This study shows that no clinically significant RF-induced heating occurs during real-time SSFP MR imaging of MR-assisted catheter tip steering for interventional MRI. The maximum increase in temperature observed after 15 minutes of continuous SSFP imaging was 0.35 °C at 15 cm from magnet isocenter, which is well-below the 4 °C increase that can cause irreversible tissue damage16,17
. As expected, temperature increases around the coils were directly proportional to the proximity of the catheter-coil tip to the magnet bore wall, as well as to the power of the MR pulse sequence (varied here by changing the RF flip angle).
Temperature rises caused by RF heating effects on the laser lithographed microcoil-catheter construct in air or saline also did not reach a significant level, with maximum tempurature rise reaching 1.84°C in the absence of saline coolant flow and with the microcoil-catheter tip positioned as far offset from isocenter as possible.
Following SSFP imaging after in vivo catheter placement, delayed contrast enhanced MRI revealed no evidence of vascular injury under normal flow and zero flow conditions. Histopathological evaluation of these carotid arteries confirmed the absence of vascular injury.
These findings indicate that RF heating previously observed in other types of conductive wires in MRI also occurs with microcoil-tipped endovascular catheters, however, to a lesser degree. The mechanism by which this RF induced heating occurs is likely due to the antenna effect, whereby the wire couples with the time-varying electric fields produced by the RF pulses. The amount of RF heating caused by conductive wires has been shown to be high in previous studies, although results have been variable and inconsistent. In vitro
temperature increases in the range of 18 to 48°C have been reported, in experiments involving a standard conducting guidewire, coaxial cable, and miniature tracking coil6,18
, levels that are incompatible with patient safety5-7
. These studies, however, are concerned with single wires whose distal end is an open circuit. In the case of microcoils used for catheter tip deflection, current carrying wires transit the length of the catheter lumen in pairs. Each pair is connected at the distal end to a conducting microcoil with low inductance compared to the coupled lead wires. This configuration is essentially a long, narrow loop antenna. Such an antenna is inefficient because the area enclosed by the loop is small. Indeed, the induced currents were found to be small so that significant cable resonances were not produced, or were damped out significantly at the wire length and field strength tested. There is no certainty that this solution will work at higher frequencies (such as 3T), or with different cable lengths. The results shown in , with differential heating in air versus saline, suggest that resonances may play a small role under the conditions tested, although the larger heat capacity of saline as opposed to air may also explain these differences. In addition, a heating study involving intravascular imaging coils, which consisted of a loop with a length of 40 mm and a width of 6 mm, demonstrated much lower heating19
, probably due to the rounded end of the coil, which prevents a high concentration of electric fields. The coils used in MRI-assisted catheter steering are similar in size and shape to intravascular imaging coils.
Despite these promising results, further steps can and should be taken to make MRI-assisted catheter tip steering safer for clinical applications. Resistive heating could be reduced by a number of modifications to the catheter-coil construct. Increasing the efficiency of the catheter, i.e., increasing the amount of deflection obtained per unit of current, may involve the use of highly flexible materials (lower elastic modulus) for catheter construction. Since power dissipation is also proportional to the resistance, conducting wires of smaller resistance per unit length, such as thicker diameter wires, or gold wires, could also be used to reduce ohmic heating. Catheter bending stiffness can also be decreased by reducing the diameter and wall thickness of the catheter, thus reducing the area moment of inertia. More coil turns increase the magnetic moment but the wire width must also be reduced to accommodate more turns in the same overall space. Both the increase in turns/wire length and the reduction in wire width increase overall resistance and heating. Also, since the magnetic torque produced by this mechanism is the cross-product of the magnetic moment and the main scanner magnetic field, doubling the magnet strength to 3T was shown to double the effective catheter deflection2
. This advantage, however, needs to be weighed carefully against the higher specific absorption rates (SAR) that are associated with higher field scanners since higher field strength may result in relatively more RF heating. The prototypical catheter-coil design used here involves winding of a magnet wire coil around the outside of the catheter. Catheter extrusion with the coils built into the walls of the catheter would increase insulation from the overlying blood and vessel walls.
The applied direct currents necessary to achieve catheter tip deflections result in resistive dissipation and more significant tissue heating than RF heating20
. As discussed in the accompanying manuscript, infusion of normal saline at clinically achievable flow rates within the lumen of the catheter during application of current is successful in transferring heat produced by applied currents both in vitro
and in vivo
. Furthermore, the transfer of unwanted heat to coolant flowing in the catheter lumen is facilitated by using a high heat conductivity material as the catheter tip substrate on the inner surface of the coils, while using an insulating material on the outside of the coils to protect the blood and vessel walls. Using a clinically achievable saline coolant flow rate (10 ml/kg/hr), preliminary in vivo
testing demonstrated no thermal injury to vessel walls from resistive or RF heating at physiologic and zero arterial flow. Although tissue temperature was not measured directly, the most important parameter in ultimate clinical application is the likelihood of thermal injury to tissues by the catheter either from resistive heating while current is applied for catheter navigation or RF heating while the catheter is being imaged in the MR scanner. At the resonant lengths between 102 cm and 92 cm from the femoral artery access site, there was no evidence for the massive RF-induced heating that has been reported in other guidewire constructs 4,6,11-13
Although we have used saline coolant flowing through the catheter lumen to decrease temperature increases related to resistive heating, this approach could also be applied in the settings where RF heating is more significant. Similarly, the flowing saline coolant approach may provide device cooling at different cable lengths and at different MR field strengths than those tested, thus potentially providing a simple solution to heat dissipation under a variety of interventional MRI conditions. Several authors have investigated the use of coaxial chokes3
, transmission lines21
, and other devices along the catheter shaft, to reduce RF heating. The use of saline flowing through the lumen can be easily applied to almost any catheter system with a simple design modification. This approach would also be compatible with a wide variety of cable or catheter lengths and field strengths, without requiring complex design adaptations for each new device.
Although this is a promising result, further studies are necessary to examine the effects of varying resonant lengths of wire 22
to better characterize RF safety.