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Whole-Body High-Field-Strength (3.0-T) MR Imaging in Clinical Practice

Part II. Technical Considerations and Clinical Applications¹

¹ From the Department of Radiology, University of Bonn, Sigmund-Freud-Str 25, Bonn 53105, Germany (C.K.K., F.T., J.G., N.M., W.W., M.v.F., C.M., H.H.S.); and Diagnosezentrum Urania, Vienna, Austria (W.D.). Received October 24, 2006; revision requested October 25; revision received November 28; accepted December 20; final version accepted March 7, 2007; final review and update by C.K.K. December 3. Address correspondence to C.K.K. (e-mail: kuhl{at}uni-bonn.de).

	ABSTRACT

TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

 
This is the second part of a two-part series on the clinical applications of high-field-strength (3.0-T) magnetic resonance (MR) imaging and spectroscopy. In this part, the current level of evidence regarding the use of higher magnetic field strengths for cardiac imaging techniques (including the assessment of cardiac anatomy and function), breast and pelvic imaging, musculoskeletal applications, pediatric imaging, and MR spectroscopy is presented. Published data are interpreted from the perspective of the clinical radiologist. Specific difficulties associated with high-field-strength MR for body imaging and for spectroscopic applications are reviewed and compared with the expected or documented added value of high-field-strength MR for clinical patient care. The overall number of studies published on clinical body high-field-strength MR is still small, and there is evidence for a clinical advantage for selected, but not all, body MR imaging applications. Even without published evidence, clinical experience suggests substantial clinical advantages for musculoskeletal and pediatric applications.

© RSNA, 2008

	INTRODUCTION

TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

 
This is the second part of a two-part series on the clinical applications of high-field-strength (3.0-T) magnetic resonance (MR) imaging and spectroscopy. While the first part (1) provided an overview of the specific physical features of high-field-strength MR imaging and their effect on image quality and efficiency and reviewed the clinical applications for neuroradiology and angiography, this part focuses on high-field-strength MR for body applications. It lists the current level of evidence regarding the use of higher magnetic field strengths for cardiac imaging (including cardiac anatomy and function), abdominal MR, breast and pelvic imaging, musculoskeletal applications, pediatrics, and MR spectroscopy. The aim is to provide an objective survey of the current use and the current published level of evidence regarding the difficulties or advantages that have been identified for body high-field-strength MR in clinical patient care.

	BODY APPLICATIONS OF HIGH-FIELD-STRENGTH MR IMAGING

TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

 
Cardiac MR Imaging
In cardiac imaging, physiologic motion by myocardial contraction and respiration dictate a window during which the data sampling can be performed. Within this short window, high-spatial-resolution coronary MR angiography, or whole-heart motion, perfusion, and late enhancement studies have to be completed. As such, clinical cardiac MR requires the highest possible signal-to-noise ratio (SNR), fast acquisition strategies, exact electrocardiographic timing and respiratory motion control, and pulse sequences that are robust against susceptibility effects caused by the adjacent air-filled pulmonary tissue. The higher radiofrequency deposition and higher artifact-to-noise level due to motion, pulsation, and susceptibility effects that can be anticipated to occur at 3.0 T need to be dealt with before the extra SNR can be used for cardiac MR. The stronger magnetohydrodynamic effects at 3.0 T cause a further elevation of T waves in the electrocardiograms, which may complicate reliable cardiac gating. The relatively small number of publications in this field indicates that it took—and still takes—quite some engineering efforts to overcome these difficulties.

For cine cardiac imaging, steady-state free precession imaging (fast imaging with steady-state precession, or FISP, balanced fast field echo) pulse sequences are commonly used. At 3.0 T, susceptibility or "off-resonance" effects may appear as dark bands or flow ghosting and can substantially degrade image quality of these pulse sequences. Already in 2003, Hinton et al (2) demonstrated that steady-state free precession imaging of short and long axis was feasible and at least equivalent to that at 1.5 T. Three years later, in 2006, Michaely et al (3) and Klumpp et al (4) confirmed that concordant cardiac function parameters are obtained at 3.0 and 1.5 T. They observed off-resonance artifacts in 12 (86%) of 14 subjects. With shorter repetition time, localized shimming, and so-called iterative frequency shifting, these artifacts could be avoided—yet at the expense of an increased overall acquisition time. The SNR and contrast-to-noise ratio (CNR) gains with respect to 1.5 T were less than expected, with a 48% increase of SNR and 9% increase of CNR. With parallel imaging, higher flip angles are feasible and provide a somewhat higher CNR gain (5–7). With a reduction factor of four, acquisition time was reduced from 6–7 minutes down to 1–1.5 minutes, a fact that may help reduce motion artifacts and improve patient compliance (Fig 1). If frequency shifting fails and/or parallel imaging is not available, spoiled gradient-echo sequences (fast low-angle shot, fast field echo) may be tried as an alternative—however, these pulse sequences may, in turn, suffer from saturation effects. In summary, it appears that cardiac cine MR imaging is feasible at 3.0 T. Although the shorter acquisition time should be clinically beneficial, an actual clinical advantage compared to cine imaging at 1.5 T has not been established.

View larger version (54K): [in this window] [in a new window] [Download PPT slide]   Figure 1a: Cardiac function (wall motion study) in a healthy volunteer. Transverse short-axis steady-state free precession MR (balanced turbo field-echo) images (repetition time msec/echo time msec, 3.3/1.66; flip angle 40°; matrix, 160 x 160; field of view, 320 mm; section thickness, 8 mm) at (a) 1.5 and (b) 3.0 T. In a, the three sections were acquired without parallel imaging in three breath holds. In b, the three sections were acquired with parallel imaging (sensitivity encoding [SENSE]), a reduction factor of four, within one 16-second breath hold, and 30 heart phases.

View larger version (57K): [in this window] [in a new window] [Download PPT slide]   Figure 1b: Cardiac function (wall motion study) in a healthy volunteer. Transverse short-axis steady-state free precession MR (balanced turbo field-echo) images (repetition time msec/echo time msec, 3.3/1.66; flip angle 40°; matrix, 160 x 160; field of view, 320 mm; section thickness, 8 mm) at (a) 1.5 and (b) 3.0 T. In a, the three sections were acquired without parallel imaging in three breath holds. In b, the three sections were acquired with parallel imaging (sensitivity encoding [SENSE]), a reduction factor of four, within one 16-second breath hold, and 30 heart phases.

View larger version (160K): [in this window] [in a new window] [Download PPT slide]   Figure 2a: Comparison of coronary angiography in the same patient. Free-breathing, electrocardiographically gated, navigator-corrected, segmented three-dimensional turbo gradient-echo (a) 1.5- and (b) 3.0-T MR angiograms (4.2/1.22, 15° flip angle) in a patient suspected of having coronary artery disease. Note the equivalent image quality. The stenoses of the proximal left anterior descending and right circumflex arteries (arrows) are depicted with at least equivalent conspicuity.

View larger version (161K): [in this window] [in a new window] [Download PPT slide]   Figure 2b: Comparison of coronary angiography in the same patient. Free-breathing, electrocardiographically gated, navigator-corrected, segmented three-dimensional turbo gradient-echo (a) 1.5- and (b) 3.0-T MR angiograms (4.2/1.22, 15° flip angle) in a patient suspected of having coronary artery disease. Note the equivalent image quality. The stenoses of the proximal left anterior descending and right circumflex arteries (arrows) are depicted with at least equivalent conspicuity.

The delineation of "late" or "delayed" (hyper)enhancement has become the method of choice to investigate myocardial viability; where it is available, it has virtually replaced scintigraphic methods. Klumpp and co-workers (4) conducted a study on a group of 40 patients with a history of myocardial infarction who underwent viability (late enhancement) imaging at either 1.5 or 3.0 T. The authors found a slightly higher image quality at 3.0 T. An intraindividual comparison of viability imaging was performed by Cheng et al (13). They investigated 16 patients with acute and chronic myocardial infarction who underwent contrast material–enhanced viability imaging at 1.5 and 3.0 T, with the same segmented turbo fast low-angle shot MR pulse sequence. Almost perfect agreement was found for myocardial hyperenhancement at 1.5 T and 3.0 T. The calculated mass and the respective transmural extent of suspected myocardial infarction were equivalent. Accordingly, there is evidence that viability imaging is as feasible at 3.0 T as it is at 1.5 T (Fig 3)—but without specific advantage compared to 1.5 T.

View larger version (172K): [in this window] [in a new window] [Download PPT slide]   Figure 3a: Comparison of cardiac viability imaging (late enhancement) in a 56-year-old male patient with transmural apical infarction. T1-weighted three-dimensional turbo gradient-echo (a) 3.0- (repetition time msec/echo time msec/inversion time msec, 4.3/1.24/240; 15° flip angle) and (b) 1.5-T (4.3/1.24, 15° flip angle) MR images adapted to the T1 of the myocardium as determined by a Look-Locker pulse sequence; 10 5-mm-thick sections acquired in a 12-second breath hold to cover the horizontal long-axis view (four-chamber view). Image quality and delineation of the infarcted myocardium (arrow) are equivalent at 3.0 and 1.5 T.

View larger version (179K): [in this window] [in a new window] [Download PPT slide]   Figure 3b: Comparison of cardiac viability imaging (late enhancement) in a 56-year-old male patient with transmural apical infarction. T1-weighted three-dimensional turbo gradient-echo (a) 3.0- (repetition time msec/echo time msec/inversion time msec, 4.3/1.24/240; 15° flip angle) and (b) 1.5-T (4.3/1.24, 15° flip angle) MR images adapted to the T1 of the myocardium as determined by a Look-Locker pulse sequence; 10 5-mm-thick sections acquired in a 12-second breath hold to cover the horizontal long-axis view (four-chamber view). Image quality and delineation of the infarcted myocardium (arrow) are equivalent at 3.0 and 1.5 T.

View larger version (105K): [in this window] [in a new window] [Download PPT slide]   Figure 4a: Comparison of cardiac adenosine stress MR perfusion imaging at (a) 3.0 T (3.7/1.8, 15° flip angle) and (b) 1.5 T (3.7/1.8, 20° flip angle) in two patients with atypical angina by using T1-weighted saturation-recovery k-space segmented gradient-echo sequence with parallel imaging (SENSE). The perfusion study in b is meant to give a representative example of the image quality that is typically obtainable at 1.5 T. There is a stress-induced perfusion deficit in the anterolateral and lateral myocardium (arrow). Note the substantially higher SNR in a.

View larger version (78K): [in this window] [in a new window] [Download PPT slide]   Figure 4b: Comparison of cardiac adenosine stress MR perfusion imaging at (a) 3.0 T (3.7/1.8, 15° flip angle) and (b) 1.5 T (3.7/1.8, 20° flip angle) in two patients with atypical angina by using T1-weighted saturation-recovery k-space segmented gradient-echo sequence with parallel imaging (SENSE). The perfusion study in b is meant to give a representative example of the image quality that is typically obtainable at 1.5 T. There is a stress-induced perfusion deficit in the anterolateral and lateral myocardium (arrow). Note the substantially higher SNR in a.

View larger version (139K): [in this window] [in a new window] [Download PPT slide]   Figure 5a: Turbo gradient-echo (turbo field-echo) MR images (3.6/2.0, 10° flip angle, parallel imaging with a reduction factor of 2.5, tagged short-axis sections, 14 phases per cardiac cycle, tag separation of 6 mm) in the same patient show myocardial tagging in systole and diastole at (a) 1.5 and (b) 3.0 T. Note that on b, the tags persist until the late diastole.

View larger version (134K): [in this window] [in a new window] [Download PPT slide]   Figure 5b: Turbo gradient-echo (turbo field-echo) MR images (3.6/2.0, 10° flip angle, parallel imaging with a reduction factor of 2.5, tagged short-axis sections, 14 phases per cardiac cycle, tag separation of 6 mm) in the same patient show myocardial tagging in systole and diastole at (a) 1.5 and (b) 3.0 T. Note that on b, the tags persist until the late diastole.

In summary, coronary vessels, myocardial motion, and myocardial viability (late enhancement) can be investigated at 3.0 T, but currently without established clinical advantage compared to 1.5 T. Myocardial perfusion and strain analysis (tagging) are indeed improved at 3.0 T, although results from clinical studies are still scarce. New approaches (vessel wall imaging, end-systole imaging) are pursued at 3.0 T that promise to further improve the early diagnosis and risk stratification in patients with arteriosclerotic and myocardial diseases.

Breast
Contrast-enhanced MR is currently the most sensitive imaging modality to detect and stage primary and recurrent invasive or intraductal cancer of the breast. It has become the new reference standard for early diagnosis, preoperative staging, and follow-up of patients after breast conserving surgery (26,27). It is increasingly used for screening women at increased risk for breast cancer.

In breast MR, cancer is identified by its early and strong enhancement. The technical requirements for contrast-enhanced breast MR are similar to those for contrast-enhanced MR angiography: For the depiction of breast cancer (as for the depiction of arterial vessels), it is necessary to acquire images with high spatial resolution over a large field of view within a relatively short period of time. The time constraints in MR angiography are due to the fact that the arterial enhancement fades rapidly and is masked by venous enhancement. In breast MR, breast cancer tends to washout the contrast agent and may be masked by the progressively enhancing normal fibroglandular tissue. Parallel imaging has been used at 1.5 T to fit higher imaging matrices into the acquisition window—however, this may yield borderline SNR, since the SENSE-related SNR penalty adds to the reduced SNR associated with higher matrix imaging. The higher SNR at higher fields should be as beneficial for breast MR as it is for contrast-enhanced angiography. Yet, unlike contrast-enhanced MR angiography, the published evidence on clinical high-field-strength breast MR is still very limited. There is a report (28) on an intraindividual study on breast MR at 1.5 and 3.0 T in 37 women with a total of 53 contrast-enhancing lesions. The results suggested a significantly higher diagnostic accuracy for breast MR at 3.0 T compared with 1.5 T. However, in that same article, we demonstrated that enhancement of lesions at 3.0 T was lower than the values obtained for the same lesions at 1.5 T, despite the longer T1 relaxation times at 3.0 T. As it appears, the heterogeneous radiofrequency transmission at 3.0 T may lead to local variations of the flip angle across the field of view—the actual excitation angles may be substantially lower than their prescribed values (29,30). In particular with two-dimensional gradient echo—a pulse sequence commonly used for dynamic subtracted breast MR—variable flip angles lead to substantial variations of the T1 contrast within the same MR image across the field of view. A spatially variable T1 contrast will yield variable enhancement of lesions depending on their location within the field of view, which is unacceptable for breast imaging. The difficulty is that these B₁ inhomogeneities may go undetected, because the MR images may appear deceivingly normal (Fig 6).

View larger version (90K): [in this window] [in a new window] [Download PPT slide]   Figure 6a: Comparison of contrast-enhanced bilateral dynamic breast MR imaging at (a) 1.5 T (290/4.6, 90° flip angle) versus (b) 3.0 T (290/2.3, 73° flip angle) in a patient with multicentric invasive duct cancer. Transverse maximum intensity projections of the first postcontrast subtracted images. Note the large breast cancer, which is located in the upper inner quadrant. In addition to the large mass, there were multiple foci which, at mastectomy, were confirmed to represent further foci of the known invasive cancer. More of these lesions are visualized in Figure 7.

View larger version (83K): [in this window] [in a new window] [Download PPT slide]   Figure 6b: Comparison of contrast-enhanced bilateral dynamic breast MR imaging at (a) 1.5 T (290/4.6, 90° flip angle) versus (b) 3.0 T (290/2.3, 73° flip angle) in a patient with multicentric invasive duct cancer. Transverse maximum intensity projections of the first postcontrast subtracted images. Note the large breast cancer, which is located in the upper inner quadrant. In addition to the large mass, there were multiple foci which, at mastectomy, were confirmed to represent further foci of the known invasive cancer. More of these lesions are visualized in Figure 7.

View larger version (29K): [in this window] [in a new window] [Download PPT slide]   Figure 7: Images in the same patient as in Figure 6. Left: Sagittal high-spatial-resolution actively fat-suppressed postcontrast three-dimensional (3D) turbo gradient-echo (10/2.6, 10° flip angle) imaging. FA = flip angle, FFE = fast field echo, FOV = field of view, TE = echo time, TR = repetition time. Middle: Sagittal diffusion-weighted parallel imaging with a reduction [SENSE] factor of two. Sshot = single shot. Right: Sagittal turbo gradient-echo T2* first-pass perfusion imaging (principles of echo shifting with a train of observation [PRESTO], SENSE factor of two) displays the parametric map revealing maximum perfusion-mediated signal intensity loss as surrogate for regional tumor perfusion volume. Note the restricted diffusion at the site of the cancer and the very strong perfusion on the perfusion map.

To date, to our knowledge, there are only two reports on high-field-strength liver MR in patients with focal liver lesions (43,44). Both are carefully designed intraindividual clinical studies that evaluate not only image quality but also the diagnostic accuracy regarding the detection and classification of focal liver lesions. Von Falkenhausen et al (43) investigated 21 patients with a total of 79 benign and malignant focal liver lesions (Fig 8); Chang et al (44) performed superparamagnetic iron oxide (SPIO)-enhanced MR in 35 patients with 55 malignant lesions. Authors of both studies used parallel imaging to avoid SAR-related difficulties. Results of those studies are concordant in that dielectric effects were rare; they were seen in one of 21 patients and two of 35 patients, which indicates that these effects may be less problematic than previously anticipated. However, the prevalence of these artifacts will depend strongly on the composition of the patient cohort—they will always be problematic if patients with ascites or other causes of large intraabdominal fluid collections are seen.

View larger version (118K): [in this window] [in a new window] [Download PPT slide]   Figure 8a: Comparison of liver imaging in a 52-year-old male patient who underwent transverse respiratory-triggered non–fat-suppressed T2-weighted turbo SE liver MR at (a) 1.5 and (b) 3.0 T without parallel imaging or flip angle modulation. Repetition time of 3000–5000 msec (depending on respiration), echo time of 80 msec, imaging matrix of 256 x 256, acquisition time of 2 minutes 34 seconds for a and 3 minutes 4 seconds for b. Note that the liver parenchyma appears darker on b, as does the spleen and bone marrow. Image quality is comparable to that of a. No increased artifact level and dielectric effects are visible.

View larger version (118K): [in this window] [in a new window] [Download PPT slide]   Figure 8b: Comparison of liver imaging in a 52-year-old male patient who underwent transverse respiratory-triggered non–fat-suppressed T2-weighted turbo SE liver MR at (a) 1.5 and (b) 3.0 T without parallel imaging or flip angle modulation. Repetition time of 3000–5000 msec (depending on respiration), echo time of 80 msec, imaging matrix of 256 x 256, acquisition time of 2 minutes 34 seconds for a and 3 minutes 4 seconds for b. Note that the liver parenchyma appears darker on b, as does the spleen and bone marrow. Image quality is comparable to that of a. No increased artifact level and dielectric effects are visible.

In the study by von Falkenhausen et al (43), the degree of artifact degradation at 3.0 T and 1.5 T was rated equivalent, including susceptibility or pulsation artifacts and image homogeneity. As a result, in that study the overall image quality was equivalent at 1.5 T in 20 of 21 patients. Probably because of the more satisfactory image quality, the detection rates of focal liver lesions were equivalent, with 76 of 79 lesions detected at 1.5 T and 77 of 79 detected at 3.0 T.

In summary, liver MR at 3.0 T, especially with an SPIO-based contrast agent, does have the potential to improve the diagnosis of focal liver lesions—provided that the degree of artifacts is decreased to a 1.5 T-equivalent level. For the time being, however, this is not always achieved and/or there may exist vendor-specific differences in how successfully artifacts are controlled (Figs 8, 9). High-field-strength liver imaging will virtually always yield nondiagnostic image quality in patients with marked ascites.

View larger version (113K): [in this window] [in a new window] [Download PPT slide]   Figure 9a: Comparison of liver imaging without and that with parallel imaging and flip angle modulation at 3.0 T in a healthy 38-year-old volunteer. Transverse T2-weighted respiratory-triggered fat-suppressed (a) turbo SE (acquisition time, 3 minutes 51 seconds) and (b) single-shot turbo SE with parallel imaging (SENSE factor of two) and flip angle modulation (flip angle sweep, 70°) acquisition time, 1 minute 15 seconds) images. Image quality is equivalent, if not improved, in b.

View larger version (124K): [in this window] [in a new window] [Download PPT slide]   Figure 9b: Comparison of liver imaging without and that with parallel imaging and flip angle modulation at 3.0 T in a healthy 38-year-old volunteer. Transverse T2-weighted respiratory-triggered fat-suppressed (a) turbo SE (acquisition time, 3 minutes 51 seconds) and (b) single-shot turbo SE with parallel imaging (SENSE factor of two) and flip angle modulation (flip angle sweep, 70°) acquisition time, 1 minute 15 seconds) images. Image quality is equivalent, if not improved, in b.

Pelvis
The MR diagnosis of early prostate cancer is difficult, and often enough, a simple digital rectal examination is more fruitful than highly sophisticated MR protocols. Systematic sextant biopsy is still considered the standard of reference—to the dismay of patients who have to undergo this crude (and seemingly atavistic) diagnostic procedure. The diagnostic accuracy afforded by MR is not ideal; published sensitivity rates even for advanced disease states (ie, with extraglandular infiltration) vary between 22% and 91% (47–51), with a mean overall sensitivity of 48%. Central and transition zone cancers are difficult or even impossible to diagnose, in particular in the presence of benign prostatic hyperplasia. Yet, even in the peripheral zone, a substantial number of cancers go undetected, for example, cancers with more permeative growth pattern. Moreover, focal prostatic atrophy or prostatitis may mimic cancer and cause false-positive results. In view of the increasing number of individuals who undergo prostate-specific antigen screening, however, there is an ever increasing need to improve the detection of prostate cancer at an earlier stage. This has been attempted by using "functional" prostate imaging, that is, dynamic contrast-enhanced imaging, proton MR spectroscopy, and diffusion-weighted imaging (52–64).

Higher magnetic field strengths should allow structural imaging of the prostate with improved spatial resolution, thus improving the detection and staging of peripheral zone cancers. In addition, functional imaging of the prostate at high field strength (dynamic contrast-enhanced imaging, MR spectroscopy, and diffusion-weighted imaging) should improve the detection of central and transition zone cancers and should improve the confidence with which any "low-signal-intensity tissue" in the peripheral zone can be predicted to represent prostate cancer.

There is one intraindividual comparative study (65) on 10 patients who all had prostate cancer and who underwent prostate imaging at 1.5 T (with endorectal plus phased-array coil) and at 3.0 T (with endorectal coil only). SNR at 3.0 T was high enough to allow T2-weighted imaging with voxel sizes of 0.13 mm³, compared to 1.21 mm³ at 1.5 T, and to speed up the sampling rate in the dynamic contrast-enhanced sequence (from 2 seconds down to 1 second). The same group published a follow-up study on very-high-spatial-resolution T2-weighted imaging in 31 patients with prostate cancer. Extracapsular extension was identified with a sensitivity and specificity of 88% and 99%, respectively (66), and even minimal capsular invasion—usually not identifiable at 1.5 T imaging—was visible in two of three patients. So although there was no direct 3.0 T versus 1.5 T comparison available, there is indirect evidence that the high-spatial-resolution studies did indeed translate into an improved staging accuracy compared to that at 1.5 T. If such high-spatial-resolution studies are offered to patients, however, it will be important to carefully use the information they provide. The prognostic implications of minimal capsular invasion are unclear to date. Men with minimally invasive disease may do equally well after surgery as those with cancers confined to the gland. It will be important to investigate the clinical effect of "minimal capsular infiltration" in order to prevent patients with allegedly inoperable (extracapsular) disease from being withheld curative surgery.

In prostate MR spectroscopy, the broader separation of metabolite resonances at 3.0 T (see Proton MR Spectroscopy) does improve the delineation of choline and creatine resonances—with the disadvantage of a disfigured citrate resonance peak, which may impair the calculation of a choline-to-citrate ratio (67–69). Fast dynamic contrast-enhanced, diffusion and diffusion-tensor imaging examinations have already been successfully performed on prostate cancer at 3.0 T (70–73). Prospective clinical trials that would compare the sensitivity and specificity of structural and functional imaging at 3.0 T versus 1.5 T for detecting small prostate cancers and for local staging of prostate cancer are, however, not yet available.

In summary, there is a clinical demand to improve prostate MR imaging, because standard 1.5-T imaging does not offer a satisfactory diagnostic accuracy. There is some anecdotal evidence that high-spatial-resolution structural 3.0-T imaging might improve the accuracy of prostate cancer staging. In addition, there is prospect that 3.0-T functional prostate imaging should help improve the detection of small, especially central or transition zone, cancers and should help prevent false-positive diagnoses.

Another approach for prostate MR at 3.0 T is to exploit the higher SNR by using phased-array coils instead of endorectal coils, which may increase the acceptance of this technique (74–76). However, the improved patient comfort occurs at the expense of the potential diagnostic advantages resulting from high-field-strength high-spatial-resolution imaging.

For the female pelvis, high-spatial-resolution imaging with voxel sizes of 0.83 mm³ has also been proposed to improve the diagnosis of gynecologic pelvic disorders (76) (Figs 10–12). Compared to the spatial resolution of state-of-the-art 1.5-T protocols (2.99 mm³), image quality and the detectability of fine anatomic details at 3.0 T were improved. Morakkabati-Spitz et al (77) reported that in four of 23 patients, some additional diagnostic information was obtained: however, this was of clinical relevance in only one patient (4%; higher confidence in excluding bladder infiltration). One may conclude that such very-high-spatial-resolution imaging may not be needed to diagnose disease and/or investigate disease extent in the majority of patients undergoing gynecologic pelvic MR examination. Alternatively, the SNR can be spent to allow very fast pelvic MR. With the use of parallel imaging with high acceleration factors, combined with flip angle modulation techniques, the scan time for a T2-weighted fast (turbo) SE pulse sequence with a full 512 x 512 imaging matrix could be decreased from 4 minutes to 22 seconds. Image quality and diagnostic accuracy were investigated in a series of 33 patients and were rated to be even improved compared with the standard protocol due to the absence of motion artifacts with the very fast protocol (77–79) (Fig 13).

View larger version (91K): [in this window] [in a new window] [Download PPT slide]   Figure 10: Pelvic MR at 3.0 T versus 1.5 T in a 35-year-old patient. Effect of field strength on the achievable spatial resolution. Left: Transverse MR at 1.5 T with the standard protocol (512 x 400 imaging matrix). Fatty tissue septae are hardly discernible. Middle: Transverse high-spatial-resolution MR at 3.0 T with parallel imaging (SENSE with a reduction factor of three). Clear resolution of fatty tissue septae. Right: Attempt to copy the high spatial resolution protocol to the 1.5 T system (transverse orientation). Multiple signals were required and resulted in an overall acquisition time of 11 minutes 5 seconds, despite the use of parallel imaging with a SENSE factor of two. Although at first glance, the image obtained with high spatial resolution at 1.5 T (right) has overall image quality seems sufficient for diagnostic purposes or is even improved compared with that of the standard protocol (left), SNR is in fact too low to exploit the high spatial resolution—as is documented by the detailed view of the fatty tissue. So although in theory, the same imaging matrices can be acquired at 1.5 T if only long enough acquisition times are used, this example shows that this may result in only "virtual" spatial resolution, which will not be useful on clinical grounds.

View larger version (104K): [in this window] [in a new window] [Download PPT slide]   Figure 11a: Comparison of standard and high-spatial-resolution imaging of the female pelvis at 3.0 T. Transverse T2-weighted turbo SE images in a 52-year-old patient with cervical cancer who underwent pelvic MR with (a) the protocol that is standard at 1.5 T (matrix, 512 x 400; section thickness, 4 mm) and (b) 3.0-T (matrix, 1024 x 600; section thickness, 4 mm) protocol. Delineation of tumor extent is equivalent. A small paracervical lymph node metastasis had been overlooked on a, whereas it was prospectively diagnosed on the basis of the assessment of b.

View larger version (108K): [in this window] [in a new window] [Download PPT slide]   Figure 11b: Comparison of standard and high-spatial-resolution imaging of the female pelvis at 3.0 T. Transverse T2-weighted turbo SE images in a 52-year-old patient with cervical cancer who underwent pelvic MR with (a) the protocol that is standard at 1.5 T (matrix, 512 x 400; section thickness, 4 mm) and (b) 3.0-T (matrix, 1024 x 600; section thickness, 4 mm) protocol. Delineation of tumor extent is equivalent. A small paracervical lymph node metastasis had been overlooked on a, whereas it was prospectively diagnosed on the basis of the assessment of b.

View larger version (159K): [in this window] [in a new window] [Download PPT slide]   Figure 12a: (a, b) Comparison of standard and high-spatial-resolution imaging of the female pelvis at 3.0 T in a 48-year-old patient with strongly hypervascularized myometrial myoma. Sagittal T2-weighted turbo SE sequence with (a) the protocol that is standard at 1.5 T and (b) with very high spatial resolution. Because of flip angle sweep (75°), the acquisition time was the same for both sequences. Note that despite the use of flip angle sweep, the image contrast is equivalent. The pathologic subserosal vessels are better delineated in b—but diagnosis can be made with equivalent confidence using both pulse sequences.

View larger version (196K): [in this window] [in a new window] [Download PPT slide]   Figure 12b: (a, b) Comparison of standard and high-spatial-resolution imaging of the female pelvis at 3.0 T in a 48-year-old patient with strongly hypervascularized myometrial myoma. Sagittal T2-weighted turbo SE sequence with (a) the protocol that is standard at 1.5 T and (b) with very high spatial resolution. Because of flip angle sweep (75°), the acquisition time was the same for both sequences. Note that despite the use of flip angle sweep, the image contrast is equivalent. The pathologic subserosal vessels are better delineated in b—but diagnosis can be made with equivalent confidence using both pulse sequences.

View larger version (146K): [in this window] [in a new window] [Download PPT slide]   Figure 13a: Comparison of (a) standard 3.0-T pelvic MR (acquisition time, 4 minutes 15 seconds) and (b) very fast 3.0-T pelvic MR with parallel imaging and flip angle modulation (acquisition time, 39 seconds; SENSE factor of two; flip angle sweep, 75°) in 29-year-old female patient with dermoid (arrow). (a, b) Transverse T2-weighted turbo SE images with 512 x 400 imaging matrix and 4-mm section thickness. Image contrast is preserved (or even improved) in b. Motion artifacts present in a are absent in b. No dielectric artifacts are present. Both pulse sequences are performed in a free-breathing mode.

View larger version (143K): [in this window] [in a new window] [Download PPT slide]   Figure 13b: Comparison of (a) standard 3.0-T pelvic MR (acquisition time, 4 minutes 15 seconds) and (b) very fast 3.0-T pelvic MR with parallel imaging and flip angle modulation (acquisition time, 39 seconds; SENSE factor of two; flip angle sweep, 75°) in 29-year-old female patient with dermoid (arrow). (a, b) Transverse T2-weighted turbo SE images with 512 x 400 imaging matrix and 4-mm section thickness. Image contrast is preserved (or even improved) in b. Motion artifacts present in a are absent in b. No dielectric artifacts are present. Both pulse sequences are performed in a free-breathing mode.

Dielectric effects have been reported by some authors (82) to substantially degrade image quality in MR of the pelvis, whereas others (65,66,77–79) did not observe them at all. Again, as with liver MR at 3.0 T, there seem to be system- or vendor-specific differences regarding the overall prevalence and the degree of dielectric effects.

Musculoskeletal Applications
Only a few clinical studies on high-field-strength musculoskeletal imaging have been published so far. This, however, is clearly not a reflection of its current clinical use. Initial publications on 3.0-T musculoskeletal imaging reported long acquisition times that were due to the fact that, for SAR-related reasons, the imaging volume had to be split into multiple packages (83–86). Meanwhile, parallel imaging and refocusing flip angle modulation techniques are available that help settle these SAR- and susceptibility-induced difficulties (Fig 14). Since these acquisition strategies (and dedicated surface coils) are available on virtually all 3.0-T systems, musculoskeletal imaging has emerged as one of the most important clinical applications of high-field-strength MR: Faster imaging with higher spatial resolution is routinely available. Imaging at 3.0 T offers a more robust fat suppression than imaging at 1.5 T, including also off-center musculoskeletal applications such as for the shoulder or the hip (87) (Fig 15). In addition, the fat saturation pulses are shorter, allowing one to acquire more sections per repetition time compared with 1.5 T (Fig 15). The substantially faster acquisition can be used to reduce the overall table time for musculoskeletal examinations—and since musculoskeletal examinations tend to be "high-volume" applications, in particular in private practices, this can be used to substantially increase patient throughput (88).

View larger version (177K): [in this window] [in a new window] [Download PPT slide]   Figure 14a: MR of the knee in a 26-year-old male patient with vertical femoral cartilage tear and joint effusion. Value of parallel imaging and dedicated multielement extremity coils for fast, high-spatial-resolution extremity imaging by using an eight-element phased-array coil at 3.0 T. (a) Sagittal intermediate-weighted turbo SE (matrix, 464 x 1024; noninterpolated in-plane resolution, 0.36 mm; SENSE reduction factor of 1.5; total acquisition time, 97 seconds) image. (b) Transverse intermediate-weighted driven equilibrium, or DRIVE, sequence with fat suppression (spectral presaturation attenuated by inversion recovery, or SPAIR) (noninterpolated in-plane pixel size, 0.55 mm; acquisition time, 2 minutes 20 seconds).

View larger version (119K): [in this window] [in a new window] [Download PPT slide]   Figure 14b: MR of the knee in a 26-year-old male patient with vertical femoral cartilage tear and joint effusion. Value of parallel imaging and dedicated multielement extremity coils for fast, high-spatial-resolution extremity imaging by using an eight-element phased-array coil at 3.0 T. (a) Sagittal intermediate-weighted turbo SE (matrix, 464 x 1024; noninterpolated in-plane resolution, 0.36 mm; SENSE reduction factor of 1.5; total acquisition time, 97 seconds) image. (b) Transverse intermediate-weighted driven equilibrium, or DRIVE, sequence with fat suppression (spectral presaturation attenuated by inversion recovery, or SPAIR) (noninterpolated in-plane pixel size, 0.55 mm; acquisition time, 2 minutes 20 seconds).

View larger version (115K): [in this window] [in a new window] [Download PPT slide]   Figure 15a: Indirect arthrography of the shoulder at 3.0 T in a 31-year-old male patient with a superior labral anterior posterior, or SLAP, lesion of the right shoulder. (a) Transverse fat-suppressed T1-weighted high-spatial-resolution gradient-echo arthrogram with isotropic 0.8-mm resolution acquired with a two-element flexible surface coil and parallel imaging (SENSE reduction factor of two; total acquisition time, 4 minutes 20 seconds). Reformatted images in (b) sagittal and (c) coronal orientation. Fat suppression is also homogeneous in off-center locations. Because of the isotropic acquisition, multiplanar reformatting is feasible in b and c.

View larger version (76K): [in this window] [in a new window] [Download PPT slide]   Figure 15b: Indirect arthrography of the shoulder at 3.0 T in a 31-year-old male patient with a superior labral anterior posterior, or SLAP, lesion of the right shoulder. (a) Transverse fat-suppressed T1-weighted high-spatial-resolution gradient-echo arthrogram with isotropic 0.8-mm resolution acquired with a two-element flexible surface coil and parallel imaging (SENSE reduction factor of two; total acquisition time, 4 minutes 20 seconds). Reformatted images in (b) sagittal and (c) coronal orientation. Fat suppression is also homogeneous in off-center locations. Because of the isotropic acquisition, multiplanar reformatting is feasible in b and c.

View larger version (75K): [in this window] [in a new window] [Download PPT slide]   Figure 15c: Indirect arthrography of the shoulder at 3.0 T in a 31-year-old male patient with a superior labral anterior posterior, or SLAP, lesion of the right shoulder. (a) Transverse fat-suppressed T1-weighted high-spatial-resolution gradient-echo arthrogram with isotropic 0.8-mm resolution acquired with a two-element flexible surface coil and parallel imaging (SENSE reduction factor of two; total acquisition time, 4 minutes 20 seconds). Reformatted images in (b) sagittal and (c) coronal orientation. Fat suppression is also homogeneous in off-center locations. Because of the isotropic acquisition, multiplanar reformatting is feasible in b and c.

In view of the excellent accuracy with which musculoskeletal disorders are diagnosed already at 1.5 T, there is probably not much room for improvement. A candidate application that may indeed benefit from high field strength is the diagnosis of cartilage disease. In recent years, new articular cartilage repair techniques such as mosaicplasty or autologous chondrocyte transplantation and other therapeutic options such as microfracturing have been introduced. These procedures necessitate an accurate presurgical imaging of the cartilage surfaces. In a porcine model and in cadaveric tali, the detection of subtle cartilage and osteochondral lesions was improved at 3.0 T (91–96). New approaches for assessing cartilage integrity are diffusion imaging and quantitative MR, both of which should demonstrate proteoglycan depletion that occurs very early in the course of cartilage degeneration. The spatial resolution that is obtainable for quantitative MR and diffusion-weighted imaging at 1.5 T is insufficient, such that so far, these examinations were performed only ex vivo by using MR microscopy systems. At 3.0 T, diffusion-weighted imaging and quantitative MR were successfully performed on porcine and bovine knee specimens, with a sub-150-µm resolution (97). Meanwhile, contrast-enhanced imaging, T2 maps, and diffusion-weighted imaging have already been used at 3.0 T to assess the repair tissue after autologous chondrocyte transplantation in 15 patients (published in three different articles, 98–100).

Another area that should benefit from 3.0 T is the evaluation of smaller joints such as the wrist (Fig 16) or the interphalangeal joints (83). Hyaline cartilage surfaces, interosseous ligaments and nerves of the hand, the carpal tunnel, and the wrist are usually not resolved on standard 1.5-T images, even if dedicated surface coils are used (101). Standard wrist imaging at 1.5 T usually takes well beyond 12–14 minutes; the resolution required for the detection of fine anatomic structures would require exceedingly long acquisition times, which would (over)stretch patient tolerance. At 3.0 T, with multielement coils and parallel imaging, it is feasible to acquire these images in less than 3 minutes.

View larger version (190K): [in this window] [in a new window] [Download PPT slide]   Figure 16: MR of the wrist at 3.0 T. High-spatial-resolution intermediate-weighted turbo SE (1581/32; turbo factor, 6; 2-mm section thickness; 80-mm field of view; in-plane pixel size, 0.3 mm; parallel imaging (SENSE with a reduction factor of 1.3; total acquisition time, 167 seconds) image.

Proton MR Spectroscopy
In MR spectroscopy, as in MR imaging, the signal should increase about linearly with the field strength (102,103). Relaxation time measurements of spectroscopic metabolites consistently reported shorter T2 relaxation times of cerebral metabolites, whereas no consistent effect was observed on metabolite T1 values (104,105).

Due to the shorter metabolite T2, MR spectroscopy with long echo time, around 280 msec—which is used for the in-phase detection of lactate—is not suitable for 3.0 T and has to be replaced by anti-phase measurements at echo time of 140 msec or by short echo-time (<40 msec) acquisitions. The anti-phase detection of lactate, however, is problematic for high-field-strength MR spectroscopy with point-resolved spatially localized spectroscopy, or PRESS, localization and may yield smaller, sometimes noninverted lactate signals (106,107).

The higher SNR at 3.0 T may be used to reduce the size of the excitation volume (voxel) and, thus, to reduce partial volume effects. This, in turn, may improve the sensitivity and specificity with which metabolic abnormalities can be identified especially in disease states that go along with more focal abnormalities (102,103). Alternatively, if the voxel size is kept unchanged, the signal gain may be used to reduce the acquisition time, for example, in uncooperative patients or in clinical situations where MR spectroscopy is added to an already extensive imaging protocol. Time-resolved, functional proton MR spectroscopic studies, for example, of cerebral glucose metabolism or of neurotransmitter kinetics (glutamate and glutamine complex) with a temporal resolution of 1 minute or less, become feasible at 3.0 T and may be combined with cerebral activation studies.

Because of the higher chemical shift dispersion at an increasing magnetic field, the absolute frequency separation (in Hertz) of metabolite lines is doubled at 3.0 T compared with 1.5 T. This should improve the spectral resolution—provided that the absolute line widths of the individual metabolites remain unchanged. This latter assumption is often taken for granted; however, it is far from being true. First, the absolute homogeneity of magnets decreases with increasing field strength—and even with sophisticated engineering tools, magnetic field inhomogeneities may persist. The reduced homogeneity will shorten T2* and lead to broader line widths even in the absence of tissue-related intravoxel susceptibility changes. Second, the field-dependent decrease of metabolite T2 will contribute to an increased line width per se. Although this is no major issue for the singlet lines of the "main" MR-visible metabolites with long T2, that is, N-acetylaspartate, total creatine, and choline compounds, it will decrease the accuracy of water suppression in muscle or liver MR spectroscopy. Third (and most important), increased susceptibility effects at tissue interfaces and around local iron deposits cause an additional T2* shortening and thus lead to patient- and region-dependent additional line broadening and signal losses, most prominent in body MR spectroscopy and in cerebral studies involving the temporal lobes (103) and other basal brain regions (108). These effects can only partly be overcome by improved shimming. With smaller voxels and with advanced shimming tools that come with the new generation of actively shielded 3.0-T magnets, spectral resolution is indeed improved by about 30%–50% with respect to that at 1.5 T—at least in "suitable" brain regions such as the parietal and occipital lobes.

If an improved spectral resolution is indeed achievable, water-suppressed MR spectroscopy at 3.0 T with short echo times can be used to assess metabolites that overlap (or are superimposed) at 1.5 T. The increased chemical shift dispersion at 3.0 T transforms them into first-order spectra with better separation of multiplets. For example, as opposed to 1.5 T, at 3 T, myo-inositol is not superimposed by the CH₂ multiplet of glutamate and glutamine complex (Fig 17). This allowed us to detect elevated myo-inositol concentrations in normal-appearing white matter in a group of 45 patients with first manifestation of possible multiple sclerosis (clinically isolated syndrome) (109). Another example is the glutamine and glutamate complex CH₂ multiplets at 2.0–2.5 ppm, which overlap broadly at 1.5 T, whereas at 3.0 T (and with spectral-editing techniques) the individual contributions of glutamate and glutamine can be quantified.

View larger version (66K): [in this window] [in a new window] [Download PPT slide]   Figure 17a: (a) Volume of interest localizations for 3.0-T single-voxel 1H MR spectroscopy (2000/31; number of signals acquired, 128) on transverse (left) and coronal fluid-attenuated inversion recovery (top right) images and N-acetylaspartate (NAA) metabolite map (bottom right) in a patient with adrenoleukodystrophy. (b) Single-voxel 1H-MR spectrum from the displayed 6.5 mL volume in right occipital white matter. There is increased myo-inositol (mIns), choline (Cho), and lactate (Lac) and strong reduction of N-acetylaspartate. tCr = total creatine.

View larger version (15K): [in this window] [in a new window] [Download PPT slide]   Figure 17b: (a) Volume of interest localizations for 3.0-T single-voxel 1H MR spectroscopy (2000/31; number of signals acquired, 128) on transverse (left) and coronal fluid-attenuated inversion recovery (top right) images and N-acetylaspartate (NAA) metabolite map (bottom right) in a patient with adrenoleukodystrophy. (b) Single-voxel 1H-MR spectrum from the displayed 6.5 mL volume in right occipital white matter. There is increased myo-inositol (mIns), choline (Cho), and lactate (Lac) and strong reduction of N-acetylaspartate. tCr = total creatine.

Pediatric Imaging
On the basis of the existing literature, the importance of high-field-strength MR for pediatric imaging seems greatly unrecognized. In fact, however, one may argue that the average small infant requires high-field-strength MR much more than does the average adult patient. A higher magnetic field strength helps achieve an acceptable image quality even with very small field of views, which, at 1.5 T, are frequently associated with only borderline SNR. Fast imaging is also very helpful in young children who may not stay still throughout a long pulse sequence and in toddlers who do not respond to sedation (Figs 18–20). The same holds true of course for unresponsive or severely ill adults or those with dementia. A number of motion-correction algorithms have been proposed to correct artifacts in restless subjects, most notably parallel lines with enhanced reconstruction, or PROPELLER (112). These advanced reconstruction algorithms can, however, only correct in-plane motion—whereas gross patient motion is usually not restricted to the imaging plane but would rather occur in all three dimensions. Last, especially in newborns and premature infants, use of 3.0 T allows fast imaging (at high spatial resolution) without long and thus possibly harmful acquisition times, in particular if no MR-compatible incubator is available to maintain body temperature. A 3.0-T imaging examination in pediatric patients should also be technically easier than adult body imaging: Because of the small fields of view that are required for imaging small patients, SAR-related difficulties and dielectric artifacts are less problematic (Fig 20). One caveat exists, however, and this relates to acoustic noise: 3.0-T systems reach substantially higher noise levels than do 1.5-T systems; in very young patients, it will be of utmost importance to maintain careful ear protection and/or choose pulse sequence protocols according to the associated noise level. Although there are virtually no clinical data on the use of high-field-strength MR in pediatric patients; in our department, pediatric imaging has become one of the major clinical indications for high-field-strength MR.

View larger version (120K): [in this window] [in a new window] [Download PPT slide]   Figure 18: Fast imaging of a nonsedated 26-month-old male toddler with primitive neuroectodermal tumor. Coronal T2-weighted turbo SE 3.0-T MR image (SENSE for parallel imaging; reduction factor, four; spatial resolution, 0.7 x 0.7 x 2 mm; 24 sections; total acquisition time, 1 minute 58 seconds).

View larger version (69K): [in this window] [in a new window] [Download PPT slide]   Figure 19a: Pediatric MR at 3.0 T in a newborn baby, 36 hours after delivery, depicts hydrocephalus. (a) Sagittal T2-weighted high-spatial-resolution turbo SE image (2-mm section thickness, 356 x 523 imaging matrix, acquisition time of 1 minute 25 seconds) and (b) high-spatial-resolution inflow MR angiogram with SENSE.

View larger version (101K): [in this window] [in a new window] [Download PPT slide]   Figure 19b: Pediatric MR at 3.0 T in a newborn baby, 36 hours after delivery, depicts hydrocephalus. (a) Sagittal T2-weighted high-spatial-resolution turbo SE image (2-mm section thickness, 356 x 523 imaging matrix, acquisition time of 1 minute 25 seconds) and (b) high-spatial-resolution inflow MR angiogram with SENSE.

View larger version (151K): [in this window] [in a new window] [Download PPT slide]   Figure 20a: Pediatric MR in a 41/2-year-old male infant with liver fibrosis. (a) Liver imaging with a breath-hold technique at 1.5 T. Massive motion artifacts degrade image quality. (b) Repeat free-breathing respiratory-triggered T2-weighted turbo SE MR at 3.0 T (32 sections, two signals acquired, total acquisition time of 19 seconds).

View larger version (108K): [in this window] [in a new window] [Download PPT slide]   Figure 20b: Pediatric MR in a 41/2-year-old male infant with liver fibrosis. (a) Liver imaging with a breath-hold technique at 1.5 T. Massive motion artifacts degrade image quality. (b) Repeat free-breathing respiratory-triggered T2-weighted turbo SE MR at 3.0 T (32 sections, two signals acquired, total acquisition time of 19 seconds).

	CONCLUSION

TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

Although substantial technical progress has been made, even today the majority of published articles focus on technical aspects alone. In part, this reflects the fast technical development and the short half-life of what is to be considered cutting-edge technology. On the other hand, this type of technology-driven science is also due to the fact that it is easier to collect data on image quality, for example, of a new pulse sequence (or field strength), than is the hassle of performing a prospective intraindividual clinical study. Accordingly, as of the time this article was written, prospective clinical studies that would systematically evaluate the diagnostic accuracy of high-field-strength MR and/or its effect on clinical patient care are lacking almost entirely. Only few studies deal with the clinical use of 3.0-T MR, and of those, only a very small number offer an intraindividual comparison with 1.5 T (Table). An added value for using 3.0 T compared with 1.5 T—even regarding "surrogate criteria" such as image quality—has been established for only a minority of these clinical applications. However, as the saying goes, "absence of evidence is not evidence of absence." Indeed, the published evidence may not reflect the current clinical use of high-field-strength MR, or, in other words: Clinical use seems to be far ahead of scientific proof. Good or bad—it appears that the new technology is embraced and used clinically. This is certainly not the first time for this to happen in the history of diagnostic imaging—there is only a very limited number of studies comparing 1.5-T imaging with MR at a lower field strength, after all. It seems natural to use an improved tool once it is available. However, in these authors' opinion, there is a clear and urgent need for scientific data to establish the advantage of high-field-strength MR in clinical patient care.

	ESSENTIALS

TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

• Functional cardiac imaging (myocardial perfusion and tagging) may be superior at 3.0 T; coronary angiography, viability imaging (late enhancement), and cine MR are currently equivalent at 1.5 T.
  
• Diagnostic accuracy of breast imaging may be increased at 3.0 T, pending the resolution of current B₁ inhomogeneities.
  
• Abdominal MR at 3.0 T is ready for clinical use but is still difficult in patients with intraabdominal fluid collections.
  
• Pelvic and musculoskeletal high-field-strength MR may benefit from higher spatial, temporal, and spectral resolution.
  
• Pediatric patients are the ones who will probably profit most from higher magnetic fields.
  

	ACKNOWLEDGMENTS

 
The authors acknowledge the support by Torsten Sommer, MD, Mike Wattjes, MD, Dariusch Hadizadeh, MD, Carsten Meyer, MD, Katharina Strach, MD, and Renate Bloemer, RT.

	FOOTNOTES

 

Abbreviations: CNR = contrast-to-noise ratio • SAR = specific absorption rate • SE = spin echo • SENSE = sensitivity encoding • SNR = signal-to-noise ratio • SPIO = superparamagnetic iron oxide

Authors stated no financial relationship to disclose.

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TOP ABSTRACT INTRODUCTION BODY APPLICATIONS OF HIGH-FIELD... CONCLUSION ESSENTIALS References

 

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