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Note: This copy is for your personal non-commercial use only. To order presentation-ready copies for distribution to your colleagues or clients, contact us at www.rsna.org/rsnarights. EDUCATION EXHIBIT 185 Optimizing Abdominal MR Imaging: Approaches to Common Problems1 Online-Only CME See www.rsna .org/education /rg_cme.html LEARNING OBJECTIVES After reading this article and taking the test, the reader will be able to: ■■Describe how MR imaging parameters affect each other and image quality. ■■List the causes of some common artifacts in abdominal MR imaging. ■■Discuss strategies for handling some common challenges of abdominal MR imaging. Roberta K.Yang, MD • Christopher G. Roth, MD • Robert J.Ward, MD Joseph O. deJesus, MD • Donald G. Mitchell, MD Abdominal magnetic resonance (MR) imaging involves many challenges and is complicated by physiologic motion not encountered to the same degree in other regions of the body. Problems that uniquely affect abdominal MR imaging include motion artifact (from respiratory, cardiac, gastrointestinal, and voluntary movement), susceptibility artifact, conductive and dielectric effects, and wraparound artifact. Techniques to minimize these artifacts often need to be addressed within the time constraints of a single breath hold. Patient motion during image acquisition is minimized by using physical restraint, respiratory gating, and reduction of acquisition time. Correction of motion-induced dephasing (through gradient moment nulling), signal averaging, and suppression of signal in moving structures all address unavoidable motion (eg, cardiac pulsation). Acquisition time is minimized by obtaining fewer phase-encoding steps, decreasing repetition time, and increasing efficiency with use of parallel imaging and multiecho acquisitions. Adjusting the echo time does not directly affect scanning time, but it does allow more time for section sampling per repetition time interval in multisection acquisitions by means of closer echo spacing and it plays a pivotal role in optimizing image quality. Familiarity with basic MR imaging principles and the ability to minimize the effects of motion and other artifacts are essential to optimizing abdominal MR imaging protocols and improving efficiency. © RSNA, 2010 • radiographics.rsna.org Abbreviations: ETL = echo train length, FOV = field of view, FSE = fast spin-echo, GMN = gradient moment nulling, RF = radiofrequency, SNR = signal-to-noise ratio, STIR = short inversion time inversion-recovery, 3D = three-dimensional RadioGraphics 2010; 30:185–199 • Published online 10.1148/rg.301095076 • Content Codes: From the Department of Radiology, Thomas Jefferson University, 1094 Main Bldg, 132 S 10th St, Philadelphia, PA 19107 (R.K.Y., C.G.R., D.G.M.); Department of Radiology, Tufts Medical Center, Boston, Mass (R.J.W.); and Department of Radiology, Chestnut Hill Hospital, Philadelphia, Pa (J.O.D.). Presented as an education exhibit at the 2008 RSNA Annual Meeting. Received April 2, 2009; revision requested May 29 and received August 31; accepted October 2. D.G.M. is a consultant with Johnson & Johnson; all other authors have no financial relationships to disclose. Address correspondence to R.K.Y. (e-mail: [email protected]). 1 The Editor has no relevant financial relationships to disclose. © RSNA, 2010 186 January-February 2010 Introduction Magnetic resonance (MR) imaging of the torso poses many challenges not commonly encountered in the imaging of other body regions. In addition to blood flow and voluntary motion, abdominal MR imaging must address (at minimum) respiratory and cardiac motion. These phenomena impose significant time constraints and must be considered when selecting pulse sequences or addressing other problems, such as susceptibility or aliasing. The relatively larger size of the abdomen generates more problems with inhomogeneities in B1 (the magnetic field generated by the radiofrequency [RF] pulse) than would be encountered in the imaging of smaller areas. Addressing these problems involves manipulating multiple parameters with complex interrelationships. Optimization of one parameter often detrimentally affects another parameter, and tradeoffs are inevitable; there is no “free lunch” in MR imaging. For example, minimizing acquisition time may diminish the signal-to-noise ratio (SNR), although this may be an acceptable compromise. Understanding these complex interrelationships between parameters is essential for optimizing abdominal MR imaging protocols. In this article, we summarize some basic MR imaging principles, address some of the technical problems encountered at abdominal MR imaging, and explore potential solutions. Motion Artifact Teaching Point Motion is a central issue in abdominal imaging by virtue of the simple fact that patients breathe, hearts pump, and bowels demonstrate peristalsis (1). Motion artifacts occur when the phase of a moving proton does not reflect its physical position along the phase-encoding axis (2). In the absence of motion, phase differences should be caused solely by the varying strength of the phase-encoding gradient. Motion artifact manifests in several ways. Ghost artifact (Fig 1a–1c) occurs when signal from a moving structure is erroneously rendered along the phase-encoding axis (3,4). Randomly moving structures (eg, in bowel peristalsis) are blurred and associated with noise propagated along the phase-encoding axis (Fig 1a) (3,4). Structures containing constant velocity flow radiographics.rsna.org can also cause blurred ghosting across this axis. Repetitive motion (eg, respiration and pulsatile arterial flow) manifests as distinct “ghosts” with regular periodicity across the phase-encoding axis. Moving fluid within defined spaces (eg, vessels) can also cause intravoxel phase dispersion (4). View-to-view and within-view phase errors are two basic causes of misregistration due to motion-induced phase differences (1). View-to-View Phase Errors View-to-view phase errors occur when the signal amplitude at a specific location is not consistent between phase-encoding steps (1). This phenomenon is associated with sporadic motion (eg, voluntary motion or arterial pulsation), which causes signal amplitude variation and ghosting artifact (Fig 1b, 1c) (3), unlike motion of constant velocity (eg, venous flow). For example, signal amplitude variation occurs when a particular voxel contains lung at one phase-encoding step and liver at the next phase-encoding step due to respiration. The distance between the discrete ghosts associated with periodic motion is proportional to repetition time, the phaseencoding matrix, motion frequency, and number of signals averaged. The intensity of the ghosts is proportional to motion amplitude, the signal intensity of the moving structure, and magnetic field strength (2–5). Within-View Phase Errors Within-view phase errors occur when phase differences are accumulated by a proton moving during an applied gradient between the time of excitation and signal acquisition (1). The “gradient moment” alters the phase of the moving proton and is usually generated by the frequencyencoding or section-select gradients (but rarely to a significant degree by the phase-encoding gradient) and alters the proton’s phase (2–4). This motion-induced phase alteration does not reflect the proton’s position along the phase-encoding axis and is erroneously rendered along this axis as ghosting (2,6). Within-view phase errors can also cause heterogeneous or decreased intravascular signal intensity by means of intravoxel phase dispersion (Fig 1d) (1,4,7). The term order of motion refers to the kinetics of a moving structure. There are several orders of motion, including first-order motion (constant RG • Volume 30 Number 1 Yang et al 187 Figure 1. Motion artifact. (a) Axial short inversion time inversion-recovery (STIR) image (repetition time msec/echo time msec = 4445/60, 90° flip angle) shows bowel motion manifesting as random signal and noise distributed across the phase-encoding axis (bracket). (b) Axial T2-weighted fast spin-echo (FSE) image (5067/73, 90° flip angle) shows breathing artifact, which manifests as repetitive, discrete ghosting of structures in the phase-encoding axis (arrows). (c) Axial T2-weighted fat-saturated fast-recovery FSE image (2300/87, 90° flip angle) shows pulsation artifact, which appears as repetitive ghosting of the heart (arrowheads) and aorta (arrows). (d) Axial T2-weighted fat-saturated fast-recovery FSE image (2300/87, 90° flip angle) shows intravoxel phase dispersion manifesting as dark lines near the edges of vessels (arrow). Protons of markedly different speeds that inhabit the same voxel (eg, in laminar flow) result in signal loss. velocity), second-order motion (constant acceleration), and third-order motion (“jerk,” or changing acceleration) (1). Gradient moment nulling (GMN) attempts to correct for phaseencoding errors by adding balancing pairs of dephasing and rephasing gradient lobes along the axis of the gradient moment so that the moving proton’s phase gain is zero when the echo is obtained, as with stationary protons in the same section (1,3,6,8). These additional gradient lobes lengthen echo time, especially as higher orders of motion are corrected (3). Because first-order motion is dominant during the brief interval that the gradient is applied (8), first-order GMN is a practical technique for addressing within-view phase errors without significantly increasing echo time or interecho spacing. GMN is most often used to reduce motion artifact on T2-weighted 188 January-February 2010 radiographics.rsna.org images, on which echo time is long enough to accommodate additional gradient lobes, and to maintain uniformly high signal intensity on timeof-flight images. Common synonyms for GMN include gradient moment rephasing, flow compensation, flow-adjustable gradients, and motion artifact suppression technique (1,6,9). to variable blood transit time. Peripheral gating can be used to reduce the cardiac phase variability on a given image, but precise binning of data to a specific portion of the cardiac cycle (eg, diastole) is not feasible (1). Motion Correction Strategies Preparatory spatial or fat-saturation pulses can suppress signal from moving tissues (4,5). For example, to avoid motion artifact from the anterior chest wall, fat saturation can be used to reduce the signal intensity of subcutaneous fat, often the most signal-rich chest wall component. STIR is another technique for suppressing signal from moving subcutaneous fat (3). Direct Correction Direct motion correction strategies include restraints, breath holding with effective instructions (translated, if necessary), comfort-promoting measures (eg, cushions and supplemental oxygen), sedation for claustrophobia, and antiperistaltic agents (eg, hyoscyamine or glucagon) to minimize bowel peristalsis (2–5,10,11). Direct motion correction techniques can be used to address both within-view and view-to-view phase errors. Respiratory and Cardiac Gating Motion gating generally corrects for view-to-view phase errors by increasing the consistency of the amplitude of each echo. Respiratory bellows or motion sensors allow monitoring of the respiratory phase (1,4). Sequences can be triggered at end-expiration, resulting in less variation in diaphragmatic displacement (12). Reordered phase encoding minimizes respiratory ghosting by rearranging the phase-encoding data from several erratic breaths to simulate one slow respiration. The phase-encoding steps are acquired in such an order that their position in kspace correlates with the degree of diaphragmatic displacement, thereby removing the periodicity of respiratory motion relative to phase encoding (10). Although this may not reduce motion-associated blur and can be difficult to implement clinically, no time is wasted because acquisition occurs throughout all phases of respiration (3,4,13). Two forms of monitoring are used for cardiac gating: electrocardiography and impedance plethysmography (peripheral gating) (1). Electrocardiographic gating can be used to analyze data acquired during a specific phase of the cardiac cycle (4). Motion artifacts, including higherorder within-view phase errors, are minimized by imaging during diastole, when blood flow is slower and more constant. Peripheral gating senses the cardiac cycle indirectly through pulsations in peripheral arterioles, the timing of which may differ unpredictably from cardiac phases due Suppression of Signal from Moving Tissue Signal Averaging Oversampling a section by increasing the number of excitations and averaging the resulting redundant signals decreases the relative contribution from motion at the expense of increased acquisition time (3,4,14). This approach works because (a) more time is spent in the quiescent phases of repetitive motion (eg, diastole and exhalation), and (b) the signal from an object varies less than the signal from motion artifact. The intensity of ghosts may be decreased owing to destructive interference between data sets (3). Increasing the number of signals averaged, a “brute force” method of motion artifact suppression, has the added benefit of increasing SNR by the factor of its square root (3,9,14); for instance, increasing the number of signals averaged from one to two would increase SNR by the square root of 2. Swapping the Phaseand Frequency-encoding Axes If ghosting along the phase-encoding axis obscures a key structure, the phase- and frequencyencoding axes can be switched so that the artifact is perpendicularly redirected away from the key structure (4). This may change imaging time if the field of view (FOV) is rectangular. Decreasing Acquisition Time The use of rapid imaging to minimize motion has the advantage of increasing patient throughput. Fast imaging minimizes the opportunity for motion and facilitates breath holding. With ultrafast sequences, an image can be acquired in less than 100 msec to “freeze” motion (4,15). Scanning time is directly proportional to the number of phase-encoding steps, the number of excitations (number of signals averaged or RG • Volume 30 Number 1 Yang et al 189 Figure 2. Schematics illustrate partial Fourier acquisition (left) and fractional echo sampling (right), both of which make use of k-space symmetry to undersample. Horizontal axis = phase, vertical axis = frequency. (Adapted, with permission, from reference 1.) sections), and repetition time and is inversely proportional to echo train length (ETL) and the parallel imaging factor (1): Teaching Point Scanning time a (PE steps × excitations × TR)/(ETL × PIF) , (1) where PE = phase-encoding, PIF = parallel imaging factor, and TR = repetition time. Shortening Repetition Time Shortening repetition time can directly decrease acquisition time, but signal intensity may also decrease. T1 contrast also changes. Optimal signal is attained when the flip angle equals the Ernst angle, which is determined by repetition time and the T1 of the tissue in question. Hence, a decrease in the flip angle allows a decrease in repetition time while maintaining T1 contrast and limiting the reduction in SNR (1). Time saved by shortening repetition time should be balanced by the consideration that additional sections can be sampled during a longer repetition time interval in multisection acquisitions (9). After the first echo is sampled in a multisection acquisition, several other sections are excited and sampled during the repetition time of the first section. Excessive shortening of repetition time decreases the time available for additional sampling and may increase the number of acquisitions required to cover the same number of sections (1). The number of sections that can be imaged within one repetition time interval is approximated as follows: Number of sections ≈ TR /([½] RF PD + TE + [½] ED), Obtaining Fewer Phase-encoding Steps Rectangular FOV.—An easy way to decrease the number of phase-encoding steps and yet maintain spatial resolution is to decrease the density of phase-encoding lines in k-space (9). This can be achieved by sampling every other phase-encoding line in k-space (1), thereby reducing the FOV in the phase-encoding axis and resulting in a rectangular FOV. Because patients are usually smallest in the anteroposterior axis and phase encoding contributes to acquisition time, phase encoding is often assigned to this axis. The FOV in the phase-encoding axis should accommodate the dimensions of the patient to avoid wraparound artifact. With three-dimensional (3D) Fourier transform acquisition, similar undersampling can be performed along the phase-encoded section-select axis. However, this general approach decreases SNR and can also cause wraparound artifact (aliasing). Partial Fourier Acquisition.—Partial Fourier ac- quisition uses the symmetry of k-space to decrease the number of phase-encoding lines required to construct an image (Fig 2) (1,15). For example, one can construct an image by sampling the lower half of k-space and mathematically deriving the upper half using the data collected from the lower half. In practice, usually slightly more than onehalf of the phase-encoding steps are acquired to minimize artifact. This method preserves spatial resolution but decreases SNR (9). Zerofill Interpolation.—Zeros are classically used (2) where ED = echo duration, PD = pulse duration, TE = echo time, and TR = repetition time (16). to fill peripheral lines of k-space and decrease pixel size (17). Supplementing sampled echoes 190 January-February 2010 with zeros fills more points in k-space and interpolates the original data matrix to a higher matrix. This can decrease the number of sampled phase-encoding lines. Because no real information is added through this interpolation, there is only an apparent increase in spatial resolution, and SNR is lower (1,15). Zerofilling can be used instead of mathematical derivations to fill in the missing data points in partial Fourier acquisition or fractional echo sampling, which speeds up image processing. Spatial resolution is better preserved when data are asymmetrically collected to include at least one edge of k-space (1). Parallel Imaging.—Parallel imaging requires a phased-array coil. Fewer phase-encoding steps are required per coil than would otherwise be needed to construct a full image. The missing data are created using the spatial sensitivities of each uncoupled coil and its geometry within the array to either fill in lines of k-space or “unwrap” what would otherwise be an image with severe aliasing (18). In essence, this is the simultaneous and independent sampling of echoes from each coil during each repetition time interval; hence, additional coils result in faster imaging (9). Certain algorithms incorporate the coil-related spatial information before Fourier transformation (19), whereas other algorithms apply coil sensitivity information after Fourier transformation (18,20). SNR is decreased in proportion to the square root of the parallel imaging factor (a measure of the degree of undersampling), but spatial resolution is maintained (9,20). An artifact associated with parallel imaging is caused by inaccurate coil sensitivity maps, producing a band of noise through the center of the image (Fig 3). radiographics.rsna.org Figure 3. Artifact associated with parallel imaging technique. Axial T2-weighted single-shot FSE image (568/220, 90° flip angle) shows a band of noise across the center of the phase-encoding axis (bracket) caused by inaccurate coil sensitivity maps during reconstruction with parallel imaging. FSE) and may reduce the number of excitations (“shots”). A short echo time can also improve SNR and minimize dephasing effects (5). Several ways to reduce echo time or reduce the time needed for echo sampling are discussed in the following sections. Decreasing the Number of Pixels in the Frequency-encoding Axis.—Reducing the number of samples obtained from each echo results in a smaller frequency-encoding matrix. If the sampling rate (receiver bandwidth) is kept constant, sampling time is reduced, allowing a shorter echo time (Fig 4a). If the frequency-encoding gradient strength is also kept constant, spatial resolution is decreased and pixel size is increased in the frequency-encoding axis, resulting in increased SNR (1). Shortening Echo Time Increasing the Receiver Bandwidth.—Increasing Echo time affects scanning time more indirectly than does repetition time. Shortening echo time allows the sampling of more sections per repetition time interval in a multisection acquisition, potentially decreasing the number of acquisitions needed to image a volume of interest (1,5). More rapid sampling of each echo allows closer interecho spacing in multiecho techniques (eg, the receiver bandwidth allows a reduced echo time, given a constant frequency-encoding matrix (1,15,21). SNR, inversely proportional to the square root of the receiver bandwidth, is decreased. The frequency-encoding matrix can be doubled by doubling either the receiver bandwidth or the sampling duration; however, if receiver bandwidth is increased without changing the frequency-encoding gradient strength or matrix, the FOV is increased while spatial RG • Volume 30 Number 1 Yang et al 191 Figure 4. Techniques for shortening echo time. Gf = frequency-encoding gradient. (a) Drawings illustrate how decreasing the echo sampling time and the number of sampled echoes results in a lower frequency-encoding matrix, thereby shortening echo time (TE). ADC = analog-to-digital conversion. (b) Drawings illustrate how the use of fractional echo sampling can also shorten echo time (TE). (Fig 4 adapted, with permission, from reference 1.) resolution is decreased (1,9). Chemical shift misregistration and susceptibility artifacts are also decreased with increasing receiver bandwidth (2). With a higher bandwidth, each pixel represents a greater frequency difference, which decreases the number of pixels affected by chemical shift artifact, a reflection of the fixed frequency difference between water and fat protons induced by the magnetic field. Faster echo sampling with a higher bandwidth decreases the time over which susceptibility-related dephasing occurs, minimizing susceptibility artifact. Fractional Echo Sampling.—Similar to partial Teaching Point Fourier acquisition, which relies on k-space symmetry in the phase-encoding axis, fractional echo sampling uses k-space symmetry in the frequency-encoding axis (Fig 2) (1,17,22). The right half of k-space can be acquired by sampling the latter half of the echo. The left half of k-space is then mathematically generated from the right half. This allows the echo to be sampled closer to the excitation pulse without interference from the initial free-induction decay curve, resulting in a shorter echo time (Fig 4b) (1,9). In practice, slightly more than one-half of the echo is acquired to reduce artifacts. Spatial resolution is preserved, but SNR is reduced. Decreasing the Number of Excitations Decreasing the Number of Sections.—Because each section in a two-dimensional Fourier transform acquisition is excited separately, a simple way to decrease the number of excitations is to decrease coverage in the section-select axis. Increasing the section thickness or the gap between sections also decreases the number of sections. Although increasing the section thickness improves SNR, through-plane resolution is reduced (1,6). In 3D Fourier transform imaging, the sectionselect axis is partitioned by phase-encoding steps. Decreasing the FOV in the section-select axis or increasing section thickness can decrease imaging time by requiring fewer phase-encoding steps. Increasing the section thickness increases the likelihood of Gibbs ringing (truncation) artifact in the through-plane direction (1). Multiecho Techniques.—With multiecho tech- niques, multiple phase-encoding steps (echoes) are acquired following each excitation pulse (1,15,23). This method shortens acquisition 192 January-February 2010 times by decreasing the number of excitation pulses (shots), or repetition time intervals, required to form an image and can be used with multisection or single-shot acquisitions. There are two general subcategories. Echoplanar imaging collects multiple gradient echoes with each shot by rapidly oscillating the frequency-encoding gradient (15). FSE imaging collects multiple spin echoes during each ETL; each echo is refocused by a 180° RF pulse (15,23).The efficiency of a multiecho technique compared with an equivalent single-echo acquisition is proportional to the number of echoes in its ETL. For example, if nine echoes are sampled in an ETL, the multiecho sequence is nine times more efficient than its single-echo counterpart (15). Any factor that shortens the sampling time for each echo also improves efficiency by decreasing the interecho spacing and increasing the number of echoes obtained within a given repetition time interval (1). The effective echo time, defined by when the lowest phase-encoding step (corresponding to central k-space) of the ETL is acquired, determines image contrast (Fig 5) (23). Increasing the ETL beyond the effective echo time causes more signal loss and edge blurring; hence, ETL is ideally kept within the effective echo time (1). Multiecho techniques can be used to achieve T2 weighting in sequences with inherently short echo times (eg, gradient-recalled-echo sequences) by acquiring multiple echoes within a long effective echo time (5). Single-Section Acquisition With single-section technique, all phase-encoding steps for one section are acquired consecutively before the next section is sampled (Fig 6a) (1). In contrast, a multisection acquisition samples several sections during each repetition time interval but requires several overlapping repetition time intervals to collect all of the phase-encoding steps (Fig 6b) (1). Although radiographics.rsna.org Figure 5. Drawing illustrates how a single-shot FSE sequence can be used to acquire all phase-encoding steps for one image after only one excitation pulse. Gf = frequency-encoding gradient, Gp = phase-encoding gradient, Gs = section-select gradient, TEef = effective echo time. (Adapted, with permission, from reference 1.) the total imaging time for several sections may not be shorter, single-section technique allows the phase-encoding steps from one section to be collected over a shorter time, thus decreasing the opportunity for motion artifact. Without any other modification, this approach usually uses a short repetition time. Single-section technique can be combined with multiecho techniques (echoplanar imaging or FSE) for faster imaging (15). Increasing the ETL to accommodate all of the phaseencoding steps in a given image yields a singleshot image—that is, an image generated by a single excitation pulse. The effective echo time is usually high, yielding T2-weighted images (Fig 5) (5,23). Single-shot FSE techniques are commonly used in abdominal imaging to create heavily T2-weighted images quickly with minimal motion artifact. Partial Fourier technique and parallel imaging independently decrease the number of phase-encoding steps to shorten the ETL of single-shot FSE imaging (15). When both are used simultaneously with single-shot FSE imaging, the ETL is decreased enough to improve FSE-associated blur and help maintain SNR (15). Some sequences even use partial Fourier technique to shorten the effective echo time and achieve moderate T2 weighting by further decreasing the ETL (5). RG • Volume 30 Number 1 Yang et al 193 Figure 6. (a) Drawings illustrate single-section technique, in which all phase-encoding steps for one section are acquired consecutively before the next section is sampled. Gp = phase-encoding gradient, TE = echo time, TR = repetition time. (b) Drawings illustrate the sampling of multiple sections during each overlapping repetition time (TR) interval. Gp = phase-encoding gradient, TE = echo time. (Fig 6 adapted, with permission, from reference 1.) Other Challenges and Potential Solutions Susceptibility Artifact The term susceptibility refers to the ability of a substance to distort the static magnetic field (24). The artifact occurs at the interface of substances with differing magnetic susceptibilities (eg, diamagnetic soft tissue and a ferromagnetic clip). The result is a local field distortion that alters the frequencies of surrounding protons, resulting in dephasing and frequency-encoding errors (4). Lower magnetic field strength decreases frequency-encoding misregistration (2,21). In addition, the number of pixels used to render frequency-encoding misregistration is minimized by increasing the frequency range represented by each pixel (eg, increasing the receiver bandwidth or the frequency-encoding gradient strength) (2,25). Susceptibility-induced field inhomogeneity makes fat saturation difficult, and fat saturation exacerbates susceptibility artifact (Fig 7a, 7b) (4). One should consider not using fat saturation when faced with severe susceptibility artifact. The STIR sequence is a good alternative, with more homogeneous fat suppression for T2-weighted sequences; STIR relies on T1 relaxation differences instead of precessional frequency (chemical shift) differences to null signal from fat, thereby minimizing susceptibility effects during fat-suppressed imaging (17). Refocusing (usually 180°) pulses correct many sources of dephasing, making FSE imaging effective for minimizing susceptibility artifact (Fig 7c) (2,25). In contrast, with single-shot echoplanar imaging, susceptibility is accentuated by the fact that an entire image is obtained with at most a single refocusing pulse (5). Decreased interecho 194 January-February 2010 radiographics.rsna.org Figure 7. Susceptibility artifact from an embolization coil (arrow). (a, b) Blooming is less conspicuous on an axial 3D non-fat-saturated gradient-echo image (3.8/1.3, 12° flip angle) (a), for which the echo time is relatively short, but is more conspicuous due to magnetic field inhomogeneities on a fat-saturated image (3.8/1.3, 12° flip angle) (b). (c) The artifact is least severe on an axial T2-weighted single-shot FSE image (1317/183, 90° flip angle), since multiple refocusing pulses rephase spins. (d, e) More blooming is seen as echo time increases from an axial opposedphase image (195/2.2, 90° flip angle) (d) to an axial in-phase image (195/4.6, 90° flip angle) (e). Teaching Point spacing and shorter echo times further decrease the time available for dephasing (Fig 7d, 7e) (2). Minimizing susceptibility artifact, particularly on gradient-recalled-echo images, hinges on shortening the echo time (2,21). This can be accomplished with fractional echo sampling and by increasing sampling bandwidth and disabling GMN (1) (see “Shortening Echo Time”). Metal artifacts result from a combination ofsusceptibility and conductivity effects (24). Con ductivity effects relate to eddy currents induced by the RF pulses (24). The artifact changes with respect to the metallic object’s orientation and position in the static magnetic field. When pos sible, metal should be placed close to the iso center of and parallel to the static field (26), the orientation of which may differ between openand closed-bore magnets. Wraparound Artifact Wraparound artifact (aliasing) occurs when the FOV is smaller than the imaged object in the phase-encoding direction (Fig 8) (2). The range Teaching Point RG • Volume 30 Number 1 Yang et al 195 Figure 8. Wraparound artifact. Three-dimensional fat-saturated gradient-recalled-echo images (5.6/1.9, 12° flip angle) (a obtained from the superior aspect and b from the inferior aspect of the imaged volume of tissue) demonstrate wraparound artifact in the section-select axis as evidenced by the right kidney mimicking a hepatic lesion (arrow). Additional wraparound artifact is seen in the phase-encoding axis (arrowhead in a). Figure 9. Wraparound artifact. Axial T2-weighted single-shot FSE images (618/180, 90° flip angle) obtained with parallel imaging show how wraparound artifact in the phase-encoding axis (arrow in a) disappears when the FOV is increased (b). of phases represented along the phase-encoding o axis is limited to 360 . Any phase above or below this range is incorrectly assigned to within the 360° range. For instance, an object with a phase of 370°, anterior to the FOV, would be reassigned to 10° (370°–360°), a location within the FOV posteriorly (2). Wraparound is generally acceptable if it occurs over the abdominal wall but can be problematic when it obscures organs of interest. On 3D Fourier transform images, wraparound artifact can also occur in the phase-en- coded section-select axis (Fig 8b). Structures on the first few images may wrap onto the last few images and vice versa (2,4). Phase-encoded wraparound is particularly problematic with parallel imaging techniques that apply coil sensitivity information after Fourier transformation. Wraparound artifact from tissue outside the phase-encoding FOV obscures the center, rather than the periphery, of the image and is caused by unsuccessful unwrapping of the image (Fig 9) (2,27). 196 January-February 2010 radiographics.rsna.org Figure 10. Effect of STIR on fat suppression. (a) T2-weighted fat-saturated FSE image (3000/92) shows how B0 field inhomogeneities contribute to incomplete fat saturation (arrows). (b) STIR image (3320/71) is slightly more noisy but has more uniform fat suppression. Phase-encoded wraparound is eliminated by enlarging the FOV to accommodate the entire object (Fig 9) (4). Although SNR is increased, spatial resolution is decreased (4). One should also consider swapping the phase- and frequency-encoding axes if phase encoding is not already along the shorter axis. Many MR imaging units have a “no phase wrap” option that doubles the number of phaseencoding steps to sample regions outside the FOV (4). The extra spatial information is not reconstructed but is used to unwrap the extraneous structures outside the region of interest. Neither the FOV nor pixel size is changed. If the number of signals averaged is not already minimized, it can be halved to accommodate the doubled number of phase-encoding steps so that imaging time is also unchanged when no phase wrap is used. Because neither the total number of echoes nor pixel size is changed, SNR is maintained. This approach is not an option with many fast techniques used in abdominal imaging, since the number of signals averaged is usually kept to a minimum. As with motion artifact, aliasing can be minimized by suppressing signal from tissue outside the region of interest. Regional coils, such as a surface coil, can be used to limit the area from which echoes are sampled. A spatial saturation pulse can be used on tissues outside the FOV (4). Wraparound artifact can theoretically occur in the frequency-encoding axis when the range of frequencies is not adequately sampled. The Nyquist-Shannon sampling theorem dictates Figure 11. Obscuration of the biliary system by ascites. Oblique coronal heavily T2-weighted thickslab MR cholangiopancreatogram (3414/826, 90° flip angle) demonstrates how ascites (arrowheads) can obscure the biliary system (arrow). that the sampling rate must be at least twice the highest frequency sampled for measurements to be accurate (4,6). Wraparound in the frequencyencoding axis is seldom encountered with newer imaging systems because the digital receiver automatically filters out unwanted high frequencies or makes use of a high bandwidth (2,6). Heterogeneous Fat Suppression Chemical shift saturation relies on the detection of the subtle frequency shift between lipid protons and water protons (3.5 ppm for a 1.5-T magnet) (1,6). Any magnetic field heterogeneity RG • Volume 30 Number 1 Yang et al 197 Figure 12. Dielectric-conductivity effects. (a, b) Axial (1570/265, 90° flip angle) (a) and coronal (3571/170, 90° flip angle) (b) T2-weighted single-shot FSE images obtained at 3.0 T show central signal intensity loss (arrow). (c) Axial 3D nonenhanced fat-saturated gradient-recalled-echo image (5.1/2.6, 13° flip angle) obtained at 3.0 T shows a low-signal-intensity band (arrow) across the left kidney, pancreas, and left hepatic lobe, and a bright band (arrowhead) across the right kidney and right hepatic lobe. (d) On an axial 3D gradient-recalledecho image (4.2/2.1, 10° flip angle) obtained at 1.5 T, the artifacts are greatly reduced. or inconsistent excitation of tissue can obscure this chemical shift. Magnetic field heterogeneity can be improved by shimming the magnet (6). Low-field-strength magnets result in a smaller chemical shift and limit the effectiveness and feasibility of fat saturation. In fact, many systems below 0.7 T do not offer spectral fat saturation. Susceptibility differences—for instance, as seen at air-tissue interfaces or owing to metal implants—distort the magnetic field and make fat saturation more difficult. Some investigators have attempted to fill in the air gaps at skin surfaces with diamagnetic material to decrease susceptibility. STIR provides more homogeneous fat suppression but suffers from poorer SNR (Fig 10). One should consider not using fat saturation when ferromagnetic material causes severe susceptibility artifact (Fig 7a, 7b). Ascites Fluid can obscure visualization of smaller structures such as lymph nodes on T2-weighted images. On heavily T2-weighted maximum intensity projection images (eg, MR cholangiopancreatograms, MR urograms), the bile ducts or ureters can be obscured (Fig 11). By increasing the patient’s girth, ascites increases the FOV and decreases spatial resolution. Ascites can limit diaphragmatic excursion, thereby decreasing the patient’s breath-holding capacity. Preimaging paracentesis may minimize these limitations. Conductive and Dielectric Effects Artifactual bands of decreased or increased signal intensity are sometimes encountered at MR imaging. These bands are thought to relate to B1 inhomogeneities (28,29). They tend to be more severe at 3.0 T (Fig 12), and it is generally advisable not to image patients with abundant fluid (eg, pregnant patients or patients with ascites) at 3.0 T. Some investigators also advise against imaging patients with a large body habitus at 3.0 T (30). Several theories to explain the phenomenon have been put forth, tested, and challenged. 198 January-February 2010 The dielectric or “standing wave” effect is one proposed mechanism. As in water, the high dielectric constant of tissues shortens the RF wavelength. RF pulses at 64 MHz, the Larmor frequency at 1.5 T, have a wavelength of 468 cm in air and 52 cm in water. RF pulses at 128 MHz, the Larmor frequency at 3.0 T, have a wavelength of 234 cm in air and 26 cm in water (31). Dielectric resonance occurs when the RF wavelength approximates the patient’s girth and electromagnetic fields oscillate near his or her natural frequency (28,30,32). The standing waves theoretically cause interference patterns with respect to the RF pulses (28,31). These waves manifest as bands of destructive and constructive interference, which are seen on the image as dark and bright zones, respectively (Fig 12). Because human girth more frequently approximates wavelengths at 3.0 T, B1 inhomogeneities are thought to occur more often at 3.0 T than at 1.5 T (21,33). Some investigators endorse placing dielectric pads to change the patient’s apparent girth to decrease this effect (21,29–31). Others have postulated that RF waves from opposite directions also cause interference; this phenomenon would be different from dielectric resonance and more pronounced at 3.0 T (32,34). Conductivity properties of the human body and of water are also suggested as a possible cause of B1 inhomogeneities, particularly in pregnant patients or patients with ascites because water is highly conductive. Repeated RF pulses, such as the multiple 180° pulses and their associated gradients used in FSE sequences, are thought to induce electrical eddy currents that magnetically oppose the B1 field, resulting in shielding and signal intensity loss (30,32,33). This effect is also worse at higher field strengths radiographics.rsna.org because the Larmor frequency and the frequency of the excitation pulses are higher. Conclusions Body MR imaging is a growing field for which new indications are constantly being developed. In light of the growing concern regarding the harmful effects of ionizing radiation, concomitant growth in the scope and volume of MR imaging is anticipated. Familiarity with basic MR imaging principles is key to the radiologist’s ability to optimize abdominal MR imaging protocols when confronted with challenges. The ability to minimize the effects of motion and other artifacts translates into more efficient protocols and improved institutional efficiency with greater throughput potential. Understanding fundamental MR imaging principles and sequence parameters facilitates management of current techniques and promotes implementation of evolving and future techniques. References 1.Mitchell DG, Cohen M. MRI principles. 2nd ed. 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RG Volume 30 Number 1 January-February 2010 Yang et al Optimizing Abdominal MR Imaging: Approaches to Common Problems Roberta K. Yang, MD, et al RadioGraphics 2010; 30:185–199 • Published online 10.1148/rg.301095076 • Content Codes: Page 186 In the absence of motion, phase differences should be caused solely by the varying strength of the phase-encoding gradient. Page 189 Scanning time (PE steps excitations TR)/(ETL parallel imaging factor, and TR = repetition time. PIF), where PE = phase-encoding, PIF = Page 191 Similar to partial Fourier acquisition, which relies on k-space symmetry in the phase-encoding axis, fractional echo sampling uses k-space symmetry in the frequency-encoding axis. Page 193 Minimizing susceptibility artifact, particularly on gradient-recalled-echo images, hinges on shortening the echo time. This can be accomplished with fractional echo sampling and by increasing sampling bandwidth and disabling GMN. Page 194 Wraparound artifact (aliasing) occurs when the FOV is smaller than the imaged object in the phaseencoding direction.