Performance evaluation of a Compton SPECT imager for determining the position and distribution of 225Ac in targeted alpha therapy: A Monte Carlo simulation based phantom study

Applied Radiation and Isotopes 154 (2019) 108893

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Performance evaluation of a Compton SPECT imager for determining the position and distribution of 225Ac in targeted alpha therapy: A Monte Carlo simulation based phantom study

T

Taewoong Leea, Minho Kima, Wonho Leeb, Byoungsoo Kima, Ilhan Lima, Kanghyon Songa, Jongguk Kima,∗ a b

Korea Institute of Radiologic and Medical Sciences, Nowon-ro 75, Seoul, 01812, South Korea School of Health and Environmental Science, Korea University, Anam-ro 145, Seoul, 02841, South Korea

H I GH L IG H T S

performance of a Compton SPECT imager for Ac in TAT was evaluated. • The in vivo monitoring of the position and distribution of Ac was discussed. • The • A Monte Carlo simulation based phantom study was applied in this research. 225

225

A R T I C LE I N FO

A B S T R A C T

Keywords: Targeted alpha therapy (TAT) Compton SPECT imaging 225 Ac radionuclide

In this study, the performance of a Compton Single Photon Emission Computed Tomography (SPECT) imager when in vivo monitoring the position and distribution of 225Ac radionuclide in targeted alpha therapy (TAT) was evaluated. When 225Ac radionuclide, which emits various γ-rays (218 and 440 keV), is used in TAT, both the photoelectric and Compton scattering events can be used for image reconstruction. Moreover, all information pertaining to the various γ-rays of the 225Ac radionuclide can be individually or simultaneously utilized in the reconstructed image. Three types of simulation phantoms and a quantitative evaluation method were used to compare the performance of the Compton SPECT imager to that of conventional SPECT imaging, which uses only photoelectric events, and the results demonstrated that the Compton SPECT imager exhibited a higher performance as the effective count for the image reconstruction was higher. To verify the accuracy of the position and distribution of the 225Ac radionuclide that had been inserted into the phantom, reconstructed images of the various γ-rays were combined with cross-sectional images of the human phantom and all combined images were found to match the predetermined simulation conditions. In conclusion, the simulation results demonstrated the feasibility of the in vivo monitoring of the position and distribution of 225Ac radionuclide using the γ-rays in TAT.

1. Introduction

2016). However, as the migration of the daughter products from the Ac radionuclide increases the toxicity to excretion organs and normal tissues, the amount of 225Ac radiopharmaceuticals that can be used is limited (Miederer et al., 2004; Essler et al., 2012; deKruijff et al., 2015). Therefore, for successful tumor treatment, it is important to monitor the position and distribution of the small amount of α-emitting radionuclide that was inserted into the phantom or patient. Many imaging systems, including α-cameras, Cherenkov imaging systems and Compton camera have been developed to monitor the position and distribution of 225Ac or other α-emitting radionuclides in preclinical research. To acquire high resolution images of the α225

Targeted alpha therapy (TAT) utilizes an α-emitting radionuclide that releases high-linear energy transfer (LET) α-particles with a kinetic energy of a few MeV (McDevitt et al., 1998; Couturier et al., 2005; Mulford et al., 2005). Due to the high-LET and short range of the αparticles, TAT can deliver a concentrated radiation dose to a tumor cell while causing minimal harm to adjacent normal tissues (Kim and Brechbiel, 2012; Baidoo et al., 2013). Recently, 225Ac radionuclide, which has a relatively long half-life (t1/2 = 10 d), has been employed for TAT (Sofou et al., 2007; Miederer et al., 2008; Kratochwil et al.,

∗

Corresponding author. E-mail address: [email protected] (J. Kim).

https://doi.org/10.1016/j.apradiso.2019.108893 Received 18 February 2019; Received in revised form 19 July 2019; Accepted 9 September 2019 Available online 10 September 2019 0969-8043/ © 2019 Elsevier Ltd. All rights reserved.

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Fig. 1. Schematic of the

225

conventional PET has been proposed and developed. In 1996, using Monte Carlo simulation, Comanor proposed the algorithms to identify the detector Compton scatter in PET module. As a result, the accuracy of the reconstructed image was improved (Comanor et al., 1996). Rafecas et al. (2003) developed a prototype Compton PET consisting of LSO-APD PET using the inter-crystal scatter (ICS) in a dual layer. The detection efficiency of the detection system using ICS significantly improved without degrading the image quality. The Compton PET composed of CZT detectors was developed by Abbaszadeh et al. (2018) A high-performance system that uses both inter- and intra-crystal scattering events significantly improved the sensitivity and contrast-tonoise ratio (CNR) in reconstructed images. To increase the detection efficiency of γ-rays in SPECT imaging, the Compton SPECT technique, which can simultaneously utilize both photoelectric events and Compton scattering events for image reconstruction, was introduced in 2018 (Lee et al., 2018). Because this approach leverages Compton scattering events to increase the effective information in the reconstructed image, the performance of the reconstructed image was significantly improved without degrading the image resolution. Moreover, the quatitative evaluation based on the signal-to-noise ratio (SNR) proved the superiority of Compton SPECT over conventional SPECT. In this study, a Compton SPECT imaging system was used to simulate the position and distribution of the various γ-rays emitted by 225Ac radionuclides when inserted into a phantom and the reconstructed images for various γ-ray energies were obtained using appropriate energy windows. To quantitatively evaluate the performance of the Compton SPECT imager, various types of phantoms, such as contrast, spatial resolution, and human-like voxel phantoms, were employed. Moreover, the performance of Compton SPECT imager was compared to that of a conventional SPECT system. To verify the precise position and distribution of 225Ac radionuclides under realistic conditions, crosssectional images of the human-like phantom were obtained and combined with the reconstructed SPECT images of the γ-rays emitted by the 225 Ac radionuclide. The results demonstrate the feasibility of the in vivo monitoring of the position and distribution of 225Ac radionuclides in TAT.

Ac radionuclide decay chain.

particles emitted from 225Ac radionuclide, Back and Jacobsson (2010) constructed an α-camera based on autoradiography combined with a charge-coupled device (CCD). Even though the reconstructed images had a high spatial resolution, the camera was limited to ex vivo planar imaging due to the short range of α-particles. Pandya et al. (2016) showed the in vivo biodistribution imaging of 225Ac radionuclide using a Cherenkov luminescence imaging (CLI) system in preclinical research; however, the CLI system had difficulty measuring the Cherenkov radiation due to the low Cherenkov emission yields from 225Ac radionuclide. A Compton camera composed of two cerium-doped gadolinium aluminium gallium garnet (GAGG) scintillator detectors was developed to image 211At in TAT. The results of the experiment showed the capability of Compton imaging for high-energy photons emitted via electron capture decay of 211At (687 keV) and via alpha decay of its daughter nuclide 211Po (570 and 898 keV) (Nagao et al., 2018). The γ-ray images of single photon emission tomography (SPECT) have been utilized for monitoring the position and distribution of the αemitting radionuclide in TAT due to their relatively higher penetration and large sensitivity. As shown in Fig. 1, the decay of 225Ac radionuclide results in four α-emissions, and the 218- and 440-keV γ-rays are emitted from 221Fr and 213Bi of the daughter nuclides of 225Ac, respectively. Swart et al. (2016). demonstrated high-resolution 213Bi SPECT imaging using dual energy windows by combining the X-ray of 79 keV of and the γ-ray of 440 keV. The resulting reconstructed image had the highest performance among the results of individual energy window settings. Robertson et al. (2017) simultaneously used both SPECT imaging of 221Fr (218 keV) and 213Bi(440 keV) in preclinical imaging tests of 225Ac radiopharmaceuticals; however, the obtained images relied on individual energy window settings. Crawford et al. (2018) acquired 209At SPECT images and demonstrated the ability of quantitative SPECT imaging to image biodistributions with high spatial resolution. However, as these studies were based on the use of small amounts of α-emitting radionuclide, the detection efficiency was limited for γ-rays. To improve the performance of the reconstructed tomography images, a Compton PET with a combination of Compton cameras and

2. Materials and methods 2.1. Design parameters and GATE simulation of the Compton SPECT A Monte Carlo simulation using the Geant4 application for tomographic emission (GATE 7.0) code was conducted to evaluate the performance of the Compton SPECT system. Four gantry heads were used to achieve the highest spatial resolution of full-volume imaging (Kimura et al., 1990). The CZT crystals were used to construct a large-volume detector array and utilize the photoelectric and Compton scattering events for image reconstruction (Pratx and Levin, 2009). A squareshaped, parallel-hole collimator that was matched to the detector pixel was utilized because of the absence of magnification of the objects (Kim et al., 2015). As shown in Fig. 2, the Compton SPECT imager consists of four gantry heads composed of a 100 × 100 array of 3 mm × 3 mm × 6 mm virtual Frisch-grid CZT detectors, which were designed based on an actual model developed at Brookhaven National Laboratory (BNL) (Bolotnikov et al., 2004, 2006). To estimate the real performance of virtual Frisch-grid CZT SPECT system obtained by performing the Monte Carlo simulation, spatial blurring and energy broadening function in the GATE simulation code were applied for the simulation. Parallel-hole collimators that were matched to the detector pixels were utilized and the size of holes (d), height (l), and pixel pitch were 1.5, 30, and 3 mm, respectively. The level of septal penetration was less than approximately 5% and the required septal thickness (t) was determined as follows (Macey et al., 1995): 2

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Fig. 2. Schematic diagram of the Compton SPECT detection system and the KTMAN-2 phantom.

t≥

6d/ μ (Ei ) , l − (3/ μ (Ei ))

peak corresponds to the energy of the emitted radiation (221Fr: 81- and 218-keV and 213Bi: 79-, and 440-keV). Because the energy resolution of the CZT detectors was not enough to discriminate the 79 keV from 81 keV in the 225Ac radionuclide, these photons were not used for image reconstruction. Because CZT detectors can be used to obtain each interaction position and deposited energy, the CZT based Compton SPECT imaging system employed both photoelectric events and Compton scattering events for image reconstruction. It is therefore essential to determine the correct sequence of Compton scattering events in order to use them for SPECT reconstruction. In all simulations, the position of the γ-rays emitted from the 225Ac radionuclide were unknown a priori. Thus, the correct sequence was determined by comparing the deposited energies at each interaction position. According to Ref. Lehner et al. (2004), a higher fraction of the correct sequence guarantees a better reconstructed image. When determining the correct sequence, if the energy of the γ-ray was less than 400 keV, it was more probable that the first interaction deposited less energy than the second (ΔE1 < ΔE2). On the other hand, if the energy of the γ-ray was higher than 400 keV, it was more probable that the first interaction deposited more energy (ΔE1 > ΔE2). Hence, as shown in Fig. 4(a), events with a lower deposited energy and events with a higher deposited energy were chosen as the effective first interaction for the 218 keV and 440 keV rays, respectively. The incorrect determination for the radiation interaction sequence provides erroneous information, which causes artifacts in the reconstructed image. For incident photon energies from 100 keV to 1 MeV, Fig. 4(b) shows the fraction of simulated events in which the first interaction deposited more energy than the second interaction. At 218 and 440 keV approximately, the fractions of 80.2% and 54.7% were correctly identified. Table 1 summarizes the ratio of the correctly and miss-identified sequences for the 218 and 440 keV energies. After the correct interaction sequence was determined, the appropriate energy window for the image reconstruction of each 218- and 440 keV events was calculated based on the energy resolution of the CZT detectors which has 1.3% at 662 keV. The energy window for the effective Compton scattering events can be calculated from the energy resolution of a Compton camera consisting of CZT detector. The energy

(1)

where d is the size of the hole in the collimator, l is the height of the collimator, and μ(Ei) is the linear attenuation coefficient (cm−1) of tungsten at an incident energy of Ei. Here, Ei was set to 440 keV. Using Eq. (1), the septal thickness was determined to be 1.5 mm. The most significant 225Ac radionuclide γ-ray emissions that can be used for image reconstruction are at 218 keV (11.4%) and 440 keV (25.9%) (Fig. 1) (Miederer et al., 2008; Robertson et al., 2017). Therefore, the GATE simulation was configured so that both γ-rays were emitted from the 221Fr and 213Bi, respectively. The energy spectrum simulated by the CZT arrays of the Compton SPECT imager are shown in Fig. 3 where it can be seen that the energy spectrum includes the radiation peaks reflecting the 225Ac radionuclide decay chain and each

Fig. 3. Energy spectrum of the γ-rays obtained from the Compton SPECT imager. 3

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Fig. 4. (a) Comparison of the γ-ray deposited energies at each interaction position. (b) Fraction of simulated Compton sequences in which the first interaction deposits more energy than the second interaction as a function of the incident photon energy.

2.2. Evaluation of the Compton SPECT performance

Table 1 Comparison of the fractions for the interaction sequences of the Compton events. Energy (keV)

218 440

Counts of Compton sequences

Fractions (%)

E1 < E2

E1 > E2

Correct identification

Miss identification

26271 45932

6486 55323

80.2 54.7

19.8 45.3

Three simulation phantoms were used to evaluate the performance of the Compton SPECT imager for the γ-rays emitted from the 225Ac radionuclide. First, the contrast and contrast-to-noise ratios (CNRs) of the reconstructed image were calculated using a contrast phantom consisting of six hot rods. As shown in Fig. 5(a), the diameter and height of all rods were 1.5 cm and 15 cm, respectively. The activity per unit volume of each successive rod increased by 1.2 times relative to the rod with the minimum source activity per unit volume. The minimum source activity per unit volume and simulated acquisition time were 0.37 MBq and 300 s, respectively. The contrast (Cs) of each rod between the source and background regions was determined as follows (Wu et al., 2010):

resolution (ΔE) of the Compton camera is defined as (Watanabe et al., 2015):

Δ E=

2 (ΔEFirst CZT detector )2 + ΔESecond CZT detector

(2a)

Cs =

Table 2 shows the appropriate energy window photoelectric and Compton events at 218- and 440 keV (Bolotnikov et al., 2004, 2006).

218 440

Energy window of photoelectric event (%)

Energy window of Compton event (%)

± 2.27 ± 1.59

± 3.2 ± 2.25

(2b)

where hs and bs represent the average pixel value calculated within the source and background regions of interest (ROIs), respectively, and the subscript s represents the size of the rods. In order to calculate the noise, the variability (Ns) between the average of the ROI values was calculated as follows (Wu et al., 2010):

Table 2 Appropriate energy window for the photoelectric and Compton events at various radiation energies. Energy (keV)

hs − bs , hs

Ns =

σh2s + σb2s All ROIs

,

(3)

where σhs and σbs represent the standard deviation (SD) of the source and background ROIs, respectively, and the denominator ( All ROIs ) indicates the average of all the ROIs. Hence, the contrast-to-noise ratio (CNR) was defined as Cs/Ns. 4

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Fig. 5. Various simulation phantoms used in the performance evaluation of the Compton SPECT imager: (a) the contrast phantom, (b) resolution phantom, and (c) KTMAN-2 phantom (axial plane).

Fig. 6. Reconstructed image from the conventional and Compton SPECT imagers for the contrast phantom.

3. Simulation results and discussions

Second, as shown in Fig. 5(b), a resolution phantom with six segments was utilized, the diameters of each segment were 3.5, 3.6, 3.7, 3.8, 3.9, and 4.0 mm, the source activity per unit volume of each segment was 0.37 MBq, and the simulated acquisition time was 300 s. The spatial resolution and relative intensity of the reconstructed image for each segment were evaluated after image reconstruction via filtered back-projection (FBP) with a Ramp filter. Third, a KTMAN-2 human phantom was used to verify the precise position and distribution of the 225 Ac radionuclide under realistic conditions. As shown in Fig. 5(c), the cross-sectional image of the KTMAN-2 phantom provides the anatomical information of the phantom, and the 225Ac radionuclide was inserted into the prostate in KTMAN-2 human phantom. The reconstructed images were acquired using a source activity per unit volume containing 0.37 MBq of 225Ac radionuclide and a simulated acquisition time of 600 s. The reconstructed images of the 225Ac radionuclide were then combined with the cross-sectional images (axial plane) of the KTMAN-2 phantom. The KTMAN-2 consisted of 48 anatomical regions and each cell had the same material information as a real human cell. The entire phantom was composed of small voxels, each of which was 2 mm × 2 mm × 5 mm, and the numbers of voxels along the X, Y, and Z axes were 300, 150, and 344, respectively (Lee et al., 2006).

3.1. Comparison of the performance for the contrast phantom The performance of the Compton SPECT imager was quantitatively evaluated based on the contrast and CNR in comparison to the performance of a conventional SPECT imager. The average pixel value of each rod and the effective counts in the reconstructed images were calculated for each rod using ROIs with diameters that were 90% of the physical diameter of the rod. The reconstructed images using FBP with a Ramp filter for various γ-ray energies are shown in Fig. 6. As the probability of Compton scattering events is higher than that of photoelectric events at high energies, the effective counts of the reconstructed images increased as the radiation energy increased. The improvement of image quality after applying Compton kinematics was the highest for Compton SPECT, as the effective counts of Compton SPECT were higher than those of conventional SPECT. Therefore, as shown in Fig. 7, the Compton SPECT imager exhibited higher contrasts and CNRs when compared to those of the conventional SPECT imager because the effective count used in the image reconstruction was higher.

5

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Fig. 7. Evaluation of the contrast and CNR from the contrast phantom image: (a) contrast and (b) contrast-to-noise ratio (CNR).

Fig. 8. Reconstructed image from the conventional and Compton SPECT imagers for the resolution phantom.

image in each rod were evaluated using a resolution phantom. The reconstructed images using FBP with a Ramp filter and their projected histograms in the cross sections are shown in Figs. 8 and 9, respectively. As shown in Fig. 8, the images from the Compton SPECT imager

3.2. Performance comparison for the resolution phantom (Resolution phantom) The spatial resolution and relative intensity of the reconstructed 6

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Fig. 9. Evaluation of the cross sections from each rod in the resolution phantom image: (a) conventional SPECT and (b) Compton SPECT.

Fig. 10. Evaluation of the relative intensities from each rod in the resolution phantom image: (a) conventional SPECT and (b) Compton SPECT.

resolved (Fig. 9). Because the higher frequency components in the edge of all the reconstructed images were included, the high-frequency signals were emphasized by the filter; as a results, the rods at the edged were better resolved than the rods inside. As shown in Fig. 10, the relative intensities were normalized to the maximum pixel value of each rod in the segments. The uniformity of the relative intensity of the Compton SPECT images was more consistent

exhibited fewer artifacts than those from the conventional SPECT imager. The dotted line (red) in the figure represent the lines of the cross section. When the diameters of the rods in the conventional (218- and 218 + 440 keV) and Compton SPECT imagers (218- and 218 + 440 keV) were greater than 3.6 mm, the rods could be easily distinguished. However, in the reconstructed images of the conventional and Compton imager for 440 keV, only rods > 3.9 mm could be 7

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Table 3 Maximum pixel value and relative standard deviation of Conventional SPECT from each rod in the resolution phantom image. Rod size (mm)

Energy (keV)

Maximum pixel value (a.u.)

Relative standard deviation (%)

3.5

218 440 218 218 440 218 218 440 218 218 440 218 218 440 218 218 440 218

17.8 14.1 31.4 17.1 14.7 31.9 17.8 13.3 31.1 17.9 15.1 32.8 17.9 14.3 31.5 18.3 14.2 31.5

14.5 14.0 11.5 14.9 12.8 11.8 11.9 12.0 9.5 10.3 11.1 8.4 9.4 11.7 7.3 8.5 9.0 6.0

3.6

3.7

3.8

3.9

4.0

+ 440

+ 440

+ 440

+ 440

+ 440

+ 440

Table 5 Comparison of the sensitivity calculated for the conventional and Compton SPECT. Energy (keV)

218 440 218 + 440

Energy (keV)

Maximum pixel value (a.u.)

Relative standard deviation (%)

3.5

218 440 218 218 440 218 218 440 218 218 440 218 218 440 218 218 440 218

24.4 20.4 48.7 24.2 19.9 47.5 23.6 20.6 45.3 24.3 21.3 48.7 24.3 20.5 46.9 23.8 21.2 46.6

12.6 13.9 10.7 11.9 13.4 10.5 11.8 10.7 8.6 9.5 11.5 6.8 8.4 9.2 6.3 7.3 7.0 4.6

3.6

3.7

3.8

3.9

4.0

+ 440

+ 440

+ 440

+ 440

+ 440

+ 440

Conventional SPECT

Compton SPECT

4.5 × 10−3 5.08 × 10−4 5.0 × 10−3

6.4 × 10−3 1.2 × 10−3 7.6 × 10−3

than those of the conventional SPECT images. The maximum pixel values and average relative standard deviation (RSD) of the relative intensities for the conventional and Compton SPECT images are shown in Tables 3 and 4, respectively. The differences in the average RSD between the Compton SPECT images and the conventional SPECT images with a multi-window (218 + 440 keV) were small; however, the Compton SPECT images had higher maximum pixel values and lower RSDs than those of the conventional SPECT images. For the RSD, a lower value indicates better performance, as RSD demonstrate the fluctuation for the signal of the selected area. Therefore, by using Compton events, the image quality and the detection efficiency of Compton SPECT are significantly increased compared to conventional SPECT. In summary, the Compton imager exhibited the highest performance.

Table 4 Maximum pixel value and relative standard deviation of Compton SPECT from each rod in the resolution phantom image. Rod size (mm)

Sensitivity in reconstructed images

3.3. Reconstructed image of the KTMAN-2 phantom A KTMAN-2 human-like voxel phantom was utilized to realistically simulate the conditions of TAT. As shown in Fig. 5(c), the 225Ac radionuclide inserted into the prostate in KTMAN-2 phantom. To verify the accuracy of the position and distribution of the 225Ac radionuclide, the reconstructed images for various γ-ray energies were combined with the cross-sectional images of the KTMAN-2 phantom. Combined images of each reconstructed γ-ray image are shown in Fig. 11. The size of the pixels in the reconstructed image were 3 mm. The sensitivity was defined as the ratio of the number of effective counts used for image reconstruction to the total counts of the source activity; Table 5 summarizes the sensitivity for each modality from reconstructed images of the 225Ac radionuclide inserted into the prostate in KTMAN-2 phantom. Overall, the effective counts of the Compton SPECT imager were found to be higher than those of a conventional SPECT imager. The sensitivities of the Compton SPECT at 218- and 440 keV were respectively 1.4 and 2.4 times higher than that of the conventional SPECT.

Fig. 11. 225Ac uptake region (prostate) inserted into the KTMAN phantom. Images were reconstructed using the FBP method combined with cross-sectional images of the phantom: (a) conventional SPECT and (b) Compton SPECT. 8

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As expected, the reconstructed image from the Compton SPECT imager, which enabled the precise monitoring of the position and distribution the 225Ac radionuclide inserted into the prostate, was significantly better than the conventional SPECT images. From all perspectives, the position and distribution of the 225Ac radionuclide in the reconstructed images matched well with the simulation conditions. However, due to the finite number of photons and limited number of angular samples, the reconstructed image obtained using the FBP was slightly marred by image artifacts. In the case of the reconstructed image at the 440 keV energy, the performance of the reconstructed images decreased due to the increase in the background noise caused by the increase in the scattered radiation in the voxel phantom. 4. Conclusion In this study, the performance of a Compton SPECT imager for the in vivo monitoring of the position and distribution of 225Ac radionuclides in TAT was evaluated. In the simulation results, the position and distribution of the 225Ac radionuclide were identified by detecting the γrays at 218 and 440 keV that were emitted from the 225Ac radionuclide inserted into a phantom. Moreover, based on a quantitative evaluation using various phantoms, the performance of Compton SPECT imaging was evaluated and compared with that of conventional SPECT imaging, and in all of the results, the Compton imaging exhibited the highest performance. The results of this study confirm the feasibility of verifying the position and distribution of 225Ac radionuclide using Compton SPECT imaging. Acknowledgments This study was supported by a grant of the Korea Institute of Radiological and Medical Science (KIRAMS), funded by Ministry of Science and ICT(MSIT), Republic of Korea (50462-2019) Appendix A. Supplementary data Supplementary data to this article can be found online at https:// doi.org/10.1016/j.apradiso.2019.108893. References Abbaszadeh, S., Chinn, G., Levin, C.S., 2018. Positioning true coincidences that undergo inter- and intra-crystal scatter for a sub-mm resolution cadmium zinc telluride-based PET system. Phys. Med. Biol. 63, 025012. Back, T., Jacobsson, L., 2010. The alpha-camera: a quantitative digital autoradiography technique using a charge-coupled device for ex vivo high-resolution bioimaging of alpha-particles. J. Nucl. Med. 51, 1616–1623. Baidoo, K.E., Yong, K., Brechbiel, M.W., 2013. Molecular pathways: targeted α-particle radiation therapy. Clin. Cancer Res. 19, 530–537. Bolotnikov, A.E., Camarda, G.S., Carini, G.A., Wright, G.W., McGregor, D.S., McNeil, W., James, R.B., 2004. New results from performance studies of Frisch-grid CdZnTe detectors. Proc. SPIE 5540, 33–45. Bolotnikov, A.E., Camarda, G.C., Carini, G.A., Fiederle, M., Li, L., McGregor, D.S., McNeil, W., Wright, G.W., James, R.B., 2006. Performance characteristics of frisch-ring CdZnTe detectors. IEEE Trans. Nucl. Sci. 53, 607–614. Comanor, K.A., Virador, P.R.G., Moses, W.W., 1996. Algorithms to identify detector compton scatter in PET modules. IEEE Trans. Nucl. Sci. 43, 2213–2218. Couturier, O., Supiot, S., Degraef-Mougin, M., Faivre-Chauvet, A., Carlier, T., Chatal, J.F., Davodeau, F., Cherel, M., 2005. Cancer radioimmunotherapy with alpha-emitting nuclides. Eur. J. Nucl. Med. Mol. Imaging 32, 601–614. Crawford, J.R., Robertson, A.K.H., Yang, H., Rodríquez-Rodríquez, C., Esquinas, P.L.,

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