# Comparison of the spectral performance between two dual-source CT systems on low-energy virtual monoenergetic images: A phantom study

Joël Greffier, Sebastian Faby, Maxime Pastor, Julien Frandon, Julien Erath, Jean-Paul Beregi, Djamel Dabli

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Joël Greffier, Sebastian Faby, Maxime Pastor, Julien Frandon, Julien Erath, et al.. Comparison of the spectral performance between two dual-source CT systems on low-energy virtual monoenergetic images: A phantom study. Physica Medica European Journal of Medical Physics, 2024, 124, pp.103429. 10.1016/j.ejmp.2024.103429. hal-05077957

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Technical Note

# Comparison of the spectral performance between two dual-source CT systems on low-energy virtual monoenergetic images: A phantom study

Joël Greffier<sup>a, *</sup>, Sebastian Faby<sup>b</sup>, Maxime Pastor<sup>a</sup>, Julien Frandon<sup>a</sup>, Julien Erath<sup>b</sup>, Jean-Paul Beregi<sup>a</sup>, Djamel Dabli<sup>a</sup>

<figure><img src="image_3.png" /></figure>

ARTICLE INFO

Keywords:

Dual-energy

Multidetector Computed tomography

Task-based image quality assessment

Iterative reconstruction

ABSTRACT

**Purpose:** To compare the spectral performance of two different DSCT (DSCT-Pulse and DSCT-Force) on virtual monoenergetic images (VMIs) at low energy levels.

**Methods:** An image quality phantom was scanned on the two DSCTs at three dose levels: 11/6/1.8 mGy. Level 3 of an advanced modeled iterative reconstruction algorithm was used. Noise power spectrum and task-based transfer function were computed on VMIs from 40 to 70 keV to assess noise magnitude and noise texture ($f_{av}$) and spatial resolution ($f_{50}$). A detectability index (d') was computed to assess the detection of one contrast-enhanced abdominal lesion as a function of the keV level used.

**Results:** For all dose levels and all energy levels, noise magnitude was significantly higher (p < 0.05) with DSCT-Pulse than with DSCT-Force (12.6 ± 2.7% at 1.8 mGy, 9.1 ± 2.9% at 6 mGy and 4.0 ± 2.7% at 11 mGy). For all energy levels, $f_{av}$ values were significantly higher (p < 0.05) with DSCT-Pulse than with DSCT-Force at 1.8 mGy (4.8 ± 3.9%) and at 6 mGy (5.5 ± 2.5%) but similar at 11 mGy (0.2 ± 3.6%; p = 0.518). For all energy levels, $f_{50}$ values were significantly higher with DSCT-Pulse than with DSCT-Force (12.7 ± 5.6% at 1.8 mGy, 17.9 ± 4.5% at 6 mGy and 13.1 ± 2.6% at 11 mGy). For all keV, similar d' values were found with both DSCT-Force and DSCT-Pulse at 11 mGy (-1.0 ± 3.1%; p = 0.084). For other dose levels, d' values were significantly lower with DSCT-Pulse than with DSCT-Force (9.1 ± 3.2% at 1.8 mGy and -6.3 ± 3.9% at 6 mGy).

**Conclusion:** Compared with the DSCT-Force, the DSCT-Pulse improved noise texture and spatial resolution, but noise magnitude was slightly higher and detectability slightly lower, particularly when the dose level was reduced.

# 1. Introduction

In recent years, numerous technical developments have appeared in routine CT clinical practice, changing the way patients receive radiological care [1-6]. Dual-energy spectral imaging is one of these developments, making it possible to improve the detection and characterization of lesions, particularly abdominal lesions [2,4,7-12]. It involves acquiring or detecting two photon spectra – one at low energy, the other at high energy – to separately assess the contribution of the photoelectric effect and the Compton effect [1,2]. This makes it possible to generate different types of images, including virtual monoenergetic images (VMIs) which enhance the contrast of tissues and lesions

containing injected iodine [1,2].

Various dual-energy CT (DECT) platforms have been developed with different technologies and spectral performances [1,13,14]. Among these DECT platforms, one CT manufacturer has developed a dual-source CT (DSCT) platform [1]. With this platform, two X-ray tube/detector pairs are used to acquire image datasets: one X-ray tube uses a low kV (X-ray tube A) and the other uses a high kV (X-ray tube B) [1,2]. Different generations of DSCT have been developed over time, with various improvements regarding tubes, available X-ray spectra and their spectral separation for DECT imaging, as well as the field of view for the B system. In 2023, a new additional DSCT was developed by the same manufacturer. Compared to the best previously available DSCT, this new

Abbreviations: CT, Computed Tomography; CTDIvol, Volume CT dose index; d', Detectability index; DECT, Dual-energy CT; DSCT, Dual-source CT; HU, Hounsfield Unit; NPS, Noise power spectrum; TTF, Task-based transfer function; VMI, Virtual monoenergetic image.

* Corresponding author at: Joël Greffier, CHU de Nîmes, Nîmes Medical Imaging Group, EA 2992, Bd. Prof Robert Debré, 30029 Nîmes, Cedex 9, France.

E-mail address: joel.greffier@chu-nimes.fr (J. Greffier).

<footnote><sup>a</sup> IMAGINE UR UM 103, Montpellier University, Department of Medical Imaging, Nîmes University Hospital, 30900 Nîmes, France</footnote>

<footnote><sup>b</sup> Department of Computed Tomography, Siemens Healthineers AG, Siemensstr. 3, 91301 Forchheim, Germany</footnote>
DSCT is equipped with different X-ray tubes and different filtration on both X-ray tubes. On the detector side it is using a smaller beam collimation, however, they employ same type of detector technology and iterative reconstruction algorithm.

Numerous clinical and patient studies have evaluated the spectral performance of DSCT, particularly on VMIs for the abdomen [15–19]. However, to our knowledge, no studies have ever evaluated the spectral performance of this new DSCT in comparison the previously available ones.

The aim of this phantom study was to compare the spectral performance of this new dual-source CT with another DSCT on virtual monoennergetic images (VMIs) at low-energy levels. To achieve this goal, a task-based image quality assessment was performed.

## 2. Materials and methods

### 2.1. Phantom

The 31 cm diameter section of the Mercury v4.0 phantom (Sun Nuclear, Melbourne, FL, USA) was used to perform the task-based image quality assessment (Fig. 1). This section was chosen to be as close as possible to the body mass index of patients undergoing abdomen CT examinations (body mass index of 27 kg/m²) [20]. The Noise Power Spectrum (NPS; Fig. 1.A) was computed on the homogeneous part of this section and the Task-based Transfer Function (TTF) on the iodine insert at 10 mg/mL in the part with the five inserts (Fig. 1.B).

### 2.2. CT scanners and scanning protocols

Acquisitions were performed on two Siemens Healthineers Dual-Source CT systems (Forchheim, Germany): the SOMATOM Force (DSCT-Force) and the new SOMATOM Pro.Pulse (DSCT-Pulse).

Acquisitions were performed using the kV pair (X-ray tube A/X-ray tube B with additional tin filter (Sn)) of 100 kV / Sn150 kV for the DSCT-Force system and 100 kV / Sn140 kV for the DSCT-Pulse. The tube current modulation system was disabled in this round phantom and the tube current was set to obtain three volume CT dose indexes (CTDI<sub>vol</sub>) of 11 mGy, 6 mGy and 1.8 mGy. These three dose levels correspond respectively to one of the two DRLs used in France, to the dose level used at our institution for spectral CT of the abdomen and to the Ultra-low dose levels of the abdomen [21]. Other acquisition parameters correspond to those used in clinical practice for dual-energy acquisitions of the abdomen (Table 1). Ten acquisitions were performed on each CT system at each dose level.

Raw data were reconstructed using Level 3 of the ADvanced Modeled

<table><caption>**Table 1**

Technical parameters used with the CT systems studied.</caption><thead><tr><th>CT scan model</th><th>SOMATOM Force</th><th>SOMATOM Pro.Pulse</th></tr></thead><tbody><tr><td>Year of launch</td><td>2015</td><td>2023</td></tr><tr><td>X-Ray Tube</td><td>Vectron®</td><td>Athlon® DS liquid metal bearing</td></tr><tr><td>Inherent filtration</td><td>5.7 mm Al at 150 kV</td><td>5.5 mm Al at 140 kV</td></tr><tr><td>Additional Tin filter thickness</td><td>0.6 mm</td><td>0.4 mm</td></tr><tr><td>Focal spot</td><td>0.8 × 1.1 mm</td><td>1.0 × 1.2 mm</td></tr><tr><td>Used tube voltage combination</td><td>100/Sn150 kV</td><td>100/Sn140 kV</td></tr><tr><td>Detector</td><td>Stellar (UFC)</td><td>Stellar (UFC)</td></tr><tr><td>Beam collimation</td><td>128 x 0.6 mm*</td><td>64 x 0.6 mm</td></tr><tr><td>Pitch</td><td>0.8</td><td>0.8</td></tr><tr><td>Used rotation time (s)</td><td>0.5</td><td>0.5</td></tr><tr><td>Displayed CTDI<sub>vol</sub> (mGy)</td><td>10.96<br>5.96<br>1.74</td><td>11.00<br>5.99<br>1.74</td></tr><tr><td>Tube current (mAs)**</td><td>189/95<br>103/52<br>31/16</td><td>102/77<br>55/43<br>16/13</td></tr></tbody></table>

Footnotes: * Beam collimation using z-flying focal spot; ** mAs are expressed for X-ray tube A (low kVp) and X-ray tube B (high kVp), respectively.

Iterative REconstruction algorithm (ADMIRE) for both CT systems. The Qr36 quantitative soft reconstruction kernel, the slice thickness of 1 mm with 1 mm increment, the 512 x 512 pixels matrix size and a 330 mm field of view were used.

For each acquisition, VMIs were reconstructed with the 4 lowest energy levels (40/50/60/70 keV) usually used in clinical practice to detect or characterize abdominal lesions.

### 2.3. Task-based image quality assessment on VMIs

A task-based image quality assessment was performed using the iQMetrix-CT software developed by the French Society of Medical Physics [22].

#### Noise power spectrum

For each energy level and each CT acquisition, the NPS was computed in 2 consecutive axial slices by placing four square regions of interest (ROIs) of 56 x 56 pixels (Fig. 1.A).

To quantify the changes in noise magnitude, the square root of the area under the NPS1D curve (HU) was measured. To quantify changes in noise texture, the average spatial frequency ($f_{av}$, mm⁻¹) of the NPS1D

<figure><img src="image_4.png" /><figcaption>**Fig. 1.** 31 cm diameter section of the Mercury v4.0 phantom used in this study. Four square regions of interest (ROIs) of 56 × 56 pixels were used for the noise power spectrum (NPS) assessment. One circular ROI was used to compute the task-based transfer function (TTF) on iodine insert at 10 mg/mL.</figcaption></figure>
<figure><img src="image_5.png" /><figcaption>Fig. 2. Noise power spectrum (NPS) curves obtained for the two dual-source CT systems and the three dose levels according to energy levels (keV).</figcaption></figure>

<figure><img src="image_6.png" /><figcaption>Fig. 3. Task-based transfer function (TTF) curves of the iodine insert at 10 mg/mL obtained for the two dual-source CT systems and all dose levels according to energy levels (keV).</figcaption></figure>

* **Contrast between the phantom's background material and iodine insert**

For both CT systems, the contrast between the phantom's background material and the iodine insert at 10 mg/ml decreased as the keV increased (Fig. 4.A). For all energy levels, similar contrast values were found according to dose level on each of the two CTs (p > 0.05).

For all dose levels and all energy levels, contrast values were significantly lower with the DSCT-Pulse system than with DSCT-Force system (2.9 % ± 1.0 %; p < 0.05).

* **Detectability index**

The values of detectability indexes (d') obtained for the simulated

contrast-enhanced lesion on the two CT systems are depicted in the Fig. 4.B. d' values decreased as the energy level increased and the dose level decreased. For the DSCT-Pulse and for all energy levels, d' values decreased significantly on average by 27.0 % ± 1.8 % from 11 mGy to 6 mGy (p < 0.05) and by 33.3 % ± 2.4 % from 6 mGy to 1.8 mGy (p < 0.05) and by 22.7 % ± 2.4 % (p < 0.05) and 31.3 % ± 2.4 % (p < 0.05) for DSCT-Force, respectively.

For all keV, similar d' values were found between DSCT-Force and DSCT-Pulse at 11 mGy (–1.0 % ± 3.1 %; p = 0.084). For other dose levels, d' values were significantly lower with DSCT-Pulse than with DSCT-Force (9.1 % ± 3.2 % at 1.8 mGy (p < 0.05) and 6.3 % ± 3.9 % at 6 mGy (p < 0.05)).

<figure><img src="image_7.png" /><figcaption>Fig. 4. Contrast values between the phantom's material background and the iodine insert at 10 mg/mL (A) and detectability index (d') values for the simulated lesion with contrast characteristics corresponding to the iodine insert at 10 mg/mL (B) according to the energy level (keV) for the two dual-source CT systems and all dose levels. Values are expressed as mean ± standard deviation (error bars).</figcaption></figure>
## 4. Discussion

In this study, we assessed the quality of virtual monoenergetic images at low energy levels on a new dual-source CT for the first time. We found that the behavior of noise, spatial resolution and detectability of a simulated contrast-enhanced abdominal lesion according to keV and dose level were similar to those of a previous DSCT. For both DSCT systems, the best detectability was found at 40 keV, whatever the dose level. Compared with the previous DSCT, the new DSCT showed a finer noise texture and a somewhat higher spatial resolution for the employed very soft kernel, but noise magnitude was slightly higher and detectability slightly lower, particularly when the dose level was reduced.

The results of this study confirm the results already published on the previous DSCT (SOMATOM Force). Whatever the dose level, the noise magnitude and detectability values decreased from 40 to 70 keV and the opposite pattern was found for noise texture and spatial resolution [25,26]. We also confirmed that image noise increased and the detectability of the simulated lesion, spatial resolution and noise texture decreased as the dose level decreased [25,26]. The same behavior was found with the new DSCT system for all image quality metrics. Comparing the performance of this new DSCT system with the previous one, we found that the noise values were higher with the new system at all keV levels. The difference in noise between the two DSCT systems increased as the dose level decreased. Regarding noise texture, we found that the f<sub>av</sub> values shifted towards higher frequencies for all dose levels and all keV with the new DSCT compared to the previous DSCT. Similar results were found for spatial resolution where the f<sub>50</sub> values for the iodine insert also shifted towards higher spatial frequencies with the new DSCT. These combined outcomes led to VMIs that were slightly noisier but with less pronounced smoothing and somewhat better spatial resolution with the new DSCT for the employed reconstruction parameters in this case. Regarding the detectability index results, we found that the d' values were similar for all energy levels at 11 mGy, which is the dose level usually used in clinical routine for dual-energy spectral imaging of the abdomen. For lower dose levels corresponding to low-dose or ultra-low dose acquisitions, the higher noise values with the new DSCT led to lower d' values compared with the previous DSCT. However, the differences in d' values are only slight and less than 10%.

The differences in results found between the two DSCT systems may be related to technical modifications made to the new version. Although both DSCT systems have the same types of energy-integrating detectors and similar reconstruction kernels and reconstruction algorithms, they have different intrinsic parameters. First, the geometry of the two systems is different. The new DSCT system has a smaller bore size, but also a narrower detector array along the x-y axis and a shallower detector array along the z axis. Second, the two X-ray tubes used on the new system are different from those used on the other DSCT and have different performance parameters, particularly in terms of mA range, power. Furthermore, the size of the focal spot and the inherent filtration thickness are also different, which should affect the quality of the photon beam. Last, the differences that seem to be most relevant are the differences in kV pairs and tin filter thickness between the two DSCTs. Two pairs of kV combination are available for DECT on the new system, whereas five are available on the other one. For X-ray tube A, 80 kV and 100 kV are available on both systems but for X-ray tube B there is no equivalence between the two DSCT systems. In fact, 140 kV is used for the new system whereas 150 kV is used for the other one. In addition, the thickness of the tin filter used for these two high kVs is also different, being thinner on the new DSCT system as it uses 140 kV instead of 150 kV. A lower kV and a thinner filter on X-ray tube B may affect the spectral separation between the two spectra of low- and high-energy photons. This in turn could affect the spectral performance and therefore explain the little higher noise values and lower detectability values on VMIs with the new DSCT system compared with the other one. These initial results on phantoms now need to be confirmed on further patient studies for different clinical indications.

This study has certain limitations. First, the acquisitions were performed with only one level of an iterative reconstruction algorithm and one quantitative reconstruction kernel. Although these two parameters were chosen because they are used in clinical routine, other parameters may show different outcomes. Second, only one pair of kV was used on each of the two DSCT systems. Other pairs of kV, may also lead to different outcomes as they could impact spectral separation and therefore spectral performance [28]. Third, we only used one phantom, which does not take into account the diversity of patients' body morphologies. Fourthly, we did not directly compare the material decomposition capability of the two DSCT systems. These analyses could be carried out as part of a future study to highlight the differences between the two scanners in terms of their ability to detect and quantify base materials. Finally, as in other previous studies [14,25], we only defined one task function to model the detection of a contrast-enhanced lesion, with contrast defined according to contrast variations linked to monoenergetic images for the iodine insert used. This method was chosen to take into account the variations in contrast of the simulated lesion according to energy level. However, the d' calculated for strong contrast enhancement lesion might not be representative for small iodine contrast lesions. As defined in a previous study [14], a clinical study on known contrast-enhanced lesions of a specific size really needs to be carried out to confirm these initial phantom results.

## 5. Conclusion

In conclusion, the new DSCT system showed higher values regarding noise texture and spatial resolution of low-energy VMIs compared to a previous DSCT for a certain very soft kernel, but noise magnitude was slightly increased and detectability was slightly lower, particularly when the dose level was reduced. For both DSCT systems, the best detectability of a simulated contrast-enhanced lesion was found at 40 keV whatever the dose level. These initial results on phantoms must now be confirmed via further patient studies for different clinical indications.

## 6. Funding sources

This research did not receive any specific grant from funding agencies in the public, commercial, or not-for-profit sectors.

## Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

## Acknowledgements

We thank Teresa Sawyers for her help in editing the manuscript. We also thank Cedric Croisille for his help in performing acquisitions on new DSCT system.

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