.
Fig. 1(A) and (B) show schematic diagrams of the transverse direction and the axial direction of our detector unit, respectively. The detector unit was composed of four scintillator blocks, 16 Si-PM arrays and a light guide. The developed detector unit had angled configuration since the light guide was bending. A detector unit had three gaps with an angle of 5.625◦ between scintillator blocks. With these configurations, we could arrange 64 scintillator blocks in a detector ring, i.e. a detector ring could be made with nearly circular shape (regular 64-sided polygon) using 16 detector units. The inner diameter of developed detector ring was 260 mm. This size was similar to those of brain PET systems [15,16], so our breast-dedicated PET detector ring could measure not only breast but also brain. 119
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by 2-dimensional (2D) filtered backprojection (FBP) algorithm or order subset expectation maximization (OSEM) algorithm. The almost same data acquisition system was used for the previously developed brain PET system [16]. 2.3. Performance measurement We measured performances of the developed detector ring with similar methods to the NEMA NU 4-2008 Standards. 2.3.1. Spatial resolution The radial, tangential and axial spatial resolutions of the detector ring were measured using a 0.25 mm diameter 22 Na point source. The lower energy threshold level was set to 350 keV. The lower threshold level was used for all measurements. The measurements were made at the center and moved off-center with 20 mm steps. The accumulated counts of each position were approximately 70k counts. The images of point source were reconstructed with FBP algorithm using with a ramp filter.
Fig. 3. Photo of nearly circular-shaped (regular 64-sided polygon) detector ring.
2.3.2. Sensitivity We measured the sensitivity profile using a 22 Na point source. The position of point source was changed in an axial direction in 5 mm steps.
The signals from the Si-PMs were fed to weight summing boards with small diameter coaxial cables. The weighted sum signals were converted to digital signals by 100 MHz analogue to digital (A–D) converters of the data acquisition system. The digitalized data were integrated for 320 ns, and positions were calculated by Anger principle with field programmable gate array (FPGA). To compensate for variation of gain of Si-PMs due to temperature changes, we included the compensation circuit in the electronics. The compensation circuit controls the bias voltage of the Si-PM arrays. We previously reported configuration and performances of the detector unit and electronics [17]. The detector units had energy resolution of 14.2%FWHM for 511 keV gamma photons and most of the pixels of scintillators were clearly resolved in two-dimensional histogram. Fig. 3 shows a photo of the developed detector ring composed of 16 detector units. We arranged 64 scintillator blocks in a nearly circular shape using 16 detector units. The axial field-of-view (FOV) was 50 mm. Fig. 4(A) shows container and a stand that was used for the positioning and light shielding for the developed detector ring. The stand can adjust the position the detector ring with levers. A photo of the detector ring set inside the container is shown in Fig. 4(B) with front-end electronics circuits. Coincidence events measured with the developed detector ring were fed to a computer in the list mode and sinogram was obtained by sorting the acquired data. The accidental coincidence measured by delayed coincidence was subtracted in the sorting process for the correction of accidental events. The data were reconstructed
2.3.3. Count rate characteristics The count rate performance was measured using a 16 cm diameter, 15 cm long cylindrical phantom which contained 40 MBq 18 F-FDG solution. We calculated the noise equivalent count rate (NECR) from the count data. 2.3.4. Temperature dependent sensitivity change We measured the temperature dependent sensitivity variation of the detector ring using a 22 Na point source with the compensation circuit for the temperature dependence of gain of the Si-PM. The point source was set at the center of FOV of the detector ring. The measurement was carried out by changing the room temperature from 20 ◦ C to 26.5 ◦ C. 2.3.5. Imaging of breast and brain phantoms We carried out imaging of breast phantom and brain phantom which contain 18 F-FDG with developed detector ring. We used the data acquired by measurement of cylindrical phantom for normalization. For imaging of breast phantom, we developed a breast phantom made of acrylic resin. Fig. 5(A) shows a photo of the breast phantom. The diameter of the phantom is 100 mm. In the phantom, there are four spheres which simulate tumor. Fig. 5(B) shows a drawing of the phantom’s transaxial plane which includes spheres. The diameters of the
Fig. 4. Photo of container and stand for detector ring (A) and detector ring set in container and front-end electronics circuits (B).
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Fig. 5. Photo of breast phantom used for imaging (A) and drawing of phantom’s transaxial plane which includes spheres (B).
Fig. 7. Sensitivity profile as function of axial position.
Fig. 6. Radial, tangential and axial spatial resolution as function of distance from center.
3.2. Sensitivity We show the sensitivity profile as a function of the axial position in Fig. 7. The peak of the sensitivity profile was 2.0% at the center of the axial FOV.
spheres were 6 mm, 5 mm, 4 mm and 3 mm. We contained 300 kBq of F18-FDG in the breast phantom. The ratio of radioactivity concentration of background to sphere was 1 to 4. We positioned the phantom at the center of FOV of the detector ring and measurement was made for 20 min. The images were reconstructed using the FBP and the OSEM algorithm (3 iterations, 6 subsets) after single slice rebinning. Assuming uniform attenuation of photons for a 90 mm diameter cylindrical subject, we corrected for photon attenuation. In regard to imaging of brain phantom, we used the 3D Hoffman brain phantom [18]. We contained 20 MBq of18 F-FDG in the phantom. We positioned the phantom at the center of FOV of the detector ring and measurement was made for 40 min. Images were reconstructed with the FBP and the OSEM algorithm (6 iterations, 9 subsets) after single slice rebinning. We assumed a 200 mm diameter cylindrical subject for attenuation correction.
3.3. Count rate characteristics Fig. 8 shows the count rate characteristics of the developed system. The maximum prompt minus delayed count rate was 53-kcps. The peak of NECR was 26-kcps at 20 MBq in the phantom. 3.4. Temperature dependent sensitivity change Fig. 9 shows the count rate variation of the detector ring with a compensation circuit as a function of room temperature. The count rate was changed within ±10% for 6.5◦ . 3.5. Imaging of breast and brain phantoms
3. Results Fig. 10(A) and (B) show transaxial images of the breast phantom reconstructed by the FBP and the OSEM algorithm, respectively. The sphere of 4 mm diameter could be imaged by both reconstruction algorithm which was obvious from the profiles of the images. There was no significant image artifact in the images reconstructed by FBP and OSEM algorithm. Images of the brain phantom reconstructed by the FBP and the OSEM algorithm were shown in Fig. 11(A) and (B), respectively. We could
3.1. Spatial resolution Fig. 6 shows the radial, tangential and axial spatial resolution of the detector ring reconstructed by the filtered back-projection (FBP) algorithm. The radial, tangential and axial spatial resolution were 2.1 mm FWHM, 2.0 mm FWHM and 1.7 mm FWHM at center of field of view (FOV), respectively. 121
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is almost no artifact in images even reconstructed by FBP. Almost no artifact in reconstructed images without software compensation for the gaps between scintillators was a great advantage for the development of a PET system. Other advantage of the developed detector ring is the relatively larger diameter that can enable to measure other part of the body such as head. The larger detector ring diameter can also reduce the pile-up effect when the subject was large and was close to the detectors. Although the diameter of the developed detector ring is larger than those of other breast-dedicated PET systems [3,4] and angular deviation was larger, our detector ring has similar spatial resolution to other high resolution breast-dedicated PET systems [19] because of the smaller channels of Si-PM and scintillator pixels. We did not employ depth of interaction (DOI) detectors, so the spatial resolutions of the developed detector ring were deteriorated due to parallax error at the edges of FOV as shown in Fig. 6. The main reason of avoiding the DOI detectors was the high cost of the DOI detectors; number of the scintillator pixels increases as the number of DOI layers increases and the cost increases almost proportionally with the number of scintillator pixels for high resolution PET systems like breast-dedicated PETs. However, in clinical situations of the breast imaging, the subject size is much smaller than the transaxial FOV and only the central area of the FOV will be used. Therefore, deterioration of the spatial resolution at the edges of FOV may not be a problem in breast imaging. We used through-silicon-via (TSV) type Si-PMs for the detector ring. One of advantages of TSV type Si-PM is that the dark noise of TSV type Si-PM is less than that of other type Si-PMs. Consequently, if axial FOV of our detector ring is increased by increasing number of Si-PMs in axial direction, deterioration of performances such as the accuracy of position calculation by Anger principle due to the increase of the dark noises may be small. The sensitivity of the developed detector ring can be improved by increasing axial FOV. Although the gain of the Si-PM seriously changed without compensation for temperature dependency [20], the sensitivity of the developed detector ring changed within ±10% with compensation circuit. This result indicates the compensation circuit could reduce the sensitivity change of the PET system. However, to achieve the sensitivity changes similar to the PMT based PET system, room temperature or detector temperature needs to be controlled. The developed detector ring is basically compatible with magnetic resonance imaging (MRI) system because Si-PM is insensitivity to the static magnetic field. Thus we may be able to use the Si-PM based breastdedicated PET detector ring in a MRI if the magnetic components are changed to nonmagnetic materials. Since our detector ring is compact
Fig. 8. Prompt, delayed, prompt minus delayed, and NECR of detector ring.
Fig. 9. Sensitivity change with compensation circuit as function of room temperature.
clearly observe the structure of the brain phantom. Image artifacts were not observed in the images reconstructed by both algorithms except for the streaking artifacts in FBP images. 4. Discussion We successfully developed a Si-PM based circular detector ring for breast-dedicated PET system. One of advantages of the detector ring
Fig. 10. Images and profiles of breast phantom reconstructed with FBP (A) and OSEM (B).
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Fig. 11. Images of brain phantom reconstructed with FBP (A) and OSEM (B).
and the outer diameter of the ring is small, developed detector ring can easily be set in narrow space of the MRI gantry. Recently, PET-MRI has been playing an important role for diagnosis [21–23] and our developed detector ring may be used for such a purpose.
resolution reconstructed images of the phantoms without artifact using a simple reconstruction algorithm. We conclude that our developed SiPM-based detector ring is promising for high resolution breast-dedicated PET system that can also be used for brain PET system.
5. Conclusion
Acknowledgment
We developed a Si-PM-based circular-shaped detector ring using 16 detector units. With the developed detector ring, we could acquire high
This work was supported by Health Labor Sciences Research Grant, Japan (201438115A). 123
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