Magnetic Resonance Imaging Acquisition Techniques for Radiotherapy Planning Gary P. Liney, PhD,*,† and Marinus A. Moerland, PhD‡ Magnetic resonance imaging (MRI) has a number of benefits for the planning of radiotherapy (RT), but its uptake into clinical practice has often been restricted to specialist research sites. There is often a lack of detailed MRI knowledge within the RT community and an apprehension of geometric distortions, both of which prevent its best utilization and merit the introduction of a standardized approach and common guidelines. This review sets out to address some of the issues involved in acquiring MRI scans for RT planning in the context of a number of clinical sites of interest and concludes with recommendations for its best practice in terms of imaging protocol and quality assurance. The article is of particular interest to the growing number of cancer therapy centers that are embarking on MRI simulation on either existing systems or their own dedicated scanners. Semin Radiat Oncol 24:160-168 C 2014 Elsevier Inc. All rights reserved.
Introduction
O
ver the last 20 years or so, there has been an exponential rise in the interest of using magnetic resonance imaging (MRI) in the radiation treatment planning of cancer. However, even with all the advantages MRI has to offer, there remains no consensus on how best to adopt this technology into clinical practice. In addition, a barrier to its development and implementation is the restricted access to MRI, usually reliant on the part-time use of radiology-based scanners that are not often suited to the specific requirements of treatment planning. There also is a lack of specialist MRI knowledge within the radiotherapy (RT) community and a mistrust of its geometric integrity that prevents the expedient use of these limited resources. For all of these reasons, a standardized approach to the acquisition of MRI in RT is warranted. This review sets out to address these issues both in the context of the increasing number of dedicated scanners and the more frequent ad hoc use of the existing diagnostic scanners. The aim of the article is
*Ingham Institute for Applied Medical Research, Liverpool, Sydney, New South Wales, Australia. †Department of Medical Physics, University of Wollongong, Wollongong, New South Wales, Australia. ‡Department of Radiation Oncology, University Medical Center Utrecht, Utrecht, The Netherlands. The authors declare no conflict of interest. Address reprint requests to Gary P. Liney, PhD, Ingham Institute for Applied Medical Research, 1, Campbell St, Liverpool, Sydney, New South Wales 2170, Australia. E-mail: [email protected]
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http://dx.doi.org/10.1016/j.semradonc.2014.02.014 1053-4296/& 2014 Elsevier Inc. All rights reserved.
to summarize the latest hardware and sequence developments for an RT audience and to point out specific aspects relating to the tumor sites of interest. It concludes with recommendations for the acquisition and quality assurance with the aim of increasing the use of this technology into RT practice.
Magnet Dedicated Systems Many of the early implementation studies with MRI were performed on low-field open systems that provided accessibility for RT immobilization equipment but which had poor image quality and homogeneity. These systems exhibited severe geometric distortions that have long lived in the mind of the RT community. Much more recently, the high-field open Panorama (Philips Medical Systems, Best, The Netherlands) was introduced, which offered a dedicated RT solution. This was a rebranding of existing technology aimed at interventional studies but was equally suited for RT with its wide lateral opening to allow imaging with fixation devices in situ. The 1.0-T magnet was augmented thanks to the vertical field geometry that allowed the use of solenoid-type radiofrequency (RF) coils and provided an increased signal-to-noise ratio (SNR) over what would be usually expected at this field strength. However, a major breakthrough in the last several years has been the introduction of conventional closed-tunnel designs with a wider 70-cm diameter bore. This extra space is sufficient to accommodate RT equipment while maintaining
MRI acquisition techniques for RT planning excellent gradient and magnet performance associated with higher field systems. Vendors have further equipped these systems with a flat tabletop and a posterior RF coil within the bed, making these scanners dedicated MRI treatment simulators (Fig. 1). As well as the technological advances in hardware, procurement of a dedicated MRI system within the RT department provides protected time slots for an increasing clinical workload as well as ongoing research development. A further consideration of a dedicated MR simulator is the installation of in-room goal-post lasers. Although these are helpful for accurately positioning the patient, their primary function is to replace the computed tomography (CT) simulator and as such will mainly find use in MR-only planning. However, installation at a later stage will require ramping down of the scanner. Finally, for staff often unfamiliar with the MRI environment, it is a good idea to mark out the 30-G line within the room, which is the proximity limit for unrestrained ferromagnetic objects. Although dedicated RT systems are certainly beneficial, it would be wrong to assume that MRI planning studies cannot be obtained using existing systems. All modern diagnostic scanners will benefit from the latest developments, which include analog to digital signal conversion at the magnet and optical transmission, which increases SNR. A high number of separate RF channels (currently 32 is typical, 64 is high end) further increases the SNR and the efficiency of the system. However, the most important considerations remain the patient positioning requirements that may restrict the standard 60-cm bore systems. Nevertheless, with simple modifications to the bed and utilization of appropriate diagnostic coils, planning in most tumor sites can be accommodated.
Field strength Most clinical systems, including the dedicated systems described earlier, are now available at both 1.5 and 3.0 T with
161 SNR increasing roughly linearly with field strength. It is anticipated that the market share of 1.5 and 3.0 T (currently around 60:40) is likely to change slightly toward higher field over the next decade with a small penetration of 7.0-T systems. For new installations, factors to consider when making the choice between field strengths are cost and how the technology will be utilized in practice. The benefit of improved soft-tissue delineation is well established at both field strengths. However, the doubling of SNR achieved at 3.0 T will be potentially important for functional imaging techniques and evaluating treatment effects. In their first incarnation, 3.0-T systems were troubled with dielectric artifacts caused by the reduced RF wavelength that creates a range of conductive and resonating effects within the patient. The resulting signal nonuniformity had previously been remedied by use of a “dielectric pad” placed on the patient. MRI vendors have now mitigated these artifacts, with the more sophisticated use of dual transmitters, meaning the full advantage of 3.0 T can now be exploited in all clinical sites. Higher field strength is beneficial in a number of other ways, notably frequency-selective fat suppression and MR spectroscopy owing to the increased chemical shift separation. However, a number of artifacts become worse1 and safety issues, such as RF heating, are of greater concern. For interventional applications, like MRI-guided brachytherapy, the choice of an ideal field strength also depends on the type of applicators, needles, or catheters to be inserted. Haack et al2 investigated applicator reconstruction for MRIbased brachytherapy for cervical cancer. Plastic and titanium applicators and needles were reconstructed with milimeter accuracy as verified on phantom CT images, although the titanium applicators and needles gave rise to susceptibility and RF artifacts in the 1.5-T MR images. These artifacts increase to 5-10 mm at 3.0 T.3 Another potential hazard for interventional applications is RF-induced heating of titanium devices. This heating is due to resonating electromagnetic RF waves along the conducting structure.4,5 The amount of heating is dependent on many factors, such as field strength, shape, and location of the conducting structure and dielectric properties of surrounding tissue, but generally the risk of heating increases with field strength. The ideal field strength for RT planning remains unproven and both 1.5- and 3-T field strengths are being pursued, but for interventional applications 1.5 T may be preferred.
Geometric Distortions System Distortions
Figure 1 A newly installed dedicated 3.0-T MR simulator (Cancer therapy centre, Liverpool, Sydney, Australia) showing the wide bore magnet (Siemens Skyra), external goal-post lasers and flat-indexed RT tabletop, which incorporates an RF coil underneath it. The 30-G line can also be seen marked out in the floor. (Color version of figure is available online.)
In MRI, spatial encoding is achieved through the combination of the static (B0) magnetic field to establish a signal of a characteristic resonance frequency and gradient coils that superimpose linear changes in this field to produce a positional dependence of frequency. Geometric distortions arise from either the inhomogeneity of the static field or the nonlinearity of the gradients, both of which become worse at increasing radial distances from the magnet isocentre. Investigations into image distortions date back from the 1980s and 1990s.6-10 In a
G.P. Liney and M.A. Moerland
162 European multicentre investigation, Lerski et al11 observed machine-dependent geometric distortions up to 5 mm and on 1 scanner up to 13 mm. The phantoms used in this study were Perspex disc-shaped phantoms, containing doped water and a grid of rods. Bakker et al12 also used grid phantoms to study distortions and they showed that in spin-echo and gradientecho imaging, gradient errors and static field inhomogeneity lead to image distortions in the direction of the frequencyencoding gradient, whereas in the phase-encoding gradient image distortions are solely caused by gradient errors. Scannerrelated distortions were measured by variation of the strength and direction of the readout gradient in imaging experiments on a grid of cylindrical sample tubes. For the 1.5-T system used in that study, static field–related errors up to 7 mm and gradient-related errors up to 4 mm were observed in a midcoronal plane within a field of view (FOV) of 40 cm. Static magnetic field–related errors were inversely proportional to gradient strength, whereas gradient magnetic field–related errors were independent of gradient strength. These observations of image distortions may have retarded the introduction of MRI into RT treatment planning. However, for a greater part these distortions are machine specific, they can be measured, they are stable in time, and thus predictable.13 On a modern scanner, typical errors would be expected to be less than 1 and 2 mm within a radius of 10 and 15 cm, respectively.14 At greater distances from the isocenter, the distortions are larger, but these machine (gradient)-specific distortions can be corrected and nowadays most MR scanners have 2-dimensional (2D) (and some 3D) image correction available. Doran et al15 showed that distortions in a large FOV (320 200 340 mm3) on a 1.5-T Siemens Vision scanner could be corrected to within the voxel resolution.
Patient-Related Distortions Magnetic field errors stem not only from imperfections of the MR scanner but also from the patient owing to the chemical shift and susceptibility phenomena. The chemical shift effect is caused by the magnetic shielding of a nucleus by the electron clouds that surround it. The resonant frequency of a nucleus depends on the chemical structure of a molecule and the location of the nucleus within the molecule. In human MRI, the relevant chemical shift is between protons in fat and water. Protons in fat experience a magnetic field, which is 3.5 ppm weaker than the magnetic field experienced by protons in water. The consequence for MRI is that the fat-containing tissues will be shifted with respect to the other tissues in the direction of the frequency-encoding gradient. The human body magnetic susceptibility distribution is a major factor perturbing the magnetic field. The magnetic susceptibility is the degree of magnetization of a material in response to the applied magnetic field. In the human body, the susceptibility differences and resulting image distortions are largest at tissueair interfaces. Magnetic field distributions and image distortions for simple objects, such as cylinders and spheres, can be solved analytically,16-18 whereas the magnetic field in and around the patient has to be calculated with numerical methods. Bhagwandien et al19 developed and validated a
method to calculate the magnetic field distribution in and around a 3D object using a technique based on the finite difference method. Field perturbations between 5 and 6 ppm were observed outside the head and between 6 and 5 ppm inside the head near the sinuses. In the neck area, field deviations ranged from 3 to 6 ppm inside and from 5 to 4 ppm outside the body.20 These deviations result in image distortions between 5 and 5 mm for a typical frequencyencoding gradient strength (1.5 mT/m at 1.5 T). Patient-related image distortions are not predictable and have to be calculated on an individual basis. Correction methods based on calculations of the field perturbations are not easy to implement in clinical practice. Chang and Fitzpatrick21 introduced the gradient-reversal technique, which combines 2 images of the same object acquired with a forward and a reversed readout gradient, from which an undistorted image can be postprocessed. Reinsberg et al22 improved this method by using maximization of mutual information for registration of the forward and reversegradient images, eliminating the need for user interaction. Jezzard and Balaban23 introduced a correction method based on measurement of the field map. Echo-planar images collected at 1.5 and 4 T on phantoms and human subjects were unwarped, resulting in images with subpixel accuracy under the assumption that only field in homogeneities account for all distortions (no eddy currents, etc). Another approach could be to acquire undistorted images with pulse sequences that apply only phase encoding of the MR signal.24 However, phase encoding is time consuming and only used in 1 or 2 dimensions in clinical sequences. Although correction methods for patient-related image distortions are available, they are not yet widely incorporated in scanner software. In practice, one could adapt the following approach. Patient-related magnetic field perturbations stem from the chemical shift and the susceptibility distribution, both of which result in local distortions of the applied magnetic field. The corresponding image distortions are proportional to the local magnetic field error and inversely proportional to the strength of the applied frequency-encoding gradient. The chemical shift between water and fat protons is approximately 3.5 ppm (220 Hz at 1.5 T and 440 Hz at 3 T), and susceptibilityinduced perturbations are at the most about twice as large. Patient-related image distortions can be reduced by applying relatively strong slice selection and frequency-encoding gradients, for example, by setting the pixel bandwidth at twice the water-fat shift, all patient-related distortions are expected to be smaller than the pixel size. In regions where anatomy is expected to be problematic, localized shimming over the intended imaging volume should be performed before the scan (noting the possibility of worsening distortions elsewhere).
Image Acquisition Patient Setup A major constraint on MR acquisition is the patient positioning or immobilization or both used to represent the treatment
MRI acquisition techniques for RT planning setup and facilitate registration with the planning CT scan. The images must be representative of the anatomy during each fraction. If deformation exists, approximating the setup as much as possible partially mitigates this issue. Although imaging in this fixed position is unnecessary in brain, it has been shown to be necessary in most sites (eg, head and neck to control neck flexion25) and may even dictate a change of imaging position for others (eg, breast and rectum). The first issue is the curvature of the patient table present on most diagnostic scanners. This can be removed using commercial tabletops or home-made flat boards, which index with RT base plates. For head and neck, a normal scanner bore diameter may prevent the use of conventional immobilization devices warranting a cut to the fixation shells or the use of alternative means to hold the patient in position. Studies in the prostate have also demonstrated the advantages of maintaining the hips in a flat position, which is crucial in not only maintaining bony anatomy but has also been shown to effect target delineation.26 MR-compatible ankle stocks and knee rests should be utilized wherever possible. These positional restrictions are less relevant if registration and guidance is based on implanted fiducial markers. For the breast, the treatment setup in a supine position and a wing board will be difficult to achieve without the use of a wide bore scanner. It is important to verify the MR compatibility of all immobilization equipment that includes both safety and image quality considerations. Carbon fiber materials should be avoided or tested rigorously for the possibility of RF heating.
RF Coils RF coils are designed to transmit and detect the MR signal either separately or in a combined transceiver fashion. Important characteristics of RF coils are their sensitivity profile (uniformity) and SNR. The design of RF coils is such that the sensitive volume is restricted to the anatomy of interest to reduce the detectable noise from the patient. In the case of the brain, standard multichannel head coils can be used to provide optimum image quality. However, in other sites, any attempt to cater for immobilization and replicate treatment positioning usually results in some compromise in image quality or coverage or both. Nevertheless, groups have managed to use a variety of diagnostic surface coils placed around fixation shells in the brain27 and head and neck28-30 and suspended above the abdomen in prostate31 to obtain images of sufficient quality for planning purposes. In the breast, the diagnostic preference is to image the prone patient, with the breasts suspended in the wells of a dedicated coil array. Supine scans may be performed but the choice of the treatment over scan position is currently an area of active research.32 When no other option is available, the integrated body coil (hidden behind the plastic magnet covers) may be used to accommodate treatment positions more easily at the expense of a significant reduction of SNR.30 Figure 2 illustrates some of these positional challenges and trade-offs. On the left, 2 head and neck scans have been overlaid, the first acquired using a RT neck rest to control positioning and the second in an arbitrary position in the diagnostic coil. On the right, 2 prostate
163 scans are compared; the first acquired in the treatment position using a flat tabletop using the integrated body coil and the second using the diagnostically preferred body array. The diagnostic scan demonstrates increased SNR but also slight internal anatomy shifts compared with the treatment position scan. Vendors are increasingly providing dedicated coil solutions for RT imaging. In head and neck, this comprises laterally placed coils that can be combined with an additional coil on top of the chest or underneath the neck or both to extend coverage in the craniocaudal direction. Bespoke coil positioning aids and adjustable supports are being introduced onto the market that hold coils in place, and in the case of the abdomen maintain close contact with the patient without compressing the skin surface, which will be important for MR-only planning (Fig. 3). Images obtained with the surface coil arrangements described previously will typically exhibit poor signal uniformity and usually some method of intensity correction is required, and vendors offer this either prospectively or retrospectively. Wherever possible, the RF coil should be multichannel (ie, individual signal elements) so that the advantages of parallel imaging can be exploited.
Scan Protocol The aim of the RT planning scan is different from that of a diagnostic protocol where usually a slice thickness of 3-5 mm is adequate and around 30-40 slices are acquired. For the RT scan, a more detailed visualization of the tumor and organs at risk is required and coverage needs to be sufficient for registration with CT. To maintain accuracy, it is desirable to reduce slice thickness to match that of corresponding CT scans and should be at most 2-3 mm depending on SNR. Coverage of between 18 and 20 cm may be required in the superiorinferior direction meaning the acquisition of 60-100 images. The consequence of this is long scan times that may result in patient or organ movement and a degradation of image quality. Motion within an acquisition is characterized by ghosting, which appears in the phase-encoding direction and can vary from subtle smearing of signal to gross artifacts. In cases where the 2D acquisition is divided over multiple concatenations to cover the required anatomy, contiguous slices can be several minutes apart leading to a bizarre steplike appearance on the reconstructions when significant movement has occurred (Fig. 4). Beyond the first step of motion management during the acquisition, a simple way to remedy this is to extend the repetition time to encompass the required number of slices in an acquisition albeit at an increased time penalty. An alternative is to use 3D sequences that excite an entire volume of tissue and use phase encoding in the “slice” direction. These acquisitions can provide isotropic resolution but can also extend scan time significantly. However, there are newer faster versions (see section Pulse sequences) that show considerable promise for RT planning, although their contrast properties have been noted to be subtly different.33 Images will usually be acquired in the axial plane, a requirement often determined by the treatment planning system (TPS), although some newer platforms are able to cope
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Figure 2 The differences between diagnostic and radiotherapy planning scans (left). Overlay of H&N images acquired in a standard diagnostic RF coil and also in a bespoke surface coil in the treatment position that demonstrates the large difference in neck flexion. Also shown are prostate images acquired using the integrated body coil and flat tabletop (center) compared with a corresponding image acquired with the preferred diagnostic posterior and anterior coil array (right). The SNR is better with the diagnostic coil, but the position of the internal anatomy is slightly different. H&N, head and neck. (Color version of figure is available online.)
with nonaxial data sets. The FOV should also be selected with consideration of the effects of geometric distortions at large distances from the isocenter, while providing enough anatomy for registration and contouring of soft tissues. On a typical closed-bore scanner, this would usually be between 20 and 25 cm, which covers the brain and head and neck. In prostate, this will be large enough to encompass femoral heads and adjacent bony landmarks (Fig. 5A), although if gold seeds are present then the FOV can be reduced or automated registration should be targeted to this smaller volume or both. Beyond this FOV, distortion will be present (eg, at the skin surface in supine breast imaging) reducing the reliability of the CT-MR-fused image in these regions. In-plane resolution should be 1 mm or better with receiver bandwidths adjusted accordingly to limit patient distortions to within a single pixel.
Pulse Sequences Unlike CT, where the limitations in scan protocol at least provide some consistency, MRI offers the user with a wide
variety of contrast weightings and sequences, each described by a myriad of vendor-specific abbreviations and acronyms. Fundamentally, these can be broadly classified into spin-echo- and gradient-echo-based sequences with the former providing the gold standard in terms of intensity distortion, and the latter being inherently faster but more liable to signal dephasing. A routine RT planning examination should be expected to provide a robust geometrically accurate data set, which will usually consist of a fast spin-echo variant with or without an additional contrast-enhanced sequence.14 A more sophisticated planning protocol may include the addition of at least one of the functional techniques discussed later. Gradient-echo imaging should be avoided in patients with metallic applicators or prosthesis. Metal artifact reduction techniques, essentially using high receiver bandwidths and short echo times, should be employed. In other instances, for example, gold fiducials, these artifacts play an essential role in their visualization but should be controlled to maintain a faithful representation of their true location especially at
Figure 3 Two commercially available RF coil solutions for radiotherapy. (Left) A head and neck suite of coils, comprising 2 lateral surface coils (6 channel) plus a 16-channel anterior coil attachment with bespoke positioning aids. When combined with a dedicated radiotherapy posterior coil, it provides 30 channels and full FOV H&N images in the treatment position (GE Healthcare, Milwaukee, WI). (Right) An 18-channel anterior body array (Siemens Healthcare, Erlangen, Germany) with coil bridges (Civco, Kalona, IA) to stand the coil off the patient. A 32-channel coil is also underneath the tabletop. H&N, head and neck. (Color version of figure is available online.)
MRI acquisition techniques for RT planning
Figure 4 This sagittally reconstructed prostate scan clearly shows motion has occurred between the 2 concatenated axial acquisitions, demonstrating a characteristic discontinuity between alternate slices as indicated by the contoured structures. (Color version of figure is available online.)
higher field strength.34 Figure 5B demonstrates the appearance of gold seeds at 3.0 T for both a gradient-echo and fast spin-echo sequence. The gradient-echo sequence shows the seed well but the artifact is larger than the true seed diameter. In this case, a short (o10 milliseconds) echo time fast spin-echo sequence was used to improve the visualization of the seed against the tissue and demonstrate a more accurate marker position. Over the preceding years, some interesting sequences and techniques have been introduced that offer a number advantages for RT planning (Table). These include speeded up implementations of 3D fast spin-echo sequences, eliminating the need for multiple acquisitions and providing data sets with isotropic resolution in reasonable scan times.35 The resolution of these scans means they can be acquired in any favorable plane to further reduce scan time or partial voluming, before subsequent reformatting into axial data for import into the TPS. Radial k-space is another recent advancement whereby
165 the trajectory of data collection is altered in a semirandom manner. This provides an effective remedy to ghosting artifacts that may occur from patient movement or internal motion (eg, bladder filling). Perhaps one of the most important developments has been parallel imaging, which reduces scan times by missing out some gradient-encoding steps and using instead the spatial information from RF coil sensitivity profiles. This is an invaluable approach and avoids the alternative of increasing slice thickness or slice gap or both to shorten scan times. Different implementations work from either the image or the raw (k-space) data and many are now self-calibrating. Theoretically, the speed-up factor is equal to the number of available coil elements, although factors of 2-3 are more typical. In- and out-of-phase imaging permits 4 separate images based on fat and water contrast to be acquired simultaneously; the wateronly image is particularly effective in providing a fat-suppressed sequence in areas of field inhomogeneity and as such is a useful tool for the RF surface coils used in RT planning. Although vendors largely offer similar products, it has been reported that variations in sequences across different systems have been shown to produce variability in tumor delineation.36 Functional imaging techniques are a particularly attractive proposition for RT planning, and a more complete review of their role is provided elsewhere.37 Traditionally, these techniques have introduced the added difficulty of some or all of the following problems; inherent sequence-specific distortion, image manipulation, and data interpretation. Diffusionweighted imaging utilizes strong bipolar gradients to sensitize the sequence to microscopic molecular diffusion and is used to measure the apparent diffusion coefficient (ADC), which is related to cellularity. These data are routinely acquired using echo-planar imaging–based sequences, which are inherently prone to artifacts and geometric distortions. New sequences either with or without radial k-space sampling have been shown to improve image quality over standard sequences and seem promising in an RT setting.38,39 MR spectroscopy has the potential to add metabolic information into the RT plan,40 but a substantial degree of expertise is required to acquire and correctly interpret the data, and the incorporation of this
Figure 5 (A) The limited coverage and FOV used in this prostate scan (color overlay) compared with the planning CT data. (B) Two different sequences showing gold seeds within the prostate at 3.0 T; on the left, a gradient-echo sequence shows the marker clearly owing to large susceptibility artifact. The fast spin-echo on the right shows it less well but with better precision.
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Table Acronyms for Pulse Sequences and Techniques From 3 MRI Vendors That are Particularly Relevant to RT Planning Generic Name
GE
Philips
Siemens
Details
3D FSE with variable flip angle Uniformity correction In- and out-ofphase imaging Parallel imaging
CUBE (previously known as XETA)
VISTA
SPACE
High-resolution volumetric imaging (particularly T2-w) in reduced scan times.
SCIC or PURE
CLEAR
Reduces intensity variation from surface coil acquisitions.
IDEAL
mDIXON
Prescan normalize DIXON
Radial k-space
PROPELLER
ASSET or ARC
The water-only scan provides excellent fat suppression image in presence of B0 and B1 inhomogeneity. SENSE mSENSE or Scan time reduction based on image or k-space data. Some selfGRAPPA calibrating others require additional scan. MultiVane BLADE Motion artifact reduction.
Abbreviations: ARC, autocalibrating reconstruction; ASSET, array sensitive encoding technique; CLEAR, constant level appearance; FSE, fast spinecho; GRAPPA, generalized autocalibrating partially parallel acquisition; IDEAL, iterative decomposition of water and fat with echo asymmetry and least squares estimation; PROPELLER, periodically rotated overlapping parallel lines with enhanced reconstruction; PURE, phased array uniformity enhancement; SCIC, surface coil intensity correction; SENSE, sensitivity encoding; SPACE, sampling perfection with applicationoptimized contrasts using different flip angle evolution; VISTA, volume isotropic t2 weighted acquisition.
information into the planning system is still in its infancy. Measurement of relaxation times, for example, R2n that has been studied in relation to hypoxia,41 has usually required some additional off-line postprocessing to produce the parameter maps. However, MRI vendors are increasingly providing software, which caters to the analysis and fusion of all these secondary data sets and must continue to do so in the context of RT, with the ability to create contours drawn on these and export them into planning systems.
Quality Assurance Before the commencement of any RT planning service, a thorough commissioning of a newly installed system or the quality assurance of a pre-existing system should be undertaken. These tests should cover all the normal quality assurance metrics that are described in detailed guidelines elsewhere.42 In addition, a series of tests specifically designed for RT planning are recommended. These will principally concern the assessments of geometric distortion and the image quality of the patient setup.28
Several commercial phantoms are available to test geometric distortion.42-44 These are usually limited in size in at least 1 dimension, and it may be necessary to build an inhouse test object to fully characterize a system over a large volume. It is worth making measurements both with and without the aforementioned gradient corrections turned on. Figure 6 shows example images acquired on a 1.5-T Philips Achieva MR scanner (Philips Medical Systems, Best, The Netherlands) with distortions before and after correction. Within a FOV of 336 mm, the software reduced distortions of 6.0 to 0.2 mm. Surface coils should be investigated for their SNR and uniformity characteristics using spherical- or cylindricalshaped flood-fill phantoms of appropriate size. Measurements should be performed both with and without any coilcorrection filters or algorithms applied and benchmarked against the diagnostic coil of choice.27,28 Finally, an end-to-end test may be included with the aim of testing the optimum MRI protocol across each step of the RT treatment chain through to registration with CT on the TPS. Any suitable CT- and MR-compatible test object
Figure 6 Images of a geometric distortion phantom acquired on a 1.5-T Philips Achieva scanner before and after gradient correction was applied. The nominal distance between the 2 columns of tubes indicated by the lines is 210 mm. The arrows depict the distances between the outer tubes on the bottom row.
MRI acquisition techniques for RT planning will suffice but more complicated anthropomorphic phantoms are worthy of some consideration as this will enable an examination of patientlike effects, such as susceptibility and water-fat shifts.45
Conclusions Over the next few years, it is expected that there will be an increase in the number of dedicated MR simulators installed within RT departments. However, most work will continue to be performed on pre-existing systems. By adopting a pragmatic approach, most MRI scanners can be used to provide a planning service in some capacity. With this in mind the following recommendations are suggested: (1) The field strength should be no less than 1.5 T and no greater than 3.0 T. (2) The magnet should be a closed-tunnel configuration, preferably with a wide bore. (3) A flat bed or tabletop insert should be used. (4) Appropriate RF coils (multichannel) should be used that cater to immobilization devices and surface coil intensity correction should always be employed. (5) Geometric distortion should be verified in all 3 directions, and this will inform the useable FOV (in plane) and slice coverage (through plane). (6) Spin-echo-based sequences (ie, fast or turbo spinecho) should be used in preference to gradient-echo for the gold-standard anatomical image. (7) Slice thickness should match the planning CT acquisition where possible. (8) Multiple (ie, concatenated) acquisitions should be examined for gross movement between adjacent slices and avoided if motion cannot be mitigated. (9) The scanner's software for correction of the gradient errors should always be turned on. (10) Pixel bandwidth should be set to a minimum of twice the water-fat shift (ie, 220 Hz at 1.5 T and 440 Hz at 3 T). (11) QA of the optimum patient setup and imaging protocol should be carried out before clinical scans. These recommendations will be helpful in standardizing the acquisition across multiple patients and different centers to ensure a more robust acceptance of MRI in RT.
Acknowledgments The authors thank the staff at the MRI department at Castle Hill hospital, Hull, UK, and especially Lynnette Cassappi and Ewa Juresik (Liverpool CTC, Sydney, Australia) for their assistance in acquiring images used in this article.
References 1. Dietrich O, Reiser MF, Schoenberg SO: Artifacts in 3-Tesla MRI: Physical background and reduction strategies. Eur J Radiol 65(1):29-35, 2008
167 2. Haack S, Nielsen SK, Lindegaard JC, et al: Applicator reconstruction in MRI 3D image-based dose planning of brachytherapy for cervical cancer. Radiother Oncol 91(2):187-193, 2009 3. Kim Y, Muruganandham M, Modrick JM, et al: Evaluation of artefacts and distortions of titanium applicators on 3.0-Tesla MRI: Feasibility of titanium applicators in MRI-guided brachytherapy for gynecological cancer. Int J Radiat Oncol Biol Phys 80:947-955, 2011 4. Yeung CJ, Karmarkar P, McVeigh ER: Minimizing RF heating of conducting wires in MRI. Magn Reson Med 58(5):1028-1034, 2007 5. van den Bosch MR, Moerland MA, Lagendijk JJ, et al: New method to monitor RF safety in MRI-guided interventions based on RF induced image artefacts. Med Phys 37(2):814-821, 2010 6. Sekihara K, Kuroda M, Kohno H: Image restoration from non-uniform magnetic field influence for direct Fourier NMR imaging. Phys Med Biol 29(1):15-24, 1984 7. O'Donnell M, Edelstein WA: NMR imaging in the presence of magnetic field inhomogeneities and gradient field nonlinearities. Med Phys 12 (1):20-26, 1985 8. Kawanaka A, Takagi M: Estimation of static magnetic field and gradient fields from NMR image. J Phys E Sci Instrum 19:871-875, 1986 9. Fraass BA, McShan DL, Diaz RF, et al: Integration of magnetic resonance imaging into radiation therapy treatment planning: I. Technical considerations. Int J Radiat Oncol Biol Phys 13(12):1897-1908, 1987 10. Cho ZH, Kim DJ, Kim YK: Total inhomogeneity correction including chemical shifts and susceptibility by view angle tilting. Med Phys 15 (1):7-11, 1988 11. Lerski RA, McRobbie DW, Straughan K, et al: Multi-centre trial with protocols and prototype test objects for the assessment of MRI equipment. EEC Concerted Research Project. Magn Reson Imaging 6(2):201-214, 1988 12. Bakker CJ, Moerland MA, Bhagwandien R, et al: Analysis of machinedependent and object-induced geometric distortion in 2DFT MR imaging. Magn Reson Imaging 10(4):597-608, 1992 13. Moerland MA: Magnetic resonance imaging in radiotherapy treatment planning, Thesis University Utrecht, 1996 14. Brunt JN: Computed tomography-magnetic resonance image registration in radiotherapy treatment planning. Clin Oncol (R Coll Radiol) 22:688-697, 2010 15. Doran SJ, Charles-Edwards L, Reinsberg SA, et al: A complete distortion correction for MR images: I. Gradient warp correction. Phys Med Biol 50 (7):1343-1361, 2005 16. Lüdeke KM, Röschmann P, Tischler R: Susceptibility artefacts in NMR imaging. Magn Reson Imaging 3(4):329-343, 1985 17. Chu SC, Xu Y, Balschi JA, et al: Bulk magnetic susceptibility shifts in NMR studies of compartmentalized samples: Use of paramagnetic reagents. Magn Reson Med 13(2):239-262, 1990 18. Bakker CJ, Bhagwandien R, Moerland MA, et al: Susceptibility artifacts in 2DFT spin-echo and gradient-echo imaging: The cylinder model revisited. Magn Reson Imaging 11(4):539-548, 1993 19. Bhagwandien R, Moerland MA, Bakker CJ, et al: Numerical analysis of the magnetic field for arbitrary magnetic susceptibility distributions in 3D. Magn Reson Imaging 12(1):101-107, 1994 20. Bhagwandien R: Object induced geometry and intensity distortions in magnetic resonance imaging, Thesis University Utrecht, 1994 21. Chang H, Fitzpatrick JM: A technique for accurate magnetic resonance imaging in the presence of field inhomogeneities. IEEE Trans Med Imaging 11:319-329, 1992 22. Reinsberg SA, Doran SJ, Charles-Edwards EM, et al: A complete distortion correction for MR images: II. Rectification of static-field inhomogeneities by similarity-based profile mapping. Phys Med Biol 50(11):2651-2661, 2005 23. Jezzard P, Balaban RS: Correction for geometric distortion in echo planar images from B0 field variations. Magn Reson Med 34(1):65-73, 1995 24. Crijns SP, Bakker CJ, Seevinck PR, et al: Towards inherently distortionfree MR images for image-guided radiotherapy on an MRI accelerator. Phys Med Biol 57(5):1349-1358, 2012 25. Hanvey S, McJury M, Tho LM, et al: The influence of MRI scan position on patients with oropharyngeal cancer undergoing radical radiotherapy. Radiat Oncol 8(1):129, 2013
168 26. Hanvey S, Sadozye AH, McJury M, et al: The influence of MRI scan position on image registration accuracy, target delineation and calculated dose in prostatic radiotherapy. Br J Radiol 85:e1256-e1262, 2012 27. Hanvey S, Glegg M, Foster J: Magnetic resonance imaging for radiotherapy planning of brain cancer patients using immobilization and surface coils. Phys Med Biol 54:5381-5394, 2009 28. Liney GP, Owen SC, Beaumont AKE, et al: Commissioning of a new widebore MRI scanner for radiotherapy planning of head and neck cancer. Br J Radiol 86:20130150, 2013 29. Webster GJ, Killgallon JE, Ho KF, et al: A novel imaging technique for fusion of high-quality immobilised MR images of the head and neck with CT scans for radiotherapy target delineation. Br J Radiol 82:497-503, 2009 30. Verduijn GM, Bartells LW, Raaijmakers CPJ, et al: Magnetic resonance imaging protocol optimization for delineation of gross tumour volume in hypopharyngeal and laryngeal tumours. Int J Radiat Oncol Biol Phys 74:630-636, 2009 31. McJury M, O'Neill A, Lawson M, et al: Assessing the image quality of pelvic MR images acquired with a flat couch for radiotherapy treatment planning. Br J Radiol 84:750-755, 2011 32. Giezen M, Kouwenhoven E, Scholten AN, et al: Magnetic resonance imaging-versus computed tomography-based target volume delineation of the glandular breast tissue (clinical target volume breast) in breast conserving therapy: An exploratory study. Int J Radiat Oncol Biol Phys 81:804-811, 2011 33. Ristow O, Steinbach L, Sabo G, et al: Isotropic 3D fast spin-echo imaging versus standard 2D imaging at 3.0 T of the knee: Image quality and diagnostic performance. Eur Radiol 19:1263-1272, 2009 34. Jonsson JH, Garpebring A, Karlsson MG, et al: Internal fiducial markers and susceptibility effects in MRI-simulation and measurement of spatial accuracy. Int J Radiat Oncol Biol Phys 82:1612-1618, 2012
G.P. Liney and M.A. Moerland 35. Proscia N, Jaffe TA, Neville AM, et al: MRI of the pelvis in women: 3D versus 2D T2-weighted technique. Am J Roentgenol 195:254-259, 2010 36. Nyholm T, Jonsson J, Söderström K, et al: Variability in prostate and seminal vesicle delineations defined on magnetic resonance images, a multi-observer, -center and -sequence study. Radiat Oncol 8(1):126, 2013 37. Metcalfe P, Liney GP, Holloway L, et al: The potential role for an enhanced role for MRI in radiation-therapy treatment planning. Technol Cancer Res Treat 2013doi:10.7785/tcrt.2012.500342 38. Babourina-Brooks B, Cowin GJ, Wang D: Diffusion-weighted imaging in the prostate: An apparent diffusion, coefficient comparison of half-Fourier acquisition single-shot turbo spin-echo and echo planar imaging. Magn Reson Imaging 30:189-194, 2011 39. Schwartz KM, Lane JI, Bolster BD, et al: The utility of diffusion-weighted imaging for cholesteatoma evaluation. Am J Neuroradiol 32:430-436, 2011 40. Payne GS, Leach MO: Applications of magnetic resonance spectroscopy in radiotherapy treatment planning. Br J Radiol 79:S16-S26, 2006 41. Hoskin PJ, Carnell DM, Taylor NJ, et al: Hypoxia in prostate cancer: Correlation of BOLD-MRI with pimonidazole immunohistochemistryinitial observations. Int J Radiat Oncol Biol Phys 68:1065-1071, 2007 42. AAPM. Quality assurance methods and phantoms for magnetic resonance imaging: Report of AAPM nuclear magnetic resonance Task Group No. 1. Med Phys 17(2):287-295, 1990 43. Gunter JL, Bernstein MA, Schuff BJ, et al: Measurement of MRI scanner performance with the ADNI phantom. Med Phys 36:219, 2009 44. http://www.leedstestobjects.com. 45. Moore CS, Liney GP, Beavis AWA: Quality assurance phantom for verification and registration of CT and MRI data sets for treatment planning of radiotherapy for head and neck cancers. J Appl Clin Med Phys 22:25-35, 2004
Introduction
O
ver the last 20 years or so, there has been an exponential rise in the interest of using magnetic resonance imaging (MRI) in the radiation treatment planning of cancer. However, even with all the advantages MRI has to offer, there remains no consensus on how best to adopt this technology into clinical practice. In addition, a barrier to its development and implementation is the restricted access to MRI, usually reliant on the part-time use of radiology-based scanners that are not often suited to the specific requirements of treatment planning. There also is a lack of specialist MRI knowledge within the radiotherapy (RT) community and a mistrust of its geometric integrity that prevents the expedient use of these limited resources. For all of these reasons, a standardized approach to the acquisition of MRI in RT is warranted. This review sets out to address these issues both in the context of the increasing number of dedicated scanners and the more frequent ad hoc use of the existing diagnostic scanners. The aim of the article is
*Ingham Institute for Applied Medical Research, Liverpool, Sydney, New South Wales, Australia. †Department of Medical Physics, University of Wollongong, Wollongong, New South Wales, Australia. ‡Department of Radiation Oncology, University Medical Center Utrecht, Utrecht, The Netherlands. The authors declare no conflict of interest. Address reprint requests to Gary P. Liney, PhD, Ingham Institute for Applied Medical Research, 1, Campbell St, Liverpool, Sydney, New South Wales 2170, Australia. E-mail: [email protected]
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to summarize the latest hardware and sequence developments for an RT audience and to point out specific aspects relating to the tumor sites of interest. It concludes with recommendations for the acquisition and quality assurance with the aim of increasing the use of this technology into RT practice.
Magnet Dedicated Systems Many of the early implementation studies with MRI were performed on low-field open systems that provided accessibility for RT immobilization equipment but which had poor image quality and homogeneity. These systems exhibited severe geometric distortions that have long lived in the mind of the RT community. Much more recently, the high-field open Panorama (Philips Medical Systems, Best, The Netherlands) was introduced, which offered a dedicated RT solution. This was a rebranding of existing technology aimed at interventional studies but was equally suited for RT with its wide lateral opening to allow imaging with fixation devices in situ. The 1.0-T magnet was augmented thanks to the vertical field geometry that allowed the use of solenoid-type radiofrequency (RF) coils and provided an increased signal-to-noise ratio (SNR) over what would be usually expected at this field strength. However, a major breakthrough in the last several years has been the introduction of conventional closed-tunnel designs with a wider 70-cm diameter bore. This extra space is sufficient to accommodate RT equipment while maintaining
MRI acquisition techniques for RT planning excellent gradient and magnet performance associated with higher field systems. Vendors have further equipped these systems with a flat tabletop and a posterior RF coil within the bed, making these scanners dedicated MRI treatment simulators (Fig. 1). As well as the technological advances in hardware, procurement of a dedicated MRI system within the RT department provides protected time slots for an increasing clinical workload as well as ongoing research development. A further consideration of a dedicated MR simulator is the installation of in-room goal-post lasers. Although these are helpful for accurately positioning the patient, their primary function is to replace the computed tomography (CT) simulator and as such will mainly find use in MR-only planning. However, installation at a later stage will require ramping down of the scanner. Finally, for staff often unfamiliar with the MRI environment, it is a good idea to mark out the 30-G line within the room, which is the proximity limit for unrestrained ferromagnetic objects. Although dedicated RT systems are certainly beneficial, it would be wrong to assume that MRI planning studies cannot be obtained using existing systems. All modern diagnostic scanners will benefit from the latest developments, which include analog to digital signal conversion at the magnet and optical transmission, which increases SNR. A high number of separate RF channels (currently 32 is typical, 64 is high end) further increases the SNR and the efficiency of the system. However, the most important considerations remain the patient positioning requirements that may restrict the standard 60-cm bore systems. Nevertheless, with simple modifications to the bed and utilization of appropriate diagnostic coils, planning in most tumor sites can be accommodated.
Field strength Most clinical systems, including the dedicated systems described earlier, are now available at both 1.5 and 3.0 T with
161 SNR increasing roughly linearly with field strength. It is anticipated that the market share of 1.5 and 3.0 T (currently around 60:40) is likely to change slightly toward higher field over the next decade with a small penetration of 7.0-T systems. For new installations, factors to consider when making the choice between field strengths are cost and how the technology will be utilized in practice. The benefit of improved soft-tissue delineation is well established at both field strengths. However, the doubling of SNR achieved at 3.0 T will be potentially important for functional imaging techniques and evaluating treatment effects. In their first incarnation, 3.0-T systems were troubled with dielectric artifacts caused by the reduced RF wavelength that creates a range of conductive and resonating effects within the patient. The resulting signal nonuniformity had previously been remedied by use of a “dielectric pad” placed on the patient. MRI vendors have now mitigated these artifacts, with the more sophisticated use of dual transmitters, meaning the full advantage of 3.0 T can now be exploited in all clinical sites. Higher field strength is beneficial in a number of other ways, notably frequency-selective fat suppression and MR spectroscopy owing to the increased chemical shift separation. However, a number of artifacts become worse1 and safety issues, such as RF heating, are of greater concern. For interventional applications, like MRI-guided brachytherapy, the choice of an ideal field strength also depends on the type of applicators, needles, or catheters to be inserted. Haack et al2 investigated applicator reconstruction for MRIbased brachytherapy for cervical cancer. Plastic and titanium applicators and needles were reconstructed with milimeter accuracy as verified on phantom CT images, although the titanium applicators and needles gave rise to susceptibility and RF artifacts in the 1.5-T MR images. These artifacts increase to 5-10 mm at 3.0 T.3 Another potential hazard for interventional applications is RF-induced heating of titanium devices. This heating is due to resonating electromagnetic RF waves along the conducting structure.4,5 The amount of heating is dependent on many factors, such as field strength, shape, and location of the conducting structure and dielectric properties of surrounding tissue, but generally the risk of heating increases with field strength. The ideal field strength for RT planning remains unproven and both 1.5- and 3-T field strengths are being pursued, but for interventional applications 1.5 T may be preferred.
Geometric Distortions System Distortions
Figure 1 A newly installed dedicated 3.0-T MR simulator (Cancer therapy centre, Liverpool, Sydney, Australia) showing the wide bore magnet (Siemens Skyra), external goal-post lasers and flat-indexed RT tabletop, which incorporates an RF coil underneath it. The 30-G line can also be seen marked out in the floor. (Color version of figure is available online.)
In MRI, spatial encoding is achieved through the combination of the static (B0) magnetic field to establish a signal of a characteristic resonance frequency and gradient coils that superimpose linear changes in this field to produce a positional dependence of frequency. Geometric distortions arise from either the inhomogeneity of the static field or the nonlinearity of the gradients, both of which become worse at increasing radial distances from the magnet isocentre. Investigations into image distortions date back from the 1980s and 1990s.6-10 In a
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162 European multicentre investigation, Lerski et al11 observed machine-dependent geometric distortions up to 5 mm and on 1 scanner up to 13 mm. The phantoms used in this study were Perspex disc-shaped phantoms, containing doped water and a grid of rods. Bakker et al12 also used grid phantoms to study distortions and they showed that in spin-echo and gradientecho imaging, gradient errors and static field inhomogeneity lead to image distortions in the direction of the frequencyencoding gradient, whereas in the phase-encoding gradient image distortions are solely caused by gradient errors. Scannerrelated distortions were measured by variation of the strength and direction of the readout gradient in imaging experiments on a grid of cylindrical sample tubes. For the 1.5-T system used in that study, static field–related errors up to 7 mm and gradient-related errors up to 4 mm were observed in a midcoronal plane within a field of view (FOV) of 40 cm. Static magnetic field–related errors were inversely proportional to gradient strength, whereas gradient magnetic field–related errors were independent of gradient strength. These observations of image distortions may have retarded the introduction of MRI into RT treatment planning. However, for a greater part these distortions are machine specific, they can be measured, they are stable in time, and thus predictable.13 On a modern scanner, typical errors would be expected to be less than 1 and 2 mm within a radius of 10 and 15 cm, respectively.14 At greater distances from the isocenter, the distortions are larger, but these machine (gradient)-specific distortions can be corrected and nowadays most MR scanners have 2-dimensional (2D) (and some 3D) image correction available. Doran et al15 showed that distortions in a large FOV (320 200 340 mm3) on a 1.5-T Siemens Vision scanner could be corrected to within the voxel resolution.
Patient-Related Distortions Magnetic field errors stem not only from imperfections of the MR scanner but also from the patient owing to the chemical shift and susceptibility phenomena. The chemical shift effect is caused by the magnetic shielding of a nucleus by the electron clouds that surround it. The resonant frequency of a nucleus depends on the chemical structure of a molecule and the location of the nucleus within the molecule. In human MRI, the relevant chemical shift is between protons in fat and water. Protons in fat experience a magnetic field, which is 3.5 ppm weaker than the magnetic field experienced by protons in water. The consequence for MRI is that the fat-containing tissues will be shifted with respect to the other tissues in the direction of the frequency-encoding gradient. The human body magnetic susceptibility distribution is a major factor perturbing the magnetic field. The magnetic susceptibility is the degree of magnetization of a material in response to the applied magnetic field. In the human body, the susceptibility differences and resulting image distortions are largest at tissueair interfaces. Magnetic field distributions and image distortions for simple objects, such as cylinders and spheres, can be solved analytically,16-18 whereas the magnetic field in and around the patient has to be calculated with numerical methods. Bhagwandien et al19 developed and validated a
method to calculate the magnetic field distribution in and around a 3D object using a technique based on the finite difference method. Field perturbations between 5 and 6 ppm were observed outside the head and between 6 and 5 ppm inside the head near the sinuses. In the neck area, field deviations ranged from 3 to 6 ppm inside and from 5 to 4 ppm outside the body.20 These deviations result in image distortions between 5 and 5 mm for a typical frequencyencoding gradient strength (1.5 mT/m at 1.5 T). Patient-related image distortions are not predictable and have to be calculated on an individual basis. Correction methods based on calculations of the field perturbations are not easy to implement in clinical practice. Chang and Fitzpatrick21 introduced the gradient-reversal technique, which combines 2 images of the same object acquired with a forward and a reversed readout gradient, from which an undistorted image can be postprocessed. Reinsberg et al22 improved this method by using maximization of mutual information for registration of the forward and reversegradient images, eliminating the need for user interaction. Jezzard and Balaban23 introduced a correction method based on measurement of the field map. Echo-planar images collected at 1.5 and 4 T on phantoms and human subjects were unwarped, resulting in images with subpixel accuracy under the assumption that only field in homogeneities account for all distortions (no eddy currents, etc). Another approach could be to acquire undistorted images with pulse sequences that apply only phase encoding of the MR signal.24 However, phase encoding is time consuming and only used in 1 or 2 dimensions in clinical sequences. Although correction methods for patient-related image distortions are available, they are not yet widely incorporated in scanner software. In practice, one could adapt the following approach. Patient-related magnetic field perturbations stem from the chemical shift and the susceptibility distribution, both of which result in local distortions of the applied magnetic field. The corresponding image distortions are proportional to the local magnetic field error and inversely proportional to the strength of the applied frequency-encoding gradient. The chemical shift between water and fat protons is approximately 3.5 ppm (220 Hz at 1.5 T and 440 Hz at 3 T), and susceptibilityinduced perturbations are at the most about twice as large. Patient-related image distortions can be reduced by applying relatively strong slice selection and frequency-encoding gradients, for example, by setting the pixel bandwidth at twice the water-fat shift, all patient-related distortions are expected to be smaller than the pixel size. In regions where anatomy is expected to be problematic, localized shimming over the intended imaging volume should be performed before the scan (noting the possibility of worsening distortions elsewhere).
Image Acquisition Patient Setup A major constraint on MR acquisition is the patient positioning or immobilization or both used to represent the treatment
MRI acquisition techniques for RT planning setup and facilitate registration with the planning CT scan. The images must be representative of the anatomy during each fraction. If deformation exists, approximating the setup as much as possible partially mitigates this issue. Although imaging in this fixed position is unnecessary in brain, it has been shown to be necessary in most sites (eg, head and neck to control neck flexion25) and may even dictate a change of imaging position for others (eg, breast and rectum). The first issue is the curvature of the patient table present on most diagnostic scanners. This can be removed using commercial tabletops or home-made flat boards, which index with RT base plates. For head and neck, a normal scanner bore diameter may prevent the use of conventional immobilization devices warranting a cut to the fixation shells or the use of alternative means to hold the patient in position. Studies in the prostate have also demonstrated the advantages of maintaining the hips in a flat position, which is crucial in not only maintaining bony anatomy but has also been shown to effect target delineation.26 MR-compatible ankle stocks and knee rests should be utilized wherever possible. These positional restrictions are less relevant if registration and guidance is based on implanted fiducial markers. For the breast, the treatment setup in a supine position and a wing board will be difficult to achieve without the use of a wide bore scanner. It is important to verify the MR compatibility of all immobilization equipment that includes both safety and image quality considerations. Carbon fiber materials should be avoided or tested rigorously for the possibility of RF heating.
RF Coils RF coils are designed to transmit and detect the MR signal either separately or in a combined transceiver fashion. Important characteristics of RF coils are their sensitivity profile (uniformity) and SNR. The design of RF coils is such that the sensitive volume is restricted to the anatomy of interest to reduce the detectable noise from the patient. In the case of the brain, standard multichannel head coils can be used to provide optimum image quality. However, in other sites, any attempt to cater for immobilization and replicate treatment positioning usually results in some compromise in image quality or coverage or both. Nevertheless, groups have managed to use a variety of diagnostic surface coils placed around fixation shells in the brain27 and head and neck28-30 and suspended above the abdomen in prostate31 to obtain images of sufficient quality for planning purposes. In the breast, the diagnostic preference is to image the prone patient, with the breasts suspended in the wells of a dedicated coil array. Supine scans may be performed but the choice of the treatment over scan position is currently an area of active research.32 When no other option is available, the integrated body coil (hidden behind the plastic magnet covers) may be used to accommodate treatment positions more easily at the expense of a significant reduction of SNR.30 Figure 2 illustrates some of these positional challenges and trade-offs. On the left, 2 head and neck scans have been overlaid, the first acquired using a RT neck rest to control positioning and the second in an arbitrary position in the diagnostic coil. On the right, 2 prostate
163 scans are compared; the first acquired in the treatment position using a flat tabletop using the integrated body coil and the second using the diagnostically preferred body array. The diagnostic scan demonstrates increased SNR but also slight internal anatomy shifts compared with the treatment position scan. Vendors are increasingly providing dedicated coil solutions for RT imaging. In head and neck, this comprises laterally placed coils that can be combined with an additional coil on top of the chest or underneath the neck or both to extend coverage in the craniocaudal direction. Bespoke coil positioning aids and adjustable supports are being introduced onto the market that hold coils in place, and in the case of the abdomen maintain close contact with the patient without compressing the skin surface, which will be important for MR-only planning (Fig. 3). Images obtained with the surface coil arrangements described previously will typically exhibit poor signal uniformity and usually some method of intensity correction is required, and vendors offer this either prospectively or retrospectively. Wherever possible, the RF coil should be multichannel (ie, individual signal elements) so that the advantages of parallel imaging can be exploited.
Scan Protocol The aim of the RT planning scan is different from that of a diagnostic protocol where usually a slice thickness of 3-5 mm is adequate and around 30-40 slices are acquired. For the RT scan, a more detailed visualization of the tumor and organs at risk is required and coverage needs to be sufficient for registration with CT. To maintain accuracy, it is desirable to reduce slice thickness to match that of corresponding CT scans and should be at most 2-3 mm depending on SNR. Coverage of between 18 and 20 cm may be required in the superiorinferior direction meaning the acquisition of 60-100 images. The consequence of this is long scan times that may result in patient or organ movement and a degradation of image quality. Motion within an acquisition is characterized by ghosting, which appears in the phase-encoding direction and can vary from subtle smearing of signal to gross artifacts. In cases where the 2D acquisition is divided over multiple concatenations to cover the required anatomy, contiguous slices can be several minutes apart leading to a bizarre steplike appearance on the reconstructions when significant movement has occurred (Fig. 4). Beyond the first step of motion management during the acquisition, a simple way to remedy this is to extend the repetition time to encompass the required number of slices in an acquisition albeit at an increased time penalty. An alternative is to use 3D sequences that excite an entire volume of tissue and use phase encoding in the “slice” direction. These acquisitions can provide isotropic resolution but can also extend scan time significantly. However, there are newer faster versions (see section Pulse sequences) that show considerable promise for RT planning, although their contrast properties have been noted to be subtly different.33 Images will usually be acquired in the axial plane, a requirement often determined by the treatment planning system (TPS), although some newer platforms are able to cope
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Figure 2 The differences between diagnostic and radiotherapy planning scans (left). Overlay of H&N images acquired in a standard diagnostic RF coil and also in a bespoke surface coil in the treatment position that demonstrates the large difference in neck flexion. Also shown are prostate images acquired using the integrated body coil and flat tabletop (center) compared with a corresponding image acquired with the preferred diagnostic posterior and anterior coil array (right). The SNR is better with the diagnostic coil, but the position of the internal anatomy is slightly different. H&N, head and neck. (Color version of figure is available online.)
with nonaxial data sets. The FOV should also be selected with consideration of the effects of geometric distortions at large distances from the isocenter, while providing enough anatomy for registration and contouring of soft tissues. On a typical closed-bore scanner, this would usually be between 20 and 25 cm, which covers the brain and head and neck. In prostate, this will be large enough to encompass femoral heads and adjacent bony landmarks (Fig. 5A), although if gold seeds are present then the FOV can be reduced or automated registration should be targeted to this smaller volume or both. Beyond this FOV, distortion will be present (eg, at the skin surface in supine breast imaging) reducing the reliability of the CT-MR-fused image in these regions. In-plane resolution should be 1 mm or better with receiver bandwidths adjusted accordingly to limit patient distortions to within a single pixel.
Pulse Sequences Unlike CT, where the limitations in scan protocol at least provide some consistency, MRI offers the user with a wide
variety of contrast weightings and sequences, each described by a myriad of vendor-specific abbreviations and acronyms. Fundamentally, these can be broadly classified into spin-echo- and gradient-echo-based sequences with the former providing the gold standard in terms of intensity distortion, and the latter being inherently faster but more liable to signal dephasing. A routine RT planning examination should be expected to provide a robust geometrically accurate data set, which will usually consist of a fast spin-echo variant with or without an additional contrast-enhanced sequence.14 A more sophisticated planning protocol may include the addition of at least one of the functional techniques discussed later. Gradient-echo imaging should be avoided in patients with metallic applicators or prosthesis. Metal artifact reduction techniques, essentially using high receiver bandwidths and short echo times, should be employed. In other instances, for example, gold fiducials, these artifacts play an essential role in their visualization but should be controlled to maintain a faithful representation of their true location especially at
Figure 3 Two commercially available RF coil solutions for radiotherapy. (Left) A head and neck suite of coils, comprising 2 lateral surface coils (6 channel) plus a 16-channel anterior coil attachment with bespoke positioning aids. When combined with a dedicated radiotherapy posterior coil, it provides 30 channels and full FOV H&N images in the treatment position (GE Healthcare, Milwaukee, WI). (Right) An 18-channel anterior body array (Siemens Healthcare, Erlangen, Germany) with coil bridges (Civco, Kalona, IA) to stand the coil off the patient. A 32-channel coil is also underneath the tabletop. H&N, head and neck. (Color version of figure is available online.)
MRI acquisition techniques for RT planning
Figure 4 This sagittally reconstructed prostate scan clearly shows motion has occurred between the 2 concatenated axial acquisitions, demonstrating a characteristic discontinuity between alternate slices as indicated by the contoured structures. (Color version of figure is available online.)
higher field strength.34 Figure 5B demonstrates the appearance of gold seeds at 3.0 T for both a gradient-echo and fast spin-echo sequence. The gradient-echo sequence shows the seed well but the artifact is larger than the true seed diameter. In this case, a short (o10 milliseconds) echo time fast spin-echo sequence was used to improve the visualization of the seed against the tissue and demonstrate a more accurate marker position. Over the preceding years, some interesting sequences and techniques have been introduced that offer a number advantages for RT planning (Table). These include speeded up implementations of 3D fast spin-echo sequences, eliminating the need for multiple acquisitions and providing data sets with isotropic resolution in reasonable scan times.35 The resolution of these scans means they can be acquired in any favorable plane to further reduce scan time or partial voluming, before subsequent reformatting into axial data for import into the TPS. Radial k-space is another recent advancement whereby
165 the trajectory of data collection is altered in a semirandom manner. This provides an effective remedy to ghosting artifacts that may occur from patient movement or internal motion (eg, bladder filling). Perhaps one of the most important developments has been parallel imaging, which reduces scan times by missing out some gradient-encoding steps and using instead the spatial information from RF coil sensitivity profiles. This is an invaluable approach and avoids the alternative of increasing slice thickness or slice gap or both to shorten scan times. Different implementations work from either the image or the raw (k-space) data and many are now self-calibrating. Theoretically, the speed-up factor is equal to the number of available coil elements, although factors of 2-3 are more typical. In- and out-of-phase imaging permits 4 separate images based on fat and water contrast to be acquired simultaneously; the wateronly image is particularly effective in providing a fat-suppressed sequence in areas of field inhomogeneity and as such is a useful tool for the RF surface coils used in RT planning. Although vendors largely offer similar products, it has been reported that variations in sequences across different systems have been shown to produce variability in tumor delineation.36 Functional imaging techniques are a particularly attractive proposition for RT planning, and a more complete review of their role is provided elsewhere.37 Traditionally, these techniques have introduced the added difficulty of some or all of the following problems; inherent sequence-specific distortion, image manipulation, and data interpretation. Diffusionweighted imaging utilizes strong bipolar gradients to sensitize the sequence to microscopic molecular diffusion and is used to measure the apparent diffusion coefficient (ADC), which is related to cellularity. These data are routinely acquired using echo-planar imaging–based sequences, which are inherently prone to artifacts and geometric distortions. New sequences either with or without radial k-space sampling have been shown to improve image quality over standard sequences and seem promising in an RT setting.38,39 MR spectroscopy has the potential to add metabolic information into the RT plan,40 but a substantial degree of expertise is required to acquire and correctly interpret the data, and the incorporation of this
Figure 5 (A) The limited coverage and FOV used in this prostate scan (color overlay) compared with the planning CT data. (B) Two different sequences showing gold seeds within the prostate at 3.0 T; on the left, a gradient-echo sequence shows the marker clearly owing to large susceptibility artifact. The fast spin-echo on the right shows it less well but with better precision.
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Table Acronyms for Pulse Sequences and Techniques From 3 MRI Vendors That are Particularly Relevant to RT Planning Generic Name
GE
Philips
Siemens
Details
3D FSE with variable flip angle Uniformity correction In- and out-ofphase imaging Parallel imaging
CUBE (previously known as XETA)
VISTA
SPACE
High-resolution volumetric imaging (particularly T2-w) in reduced scan times.
SCIC or PURE
CLEAR
Reduces intensity variation from surface coil acquisitions.
IDEAL
mDIXON
Prescan normalize DIXON
Radial k-space
PROPELLER
ASSET or ARC
The water-only scan provides excellent fat suppression image in presence of B0 and B1 inhomogeneity. SENSE mSENSE or Scan time reduction based on image or k-space data. Some selfGRAPPA calibrating others require additional scan. MultiVane BLADE Motion artifact reduction.
Abbreviations: ARC, autocalibrating reconstruction; ASSET, array sensitive encoding technique; CLEAR, constant level appearance; FSE, fast spinecho; GRAPPA, generalized autocalibrating partially parallel acquisition; IDEAL, iterative decomposition of water and fat with echo asymmetry and least squares estimation; PROPELLER, periodically rotated overlapping parallel lines with enhanced reconstruction; PURE, phased array uniformity enhancement; SCIC, surface coil intensity correction; SENSE, sensitivity encoding; SPACE, sampling perfection with applicationoptimized contrasts using different flip angle evolution; VISTA, volume isotropic t2 weighted acquisition.
information into the planning system is still in its infancy. Measurement of relaxation times, for example, R2n that has been studied in relation to hypoxia,41 has usually required some additional off-line postprocessing to produce the parameter maps. However, MRI vendors are increasingly providing software, which caters to the analysis and fusion of all these secondary data sets and must continue to do so in the context of RT, with the ability to create contours drawn on these and export them into planning systems.
Quality Assurance Before the commencement of any RT planning service, a thorough commissioning of a newly installed system or the quality assurance of a pre-existing system should be undertaken. These tests should cover all the normal quality assurance metrics that are described in detailed guidelines elsewhere.42 In addition, a series of tests specifically designed for RT planning are recommended. These will principally concern the assessments of geometric distortion and the image quality of the patient setup.28
Several commercial phantoms are available to test geometric distortion.42-44 These are usually limited in size in at least 1 dimension, and it may be necessary to build an inhouse test object to fully characterize a system over a large volume. It is worth making measurements both with and without the aforementioned gradient corrections turned on. Figure 6 shows example images acquired on a 1.5-T Philips Achieva MR scanner (Philips Medical Systems, Best, The Netherlands) with distortions before and after correction. Within a FOV of 336 mm, the software reduced distortions of 6.0 to 0.2 mm. Surface coils should be investigated for their SNR and uniformity characteristics using spherical- or cylindricalshaped flood-fill phantoms of appropriate size. Measurements should be performed both with and without any coilcorrection filters or algorithms applied and benchmarked against the diagnostic coil of choice.27,28 Finally, an end-to-end test may be included with the aim of testing the optimum MRI protocol across each step of the RT treatment chain through to registration with CT on the TPS. Any suitable CT- and MR-compatible test object
Figure 6 Images of a geometric distortion phantom acquired on a 1.5-T Philips Achieva scanner before and after gradient correction was applied. The nominal distance between the 2 columns of tubes indicated by the lines is 210 mm. The arrows depict the distances between the outer tubes on the bottom row.
MRI acquisition techniques for RT planning will suffice but more complicated anthropomorphic phantoms are worthy of some consideration as this will enable an examination of patientlike effects, such as susceptibility and water-fat shifts.45
Conclusions Over the next few years, it is expected that there will be an increase in the number of dedicated MR simulators installed within RT departments. However, most work will continue to be performed on pre-existing systems. By adopting a pragmatic approach, most MRI scanners can be used to provide a planning service in some capacity. With this in mind the following recommendations are suggested: (1) The field strength should be no less than 1.5 T and no greater than 3.0 T. (2) The magnet should be a closed-tunnel configuration, preferably with a wide bore. (3) A flat bed or tabletop insert should be used. (4) Appropriate RF coils (multichannel) should be used that cater to immobilization devices and surface coil intensity correction should always be employed. (5) Geometric distortion should be verified in all 3 directions, and this will inform the useable FOV (in plane) and slice coverage (through plane). (6) Spin-echo-based sequences (ie, fast or turbo spinecho) should be used in preference to gradient-echo for the gold-standard anatomical image. (7) Slice thickness should match the planning CT acquisition where possible. (8) Multiple (ie, concatenated) acquisitions should be examined for gross movement between adjacent slices and avoided if motion cannot be mitigated. (9) The scanner's software for correction of the gradient errors should always be turned on. (10) Pixel bandwidth should be set to a minimum of twice the water-fat shift (ie, 220 Hz at 1.5 T and 440 Hz at 3 T). (11) QA of the optimum patient setup and imaging protocol should be carried out before clinical scans. These recommendations will be helpful in standardizing the acquisition across multiple patients and different centers to ensure a more robust acceptance of MRI in RT.
Acknowledgments The authors thank the staff at the MRI department at Castle Hill hospital, Hull, UK, and especially Lynnette Cassappi and Ewa Juresik (Liverpool CTC, Sydney, Australia) for their assistance in acquiring images used in this article.
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