Imaging PhysicsFree Access

MR Imaging and Cochlear Implants with Retained Internal Magnets: Reducing Artifacts near Highly Inhomogeneous Magnetic Fields

Published Online:https://doi.org/10.1148/rg.2018170135

Abstract

The number of patients receiving cochlear implants and auditory brainstem implants for severe to profound sensorineural hearing loss has rapidly increased. These implants consist of an internal component implanted between the skull and the temporal scalp and an external removable speech processor unit. A small magnet within the internal component is commonly used to hold the external speech processor unit in place. Several cochlear implant models have recently received U.S. Food and Drug Administration and European Economic Area regulatory approval to allow magnetic resonance (MR) imaging examinations to be performed under certain specified conditions. The small internal magnet presents a challenge for imaging of the head and neck near the implant, creating a nonlinear magnetic field inhomogeneity and significant MR imaging artifacts. Fat-saturation failures and susceptibility artifacts severely degrade image quality. Typical artifacts at diffusion-weighted imaging and accelerated imaging are exacerbated. Each examination may require impromptu adjustments to allow visualization of the tissue or contrast of interest. Patients may also be quite uncomfortable during the examination, as a result of either imposed magnetic forces or a tight head wrap that is often applied to minimize internal magnet movement. Translational forces and torque sometimes displace the implanted magnet even when a head wrap is used. Diseases such as neurofibromatosis type 2 that are associated with bilateral vestibular schwannomas and hearing loss often require lifelong tumor surveillance with MR imaging. A collaborative team of radiologists, technologists, and/or medical physicists or MR imaging scientists, armed with strategies to mitigate artifacts near implanted magnets, can customize the examination for better visualization of tissue and consistent comparison examinations over time.

©RSNA, 2018

SA-CME LEARNING OBJECTIVES

After completing this journal-based SA-CME activity, participants will be able to:

  • ■ Discuss the risks of proceeding with MR imaging for patients with a cochlear implant.

  • ■ Identify common MR imaging artifacts arising from cochlear implants with typical neuroimaging sequences.

  • ■ Describe strategies to reduce or shift artifacts from an anatomic region of interest.

Introduction

Cochlear implants and auditory brainstem implants (ABIs) have dramatically improved the quality of life for many patients with severe to profound sensorineural hearing loss. When these patients require a magnetic resonance (MR) imaging examination, the implanted components of the device introduce complexity to both patient preparation and image acquisition. The internal components of the implant—the active circuitry, electrodes, receiver antenna, and magnet—interact with various components of the MR imaging unit. Safety and image quality concerns exist when performing an MR imaging examination on a patient with unilateral or bilateral cochlear implants (1,2).

Many early cochlear implant models have never been MR imaging safety tested and are still considered MR Unsafe (using the ASTM International standard) in any MR imaging environment (3). However, several manufacturers now offer cochlear implants containing an internal magnet that have both U.S. Food and Drug Administration (FDA) approval and European Economic Area (EAA) CE marking as ASTM MR Conditional with specific conditions to image at 1.5 T, and at least one device is also MR Conditional for 3 T. Cochlear implants that contain a small magnet create a particular challenge at MR imaging. The patient may experience significant discomfort due to movement of the implant magnet and may need additional preparation before the examination. The local magnetic field from the implant also makes image artifacts notably worse. Signal loss can extend more than 10 cm from the implant when using a spin echo–based sequence, an imaging technique that is typically robust in the presence of magnetic field inhomogeneities (4). Artifacts from more-advanced imaging techniques, such as parallel imaging or long echo train imaging, may be subtle and located in regions away from the primary artifact.

Evaluation of tumors near the internal auditory canal and cerebellopontine angle requires fat-saturated imaging techniques, balanced gradient-echo acquisitions, and, potentially, diffusion-weighted imaging. All of these imaging techniques are highly sensitive to magnetic field inhomogeneities, and the implanted magnet often produces profound artifacts that may obscure adjacent regions of anatomy.

In this article, we describe the interaction of cochlear implants with the MR imaging environment. We also describe patient populations with cochlear implants that require routine MR imaging, how to prepare the patient to enter the MR imaging unit, and strategies to best perform a diagnostic MR imaging examination. Phantom studies demonstrate artifacts common to cochlear implants and are presented with corresponding clinical images to help illustrate artifact reduction strategies. MR imaging of patients with retained cochlear implant magnets is technically demanding, but with close collaboration among the radiologist, MR imaging technologist, and medical physicist or MR imaging scientist, the examination can be carefully tuned to the patient’s needs.

Cochlear Implants and the MR Imaging Environment

Cochlear implants are designed to bypass defective inner ear hair cells, sending electrical signals directly to the terminal ends of the cochlear nerve located within the cochlear modiolus. ABIs are of similar design but instead transmit signals directly to the cochlear nucleus in the brainstem and are most commonly used in patients with an absent cochlear nerve. As of 2012, more than 300 000 patients worldwide have received cochlear implants for rehabilitation of severe to profound sensorineural hearing loss (1). A complete cochlear implant or ABI system consists of both internal and external components. The external speech processor detects sound and transmits signals to the internal device transcutaneously. The internal device contains a radiofrequency (RF) antenna to receive signals, a magnet to secure the external speech processor in the proper position, an electronics package to transmit the processed signal, and an electrode to convey the signal to the cochlea or brainstem (Fig 1).

Figure 1a.

Figure 1a. Internal components of cochlear implants seen at radiography. (a) Oblique posteroanterior view shows the RF antenna (black arrow) to receive signals from an external speech processor, which is secured by an optional magnet (arrowhead). The processed signal is transmitted through an electrode (white arrow) to bypass defective inner ear hair cells, sending electrical signals directly to the terminal ends of the cochlear nerve located within the cochlear modiolus. (b, c) Lateral (b) and posteroanterior (c) views show cochlear implants implanted bilaterally.

Figure 1b.

Figure 1b. Internal components of cochlear implants seen at radiography. (a) Oblique posteroanterior view shows the RF antenna (black arrow) to receive signals from an external speech processor, which is secured by an optional magnet (arrowhead). The processed signal is transmitted through an electrode (white arrow) to bypass defective inner ear hair cells, sending electrical signals directly to the terminal ends of the cochlear nerve located within the cochlear modiolus. (b, c) Lateral (b) and posteroanterior (c) views show cochlear implants implanted bilaterally.

Figure 1c.

Figure 1c. Internal components of cochlear implants seen at radiography. (a) Oblique posteroanterior view shows the RF antenna (black arrow) to receive signals from an external speech processor, which is secured by an optional magnet (arrowhead). The processed signal is transmitted through an electrode (white arrow) to bypass defective inner ear hair cells, sending electrical signals directly to the terminal ends of the cochlear nerve located within the cochlear modiolus. (b, c) Lateral (b) and posteroanterior (c) views show cochlear implants implanted bilaterally.

Historically, the internal magnet had to be routinely removed before performing MR imaging and replaced after the examination was complete. This strategy reduces associated artifacts and discomfort from magnet movement. However, drawbacks include the need for two minor surgical procedures (magnet removal and replacement), an increased risk of infection, and a period of nonuse while the incision site heals. More recently, many centers have begun performing MR imaging with a retained internal magnet.

To reduce the risk for magnet displacement when the internal magnet is left in place, a head wrap or splint is often used during MR imaging (5). An internal magnet system that allows the internal magnet to freely rotate and align with the constant magnetic induction field (B0) has also been developed (Synchrony; MED-EL, Innsbruck, Austria), thereby reducing torque effects. Even with such precautions, patients may experience significant discomfort due to the movement of the magnet in the MR imaging environment. Some implant manufacturers make recommendations regarding the suitability of proceeding with MR imaging on the basis of the bone thickness adjacent to the magnet. When considering MR imaging for pediatric patients with a cochlear implant or ABI, tolerance of a tight head wrap, potential for discomfort, and bone thickness may be part of the decision of whether to proceed with MR imaging.

Teaching Point The implant manufacturer’s MR imaging conditions for any particular implant model are highly specific to that model and may intermittently be updated to reflect new testing or experience. The best clinical practice is to verify the implant model and current MR Conditional guidance each time a patient with a cochlear implant presents for MR imaging.
The implant manufacturer’s MR imaging conditions provide recommendations for acceptable B0, RF deposition rates, gradient slew rates (dB/dt), and spatial gradient of the magnetic field (dB/dx) near the mouth of the bore, as well as how best to prepare the implant and the patient for entering the specified MR imaging environment.

The three main subsystems of the MR imaging unit, the static magnetic field (B0), the RF field, and changing gradient magnetic fields, may interact with the internal components of the implant, and efforts can be made to minimize negative effects. These MR imaging subsystems and their associated safety precautions are described in depth by the International Electrotechnical Commission standard IEC 60601-2-33, edition 3.2 (3). This standard also mandates what safety information the MR imager manufacturer must provide in the instructions for use, including data for determining whether the particular imaging unit can accommodate the MR imaging conditions specified by the implanted device.

The B0 may cause the patient to experience a pulling sensation at the site of the magnet and torque, perceived as pressure, as the magnet rotates to align with the strong B0, both of which may subject the patient to considerable local discomfort. If the imaging unit table is of dockable design, patients with cochlear implants may be positioned for the examination on the table outside of the MR imaging room, away from the fringe field of the magnet. The fringe field of the MR imaging unit within 1 meter of the mouth of the bore produces the strongest translational force on the internal magnet. The technologist should avoid repositioning the patient near the mouth of the imaging unit bore, where the static magnetic field has the largest spatial gradient and causes the most discomfort due to force and torque imparted on the implanted magnet.

Once the patient is advanced into the bore, RF energy during active imaging may cause localized heating within the electrode and conductive loops of the antenna. Limiting RF deposition according to the implant manufacturer’s recommendations can mitigate these effects. The changing gradient magnetic field could induce currents within the implant circuitry; as active electronic devices, these effects may lead to the patient perceiving clicks and odd noises during image acquisition.

Despite the precautions specified in the MR Conditional labeling, surgical revision occasionally is required after completion of the MR imaging examination to correct magnet displacement (1,7,8). If the internal magnet has been removed and replaced multiple times, the silicone flange that secures the magnet may become compromised or fatigued, and the magnet may be more likely to displace or flip from its designated position when approaching the static magnetic field of the MR imaging unit.

Teaching Point Some degree of patient discomfort during MR imaging is not uncommon, and, given the risk that surgical revision may occasionally be needed after the examination, a risk-benefit analysis should be performed before each MR imaging examination for patients with a cochlear implant or ABI. If alternative diagnostic modalities such as computed tomography (CT) or ultrasonography may provide answers to the clinical question, these could be considered before attempting MR imaging.

Hearing Loss–related Diseases Best Monitored with MR Imaging

Hearing loss may stem from progressive disease processes that require frequent MR imaging surveillance and are often close to the cochlear implant or ABI device. Familiarity with disease manifestations, MR imaging techniques used to evaluate disease progression, and artifacts associated with these imaging techniques can help the team prepare an efficient MR imaging examination for a patient who may be very uncomfortable for the duration of the examination. Patients with neurofibromatosis type 2 (NF2), an inherited autosomal-dominant syndrome characterized by multiple schwannomas, meningiomas, and ependymomas, are often candidates for a cochlear implant or ABI. The hallmark of NF2 is bilateral vestibular schwannomas, which commonly result in progressive sensorineural hearing loss. As NF2-associated tumors also occur at other intracranial, skull base, and spinal locations, these anatomic regions also generally require frequent MR imaging surveillance (Fig 2).

Teaching Point Because of the proximity of the cochlear implant or ABI to the posterior fossa, patients with NF2 pose special challenges in performing diagnostic MR imaging, owing to artifacts and image degradation produced by the retained magnet (9). The artifacts may completely obscure the anatomy most in need of monitoring. These patients typically require at least yearly MR imaging examinations to evaluate for possible disease progression or treatment effects.

Figure 2a.

Figure 2a. Contrast-enhanced T1-weighted spin-echo images show hearing loss–related disease in two patients. Artifacts arising from cochlear implants may interfere with evaluation of tumors near the cerebellopontine angle. (a) Coronal image in a 39-year-old man shows an extra-axial mass consistent with vestibular schwannoma, a benign tumor of the Schwann cells responsible for myelin production about the vestibulocochlear nerve, which frequently causes symptoms of sensorineural hearing loss, tinnitus, imbalance, and occasionally facial numbness. (b, c) Axial brain (b) and sagittal cervical spine (c) images show evidence of NF2 in a 31-year-old man being evaluated for progression of the disease.

Figure 2b.

Figure 2b. Contrast-enhanced T1-weighted spin-echo images show hearing loss–related disease in two patients. Artifacts arising from cochlear implants may interfere with evaluation of tumors near the cerebellopontine angle. (a) Coronal image in a 39-year-old man shows an extra-axial mass consistent with vestibular schwannoma, a benign tumor of the Schwann cells responsible for myelin production about the vestibulocochlear nerve, which frequently causes symptoms of sensorineural hearing loss, tinnitus, imbalance, and occasionally facial numbness. (b, c) Axial brain (b) and sagittal cervical spine (c) images show evidence of NF2 in a 31-year-old man being evaluated for progression of the disease.

Figure 2c.

Figure 2c. Contrast-enhanced T1-weighted spin-echo images show hearing loss–related disease in two patients. Artifacts arising from cochlear implants may interfere with evaluation of tumors near the cerebellopontine angle. (a) Coronal image in a 39-year-old man shows an extra-axial mass consistent with vestibular schwannoma, a benign tumor of the Schwann cells responsible for myelin production about the vestibulocochlear nerve, which frequently causes symptoms of sensorineural hearing loss, tinnitus, imbalance, and occasionally facial numbness. (b, c) Axial brain (b) and sagittal cervical spine (c) images show evidence of NF2 in a 31-year-old man being evaluated for progression of the disease.

Patient Preparation for MR Imaging

Patients should be instructed to keep their reading glasses (if applicable) and the cochlear implant or ABI external speech processor with them to aid communication until they are ready to enter zone IV (the MR imaging room), as they may be functionally deaf without the external speech processor, which is labeled as MR Unsafe for all current models (10).

The MR imaging technologist should provide additional time before the examination to coach the patient on the expected sensations associated with the implant and to apply a splint with the assistance of a member of the otolaryngology team, if required. Depending on the required splint and head wrap design, significant pressure at the site of the internal magnet may be uncomfortable for the patient. The splint should be applied as a last step before the patient enters the imaging room to maximize the available imaging time. If directed by a physician, a local anesthetic may provide some relief for discomfort from the splint. A pneumatic call system is often used within the MR imaging unit to enable communication between the patient and the technologist. We recommend demonstrating and testing the call system once the patient is in position. Testing the call button serves three purposes: it confirms that the call system is functional, that the patient understands how to use it, and that the patient is physically capable of using it.

After completion of the examination, be prepared to immediately remove the splint. The external speech processor can be used as a basic test of functionality and to assess whether the cochlear implant magnet has been displaced. Palpation of the scalp at the site of the implant may also provide an indication of magnet position. In rare cases, radiography may be helpful for answering questions of whether the magnet remains in the proper location within the silicone sleeve (7). Pain at the implant site after the examination has been completed is often the first sign of a displaced magnet. If a patient reports pain or discomfort when using the device, this should be carefully investigated to ensure that the magnet is seated properly.

Typical MR Imaging Protocols and Associated Artifacts

With the internal magnet of the cochlear implant or ABI degrading the homogeneous B0 field necessary for MR imaging signal localization, significant artifacts should be anticipated even with the most robust acquisitions, including spin-echo–based techniques. Figure 3 shows typical acquisitions from a routine brain protocol on a 1.5-T MR imager and the artifacts generated from the presence of a cochlear implant with an internal magnet. The highly nonlinear magnetic field imposed by the internal magnet is beyond the correction capabilities of the linear gradient shim fields available from any MR imaging unit manufacturer. Several artifacts from the inhomogeneous B0 field are evident: intravoxel dephasing, where protons within a single voxel precess at different frequencies, resulting in signal loss; geometric distortion of the section selection pulse; and in-plane distortion for lower-bandwidth acquisitions, causing signal pile up or hot spots at some locations and stretched-out low signal in other locations (11). Distortion of the location of section selection is most evident for this examination on the sagittal T1-weighted spin-echo image (Fig 3a), where the magnetic field on the superior side of the internal cochlear implant magnet adds to B0, shifting the section selection to tissue medial to the intended section, and the magnetic field inferior to the cochlear implant magnet subtracts from B0, shifting the section selection lateral to the intended section (12,13). The final image displays a gross discontinuity where nonadjacent tissue is reconstructed on the image near the location of the cochlear implant magnet.

Figure 3a.

Figure 3a. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

Figure 3b.

Figure 3b. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

Figure 3c.

Figure 3c. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

Figure 3d.

Figure 3d. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

Figure 3e.

Figure 3e. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

Figure 3f.

Figure 3f. Artifacts at routine brain MR imaging in patients with cochlear implants with retained magnets. (Images in a, b, d, and e are from the same patient.) (a) The section selection pulse is geometrically distorted for the sagittal T1-weighted spin-echo acquisition (repetition time [TR] msec/echo time [TE] msec = 683/13) owing to the cochlear implant magnet adding to B0 superior to the implant and subtracting from B0 inferior to the implant. Distorted signal adds constructively, creating signal pile ups and over-ranges in certain locations. (b) Axial T2-weighted fluid-attenuated inversion-recovery (FLAIR) image (11 000/148) exhibits intravoxel dephasing and loss of cerebrospinal fluid suppression near the magnet. Inadequate cerebrospinal fluid suppression comes as a result of failed section selection for the inversion pulse. (c) Axial gradient-echo image (400/15) exhibits extreme intravoxel dephasing, even at relatively short TEs. (d) Axial diffusion-weighted echo-planar image (9000/81) exhibits extreme geometric distortion due to the low bandwidth readout in the phase-encoding direction. (e) Axial T2-weighted FSE image (4200/98) also displays large areas of intravoxel dephasing. (f) Even without an internal magnet, expect moderate metal artifact from the cochlear implant RF antenna, as seen on this axial T2-weighted FSE image (3900/98).

MR imaging datasets acquired with cochlear implants or ABIs present may not be suitable for precision surgical planning, such as might be needed for stereotactic radiosurgery or proton-beam surgical procedures. Even without the internal magnet, moderate metallic artifact from the implant antenna should be anticipated. Owing to the anticipated artifact, experienced implant surgeons frequently implant new cochlear implant or ABI systems as far away as possible from intracranial pathologic conditions that may require ongoing MR imaging surveillance or treatment. For example, implanting the device farther posterior and superior may allow better long-term monitoring of the internal auditory canal (14).

Recommended MR Imaging Protocol Modifications for Cochlear Implants and ABIs

Use of a Phantom with Attached Cochlear Implant for Protocol Testing

A phantom with an attached cochlear implant containing a small magnet is helpful to demonstrate MR imaging artifacts and the results of protocol adjustments in an object of known geometry and signal intensity. At our institution, a spherical phantom was imaged with an attached cochlear implant demonstration model (Nucleus CI00000 demonstration model; Cochlear, Sydney, Australia). The removable magnet was placed within the cochlear implant silicone flange, attached to the phantom, and imaged within an 8-channel brain coil on a 1.5-T MR imager (MR450W; GE Healthcare, Waukesha, Wis) to demonstrate the source or appearance of artifact for specific imaging techniques. Without the attached cochlear implant, the signal within the phantom would appear uniform in a perfectly circular shape within each two-dimensional image. To prevent migration of the small button magnet while within the MR imaging unit, the device was double-bagged and strapped to the phantom, and then the phantom with the attached device was placed inside an additional closed plastic container. In the following sections, phantom and clinical imaging with standard protocols are compared with images obtained when using the artifact mitigation strategies.


Fat Suppression

Fat-saturation techniques rely on the chemical shift of the proton resonance of fat relative to water. A chemical shift–selective (CHESS) RF pulse tuned to fat (220-Hz frequency shift at 1.5 T) saturates and dephases the fat signal (15). The CHESS pulse depends on a very homogeneous B0 field; any tissue that is 220-Hz off resonance is also saturated. Even in the absence of an implant, anatomically induced B0 offsets due to anatomic shape or differing magnetic susceptibilities can be challenging for CHESS pulses, in some cases causing the CHESS pulse to instead suppress water signal, or shifting the pulse off resonance in the opposite direction such that it no longer suppresses fat or water. Figure 4 shows a map of the magnetic field within the phantom, where the signal intensity swaps from white to black at 220-Hz intervals. A fat-saturation sequence demonstrates signal loss corresponding to the B0 off resonance at the same locations within the phantom as the signal intensity swaps. Likewise, in vivo, we observe regions of suppressed water signal and bright fat signal throughout an image that ideally would have uniformly dark fat (Fig 4c, 4d).

Figure 4a.

Figure 4a. Chemically selective fat saturation that is ineffective owing to B0 inhomogeneity. (a) With a multiecho phase map (35/4, 9) acquired in the axial plane using a phantom, cochlear implant–induced magnetic field inhomogeneities were mapped across the phantom such that phase swaps from black to white occurred at 220-Hz intervals, corresponding to the chemical shift between fat and water. (b) With an axial fat-saturated T1-weighted acquisition (573/10) in the phantom, regions of signal suppression were visible at corresponding locations to the phase swaps, despite the fact the phantom contained no fat. (c, d) Obscured left vestibular schwannoma (arrows in d depict tumor position) in a 59-year-old woman with NF2. On axial (c) and coronal (d) fat-saturated T1-weighted images (600/12), fat saturation fails in many places across the field. Most notably, the radiologist was unable to evaluate the internal auditory canal given the profound artifact and signal suppression.

Figure 4b.

Figure 4b. Chemically selective fat saturation that is ineffective owing to B0 inhomogeneity. (a) With a multiecho phase map (35/4, 9) acquired in the axial plane using a phantom, cochlear implant–induced magnetic field inhomogeneities were mapped across the phantom such that phase swaps from black to white occurred at 220-Hz intervals, corresponding to the chemical shift between fat and water. (b) With an axial fat-saturated T1-weighted acquisition (573/10) in the phantom, regions of signal suppression were visible at corresponding locations to the phase swaps, despite the fact the phantom contained no fat. (c, d) Obscured left vestibular schwannoma (arrows in d depict tumor position) in a 59-year-old woman with NF2. On axial (c) and coronal (d) fat-saturated T1-weighted images (600/12), fat saturation fails in many places across the field. Most notably, the radiologist was unable to evaluate the internal auditory canal given the profound artifact and signal suppression.

Figure 4c.

Figure 4c. Chemically selective fat saturation that is ineffective owing to B0 inhomogeneity. (a) With a multiecho phase map (35/4, 9) acquired in the axial plane using a phantom, cochlear implant–induced magnetic field inhomogeneities were mapped across the phantom such that phase swaps from black to white occurred at 220-Hz intervals, corresponding to the chemical shift between fat and water. (b) With an axial fat-saturated T1-weighted acquisition (573/10) in the phantom, regions of signal suppression were visible at corresponding locations to the phase swaps, despite the fact the phantom contained no fat. (c, d) Obscured left vestibular schwannoma (arrows in d depict tumor position) in a 59-year-old woman with NF2. On axial (c) and coronal (d) fat-saturated T1-weighted images (600/12), fat saturation fails in many places across the field. Most notably, the radiologist was unable to evaluate the internal auditory canal given the profound artifact and signal suppression.

Figure 4d.

Figure 4d. Chemically selective fat saturation that is ineffective owing to B0 inhomogeneity. (a) With a multiecho phase map (35/4, 9) acquired in the axial plane using a phantom, cochlear implant–induced magnetic field inhomogeneities were mapped across the phantom such that phase swaps from black to white occurred at 220-Hz intervals, corresponding to the chemical shift between fat and water. (b) With an axial fat-saturated T1-weighted acquisition (573/10) in the phantom, regions of signal suppression were visible at corresponding locations to the phase swaps, despite the fact the phantom contained no fat. (c, d) Obscured left vestibular schwannoma (arrows in d depict tumor position) in a 59-year-old woman with NF2. On axial (c) and coronal (d) fat-saturated T1-weighted images (600/12), fat saturation fails in many places across the field. Most notably, the radiologist was unable to evaluate the internal auditory canal given the profound artifact and signal suppression.

Teaching Point When imaging near the cochlear implant or ABI magnet, fat-suppression techniques such as short tau inversion time inversion-recovery (STIR) or robust Dixon techniques should be substituted for fat-saturation techniques.
STIR relies on the short T1 relaxation time of fat, using an inversion-recovery sequence to measure data when fat is dark (16). Dixon techniques collect data at multiple TEs to resolve signal originating from the 220-Hz fat-water chemical shift (1719). The Dixon reconstruction algorithm for fat- and water-only images may also struggle with severe B0 inhomogeneity, but if both fat and water datasets are saved, any spurious fat/water signal swaps could be sorted through by the interpreting radiologist. Figure 5 shows Dixon images acquired in the same patient as in Figure 4 at a subsequent MR imaging examination. Hybrid fat-suppression techniques that rely on both fat saturation and the short T1 signal properties of fat may more reliably suppress fat signal but will still saturate water signal at some locations. Although fat saturation with a CHESS pulse is commonly used for fat suppression within the head and neck, CHESS is the fat-suppression technique most sensitive to magnetic field inhomogeneity.

Figure 5a.

Figure 5a. Multiecho techniques for fat suppression. (a) Axial Dixon FSE imaging (700/13) resolves signals from fat and water and reconstructs them into separate images. (b, c) The algorithm to reconstruct the separate images can get confused by B0 field inhomogeneities and map tissue into the wrong dataset, as observed on this coronal Dixon FSE image (467/13) (b). The MR imaging technologist should be careful not to prematurely discard the fat images, as the fat and water images can be recombined for a composite image (c) if errors in the reconstruction occur. Dixon FSE imaging enabled easier evaluation of the vestibular schwannoma (arrow in b and c) in comparison with the chemically selective fat-saturation images obtained in the same patient 1 year earlier (Fig 4).

Figure 5b.

Figure 5b. Multiecho techniques for fat suppression. (a) Axial Dixon FSE imaging (700/13) resolves signals from fat and water and reconstructs them into separate images. (b, c) The algorithm to reconstruct the separate images can get confused by B0 field inhomogeneities and map tissue into the wrong dataset, as observed on this coronal Dixon FSE image (467/13) (b). The MR imaging technologist should be careful not to prematurely discard the fat images, as the fat and water images can be recombined for a composite image (c) if errors in the reconstruction occur. Dixon FSE imaging enabled easier evaluation of the vestibular schwannoma (arrow in b and c) in comparison with the chemically selective fat-saturation images obtained in the same patient 1 year earlier (Fig 4).

Figure 5c.

Figure 5c. Multiecho techniques for fat suppression. (a) Axial Dixon FSE imaging (700/13) resolves signals from fat and water and reconstructs them into separate images. (b, c) The algorithm to reconstruct the separate images can get confused by B0 field inhomogeneities and map tissue into the wrong dataset, as observed on this coronal Dixon FSE image (467/13) (b). The MR imaging technologist should be careful not to prematurely discard the fat images, as the fat and water images can be recombined for a composite image (c) if errors in the reconstruction occur. Dixon FSE imaging enabled easier evaluation of the vestibular schwannoma (arrow in b and c) in comparison with the chemically selective fat-saturation images obtained in the same patient 1 year earlier (Fig 4).

CISS or FIESTA-C

Balanced steady-state free precession–based imaging techniques have band artifacts that appear at off-resonance frequencies that depend on the TR. A shorter TR spreads the band artifacts farther apart. Even with phase-cycling schemes, such as constructive interference steady state (CISS) (Siemens Healthineers; Erlangen, Germany) or fast imaging employing steady-state acquisition with cycled phases (FIESTA-C) (GE Healthcare), designed to reduce signal loss within the band artifacts, residual bands remain near the large off-resonance band created by the implant magnet (20,21). The band artifacts may interfere with visualization of the anatomy of interest. As these imaging techniques are workhorse sequences for evaluation of the internal auditory canal, such artifacts are problematic. Figure 6 demonstrates band artifacts in a phantom and in vivo.

Figure 6a.

Figure 6a. Band artifacts from steady-state free precession techniques. (a) Band artifacts stand out against the uniform background of an axial FIESTA-C phantom-based image (5.5/2.6), even with the phase cycling scheme applied to reduce band artifacts. Slight adjustments in TR, head position, or shim values may shift artifacts away from the region of interest. (b, c) In a 61-year-old woman with NF2 (same MR imaging examination as in Fig 5), band artifact (between arrows in b) reduces the conspicuity of the tumor in the internal auditory canal on an axial FIESTA-C image (7.5/3.4) (b); after shim adjustments, the tumor is more easily visualized with identical acquisition parameters (c).

Figure 6b.

Figure 6b. Band artifacts from steady-state free precession techniques. (a) Band artifacts stand out against the uniform background of an axial FIESTA-C phantom-based image (5.5/2.6), even with the phase cycling scheme applied to reduce band artifacts. Slight adjustments in TR, head position, or shim values may shift artifacts away from the region of interest. (b, c) In a 61-year-old woman with NF2 (same MR imaging examination as in Fig 5), band artifact (between arrows in b) reduces the conspicuity of the tumor in the internal auditory canal on an axial FIESTA-C image (7.5/3.4) (b); after shim adjustments, the tumor is more easily visualized with identical acquisition parameters (c).

Figure 6c.

Figure 6c. Band artifacts from steady-state free precession techniques. (a) Band artifacts stand out against the uniform background of an axial FIESTA-C phantom-based image (5.5/2.6), even with the phase cycling scheme applied to reduce band artifacts. Slight adjustments in TR, head position, or shim values may shift artifacts away from the region of interest. (b, c) In a 61-year-old woman with NF2 (same MR imaging examination as in Fig 5), band artifact (between arrows in b) reduces the conspicuity of the tumor in the internal auditory canal on an axial FIESTA-C image (7.5/3.4) (b); after shim adjustments, the tumor is more easily visualized with identical acquisition parameters (c).

A few strategies to shift the artifact away from the anatomy of interest include adjusting imaging parameters to change the TR, adjusting the shim using a local shim volume during tuning, and manually overriding the applied shim. When manually adjusting the shim, depending on the implant location, often a simple doubling of the applied shim in the left-to-right direction is adequate to move the artifact.

Parallel or Accelerated Imaging Artifacts

Parallel or accelerated imaging relies on the location of the receiver coil elements to help spatially encode the image and is an integral part of many modern MR imaging sequences (22). Parallel imaging reconstruction uses calibration scans to map the relative signal detected by each coil element. The calibration scan may be acquired separately or integrated into the sequence itself. Typically, the calibration scan is a low-resolution gradient-echo acquisition that suffers from intravoxel dephasing. With signal dropout near the implant magnet, the parallel imaging reconstruction fails to spatially encode signal in that region, creating aliased signal on the reconstructed image at predictable locations based on the acceleration factor. Figure 7 shows a phantom image acquired with and without parallel imaging and an in vivo image demonstrating the parallel imaging artifacts. With an acceleration factor of 2, the aliased signal is shifted in the phase-encoding direction to a location half of the field of view from where the signal originated. Depending on the imaging parameters, techniques such as generalized autocalibrating partially parallel acquisition (GRAPPA) with integrated oversampling instead of a calibration scan may have fewer artifacts (23).

Teaching Point If possible, avoid accelerated acquisitions when imaging a patient with a cochlear implant or ABI, as receiver coil elements near the internal magnet do not contribute to the spatial encoding as might otherwise be expected.

Figure 7a.

Figure 7a. Accelerated parallel imaging artifacts. (a) Axial unaccelerated FSE phantom-based image (3900/93) shows obvious susceptibility artifact near the implant. (b) Axial gradient-echo calibration phantom-based image (150/2), which is used to map the location and sensitivity of each receiver coil element, is missing data critical to helping spatially encode the data. (c) The addition of sensitivity encoding (SENSE) to accelerate the acquisition creates additional large artifacts (arrows) on the right side of the phantom. (d) Axial in vivo accelerated T2-weighted FSE image (3917/63) shows similar patterns of artifact (arrows).

Figure 7b.

Figure 7b. Accelerated parallel imaging artifacts. (a) Axial unaccelerated FSE phantom-based image (3900/93) shows obvious susceptibility artifact near the implant. (b) Axial gradient-echo calibration phantom-based image (150/2), which is used to map the location and sensitivity of each receiver coil element, is missing data critical to helping spatially encode the data. (c) The addition of sensitivity encoding (SENSE) to accelerate the acquisition creates additional large artifacts (arrows) on the right side of the phantom. (d) Axial in vivo accelerated T2-weighted FSE image (3917/63) shows similar patterns of artifact (arrows).

Figure 7c.

Figure 7c. Accelerated parallel imaging artifacts. (a) Axial unaccelerated FSE phantom-based image (3900/93) shows obvious susceptibility artifact near the implant. (b) Axial gradient-echo calibration phantom-based image (150/2), which is used to map the location and sensitivity of each receiver coil element, is missing data critical to helping spatially encode the data. (c) The addition of sensitivity encoding (SENSE) to accelerate the acquisition creates additional large artifacts (arrows) on the right side of the phantom. (d) Axial in vivo accelerated T2-weighted FSE image (3917/63) shows similar patterns of artifact (arrows).

Figure 7d.

Figure 7d. Accelerated parallel imaging artifacts. (a) Axial unaccelerated FSE phantom-based image (3900/93) shows obvious susceptibility artifact near the implant. (b) Axial gradient-echo calibration phantom-based image (150/2), which is used to map the location and sensitivity of each receiver coil element, is missing data critical to helping spatially encode the data. (c) The addition of sensitivity encoding (SENSE) to accelerate the acquisition creates additional large artifacts (arrows) on the right side of the phantom. (d) Axial in vivo accelerated T2-weighted FSE image (3917/63) shows similar patterns of artifact (arrows).

Metal Artifact–reduction Techniques

Metal artifact–reduction techniques such as multi acquisition with variable resonance image combination (MAVRIC) are designed to handle B0 inhomogeneity near large metallic implants of varying magnetic susceptibility (2426). Typically, these techniques acquire data at several frequency offsets to map the off-resonance data back onto the primary image. Variations of the technique may use view-angle tilting to also correct the distorted section selection. These techniques do require additional acquisition time to capture the frequency offsets and are tuned to the frequency offsets of common orthopedic implants, which are not a large enough range to resolve the B0 inhomogeneity induced by the implant magnet. However, the size of the artifact is markedly reduced and may allow evaluation of anatomic structures that would otherwise be obscured by artifact. In comparison, three-dimensional T1-weighted fast spin-echo (FSE) sequences using variable flip-angle techniques are highly sensitive to off resonance as the refocusing pulses get out of phase (Fig 8) (27).

Figure 8a.

Figure 8a. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Figure 8b.

Figure 8b. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Figure 8c.

Figure 8c. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Figure 8d.

Figure 8d. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Figure 8e.

Figure 8e. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Figure 8f.

Figure 8f. Use of MAVRIC at three-dimensional T1-weighted imaging to correct for metal artifacts. Lateral (top) and midline (bottom) two-dimensional spin-echo (468/9) (a, b), MAVRIC (655/7) (c, d), and CUBE (GE Healthcare) (600/15) (e, f) sagittal images demonstrate how MAVRIC reduces the size of the signal void near the implant. T1-weighted CUBE, an alternative three-dimensional acquisition, depends on maintaining the phase coherence of a long train of flip angles and performs very poorly around magnets.

Conventional wisdom suggests increasing the receiver bandwidth to resolve structures near metal or other B0 inhomogeneities. However, the highly nonlinear B0 field near the small cochlear implant or ABI magnet changes rapidly, so there is only a modest benefit of about a centimeter to increasing the receiver bandwidth, at the cost of reduced signal-to-noise ratio across the image. If the operator is not careful about how the receiver bandwidth changes other imaging parameters, the operator may be surprised with the resulting image contrast. For example, the echo train length for an FSE sequence may become too short to obtain the desired effective TE, providing more of a proton-density–weighted appearance than a T2-weighted appearance. Figure 9 shows the FSE images of the phantom with receiver bandwidths of 16 kHz and 50 kHz. The echo train length after each 90° excitation pulse had to be doubled to obtain the desired TE. With only slight benefit near the magnet, but overall image-quality reduction, increasing the receiver bandwidth adds little value and risks unforeseen drawbacks across the rest of the image. Increasing the receiver bandwidth should generally only be used as a problem-solving tool or to reduce signal pile up to evaluate tissue near the implant.

Figure 9a.

Figure 9a. Bandwidth adjustment to reduce susceptibility artifacts. An axial phantom-based FSE image (3200/99) acquired at a 16-kHz receiver bandwidth (a) has a slightly larger signal void than a phantom image (3200/99) acquired at a 50-kHz receiver bandwidth (b). Increasing the bandwidth may recoup approximately 1 cm of signal around the susceptibility artifact but comes at the expense of signal-to-noise ratio and potentially altered tissue contrast. Unless there is tissue immediately adjacent to the artifact that requires close evaluation, the bandwidth should not be adjusted too far from its standard setting to maintain image quality across the rest of the image.

Figure 9b.

Figure 9b. Bandwidth adjustment to reduce susceptibility artifacts. An axial phantom-based FSE image (3200/99) acquired at a 16-kHz receiver bandwidth (a) has a slightly larger signal void than a phantom image (3200/99) acquired at a 50-kHz receiver bandwidth (b). Increasing the bandwidth may recoup approximately 1 cm of signal around the susceptibility artifact but comes at the expense of signal-to-noise ratio and potentially altered tissue contrast. Unless there is tissue immediately adjacent to the artifact that requires close evaluation, the bandwidth should not be adjusted too far from its standard setting to maintain image quality across the rest of the image.

Echo-planar and Diffusion-weighted Sequences

Echo-planar imaging is notoriously sensitive to B0 off resonance. With its rastered acquisition collecting multiple lines of k-space following a single RF excitation, the phase-encoding direction is essentially doing a very-low-bandwidth acquisition (28). This manifests as significant geometric distortion and intravoxel dephasing adjacent to normal anatomic structures of differing susceptibility. In the presence of a cochlear implant or ABI, the diffusion-weighted echo-planar images are often rendered nearly uninterpretable, with only small regions of clear anatomy visible through the smeared mess. However, given the brief acquisition time and the unique ability to depict early infarction, most radiologists would still recommend attempting diffusion-weighted echo-planar imaging.

Other data acquisition schemes, such as periodically rotated overlapping parallel lines with enhanced reconstruction (PROPELLER; GE Healthcare) or BLADE (Siemens Healthineers), acquire the data in multiple shots, traversing k-space in a unique fashion (29,30). The pattern of artifact is quite different from that produced at echo-planar imaging and may prove useful should the patient need assessment for stroke. Figure 10 shows diffusion-weighted images acquired with echo-planar imaging and with PROPELLER in a patient with a right-side cochlear implant.

Figure 10a.

Figure 10a. Axial diffusion-weighted imaging in a patient with a right-side cochlear implant. Images of two section locations from a diffusion-weighted dataset acquired with echo-planar imaging (9200/81, b = 1000 sec/mm2) (a, b) are compared with images of the same two sections acquired with a diffusion-weighted PROPELLER acquisition (6500/86, b = 800 sec/mm2) (c, d). The PROPELLER images still have a band of signal loss, but much more anatomy can be evaluated.

Figure 10b.

Figure 10b. Axial diffusion-weighted imaging in a patient with a right-side cochlear implant. Images of two section locations from a diffusion-weighted dataset acquired with echo-planar imaging (9200/81, b = 1000 sec/mm2) (a, b) are compared with images of the same two sections acquired with a diffusion-weighted PROPELLER acquisition (6500/86, b = 800 sec/mm2) (c, d). The PROPELLER images still have a band of signal loss, but much more anatomy can be evaluated.

Figure 10c.

Figure 10c. Axial diffusion-weighted imaging in a patient with a right-side cochlear implant. Images of two section locations from a diffusion-weighted dataset acquired with echo-planar imaging (9200/81, b = 1000 sec/mm2) (a, b) are compared with images of the same two sections acquired with a diffusion-weighted PROPELLER acquisition (6500/86, b = 800 sec/mm2) (c, d). The PROPELLER images still have a band of signal loss, but much more anatomy can be evaluated.

Figure 10d.

Figure 10d. Axial diffusion-weighted imaging in a patient with a right-side cochlear implant. Images of two section locations from a diffusion-weighted dataset acquired with echo-planar imaging (9200/81, b = 1000 sec/mm2) (a, b) are compared with images of the same two sections acquired with a diffusion-weighted PROPELLER acquisition (6500/86, b = 800 sec/mm2) (c, d). The PROPELLER images still have a band of signal loss, but much more anatomy can be evaluated.

Conclusion

Cochlear implants or ABIs with retained internal magnets present significant MR imaging challenges, particularly when the disease to be imaged is close to the implant. Extreme care must be used to position the patient for the MR imaging examination and to secure the implant when an internal magnet is included. The implant manufacturer’s MR imaging conditions for each individual implant model should be carefully followed, as subtle nuances in implant design may affect the likelihood of the internal magnet migrating or the components heating during the examination. Depending on the implant location and orientation, the patient’s perception of pain near the implant, and the need for a splint or head wrap, each patient has her or his own tolerance level for remaining in the imager. Not all patients with cochlear implants can tolerate the sensation of pressure at the site of their implant, and some may choose to abort the examination.

Moderate to severe artifacts are to be expected, both near the small internal magnet and displaced away from the magnet, depending on the image acquisition technique. Even with the help of a dedicated MR imaging physicist or scientist, the artifact will never be completely resolved. During surgical planning, experienced cochlear implant surgeons adjust the planned location of the implant with the hope of limiting artifact near disease that requires ongoing medical attention.

When imaging the head or neck near an implant with a retained magnet, STIR or Dixon techniques provide more-reliable fat suppression than CHESS fat-saturation pulses. If possible, parallel or accelerated imaging techniques that rely on coil sensitivities for spatial encoding should be avoided. Metal artifact–reduction techniques may help reduce the artifact but are unlikely to fully compensate for a magnetic material. Consider acquisitions other than echo-planar imaging to avoid gross geometric distortion and signal loss. Most importantly, when adjusting imaging parameters to reduce artifacts, always carefully check that imaging is still within the implant manufacturer’s MR imaging conditions for RF deposition, and avoid parameter changes that substantially alter the desired image contrast. Being prepared with strategies to reduce or move the B0 inhomogeneity artifact can shorten the duration of the examination and allow diagnostic images to be obtained.

Disclosures of Conflicts of Interest.— M.L.C.Activities related to the present article: disclosed no relevant relationships. Activities not related to the present article: personal fees from Advanced Bionics and MED-EL and nonfinancial support from Cochlear. Other activities: disclosed no relevant relationships.

Acknowledgment

The authors thank Sonia Watson, PhD, for her thoughtful assistance in manuscript preparation.

Presented as an education exhibit at the 2016 RSNA Annual Meeting.

For this journal-based SA-CME activity, the author M.L.C. has provided disclosures (see “Disclosures of Conflicts of Interest”); all other authors, the editor, and the reviewers have disclosed no relevant relationships.

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Article History

Received: May 15 2017
Revision requested: Aug 11 2017
Revision received: Sept 8 2017
Accepted: Sept 15 2017
Published online: Jan 10 2018
Published in print: Jan 2018