Improving Safety of MRI in Patients with Deep Brain Stimulation Devices
MRI is a valuable clinical and research tool for patients undergoing deep brain stimulation (DBS). However, risks associated with imaging DBS devices have led to stringent regulations, limiting the clinical and research utility of MRI in these patients. The main risks in patients with DBS devices undergoing MRI are heating at the electrode tips, induced currents, implantable pulse generator dysfunction, and mechanical forces. Phantom model studies indicate that electrode tip heating remains the most serious risk for modern DBS devices. The absence of adverse events in patients imaged under DBS vendor guidelines for MRI demonstrates the general safety of MRI for patients with DBS devices. Moreover, recent work indicates that—given adequate safety data—patients may be imaged outside these guidelines. At present, investigators are primarily focused on improving DBS device and MRI safety through the development of tools, including safety simulation models. Existing guidelines provide a standardized framework for performing safe MRI in patients with DBS devices. It also highlights the possibility of expanding MRI as a tool for research and clinical care in these patients going forward.
© RSNA, 2020
After reading the article and taking the test, the reader will be able to:
■ Identify the main risks of performing MRI for deep brain stimulation (DBS) devices
■ List the common dosimetric indexes used to define specific heating-related thresholds for DBS devices and discuss their limitations
■ Identify techniques and tools to improve MRI safety for patients with DBS devices
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Performance of MRI with deep brain stimulation devices carries risks, but developed tools aim to improve safety and expand the role of MRI in clinical care and research for these patients.
■ The most studied risks of performing MRI with deep brain stimulation (DBS) devices include device heating, induced currents, implantable pulse generator dysfunction, and magnetic field–induced device movement.
■ Device heating and resultant neuronal damage remains the most serious risk for performing MRI with modern DBS devices.
■ MRI performed under DBS vendor guidelines for MRI report no adverse events.
■ Following local safety testing to characterize the risks, acquiring optimal functional neuroimaging data safely outside prescribed DBS vendor guidelines for MRI could provide insight into pathologic brain states, as well as the structural and functional changes associated with active DBS therapy.
■ At present, tools to improve DBS safety including safety simulation models aimed to expand clinical and research use of MRI in these patients.
Deep brain stimulation (DBS) modulates aberrant neural circuits implicated in a broad range of neurologic disorders (1). Most commonly used in movement disorders such as Parkinson disease, dystonia, and tremor, DBS is also investigated for use in psychiatric and cognitive disorders such as depression and Alzheimer disease (2). Estimates show 150 000 patients have undergone surgical procedures for DBS worldwide (3). Approximately 70% of these patients will require an MRI within 10 years of implantation (4). However, past adverse MRI-related incidents led to strict guidelines restricting access to imaging for patients with DBS devices (5). In addition to constraining the diagnostic utility of MRI in this population, these regulations substantially impede MRI-based research.
The DBS literature reports five cases of injury involving radiofrequency (RF) currents—three MRI-related (Table 1), which prompted a 2005 U.S. Food and Drug Administration warning regarding patients with DBS devices undergoing MRI (6). To prevent additional injuries, DBS hardware vendors (ie, Medtronic, Boston Scientific, and Abbott) established MRI guidelines for imaging patients with DBS devices (5,7). These vendors approved and produced MRI-conditional DBS devices. Vendor guidelines restrict MRI coil types and gradient settings, as well as scan only at specific magnetic field strengths (ie, 1.5 T) and specific heating-related thresholds (eg, specific absorption rate [SAR] ≤0.1 W/kg and root-mean-square value of the MRI effective component of the RF magnetic [B1] field [B1+RMS] <2 μT) (8,9). These guidelines mainly aim to prevent MRI-induced device heating, which could lead to substantial brain damage (6). Recently, some constraints limiting brain MRI of patients with DBS devices have been relaxed, with certain newer DBS device models deemed full-body eligible (8–10).
Since the first reported use of MRI in patients with DBS devices, phantom models have been the most common tool to improve our understanding of DBS device and MRI safety (Fig 1). Phantom models allow investigators to simulate, albeit with limitations, how DBS devices behave during MRI scanning under safe and controlled conditions (5). However, a growing number of investigators now embrace newer techniques such as computer simulation to improve safety.
Enhanced knowledge of MRI safety in patients with DBS devices could lead to more proportionate safety guidelines, thereby improving patient care. Tagliati et al (11) found that nearly half the centers they surveyed were not performing brain MRI in patients with DBS devices. Moreover, only 13% of centers reported performing MRI of other body parts. These numbers are particularly worrisome given that MRI has become the reference standard for imaging many neurologic pathologic conditions (including surgical emergencies such as suspected spinal cord or cauda equina compression), as well as a range of musculoskeletal and abdominal pathologic conditions. In fact, 66%–75% of patients diagnosed with Parkinson disease, essential tremor, or dystonia will need an MRI in the 10 years following an operation for DBS (4). Most of these MRI examinations (62%) would be body MRI unrelated to the primary neurologic or neuropsychiatric diagnosis. Furthermore, higher field strength (eg, 3.0 T and 7.0 T) and less common sequences (eg, functional MRI and diffusion-weighted imaging) continue to show potential for clinical and research use. Investigation of medical implant safety under these conditions generates considerable interest (12–15). For example, functional MRI allows visualization of brain activity changes as a result of stimulation, revealing clinically efficacious networks in Parkinson disease (16,17) and obsessive-compulsive disorder (18). Also, functional MRI demonstrates the motor and nonmotor cortical segregation of the subthalamic nucleus, which is a structure targeted for Parkinson disease (19). Similarly, diffusion-weighted imaging commonly assesses stroke, assists neurosurgical planning, and provides disease markers (20,21). However, DBS vendor guidelines remain highly restrictive and often prohibit using both routine clinical (eg, DBS electrode localization ) and research (eg, arterial spin-labeling ) protocols (23,24).
MRI-related Risks of DBS Devices
Standards related to MRI safety testing of neurostimulators are based on the American Society for Testing and Materials International (25–27), International Organization for Standardization Technical Standard or ISO/TS 10974 (28), and other documents (29). A detailed description of these constantly evolving standards is beyond the scope of this review. Below is a summarization of the most common MRI risks studied in the literature (26). However, other potential risks remain, such as gradient field–induced vibrations.
Heating of DBS devices and the potential for subsequent brain damage constitutes the main risk when performing MRI in patients with neurostimulators (30–32). RF pulses applied during MRI elicit a detectable signal in bodily tissue and thereby acquire images of the target structures. These pulses may induce high currents in DBS electrodes and extension wires (5,33,34). Specifically, provided a conductive system (the extension wire and lead in the case of DBS) of appropriate length, the rapidly changing magnetic fields during RF excitation induce a current (ie, the antenna effect) (35). Then, the induced current dissipates as heat at the electrode tip. This is the location where the electrical current flux density is highest (and with highest resistance) (36). Because permanent brain damage can occur at temperatures exceeding 45°C (113°F) (ie, 7°C–8°C [45°C–46°F] in excess of normal body temperature) (37), temperature increases of less than 2°C (36°F) are felt to be acceptable and within a sufficient margin of safety (12,38–40). In addition to improving patient safety, minimizing MRI-induced heating could also improve data quality. For example, because blood oxygen level–dependent signal may be altered by a change in temperature, the minimization of DBS device heating may improve functional MRI data quality (39,41).
DBS device heating is related to both intrinsic and extrinsic factors (Fig 2). Extrinsic factors include the amount of RF energy applied, related to the SAR and B1+RMS (5). These metrics are used to estimate implant heating (42). The magnetic field strength and the type of MRI coils also constitute extrinsic factors. When using body-transmit coils, a larger proportion of the DBS device hardware is exposed to the RF pulses compared with head-transmit coils; as such, body-transmit coils cause more device heating (12,33,39,43). Intrinsic factors, such as the type of neurostimulator, may also influence temperature rise. Different brands and models of DBS device components (ie, electrodes, extension wire, and implantable pulse generator [IPG]) may exhibit different temperature increases based on their electrical characteristics (32). The geometry of the implanted device within the patient can also contribute to heating (12,33). The configuration of the implant is often institution-specific and varies in terms of electrode quantity, placement of excess extension wires, and IPG positioning. Finally, patients’ position in reference to the MRI bore and coils may further influence heating (39).
Measures of Heating
MRI guidelines specify heating-related thresholds (ie, SAR and B1+RMS) to prevent brain injury. Historically, SAR, a dosimetric index characterizing the thermogenic aspects of an electromagnetic field, commonly estimates energy deposition during MRI (7). However, SAR has inherent limitations, namely that its calculations vary by MRI system manufacturer and is derived by using various evolving assumptions and models of the human body by manufacturers (7,44,45). These limitations might explain why B1+RMS—a measure of the average magnetic field generated by the RF transmit coil—may provide an alternative dosimetric index to SAR (42). Unlike SAR, B1+RMS is independent of the patient’s weight, vendor-specific assumptions, and MRI hardware and software. It is a requirement of the International Electrotechnical Commission that newer MRI software provide this measure (29). However, B1+RMS lacks peer-reviewed evidence to confirm its role as a viable alternative (45). Although SAR and B1+RMS are useful for establishing precautionary measures regarding global tissue heating in patients, they are less informative when DBS implants are present. SAR and B1+RMS measure the energy deposited globally in the patient and not specifically at the implant, as they assume heating is uniform within the field of view (46–48). Crucially, both indexes fail to account for the localized high-RF electric fields (and potentially heating) at the electrode tips and thus should be interpreted cautiously.
To prevent dangerously high levels of DBS device heating, dosimetric indexes (ie, global SAR and B1+RMS) restrict the MRI acquisition parameters. To follow the recommended MRI guidelines, routinely used clinical MRI protocols may need modification to decrease their SAR and/or B1+RMS. This can involve reducing the number of slices, flip angles, and number of echoes, or increasing the repetition time (49). Such modifications typically necessitate the presence of those with a specialized knowledge base (ie, radiologists and physicists) and may result in compromised image quality (ie, signal-to-noise ratio and contrast-to-noise ratio) and/or increased acquisition time (49,50). Moreover, reducing the number of slices requires radiologists’ supervision to ensure the area in question remains in the field of view. Pulse sequences with larger and more frequent RF pulses have higher SAR and B1+RMS. Spin-echo images using large pulses (eg, fluid-attenuated inversion recovery) and fast spin-echo images using rapid pulses (eg, T2-weighted sequences) deliver more RF power, increasing their likelihood of causing heating. Conversely, gradient-echo images (eg, gradient-recalled echo–echo-planar imaging) use smaller and more frequent RF pulses, generally considered safer for patients with DBS implants (49). In sum, with specialized knowledge and clinical expertise, MRI usage in patients with DBS may be slightly expanded with modifications of routine clinical MRI protocols to remain within approved limits of SAR and/or B1+RMS.
MRI can induce currents in DBS hardware through two processes. First, RF pulses and gradient switchings cause time-varying magnetic fields capable of inducing impulses in the circuit, consisting of the pulse generator, leads, electrodes, and brain matter (35). Second, the antenna effect is also responsible for induced currents in addition to heating (35,36). Because RF currents have a carrier frequency in the megahertz range, they would not normally induce neuronal activity (ie, action potentials) (39,51). Gradient switching, however, may induce currents at lower frequencies (low kilohertz range) associated with neuronal firing (39). Depending on the position of the electrodes relative to neuroanatomic structures and fiber bundles, these currents could cause patient discomfort (eg, paresthesias, muscle spasms) or potentially more severe adverse effects such as seizures (52). Unintended stimulation and tissue damage produced by gradient-induced lead voltage, as well as tissue damage due to rectification produced by RF-induced lead voltage, should be monitored as per the standards (25,27,28).
Magnetic fields produced by MRI hardware could interfere with IPG function (27). Older IPG models (eg, Itrel II; Medtronic, Minneapolis, Minn) relied on a magnetic reed switch to turn the device on or off. Depending on the patients’ position relative to the MRI magnetic field, these reed switches could theoretically become spontaneously activated during MRI scanning, with the potential to harm patients and damage DBS components or alter settings (53). Whereas the most commonly used newer IPG models may not have magnetic reed switches, device malfunction induced by static magnetic fields, gradients, or RF fields should nonetheless be tested for following imaging as per the standards (25,27,28).
Magnetic Field–induced Device Movement
As modern components of MRI-conditional neurostimulators are nonmagnetic or diamagnetic, strong mechanical forces are not expected (5). Older IPG models (ie, Itrel; Medtronic) contained residual magnetic materials, such as sealing chips, ferrite core antennae, and reed switches (53,54). The magnetic component in newer IPG models should be minimal because leads and extension wires are nonmagnetic (55–57). Because the magnetic force in the MRI isocenter is theoretically minimal (owing to the high homogeneity of the magnetic field), one would expect the mildly magnetic IPG components located at the periphery to be most sensitive to mechanical forces (54).
Phantom Safety Studies
Since 1992, more than 15 studies described the use of phantom models to investigate the safety of DBS during MRI (Table E1 [online]). Device heating and, to a lesser extent, IPG function and device movement were most commonly assessed. Except for one study that examined thoracic area MRI (34), phantom studies have only assessed brain MRI safety. As such, the following section concerns brain MRI.
Per the American Society for Testing and Materials F2182 and ISO/TS 10974 standards (25,28), phantoms, defined as standardized containers filled with a medium that simulates the electrical and thermal properties of the human body (Fig 3), should measure RF-induced heating near the DBS device during MRI. These standards are subject to improvement for better estimate of heating (58). The most commonly used medium in phantom studies has been semisolid gel containing polyacrylic acid. The gelling agent prevents the transportation of phantom material that is locally heated by thermal convection (59). In addition, it simulates the permittivity and conductivity of various tissues in the frequency range of interest (60). The temperature should be recorded at the location suspected to have the highest amount of heating. In the case of a DBS device, this is the most distal uninsulated contact of the stimulating electrode. If substantial heating is detected, then computational modeling may refine the safety testing by simulating conditions in vivo. However, phantom models are limited by their lack of perfusion—thereby failing to replicate heat dissipation and the thermoregulatory effects of blood flow—and their inability to reproduce the thermodynamics of the human brain (6). Nonetheless, these experiments provide invaluable opportunities to explore implant safety outside of DBS vendor guidelines for MRI.
Although DBS devices have evolved, they continue to be implanted in similar configurations (ie, linear conductive wire), meaning that heating due to the antenna effect remains a major concern. Unacceptable temperature rises (ie, >2°C [36°F]) at the electrode tips have been reported in phantom models with pulse sequences using excessive SAR (37,61), with higher magnetic field strength (12,51), and with different DBS device configurations (62). Because of the constantly evolving nature of MRI hardware, DBS devices, and surgical techniques, the data from these studies may not be readily applicable to today’s conditions.
Extrinsic factors related to DBS device heating vary across phantom studies. First, as expected, higher field strengths are generally associated with greater temperature rises. However, the difference in temperature rises between 3.0 T and 1.5 T was less than 1°C (34°F) in recent studies (16,39). Interestingly, variations in MRI hardware and software systems contribute to inconsistencies in temperature rises. Notably, Baker et al (7,44) demonstrated that two different-generation MRI systems from the same manufacturer, using similar coils and RF-deposited energy, nonetheless triggered different temperature changes; they concluded that console-reported SAR in patients with DBS devices is an unreliable predictor of heating. Second, studies examined the relationship between coil type (ie, head-transmit or body-transmit coil) and temperature rise. Strikingly, Rezai et al (33) reported a temperature rise as high as 25.3°C (77.5°F) with a body-transmit coil, compared with a rise of 7.1°C (44.8°F) with a head-transmit coil. These results were not reproduced in several recent studies, which showed a difference of less than 1°C (34°F) between both coils (12,16,39,40). The discrepancy might be explained by the fact that Rezai et al (33) studied sequences that used more RF energy than would be used with clinical sequences. This study design modeled extreme scenarios, whereas the more recent reports investigated clinically applicable sequences. Differences in DBS device models may also contribute to these inconsistencies. Finally, phantom models show considerable variation across studies in terms of shape and filling medium. Specifically, the concentration of polyacrylic acid used to make the semisolid flesh-simulating gel reportedly influences device heating (59). To avoid underestimating device heating, investigators used increasingly higher concentrations of polyacrylic acid over the years.
Phantom model studies had inconsistencies in intrinsic factors related to DBS device heating. First, many experiments date back to the early 1990s, testing obsolete DBS models no longer routinely used in clinical practice (33,37,38,44,51,53,59,61–65). Moreover, all but one study (53) exclusively examined Medtronic DBS devices (Table E1 [online]). This limits the applicability of this data to other brands and more recent models because electrical conductivity (and thus heating) properties may vary across DBS devices (5). Second, DBS device configurations may vary in terms of electrode number and IPG, as well as the placement of excess extension wires. Many studies highlight the crucial role of configuration and geometry of devices and their position in the RF electrical fields relative to the coil. For example, phantom model studies commonly report a slightly higher rise in temperature at the left electrode when compared with the right. This might reflect the different position of the left extension wire, which typically connects a left-sided electrode to a right-sided IPG (12,33,51,62). Coiling the excess extension wire exclusively behind the IPG had marked temperature rises as high as 25.3°C (77.5°F); by contrast, partially coiling it at the cranium had maximal rises of only 6.1°C (43°F) (33). Small concentric loops placed on the cranium reduced DBS device heating (66). Nazzaro et al (62) showed that a DBS device configuration with bilateral IPG had higher temperature rise at the electrodes. Finally, short-circuit or open-circuit malfunction created by compromised DBS device integrity (eg, extension breakage) causes sparking (35). DBS vendor guidelines require confirming “normal” DBS device impedances (ie, bipolar current >250 Ω or monopolar current <2000 Ω) (9). Safety data from one group may not apply to different MRI hardware or institutions. Therefore, perform safety testing with MRI hardware and typical DBS configurations on an institution and hardware-specific basis.
Although the use of heating-related thresholds such as SAR and B1+RMS are general predictors of heating, it is unclear to what degree these metrics apply when implants are present. Phantom model studies report a linear relationship between SAR and temperature rise, irrespective of the DBS device and MRI hardware used (7,44). Pulse sequences with higher SAR, such as T2-weighted sequences, tended to cause relatively higher rises in temperature in most studies (12,16,37,38,51). These data suggest a degree of consistency within highly specific environments but confounded by a high degree of variability in induced heat per unit of SAR (Δtemperature rise/SAR) across MRI scanners and software (7,44), and DBS configurations (single compared with bilateral IPG) (62). Therefore, SAR alone cannot estimate absolute temperature rise for different DBS devices and MRI hardware (32). Taking the limitation of measuring implant heating into consideration, SAR less than 1 W/kg during brain MRI has acceptable temperature rises of less than 2°C (36°F) (12,38–40,51,65). Although seldom reported (12,16), a whole-body measure such as B1+RMS likely has similar limitations as SAR. Although supposedly less restrictive (8,9), the presumed advantages of B1+RMS over SAR must be confirmed by additional safety studies.
MRI-induced currents in DBS devices could present additional risks for patients. Phantom studies that investigated IPG output during 1.5-T (39,51,53) or 3.0-T (12) MRI found that gradient switchings can induce up to 1.5 V (39). This value is low when one considers that patients with Parkinson disease or psychiatric disorders commonly receive 3–5 V or 5–7 V, respectively, from their DBS devices (16,67). Low voltages induced by gradient switchings would be unlikely to cause patient discomfort, even when (rarely) superimposed on the IPG pulses. Nevertheless, this mild induced voltage and potential neuronal activity could reduce the sensitivity of an experiment (eg, by using functional MRI to compare “on” vs “off” DBS states). The larger currents induced by RF excitation pulses (up to 7.0 V) (35) should not trigger neuronal activity because their frequency is in the megahertz range (35,39,51).
Prior Medtronic IPG models most commonly reported spontaneous switching between “on” and “off” states and mild movements of the IPG (35,53,65). However, recent studies using Activa models (Medtronic) demonstrated stable IPG output with no indication that gross translational or torque force on the IPG occurred (12,16,39). This finding is important for experiments (eg, using functional MRI) that assume the MRI environment does not disrupt IPG output.
Recent phantom model studies using modern DBS devices and MRI hardware led to safe image acquisition in patients with DBS at 3.0-T MRI by using body-transmit coils (12,16,40). Higher field strengths (>1.5 T) combined with a body-transmit coil crucially provide optimal signal-to-noise ratio and minimize image distortion (39). Phantom models also showed that sequences with research applications, such as functional MRI (12,16,35,40,51,62,64) and diffusion-weighted imaging (12), could be used in patients with DBS in specific conditions. However, arterial-spin labeling, which can measure cerebral blood flow, could not be safely acquired in a recent phantom study (12).
Given that DBS devices continue to use similar configurations (ie, linear conductive wire), heating and, to a lesser extent, induced currents remain the major risks. Generally, lower field strength (ie, 1.5 T) and transmit-receive head coils trigger less DBS device heating than do 3.0-T and body-transmit coils. However, recent phantom studies show that this increased heating is much smaller than suggested by older studies. Heating also varies unpredictably across MRI scanners, software, and DBS device configurations. Conversely, mechanical forces and IPG dysfunction appear to be less of a concern with modern DBS device models due to advances in hardware technology. These findings are limited to modern Medtronic devices. Finally, data from most phantom studies do not readily apply to today’s conditions due to constantly evolving MRI hardware, DBS devices, and surgical techniques.
Animal and Human Studies
DBS safety studies in animals provide the opportunity to perform postmortem examinations following MRI scanning to assess for peri-electrode brain damage. Human safety data are also crucial, providing the necessary clinical translation piece after adequate safety data has been gathered by using other methods.
The animal model safety literature has many major limitations. Most notably, no studies have simultaneously investigated both temperature rise and histopathologic changes in peri-electrode brain tissue (Table E2 [online]). Nevertheless, this body of work again demonstrates the importance of DBS device geometry relative to the RF coil. Unconventional device positions, such that DBS leads lie perpendicular to the MRI magnetic field (68) or that the cranial extension loop is placed laterally on the head (69), triggered markedly higher temperature increases. These results highlight the importance of standardized device implantation and patient positioning during MRI.
Human safety studies provide more insights (Table E3 [online]), despite being limited to Medtronic devices. Most of these studies were conducted between 2005 and 2011, coinciding with the Food and Drug Administration warning release and archiving regarding DBS MRI safety (6). The majority performed MRI in patients in accordance with DBS vendor guidelines, presumably with the intent of confirming the safety of said guidelines. Tagliati et al (11) reported no adverse events in more than 3000 patients with DBS scanned at 1.5 T (although one center used 1.0 T). Furthermore, no studies listed in Table E3 (online) reported adverse events. However, some studies used MRI parameters outside of DBS vendor guidelines, using sequences with high SAR (22,70,71), atypical IPG placement (such as in the abdomen) (72), or higher field strength (ie, 3.0 T) (12,16,40,73). Studies performing 3.0-T MRI reported no adverse events in 88 patients (16,40,64,73,74), among which 27 were performed with a body-transmit coil. Furthermore, the extent of the MRI artifact related to the DBS device was also limited on functional MRI sequences (Figs 4, 5) (73). Thus far, most functional MRI studies in patients with DBS have been limited to 1.5 T (19,75–80). The possibility of acquiring data—in particular, functional MRI—with optimal MRI hardware in patients with DBS may not only help to expand clinical MRI applications, but also could facilitate important discoveries in the field of neuromodulation. Given the constant evolution of MRI systems and DBS devices, it is difficult to differentiate between inherent risks and lack of safety testing by the implant manufacturer. As with most aspects of medicine, clinicians should always weigh the advantages and disadvantages of a diagnostic test. If the MRI is likely to provide substantial benefits for the patient with DBS, then some degree of risk may be acceptable (81).
Studies reported IPG dysfunction (absent of patient injuries) in older IPG models, most commonly the Itrel model (Medtronic) (54,72,82) and noting spontaneous activation or deactivation of IPG. Magnet-related activation of IPG reportedly occurred spontaneously during MRI, up to 427 times during a single MRI session (54). In rare instances, it required the replacement of IPG (11). Although more recent studies have not specifically assessed IPG output, the lack of reported adverse effects or changes in patients’ clinical status is reassuring and suggests modern IPG output is stable during MRI (12,16,40,73). Nevertheless, IPG status should be routinely checked after MRI. Stability of the DBS device impedances ensure the electrical circuit integrity of the system and the stability of peri-electrode tissue (eg, absence of gross edema or hemorrhage) (12,73). Finally, few recent studies explicitly report performing MRI while modern IPGs were turned on using patients’ optimal stimulation settings, most often monopolar stimulation (12,16,73). At present, only one vendor allows IPG to be turned on to bipolar stimulation during MRI (9). This is problematic because turning off the IPG or using suboptimal settings typically exacerbates patients’ symptoms. Clarifying the safety and stability of IPG output requires further studies, regardless of current type.
Techniques and Tools to Improve Safety
Technologic advancements create opportunities to use techniques and tools to improve MRI safety for patients with DBS (Table 2). Since the Food and Drug Administration warning was archived in 2011, a dramatic increase in studies report innovative ways to improve MRI safety for patients with neurostimulators (Fig 1).
Table 2 categorizes technique and tool-related studies into optimized MRI acquisition parameters, DBS hardware modification, MRI hardware (including coils), and simulation models (Table 2). Few studies aim to change DBS device design, possibly due to the proprietary nature of medical devices and the technologic and economic barriers precluding most researchers from building in-house devices. For example, Elwassif et al (83) incorporated device components to act as heat sinks (ie, modification of the thermal conductivity of support material), which can disperse potentially hazardous temperature rises. In addition to safety considerations, DBS electrode design may undergo modification to decrease the MRI-related susceptibility artifact (84,85). Other studies have reported MRI hardware and coil modifications to improve MRI safety. Given the influence of the spatial relationship between MRI coils and the DBS device, investigators have developed modified coils resulting in lower SAR and heating, specifically proposing the use of a rotatable linearly polarized birdcage transmitter (86,87) and an optimized parallel RF transmission with a different coil design (88–91). One group dedicated their efforts into customizing MRI acquisition parameters to limit the use of RF pulses—and thus SAR—while maintaining image quality (92–94). Finally, computational models provide an opportunity to explore the multivariable problem of RF-associated heating and further elucidate the interaction between electromagnetic fields and biologic tissues (95,96). Generally, these studies aim to either provide tools that computationally simulate and estimate temperature rise at the electrode or establish a means through which to minimize DBS device heating on an individual patient basis. Individualized models of patients with DBS (97–99), high-spatial-resolution anatomic templates (100), and the angle of DBS lead insertion (101) have been investigated with computerized simulation. However, these tools require further studies to confirm their reliability in maintaining patient safety.
Summary of Best Practice Guidelines for Imaging Patients with DBS
Patients with DBS devices commonly present for MRI examinations. Fortunately, newer DBS devices have less restrictive MRI guidelines and more practical vendor specifications. However, radiologists commonly have requests for MRI examinations outside of the labeling requirements of a specific DBS device. In these circumstances, the clinician must determine whether (and how) to proceed with the MRI examination. Based on the guidelines of DBS vendors (8–10), reviewed literature, and our experience, we include a summary of recommendations of best practices for MRI in patients with DBS devices (Fig 6). These recommendations are not intended to account for all circumstances but provide a framework through which radiology groups can approach MRI in patients with DBS devices.
Heating remains the primary risk associated with DBS devices. This requires an accurate and reliable way to predict implant heating, rather than simply providing whole-body measures of heating, such as SAR or B1+RMS (32,42,99).
In the future, institutions that adopt patient-specific computerized models could then be applied in simulation software (102). The simulation tools currently under development align with providing individualized safety risks (97–99). If needed, then radiologists could modify routinely used protocols to remain within the safe limits determined by the bespoke simulation software, rather than limiting patients to the same MRI field strengths and low-SAR pulse sequences as if they were a homogeneous group.
Lowering the risk of heating altogether is not an easy task, as it would primarily involve changing the design of DBS devices and minimizing wire length. Drastic reductions in heating could potentially be achieved with technologic advancements such as much smaller, cranially placed IPG or wireless electrodes (103,104).
Finally, MRI safety advances in patients with DBS have opened research possibilities. Specifically, prospective acquisition of suboptimal functional MRI data (ie, 1.5 T, head-transmit coil) in patients with Parkinson disease and DBS showed promise as a marker of efficacy, suggesting that state-of-the-art functional neuroimaging data (ie, 3.0 T, body-transmit coil) may potentially represent an unparalleled research tool to probe brain functions. The risks of DBS devices in ultra-high-field-strength MRI (ie, 7.0 T) remains unevaluated (105). However, investigating the safety of ultra-high-field-strength MRI in patients with DBS is garnering interest because of its potential to be used as a powerful research and clinical tool.
Following MRI-related patient injuries in the early 2000s, considerable efforts have been made by both vendors and the scientific community to advance our understanding of deep brain stimulation (DBS) device and MRI safety and to provide further safeguards. Notably, heating remains a serious issue when venturing outside the prescribed guidelines.
At present, investigators are using innovative methods, particularly computational models, to assess MRI safety of neurostimulators in a more comprehensive and generalizable manner. Tailoring MRI safety guidelines to the individual may be the sensible next step for realizing the full potential of MRI in patients with a DBS device.Disclosures of Conflicts of Interest: A.B. disclosed no relevant relationships. C.T.C. disclosed no relevant relationships. K.N. disclosed no relevant relationships. G.J.B.E. disclosed no relevant relationships. C.N. disclosed no relevant relationships. J.G. disclosed no relevant relationships. M.R. disclosed no relevant relationships. A.L. disclosed no relevant relationships. A.J.M. Activities related to the present article: disclosed no relevant relationships. Activities not related to the present article: has grants/grants pending for MRI interventions (ClearPoint Neuro); is a consultant for UniQure. Other relationships: disclosed no relevant relationships. W.K. Activities related to the present article: disclosed no relevant relationships. Activities not related to the present article: is employed by the University of Toronto. Other relationships: disclosed no relevant relationships. C.J.S. disclosed no relevant relationships. I.H. Activities related to the present article: disclosed no relevant relationships. Activities not related to the present article: was employed by General Electric; is employed by the National Institutes of Health. Other relationships: disclosed no relevant relationships. A.R.R. disclosed no relevant relationships. A.M.L. Activities related to the present article: disclosed no relevant relationships. Activities not related to the present article: is a consultant for Abbott, Boston Scientific, Functional Neuromodulation, and Medtronic. Other relationships: disclosed no relevant relationships.
We would like to thank our illustrator, Andrew O’Connor, BA, for helping with the design of Figure 2.
Supported by the RR Tasker Chair in Functional Neurosurgery at University Health Network. K.N. supported by the Jane & Howard O. Jones Bursary Fund. C.N. supported by the German Research Foundation (Deutsche Forschungsgemeinschaft, DFG NE 2276/1-1). The opinions expressed in this article are the authors’ own and do not reflect the views of the National Institutes of Health, the Department of Health and Human Services, or the United States government.
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Article HistoryReceived: Oct 11 2019
Revision requested: Nov 20 2019
Revision received: Feb 13 2020
Accepted: Feb 20 2020
Published online: June 23 2020
Published in print: Aug 2020