Experimental evaluation of the GE NM/CT 870 CZT clinical SPECT system equipped with WEHR and MEHRS collimator

Abstract Purpose A high‐energy‐resolution whole‐body SPECT‐CT device (NM/CT 870 CZT; C‐SPECT) equipped with a CZT detector has been developed and is being used clinically. A MEHRS collimator has also been developed recently, with an expected improvement in imaging accuracy using medium‐energy radionuclides. The objective of this study was to compare and analyze the accuracies of the following devices: a WEHR collimator and the MEHRS collimator installed on a C‐SPECT, and a NaI scintillation detector‐equipped Anger‐type SPECT (A‐SPECT) scanner, with a LEHR and LMEGP. Methods A line phantom was used to measure the energy resolutions including collimator characteristics in the planar acquisition of each device using 99mTc and 123I. We also measured the system's sensitivity and high‐contrast resolution using a lead bar phantom. We evaluated SPECT spatial resolution, high‐contrast resolution, radioactivity concentration linearity, and homogeneity, using a basic performance evaluation phantom. In addition, the effect of scatter correction was evaluated by varying the sub window (SW) employed for scattering correction. Results The energy resolution with 99mTc was 5.6% in C‐SPECT with WEHR and 9.9% in A‐SPECT with LEHR. Using 123I, the results were 9.1% in C‐SPECT with WEHR, 5.5% in C‐SPECT with MEHRS, and 10.4% in A‐SPECT with LMEGP. The planar spatial resolution was similar under all conditions, but C‐SPECT performed better in SPECT acquisition. High‐contrast resolution was improved in C‐SPECT under planar condition and SPECT. The sensitivity and homogeneity were improved by setting the SW for scattering correction to 3% of the main peak in C‐SPECT. Conclusion C‐SPECT demonstrates excellent energy resolution and improved high‐contrast resolution for each radionuclide. In addition, when using 123I, careful attention should be paid to SW for scatter correction. By setting the appropriate SW, C‐SPECT with MEHRS has an excellent scattered ray removal effect, and highly homogenous imaging is possible while maintaining the high‐contrast resolution.


| INTRODUCTION
Recent studies have reported the effectiveness of a gamma camera equipped with a cadmium-zinc-telluride (CZT) detector. Moreover, a device that was developed specifically for the heart has been used clinically mainly in the field of cardiology. [1][2][3][4][5][6][7][8] Upon application of high voltage to the electrodes sandwiching the crystal, the CZT detector measures the electric charge by collecting the generated electrons/holes when radiation enters the crystal and interacts with it. The energy required to generate an electron-hole pair is about 5 eV; this energy can be detected by various sensors and precisely measured, yielding an energy resolution higher than that of the Anger-type detector. 3 In addition, because the band gap in the CZT detector is large, it can be used at room temperature, and the high atomic number of CZT assists in efficient photoelectric absorption, leading to an improved system sensitivity compared with that of the Anger-type detector.
Thus, broad clinical applications of CZT-based imaging can be expected. Therefore, a whole-body single-photon emission computed tomography (SPECT)-computed tomography (CT) device using a CZT detector (C-SPECT) with these excellent system characteristics was developed. This device consists of a two-detector SPECT scanner and a 16-row CT scanner. Thirteen 4-cm × 4-cm CZT units are arranged in each detector in the x-direction and 10 in the y-direction. This device has an effective visual field of 51 cm × 39 cm. In addition, as compared with a conventional NaI scintillation detector-equipped Anger-type SPECT (A-SPECT), the improved contrast resolution afforded by higher energy resolution of this pixel-type detector may be effective in the clinical setting. Clinical studies using this device have included the verification of dual isotope imaging and short time acquisition using 99m Tc and 123 I in myocardial blood flow tests. 9,10 However, no studies have focused on the evaluation of the performance of the device itself. In addition to the wide-energy high-resolution (WEHR) collimator that is standard for this device, a medium-energy high-resolution sensitivity (MEHRS) collimator was recently developed. This collimator is expected to improve imaging accuracy of medium-energy radionuclides. Thus, verifying the performance of this device in clinical applications is of utmost importance. In this study, we analyzed the system performance of the C-SPECT device using 99m Tc (a low-energy radionuclide) and 123 I (a low-and medium-energy radionuclide), both of which are widely used in clinical practice. 11 The system performance assuming clinical indications was evaluated by comparing the device with a two-detector SPECT equipped with an NaI scintillation detector.  Table 1. Although the parameters of LMEGP were not disclosed, the system spatial resolution and sensitivity were LEHR × 0.70 and LEHR × 1.71 at the manufacturing stage respectively.  Radionuclide, 99m Tc-incardronate (Nihon Medi-Physics Co Ltd, Tokyo, Japan) and 123 I-IMP perfuzamine (Nihon Medi-Physics Co Ltd, Tokyo, Japan).

2.B | Planar acquisition
Energy resolution, spatial resolution, and system sensitivity measurements were performed using a method that conformed to the NEMA standards 13 to the best possible extent.

2.B.1 | Energy resolution
We sealed 84.1 (MBq /ml) of 99m Tc and 87.7 (MBq /ml) of 123 I in a selfmade line phantom with a diameter of 1.5 mm and length of 30.0 cm.
For C-SPECT_WEHR and A-SPECT_LEHR with 99m Tc, and for C-SPECT_WEHR, C-SPECT_MEHRS, and A-SPECT_LMEGP with 123 I, planar images were acquired to achieve 10 kilo counts (kct) at 140 and 159 keV peak energy for 99m Tc and 123 I respectively. Furthermore, the counting rate was set to ≤20 kct/s. Acquisitions of two placements of the line phantom were performed: one without a scatterer (scatter (−)), and one with a water-equivalent phantom (Kyoto Kagaku Co., Kyoto, Japan) placed 5 cm behind and 10 cm in front of the line phantom (scatter (+)). The conditions of data acquisition were as follows: C-SPECT was performed in a 512 × 512 matrix at 2 × magnification with a pixel size of 0.55 mm. A-SPECT was performed at 2 × magnification in a 512 × 512 matrix with a pixel size of 0.59 mm. The distance between the radiation source and the collimator was set to 10 cm.
From the obtained γ-ray photopeak energy spectrum, we calculated the energy resolution, including the collimator characteristics using the following Eq. (1), the photopeak energy Ep ( 99m Tc: 140 keV, 123 I:159 keV), and the peak width Δ Ep (full-width half-maximum; FWHM), equivalent to half the count at the peak. Tc for each collimator. Next, using the data obtained, 10 rows from the center under each acquisition condition of the line phantoms were added, and the FWHM was calculated from the profile curve.

2.B.3 | System sensitivity
Aqueous solutions of 196 MBq 99m Tc and 183 MBq 123 I were sealed in a circular container with a diameter of 9 cm and a height of 0.9 cm. Each collimator was installed, and data were acquired with a radiation source-collimator distance of 100 mm and a total count (Ct) of 4000 kct. The conditions of data acquisition were 2 × magnification with a 512 × 512 matrix under each condition. Next, using the acquired data, we set a rectangular region of interest (ROI) surrounding the entire image. We then measured the Ct of the entire image and calculated the in-air system sensitivity, Ssys (count/s/ MBq), from Eqs. (2) and (3).
Here, Rt is the count rate corrected for radioactive decay (cps), T is the acquisition start time, T acq is the acquisition time (seconds), which statistical noise could be sufficiently removed. The data acquisition conditions were set to 512 × 512 at 1 × magnification for each device. The pixel size was set to 1.1 mm for C-SPECT and 1.2 mm for A-SPECT, and each EW was set at 15%. Next, from the data obtained, the profile lines of 27 bars at a lead bar spacing of 2.0 mm, 23 bars at 2.5 mm, 18 bars at 3.0 mm, and 15 bars at 3.5 mm were calculated. In addition, the reference image was subjected to two-dimensional Fourier transformation for frequency evaluation. The radial-directional intensity distribution function Pr(n) was calculated to conduct a one-dimensional evaluation of the two-dimensional frequency distribution under each condition. 16 2.C | SPECT acquisition 2.C.1 | SPECT resolution evaluation As SC of C-SPECT, the lower sub window (SW) of 7% (119.7keV -129.5keV) in 99m Tc ( 99m Tc_SC7) and the lower SW of 7% (136.0keV -147.1keV) and upper SW of 7% (170.9keV -182.0keV) in 123 I ( 123 I_SC7) were also acquired with respect to the main peak.
Furthermore, the lower SW of 3% (125.3keV -129.5keV) in 99m Tc ( 99m Tc_SC3) and the lower SW of 3% (142.3keV -147.1keV) and upper SW of 3% (170.9keV -175.7keV) in 123 I ( 123 I_SC3) were also acquired with respect to the main peak. As SC of A-SPECT, 99m Tc_SC7 and 123 I_SC7 were also acquired with respect to the main peak.
Next, the six sealed pieces in the columnar part were filled with  Next, a 44-kBq/mL 123 I aqueous solution was sealed, and the data were acquired for 60 min under the same conditions as those for 99m Tc. Next, the obtained data were reconstructed in the same manner as used for the evaluation of the radioactivity concentration linearity. Furthermore, Pr(n) was calculated from the reconstructed axial image. . The same process was then repeated with a sealed 44-kBq/mL 123 I aqueous solution. Note that Ct is the average count value within the ROI, and SD is the standard deviation within the ROI. Figure 2 shows the energy spectrum for each device and collimator.

3.A | Planar acquisition
C-SPECT_WEHR had high relative counts in the Compton region for each radionuclide. With 123 I, the tendency was the strongest for C-SPECT_WEHR, but the relative count in the Compton region was greatly reduced for C-SPECT_MEHRS and was lower than that for A-SPECT _LMEGP. Moreover, the same tendency was observed with the inclusion of scatterers. With 99m Tc, the energy resolution was 5.6% for C-SPECT_WEHR and 9.9% for A-SPECT_LEHR. With 123 I, it was 9.1% for C-SPECT_WEHR, 5.5% for C-SPECT_MEHRS, and 10.4% for A-SPECT_LMEGP. Table 2   Moreover, A-SPECT_LMEGP had the lowest resolution. Furthermore, the FWHM was the same under each condition for both pixel sizes.
The same tendency was observed for the 15-mm diameter line phantom. For the evaluation of system sensitivity, Table 3 shows that with 99m Tc, the value was highest in A-SPECT_LEHR and lowest in C-SPECT_WEHR10. With 123 I, A-SPECT_LMEGP and C-SPECT_-WEHR were equivalent to each other and higher than C-SPECT_MEHRS. Figure 3 shows the bar phantom image and line profile of each count as a high-contrast resolution evaluation. In C-SPECT at 500 and 1000 kct, a bar of 3.0 mm matched the actual number of bars with the number of count peaks shown in the line profile. A-SPECT was inseparable for all bar widths. In addition, from the frequency evaluation, C-SPECT_WEHR15 showed falling signal strength from about 1.60 cycles/cm onward (Fig. 4). However, A-SPECT_LEHR showed a constant signal strength from 1.60 cycles/cm onward. Table 4 shows the results of the SPECT spatial resolution evaluation.   (Table 5). With 123 I, the slope was the highest in A-SPECT_LMEGP and lowest in 99m Tc_SC7 of C-SPECT_WEHR. The scatter fraction was lowest in C-SPECT_MEHRS and highest in C-SPECT_WEHR (Table 5). Furthermore, 123 I_SC3 of C-SPECT_-WEHR showed the highest value. The coefficient of determination of the regression line was close to 1.0 under all conditions (Table 5).

3.B | SPECT acquisition
For the evaluation of high-contrast resolution, Fig. 7 shows the Hot-Rod axial image, and Fig. 8 shows the Pr(n). Figure

| DISCUSSION
In this study, the performance of the C-SPECT system was com- which the threshold of the detector excitation is low and more signal energy can be acquired and (b) the collimator characteristic, in which one hole in the WEHR is geometrically equivalent to one of the pixel detectors. As such, it is possible to capture γ-rays with high accuracy up to the edge of the detector ( Fig. 1; Table 1). Next, in the evaluation using 123 I, the energy resolution of C-SPECT_MEHRS improved by 1.67 times that of C-SPECT_WEHR and by 1.90 times that of A-SPECT_LMEGP. This can be understood from the energy distribution in Fig. 2. First, the difference in the energy spectrum between the high-energy region and the Compton region in WEHR and MEHRS in C-SPECT (Fig. 2) is likely a result of the influence of 529 keV γ- The system spatial resolution of the planar acquisition (Table 2) with 99m Tc was improved with a smaller pixel size in A-SPECT. However, it did not change with the pixel size of 2.46 mm or less in C-SPECT, regardless of EW. This is likely due to the fact that when the element size is 2.46 mm (native pixel) or smaller, a blur component is included to construct an image by interpolation processing in each element, and the spatial resolution deteriorates compared with A-SPECT. Furthermore, images with pixel dimensions different from the native element dimensions must necessarily be interpolated.
With 123 I, C-SPECT and A-SPECT showed a similar tendency in relation to the pixel size. Furthermore, C-SPECT_WEHR had a higher resolution than C-SPECT_MEHRS. This is because WEHR is designed to have a larger hole length and a smaller hole diameter than MEHRS, resulting in improved spatial resolution.
Next, the evaluation of system sensitivity (Table 3) with 99m Tc showed that A-SPECT_LEHR had a higher sensitivity than C-SPECT_WEHR. The reason for this is that the LEHR is designed to have a smaller hole diameter than the WEHR for improved resolution, but to maintain acceptable sensitivity, the septa are designed to be thinner and the hole length is designed to be approximately 46% lower than WEHR (Table 1). For 123 I, C-SPECT_WEHR exhibited a higher sensitivity than C-SPECT_MEHRS, which likely occurs because of the inclusion of more scatter from 529 keV septal penetrating photons in the 159 keV energy window [ Fig. 2(d)]. Furthermore, the WEHR square holes are better matched to the detector element dimension, which probably improves the sensitivity relative to the MEHRS hexagonal holes. However, in C-SPECT, there is an increase in the number of scattering components due to hole tailing and high-energy γ-rays. WEHR is also greatly affected by this factor because of its structure. However, MEHRS can reduce this impact.
As such, it is likely that the sensitivity improved because WEHR contained many scattering components, as Ct.   to an improved radial spatial resolution.
Next, with regard to the linearity of the radioactivity concentration, the slope of the regression line in Fig. 6 and Table 5  Therefore, removal of more scattering components than MEHRS through SC using the TEW method likely led to a low Ct and decreased the sensitivity of C-SPECT_WEHR. Furthermore, by setting the SW for TEW scattering correction to 3% of the main peak, increase in sensitivity in C-SPECT_WEHR is larger than that in C-SPECT_MEHRS. Therefore, in C-SPECT_WEHR, the scattered ray content greatly changes depending on the setting of the SW. were mainly low-frequency. This is because the high-frequency components beyond the fundamental frequency were removed. Therefore, with 99m Tc, the fact that C-SPECT_WEHR showed a higher signal strength than A-SPECT_LEHR in the high-frequency region after 0.60 cycle/cm was not due to the noise component but rather to the edge component of contour formation. With 123 I, in C-SPECT_MEHRS, the signal strength was high in the high-frequency range and was close to that of C-SPECT_WEHR. This trend is the same as that of C-SPECT_WEHR and shows the possibility to profile and separate the 8-mm rod more than A-SPECT. Furthermore, the tendency of 123 I_SC3 of each collimator was close to that of 123 I_SC7, indicating that the contrast resolution was retained.
In the homogeneity evaluation ( Fig. 9), 99m Tc was similar under each condition but slightly improved in 99m Tc_SC3 of C-SPECT_-WEHR15. This finding does not contradict the sensitivity results in Fig. 6(a). With 123 I, the findings do not contradict the sensitivity results shown in Fig. 6(b). The homogeneity was improved in 123 I_SC3. The reason that the homogeneity was improved by the setting of 123 I_SC3 is that C-SPECT_MEHRS was able to acquire the primary photons with high accuracy rather than the increased scattered radiation (Table 5b). At the same time, C-SPECT_WEHR and A-SPECT_LMEGP showed improved homogeneity due to increased scattered radiation. It can be inferred from these results that C-SPECT_MEHRS is suitable for 123 I imaging because of improved homogeneity due to excellent scattered ray removal. ITO ET AL.

| 175
Our results further suggest that it is necessary to pay attention when conducting a direct comparison of each radionuclide because in clinical applications, the sealed concentrations differ for each radionuclide. Additionally, attention must be paid to the comparison of collimator design because there is a difference in the C-SPECT and A-SPECT system configurations. However, C-SPECT showed superior energy resolution in each radionuclide when compared with A-SPECT, leading to an improved contrast resolution. In addition, Peng et al reported a method for improving the crosstalk in the TEW method in SC with a CZT detector-equipped device for the heart. 26 In our results, the sensitivity and homogeneity were affected by the difference in the SC TEW method settings. However, by setting the SW for TEW scattering correction to 3% of the main peak in C-SPECT, the contrast resolution is superior to that of A-SPECT, imaging with similar homogeneity is possible, and improvement in quantification can be expected.
This study has some limitations. First, the availability of the evaluated radionuclides was limited. For MEHRS, in particular, the applicability of medium-energy radionuclides 111 In and 177 Lu could be considered. However, given the characteristics of each radionuclide, the evaluation parameters will likely differ from those used in this study. Verification will be necessary with each imaging technique.
Second, in this study, statistical noise was reduced in terms of system performance evaluation, but there is still a large influence of statistical noise in clinical practice. Therefore, it is necessary to verify each imaging process by including the noise components. Another limitation is related to the correction methods. As SPECT and CT are being used globally, attenuation correction of γ-rays based on CT attenuation maps is considered useful. 27,28 SPECT images with corrections that improve positional resolution and reduce deterioration caused by collimator opening width and source distance combined with order subset expectation maximization are also clinically useful. 29,30 When these corrections are combined, the target noise components change, and setting of the optimal condition based on an understanding of the noise for each correction is required. Furthermore, it is necessary to verify the scatter correction factor in each imaging to apply a more accurate scatter correction.

| CONCLUSION
We verified the system performance of C-SPECT using 99m Tc and 123 I. Each C-SPECT system had superior energy resolution when compared with the A-SPECT system and showed improved contrast resolution with planar and SPECT acquisition. In addition, with the newly developed MEHRS collimator, this study showed excellent scattered ray removal and high-energy resolution on imaging with 123 I formulation, demonstrating its clinical applicability. Furthermore, the verification results show that proper setting of the SW for TEW scatter correction improves the sensitivity and uniformity while maintaining contrast resolution. Clinical versatility can be enhanced by verifying each imaging technique in future studies.

ACKNOWLEDG MENTS
We thank the staff at the Department of Radiology, Saiseikai Yokohamashi Tobu Hospital, and Hideyasu Hosono (GE Healthcare Japan, Tokyo, Japan), for providing technical support.

CONFLI CT OF INTEREST
The authors have no relevant conflicts of interest to disclose.