- Email: [email protected]
A simulation study of a C-shaped in-beam PET system for dose verification in carbon ion therapy
Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
Contents lists available at SciVerse ScienceDirect
Nuclear Instruments and Methods in Physics Research A journal homepage: www.elsevier.com/locate/nima
A simulation study of a C-shaped in-beam PET system for dose verification in carbon ion therapy Su Jung An a,b, Cheol-Ha Beak a,b, Kisung Lee c, Yong Hyun Chung a,b,n a
Department of Radiological Science, College of Health Science, Yonsei University, 1 Yonseidae-gil, Wonju, Kangwon-do 220-710, Republic of Korea Institute of Health Science, Yonsei University, Wonju, 220-710, Republic of Korea c Department of Radiologic Science, Korea University, Seoul, 136-703, Republic of Korea b
a r t i c l e i n f o
abstract
Article history: Received 4 April 2012 Received in revised form 8 August 2012 Accepted 23 September 2012 Available online 28 September 2012
The application of hadrons such as carbon ions is being developed for the treatment of cancer. The effectiveness of such a technique is due to the eligibility of charged particles in delivering most of their energy near the end of the range, called the Bragg peak. However, accurate verification of dose delivery is required since misalignment of the hadron beam can cause serious damage to normal tissue. PET scanners can be utilized to track the carbon beam to the tumor by imaging the trail of the hadroninduced positron emitters in the irradiated volume. In this study, we designed and evaluated (through Monte Carlo simulations) an in-beam PET scanner for monitoring patient dose in carbon beam therapy. A C-shaped PET and a partial-ring PET were designed to avoid interference between the PET detectors and the therapeutic carbon beam delivery. Their performance was compared with that of a full-ring PET scanner. The C-shaped, partial-ring, and full-ring scanners consisted of 14, 12, and 16 detector modules, respectively, with a 30.2 cm inner diameter for brain imaging. Each detector module was composed of a 13 13 array of 4.0 mm 4.0 mm 20.0 mm LYSO crystals and four round 25.4 mm diameter PMTs. To estimate the production yield of positron emitters such as 10C, 11C, and 15O, a cylindrical PMMA phantom (diameter, 20 cm; thickness, 20 cm) was irradiated with 170, 290, and 350 AMeV 12C beams using the GATE code. Phantom images of the three types of scanner were evaluated by comparing the longitudinal profile of the positron emitters, measured along the carbon beam as it passed a simulated positron emitter distribution. The results demonstrated that the development of a C-shaped PET scanner to characterize carbon dose distribution for therapy planning is feasible. & 2012 Elsevier B.V. All rights reserved.
Keywords: Range verification Hadron therapy GATE In-beam PET
1. Introduction The goal of radiation therapy is to deliver the maximum dose of radiation to the tumor while minimizing the exposure of
normal tissue [1]. To achieve this goal in cancer treatment, hadron radiotherapy using carbon ion beams has evolved into a practical modality due to its unique energy deposition at the Bragg peak and its high relative biological effectiveness (RBE) [2]. According
Fig. 1. Simulated PET scanner configurations in which the full-ring, partial-ring, and C-shaped scanners consisted of 16, 12, and 14 detector modules, respectively.
n Corresponding author at: Department of Radiological Science, College of Health Science, Yonsei University, 1 Yonseidae-gil, Wonju, Kangwon-do 220-710, Republic of Korea. Tel.: þ82 33 760 2477; fax: þ82 33 760 2562. E-mail address: [email protected] (Y. Hyun Chung).
0168-9002/$ - see front matter & 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.nima.2012.09.050
38
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
to the particle therapy co-operative group (PTCOG), as of March 2012, six heavy ion centers were operating in Japan, Germany and China, and 20 proton or carbon therapy centers are under construction world-wide [3]. In Korea, the Korea Heavy Ion Medical Accelerator (KHIMA) Development Project is being carried out by the Korea Institute of Radiological and Medical Sciences (KIRAMS). To make the best use of hadron therapy, a real time dose verification method is highly desirable because any misalignment of the carbon beam can cause serious damage to critical organs surrounding the tumor [4–6]. However, the primary particles in a hadron beam are completely absorbed in the body, which precludes the use of real-time monitoring of dose. As a tool for in vivo and non-invasive monitoring of the treatment dose distribution, PET scanners have been suggested because they can image the trail of the hadron-induced positron emitters produced by nuclear fragmentation reactions of the fast projectiles with target nuclei [7–10]. Many research groups have focused on three types of PET scanners (in-beam, in-room, and off-line) and have confirmed the feasibility for monitoring deviations in the delivered dose [11,12]. Although the off-line and in-room PET methods are easier to implement with commercial PET/CT scanners, the measured distribution of positron emitters is influenced by the time between irradiation and imaging. Due to practical limitations of the off-line and in-room techniques, dedicated in-beam PET
Fig. 2. The simulated geometry of image acquisition with five point sources at different positions on the carbon beam line.
combined with therapy beam-line is the most clinically useful technique for on-line investigation of the positron emitters induced by hadron irradiation. To avoid interference between the PET detectors and the hadron beam, various in-beam PET scanners have been devised, such as a partial-ring PET with limited angular coverage [1] and an OpenPET design with gaps between detector rings [13]. We proposed partial-ring and C-shaped [14] in-beam PET scanners for whole body imaging and dose verification in carbon beam therapy. In this study, small prototype scanners for brain imaging were modeled and their performances were evaluated via GATE simulation to show a proof-of-concept of whole body imaging. GATE is a Monte Carlo simulation code based on GEANT4 enabling to the modeling of emission tomography, transmission tomography and radiation therapy [15]. The dose distribution of the carbon beam and the yields and spatial distribution of the induced positron emitters were simulated for various carbon energies in a cylindrical phantom. Then, the positron emitters were imaged with the partial-ring and C-shaped PET scanners, and the dose distribution and the profile of PET images were compared. Finally, the performance of the partial-ring and C-shaped PET scanners was compared to that of the full-ring PET. We aim to apply our in-beam PET to Korea Heavy Ion Medical Accelerator (KHIMA).
Fig. 3. Simulation arrangement of a PET scanner with the incident 12C ion beam from the left. The sphere shows a 20 cm PMMA phantom, and the 12C beam was simulated as being delivered horizontally to the phantom.
Table 1 Hadronic models used in the GATE simulations. Hadronic process
particle
Geant4 processes
Geant4 models
Geant4 datasets
Energy range
Elastic scattering
Generic Ion All other particles
G4HadronElasticProcess G4ProtonInelasticProcess
G4LElastic G4HadronElastic
G4HadronElasticDataSet G4HadronElasticDataSet
– –
Inelastic process for protons
Protons
G4ProtonInelasticProcess
G4BinaryCascade
G4ProtonInelasticCrossSection
0 20 GeV
Inelastic process for ions
Generic ion Deuteron Triton Alpha
G4IonInelasticProcess G4DeuteronInelasticProcess G4TritonInelasticProcess G4AlphaInelasticProcess
G4QMDReaction G4QMDReaction G4QMDReaction G4QMDReaction
G4IonsShenCrossSection G4IonsShenCrossSection G4IonsShenCrossSection G4IonsShenCrossSection
0–20 GeV 0–20 GeV 0–20 GeV 0–20 GeV
Inelastic process for pions
pi þ , pi
G4PionPlusInelasticProcess G4PionMinusInelasticProcess
G4LEPionMinusInelastic G4PionPlusInelastic
G4HadronInelasticDataSet
–
Radiative capture (neutrons)
Neutron
G4HadronCaptureProcess
G4NeutronHPCapture G4LCapture
G4NeutronHPCaptureData G4HadronCaptureDataSet
0–20 MeV 14 MeV–20 GeV
Inelastic scattering for neutrons
Neutron
G4NeutronInelasticProcess
G4NeutronHpInelastic G4BinaryCascade
G4NeutronHPInelasticData G4NeutronInelasticCrossSection
0–20 MeV 14 MeV–20 GeV
Fission
Neutron
G4HadronFissionProcess
G4NeutronHPFission G4LFission
G4NeutronHPFissionData G4HadronFissionDataSet
0–20 MeV 14 MeV–20 GeV
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
2. Materials and methods 2.1. Design of dedicated in-beam PET scanner 2.1.1. Description of in-beam PET scanner design The PET scanners considered in this study have a ring diameter of 30.2 cm and an axial field of view of 5 cm for brain imaging. The full-ring system consists of 16 detector modules, and each module is composed of an LYSO array of 13 13 elements with a pixel size of 4 4 20 mm3 and four round PMTs 25.4 mm in diameter. To avoid blockage of the carbon beam by the detector modules, partial-ring and C-shaped PETs were designed. For the partial-ring design, two detector modules located in the incident beam path were removed from both ends, while for the C-shaped design, two modules on the incident beam side only were removed. Fig. 1 shows the simulated PET scanner configurations in which the full-ring, partial-ring, and C-shaped scanner consist of 16, 12, and 14 detector modules, respectively. 2.1.2. The performance test of in-beam PET scanners The spatial resolution of the reconstructed image was measured with a 10 MBq Na-22 point source (0.5 mm diameter), positioned centrally in the axial direction and at radial offsets ranging from 10 to 10 cm with 5-cm interval as shown in Fig. 2. The radial and tangential resolution components of the reconstructed images were measured by fitting Gaussian functions to the respective profiles along the center of mass of the reconstructed images of the point source and calculating their respective full-width at half maximum (FWHM). PET images were reconstructed using a filtered back-projection method. The sensitivity was measured with the same point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction across the entire field of view (FOV). The true coincidence counts were recorded at each position. These measurements were repeated for energy window of 250–750, 350–650 and 410–613 keV. Then, we compared the performance of the two proposed PET scanners to each other and to that of the full-ring PET [16].
39
We then simulated carbon ion beam irradiation on a PMMA phantom with 20 cm diameter and 20 cm height for estimating the yield and distribution of the positron emitter production. The carbon beam was delivered horizontally to the phantom, as shown in Fig. 3. Three beam energies were selected: 170, 290, and 350 AMeV, corresponding to the KHIMA applications. The beam profile in the transverse direction was assumed to be a Gaussian shape with a FWHM of 10 mm, and the beam energy spread was performed with a FWHM of 0.2%. The intensity of the carbon beam was 1 108 particles per second (pps). Three positron emitters, 11C, 10C, and 15O, were analyzed. This simulation output was used as an input source in the subsequent PET image acquisition simulations. To evaluate the feasibility of PET imaging for dose verification, the distribution of carbon-induced positron emitters in the PMMA phantom was imaged with the proposed PET scanners. The source activity in the PMMA phantom was calculated based on these assumptions. – The carbon beam has a tunable time structure given by a repetition of 2 s spill length and 3 s pause. – Total therapy time is 2 min. – PET images are acquired during the half-life of 11C nuclei.
Phantom images were reconstructed by filtered back-projection, and the longitudinal profiles of the reconstructed images were calculated. Then, the effectiveness of dose verification was evaluated by comparing the image profile measured by PET and the distribution of the positron emitter production.
2.2. Production of positron-emitting nuclei by carbon ion and PET image acquisition In this work, GATE version 6.1 based on GEANT4 version 9.4 was used. As recommended by Open GATE collaboration, the Standard Physics List option 3 configurations were used to describe the energy loss of primary and secondary charged particles. Table 1 summarizes the Hadronic models implemented in the simulations. ‘‘DoseActor’’ was used to calculate the dose distributions and ‘‘ProductionAndStoppingActor’’ was used to estimate the 11C, 15O and 10C distributions. To validate the simulator, yields of positron-emitting nuclei produced by 212.12, 259.5 and 343.46 AMeV carbon beams in a PMMA phantom (9 9 30 cm3) were simulated and compared to simulated and measured data from other groups [5,6].
Fig. 5. The radial and tangential resolution were measured with a 10 MBq Na-22 point source, positioned centrally in the axial direction and at radial offsets ranging from 10 to 10 cm.
Fig. 4. The sinograms and reconstructed images of point sources at different positions.
40
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
Fig. 6. The sensitivities were measured with a 10 MBq Na-22 point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction and the three different energy window settings. (a) Radial direction sensitivity (250–750 keV), (b) radial direction sensitivity (350–650 keV), (c) radial direction sensitivity (410–613 keV), (d) axial direction sensitivity (250–750 keV), (e) axial direction sensitivity (350–650 keV) and (f) axial direction sensitivity (410–613 keV).
Table 2 Calculated yields of positron-emitting nuclei produced by 212.12, 259.5, and 343.46 AMeV 212.12 AMeV
11 10
C C O
15
12
C ions (per beam particle, in %).
259.5 AMeV
343.46 AMeV
MCHIT
Experiment
GATE6
MCHIT
Experiment
GATE6
MCHIT
Experiment
GATE6
11.9 1.97 2.38
10.57 1.3 0.87 0.3 2.17 0.3
9.73 1.38 1.71
16.83 2.79 3.69
14.7 71.6 1.2 70.3 3.1 70.4
13.11 1.85 2.36
25.25 4.27 6.09
19.9 7 2.4 1.5 7 0.3 5.07 0.4
18.43 2.57 3.48
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
41
3. Results 3.1. Performance of 3 types of PET scanner Fig. 4 shows sinograms and reconstructed images of point sources at different positions obtained by the three PET scanners. With the partial-ring and C-shaped scanners, images of the point sources in the near-peripheral region of FOV were blurred, while the full-ring scanner shows relatively uniform imaging characteristics over the entire FOV. The radial and tangential spatial resolutions are illustrated in Fig. 5 for different radial offsets across the FOV. The values of sensitivity, measured with a 10 MBq Na-22 point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction and the three different energy window settings, are illustrated in Fig. 6. The radial sensitivity depends on the number of detector modules and the scanner geometry. 3.2. Carbon beam induced positron emitters Table 2 shows the simulated yields of positron emitters produced by 212.12, 259.5, and 343.46 AMeV of 12C ions. MCHIT and experimental data from a previous study [5,6] are also shown in Table 2 for comparison. The yield of the positron emitters is defined as the number of positron emitters per incident carbon ion. The GATE outputs agreed well with the experimental data, reflecting that the GATE model is quite successful in calculating the absolute yields of positron emitters. After simulator verification, the yields of 11C, 10C, and 15O produced by 170, 290, and 350 AMeV carbon beams (selected by the KHIMA), respectively, in the PMMA cylinder phantom were simulated, and the results are shown in Table 3. The yield of 11C is almost 6 times higher than those of 15O and 10C, and the amount of positron emitters increases with the incident beam energy. The contribution from 11C nuclei dominates. Fig. 7 shows the spatial distribution of positron emitters with the corresponding depth–dose distribution of the carbon beam. The Bragg peaks are located at depths of 56 mm, 140 mm, and 190 mm for the 170, 290, and 350 AMeV carbon beams, respectively. The relative errors between peak of positron activity and Bragg peak are 3.6%, 5.7% and 6.3% in 170, 290 and 350 AMeV, respectively. The peak of positron activity is located very close to the Bragg peak and is a very attractive feature of PET to indicate the possibility of 12C beam range monitoring. 3.3. Acquisition of PET images Reconstructed PET images and their longitudinal profiles are shown in Fig. 8. In order to evaluate the performance of the scanners, the simulated dose distribution and b þ activity distribution are also shown. There exists a decisive correlation between the PET image profiles and the dose distributions, showing that PET is a valuable tool for range verification of carbon therapy. Table 3 Calculated yields of positron-emitting nuclei produced by 170, 290, and 350 AMeV 12 C ions (per beam particle, in %).
11 10
C C O
15
170 AMeV
290 AMeV
350 AMeV
6.86 0.99 1.17
15.06 2.08 2.79
18.80 2.62 3.57
Fig. 7. Calculated depth distributions of deposited energy (solid curves) and positron activity for (a) 170, (b) 290, and (c) 350 AMeV 12C nuclei in the PMMA phantom. The distributions of 11C, 10C, and 15O nuclei are shown by dotted, dash dotted, and dash dot dotted histograms, respectively. Their sum is represented by the dashed line histogram.
The range of 170 AMeV carbon beam was located left of center of the FOV. In this case, there was no significant difference between the images obtained by the partial-ring and C-shaped
42
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
Fig. 8. Reconstructed PET images and their longitudinal profiles (dotted line) for (a) 170, (b) 290, and (c) 350 AMeV 12C nuclei in the PMMA phantom using GATE simulations. The simulated dose distribution and b þ activity distribution are also shown for comparison (solid and dashed lines, respectively).
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
scanners. However, when the range of the carbon beam was behind the center of FOV (such as for the 290 and 350 AMeV carbon beams), the C-shaped scanner showed better image quality. Consequently, the C-shaped scanner is more suitable as an in-beam PET scanner.
4. Discussion and conclusions We proposed C-shaped and partial-ring scanners for dose verification in carbon ion therapy. In the simulation study, the C-shaped scanner showed better spatial resolution and sensitivity. Although the imaging performance of the C-shaped PET scanner was worse than that of the full-ring scanner, especially at the periphery of the FOV, it is a more feasible design because it does not interrupt the carbon therapy beam. The peak of the image profile is close to, but not exactly at, the position of the Bragg peak because the dose distribution and the positron activity distribution are generated by different processes. The dose distribution is formed by a Coulomb interaction between the carbon ions and the target ions, while the activity distribution is created by a nuclear reaction between the projectiles and the target nuclei [13,17]. However, the two curves have similar peak positions. The projectile fragments (such as 11 C and 10C) were created by peripheral nucleus–nucleus collisions and have similar velocities. At the same initial velocity, the range of the ions is proportional to A/Z2, where A and Z represent the ion mass and charge, respectively. Therefore, the ranges of 11C and 10C are shorter than the range of projectile ions. However, the maximum of the positron emitter distribution is located very close to the Brag peak because 11C and 10C nuclei can be created at any depth within the path of 12C, except at the end of the beam range where 12C is not energetic enough for a fragmentation reaction. This phenomenon explains the gap between the solid and dotted lines in Fig. 8. In Fig. 8, small peaks are observed at 25, 70, and 100 mm due to the gaps between detector modules. A compensation algorithm to remove the effects of the gaps is required. With the partial-ring and C-shaped scanners, the position of the maximum yield of positron emitters was successfully measured, but the positron yields between the beam entrance and the peak were not imaged well (unlike with the full-ring scanner) because of the loss of information due to the removed detector modules. This study was focused on the hardware design to monitor the Bragg peak, but the image correction and reconstruction methods can be improved to achieve better imaging performance.
43
PET images from the C-shaped and partial-ring structure can be improved by using statistical iterative reconstruction methods in combination with time of flight (TOF) information. Our group is developing the OSEM reconstruction method based on prior knowledge. Because most of the annihilation photons are generated along the path of carbon beam, we can approximately estimate the origin of photon pairs from the detected line of response. In conclusion, the results of this simulation study demonstrate that development of C-shaped PET scanners to characterize the carbon dose distribution for therapy planning is feasible.
Acknowledgments This research was supported by the Heavy Ion Medical Accelerator Center (HIMAC), Korea Institute of Radiological & Medical Sciences (KIRAMS), 2011K000567.
References [1] S. Surti, W. Zou, M.E. Daube-Witherspoon, et al., Physics in Medicine and Biology 56 (2011) 2667. ¨ ¨ [2] M. Kramer, O. Jakel, T. Haberer, et al., Physics in Medicine and Biology 45 (2000) 3299. [3] Particle Therapy Co-operative Group. Available from: /http://ptcog.web.psi.chS. [4] W. Enghardt, P. Crespo, F. Fiedler, et al., Nuclear Instruments & Methods in Physics Research Section A 525 (2004) 284. [5] Igor Pshenichnov, Igor Mishustin, Walter Greiner, Physics in Medicine and Biology 51 (2006) 6099. [6] K. Parodi, Ph.D. Dissertation, Technische Universit€, Dresden, 2004. [7] W. Enghardt, W.D. Fromm, H. Geissel, et al., Physics in Medicine and Biology 37 (1992) 2127. [8] M. Lazzeroni, A. Brahme, Physics in Medicine and Biology 56 (2011) 1585. [9] J. Pawelke, W. Enghardt, Haberer Th., et al., IEEE Transactions on Nuclear Science NS-44 (1997) 1492. [10] K. Parodi, T. Bartfeld, T. Haberer, International Journal of Radiation Oncology Biology Physics 71 (2008) 945. [11] D.W. Litzenberg, D.A. Roberts, M.Y. Lee, et al., Medical Physics 26 (1999) 992. [12] T. Yamaya, E. Yoshida, T. Inaniwa, et al., Physics in Medicine and Biology 56 (2011) 1123. [13] F. Fiedler, P. Crespo, K. Parodi, et al., IEEE Transactions on Nuclear Science NS-53 (2006) 2252. [14] M. Furuta, K. Kitamura, J. Ohi et al., Nuclear Science Symposium Conference Record, NSS 09, IEEE, 2009, p. 2548. [15] S. Jan, D. Benoit, E. Becheva, et al., Physics in Medicine and Biology 56 (2011) 881. [16] M.C. a~ nadas, P. Arce, P. Rato Mendes, et al., Physics in Medicine and Biology 56 (2011) 273. [17] F. Sommerer, F. Cerutti, K. Parodi, et al., Physics in Medicine and Biology 54 (2009) 3979.
Contents lists available at SciVerse ScienceDirect
Nuclear Instruments and Methods in Physics Research A journal homepage: www.elsevier.com/locate/nima
A simulation study of a C-shaped in-beam PET system for dose verification in carbon ion therapy Su Jung An a,b, Cheol-Ha Beak a,b, Kisung Lee c, Yong Hyun Chung a,b,n a
Department of Radiological Science, College of Health Science, Yonsei University, 1 Yonseidae-gil, Wonju, Kangwon-do 220-710, Republic of Korea Institute of Health Science, Yonsei University, Wonju, 220-710, Republic of Korea c Department of Radiologic Science, Korea University, Seoul, 136-703, Republic of Korea b
a r t i c l e i n f o
abstract
Article history: Received 4 April 2012 Received in revised form 8 August 2012 Accepted 23 September 2012 Available online 28 September 2012
The application of hadrons such as carbon ions is being developed for the treatment of cancer. The effectiveness of such a technique is due to the eligibility of charged particles in delivering most of their energy near the end of the range, called the Bragg peak. However, accurate verification of dose delivery is required since misalignment of the hadron beam can cause serious damage to normal tissue. PET scanners can be utilized to track the carbon beam to the tumor by imaging the trail of the hadroninduced positron emitters in the irradiated volume. In this study, we designed and evaluated (through Monte Carlo simulations) an in-beam PET scanner for monitoring patient dose in carbon beam therapy. A C-shaped PET and a partial-ring PET were designed to avoid interference between the PET detectors and the therapeutic carbon beam delivery. Their performance was compared with that of a full-ring PET scanner. The C-shaped, partial-ring, and full-ring scanners consisted of 14, 12, and 16 detector modules, respectively, with a 30.2 cm inner diameter for brain imaging. Each detector module was composed of a 13 13 array of 4.0 mm 4.0 mm 20.0 mm LYSO crystals and four round 25.4 mm diameter PMTs. To estimate the production yield of positron emitters such as 10C, 11C, and 15O, a cylindrical PMMA phantom (diameter, 20 cm; thickness, 20 cm) was irradiated with 170, 290, and 350 AMeV 12C beams using the GATE code. Phantom images of the three types of scanner were evaluated by comparing the longitudinal profile of the positron emitters, measured along the carbon beam as it passed a simulated positron emitter distribution. The results demonstrated that the development of a C-shaped PET scanner to characterize carbon dose distribution for therapy planning is feasible. & 2012 Elsevier B.V. All rights reserved.
Keywords: Range verification Hadron therapy GATE In-beam PET
1. Introduction The goal of radiation therapy is to deliver the maximum dose of radiation to the tumor while minimizing the exposure of
normal tissue [1]. To achieve this goal in cancer treatment, hadron radiotherapy using carbon ion beams has evolved into a practical modality due to its unique energy deposition at the Bragg peak and its high relative biological effectiveness (RBE) [2]. According
Fig. 1. Simulated PET scanner configurations in which the full-ring, partial-ring, and C-shaped scanners consisted of 16, 12, and 14 detector modules, respectively.
n Corresponding author at: Department of Radiological Science, College of Health Science, Yonsei University, 1 Yonseidae-gil, Wonju, Kangwon-do 220-710, Republic of Korea. Tel.: þ82 33 760 2477; fax: þ82 33 760 2562. E-mail address: [email protected] (Y. Hyun Chung).
0168-9002/$ - see front matter & 2012 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.nima.2012.09.050
38
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
to the particle therapy co-operative group (PTCOG), as of March 2012, six heavy ion centers were operating in Japan, Germany and China, and 20 proton or carbon therapy centers are under construction world-wide [3]. In Korea, the Korea Heavy Ion Medical Accelerator (KHIMA) Development Project is being carried out by the Korea Institute of Radiological and Medical Sciences (KIRAMS). To make the best use of hadron therapy, a real time dose verification method is highly desirable because any misalignment of the carbon beam can cause serious damage to critical organs surrounding the tumor [4–6]. However, the primary particles in a hadron beam are completely absorbed in the body, which precludes the use of real-time monitoring of dose. As a tool for in vivo and non-invasive monitoring of the treatment dose distribution, PET scanners have been suggested because they can image the trail of the hadron-induced positron emitters produced by nuclear fragmentation reactions of the fast projectiles with target nuclei [7–10]. Many research groups have focused on three types of PET scanners (in-beam, in-room, and off-line) and have confirmed the feasibility for monitoring deviations in the delivered dose [11,12]. Although the off-line and in-room PET methods are easier to implement with commercial PET/CT scanners, the measured distribution of positron emitters is influenced by the time between irradiation and imaging. Due to practical limitations of the off-line and in-room techniques, dedicated in-beam PET
Fig. 2. The simulated geometry of image acquisition with five point sources at different positions on the carbon beam line.
combined with therapy beam-line is the most clinically useful technique for on-line investigation of the positron emitters induced by hadron irradiation. To avoid interference between the PET detectors and the hadron beam, various in-beam PET scanners have been devised, such as a partial-ring PET with limited angular coverage [1] and an OpenPET design with gaps between detector rings [13]. We proposed partial-ring and C-shaped [14] in-beam PET scanners for whole body imaging and dose verification in carbon beam therapy. In this study, small prototype scanners for brain imaging were modeled and their performances were evaluated via GATE simulation to show a proof-of-concept of whole body imaging. GATE is a Monte Carlo simulation code based on GEANT4 enabling to the modeling of emission tomography, transmission tomography and radiation therapy [15]. The dose distribution of the carbon beam and the yields and spatial distribution of the induced positron emitters were simulated for various carbon energies in a cylindrical phantom. Then, the positron emitters were imaged with the partial-ring and C-shaped PET scanners, and the dose distribution and the profile of PET images were compared. Finally, the performance of the partial-ring and C-shaped PET scanners was compared to that of the full-ring PET. We aim to apply our in-beam PET to Korea Heavy Ion Medical Accelerator (KHIMA).
Fig. 3. Simulation arrangement of a PET scanner with the incident 12C ion beam from the left. The sphere shows a 20 cm PMMA phantom, and the 12C beam was simulated as being delivered horizontally to the phantom.
Table 1 Hadronic models used in the GATE simulations. Hadronic process
particle
Geant4 processes
Geant4 models
Geant4 datasets
Energy range
Elastic scattering
Generic Ion All other particles
G4HadronElasticProcess G4ProtonInelasticProcess
G4LElastic G4HadronElastic
G4HadronElasticDataSet G4HadronElasticDataSet
– –
Inelastic process for protons
Protons
G4ProtonInelasticProcess
G4BinaryCascade
G4ProtonInelasticCrossSection
0 20 GeV
Inelastic process for ions
Generic ion Deuteron Triton Alpha
G4IonInelasticProcess G4DeuteronInelasticProcess G4TritonInelasticProcess G4AlphaInelasticProcess
G4QMDReaction G4QMDReaction G4QMDReaction G4QMDReaction
G4IonsShenCrossSection G4IonsShenCrossSection G4IonsShenCrossSection G4IonsShenCrossSection
0–20 GeV 0–20 GeV 0–20 GeV 0–20 GeV
Inelastic process for pions
pi þ , pi
G4PionPlusInelasticProcess G4PionMinusInelasticProcess
G4LEPionMinusInelastic G4PionPlusInelastic
G4HadronInelasticDataSet
–
Radiative capture (neutrons)
Neutron
G4HadronCaptureProcess
G4NeutronHPCapture G4LCapture
G4NeutronHPCaptureData G4HadronCaptureDataSet
0–20 MeV 14 MeV–20 GeV
Inelastic scattering for neutrons
Neutron
G4NeutronInelasticProcess
G4NeutronHpInelastic G4BinaryCascade
G4NeutronHPInelasticData G4NeutronInelasticCrossSection
0–20 MeV 14 MeV–20 GeV
Fission
Neutron
G4HadronFissionProcess
G4NeutronHPFission G4LFission
G4NeutronHPFissionData G4HadronFissionDataSet
0–20 MeV 14 MeV–20 GeV
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
2. Materials and methods 2.1. Design of dedicated in-beam PET scanner 2.1.1. Description of in-beam PET scanner design The PET scanners considered in this study have a ring diameter of 30.2 cm and an axial field of view of 5 cm for brain imaging. The full-ring system consists of 16 detector modules, and each module is composed of an LYSO array of 13 13 elements with a pixel size of 4 4 20 mm3 and four round PMTs 25.4 mm in diameter. To avoid blockage of the carbon beam by the detector modules, partial-ring and C-shaped PETs were designed. For the partial-ring design, two detector modules located in the incident beam path were removed from both ends, while for the C-shaped design, two modules on the incident beam side only were removed. Fig. 1 shows the simulated PET scanner configurations in which the full-ring, partial-ring, and C-shaped scanner consist of 16, 12, and 14 detector modules, respectively. 2.1.2. The performance test of in-beam PET scanners The spatial resolution of the reconstructed image was measured with a 10 MBq Na-22 point source (0.5 mm diameter), positioned centrally in the axial direction and at radial offsets ranging from 10 to 10 cm with 5-cm interval as shown in Fig. 2. The radial and tangential resolution components of the reconstructed images were measured by fitting Gaussian functions to the respective profiles along the center of mass of the reconstructed images of the point source and calculating their respective full-width at half maximum (FWHM). PET images were reconstructed using a filtered back-projection method. The sensitivity was measured with the same point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction across the entire field of view (FOV). The true coincidence counts were recorded at each position. These measurements were repeated for energy window of 250–750, 350–650 and 410–613 keV. Then, we compared the performance of the two proposed PET scanners to each other and to that of the full-ring PET [16].
39
We then simulated carbon ion beam irradiation on a PMMA phantom with 20 cm diameter and 20 cm height for estimating the yield and distribution of the positron emitter production. The carbon beam was delivered horizontally to the phantom, as shown in Fig. 3. Three beam energies were selected: 170, 290, and 350 AMeV, corresponding to the KHIMA applications. The beam profile in the transverse direction was assumed to be a Gaussian shape with a FWHM of 10 mm, and the beam energy spread was performed with a FWHM of 0.2%. The intensity of the carbon beam was 1 108 particles per second (pps). Three positron emitters, 11C, 10C, and 15O, were analyzed. This simulation output was used as an input source in the subsequent PET image acquisition simulations. To evaluate the feasibility of PET imaging for dose verification, the distribution of carbon-induced positron emitters in the PMMA phantom was imaged with the proposed PET scanners. The source activity in the PMMA phantom was calculated based on these assumptions. – The carbon beam has a tunable time structure given by a repetition of 2 s spill length and 3 s pause. – Total therapy time is 2 min. – PET images are acquired during the half-life of 11C nuclei.
Phantom images were reconstructed by filtered back-projection, and the longitudinal profiles of the reconstructed images were calculated. Then, the effectiveness of dose verification was evaluated by comparing the image profile measured by PET and the distribution of the positron emitter production.
2.2. Production of positron-emitting nuclei by carbon ion and PET image acquisition In this work, GATE version 6.1 based on GEANT4 version 9.4 was used. As recommended by Open GATE collaboration, the Standard Physics List option 3 configurations were used to describe the energy loss of primary and secondary charged particles. Table 1 summarizes the Hadronic models implemented in the simulations. ‘‘DoseActor’’ was used to calculate the dose distributions and ‘‘ProductionAndStoppingActor’’ was used to estimate the 11C, 15O and 10C distributions. To validate the simulator, yields of positron-emitting nuclei produced by 212.12, 259.5 and 343.46 AMeV carbon beams in a PMMA phantom (9 9 30 cm3) were simulated and compared to simulated and measured data from other groups [5,6].
Fig. 5. The radial and tangential resolution were measured with a 10 MBq Na-22 point source, positioned centrally in the axial direction and at radial offsets ranging from 10 to 10 cm.
Fig. 4. The sinograms and reconstructed images of point sources at different positions.
40
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
Fig. 6. The sensitivities were measured with a 10 MBq Na-22 point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction and the three different energy window settings. (a) Radial direction sensitivity (250–750 keV), (b) radial direction sensitivity (350–650 keV), (c) radial direction sensitivity (410–613 keV), (d) axial direction sensitivity (250–750 keV), (e) axial direction sensitivity (350–650 keV) and (f) axial direction sensitivity (410–613 keV).
Table 2 Calculated yields of positron-emitting nuclei produced by 212.12, 259.5, and 343.46 AMeV 212.12 AMeV
11 10
C C O
15
12
C ions (per beam particle, in %).
259.5 AMeV
343.46 AMeV
MCHIT
Experiment
GATE6
MCHIT
Experiment
GATE6
MCHIT
Experiment
GATE6
11.9 1.97 2.38
10.57 1.3 0.87 0.3 2.17 0.3
9.73 1.38 1.71
16.83 2.79 3.69
14.7 71.6 1.2 70.3 3.1 70.4
13.11 1.85 2.36
25.25 4.27 6.09
19.9 7 2.4 1.5 7 0.3 5.07 0.4
18.43 2.57 3.48
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
41
3. Results 3.1. Performance of 3 types of PET scanner Fig. 4 shows sinograms and reconstructed images of point sources at different positions obtained by the three PET scanners. With the partial-ring and C-shaped scanners, images of the point sources in the near-peripheral region of FOV were blurred, while the full-ring scanner shows relatively uniform imaging characteristics over the entire FOV. The radial and tangential spatial resolutions are illustrated in Fig. 5 for different radial offsets across the FOV. The values of sensitivity, measured with a 10 MBq Na-22 point source, stepped at 1-cm increments in the axial direction and at 5-cm increments in the radial direction and the three different energy window settings, are illustrated in Fig. 6. The radial sensitivity depends on the number of detector modules and the scanner geometry. 3.2. Carbon beam induced positron emitters Table 2 shows the simulated yields of positron emitters produced by 212.12, 259.5, and 343.46 AMeV of 12C ions. MCHIT and experimental data from a previous study [5,6] are also shown in Table 2 for comparison. The yield of the positron emitters is defined as the number of positron emitters per incident carbon ion. The GATE outputs agreed well with the experimental data, reflecting that the GATE model is quite successful in calculating the absolute yields of positron emitters. After simulator verification, the yields of 11C, 10C, and 15O produced by 170, 290, and 350 AMeV carbon beams (selected by the KHIMA), respectively, in the PMMA cylinder phantom were simulated, and the results are shown in Table 3. The yield of 11C is almost 6 times higher than those of 15O and 10C, and the amount of positron emitters increases with the incident beam energy. The contribution from 11C nuclei dominates. Fig. 7 shows the spatial distribution of positron emitters with the corresponding depth–dose distribution of the carbon beam. The Bragg peaks are located at depths of 56 mm, 140 mm, and 190 mm for the 170, 290, and 350 AMeV carbon beams, respectively. The relative errors between peak of positron activity and Bragg peak are 3.6%, 5.7% and 6.3% in 170, 290 and 350 AMeV, respectively. The peak of positron activity is located very close to the Bragg peak and is a very attractive feature of PET to indicate the possibility of 12C beam range monitoring. 3.3. Acquisition of PET images Reconstructed PET images and their longitudinal profiles are shown in Fig. 8. In order to evaluate the performance of the scanners, the simulated dose distribution and b þ activity distribution are also shown. There exists a decisive correlation between the PET image profiles and the dose distributions, showing that PET is a valuable tool for range verification of carbon therapy. Table 3 Calculated yields of positron-emitting nuclei produced by 170, 290, and 350 AMeV 12 C ions (per beam particle, in %).
11 10
C C O
15
170 AMeV
290 AMeV
350 AMeV
6.86 0.99 1.17
15.06 2.08 2.79
18.80 2.62 3.57
Fig. 7. Calculated depth distributions of deposited energy (solid curves) and positron activity for (a) 170, (b) 290, and (c) 350 AMeV 12C nuclei in the PMMA phantom. The distributions of 11C, 10C, and 15O nuclei are shown by dotted, dash dotted, and dash dot dotted histograms, respectively. Their sum is represented by the dashed line histogram.
The range of 170 AMeV carbon beam was located left of center of the FOV. In this case, there was no significant difference between the images obtained by the partial-ring and C-shaped
42
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
Fig. 8. Reconstructed PET images and their longitudinal profiles (dotted line) for (a) 170, (b) 290, and (c) 350 AMeV 12C nuclei in the PMMA phantom using GATE simulations. The simulated dose distribution and b þ activity distribution are also shown for comparison (solid and dashed lines, respectively).
S. Jung An et al. / Nuclear Instruments and Methods in Physics Research A 698 (2013) 37–43
scanners. However, when the range of the carbon beam was behind the center of FOV (such as for the 290 and 350 AMeV carbon beams), the C-shaped scanner showed better image quality. Consequently, the C-shaped scanner is more suitable as an in-beam PET scanner.
4. Discussion and conclusions We proposed C-shaped and partial-ring scanners for dose verification in carbon ion therapy. In the simulation study, the C-shaped scanner showed better spatial resolution and sensitivity. Although the imaging performance of the C-shaped PET scanner was worse than that of the full-ring scanner, especially at the periphery of the FOV, it is a more feasible design because it does not interrupt the carbon therapy beam. The peak of the image profile is close to, but not exactly at, the position of the Bragg peak because the dose distribution and the positron activity distribution are generated by different processes. The dose distribution is formed by a Coulomb interaction between the carbon ions and the target ions, while the activity distribution is created by a nuclear reaction between the projectiles and the target nuclei [13,17]. However, the two curves have similar peak positions. The projectile fragments (such as 11 C and 10C) were created by peripheral nucleus–nucleus collisions and have similar velocities. At the same initial velocity, the range of the ions is proportional to A/Z2, where A and Z represent the ion mass and charge, respectively. Therefore, the ranges of 11C and 10C are shorter than the range of projectile ions. However, the maximum of the positron emitter distribution is located very close to the Brag peak because 11C and 10C nuclei can be created at any depth within the path of 12C, except at the end of the beam range where 12C is not energetic enough for a fragmentation reaction. This phenomenon explains the gap between the solid and dotted lines in Fig. 8. In Fig. 8, small peaks are observed at 25, 70, and 100 mm due to the gaps between detector modules. A compensation algorithm to remove the effects of the gaps is required. With the partial-ring and C-shaped scanners, the position of the maximum yield of positron emitters was successfully measured, but the positron yields between the beam entrance and the peak were not imaged well (unlike with the full-ring scanner) because of the loss of information due to the removed detector modules. This study was focused on the hardware design to monitor the Bragg peak, but the image correction and reconstruction methods can be improved to achieve better imaging performance.
43
PET images from the C-shaped and partial-ring structure can be improved by using statistical iterative reconstruction methods in combination with time of flight (TOF) information. Our group is developing the OSEM reconstruction method based on prior knowledge. Because most of the annihilation photons are generated along the path of carbon beam, we can approximately estimate the origin of photon pairs from the detected line of response. In conclusion, the results of this simulation study demonstrate that development of C-shaped PET scanners to characterize the carbon dose distribution for therapy planning is feasible.
Acknowledgments This research was supported by the Heavy Ion Medical Accelerator Center (HIMAC), Korea Institute of Radiological & Medical Sciences (KIRAMS), 2011K000567.
References [1] S. Surti, W. Zou, M.E. Daube-Witherspoon, et al., Physics in Medicine and Biology 56 (2011) 2667. ¨ ¨ [2] M. Kramer, O. Jakel, T. Haberer, et al., Physics in Medicine and Biology 45 (2000) 3299. [3] Particle Therapy Co-operative Group. Available from: /http://ptcog.web.psi.chS. [4] W. Enghardt, P. Crespo, F. Fiedler, et al., Nuclear Instruments & Methods in Physics Research Section A 525 (2004) 284. [5] Igor Pshenichnov, Igor Mishustin, Walter Greiner, Physics in Medicine and Biology 51 (2006) 6099. [6] K. Parodi, Ph.D. Dissertation, Technische Universit€, Dresden, 2004. [7] W. Enghardt, W.D. Fromm, H. Geissel, et al., Physics in Medicine and Biology 37 (1992) 2127. [8] M. Lazzeroni, A. Brahme, Physics in Medicine and Biology 56 (2011) 1585. [9] J. Pawelke, W. Enghardt, Haberer Th., et al., IEEE Transactions on Nuclear Science NS-44 (1997) 1492. [10] K. Parodi, T. Bartfeld, T. Haberer, International Journal of Radiation Oncology Biology Physics 71 (2008) 945. [11] D.W. Litzenberg, D.A. Roberts, M.Y. Lee, et al., Medical Physics 26 (1999) 992. [12] T. Yamaya, E. Yoshida, T. Inaniwa, et al., Physics in Medicine and Biology 56 (2011) 1123. [13] F. Fiedler, P. Crespo, K. Parodi, et al., IEEE Transactions on Nuclear Science NS-53 (2006) 2252. [14] M. Furuta, K. Kitamura, J. Ohi et al., Nuclear Science Symposium Conference Record, NSS 09, IEEE, 2009, p. 2548. [15] S. Jan, D. Benoit, E. Becheva, et al., Physics in Medicine and Biology 56 (2011) 881. [16] M.C. a~ nadas, P. Arce, P. Rato Mendes, et al., Physics in Medicine and Biology 56 (2011) 273. [17] F. Sommerer, F. Cerutti, K. Parodi, et al., Physics in Medicine and Biology 54 (2009) 3979.