A prototype RF dosimeter for independent measurement of the average specific absorption rate (SAR) during MRI.

THERE ARE CURRENTLY over 20,000 MRI scanners performing over 60 million studies world-wide (1). About half of these scans are performed in the United States, where the number of scans has been doubling about every three years for the past decade. In the United States, over one-half of the existing scanners operate at 1.5T (2), while most luminary institutions active in medical imaging research have added 3T MRI systems, and 4T (3) and 7T (4) whole body systems are also commercially available. While MRI is not considered a significant risk (5), it is not hazard-free. The potential for radio frequency (RF) power deposition and heating in human MRI was realized as early as 1978 (6). The mechanism for heating is the induction of eddy currents in the body by the time-dependent RF magnetic field in accordance with Faraday’s Law, due to the finite conductivity of the body. Models comprised of homogeneous cylinders of tissue of measured conductivity and dielectric constant, were solved analytically for various configurations (6-8). The key results were: 1) local peak and average specific power absorption rates (SARs) in W/kg or W/cm3 can be determined from the known RF pulse width, duty cycle, flip-angle, and sample size using relatively simple formulae (6-8); 2) SAR varies approximately quadratically with MRI frequency or field strength, and with sample radius at lower frequencies (<100 MHz); and 3) peak SAR is related to the average SAR by simple numerical factors (7). That excess RF power deposition can cause heating and burns is evidenced by voluntary reports to the U.S. Food and Drug Administration (FDA) of injuries received during clinical MRI, as summarized in Table 1 for 2005 and 2006 (9). Many of these injury reports are not linked to patient proximity to leads, metal, separate receiver coils, or the magnet bore, which may pose additional risk. Table 1 Adverse Events Involving RF Burns in MRI Scanners Reported to the FDA, 2005-2006* The safety of RF exposure during clinical MRI is regulated via government and industry guidelines (10,11). RF exposure is also a factor in assessing the safety of MRI in human research overseen by Institutional Review Boards (IRBs). Consequently, accurate RF dosimetry is central to the safe operation of thousands of MRI scanners and millions of human MRI scans. Indeed, issues relating to SAR are listed among the top three unsolved problems and unmet needs by each of three study groups of the International Society of Magnetic Resonance in Medicine (ISMRM) (12), underscoring the need for providing accurate and independent methods of measuring it. In Europe, the International Electrotechnical Commission (IEC) defines the whole-body SAR as the absorbed RF power (PA in Watts) divided by patient mass (m), and the partial body SAR is calculated based on the body mass in the coil, which may be modeled by homogeneous cylinders (11). The local SAR in 1 g (9) or 10 g (11) of tissue is determined from experimentally-validated models or by experiments on phantoms (11). For homogeneous cylinders and spheres, the models yield a local SAR two to four times the average SAR (7). Heterogeneous models with quadrature excitation yield ratios of 4.5-6 for the head at 63-175 MHz (13-15), and 10-16 for male and female torsos at 1.5T and 3T (16,17). Thus, once an average SAR is determined via PA/m, a model-based local maximum can be obtained by simple multiplication by the corresponding model factor, at least for the above frequency ranges. One simple method of measuring SAR that satisfies the IEC guidelines and does not require MRI parameters, determines PA from the incident (forward minus reflected) root-mean-square (rms) power PI at the transmitter coil, but requires the coil’s quality factors (Q) measured with (QL), and without (QU), the subject load (18,19). The amount of power deposited in the subject is: PA=PI(1−QL∕QU). (1) Although determining these parameters at each exam would provide a straightforward approach to measuring SAR, various expediencies are commonly adopted for reporting and limiting SAR in many commercial clinical MRI scanners. For example, the Qs and PI are generally not measured at the coil terminals by the scanner during setup for a patient study, particularly in older scanners. Instead, fixed factory-determined parameters characterizing the selected coil-type are typically entered into scanner configuration files to calculate the SAR for a desired duty cycle, with conservative safety factors added to accommodate variations that may occur in the field. These “scanner SAR” values are utilized by the scanner for limiting pulse sequence parameters, and are generally accessible to the MRI operator. The accuracy with which such “scanner SAR” values measure the true average SAR may be compromised when coil Qs and power losses in the transmit line change with time, and/or when such conservative safety factors are introduced into the scanner’s SAR calculation by the manufacturer. This uncertainty is further compounded by those cases in which RF burns are sustained by patients undergoing MRI (9,20) (Table 1). Although such cases represent a tiny fraction of the millions of clinical MRI scans performed annually, they are evidence that “scanner SAR” was incorrect, or at least not limiting, at the time of the injury. Thus, the accuracy of “scanner SARs” reported by MRI scanners today, is a concern, and in any case, the SAR is currently not verifiable independent of the scanner. Indeed, if one seeks to independently measure deposited power via Eq. [1], access to the coil for power as well as QL and QU measurements is generally unavailable to scanner users, researchers, or medical physicists running routine checks on clinical units. Further problems arise with the testing of implanted devices, whose use in patients in need of diagnostic MRI grows exponentially. Devices that test safe at a specified “scanner SAR” level, may not be reliable if the “scanner SAR” includes unknown factors, or is scanner-dependent, as has been recently reported (21,22). A device testing as safe at a “scanner SAR” level of 4 W/kg, for example, may in fact have only been tested at, say, 2 W/kg if SAR is being conservatively overstated. In summary: 1) if patients report burning sensations during MRI (Table 1); or 2) if internal device/implant makers wish to determine that their devices are safe at a certain SAR exposure level; or 3) if questions arise during IRB assessment concerning the safe conduct of research studies, there are currently no convenient scanner-independent means of determining whether the machine is operating within regulatory SAR guidelines, other than performing thermal testing of instrumented phantoms (23). We propose and demonstrate herein a novel prototype RF dosimeter for measuring average head or body SAR, entirely independent of the MRI scanner. Peak SAR can then be determined from average SAR by applying an appropriate model factor as above (7,13-17). The dosimeter has a transducer that is placed in the magnet during an MRI scan, from which the time-averaged SAR is read. Because the scanner automatically sets the flip-angle based on MRI signal and load, the transducer contains a small signal-generating sample which the scanner uses for setup. The sample is surrounded by resistively-loaded orthogonal conducting loops, such that the transducer provides a load to the scanner that is equivalent to the head or torso. Such a transducer can also be used for SNR measurements as described previously (24). The time-averaged (rms) RF voltage is measured across the loop loads, from which PA and the equivalent body or head SAR are determined after calibration.