Artifacts on MRI images obtained in patients with metallic implants are produced by the large differences between the magnetic properties of human tissues and those of the implanted metals [14]. The artifacts are more marked when the differences in magnetic susceptibilities between the metallic object and the surrounding matter are substantial, creating local magnetic field inhomogeneities, altering the phase and frequency of local spins. Thus, the spins are subsequently mapped to an erroneous location within the image. The results are distortion of the shape of the metallic object along the axes of frequency encoding and section selection, and loss of signal within the metallic object. A rim of high signal intensity appears around the metallic object as a result of the mismapping of a disproportionate number of spins to that location [15, 16].
Factors that influence the production of metal-related artifacts at MRI imaging include the composition, size, and orientation of the metallic object with regard to the direction of the external magnetic field; the type of pulse sequences applied; and the sequence parameters, including magnetic field strength, voxel size (determined by field of view, image matrix, and section thickness), and echo train length. To demonstrate these factors, we performed MRI imaging in a phantom and a patient. Reduction of artifacts can be made paying attention to patient positioning, choosing adequate imaging parameters and selecting different pulse sequences.
Composition, size, and orientation
Titanium implants are non-ferromagnetic and produce much less severe artifacts than do ferromagnetic implants made of stainless steel [4, 16–19]. In addition, the artifact size is affected by the implant size with larger implants producing more obtrusive artifacts [18, 19].
When the direction of the main magnetic field (z axis of the scanner) is aligned parallel with the longitudinal axis of the hardware device, there is a significant relative reduction in artifact [4, 16, 18–21]. The limitations in patient positioning resulting from restrictive MRI scanner bore diameters can be overcome in magnets with an “open” configuration.
Pulse sequence selection
MRI degradation caused by metal hardware is primarily the result of a series of MRI artifacts produced by the ferromagnetic properties of the metallic device [22]). These artifacts include intravoxel dephasing, diffusion-related signal loss, slice thickness variation, misregistration artifacts, and inhomogeneous or paradoxical tissue-selective signal suppression with spectral (frequency-selective) fat-saturation techniques. The type and severity of metal-related artifacts are linked to the pulse sequence and operator selection of individual sequence parameters. Some MRI sequences are more susceptible to artifacts than others.
Gradient recalled echo (GRE) sequences are extremely sensitive to the presence of metal . Intravoxel dephasing is the predominant cause of signal loss on GRE imaging [23], resulting in a dark or black area (signal loss) around the metal on the processed images. Shortening the MRI parameter echo time (TE) and decreasing voxel size can be used to reduce the degree of intravoxel dephasing seen on GRE acquisition .
Other MRI sequences are less susceptible to intravoxel dephasing. Spin-echo (SE) sequences use refocusing 180° radiofrequency pulses that correct for static/fixed magnetic field inhomogeneities, resulting in a dramatic reduction in intravoxel dephasing signal loss . Misregistration artifacts in the vicinity of a metallic implant cause distinctive signal alterations, including signal voids and loss and signal increases resulting in hypointense and hyperintense signal artifacts around the implanted metal. Misregistration is a major source of metal-related signal artifacts on SE and FSE imaging sequences. The magnitude of local misregistration artifact is also inversely proportional to the frequency encoding (readout) gradient strength used in the imaging acquisition [25]. Misregistration artifacts occur only in the frequency encoding direction and are not seen in the phase encoding direction [4, 16, 26–28]. Orientation of the frequency and phase encoding gradients can be selected such that misregistration artifacts are directed away from areas of anticipated clinical diagnostic interest. The magnitude of misregistration artifact is inversely proportional to frequency encoding gradient strength, and increasing the frequency encoding gradient strength also may decrease these artifacts. This can be achieved practically on MRI acquisitions by widening or increasing the receiver bandwidth.
The presence of large magnetic field inhomogeneities adjacent to ferromagnetic materials also results in increased dephasing per unit distance of travel in randomly diffusing spinning protons. This artifact determines signal loss, represented as a dark or black region around the orthopedic hardware on the final images. This diffusion-related signal loss is most pronounced on long TE (T2-weighted) acquisition sequences and, in contrast to intravoxel dephasing signal loss, is not recoverable with an 180° refocusing pulse and affects GRE, SE, and FSE sequences [23, 24]. FSE imaging sequences employ multiple sequential refocusing 180° pulses in acquiring multiple lines of information (K-space data) within a single repetition time (TR): it refocuses spins at a faster interval than with conventional spin echo imaging, causing a small reduction in diffusion-related signal loss. This effect may be increased further, resulting in artifact reduction, with the use of smaller interecho spacing [16, 29, 30]. FSE also reduces artifact arising from malrotation [31], which is a relatively small contributor to signal loss around metal.
Frequency-selective fat saturation relies on the different resonance frequencies of hydrogen protons within water and fat. Fat signal suppression is achieved with the application of a narrow-bandwidth radiofrequency pulse limited to the spectral frequency of fat, and the magnetic field must be homogeneous within the imaging volume in order to obtain uniform fat suppression within the field of view. Variation of the regional magnetic field surrounding metallic devices or debris creates an inhomogeneous magnetic field, with resultant areas of suboptimal fat saturation [32]. Short inversion time inversion recovery (STIR) imaging is an effective alternative method of fat signal suppression and is less dependent on the homogeneity of the main magnetic field . The major disadvantage in STIR sequencing is decreased signal-to-noise ratio, resulting in a grainy appearance to the final images with loss of tissue signal resolution
An additional technique, view angle tilting, has been advocated more recently to decrease metallic artifacts [35, 36]. It technically involves application of a “compensatory gradient” during imaging acquisitions, correcting for inhomogeneous perturbations in the local magnetic field in the vicinity of a metallic device. It should be used in conjunction with increased receiver bandwidth, readout gradient strength, and reduced voxel size. View angle tilting results in image blurring across the entire imaging field of view, which must be partially compensated for by imaging parameter alterations (increases in the phase and frequency encoding gradients). The major drawback of this new MRI sequence is a resultant low image signal-to-noise ratio, although this has been shown not to compromise overall image interpretation [16, 35, 36].
There is an additional MRI method, single-point imaging (SPI). SPI is immune to the susceptibility artifacts observed with conventional MRI, which require several milliseconds for signal preparation and acquisition, failing to image metal materials because signal from these systems vanishes before it can be recorded completely. In SPI methods, this problem is solved by acquiring only one point of the free induction decay as soon as possible after excitation. SPI requires large gradient amplitudes and long scanning times; these problems are being addressed with good results [2, 16].
Sequence parameters
Slice thickness variations may be complex and account for significant signal loss and image distortion, particularly adjacent to metal devices with a complex three-dimensional geometry [26]. For imaging in the vicinity of a metal device, slice thickness should be minimized as much as possible because small voxel size in MRI in the vicinity of metal increases image quality. Small voxel size increases the spatial definition of metal-induced artifacts, reducing the apparent size of signal void present [4, 24, 25, 37]. As mentioned previously, smaller voxel size additionally may help decrease the degree of diffusion-related signal loss [16, 23, 24]. Voxel size may also be reduced without increasing examination time by increasing the number of frequency encoding steps used. Decreasing voxel size through this method is associated, however, with a tradeoff of progressive loss of image signal-to-noise ratio, which may require a compensatory increase in the number of excitations with associated increased imaging time.
Clinical applications of high-field strength MRI of evaluation of instrumented spine
Instrumented postero-lateral, or less frequently anterior, spinal fusion was performed with pedicular screws, longitudinal bars and cages to treat degenerative pathology (ostheoarthritis, spondylolisthesis, disc herniations, spinal stenosis), fractures, or severe scoliosis.
The major complications imaged are prosthesis loosening, periprosthetic fracture, and infection. Other complications that can be investigated by MRI include heterotopic ossification and prosthetic fracture. Adjacent or surrounding structures can also be evaluated, with common pathologies such as avascular necrosis, tumor recurrence, and internal derangement of joints potentially identified.
Prosthesis loosening shows fluid collections in the setting of mechanical loosening that had low T1 signal intensity and high T2 signal intensity surrounding the femoral stem with no enhancement. MRI findings of small particle disease consist of focal periprosthetic intraosseous masses. These masses are intermediate signal on proton density sequences [15]; low signal intensity on T1 sequences; and heterogeneous, predominately low to intermediate signal intensity, on T2 sequences with rim enhancement (occasional internal enhancement also can be seen) [23]. Pyogenic infection shows fluid signal on T1 and T2, peripheral rim enhancement, and surrounding edema (high T2 signal within adjacent tissues). Heterotopic ossification on MRI is visualized as areas or masses with well-circumscribed hypointense cortices and internal adipose signal intensity.
