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Influence of Body Coil Drive Mode on the MR Safety of Interventional Devices1Division of Medical Physics, Department of Radiology, Medical Center - University of Freiburg, University of Freiburg, Freiburg, Germany, 2Interventional MR Clinical R&D Institute, Ankara University, Ankara, Turkey, 3National MR Research Center (UMRAM), Bilkent University, Ankara, Turkey
Synopsis
Keywords: Safety, Safety, Low Field, RF-induced heating, Interventional
Motivation: RF-induced heating of interventional devices is reduced at lower magnetic fields, but dangerous heating conditions can still occur during MR-guided interventions.
Goal(s): To investigate the effect of body coil drive mode on RF-induced heating of interventional devices at 0.55T and 1.5T.
Approach: Numerical simulations, E-field, transfer function and temperature measurements are used to evaluate the dependency of RF-induced heating of commonly used interventional devices on the phase and amplitude settings of body coil ports.
Results: RF-induced heating can be reduced using individually adapted body coil drive modes.
Impact: Dual-drive body coil technology can be used to generate implant-friendly electromagnetic fields, which can enable safe use of clinical intravascular devices during MR-guided interventions at 0.55T and 1.5T.
Introduction
RF-induced heating of implants and devices in the MRI is a result of the coupling of external electromagnetic (EM) transmit fields with elongated metallic structures[1], [2]. The heating depends on the incident E field[3], EM properties of device and surrounding tissue, input impedance[4], [5] and insertion length[6], [7]. Low-field MRI systems (e.g., at 0.55T) are expected to cause less heating than 1.5T or 3T systems[8]; however, since the heating depends on the target organ and patient size[7], additional measures are needed to reduce heating of devices during interventions. To use existing interventional devices, the MRI must be modified to be ‘device-safe’. Device-safe MRI methods are mostly based on modification of the transmit (Tx) coil’s EM fields[9]–[16] or RF pulses[17] using parallel Tx or create a zero-E field plane of a linearly polarized birdcage coil[9], [18]–[20]. Other methods include design of low-specific-absorption-rate (SAR) protocols[21] and low field systems [8]. A numerical analysis of RF exposure on implant heating was also investigated in[22]. As 2-channel parallel transmit technologies have become available even for the birdcage body coils in conventional MR systems at 1.5T and 3T, in this study we investigate their effect on RF-induced heating of interventional devices at 0.55T and 1.5T.Methods
General workflow for the numerical simulation and measurements is presented in Fig. 1.Test setup: A modified Tier-3 approach was applied [7], [23], where high resolution E-field maps were measured for hot-spot detection during RF excitation at fLarmor(0.55T)=23.66MHz, fLarmor(1.5T)=63.85MHz, and TF measurements were performed with an electro-optic sensor (EOS) [24] at the hot spot [6]. TFs were calibrated using various dipole excitation patterns [23]. An RF test-bench was constructed for temperature measurements in 0.55T (Ø=790mm, Length=500mm) and 1.5T (Ø=790mm, Length=500mm) quadrature body coils. In addition to the circularly polarized (CP) mode, two elliptically polarized (EP) modes were generated by a variable phase shift between the feed ports.
Device selection: Realistic device trajectories for coronary artery (CA), hepatic artery (HA), and intracranial (IC) catheterizations were obtained [7]. Temperature measurements were performed using a fiber-optic temperature sensor (FOTEMP, Optocon) in a gel phantom (30g/L-HEC&1.55g/L-NaCl[25], εr = 82.5/81.1, σ = 0.63/0.64 S/m @23.7/63.9MHz). For CA, HA, and IC devices, insertion lengths of [80,85,90]cm, [60,65,70]cm, and [110,120,130]cm were tested. For CA and HA engagement, two guidewires(GW), a guiding-catheter(GC), and a microcatheter-(uC) were selected. In addition to a GW, an intrasaccular flow disruptor for aneurysm cavities was evaluated [26], [27].
Numerical simulations: Body coils were simulated using FDTD (Sim4Life, ZMT, Zürich), and tangential E fields along the trajectories were extracted for various relative amplitude&phase settings of the body coil feed ports. EP modes with minimum B1+ homogeneity (i.e., maximum deviation of 20%) were identified. The RF input power was scaled to achieve mean B1+ of 11.75µT within the center of the phantom.
Results
Consistent with previous findings [23], hotspot locations were identical at 0.55T and 1.5T. Simulated E fields are shown in Fig. 2 for the CP and two EP modes for 0.55T & 1.5T. E field phases and amplitudes along the device trajectories are plotted for the drive modes with minimum B1+ homogeneity in Fig. 3 together with estimated SAR in comparison to the CP mode. Although EP modes exists with an up to 4-fold lower SAR than the CP mode, also modes with 2.5-fold higher SAR have been observed (Fig. 3c). Temperature measurements and TF-based temperature estimations at 0.55T (Fig. 4) and at 1.5T (Fig. 5) show a strong dependency on the insertion length of the device. IC catheterization has higher average temperature increases (ΔTmax = 4.4K) than CA (ΔTmax = 0.8K) and HA (ΔTmax = 0.7K) at 0.55T, as during an IC the IL approaches the resonant length. At 0.55T, temperature rises were consistently lower than 1.5T except for the HA case. TF-based temperature rise and RF-induced heating measurements were comparable, except for 1.5T, where up to 100% differences are observed. Differences between CP and EP modes were higher in simulations than in the measurements, as only the phase difference was adjusted in the measurements.Discussion
Using optimized drive modes, RF-induced device heating can be significantly reduced during an intervention. Parallel transmit options for body coils are valuable also at low fields, where B1 shimming is usually not required. Note, that EP modes selected in FDTD simulations were not identical, since EP mode for temperature measurements were empirically selected by adjusting the phase of one port using different cable lengths. Inconsistencies between simulation and measurement at 1.5T are due to inaccuracies in our model, which will be improved in future versions.Acknowledgements
Technical support by M. Malzacher, J. Nistler, M. Leghissa, B. Kraus of Siemens Healthineers; and grant support by BMBF (13GW0356B), and DFG under CRC 1425 (Project P15) are gratefully acknowledged.
The mention of commercial products, their sources, or their use in connection with material reported herein is not to be construed as either an actual or implied endorsement of such products by the authors or their institutions.
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