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. Author manuscript; available in PMC: 2022 Apr 1.
Published in final edited form as: Radiology. 2021 Dec 14;303(1):130–138. doi: 10.1148/radiol.212579

First clinical photon-counting-detector CT system: Technical evaluation

Kishore Rajendran 1, Martin Petersilka 2, André Henning 3, Elisabeth R Shanblatt 4, Bernhard Schmidt 5, Thomas G Flohr 6, Andrea Ferrero 7, Francis Baffour 8, Felix E Diehn 9, Lifeng Yu 10, Prabhakar Rajiah 11, Joel G Fletcher 12, Shuai Leng 13, Cynthia H McCollough 14
PMCID: PMC8940675  NIHMSID: NIHMS1762704  PMID: 34904876

Abstract

Background:

The first clinical CT system to use photon-counting-detector (PCD) technology has become available for patient care.

Purpose:

To assess the technical performance of the PCD-CT system using phantoms and representative participant exams.

Materials and Methods:

Institutional review board approval and written informed consent from four participants were obtained. Technical performance of a dual-source PCD-CT system was measured for standard and high resolution (HR) collimations. Noise power spectrum (NPS), modulation transfer function (MTF), section sensitivity profile (SSP), iodine CT number accuracy in virtual monoenergetic images (VMI), and iodine concentration accuracy were measured. Four participants were enrolled (between May 2021 and August 2021) in this prospective study and scanned using similar or lower radiation doses compared to same-day exams performed using energy-integrating-detector (EID) CT.

Results:

All standard technical performance measures met accreditation requirements. Relative to filtered-back-projection reconstructions, images from iterative reconstruction had lower noise magnitude but preserved NPS shape and peak-frequency. Maximum in-plane spatial resolutions of 125 and 208 microns were measured for PCD-HR and PCD-standard scans, respectively. Minimum values for SSP full-width-half-maximum measurements were 0.34-mm (0.2 mm nominal section thickness) and 0.64 mm (0.4-mm nominal section thickness) for PCD-HR and PCD-standard scans, respectively. In a PCD-CT 120-kV standard scan of a 40-cm phantom, VMI iodine CT numbers had a mean percent error of 5.7% and iodine concentration had root-mean-squared-error of 0.5 mg/cc, comparable to previously reported values for EID-CT. VMI, iodine map, and virtual non-contrast images were created for a coronary CT angiogram acquired with 66-ms temporal resolution. Participant PCD-CT images showed up to 47% lower noise and/or improved spatial resolution compared to EID-CT.

Conclusions:

Technical performance of a new clinical photon-counting-detector CT is improved relative to current state-of-the-art CT system. The dual-source photon-counting-detector geometry facilitated 66-ms-temporal-resolution multi-energy cardiac imaging. Study-participant images illustrated the impact of the improved technical performance.

1. Introduction

Photon-counting-detector (PCD) CT systems have demonstrated suppression of electronic noise (1); simultaneous single-tube-potential, multi-energy imaging (2–4); high spatial resolution without loss of dose efficiency (5–7); radiation dose reduction compared to energy-integrating-detector (EID) CT for the same image quality (7, 8); and improved iodine contrast-to-noise ratio (9). PCDs directly convert x-rays to electrical signal, facilitating small detector pixel designs without loss of geometric dose efficiency and allowing simultaneous multi-energy CT through the use of multiple energy thresholds. Advances in the detector technology have facilitated translation of PCDs to human imaging at clinical doses and dose rates (10, 11). Investigational PCD-CT systems with 0.15–0.27 mm detector pixel sizes (isocenter) and 2–8 energy thresholds have been reported, with human imaging studies demonstrating potential clinical benefits (8, 12–14).

With the introduction of the first clinical PCD-CT, the beneficial characteristics of PCDs can be realized in clinical practice. However, a comprehensive performance assessment is needed to confirm current clinical requirements are met and to provide performance benchmarks relative to EID-CT. The purpose of this study was to assess the technical performance of a clinical PCD-CT system using phantoms and representative images of study participants.

2. Methods and Materials

Siemens Healthineers provided support for this study under a sponsored research agreement with Mayo Clinic. Authors employed by Mayo Clinic who did not receive financial support from Siemens maintained control of all data and information presented in this article.

Institutional review board approval and written informed consent was obtained for study participants.

2.1. Photon-counting-detector CT system

The first clinical PCD-CT system (NAEOTOM Alpha, Siemens Healthineers, Germany) uses a dual-source geometry and 0.25s gantry rotation time to provide 66-ms temporal resolution (isocenter). The primary PCD array has a 50-cm field-of-view (FOV) and the secondary PCD array, used only for cardiac or high helical pitch (up to 3.2) scanning, has a 36-cm FOV (Supplemental Materials, Figure S1).

The most current EID dual-source CT system (SOMATOM Force, Siemens Healthineers) operates its two source-detector pairs at the same tube potential to achieve 66-ms temporal resolution and high helical pitch scanning, or at two different tube-potential/tin-filter combinations for dual-energy acquisition at 125-ms temporal resolution. The energy-resolving detectors of the PCD-CT acquire multi-energy data using a single PCD array and single tube potential. In dual-source mode, multi-energy data are acquired at 66-ms temporal resolution over a 36-cm FOV. Detector pixel and energy-threshold configurations are described in Supplemental Materials (A1, Figure S1).

2.2. Quantitative evaluation of image quality in phantoms

Standard and advanced image quality performance metrics were assessed for PCD-CT using phantoms (Table 1) and compared to published results for EID-CT (15, 16).

Table 1.

Experimental details for image quality performance assessment

ACR CT accreditation phantom Tungsten wire (25 µm diameter) Gold foil (25 µm thickness) Water phantom (20 cm diameter) Iodine phantom (40 cm diameter)
Acquisition mode1 Standard Standard and HR Standard and HR Standard and HR Standard
Protocol type Body Body Head Head Body
Tube potential (kV) 120 120 120 120 120
Tube-current-time product (mAs) 150 400 400 207 (standard), 204 (HR) 362
Rotation time (s) 0.5 0.5 0.5 0.5 0.5
Helical pitch 0.6 0.6 0.6 0.6 0.6
Reconstruction technique QIR QIR QIR WFBP, QIR QIR
Kernel Br44 Br76 (standard), Br96 (HR) Hr76 (standard), Hr84 (HR) Hr40 Qr40 (3)
CTDIvol (mGy) 11.8 31.6 (standard), 32 (HR) 50.7 (standard) 34.3 (HR) 35 mGy (standard and HR) 28.9
QIR strength 3 3 3 1, 3 3.0 / 3.0
Section thickness/interval (mm) 3.0 / 3.0 0.4 / 0.3 0.4 / 0.2 (standard), 0.2 / 0.1 (HR) 3.0 / 3.0 20
Field of view (cm) 21 5 5 21 512×512
Image plane Axial Axial Axial Axial Axial
Matrix size 512×512 1024×1024 1024×1024 512×512 Standard (144 × 0.4 mm)
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HR: High Resolution; CTDIvol: Volume CT dose index; WFBP: Weighted Filtered Back Projection; QIR: Quantum Iterative Reconstruction; Br: Body regular; Hr: Head regular; Qr: Quantitative regular.

1

Standard mode collimation: 144 × 0.4 mm; HR (high resolution) collimation:120 × 0.2 mm

Standard image quality evaluation

CT number accuracy and uniformity, section thickness, low-contrast and high-contrast (spatial) resolution were evaluated using the American College of Radiology (ACR) CT accreditation phantom (CT ACR-464, Sun Nuclear Corporation, USA) and measurement results compared to published limits (17).

Advanced measures of spatial resolution

Using standard and high-resolution (HR) collimations, a tungsten wire was scanned, and images reconstructed with the sharpest available kernels. The modulation transfer function (MTF), which characterizes in-plane spatial resolution as a function of spatial frequency, was calculated using previously validated methods (18).

For standard and HR collimations, a gold foil was scanned, and images reconstructed with the smallest section thicknesses with 50% overlapping increment. Section sensitivity profiles (SSP) were plotted and full-width-half-maximum values calculated to quantify longitudinal spatial resolution (18).

Image noise and noise power spectra

Image noise (standard deviation of pixel values in a 5-cm-diameter region-of-interest, averaged across 15 images) and noise power spectra (NPS) were measured in a water phantom for standard and HR collimations using weighted-filtered-back-projection (WFBP) and quantum-iterative-reconstruction (QIR) images (19). Low-energy threshold images (referred to as T3D by the manufacturer) were reconstructed, which include photon-energies from 20-keV to 120-keV. The shape and peak spatial frequencies of the WFBP and QIR NPS were compared to quantify any noise texture differences.

Quantitative multi-energy CT performance

A 40-cm water phantom containing solid iodine inserts (Gammex, Sun Nuclear Corporation, USA) of concentrations 2, 5, 10 and 15-mg/cc was scanned. Virtual monoenergetic images (VMI) from 40-keV to 140-keV were synthesized and iodine VMI CT numbers were compared against manufacturer reference values. Additionally, iodine maps were generated, and root-mean-squared-error of the estimated iodine concentrations calculated. Virtual non-contrast (VNC) images were generated, and mean CT numbers measured.

2.3. Proof-of-principle study involving participants

Participants undergoing clinically indicated EID-CT were prospectively enrolled. Four clinical indications (CT angiography of the coronary arteries and abdomen/pelvis; whole-body low-dose CT for skeletal surveillance; and temporal bone imaging) were selected to illustrate in humans the effects of PCD-CT’s improved technical performance in phantoms. PCD-CT scans were performed at the same or lower radiation dose relative to the clinical exam (Table 2).

Table 2:

Protocol parameters used for each participant’s clinical EID-CT and research PCD-CT exams

Modality1 Coronary CT angiography Abdomen-pelvis CT angiography Low-dose CT bone survey Temporal bone
PCD-CT EID-CT PCD-CT EID-CT PCD-CT EID-CT PCD-CT EID-CT
Scanner2 NAEOTOM Alpha SOMATOM Force NAEOTOM Alpha SOMATOM Edge+ NAEOTOM Alpha SOMATOM Edge+ NAEOTOM Alpha SOMATOM Force
Collimation 144 × 0.4 mm 192 × 0.6 mm 120 × 0.2 mm 128 × 0.6 mm 120 × 0.2 mm 128 × 0.6 mm 120 × 0.2 mm 64 × 0.6 mm
Scan mode(s)3 DS, ME, FL, ECG-P DS, SE, ECG-R HR SE HR SE HR UHR, SE
Tube potential (kV) 120 90 120 100 120 120 120 120
Tube-current-time-product (mAs) 109 224 123 267 52 63 190 337
CTDIvol (mGy)4 7.6 31.8 9.9 10.6 4.2 4.2 32.6 46.4
Rotation time (s) 0.25 0.25 0.25 0.5 0.25 0.5 1 0.25
Helical Pitch 3.2 0.16 0.8 0.6 0.85 1 0.5 0.35
Contrast media5 / saline flush volume (mL) 90 / 45 110 / 40 100 / 30 100 / 30 - (non-contrast) - (non-contrast) - (non-contrast) - (non-contrast)
Reconstruction type6 QIR ADMIRE QIR ADMIRE QIR ADMIRE QIR ADMIRE
Reconstruction kernel7 Bv48, Qr40 Bv40 Br44, Br72 Br43 Br64 Br62 Hr76, Hr84 Ur77, Ur85
IR strength 4 4 2,4 2 3 3 3 3
Section thickness / interval (mm) 0.6 / 0.4 0.6 / 0.4 1.0 / 0.8 2.0 / 1.2 2.0 / 1.0, 1.0 / 0.5 2.0 / 1.0 0.2 / 0.1 0.4 / 0.3
Field of view (cm) 20 20 42 42 50 50 16 16
Image plane Axial, oblique coronal Axial, oblique coronal Axial Axial Axial Axial Oblique axial Oblique axial
Matrix size 1024×1024 512×512 1024×1024 512×512 512×512, 1024×1024 512×512 1024×1024 1024×1024
Image type(s) created8 VMI, VNC, iodine map 90 kV SE T3D 100 kV SE T3D 120 kV SE T3D 120 kV UHR
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1

PCD-CT: Photon-counting detector CT; EID-CT: Energy-integrating detector CT

2

All scanners manufactured by Siemens Healthineers GmbH

3

DS: Dual source; ME: Multi-energy; FL: Flash (3.2 pitch); ECG-P: Prospective ECG triggering; SE: Single energy; ECG-R: Retrospective ECG gating; HR: High Resolution (does not use a highly-attenuating, post-patient comb or grid filter); UHR: Ultra-High Resolution (uses a highly-attenuating, post-patient comb filter to reduce effective detector pixel size at the expense of radiation dose efficiency)

4

Volume CT Dose Index; based on 32-cm diameter phantom for all protocols involving torso imaging; based on 16-cm diameter phantom for temporal bone (head) imaging

5

Iodinated contrast media (Omnipaque 350, GE Healthcare, WI, USA)

6

QIR: Quantum Iterative Reconstruction; ADMIRE: ADvanced Model-based Iterative REconstruction

7

Kernel description codes: Bv: Body vascular; Br: Body regular; Hr: Head regular; Ur: Ultra-high-resolution (uses comb filter) regular; Higher numerical values indicate increasing sharpness; Routine kernels used in clinical practice selected for EID-CT; Some additional kernels used for the purpose of illustrating specific benefits of photon-counting detector (PCD) CT.

8

T3D: Manufacturer’s name for images created using photon energies from the low-energy threshold (20 keV for all PCD scans) to the maximum possible energy (120 keV for all PCD scans); VMI: Virtual monoenergetic image; VNC: Virtual non-contrast image

3. Results

3.1. Quantitative evaluation of image quality in phantoms

Standard image quality evaluation

Standard image-quality measures conformed with limits set by the ACR CT accreditation program and the state of Minnesota. Measurements and images are shown in Supplemental Materials (Figure S2, Table S2).

Advanced measures of spatial resolution

MTF curves showed cutoff frequencies of 40 line-pairs-per-cm (lp/cm) for body-regular (Br) 96 kernel (125-micron limiting resolution) for HR-PCD and 24 lp/cm for Br76 kernel (208-micron limiting resolution) for standard-PCD, compared to 18 lp/cm (Br69, 278 micron limiting resolution) for EID-CT (16). SSP-full-width-half-maximum values for the smallest available section thicknesses were 0.34-mm (0.2-mm nominal), 0.64 mm (0.4-mm nominal) and 0.72 mm (0.6-mm nominal) for HR-PCD, standard-PCD, and EID (16) scans, respectively (Figure 1).

Figure 1:

Figure 1:

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Modulation transfer functions (left) and section sensitivity profiles (right) for the standard (STD) and high-resolution (HR) collimations of the photon-counting-detector CT system. A limiting in-plane spatial resolution of 125 microns (µm) and a full-width-half-maximum value (section sensitivity profile) of 0.34 mm were achieved using the HR collimation.

Image noise and noise power spectra.

HR images showed 9.51% (4.85 HU, WFBP) lower image noise than standard images (5.36 HU, WFBP) (Figure 2, A). QIR images with strengths 1 and 3 had 11% and 34% lower image noise than WFBP, respectively, for both collimations (see Table T3, Supplemental Materials). NPS curves showed 22% and 58% lower values at peak frequency in QIR-strength-1 and QIR-strength-3 respectively, relative to WFBP in both collimations, and shape and peak frequency were unchanged from WFBP (Figure 2, B, Figure 2, C).

Figure 2:

Figure 2:

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Image noise was reduced 11.5% and 34.5% for quantum iterative reconstruction (QIR) images using strengths 1 and 3, respectively (labelled QIR1 and QIR3), relative to weighted-filtered-back-projection (WFBP) as shown in A. Horizontal dotted lines assist the reader in comparing noise levels between the collimation settings. Noise power spectra for standard (B) and high resolution (HR) (C) collimations demonstrate QIR reduced the noise magnitude (height of the curve) compared to WFBP images for both collimations and strengths without shifting the spatial frequency of the peak of the noise power spectrum (NPS) curves, indicating that only the amount of noise, and not the noise texture, is impacted with use of QIR.

Quantitative multi-energy CT performance

Measured iodine VMI CT numbers for PCD-CT matched expected values (Figure S3, Supplemental Materials), with mean percentage error of 5.7% compared to 5% on dual-source EID-CT (15). Root-mean-squared-error for iodine concentration estimates was 0.5 mg/cc, compared to 0.1–0.4 mg/cc root-mean-squared-error on EID-CT (15). Mean CT number in VNC image corresponding to iodine inserts was −7 HU.

3.2. Proof-of-principle study involving participants

Four participants were scanned using PCD-CT after their routine EID-CT exam at Mayo Clinic between May-2021 and August-2021. These images provide a visual connection between phantom measurements and the appearance of participant images.

PCD coronary CT angiography of a 71-year-old man (Figure 3) demonstrated multi-energy, 66-ms temporal-resolution cardiac imaging. 45 and 55 keV VMIs showed higher iodine signal (45-keV: 1164 HU, 55-keV: 800 HU) compared to an EID-CT at 90-kV (724 HU) that used 22% more iodine contrast (PCD-CT: 90 mL versus EID-CT: 110 mL). PCD-CT iodine maps showed clear delineation of the left coronary artery without motion blur. VMI, iodine maps, and VNC images could not be created from the 66-ms temporal resolution EID-CT data because of the lack of multi-energy information.

Figure 3:

Figure 3:

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Cardiac energy-integrating-detector (EID)-CT (A) and photon-counting-detector (PCD)-CT images (B, C) of a 71-year-old man performed with dual-source CT to achieve 66-ms temporal resolution. While the EID-CT exam is limited to single energy data (A) at this temporal resolution, the multi-energy capabilities of the PCD-CT system allows creation of low energy (45, 55 keV) virtual mono-energetic images (VMIs) (B), which show increased iodine signal compared to EID-CT despite an 18% decrease (A: 110 mL versus B/C: 90 mL) in administered contrast volume (mean CT numbers for the regions of interest shown in black are given in the left column of images). The use of VMIs adds to the inherently higher iodine contrast-to-noise ratio possible with PCD-CT, providing clearer delineation of a branch of the left coronary artery (white arrowheads, right column). Increasing the VMI energy (65 keV or higher) decreases calcium blooming relative to EID-CT (white arrow, right column, A and B). Absolute iodine concentration can be measured using the iodine map images and the virtual-non-contrast images can be used to visualize calcifications having similar attenuation to the iodinated blood (white arrow, right column, C). Reconstruction kernels used: Bv40 (EID-CT 90-kV, A), Bv48 (PCD-CT VMIs, B), Qr40 (PCD-CT iodine map and VNC, C). Display window/level: 2000/200 HU for EID-CT and PCD-CT VMIs, 30/15 mg.cc-1 for iodine map and 1000/100 HU for VNC image.

For abdomen-pelvis CT angiographic images of an 85-year-old man (Figure 4), PCD-CT HR images reconstructed with Br44 (1.0-mm thickness) showed comparable image noise to EID-CT images reconstructed with a very similar Br43 kernel (2.0-mm thickness). PCD-CT allowed sharper reconstructions (Br72 kernel) than EID-CT due to the inherently smaller detector elements. Containing higher spatial frequencies (0% MTF=17 lp/cm), PCD-CT-Br72 image noise was higher (125 HU) than that of PCD-CT-Br44 (31 HU, 0% MTF=9.6 lp/cm).

Figure 4.

Figure 4

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Energy-integrating-detector CT (EID-CT) 100 kV images (left column) and high-resolution photon-counting-detector (PCD) CT T3D images (includes photon energies from 20 to 120 keV) (center and right columns) from an abdomen-pelvis CT angiogram of a human subject with aortobiiliac endovascular stent graft. Both scans used the same radiation dose (10 mGy). The top row shows an enlarged view of the rectangular region of interest displayed in the bottom row. The PCD-CT image in the right column was created using a sharper resolution (Br72 kernel) than is possible to achieve on EID-CT to reduce stent blooming and improve delineation of the struts at the expense of increased image noise. Display window/level = 1800/440 HU.

A low-dose CT skeletal survey for multiple myeloma in a 74-year-old man showed a lytic lesion in the vertebral body (Figure 5). For the same radiation dose and section thickness, PCD-CT images showed 47% lower noise than EID-CT (36.8 HU versus 69.4 HU). Increasing matrix size to 10242 and reducing section thickness to 1-mm yielded similar image noise between PCD-CT and EID-CT (69 HU versus 70 HU), but further improved lesion conspicuity.

Figure 5:

Figure 5:

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Whole-body low-dose CT skeletal survey performed on a human subject with multiple myeloma. Energy-integrating-detector CT (EID-CT) scan (left) and high-resolution photon-counting-detector CT (PCD-CT) scan (center and right) were acquired using the same radiation dose (4.2 mGy). PCD-CT images (center) using the same section thickness (2 mm) as EID-CT images (left) demonstrated 47% lower noise (69.4 HU vs. 36.8 HU). Use of thinner (1 mm) section thickness (right) and 1024×1024 matrix improved resulted in image noise comparable to EID-CT and provided improved delineation of a vertebral lesion (arrow). Zoomed images are displayed to show the lytic lesion. Display window/level = 1500/150 HU.

Right temporal bone images of a 72-year-old man reconstructed using similar kernels (EID-CT/Ur77, PCD-CT/Hr76) showed improved stapes visualization for PCD-CT (0.2-mm thickness) compared to EID-CT (0.4-mm thickness) (Figure 6.) At sharper kernels, EID-CT/Ur85 showed 124% higher noise than EID-CT/Ur77 (83.5 HU versus 187.2 HU), impeding visualization of the stapes and surrounding structures. PCD-CT/Hr84 images showed 58% higher image noise versus PCD-CT/Hr76 (144.8 HU versus 91.7 HU) that did not substantially deteriorate stapes visualization.

Figure 6:

Figure 6:

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Right temporal bone images of a human subject scanned on energy-integrating detector CT (EID-CT) (Ur77 and Ur85 kernels, 0.4-mm section thickness) and photon-counting detector (PCD) CT (Hr76 and Hr84 kernels, 0.2-mm section thickness). PCD-CT show better trabecular detail and delineation of the stapes compared to EID-CT (images A/B versus C/D), despite use of 30% (46.4 versus 32.6 mGy) lower dose for PCD-CT. Display window/level = 3800/700 HU.

4. Discussion

The evaluated photon-counting detector CT system demonstrated 125-µm limiting in-plane spatial resolution, the smallest available for a clinical CT system. Iodine CT numbers in virtual-monoenergetic images matched expected values (5.7% mean percent error), and a root-mean-squared-error of 0.5 mg/cc was measured for iodine concentrations. Participant PCD-CT images illustrated improved spatial resolution and the potential for lower radiation dose (30% in a temporal bone exam with 50% reduction of section thickness) and/or lower image noise (47% lower in a skeletal survey exam with matched radiation dose) compared to a similarly configured EID-CT scanner.

Combining dual-source geometry and 0.25s rotation time with PCDs enabled multi-energy data acquisition at 66-ms temporal resolution (isocenter). Multi-energy image types, including VMI, VNC, and iodine maps, were therefore available without sacrificing temporal resolution (with dual-source EID-CT, dual-energy acquisitions require each source/detector-pair to be operated at a different tube potential, limiting the temporal resolution to 125-ms). Low-keV VMIs, which enable lower contrast doses for patients with renal insufficiency (20), improve iodine signal in data acquired with suboptimal timing (21), and improve visualization of small low-contrast structures (20), can thus be obtained at the best temporal resolution reported to date for cardiac CT. Dual-source PCD-CT also allowed multi-energy imaging using helical pitch values up to 3.2, which can reduce patient dose by up to a factor of two (22).

PCD-CT HR collimation can be used to generate thinner image sections with image noise comparable to thicker EID-CT sections. Alternatively, sharper kernels can be used to improve visualization of fine detail, including stents. For high-contrast applications (e.g. temporal bone), PCD-specific kernels and 0.2-mm section thickness improve spatial resolution without substantially increasing image noise.

For screening or other exams performed at reduced doses relative to routine applications, PCD-HR collimation can be leveraged to reduce image noise at a prescribed spatial resolution (7, 8, 23). Alternatively, the 0.2 mm detector pixels can be used to improve anatomic conspicuity by using reduced section thicknesses and/or larger image matrices to reduce partial-volume effects without increasing image noise relative to EID-CT. Low-dose CT exams also benefit from PCDs’ intrinsic ability to eliminate electronic noise (1).

A clinical EID-CT system with 0.25×0.25-mm2 detector-pixel (isocenter), the smallest among current EID-CT systems, is reported to have a 10%-MTF of 19.6-lp/cm using the sharpest reconstruction kernel assessed (24), compared to a 10%-MTF value of 36.1-lp/cm on the evaluated PCD-CT. The minimum SSP-FWHM of this EID-CT system is reported to be 0.45-mm (0.25-mm nominal section thickness) (25), compared to 0.34-mm (0.20-mm nominal thickness) on the evaluated PCD-CT. PCD-CT also allows simultaneous single-tube-potential multi-energy imaging, count weighting rather than energy weighting of detected photons resulting in improved contrast-to-noise ratio, and higher geometric dose-efficiency by avoiding the need for inter-detector-pixel septa.

Our study had the following limitations. PCD-CT data were acquired using only the manufacturer-recommended tube potential. Studies at additional tube potentials are needed to determine optimal settings for different clinical tasks. Since our purpose was to assess technical performance of the PCD-CT system using phantoms, a reader study using patient images was not performed. Instead, representative PCD-CT images of participants are shown to provide a visual correlation to phantom results. Reader studies to assess diagnostic performance, reader confidence, and clinical impact for different diagnostic tasks and clinical cohorts are necessary.

In conclusion, the first clinical photon-counting-detector CT system demonstrated superior spatial resolution, improved noise properties, and improved multi-energy temporal resolution relative to similarly configured, energy-integrating-detector CT. Visual correlation of phantom results was shown in humans for four clinical applications.

Supplementary Material

1

Table S1: Additional technical specifications for the evaluated PCD-CT system

Table S2: American College of Radiology (ACR) CT accreditation phantom measurements

Table S3: CT number and image noise measurements in a 20-cm water phantom

Figure S1: (A) Schematic of the dual-source photon-counting-detector (PCD) CT system. The A and B detector arrays (labeled Quantum detector A and B) both use PCDs and have 50-cm and 36-cm scanning field of views, respectively. (B) Detector pixel layout shows 0.150 × 0.176 mm2 subpixels (measured at isocenter). Due to the presence of collimator blades for every 6 subpixels in the z-direction, the minimum effective section thickness is 0.2 mm for the high resolution (HR) 120×0.2 mm collimation and 0.4 mm for the standard (144×0.4 mm) collimation.

Figure S2: Virtual mono-energetic images at 70 keV of the ACR CT accreditation phantom. Table S2 provides region-of-interest measurements and corresponding limits set by the ACR CT accreditation program. For high contrast (spatial) resolution, up to 8 line-pairs were resolvable (arrow, image number 4) for our routine abdomen scan protocol.

Figure S3: Quantitative accuracy of CT numbers from 40 to 140 keV virtual mono-energetic images for four rods having different concentrations of iodine (2, 5, 10 and 15 mg/cc) for scan acquired in a 40-cm body-size phantom at 120 kV PCD-standard mode. Measured CT number closely matched reference values (mean percent error = 5.7%).

Summary statement:

The first clinical photon-counting-detector CT system demonstrated superior spatial resolution relative to current clinical CT systems and improved noise properties and multi-energy temporal resolution relative to similarly configured, energy-integrating-detector CT.

Key results.

  1. The high-resolution mode demonstrated 125-micron in-plane spatial resolution and 0.3 mm longitudinal resolution, the smallest reported to date for a clinical CT system.

  2. The photon-counting detector CT system provided 66-ms temporal resolution multi-energy imaging when operated in dual source geometry.

  3. Noise reduction (up to 47%) or dose reduction (up to 30%) were illustrated in different study participants using the photon-counting detector CT system relative to a similar CT system equipped with energy-integrating-detector.

Source of Funding

This study was supported by the National Institutes of Health under award numbers R01 EB028590 and supported in kind by Siemens Healthineers GmbH, who own the evaluated system under the terms of a sponsored research agreement with the Mayo Clinic. The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institute of Health. The device described is CE labeled and has received 510k clearance by the U.S. Food and Drug Administration.

Abbreviations

PCD

Photon-counting detector

EID

Energy-integrating detector

HR

High (spatial) resolution

NPS

Noise power spectrum

VMI

Virtual monoenergetic image

VNC

Virtual non-contrast image

MTF

Modulation transfer function

SSP

Section sensitivity profile

QIR

Quantum iterative reconstruction

WFBP

Weighted filtered back projection

Footnotes

Data-sharing statement

Data generated or analyzed during the study are available from the corresponding author by request.

Contributor Information

Kishore Rajendran, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Martin Petersilka, Siemens Healthineers, Forchheim, Germany.

André Henning, Siemens Healthineers, Forchheim, Germany.

Elisabeth R Shanblatt, Siemens Medical Solutions, Malvern, PA, USA.

Bernhard Schmidt, Siemens Healthineers, Forchheim, Germany.

Thomas G. Flohr, Siemens Healthineers, Forchheim, Germany.

Andrea Ferrero, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Francis Baffour, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Felix E. Diehn, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Lifeng Yu, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Prabhakar Rajiah, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Joel G. Fletcher, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Shuai Leng, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

Cynthia H. McCollough, Department of Radiology, Mayo Clinic, Rochester, MN 55905 USA.

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Associated Data

This section collects any data citations, data availability statements, or supplementary materials included in this article.

Supplementary Materials

1

Table S1: Additional technical specifications for the evaluated PCD-CT system

Table S2: American College of Radiology (ACR) CT accreditation phantom measurements

Table S3: CT number and image noise measurements in a 20-cm water phantom

Figure S1: (A) Schematic of the dual-source photon-counting-detector (PCD) CT system. The A and B detector arrays (labeled Quantum detector A and B) both use PCDs and have 50-cm and 36-cm scanning field of views, respectively. (B) Detector pixel layout shows 0.150 × 0.176 mm2 subpixels (measured at isocenter). Due to the presence of collimator blades for every 6 subpixels in the z-direction, the minimum effective section thickness is 0.2 mm for the high resolution (HR) 120×0.2 mm collimation and 0.4 mm for the standard (144×0.4 mm) collimation.

Figure S2: Virtual mono-energetic images at 70 keV of the ACR CT accreditation phantom. Table S2 provides region-of-interest measurements and corresponding limits set by the ACR CT accreditation program. For high contrast (spatial) resolution, up to 8 line-pairs were resolvable (arrow, image number 4) for our routine abdomen scan protocol.

Figure S3: Quantitative accuracy of CT numbers from 40 to 140 keV virtual mono-energetic images for four rods having different concentrations of iodine (2, 5, 10 and 15 mg/cc) for scan acquired in a 40-cm body-size phantom at 120 kV PCD-standard mode. Measured CT number closely matched reference values (mean percent error = 5.7%).

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