Atherosclerotic cardiovascular disease is the leading cause of death worldwide. Although tremendous efforts have been made to prevent, diagnose, and treat the disease over the past few decades, mortality continues to increase (1,2). Atherosclerosis primarily begins when plasma cholesterol levels become elevated. This scenario, among other factors, leads to the formation of atherosclerotic plaques (3). Plaque in the artery wall narrows the artery lumen and may result in myocardial infarction or cerebral stroke. Studies have demonstrated that most acute infarctions are correlated with vulnerable plaques (4). Vulnerable plaques are characterized by luminal surface ulceration, intraplaque hemorrhage, a fissured fibrous cap, and lipid-rich necrotic core, neovascularization, and inflammation, plaque remodeling, and increased size (5). Rupture of a vulnerable plaque leads to platelet accumulation and artery occlusion, resulting in organ infarction. Magnetic resonance imaging (MRI) is useful for detecting carotid artery plaques and their components, and the imaging modality has helped prevent cerebral infarction (6-10).
In recent years, deep coronary artery wall MRI has been introduced for detecting coronary artery plaques (11). Due to the limitations of the current MRI technique, a voxel size of 0.7 mm × 0.7 mm × 2 mm is usually used in human coronary artery wall imaging at 3.0 Tesla (T) (12,13). However, the voxel size is larger than the artery wall, which has a thickness of approximately 0.2–0.3 mm (14). Therefore, the vulnerability of the plaque is difficult to determine directly.
Performing MRI on deep arteries can provide low-resolution images, which presents a challenge in plaque imaging. An intravascular MRI technique was developed, which involved the placement of a magnetic resonance (MR) radio frequency (RF) coil into the target vessel to generate intravascular high-resolution wall imaging to address this challenge (15,16). Due to the limited space in vessels, the intravascular RF coil design and the miniature coaxial transmission line coils (17,18), including loopless MR RF antenna, were advantageous because they were able to fit into smaller vessels and were more easily manipulated (19). The loopless MR antenna, also known as the MRIG, added additional functionality by acting as a conventional guidewire for MR-guided intravascular intervention, such as balloon angioplasty (20). The MRIG was further refined to 0.014 inches (about 0.35 mm) in diameter and used on 1.5 T MRI for coronary artery imaging (21,22).
Studies have shown that intravascular imaging can acquire high-resolution vascular images of 0.08 mm in ex vivo tissues with a 2.2 mm loopless antenna detector (23). Given the improved sensitivity, the higher magnetic field strength is expected to be beneficial in high-resolution image acquisition (24-30). To the best of our knowledge, no reports exist on the feasibility of high-resolution imaging with an MRIG as thin as 0.014 inches at 3.0 T MRI.
This study aimed to validate the feasibility of generating a high-resolution intravascular 3.0 T MRI of the coronary artery wall for further plaque imaging.
In this study, a 3.0 T “receive only” MRIG was manufactured from a 0.014 inch diameter coaxial cable (UL1005, Sumitomo Electric Inc., Osaka, Japan) to fit inside the coronary artery. The coaxial cable conductors were made of Sn-Cu alloy (Sn: 8.8×106 S/m, Cu: 5.8×107 S/m). We used a soft coaxial cable as the material of the MRIG, as it had a higher conductivity compared to the Nitinol-based MRIG (Nitinol: 1×106 S/m) (31) and was safer for coronary artery intervention. The inner conductor with an insulation layer was extended 10 cm beyond the outer conductor shield, and the total length of the MRIG was longer than 40 cm. The insulation layer broadened the longitudinal region over which images could be acquired (32) and improved the strength of the inner conductors.
The proximal end of the coaxial cable was connected to a 3.0 T MR scanner (GE Discovery 750W, GE Healthcare, Best, USA) through a matching/tuning circuit (Figure 1A). Regarding the circuit, the S11 of the MRIG (Figure 1B) was adjusted to below −20 decibel (dB) at the operating frequency of 127.7 megahertz (MHz). As shown in Figure 1A, the capacitances C2 and C3 at 13 picofarads (pF), and the inductors L1 and L2 at 120 nanohenry (nH) formed a parallel resonant circuit. The bias voltage was transmitted via the coaxial cable, and the switching diodes D1 and D2 were driven by the bias voltage. When the switches were on, the 2 detuning circuits showed high impedance. It likely protected the scanner from unexpected high voltage surges. By adjusting the capacitances of C4, C5, and C6, a wide range of frequencies covering the operating frequency was achieved (Figure 2). We set the capacitance of C1 to 1 microfarad (μF) to isolate the direct current (DC) signal. In addition, the inductors L3, L4, and L5 at 2.7 μH were used to isolate the alternating current (AC) signal.
In vitro study
To test the functionality of the MRIG, in vitro experiments were performed. We constructed a plastic box containing a plastic tube inside to mimic the vessel, and filled the plastic tube with 0.9% normal saline (Figure 3). The outside diameter of the tube measured approximately 3 mm, and the thickness was 0.5 mm. The plastic box and the tube were both made of polyethylene, through which no MR signal could be generated. We placed the MRIG into the plastic tube “vessel”. Axial T2-weighted MRI of the plastic tube “vessel” was then acquired using a fast spin-echo (FSE) sequence where repetition time (TR) =3,000 ms, echo time (TE) =68 ms, echo train length =6, field of view (FOV) =100 mm × 100 mm, matrix =512×512, slice thickness =2 mm, and number of averages =4. With the same MR parameters and position, we acquired phantom imaging with a commercial 16-channel surface coil placed approximately 5 cm superior to the plastic tube.
Ex vivo study
For the ex vivo study, we placed a fresh adult swine heart into a plastic box filled with 0.9% normal saline. After placing the MRIG into the left anterior descending (LAD) branch of the coronary artery, an MRI of the coronary artery wall was acquired. Axial T2-weighted imaging of the LAD branch was performed using a FSE sequence with TR =3,000 ms, TE =68 ms, echo train length =6, FOV =100 mm × 100 mm, matrix =512×512, slice thickness =2 mm, and number of averages =4. Maintaining the same position, we adjusted the FOV and matrix size to 110 mm × 110 mm and 160×160, respectively. We then acquired high-resolution and conventional resolution images of the same LAD branch of the coronary artery. The commercial surface coil was placed about 5 cm superior to the LAD branch to acquire wall imaging using the same settings. The experiments were repeated 4 times, and images were acquired from 4 different swine hearts using the same guidewires.
After obtaining all the images, the noise was acquired by maintaining all conditions and manually setting the RF gain to 0. The signal-to-noise (SNR) of each voxel was calculated using the following formula (33):
where S is the signal intensity of each channel and ψ is the noise resistance. A reference line was drawn through the tube center to demonstrate the spatial SNR distribution visually in the phantom study, and the value that varied with distance from the center was plotted.
In the ex vivo study, the highest achievable SNR was measured by placing the region of interest (ROI) at the brightest part of the coronary artery walls. To calculate the average SNR of the swine arterial wall (Sw), we measured the average SNR (St) of the total area and the average SNR (Sl) of the vessel lumen. The Sw was calculated using the following formula (34):
Where At is the total vessel area, and Al is the vessel lumen area.
After achieving MRI, the coronary artery segments were harvested and cryosectioned at 5 µm. We then performed histological examinations with hematoxylin and eosin (HE) staining to detect the integrity of the artery wall.
In the in vitro experiments, both the MRIG and commercial surface coil functioned well as receiver coils when used with the 3.0 T MR scanner. In both images, the plastic tube “vessel” was able to be identified at the resolution of 0.2 mm × 0.2 mm × 2 mm. The vessel wall imaging by the MRIG could be clearly identified, whereas the plastic tube “vessel” imaged by the surface coil was blurred (Figure 3A,B). As shown in Figure 3C, a series of SNR values along the reference line through the plastic tube “vessel” was achieved. The peak SNR of the MRIG and the surface coil-generated imaging was 213.6 and 19.8, respectively. The signal decreased rapidly as the distance from the MRIG increased. The calculated curve indicated that the SNR was higher than the surface coil in the range of ±0.5 cm near the intersection of the inner conductor and the outer conductor of the MRIG.
In the ex vivo experiments, LAD branch wall imaging was acquired using the MRIG and the surface coil (Figure 4). As shown in the images, the MRIG-acquired LAD branch wall was able to be identified at the resolution of 0.2 mm × 0.2 mm × 2 mm. The surface coil-acquired LAD branch wall was blurred. Compared to the high-resolution images, the conventional resolution images of 0.7 mm × 0.7 mm × 2 mm lacked the detail of the high-resolution images. The highest achievable SNR of the artery wall was 56.7±20.1 compared to the SNR of 19.1±4.1 that was obtained by the surface coil. The average SNR (Sw) of the artery wall obtained by the MRIG was 21.1±5.40, compared to the SNR of 8.4±2.19 obtained by the surface coil (Figure 5). In Figure 6, as expected, the average SNR in ROIs obtained using the MRIG were 16.9, 15.2, and 15.5, which was much higher than the SNR of 6.2, 7.3, and 6.6, obtained using the surface coil.
The HE staining in Figure 7 showed no evidence of injury at the target vessel walls.
It is difficult to build an MRIG inserted into a vessel smoothly without scratching the vessel wall and causing spasms. To achieve this, the tip of the MRIG needs to be soft while maintaining a certain degree of toughness. To this end, we stripped off the outer conductor of a commercial coaxial cable but maintained its insulating layer. This insulating layer wrapped around the inner conductor to sufficiently maintain its mechanical properties.
In conclusion, we validated the feasibility of safely generating 0.2 mm × 0.2 mm × 2 mm MR images of the coronary artery wall in rich detail using a 0.014-inch MRIG. The technique helps to lay the foundation of clinical practice for high-resolution MRI of coronary artery plaques.
Funding: This work was supported by the National Natural Science Foundation of China (Grant No.: 81671789); the Project of Health Commission of Shanxi Province (No.: 2018109); National Key R&D Program of China 2021YFE0204400; the Strategic Priority Research Program of the Chinese Academy of Sciences (Grant No. XDB25000000); city grant RCYX20200714114735123; Guangdong Province grant 2020B1212060051; and Youth Innovation Promotion Association of CAS No. 2017415.
Conflicts of Interest: All authors have completed the ICMJE uniform disclosure form (available at https://dx.doi.org/10.21037/qims-21-286). Dr. XZ serves as an Associate Editor of Quantitative Imaging in Medicine and Surgery. Dr. YM reports that his institution is funded by the National Natural Science Foundation of China (Grant No.: 81671789) and the Project of Health Commission of Shanxi Province (No.: 2018109). Dr. YL reports that his institution is funded by the National Key R&D Program of China 2021YFE0204400; the Strategic Priority Research Program of the Chinese Academy of Sciences (Grant No. XDB25000000); city grant RCYX20200714114735123; Guangdong Province grant 2020B1212060051; and Youth Innovation Promotion Association of CAS No. 2017415. The other authors have no conflicts of interest to declare.
Ethical Statement: The authors are accountable for all aspects of the work by ensuring that questions related to the accuracy or integrity of any part of the work are appropriately investigated and resolved. Ethical approval and written informed consent were not applicable because no human subjects were involved in our experiments.
Open Access Statement: This is an Open Access article distributed in accordance with the Creative Commons Attribution-NonCommercial-NoDerivs 4.0 International License (CC BY-NC-ND 4.0), which permits the non-commercial replication and distribution of the article with the strict proviso that no changes or edits are made and the original work is properly cited (including links to both the formal publication through the relevant DOI and the license). See: https://creativecommons.org/licenses/by-nc-nd/4.0/.
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