Calibrated scintigraphic imaging procedures improve quantitative assessment of the cardiac sympathetic nerve activity

The 123I-labeled meta-iodobenzylguanidine (MIBG) is an analogue of noradrenaline that can evaluate cardiac sympathetic activity in scintigraphy. Quantitative analysis of 123I-MIBG images has been verified in patients with heart failure and neurodegenerative diseases. However, quantitative results differ due to variations in scintigraphic imaging procedures. Here, we created and assessed the clinical feasibility of a calibration method for 123I-MIBG imaging. The characteristics of scintigraphic imaging systems were determined using an acrylic calibration phantom to generate a multicenter phantom imaging database. Calibration factors corresponding to the scintigraphic imaging procedures were calculated from the database and applied to a clinical study. The results of this study showed that the calibrated analysis eliminated inter-institutional differences among normal individuals. In summary, our standardization methodology for 123I-MIBG scintigraphy could provide the basis for improved diagnostic precision and better outcomes for patients.

HMR variation under acquisition conditions. We confirmed the HMR in 123 I-MIBG planar images obtained over periods of 1-10 min. Although mean HMR did not significantly differ among acquisition periods of 1, 2, 3, 4, 5, 7, and 10 min (2.32 ± 0.063, 2.32 ± 0.035, 2.34 ± 0.040, 2.35 ± 0.029, 2.33 ± 0.017, 2.34 ± 0.024, and 2.35 ± 0.010, respectively), the standard deviations (SD) of the HMR gradually decreased over longer acquisition periods (Fig. 2a). The quality of phantom images seemed most stable during acquisition for 5 and 7 min. The HMR were also stable at any gamma camera position from the phantom surface (Fig. 2b). Image quality was degraded in both LE high-resolution (LEHR) and LME general-purpose (LMEGP) collimators when the distance from the phantom surface was increased. The HMR was higher when images were acquired at the energy windows of 159 keV ± 7.5% than at 159 keV ± 10%, indicating that the setting of primary energy window of 123 I affected the HMR (Fig. 2d). The HMR values with 256 and 512 matrices did not significantly differ except under two imaging conditions (Fig. 2e).

Scientific Reports
To confirm the flexibility of conversion coefficients to account for the variation in image acquisitions, we calculated average multicenter conversion coefficients from the phantom image database consisted of 1459 phantom image sets and compared these conversion coefficients and those from the database consisted of 705 image sets in Supplementary Table 1. The average conversion coefficients were statistically equivalent between two databases in CHR, LEHR, LEGP, ELEGP, LMEGP, and MEGP collimators except for MELP collimator. Consequently, we additionally determined the average multicenter conversion coefficients of combinations of gamma cameras with collimators in 1459 phantom image sets (Supplementary Table 2).
Clinical application of the method for standardizing 123 I-MIBG scintigraphy. Normal HMR values of two hospitals were evaluated with or without standardization (Fig. 4a). The HMR for early and delayed 123 I-MIBG scintigraphic images were corrected with institutional and multicenter conversion coefficients. The institutional conversion coefficients were 0.631 and 0.840 for hospitals A and B, respectively. The multicenter conversion coefficients were 0.621 and 0.838 for hospitals A and B, respectively. Although the averaged normal values in hospitals A and B significantly differed before standardization, the standardized normal values did not significantly differ between these hospitals in both early and delayed 123 I-MIBG images (Fig. 4b,c). Furthermore, HMR corrected with institutional and multicenter conversion coefficients did not significantly differ.
To evaluate the accuracy of the calibration factor, we calculated net reclassification improvement 34 (NRI) in normal subjects and patients with heart failure. Of the 12 patients (24 image sets) who were diagnosed with heart failure, classification was improved in four images when the calibration factor was applied to H/M ratio. However, of the 21 normal subjects (42 image sets), classification was only worsened in one image. The NRI in all subjects showed 14.3% (p = 0.099). Uncorrected and corrected HMR values for individual heart failure patients are shown in Supplementary Table 3.

Discussion
The major findings of the present simulation study were that collimator design, including collimator length, hole diameter, and septal thickness, affect 123 I-MIBG image quality and HMR. The phantom studies revealed that the energy-window setting for 123 I is an important factor for reducing HMR variation. However, HMR variations due to acquisition time, matrix size, and distance between gamma camera and phantom surfaces were limited. The conversion coefficients represented the characteristics of gamma cameras and collimators in the multicenter phantom study, and differed among manufacturers of collimators even those with the same name. Moreover, conversion coefficients significantly differed in some combinations of gamma cameras with collimators, even within Table 1. Average multicenter conversion coefficients for combinations of gamma cameras and collimators obtained from 705 image sets. CHR cardiac high-resolution, ELEGP extended low-energy general-purpose, LEGP low-energy general-purpose, LEHR low-energy high-resolution, LMEGP low-medium-energy generalpurpose, MEGP medium-energy general-purpose, MELP ME low-penetration. Monte Carlo simulation provided reasonable 123 I-MIBG phantom planar images even considering the effects of the 529 keV photons added to the 159 keV photons. Although the fractions of 529 and159 keV photons of 123 I were 1.39% and 83.3%, respectively, the high-energy photons hampered quantitative analysis of HMR and degrade 123 I-MIBG planar image quality. Since these 529 keV photons easily penetrated thin collimator septa, a peak appeared in the energy spectrum with the LEHR collimator. Photons that penetrated the septum or scattered, also degraded 123 I-MIBG planar image quality and reduced the HMR. In addition to septal thickness, collimator length and hole diameter are also important components that determine both 123 I-MIBG image quality and HMR. Thick collimator septa, small hole diameters, and long collimators are most appropriate. Considering these effects of 529 keV photons, the MEGP collimator is adequate for 123 I-MIBG imaging.
We previously determined conversion coefficients for several collimator groups in 225 experiments at 84 institutions 26  The present findings showed that although the conversion coefficients were equivalent between the two vendors (Fig. 3b), they were affected by the energy window setting of 123 I (Fig. 2d). We applied a single energy window setting in the present study, whereas windows were set at 159 keV ± 10% and 159 keV ± 7.5% in the previous study. In Supplementary Table 1, when we compared two imaging databases acquired with single and various energy window settings, the average values of conversion coefficients were significantly different in for the MELP collimator condition.
Imaging conditions need standardization in addition to HMR for 123 I-MIBG image acquisition. A tremendous amount of data regarding imaging protocols has been accumulated in the multicenter phantom image database with respect to the imaging matrix, energy window setting of 123 I, and acquisition time. Moreover, 145 image datasets scatter-corrected using 123 I dual- 35,36 and triple-energy 23,37 windows were included in the phantom database. Since the clinical usage of 123 I-MIBG was approved in 1992 in Japan, many studies have investigated the HMR quantitation 23,38-41 , which has led to a wide variety of imaging conditions and correction methods. In addition, 123 I-MIBG phantom experiments conducted in the Netherlands, Belgium, the UK, Austria, and www.nature.com/scientificreports/ Italy 28-30,32,33 have also generated a considerable amount of data. Our phantom-based standardization methodology allows international comparisons of HMR. Our study has several limitations. We used an institutional 123 I-MIBG imaging procedure for the phantom scans. Therefore, the imaging procedure was not unified in the multicenter phantom study. However, eligible phantom image sets were selected according to the selection criteria of the multicenter 123 I-MIBG phantom image database (Fig. 2c). Although we provided multicenter conversion coefficients to standardize HMR, they could only be used at the energy-window setting of 159 keV ± 10%. Since the number of conversion coefficients for the energy-window setting of 159 keV ± 7.5% was limited (Supplementary Table 4), additional multicenter 123 I-MIBG phantom imaging studies are needed to accumulate conversion coefficients for this setting. The clinical validation study confirmed the feasibility of our method only for patients with normal 123 I-MIBG distribution. A multicenter clinical trial should be conducted using institutional and multicenter conversion coefficients.
In conclusion, our standardization methodology for 123 I-MIBG scintigraphy allowed determination of the characteristics of gamma cameras and collimator combinations in the multicenter phantom study. The clinical validation study showed that normal HMR derived from two different institutions did not significantly differ after standardization.

Material and methods
Quantitative analysis in 123 I-MIBG imaging. The HMR was used to calculate cardiac 123 I-MIBG accumulation in planar images as cardiac 123 I-MIBG uptake divided by background of 123 I-MIBG distribution using ROI positioned over the heart and over the upper mediastinum 14 . Fully and semi-automated ROI setting algorithms were applied to the phantom and clinical studies 20 , respectively. The HMR were automatically calculated using both algorithms.

Calibration phantom for planar 123 I-MIBG imaging. A flat, polymethyl methacrylate phantom (Taisei
Medical, Co. Ltd, Osaka, Japan) was developed to calibrate HMR under various imaging conditions with collimators 23,33 (Fig. 1a). The volume (width × depth × height) of this phantom is 380 × 380 × 50 mm 3 , and it can mimic planar 123 I-MIBG distribution in the heart, mediastinum, liver, lungs, and thyroid gland. Anterior and posterior planar 123 I-MIBG images were acquired from both sides of the phantom. The designated HMR of the anterior and posterior views were 2.60 and 3.50, respectively. Details of the phantom design have been published elsewhere 23 . Calibration factor for gamma camera and collimator system. The calibration factor was calculated from the HMR derived from anterior (HMR Ant ) and posterior (HMR Post ) planar 123 I-MIBG phantom images using dedicated software and is defined as a conversion coefficient calculated as: where, 2.60 and 3.50 are the respective designated HMR in anterior and posterior views of the calibration phantom. An institutional conversion coefficient (CC i ) was derived from the anterior and posterior phantom images after image acquisition under institutional 123 I-MIBG planar imaging conditions.

Conversion to standardized HMR using the calibration factor. Since the European Association
Nuclear Medicine and the European Council of Nuclear Cardiology have proposed using MEGP collimators for 123 I-MIBG imaging 16 , all HMR were converted into that for a MEGP collimator. A standardized conversion coefficient (CC std ) has already been defined as 0.88 26 . The CC i and CC std allow for the conversion of all institutional HMR (HMR i ) into standardized HMR (HMR std ) using the equation 26 : Monte Carlo simulation for 123 I-MIBG imaging. A digital phantom image was created from the acrylic calibration phantom image acquired using X-ray CT. Density and source maps of the phantom were generated for the following simulations. The simulation of imaging nuclear detectors (SIMIND; Lund University, Lund, Sweden) Monte Carlo program 42 allowed 123 I-MIBG planar imaging simulations using various types of collimators. Combinations of the following collimator conditions were examined: collimator hole diameters of 1, 2, 3, 4, and 5 mm; septal thicknesses of 0.10, 0.45, 0.80, 1.15, and 1.50 mm, and collimator lengths of 20, 30, 40, 50, and 60 mm. We generated 123 I-MIBG planar images using a total of 8.65 × 10 8 photons. The number of detected photons ranged from 418 to 15,321 per second. Planar MIBG imaging was simulated with 256 × 256 matrices, and the energy window of 123 I was set at 159 keV ± 7.5%.

HMR variations during various acquisition periods.
Anterior MIBG planar images were acquired using a dual-head gamma camera (e.cam; Toshiba Medical Systems, Tokyo, Japan) and an LMEGP collimator over periods of 1, 2, 3, 4, 5, 7, and 10 min from the phantom containing 55.5 MBq of 123 I-MIBG. Five image datasets were acquired during each period. Planar imaging was conducted with a 256 × 256 matrix, and a pixel size of 1.65 mm. A photopeak window of 123 I was centered at 159 keV with a 15% energy window. This study proceeded at Narita Memorial Hospital, Aichi, Japan.
HMR variation according to imaging matrix. The HMR from 256 and 512 matrices were compared in the multicenter phantom image database that comprised 705 image datasets from 309 institutions obtained with an energy window setting of 159 keV ± 10%. The image datasets from GE (n = 351), Picker (n = 80), Siemens (n = 132) and Toshiba (n = 88) were compared.

Conversion coefficient for combinations of gamma cameras and collimators.
Based on the multicenter 123 I-MIBG phantom image database with the image selection criteria, mean conversion coefficients for combinations of gamma cameras and collimators were determined using 705 image sets from 309 institutions. The 9 types of gamma cameras were Discovery/Optima (n = 121), Infinia (n = 151), Millennium MG (n = 33), and Millennium VG (n = 33) manufactured by GE; BrightView (n = 36) by Philips; PRISM (n = 73) by Picker; e.cam/Symbia (n = 110) and EvoExcel/IntevoExcel (n = 12) by Siemens, and e.cam/Symbia (n = 79) by Toshiba. Additional multicenter 123 I-MIBG phantom image datasets for the EvoExcel/IntevoExcel system were accumulated due to the absence of these image datasets in the phantom database. The number of additional image datasets was 27 from 14 institutions.
Clinical validation image dataset. We applied the calibration method of HMR to an anonymized clinical image dataset. The Japanese Society of Nuclear Medicine working group (JSNM-WG) activity collected planar images from patients who were determined as normal cardiac 123 I-MIBG uptake in 2007 and 2015 [43][44][45] . All personal information of 123 I-MIBG images was excluded and anonymized 123 I-MIBG images formatted with digital imaging and communications in medicine were provided as a research database. We obtained the permission for the secondary use of the databases as a research purpose in accordance with JSNM-WG regulation. Details of the patient characteristics have been published elsewhere 43 . In the anonymized clinical image dataset, male and female (n = 8 each) 123 I-MIBG images were collected from hospital A, and eight male and six female images were collected from hospital B (Fig. 4a) www.nature.com/scientificreports/ respectively, and e.cam and LMEGP manufactured by Siemens at hospital B, respectively. The acquisition time, imaging matrix, and energy-window setting for 123 I were 5 min, 256, and 159 keV ± 10%, respectively at both hospitals. Early and delayed planar images were acquired at 15 min and 4 h after injecting 123 I-MIBG in both hospitals, respectively. The institutional conversion coefficients were obtained from 123 I-MIBG phantom scans at each hospital. Mean multicenter conversion coefficients matching the combinations of gamma cameras and collimators at hospitals A and B were calculated from the 123 I-MIBG phantom image database. For the calculation of net reclassification improvement with the standardization procedure in clinical subjects, we used clinical image datasets collected from the hospital A (16 subjects, 32 images for early and delayed conditions) and Kanazawa University Hospital, Kanazawa, Japan (17 subjects, 34 images). Regarding clinical 123 I-MIBG imaging condition in Kanazawa University Hospital, the acquisition time, imaging matrix, and energywindow setting for 123 I were 5 min, 256, and 159 keV ± 10%, respectively. Early and delayed planar images using an LMEGP collimator were acquired at 20 min and 3 h after injecting 123 I-MIBG. Of the 33 patients, 12 patients were diagnosed with heart failure, and 21 subjects were diagnosed with a normal heart. The reclassification table was generated to compare standardized HMR values using multicenter conversion coefficients with uncorrected HMR values. These HMR values were classified into two patient groups using the thresholds of 2.17 and 2.49 that were determined by the receiver operating characteristic analysis in unstandardized and standardized conditions, respectively.

Statistical analysis.
All continuous values are expressed as means ± SD. The Shapiro-Wilk testing for the evaluation of normality was performed in the continuous dataset. Differences in continuous variables were analyzed using Student t-tests and Wilcoxon singed rank tests. Multiple comparisons of continuous variables were assessed using Tukey-Kramer tests. Differences in paired continuous data were analyzed using paired t-tests. All statistical tests were two-tailed, and values with p < 0.05 were considered significant. All data were statistically analyzed using JMP version 11.2.1 (SAS Institute Inc., Cary, NC, USA).

Data availability
The multicenter phantom data that support the findings of this study are available from the corresponding authors (KO and KN) upon reasonable request. Regarding the clinical normal database for 123 I-MIBG scintigraphy, the planar 123 I-MIBG imaging data that support the findings of this study are available from the corresponding author (KN) upon reasonable request.