An Interventional MRI guidewire combining profile and tip conspicuity for catheterization at 0.55T

Dursun Korel Yildirim; Dogangun Uzun; Christopher G Bruce; Jaffar M Khan; Toby Rogers; William H Schenke; Rajiv Ramasawmy; Adrienne Campbell-Washburn; Daniel Herzka; Robert J Lederman; Ozgur Kocaturk

doi:10.1002/mrm.29466

. Author manuscript; available in PMC: 2024 Feb 1.

Published in final edited form as: Magn Reson Med. 2022 Oct 5;89(2):845–858. doi: 10.1002/mrm.29466

An Interventional MRI guidewire combining profile and tip conspicuity for catheterization at 0.55T

Dursun Korel Yildirim ¹, Dogangun Uzun ^1,², Christopher G Bruce ¹, Jaffar M Khan ¹, Toby Rogers ¹, William H Schenke ¹, Rajiv Ramasawmy ¹, Adrienne Campbell-Washburn ¹, Daniel Herzka ¹, Robert J Lederman ^1,^*, Ozgur Kocaturk ^2,^*

PMCID: PMC9712240 NIHMSID: NIHMS1834438 PMID: 36198118

Abstract

Purpose:

We describe a clinical grade, “active”, monopole antenna-based metallic guidewire that has a continuous shaft-to-tip image profile, a pre-shaped tip-curve, standard 0.89mm (0.035”) outer diameter, and a detachable connector for catheter exchange during cardiovascular catheterization at 0.55T.

Methods:

Electromagnetic simulations were performed to characterize the magnetic field around the antenna whip for continuous tip visibility. The active guidewire was manufactured using medical grade materials in an ISO Class 7 cleanroom. RF-induced heating of the active guidewire prototype was tested in one gel phantom per ASTM 2182-19a, alone and in tandem with clinical metal-braided catheters. Real-time MRI visibility was tested in one gel phantom and in-vivo in two swine. Mechanical performance was compared with commercial equivalents.

Results:

The active guidewire provided continuous “profile” shaft and tip visibility in-vitro and in-vivo, analogous to guidewire shaft-and-tip profiles under X-ray. The MRI signal signature matched simulation results. Maximum unscaled RF-induced temperature rise was 5.2°C and 6.5°C (3.47 W/kg local background SAR), alone and in tandem with a steel-braided catheter, respectively. Mechanical characteristics matched commercial comparator guidewires.

Conclusion:

The active guidewire was clearly visible via real-time MRI at 0.55T, and exhibits a favorable geometric sensitivity profile depicting the guidewire continuously from shaft-to-tip including a unique curved-tip signature. RF-induced heating is clinically acceptable. This design allows safe device navigation through luminal structures and heart chambers. The detachable connector allows delivery and exchange of cardiovascular catheters while maintaining guidewire position. This enhanced guidewire design affords the expected performance of X-ray guidewires during human MRI catheterization.

Keywords: Interventional MRI, active MRI guidewire, curved tip active MRI guidewire, low-field MRI, MRI devices, MRI safety

1. Introduction

Guidewires are workhorse tools for interventional cardiology. They are used for navigation through vascular structures, accessing target lesions, facilitating passage through diseased segments, and delivering and exchanging diagnostic and therapeutic devices.¹ Real-time MRI is an attractive alternative to X-Ray fluoroscopy for the image guidance of interventional procedures because it provides high soft tissue contrast without ionizing radiation, simultaneous multi-planar real-time imaging with frame rates up to 10 frames per second and intra-operational assessment of hemodynamics and tissue characteristics.^2–4 Additionally, intravascular MRI enables high-resolution vessel wall imaging during interventional procedures, which cannot be seen by conventional X-ray imaging.^5,6 However, the clinical adoption of interventional MRI (iMRI) is hampered by a paucity of iMRI devices and in particular, a clinically safe and conspicuous guidewire.^2,7

iMRI devices require interactions with the main magnetic field (B₀) and surrounding proton spins, to be rendered visible under MRI.⁷ iMRI devices are classified as passive,⁸ semi-active^9,10 and active devices¹¹ in terms of device visualization. Active devices incorporate RF antennas and electronic components embedded into the device body and their signal is directly carried to the MR scanner, providing high SNR device visibility, significantly superior to that achievable in passive devices.^11,12 Their signal can be separated from other receive channels during image processing and can be colorized and overlaid onto the anatomical image for augmented conspicuity.¹³ However, long metallic components used in active iMRI device designs are prone to RF-induced heating during real-time MRI.^14,15 Interventional X-Ray fluoroscopy devices use conventional metals such as stainless-steel to provide adequate mechanical characteristics. In addition to RF-induced heating risk, the strong magnetic field employed by MRI systems limits the use of such metals for iMRI device designs because of magnetic displacement forces and prohibitive magnetic susceptibility imaging artefact. As a result, design and engineering of safe and conspicuous iMRI guidewires becomes challenging because of the electrical and mechanical requirements, and size limitations.

Attempts to replace metallic components with polymer alternatives to avoid RF-induced heating risk of iMRI guidewires¹⁶ have been hampered by poor mechanical performance. Alternatively, passive guidewire and catheter designs were introduced using segmented and electrically insulated yet mechanically coupled metallic components.^17,18 However, conspicuity of these passive devices cannot compete with active designs.

Considering the size limitations, the loopless antenna structure can be exploited to design active iMRI guidewires thanks to its simplicity and high near-field SNR. One important advantage of a loopless antenna-based iMRI guidewire design is that the guidewire can be made as thin as 0.36 mm (0.014”), which is specifically important for coronary interventions.¹⁹ Previously, monopole loopless antenna-based active guidewire designs were introduced by extending the inner conductor of a coaxial transmission line, providing continuous device visibility.^6,19–22 Physician-operators need to visualize the whole guidewire and the exact tip location to safely operate within the vascular structures. One drawback of the loopless antenna design is that the current intensity along the antenna pole (whip) and the sensitivity profile of the antenna decrease to zero towards the distal tip which renders the distal tip of the guidewire invisible under MRI, risking the operational safety and efficacy.²³ Potential solutions have included packing the whip as a coil,⁵ tapering the distal part of the whip,²⁴ incorporating a solenoid loop coil at the distal end as a separate channel,²⁵ inductively loading the antenna by using a backwards wound coil at the distal end,^26,27 and exposing a portion of the distal tip and actively controlling the current flow on the guidewire.²⁸ However, these methods either restricted the usable device length or overcomplicated the design. Moreover, mechanical and electrical safety risks of the proposed designs require further elaboration.^29,30 As a result, there are still no commercial, electrically safe active guidewires available having acceptable mechanical performance and conspicuity.

On the other hand, the recent introduction of high-performance, low-field MRI allows a ~9-fold reduction in specific absorption rate (SAR) for approximately equal B_1rms,^31,32 and an increased electrical wavelength^14,22 compared to 1.5T systems, for a given imaging sequence, enabling safe-by-design iMRI devices.^12,33

In this study, we introduce a novel, 0.89mm (0.035”) outer diameter (OD) actively visualized, monopole antenna-based metallic guidewire design for human interventional catheterization procedures at 0.55T. Distinguishing from the previous studies, the proposed design here incorporates a clinically required, non-traumatizing curved tip geometry, a unique combination of continuous tip and full shaft visibility, clinically acceptable mechanical characteristics and a detachable RF connector which allows catheter delivery and exchange over the active guidewire.

2. Methods

Design specifications of the active guidewire prototype were defined by the clinical cardiology users in our group. The final prototype was expected to have a 162.5cm total length and 0.89mm OD with a detachable RF connector to deliver and exchange commonly used cardiovascular catheters over the guidewire. Mechanical performance of the active guidewire was expected to be comparable with commonly used commercial equivalents. A non-traumatizing curved tip geometry with a unique tip and continuous shaft visibility was required to allow safe navigation and steerability through various luminal structures. A series of electromagnetic simulations were performed to define the optimum tip coil specifications and provide continuous visibility at the distal tip of the active guidewire.

2.1. Electromagnetic simulations

A comprehensive explanation of the monopole antenna-based active iMRI guidewire concept²⁰ and its detailed numerical analysis³⁴ were introduced before. It was demonstrated that inductive loading using a solenoid coil, electrically connected to the distal tip, recovers the weakened monopole antenna signal by increasing the charge concentration towards the tip.^26,35 Other studies showed that insulating the whip alters the effective wavelength of the monopole antenna and provides a more homogenous current intensity throughout the whip resulting in a higher SNR,^5,24 and using a backwards wound tip coil at the monopole antenna end provides a distinct tip signal at 1.5 T.^26,27 It was also shown that RF magnetic (B) field distribution around a solenoid coil and the signal signature on the MR image can be altered adjusting the solenoid coil pitch (p).^12,36,37 Here, we hypothesize that a full shaft and continuous tip visibility can be provided by a monopole antenna-based active MRI guidewire by optimizing the tip coil parameters. Finite difference time domain method electromagnetic simulations were performed using Sime4Life 6.0 (Zurich MedTech AG, Switzerland) to study the B-field along the whip and around the tip coil and to optimize the tip coil length to provide a continuous tip signal.

The active guidewire prototype was modeled as a monopole antenna by extending the inner conductor of a coaxial structure. All conductor materials were defined as nitinol with 80Ω/m impedance at 23.66 MHz. The polymer material insulating the inner and outer conductors and covering the active guidewire was defined as a dielectric material with electrical conductivity of 0 S/m and relative permittivity of 3.2. A finite volume encapsulating the antenna whip was defined as a gel material with electrical conductivity of 0.47 S/m and relative permittivity of 78, mimicking human body loading conditions. This definition simulates partial immersion of the active guidewire in the gel phantom. The active guidewire was aligned parallel to B₀ since this is the most common use condition of a guidewire and it was excited at the “junction” where the inner conductor emerges from the hypotube, at 23.66 MHz with a 1A sinusoidal coaxial source to exploit the reciprocity principle. Only the wire diameter of 0.076mm and 0.051mm were simulated for the tip coil since they are the thinnest insulated MP35N and 35N LT wires, respectively, that are commercially available. Normalized solenoid pitch and coil OD were kept 0.2 and 0.762mm, respectively. Coil length was adjusted between 0 and 60mm in 10mm increments during the simulations. B₁⁺ field distributions were extracted on a parallel line 1mm apart from the whip. Similarly, the B-field phase was extracted on the same line but only at the last 80 mm of the whip. Finally, the same active guidewire geometry was simulated at 63.86 and 127.74 MHz to estimate the maximum local SAR and RF-induced heating rise at the hot spot at 1.5T and 3T. Time averaging based on the imaging sequence was not performed for calculating the maximum SAR. Electromagnetic simulations were performed for a straight tip guidewire design since implementing a curved helical geometry with a large number (>10⁹) disproportionally small elements dramatically increases the simulation time and required computational resources with little expected impact on findings. Tip coil length was optimized experimentally for the curved tip geometry, based on the simulation results.

2.2. Active iMRI guidewire design

The active guidewire prototype incorporating a modified monopole loopless antenna was fabricated in an ISO Class 7 cleanroom using medical grade MRI compatible materials. A custom, ground nitinol rod and a superelastic nitinol hypotube with 0.787mm OD and 0.610mm ID were used to form the monopole antenna design. These metal components were coated with thermoplastic Pebax tubing using a hot air source (210-A, BEAHM Design Inc., CA). The inner conductor was extended 37.5cm (~λ/4 at 23.66 MHz in human body loading conditions)¹⁴ out of nitinol hypotube to maximize the RF receiver antenna performance. The inner conductor diameter was first sharply increased after the junction and then gradually reduced along the whip to avoid abrupt stiffness changes between the hypotube and the extended inner conductor (Figure 1A). A tight pitch tip coil with 0.762mm coil OD, made of a 0.051mm OD Polytetrafluoroethylene (PTFE) insulated 35N LT wire (Fort Wayne, IN) was placed at the distal end of the whip and soldered to the whip only at the distal end, and the proximal end remained floating. Finally, the overall guidewire length was coated with Pebax thermoplastic polymer for electrical insulation purposes. (Figure 1A and 1B). The optimum solenoid coil length was 50mm, based on simulation results of the straight whip geometry, and experimentally decreased to 14mm for the curved tip geometry.

Figure 1. — (A) Technical drawing of the active guidewire prototype. Point A: Distal tip and tip coil. Point B: Junction. The whip is between points A and B. The shaft is between points B and C. (B) Final assembly including the detachable handle. (C) Opened-view of detachable handle showing the modified torquer and tune and match and leakage current blockage circuit.

A mmcx connector (262125, Amphenol RF, CT) was soldered to a commercially available torquer device (97327, Qosina, NY). The torquer device was used as a detachable coaxial RF connector by clamping on the active guidewire at the proximal end where the polymer coating is denuded (Figure 1C). In this way, the 0.89mm guidewire OD was maintained when the handle was detached. The monopole antenna impedance was matched to 50Ω at 23.66 MHz via a tune and match circuit using a vector network analyzer (VNA) (E5080B, Keysight, CA). A solenoid cable trap BALUN was used to eliminate shield common mode currents and a PIN diode was used to actively detune the monopole antenna during RF transmission.²² In addition, a leakage current block was added to the tune and match circuit to mitigate the micro shock risk by avoiding current leakage from the MRI RF receiver chain to the active guidewire. The active guidewire prototype was limited to function only in receive mode and no RF energy was allowed to be transmitted over it.

Additionally, three straight guidewire prototypes (one with no tip coil, two with 50mm long, 0.762mm OD tight pitch tip coils made of the insulated MP35N and 35N LT coil wires [Fort Wayne, IN], respectively) and a curved tip guidewire prototype with a 24mm long, 0.762mm OD tight pitch tip coil made of the insulated MP35N coil were manufactured to assess the effects of invisible-tip and distinct-tip imaging signatures.

2.3. In-vitro characterization

Electrical characteristics of the active guidewire prototype in air and in an ASTM gel phantom with electrical conductivity of 0.47 S/m and relative permittivity of 78, were assessed using the VNA. The leakage current tests were performed using a power supply (Vega 650, TDK-Lambda, UK) used in Siemens Aera scanners, a vendor provided 4-channel flex coil adaptor and a leakage current tester (LT-601HC, ED&D Inc., NC). In addition to the electrical tests, the entire guidewire assembly and the detachable RF connector were exposed to ethylene oxide (EtO) sterilization for 12 hours to observe the effects of EtO sterilization on the electronic components and guidewire performance.

A real-time balanced steady-state free precession (bSSFP) sequence (TE/TR: 1.05/2.50ms, flip angle: 45°/60°/75°, FOV: 320 × 320mm, bandwidth: 1042Hz/Px, matrix: 192×144, slice thickness: 6mm, 3 fps) was used during in-vitro tests.¹² In-vitro RF-induced heating and imaging performance of the active guidewire prototype was tested using a high-performance low-field scanner (prototype MAGNETOM Aera, Siemens Healthcare, Erlangen, Germany) at 0.55T.

RF-induced heating performance of the active guidewire prototype was tested in a gel phantom per ASTM 2182–19a.³⁸ The distal 50cm portion of the active guidewire prototype was immersed in the gel phantom, aligned parallel to B₀, 2cm away from the right edge of the phantom and 3cm below the gel surface where the SAR is maximum (as observed in previous experiments)^12,39,40 and heat dissipation is minimum. Scans were performed in the “First Level” SAR mode (max. 4W/kg) with 45°, 60° and 75° flip angles. Background SAR was measured using an MRI compatible SAR probe (EX3DV4, SPEAG, Zurich, Switzerland) calibrated for the real-time bSSFP sequence in the ASTM gel phantom. 0.080mm OD fiberoptic temperature probes (OPT M-170, OpSens, Canada) were used for temperature measurements as previously described.¹⁸ First, the “hot spot” where the maximum heating occurs, was located by profiling the temperature rise alongside the active guidewire prototype during MRI scan while withdrawing the temperature probe. Next, RF-induced heating tests were performed as the temperature probe was positioned at the hot spot. Baseline temperature was recorded for 30 s. before each scan and then the active guidewire prototype was scanned for 90 s., which represents the start of saturation after the clinically significant temperature rise. Measurements were repeated using the active guidewire prototype in tandem with a 316L steel-braided 6 Fr clinical catheter (Infiniti, Cordis, FL) in various tandem configurations to assess the metal interaction. Finally, the active guidewire prototype was scanned with 75° flip angle for 15 minutes, both alone and in tandem with the metal-braided catheter in the configuration causing maximum heating, to simulate a challenging, prolonged scan. Test results were scaled to 1W/kg using the background SAR values.

In-vitro MRI visibility performance of the active guidewire prototype was tested using the same bSSFP sequence but with only 45° flip angle in the ASTM gel phantom since a 45° flip angle provides adequate tissue and device SNR to safely perform interventional cardiovascular procedures at 0.55T.¹² The active guidewire prototype was placed parallel to B₀, 5cm away from the right edge of the phantom and 4cm below the gel surface. It was scanned alone and in tandem with a list of commercially available, commonly used catheters (Table 1). Additionally, alternative guidewire prototypes having (1) no tip coil and (2) a MP35N wire tip coil were scanned alone under the same conditions. Static MR images were reconstructed using only the active channel signal received from the active guidewire antenna and all other imaging coils were deactivated. All prototypes were scanned with the same RF receiver gain. The pseudo-replica method⁴¹ with N=100 was used to generate pixel-wise SNR maps of the prototypes with tip coils, using Matlab (2020a, Mathworks, MA) to assess the signal profile along the antenna whip. Real-time in-vitro MRI visibility of the active guidewire prototype was tested using a prototype interactive real-time MRI interface (MonteCarlo, Siemens, Germany) with a real-time bSSFP sequence (TE/TR:2.0/4.0ms, FA:45°, FoV: 320×320mm, bandwidth: 1042 Hz/Px, matrix: 192×144, ST:8mm, 3 fps). First, the active guidewire prototype alone was manipulated in a 3D-printed left heart model. Next, the 316L steel-braided 6 FR pig-tail catheter (SuperTorque, Cordis, FL) was operated using the active guidewire prototype. An eighteen-element spine coil was used for phantom imaging. The active channel signal was colorized on the MR image.

Table 1.

List of commercially available catheters with which the active guidewire prototype was tested in tandem.

Commercially available catheter	Braiding type

Cordis, Super Torque Angiographic Catheter, 5 Fr.	Metallic
Cook, Flow Directed Balloon Catheter.	None
Medtronic, Pulmonary Wedge Pressure Catheter.	None
Vascor, Balloon Wedge Pressure Catheter.	None
Edwards, True Size Monitoring Catheter.	None
Edwards, True Size Hi-Shore Monitoring “T” Tip Catheter.	None
Arrow, Balloon Wedge Pressure Catheter.	None
Edwards, True Size Monitoring “S” Tip Catheter.	None
Cordis, Super Torque JR4.	Metallic
Cordis, Super Torque MPA.	Metallic
Medtronic, Woven, NIH.	Non-metallic
Medtronic, Woven, Goodale-Lubin nylon-braided Catheter.	Non-metallic

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Mechanical performance of the active guidewire prototype was tested for torque transmission, stiffness and tracking, and compared to commercially available, commonly used clinical guidewires including a Glidewire (Terumo, NJ), a HiWire (Cook, IN) and a Nitrex (stiff) (Medtronic, Dublin, Ireland). Mechanical tests were performed as described before (Supplemental material, Figure S1).¹⁸ The stiffness test was performed at the whip portion of the active guidewire prototype. Additionally, stiffness of the active guidewire prototype was tested by observing the straightening of the 316L-braided pigtail catheter by the guidewire advancement, in comparison with the commercial nitinol Glidewire.

2.4. In-vivo characterization

The real-time in-vivo MRI visibility performance of the active guidewire prototype was tested in two swine under general anesthesia via percutaneous transfemoral venous and arterial access. Animal experiments were approved by the National Heart, Lung and Blood Institute (NHLBI) Animal Use and Care Committee and performed according to contemporary National Institutes of Health (NIH) standards. First, the guidewire prototype alone was advanced retrograde across the aortic valve into the left ventricle. Next, the active guidewire prototype was advanced through the inferior vena cava, in tandem with a balloon catheter (Arrow, Teleflex, PA) into the pulmonary artery branches. The same real-time MRI setup with same sequence and parameters used for the in-vitro tests were used during in-vivo experiments. No projection-based tracking method was applied and the real-time device tracking was performed by manually adjusting imaging planes.

3. Results

Figure 2 and 3 show the electromagnetic simulation results of the active guidewire prototype with no tip coil, and tip coils made of insulated 0.076mm OD (MP35N) and 0.051mm OD (35N LT) wires. B₁⁺ field distributions show that the tip coil configuration alters the signal signature of a monopole antenna-based active guidewire. The estimated maximum local SAR at the hot spot was 158.6, 1298.5 and 1448.3 W/kg at 0.55T, 1.5T and 3T, respectively.

Figure 3. — B₁⁺ magnitude plots show that 50mm long insulated 35N LT tip coil provides the highest antenna sensitivity which translates to the highest expected SNR signature on the MR image. B-field phase graph shows a phase shift that varies non-linearly along the shaft towards the tip coil, generating a distinct tip signal signature.

The characteristic impedance of the coaxial part of the active guidewire prototype (r_o=0.787mm, r_i=0.127mm) was calculated as 61.13Ω. The input impedance of the modified monopole antenna was 22.2 - j15.4Ω in air and 37.6 - j4.9Ω in the gel phantom. The leakage current was measured as 45.1 and 9.8 μA before and after the leakage current block, respectively.

RF-induced heating test results are shown in Figure 4. The background SAR was measured as 1.42 W/kg, 2.37 W/kg and 3.47W/kg for 45°, 60° and 75° flip angles, respectively. The hot spot was at the distal tip of the active guidewire prototype. The temperature at the hot spot began to rise at the onset of imaging and saturated quickly after the first 60 seconds. The maximum unscaled temperature rise during a 75° flip angle, 90 seconds long scan was 4.64 and 5.98 °C when the active guidewire prototype was alone and in tandem with the metal braided catheter, respectively. Maximum in tandem temperature rise was observed when the distal tip of the metal braided catheter was aligned with the junction. The maximum unscaled temperature rise at the end of a 75° flip angle, 15 minutes long scan was 5.21 and 6.54 °C when the active guidewire prototype was alone and in tandem with the metal braided catheter, respectively. The maximum temperature rise scaled to 1W/kg background SAR was 1.50 and 1.88 °C when the active guidewire was alone and in tandem with the metal braided catheter, respectively.

Figure 5 shows in-vitro MR images and corresponding SNR maps of the active guidewire prototypes aligned parallel to B₀ and alone in the gel phantom. The 0.051mm OD coil wire provides a continuous tip signal while the 0.076mm OD coil wire provides a distinct tip signal. In-vitro MR images of the active guidewire prototype in tandem with commercially available clinical catheters are shown in Figure 6. In-vitro real-time MRI videos of the active guidewire prototype provided within the Supplemental material (Video S1 and S2) show that the continuous tip visibility provides information also about the wire rotation.

Figure 6. — In-vitro MRI of the active guidewire prototype in tandem with commercially available clinical catheters. Circular dark artifacts on images are signal voids caused by mounting pegs and/or air filled balloons of the catheters. Since the tested catheters have curved geometries, it is not always possible to align the entire active guidewire parallel to one plane which explains the SNR fluctuations along the active guidewire prototype.

Mechanical test results are compared with commonly used commercial equivalents and shown in Figure 7. The active guidewire stiffness at the whip portion was 6 N/m. Additional stiffness test results in tandem with the metal braided pigtail catheter is provided within the Supplemental material (Figure S2).

Figure 8 shows in-vivo MR images of the active guidewire prototype alone in the aortic arch and in tandem with the balloon catheter in the inferior vena cava. The active guidewire prototype provided in-vivo full shaft and continuous tip visibility. Corresponding in-vivo real-time MRI videos are provided within the Supplemental Material (Video S3 and S4).

4. Discussion and Conclusion

When there is no tip coil, the current intensity along the whip, and B-field strength, approaches zero towards the tip; this renders the distal tip invisible under MRI (Figure 2 and 3). When the whip length is extended using an insulated MP35N wire (OD=0.076mm) and the extension is coiled backwards onto the whip, the current intensity is pushed towards the open end of the tip coil increasing the current intensity at the distal end of the whip.^5,22 The current in a helical wire can be decomposed into longitudinal azimuthal components. Likewise, the azimuthal current can be decomposed into longitudinal and radial components.⁴² Since the tip coil and the straight whip wire are extended in opposite directions, the longitudinal currents on these two wires counteract. Counteracting longitudinal currents may cancel each other out, resulting in a lack of transverse B-field and displacing the signal void to the tip. After the signal void, the transverse B-field of the whip stands alone and provides the MRI visibility again. In this case a non-linear B-field phase shift along the tip coil, dramatically increasing towards the distal end is also observed (Figure 3). Such signal nulls can also be observed with short (<40mm) 35N LT (OD=0.051mm) tip coils. When a longer tip coil made of a 0.051mm OD 35N LT wire is used, a longer conductor can be packed in the same volume extending the whip length by multiple orders of λ/4 (~1430mm long wire [~3.8λ/4] can be packed in 50mm length). In this case, ignoring the parasitic capacitance between coil turns, a higher current intensity can be achieved towards the open end of the coil. As a result, the transverse magnetic field along the tip coil is maintained without signal nulls providing a continuous tip and full shaft visibility, and the B-field phase shift along the coil is reduced (Figure 3). Simulation results demonstrated that even though 40mm and 60mm long coils also provide a continuous tip signal, the maximum full shaft and continuous tip SNR can be achieved using a 50mm long 35N LT tip coil. The transmit reciprocity approach excluded immersed common mode interactions during electromagnetic simulations.

The tip coil length was experimentally shortened to 14mm to obtain a continuous tip signal for the guidewire design with curved tip geometry. B₁⁺ and B₁⁻ fields are the rotational components of the projection of the B-field on a plane normal to B₀. The shape of a coil and its orientation with respect to B₀ affect the B₁⁺ distribution around it altering the MRI visibility of the coil.⁴³ The requirement for shortening the tip coil of the active guidewire design with the curved tip geometry was attributed to the changing geometry and orientation of the tip coil.

The maximum RF-induced temperature rise, scaled to 1W/kg local background SAR, was 1.50 and 1.88 °C when the active guidewire prototype is alone and in tandem with a metal braided catheter, respectively, during 15 minute long scans. The maximum unscaled temperature rise remained safely under 6°C when the active guidewire prototype is alone and slightly exceeded 6°C (6.54°C) when it is in tandem with a metal braided catheter during the same scans (Figure 4). Maximum RF-induced heating was observed at the upper limits of the scanner allowed local SAR (75° flip angle, 3.47W/kg), with minimum heat conduction in a gel phantom and when the metal braided catheter was aligned right at the junction. This is attributed to the adjusted effective wavelength along the coaxial structure affecting the active detuning performance and to the common mode interactions between the active guidewire and the metal braiding of the catheter. Based on the simulation results, the maximum RF-induced temperature rise is estimated to be ~42.5 and ~47.5°C at 1.5T and 3T, respectively due to the linear proportionality between SAR and RF-induced heating. This is attributed to the reduced wavelength and increased deposited RF energy. No experimental tests were performed at 1.5 or 3T and the actual in-vitro temperature rise is expected to be lower due to heat convection/conduction. The safety limit for clinical heating is inferred from plasma protein denaturation beginning approximately 6°C above body temperature.⁴⁴ However, we hypothesize that the RF-induced temperature rise at 0.55T will safely remain under 6°C during a human catheterization due to the central in-vivo location of the device, cooling from perfusion and more heterogenous E-field distribution within the body. No adverse change in the hemodynamics was observed during the in-vivo animal experiments supporting our hypothesis. The RF-induced heating tests were performed in-vitro per ASTM 2182–19a during this study. To ensure complete in-vivo electrical safety, a more comprehensive RF safety assessment via the transfer function approach⁴⁵ per ISO/TS 10974:2018⁴⁶ is being developed by our group with preliminary results confirming the clinically safe (< 2°C) in-vivo RF-induced heating.

The mechanical performance of the proposed design was comparable to commonly-used commercial equivalents (Figure 7). The tapered nitinol inner conductor provided a smooth transition between the coaxial shaft and the whip. A pigtail catheter was straightened, meeting user input requirements, and a broad selection of catheters was delivered over the proposed design without compromising the MRI visibility performance (Videos S2 and S4, and Figures 6 and S2). In this study, the component dimensions were chosen for a slightly higher stiffness. However, those dimensions can be altered considering the electrical characteristics to achieve desired mechanical characteristics for specific applications.

In-vitro MR images of active guidewire prototypes matched electromagnetic simulation results for the straight tip geometry. The curved tip active guidewire design with the distinct tip signal lacks the curve information which is important for the assessment of the tip deflection and for eliminating the perforation risk. A 45° flip angle provided sufficient anatomical and device SNR to safely access target anatomy, minimizing the RF-induced heating risk. The unique curved tip geometry with continuous tip visibility allowed access to the left ventricle through the aortic arch crossing the aortic valve (Video S3). The active guidewire visibility was not affected by changing orientation with respect to B₀. The detachable RF connector in combination with the 162.5cm device length allowed intra-procedural delivery and exchange of cardiovascular catheters over the active guidewire prototype while maintaining guidewire position, another key user requirement (Videos S2 and S4). The continuous tip signal allowed the assessment of the active guidewire rotation and tip deflection in the inferior vena cava (Video S4). The near-field sensitivity of loopless monopole antennas diminishes approximately as the inverse of radial distance from the antenna and active interventional MRI devices provide a device conspicuity by exploiting the hyper-intense near-field MRI signal.²² The gradual fall-off of the antenna sensitivity results in a nebulous signal signature near the edges of the antenna sensitivity profile. However, the nebulous signal signature can be sharpened by adjusting the channel specific window width and window level values. After 12-hour long EtO sterilization, no adverse effect was observed in the material integrity or active guidewire performance during the in-vitro and in-vivo experiments.

This study introduces a monopole antenna-based actively visualized, metallic guidewire suitable for MRI-guided high performance human cardiovascular catheterization procedures at 0.55T. The proposed device has a 0.89mm (0.035”) OD, a fixed-curve distal tip required by users, and a detachable RF connector to allow in-situ exchange of catheter tools including balloons and stents. We demonstrate, for the first time, mechanical performance comparable to commercial equivalents with full shaft and continuous tip curvature visibility under MRI, without hazardous RF-induced heating. The proposed “safe-by-design” device is ready for clinical testing pending United States Food and Drug Administration (FDA) licensure and expected to respond to an urgent need in clinics^2,47–49 We expect this novel design to help other investigators to develop clinical-grade diagnostic and therapeutic MRI-catheter tools.

Supplementary Material

vS2

Video S2. In-vitro real-time MRI of the active guidewire prototype in tandem with a 5Fr pigtail catheter. The active guidewire prototype can mechanically straighten out the clinical catheter while providing a full shaft and continuous tip visibility.

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supinfo

Figure S1. Mechanical test setups for (A) the stiffness, (B) torque transmission and (C) trackability tests of the active guidewire prototype. The active guidewire prototype is fixed on the custom apparatus forming an arch for the stiffness test. A paper indicator is attached to the proximal end of the active guidewire prototype for the torque transmission test and the tip curve itself was used as the distal end indicator. The active guidewire prototype was manipulated by the rubber coated actuators of the dedicated trackability test setup.

Figure S2. Active guidewire prototype has a comparable stiffness with the commercially available guidewire and can completely straighten out a metal braided pigtail catheter.

NIHMS1834438-supplement-supinfo.docx^{(658.3KB, docx)}

vS3

Video S3. In-vivo real-time MRI of the active guidewire prototype in tandem with a balloon catheter in the inferior vena ceva. The active guidewire signal is not affected by the presence of the tandem catheter.

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vS1

Video S1. In-vitro real-time MRI of the active guidewire prototype, alone in a 3D printed aortic arch model. The continuous tip signal of the curved distal tip provides information about the orientation of the active guidewire.

Download video file^{(708.4KB, avi)}

vS4

Video S4. In-vivo real-time MRI of the active guidewire prototype, alone in the aortic arch. Full shaft and continuous tip visibility of the active guidewire is not affected by the changing orientation.

Download video file^{(597.8KB, avi)}

ACKNOWLEDGEMENT

We thank Katherine Lucas and Victoria Haley for their invaluable support during the in-vivo animal experiments. We also thank Transmural Systems for their industrial collaboration. Research reported in this publication was supported by National Heart Lung and Blood Division of Intramural Research award Z01-HL006041 (to RJL). The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institutes of Health. We would like to acknowledge the assistance of Siemens Healthineers in the modification of the MRI system for operation at 0.55T and for providing the MonteCarlo prototype under an existing cooperative research agreement (CRADA) between NHLBI and Siemens Healthineers.

REFERENCES

1.Kipling M, Mohammed A, Medding RN. Guidewires in clinical practice: applications and troubleshooting. Expert Rev Med Devices. 2009;6(2):187–195. [DOI] [PubMed] [Google Scholar]
2.Nageotte SJ, Lederman RJ, Ratnayaka K. MRI Catheterization: Ready for Broad Adoption. Pediatr Cardiol. 2020;41(3):503–513. [DOI] [PMC free article] [PubMed] [Google Scholar]
3.Campbell-Washburn AE, Tavallaei MA, Pop M, et al. Real-time MRI guidance of cardiac interventions. J Magn Reson Imaging. 2017;46(4):935–950. [DOI] [PMC free article] [PubMed] [Google Scholar]
4.Kolandaivelu A, Bruce CG, Ramasawmy R, et al. Native contrast visualization and tissue characterization of myocardial radiofrequency ablation and acetic acid chemoablation lesions at 0.55 T. J Cardiovasc Magn Reson. 2021;23(1):50. [DOI] [PMC free article] [PubMed] [Google Scholar]
5.Susil RC, Yeung CJ, Atalar E. Intravascular extended sensitivity (IVES) MRI antennas. Magn Reson Med. 2003;50(2):383–390. [DOI] [PubMed] [Google Scholar]
6.Erturk MA, El-Sharkawy AMM, Bottomley PA. Interventional loopless antenna at 7 T. Magnetic Resonance in Medicine. 2012;68(3):980–988. [DOI] [PMC free article] [PubMed] [Google Scholar]
7.Rogers T, Lederman RJ. Interventional CMR: Clinical applications and future directions. Curr Cardiol Rep. 2015;17(5):31. [DOI] [PMC free article] [PubMed] [Google Scholar]
8.Settecase F, Martin AJ, Lillaney P, Losey A, Hetts SW. Magnetic Resonance-Guided Passive Catheter Tracking for Endovascular Therapy. Magn Reson Imaging Clin N Am. 2015;23(4):591–605. [DOI] [PMC free article] [PubMed] [Google Scholar]
9.Baysoy E, Yildirim DK, Ozsoy C, Mutlu S, Kocaturk O. Thin film based semi-active resonant marker design for low profile interventional cardiovascular MRI devices. MAGMA. 2017;30(1):93–101. [DOI] [PubMed] [Google Scholar]
10.Quick HH, Zenge MO, Kuehl H, et al. Interventional magnetic resonance angiography with no strings attached: wireless active catheter visualization. Magn Reson Med. 2005;53(2):446–455. [DOI] [PubMed] [Google Scholar]
11.Wang W Magnetic Resonance-guided Active Catheter Tracking. Magn Reson Imaging Clin N Am. 2015;23(4):579–589. [DOI] [PMC free article] [PubMed] [Google Scholar]
12.Yildirim DK, Bruce C, Uzun D, et al. A 20-gauge active needle design with thin-film printed circuitry for interventional MRI at 0.55T. Magn Reson Med. 2021. [DOI] [PMC free article] [PubMed] [Google Scholar]
13.Guttman MA, Ozturk C, Raval AN, et al. Interventional cardiovascular procedures guided by real-time MR imaging: an interactive interface using multiple slices, adaptive projection modes and live 3D renderings. J Magn Reson Imaging. 2007;26(6):1429–1435. [DOI] [PMC free article] [PubMed] [Google Scholar]
14.Yeung CJ, Susil RC, Atalar E. RF safety of wires in interventional MRI: using a safety index. Magn Reson Med. 2002;47(1):187–193. [DOI] [PubMed] [Google Scholar]
15.Yeung CJ, Karmarkar P, McVeigh ER. Minimizing RF heating of conducting wires in MRI. Magn Reson Med. 2007;58(5):1028–1034. [DOI] [PubMed] [Google Scholar]
16.Knight DS, Kotecha T, Martinez-Naharro A, et al. Cardiovascular magnetic resonance-guided right heart catheterization in a conventional CMR environment - predictors of procedure success and duration in pulmonary artery hypertension. J Cardiovasc Magn Reson. 2019;21(1):57. [DOI] [PMC free article] [PubMed] [Google Scholar]
17.Basar B, Campbell-Washburn AE, Rogers T, et al. Stiffness-matched segmented metallic guidewire for interventional cardiovascular MRI. Journal of Cardiovascular Magnetic Resonance. 2015;17(1):P414. [DOI] [PMC free article] [PubMed] [Google Scholar]
18.Yildirim KD, Basar B, Campbell-Washburn AE, Herzka DA, Kocaturk O, Lederman RJ. A cardiovascular magnetic resonance (CMR) safe metal braided catheter design for interventional CMR at 1.5 T: freedom from radiofrequency induced heating and preserved mechanical performance. J Cardiovasc Magn Reson. 2019;21(1):16. [DOI] [PMC free article] [PubMed] [Google Scholar]
19.Qiu B, Karmarkar P, Brushett C, et al. Development of a 0.014-inch magnetic resonance imaging guidewire. Magn Reson Med. 2005;53(4):986–990. [DOI] [PubMed] [Google Scholar]
20.Ocali O, Atalar E. Intravascular magnetic resonance imaging using a loopless catheter antenna. Magn Reson Med. 1997;37(1):112–118. [DOI] [PubMed] [Google Scholar]
21.Serfaty JM, Yang X, Aksit P, Quick HH, Solaiyappan M, Atalar E. Toward MRI-guided coronary catheterization: visualization of guiding catheters, guidewires, and anatomy in real time. J Magn Reson Imaging. 2000;12(4):590–594. [DOI] [PubMed] [Google Scholar]
22.El-Sharkawy AMM, Qian D, Bottomley PA. The performance of interventional loopless MRI antennae at higher magnetic field strengths. Medical Physics. 2008;35(5):1995–2006. [DOI] [PMC free article] [PubMed] [Google Scholar]
23.Peterson DM, Beck BL, Duensing CR, Fitzsimmons JR. Common mode signal rejection methods for MRI: Reduction of cable shield currents for high static magnetic field systems. Concept Magn Reson B. 2003;19b(1):1–8. [Google Scholar]
24.Qian D, El-Sharkawy AM, Atalar E, Bottomley PA. Interventional MRI: tapering improves the distal sensitivity of the loopless antenna. Magn Reson Med. 2010;63(3):797–802. [DOI] [PMC free article] [PubMed] [Google Scholar]
25.Kocaturk O, Kim AH, Saikus CE, et al. Active two-channel 0.035” guidewire for interventional cardiovascular MRI. J Magn Reson Imaging. 2009;30(2):461–465. [DOI] [PMC free article] [PubMed] [Google Scholar]
26.Sonmez M, Saikus CE, Bell JA, et al. MRI active guidewire with an embedded temperature probe and providing a distinct tip signal to enhance clinical safety. J Cardiovasc Magn Reson. 2012;14:38. [DOI] [PMC free article] [PubMed] [Google Scholar]
27.Karmarkar PV, Kraitchman DL, Izbudak I, et al. MR-trackable intramyocardial injection catheter. Magn Reson Med. 2004;51(6):1163–1172. [DOI] [PubMed] [Google Scholar]
28.Etezadi-Amoli M, Stang P, Kerr A, Pauly J, Scott G. Interventional device visualization with toroidal transceiver and optically coupled current sensor for radiofrequency safety monitoring. Magn Reson Med. 2015;73(3):1315–1327. [DOI] [PMC free article] [PubMed] [Google Scholar]
29.Starmer CF, McIntosh HD, Whalen RE. Electrical hazards and cardiovascular function. N Engl J Med. 1971;284(4):181–186. [DOI] [PubMed] [Google Scholar]
30.Monies-Chass I, Vilensky A, Mordechowitz B, Birkhahn J. Hidden risk in operating room. Micro-shock. Acta Anaesthesiol Belg. 1986;37(1):39–44. [PubMed] [Google Scholar]
31.Yeung CJ, Atalar E. RF transmit power limit for the barewire loopless catheter antenna. J Magn Reson Imaging. 2000;12(1):86–91. [DOI] [PubMed] [Google Scholar]
32.Bottomley PA, Edelstein WA. Power deposition in whole-body NMR imaging. Med Phys. 1981;8(4):510–512. [DOI] [PubMed] [Google Scholar]
33.Campbell-Washburn AE, Ramasawmy R, Restivo MC, et al. Opportunities in Interventional and Diagnostic Imaging by Using High-Performance Low-Field-Strength MRI. Radiology. 2019;293(2):384–393. [DOI] [PMC free article] [PubMed] [Google Scholar]
34.Atlamazoglou PE, Uzunoglu NK. A Galerkin moment method for the analysis of an insulated antenna in a dissipative dielectric medium. IEEE Transactions on Microwave Theory and Techniques. 1998;46(7):988–996. [Google Scholar]
35.Fournier M, Pomerleau A. Experimental study of an inductively loaded short monopole antenna. IEEE Transactions on Vehicular Technology. 1978;27:1–6. [Google Scholar]
36.Alipour A, Meyer ES, Dumoulin CL, et al. MRI Conditional Actively Tracked Metallic Electrophysiology Catheters and Guidewires With Miniature Tethered Radio-Frequency Traps: Theory, Design, and Validation. IEEE Trans Biomed Eng. 2020;67(6):1616–1627. [DOI] [PMC free article] [PubMed] [Google Scholar]
37.Ladd ME, Erhart P, Debatin JF, et al. Guidewire antennas for MR fluoroscopy. Magn Reson Med. 1997;37(6):891–897. [DOI] [PubMed] [Google Scholar]
38.Gruber B, Froeling M, Leiner T, Klomp DWJ. RF coils: A practical guide for nonphysicists. J Magn Reson Imaging. 2018. [DOI] [PMC free article] [PubMed] [Google Scholar]
39.Yaras YS, Yildirim DK, Herzka DA, et al. Real-time device tracking under MRI using an acousto-optic active marker. Magnetic Resonance in Medicine. 2020. [DOI] [PMC free article] [PubMed] [Google Scholar]
40.Yaras YS, Yildirim DK, Kocaturk O, Degertekin FL. Sensitivity and phase response of FBG based acousto-optic sensors for real-time MRI applications. OSA Contin. 2020;3(3):447–458. [DOI] [PMC free article] [PubMed] [Google Scholar]
41.Robson PM, Grant AK, Madhuranthakam AJ, Lattanzi R, Sodickson DK, McKenzie CA. Comprehensive quantification of signal-to-noise ratio and g-factor for image-based and k-space-based parallel imaging reconstructions. Magn Reson Med. 2008;60(4):895–907. [DOI] [PMC free article] [PubMed] [Google Scholar]
42.Tokaya JP, Raaijmakers AJE, Luijten PR, Bakker JF, van den Berg CAT. MRI-based transfer function determination for the assessment of implant safety. Magn Reson Med. 2017;78(6):2449–2459. [DOI] [PubMed] [Google Scholar]
43.Celik H, Uluturk A, Tali T, Atalar E. A catheter tracking method using reverse polarization for MR-guided interventions. Magnetic Resonance in Medicine. 2007;58(6):1224–1231. [DOI] [PubMed] [Google Scholar]
44.Vazquez R, Larson DF. Plasma protein denaturation with graded heat exposure. Perfusion. 2013;28(6):557–559. [DOI] [PubMed] [Google Scholar]
45.Park SM, Kamondetdacha R, Nyenhuis JA. Calculation of MRI-induced heating of an implanted medical lead wire with an electric field transfer function. J Magn Reson Imaging. 2007;26(5):1278–1285. [DOI] [PubMed] [Google Scholar]
46.ISO/TS 10974:2018. Assessment of the safety of magnetic resonance imaging for patients with an active implantable medical device. Geneva, Switzerland. [Google Scholar]
47.Rogers T, Ratnayaka K, Khan JM, et al. CMR fluoroscopy right heart catheterization for cardiac output and pulmonary vascular resistance: results in 102 patients. J Cardiovasc Magn Reson. 2017;19(1):54. [DOI] [PMC free article] [PubMed] [Google Scholar]
48.Thompson SM, Gorny KR, Koepsel EMK, et al. Body Interventional MRI for Diagnostic and Interventional Radiologists: Current Practice and Future Prospects. Radiographics. 2021;41(6):1785–1801. [DOI] [PubMed] [Google Scholar]
49.Mukherjee RK, Chubb H, Roujol S, Razavi R, O’Neill MD. Advances in Real-Time MRI-Guided Electrophysiology. Curr Cardiovasc Imaging Rep. 2019;12(2):6–6. [DOI] [PMC free article] [PubMed] [Google Scholar]

Associated Data

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

Supplementary Materials

vS2

Download video file^{(418.2KB, avi)}

supinfo

Figure S2. Active guidewire prototype has a comparable stiffness with the commercially available guidewire and can completely straighten out a metal braided pigtail catheter.

NIHMS1834438-supplement-supinfo.docx^{(658.3KB, docx)}

vS3

Download video file^{(374.3KB, avi)}

vS1

Download video file^{(708.4KB, avi)}

vS4

Download video file^{(597.8KB, avi)}

[R1] 1.Kipling M, Mohammed A, Medding RN. Guidewires in clinical practice: applications and troubleshooting. Expert Rev Med Devices. 2009;6(2):187–195. [DOI] [PubMed] [Google Scholar]

[R2] 2.Nageotte SJ, Lederman RJ, Ratnayaka K. MRI Catheterization: Ready for Broad Adoption. Pediatr Cardiol. 2020;41(3):503–513. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R3] 3.Campbell-Washburn AE, Tavallaei MA, Pop M, et al. Real-time MRI guidance of cardiac interventions. J Magn Reson Imaging. 2017;46(4):935–950. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R4] 4.Kolandaivelu A, Bruce CG, Ramasawmy R, et al. Native contrast visualization and tissue characterization of myocardial radiofrequency ablation and acetic acid chemoablation lesions at 0.55 T. J Cardiovasc Magn Reson. 2021;23(1):50. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R5] 5.Susil RC, Yeung CJ, Atalar E. Intravascular extended sensitivity (IVES) MRI antennas. Magn Reson Med. 2003;50(2):383–390. [DOI] [PubMed] [Google Scholar]

[R6] 6.Erturk MA, El-Sharkawy AMM, Bottomley PA. Interventional loopless antenna at 7 T. Magnetic Resonance in Medicine. 2012;68(3):980–988. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R7] 7.Rogers T, Lederman RJ. Interventional CMR: Clinical applications and future directions. Curr Cardiol Rep. 2015;17(5):31. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R8] 8.Settecase F, Martin AJ, Lillaney P, Losey A, Hetts SW. Magnetic Resonance-Guided Passive Catheter Tracking for Endovascular Therapy. Magn Reson Imaging Clin N Am. 2015;23(4):591–605. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R9] 9.Baysoy E, Yildirim DK, Ozsoy C, Mutlu S, Kocaturk O. Thin film based semi-active resonant marker design for low profile interventional cardiovascular MRI devices. MAGMA. 2017;30(1):93–101. [DOI] [PubMed] [Google Scholar]

[R10] 10.Quick HH, Zenge MO, Kuehl H, et al. Interventional magnetic resonance angiography with no strings attached: wireless active catheter visualization. Magn Reson Med. 2005;53(2):446–455. [DOI] [PubMed] [Google Scholar]

[R11] 11.Wang W Magnetic Resonance-guided Active Catheter Tracking. Magn Reson Imaging Clin N Am. 2015;23(4):579–589. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R12] 12.Yildirim DK, Bruce C, Uzun D, et al. A 20-gauge active needle design with thin-film printed circuitry for interventional MRI at 0.55T. Magn Reson Med. 2021. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R13] 13.Guttman MA, Ozturk C, Raval AN, et al. Interventional cardiovascular procedures guided by real-time MR imaging: an interactive interface using multiple slices, adaptive projection modes and live 3D renderings. J Magn Reson Imaging. 2007;26(6):1429–1435. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R14] 14.Yeung CJ, Susil RC, Atalar E. RF safety of wires in interventional MRI: using a safety index. Magn Reson Med. 2002;47(1):187–193. [DOI] [PubMed] [Google Scholar]

[R15] 15.Yeung CJ, Karmarkar P, McVeigh ER. Minimizing RF heating of conducting wires in MRI. Magn Reson Med. 2007;58(5):1028–1034. [DOI] [PubMed] [Google Scholar]

[R16] 16.Knight DS, Kotecha T, Martinez-Naharro A, et al. Cardiovascular magnetic resonance-guided right heart catheterization in a conventional CMR environment - predictors of procedure success and duration in pulmonary artery hypertension. J Cardiovasc Magn Reson. 2019;21(1):57. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R17] 17.Basar B, Campbell-Washburn AE, Rogers T, et al. Stiffness-matched segmented metallic guidewire for interventional cardiovascular MRI. Journal of Cardiovascular Magnetic Resonance. 2015;17(1):P414. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R18] 18.Yildirim KD, Basar B, Campbell-Washburn AE, Herzka DA, Kocaturk O, Lederman RJ. A cardiovascular magnetic resonance (CMR) safe metal braided catheter design for interventional CMR at 1.5 T: freedom from radiofrequency induced heating and preserved mechanical performance. J Cardiovasc Magn Reson. 2019;21(1):16. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R19] 19.Qiu B, Karmarkar P, Brushett C, et al. Development of a 0.014-inch magnetic resonance imaging guidewire. Magn Reson Med. 2005;53(4):986–990. [DOI] [PubMed] [Google Scholar]

[R20] 20.Ocali O, Atalar E. Intravascular magnetic resonance imaging using a loopless catheter antenna. Magn Reson Med. 1997;37(1):112–118. [DOI] [PubMed] [Google Scholar]

[R21] 21.Serfaty JM, Yang X, Aksit P, Quick HH, Solaiyappan M, Atalar E. Toward MRI-guided coronary catheterization: visualization of guiding catheters, guidewires, and anatomy in real time. J Magn Reson Imaging. 2000;12(4):590–594. [DOI] [PubMed] [Google Scholar]

[R22] 22.El-Sharkawy AMM, Qian D, Bottomley PA. The performance of interventional loopless MRI antennae at higher magnetic field strengths. Medical Physics. 2008;35(5):1995–2006. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R23] 23.Peterson DM, Beck BL, Duensing CR, Fitzsimmons JR. Common mode signal rejection methods for MRI: Reduction of cable shield currents for high static magnetic field systems. Concept Magn Reson B. 2003;19b(1):1–8. [Google Scholar]

[R24] 24.Qian D, El-Sharkawy AM, Atalar E, Bottomley PA. Interventional MRI: tapering improves the distal sensitivity of the loopless antenna. Magn Reson Med. 2010;63(3):797–802. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R25] 25.Kocaturk O, Kim AH, Saikus CE, et al. Active two-channel 0.035” guidewire for interventional cardiovascular MRI. J Magn Reson Imaging. 2009;30(2):461–465. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R26] 26.Sonmez M, Saikus CE, Bell JA, et al. MRI active guidewire with an embedded temperature probe and providing a distinct tip signal to enhance clinical safety. J Cardiovasc Magn Reson. 2012;14:38. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R27] 27.Karmarkar PV, Kraitchman DL, Izbudak I, et al. MR-trackable intramyocardial injection catheter. Magn Reson Med. 2004;51(6):1163–1172. [DOI] [PubMed] [Google Scholar]

[R28] 28.Etezadi-Amoli M, Stang P, Kerr A, Pauly J, Scott G. Interventional device visualization with toroidal transceiver and optically coupled current sensor for radiofrequency safety monitoring. Magn Reson Med. 2015;73(3):1315–1327. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R29] 29.Starmer CF, McIntosh HD, Whalen RE. Electrical hazards and cardiovascular function. N Engl J Med. 1971;284(4):181–186. [DOI] [PubMed] [Google Scholar]

[R30] 30.Monies-Chass I, Vilensky A, Mordechowitz B, Birkhahn J. Hidden risk in operating room. Micro-shock. Acta Anaesthesiol Belg. 1986;37(1):39–44. [PubMed] [Google Scholar]

[R31] 31.Yeung CJ, Atalar E. RF transmit power limit for the barewire loopless catheter antenna. J Magn Reson Imaging. 2000;12(1):86–91. [DOI] [PubMed] [Google Scholar]

[R32] 32.Bottomley PA, Edelstein WA. Power deposition in whole-body NMR imaging. Med Phys. 1981;8(4):510–512. [DOI] [PubMed] [Google Scholar]

[R33] 33.Campbell-Washburn AE, Ramasawmy R, Restivo MC, et al. Opportunities in Interventional and Diagnostic Imaging by Using High-Performance Low-Field-Strength MRI. Radiology. 2019;293(2):384–393. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R34] 34.Atlamazoglou PE, Uzunoglu NK. A Galerkin moment method for the analysis of an insulated antenna in a dissipative dielectric medium. IEEE Transactions on Microwave Theory and Techniques. 1998;46(7):988–996. [Google Scholar]

[R35] 35.Fournier M, Pomerleau A. Experimental study of an inductively loaded short monopole antenna. IEEE Transactions on Vehicular Technology. 1978;27:1–6. [Google Scholar]

[R36] 36.Alipour A, Meyer ES, Dumoulin CL, et al. MRI Conditional Actively Tracked Metallic Electrophysiology Catheters and Guidewires With Miniature Tethered Radio-Frequency Traps: Theory, Design, and Validation. IEEE Trans Biomed Eng. 2020;67(6):1616–1627. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R37] 37.Ladd ME, Erhart P, Debatin JF, et al. Guidewire antennas for MR fluoroscopy. Magn Reson Med. 1997;37(6):891–897. [DOI] [PubMed] [Google Scholar]

[R38] 38.Gruber B, Froeling M, Leiner T, Klomp DWJ. RF coils: A practical guide for nonphysicists. J Magn Reson Imaging. 2018. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R39] 39.Yaras YS, Yildirim DK, Herzka DA, et al. Real-time device tracking under MRI using an acousto-optic active marker. Magnetic Resonance in Medicine. 2020. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R40] 40.Yaras YS, Yildirim DK, Kocaturk O, Degertekin FL. Sensitivity and phase response of FBG based acousto-optic sensors for real-time MRI applications. OSA Contin. 2020;3(3):447–458. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R41] 41.Robson PM, Grant AK, Madhuranthakam AJ, Lattanzi R, Sodickson DK, McKenzie CA. Comprehensive quantification of signal-to-noise ratio and g-factor for image-based and k-space-based parallel imaging reconstructions. Magn Reson Med. 2008;60(4):895–907. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R42] 42.Tokaya JP, Raaijmakers AJE, Luijten PR, Bakker JF, van den Berg CAT. MRI-based transfer function determination for the assessment of implant safety. Magn Reson Med. 2017;78(6):2449–2459. [DOI] [PubMed] [Google Scholar]

[R43] 43.Celik H, Uluturk A, Tali T, Atalar E. A catheter tracking method using reverse polarization for MR-guided interventions. Magnetic Resonance in Medicine. 2007;58(6):1224–1231. [DOI] [PubMed] [Google Scholar]

[R44] 44.Vazquez R, Larson DF. Plasma protein denaturation with graded heat exposure. Perfusion. 2013;28(6):557–559. [DOI] [PubMed] [Google Scholar]

[R45] 45.Park SM, Kamondetdacha R, Nyenhuis JA. Calculation of MRI-induced heating of an implanted medical lead wire with an electric field transfer function. J Magn Reson Imaging. 2007;26(5):1278–1285. [DOI] [PubMed] [Google Scholar]

[R46] 46.ISO/TS 10974:2018. Assessment of the safety of magnetic resonance imaging for patients with an active implantable medical device. Geneva, Switzerland. [Google Scholar]

[R47] 47.Rogers T, Ratnayaka K, Khan JM, et al. CMR fluoroscopy right heart catheterization for cardiac output and pulmonary vascular resistance: results in 102 patients. J Cardiovasc Magn Reson. 2017;19(1):54. [DOI] [PMC free article] [PubMed] [Google Scholar]

[R48] 48.Thompson SM, Gorny KR, Koepsel EMK, et al. Body Interventional MRI for Diagnostic and Interventional Radiologists: Current Practice and Future Prospects. Radiographics. 2021;41(6):1785–1801. [DOI] [PubMed] [Google Scholar]

[R49] 49.Mukherjee RK, Chubb H, Roujol S, Razavi R, O’Neill MD. Advances in Real-Time MRI-Guided Electrophysiology. Curr Cardiovasc Imaging Rep. 2019;12(2):6–6. [DOI] [PMC free article] [PubMed] [Google Scholar]

PERMALINK

An Interventional MRI guidewire combining profile and tip conspicuity for catheterization at 0.55T

Dursun Korel Yildirim, PhD

Dogangun Uzun, MSc

Christopher G Bruce, MB ChB

Jaffar M Khan, BM BCh PhD

Toby Rogers, BM BCh PhD

William H Schenke, BS

Rajiv Ramasawmy, PhD

Adrienne Campbell-Washburn, PhD

Daniel Herzka, PhD

Robert J Lederman, MD

Ozgur Kocaturk, PhD

Abstract

Purpose:

Methods:

Results:

Conclusion:

1. Introduction

2. Methods

2.1. Electromagnetic simulations

2.2. Active iMRI guidewire design

Figure 1.

2.3. In-vitro characterization

Table 1.

2.4. In-vivo characterization

3. Results

Figure 2.

Figure 3.

Figure 4.

Figure 5.

Figure 6.

Figure 7.

Figure 8.

4. Discussion and Conclusion

Supplementary Material

ACKNOWLEDGEMENT

REFERENCES

Associated Data

Supplementary Materials

ACTIONS

PERMALINK

RESOURCES

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