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Body and Cardiovascular MR Imaging at 3.0 T¹

¹ From the Department of Radiology, New York University Medical Center, 530 First Ave, New York, NY 10016. Received March 31, 2006; revision requested May 31; revision received July 19; accepted August 23; final version accepted November 11; final review and update by V.S.L April 6, 2007. Address correspondence to V.S.L. (e-mail: vivian.lee{at}med.nyu.edu).

	ABSTRACT

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
Potential advantages of magnetic resonance (MR) imaging at 3 T include higher signal-to-noise ratios, better image contrast, particularly in gadolinium-enhanced applications, and better spectral separation for spectroscopic applications. In terms of clinical imaging, these advantages can mean higher-spatial-resolution images, faster imaging, and improved MR spectroscopy. However, achieving superior imaging and spectroscopic quality at 3 T can be challenging. This review discusses many of the problems encountered in body and cardiovascular MR imaging at 3 T, such as increased susceptibility, B₁ field inhomogeneity, and increased specific absorption rate. The article also considers solutions that are being pursued, such as parallel imaging, variable-rate selective excitation, and variable flip angle sequences. A review of the most commonly used pulse sequences provides practical tips on how these can be optimized for 3-T imaging. In the coming few years, substantial improvements in 3-T technology for clinical imaging and spectroscopy will undoubtedly be seen. An understanding of the basic principles on which these developments are based will help radiologists translate the advances into better imaging studies and, ultimately, better patient care.

© RSNA, 2007

	INTRODUCTION

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
Motivated by a combination of higher signal-to-noise ratio (SNR) potential, better spectral separation, and some improvements in image contrast, the proliferation of 3-T whole-body magnetic resonance (MR) systems for clinical imaging has introduced new challenges for the user (1,2). To realize the potential for improved imaging at higher field strengths, users of the systems must become familiar with some of the important differences between 1.5 T and 3 T. Some of these differences lead to obvious modifications in 1.5-T protocols for 3-T applications, such as the adjustment of echo times (TEs) for in-phase and opposed-phase T1-weighted gradient-echo imaging. Others lead to pursuit of specific strategies at 3 T, such as parallel imaging to take advantage of higher SNR and variable flip angle pulse sequences to reduce specific absorption rate (SAR). The most challenging problems, such as radiofrequency (RF) or B₁ heterogeneity, give rise to the need for new engineering and technical developments that are rapidly being pursued in research and commercial laboratories. While most of the early work at high field strength was concentrated on brain imaging, particularly functional MR imaging, the recent availability of 3-T whole-body systems with phased-array coils has generated considerable interest in body and cardiovascular MR imaging techniques.

This article starts with a basic review of some of the general principles of MR physics at 3 T, including a comment about safety considerations. It then discusses the practical implications of these principles on commonly used body and cardiovascular MR imaging sequences. Finally, future directions of high-field-strength body MR imaging are briefly considered.

	SAFETY OVERVIEW

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
Short-term exposures to static magnetic fields have not been shown to result in harmful biologic effects in humans. The U.S. Food and Drug Administration has categorized clinical MR systems with a static magnetic field of up to 8 T as posing "nonsignificant risk" for patients older than 1 month (current Food and Drug Administration guidelines can be found at http://www.fda.gov). The safety of medical devices and implants is a topic that is proliferating in the literature. In general, most devices that are safe at 1.5 T are proving also to be safe at 3 T (3). In particular, recently approved drug-eluting coronary stents have now been approved for immediate MR imaging following insertion, at both 1.5 T and 3 T (4). The reader is referred to several continuously updated internet Web sites for reference in specific cases (eg, http://www.radiology.upmc.edu/MRsafety and http://www.mrisafety.com).

One important safety consideration at higher field strengths that affects pulse sequence design and the realization of improved SNR at 3 T is the SAR. This rate is estimated based on RF energy deposition and must be carefully monitored to avoid excess heating of subjects in the magnet bore. SAR is proportional to the square of the main magnetic field. Consequently, at 3 T, RF energy deposition is four times greater than at 1.5 T. The U.S. Food and Drug Administration limits SAR on the basis of concerns regarding the potential for tissue damage caused by excessive heat exposure. The specific limits are no more than 4 W/kg as averaged over the entire body for a duration greater than or equal to 15 minutes and no more than 8 W/kg in any gram of tissue in the head or torso for a duration of greater than or equal to 5 minutes (5). These limits are guidelines, and each manufacturer of high-field-strength magnets sets its own specific limits, which are typically more conservative. As will be considered below, the higher SAR at 3 T constrains many types of RF-intensive pulse sequences, necessitating new strategies for pulse sequence design and implementation.

	GENERAL PRINCIPLES OF MR PHYSICS AT 3 T

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
B₀ and Susceptibility
B₀ defines the main static field of the MR system. A well-shimmed 3-T whole-body MR system can provide field homogeneity specifications that are comparable with or even improved over those of 1.5-T systems (6). One vendor reports a 0.25 ppm root mean square voltage error, or V_RMS, at 40-cm diameter of sensitive volume, or DSV, and 1 ppm V_RMS at 50-cm DSV at 3 T (Siemens Medical Solutions, Erlangen, Germany) (7). Other vendors offer 0.5 ppm V_RMS or less at 40-cm DSV at 3 T (GE Healthcare, Waukesha, Wis [8] and Philips Medical Systems, Best, the Netherlands [9] [Romhild Hoogeveen, personal communication]).

Once a subject is placed into the bore of the magnet, differences in magnetic susceptibility of tissues lead to local alterations of the magnetic field. Susceptibility effects are larger at 3 T than 1.5 T because of a linear dependence of the magnetization on the field strength. There are several consequences of this increased susceptibility. The shortened T2* relaxation times that result from local field distortions cause greater signal loss at a given TE at 3 T than at 1.5 T. Given that spatial localization relies on gradient-induced frequency and phase encoding, susceptibility-induced field alterations also lead to image distortions. For some applications such as functional MR imaging of the brain, where the blood oxygenation level–dependent, or BOLD, effect relies on the susceptibility effects of deoxyhemoglobin, this greater sensitivity at 3 T is highly desirable. Susceptibility-weighted imaging of the brain, and even some abdominal imaging applications, may benefit from greater susceptibility. For example, improved sensitivity to iron deposition in the liver and other organs may be useful in the detection of diseases such as hemochromatosis. The signal-to-noise and contrast-to-noise effects of iron-based contrast agents depend on the sequences and doses and are the subject of active investigation (10,11).

However, for most body and cardiovascular applications, greater susceptibility creates undesirable image distortions and signal loss—for example, in regions such as bowel-gas and cardiac-lung interfaces and around implanted metallic clips (12) or prostheses (Fig 1).

View larger version (71K): [in this window] [in a new window] [Download PPT slide]   Figure 1: Transverse T1-weighted gradient-echo MR images (repetition time, 167 msec; flip angle, 90°) illustrate greater susceptibility effects (arrow) associated with anterior abdominal wall surgical clips at (right) 3 T (TE, 1.58 msec) compared with at (left) 1.5 T (TE, 2.38 msec) despite the considerably shorter TE used in the same subject.

View larger version (85K): [in this window] [in a new window] [Download PPT slide]   Figure 2: Oblique coronal (short-axis) balanced steady-state free precession (SSFP) MR images (repetition time msec/TE msec, 2.8/1.1; 50° flip angle) obtained through the heart at 3 T (left) before and (right) after local shimming demonstrate the sensitivity of the images to inhomogeneities in the main magnetic field, which are improved once local shimming is performed.

In cardiovascular applications, SSFP sequences that rely on balanced gradients to maintain phase coherence of transverse magnetization, such as balanced fast field echo, true fast imaging with steady-state precession, or fast imaging excitation with steady-state acquisition, are particularly sensitive to magnetic field inhomogeneities. Frequency mismatches lead to banding artifacts and other types of signal loss (Fig 3). Matching the excitation frequency with the resonance frequency can be achieved by using a frequency scout sequence, with considerable reduction in artifacts (Fig 4). Unfortunately, frequency matching does not fully solve the problem. Artifacts relating to flow and motion may persist (Fig 4). A homogeneous field remains of utmost importance for those applications.

View larger version (126K): [in this window] [in a new window] [Download PPT slide]   Figure 3: Transverse MR images obtained through the abdomen demonstrate banding artifacts associated with SSFP imaging (3.0/1.5; 50° flip angle). A, By deliberately introducing a magnetic field gradient across the image plane, severe banding artifacts can be observed on images acquired with SSPF sequences. These artifacts can be greatly reduced by improving magnetic field homogeneity, as shown in B and C. D, With a perfectly shimmed imaging volume, banding artifacts can be entirely eliminated.

View larger version (79K): [in this window] [in a new window] [Download PPT slide]   Figure 4: Oblique transverse (horizontal long axis) views of the heart on electrocardiographically gated balanced SSFP MR images (64.8/1.4; 53° flip angle) obtained (left) before and (right) after resonance frequency matching improves artifacts associated with field inhomogeneity effects in phase-coherent sequences such as balanced SSFP. Frequency mismatching can be particularly problematic at the lung-heart interface at 3 T. Note that minimal artifacts relating to flow and motion (arrow) persist.

With the increase in precessional frequency, chemical shift between compounds, which is similarly proportional to the magnetic field strength, also increases. This greater spectral separation of interrogated compounds, combined with greater SNR, provides great promise for MR spectroscopy at high field strengths (Fig 5) (13–15).

View larger version (28K): [in this window] [in a new window] [Download PPT slide]   Figure 5: MR spectroscopy shows greater SNR and spectral separation at (right) 3 T compared with at (left) 1.5 T in the same lactating subject despite considerably shorter acquisition time at 3 T. Note that at 3 T, two distinct resonant lactose peaks (a doublet) are demonstrated. MR spectroscopy parameters: for 1.5 T, 1500/135, voxel size of 18 cm3, 128 signals acquired, and acquisition time of 3 minutes 12 seconds; for 3 T, 1500/135, voxel size of 10 cm3, 64 signals acquired, and acquisition time of 1 minute 36 seconds. Cho = choline.

Second, given the greater separation of fat and water at 3 T, a chemical shift artifact of the first kind, owing to mismapping of the frequency-encoded signal of fat into water voxels, can be seen at higher receiver bandwidths than at 1.5 T. A receiver bandwidth of 220 Hz/voxel will result in a chemical shift artifact that is 2 voxels wide at 3 T (Fig 6), compared with 1 voxel wide at 1.5 T. Use of higher receiver bandwidths can minimize chemical shift artifacts of the first kind; however, it also results in lowered SNR. Last, for applications that depend on frequency-selective fat suppression pulses, the greater spectral separation lends itself to more successful suppression, provided reasonable B₀ field homogeneity is maintained (Fig 7).

View larger version (111K): [in this window] [in a new window] [Download PPT slide]   Figure 6: Transverse T1-weighted spoiled gradient-echo MR images (217/4.9; 90° flip angle) obtained through the abdomen demonstrate that chemical shift artifact, resulting from a shift in spatial localization of fat, at 3 T is most prominent (arrows) at a bandwidth (BW) of 220 Hz/voxel (bottom right) and also clearly visible at 400 Hz/voxel (bottom left), while at higher bandwidths, artifact is negligible.

View larger version (106K): [in this window] [in a new window] [Download PPT slide]   Figure 7: Transverse frequency-selective fat-suppressed fast spin-echo (FSE) MR image obtained through the abdomen at 3 T illustrates excellent fat suppression, as well as loss of signal intensity (arrow) in the left lobe of the liver, attributable to dielectric effect.

RF eddy current effects also have implications for SAR calculations. With placement of a body in the coil, heterogeneous B₁ means that the distribution of the SAR tends to vary across the body and may be harder to predict or generalize across individuals. From a practical standpoint, inhomogeneous RF excitation because of B₁ inhomogeneity results in visible inhomogeneity across the image that can obscure visualization of certain parts of the body, such as the left lobe of the liver (Fig 7). Moreover, Edelman et al (16) have observed worsened signal nonuniformity, despite higher SNR, with single-shot FSE in the abdomen at 3.0 T compared with 1.5 T. They suggest that the lengthy train of refocusing RF pulses may make the single-shot sequence particularly sensitive to B₁ inhomogeneity.

Several different approaches to reducing this heterogeneity are being explored. The least technical involves the use of dielectric padding over the abdomen, which is intended to counter the inductive effects of tissues. Additionally, modified transmitter adjustment methods that increase reference transmitter voltages also are being used to improve image quality. More elegant and robust solutions include the development of new pulse sequences and RF excitation pulses to generate uniform excitation over the desired imaging volume. Additionally, new phased-array transmit coil designs are under investigation to permit customized excitation pulses for uniformity (17).

T1 and T2 Relaxation Times
The magnetic field dependence of tissue relaxation times is well documented in the brain and musculoskeletal system (18–21). Takahashi et al (22) determined the relaxation time of phantoms with different concentrations of gadolinium-based contrast material at both 1.5 T and 4 T and showed that T1 relaxation times were prolonged (1.10–1.47 times) at 4 T compared with those at 1.5 T, while T2 values were identical or slightly shortened. In human volunteers, de Bazelaire et al (23) measured relaxation times of abdominal organs at 1.5 T and 3 T and showed an overall increase in T1 relaxation times and a slight decrease in T2 relaxation times at 3 T compared with 1.5 T, depending on the organ. These findings were also confirmed by Stanisz et al (24).

Owing to the lengthening of T1 relaxation times for most abdominal organs, there is a potential reduction of T1 contrast at 3 T. Consequently, for fast gradient-echo imaging, the repetition time should be shorter or flip angle larger to optimize image contrast and SNR. T1 relaxation times for gadolinium-based contrast agents, however, are less affected at higher field strengths, and this results in greater image contrast on gadolinium-enhanced images at 3 T compared with those at 1.5 T (Fig 8).

View larger version (126K): [in this window] [in a new window] [Download PPT slide]   Figure 8: Oblique sagittal thin maximum intensity projection images from gadolinium-enhanced MR angiography of the thoracic aorta (3.5/1.3; 25° flip angle) at (left) 1.5 T and (right) 3 T demonstrate higher image contrast at higher field strength. Contrast-to-noise measurements taken from source image in region 1 were 71.9 and 103.5 at 1.5 T and 3 T, respectively. Similarly, for region 2, the corresponding values were 50.5 and 131.5, respectively.

Specific Absorption Rate
When imaging human subjects, RF power deposition is constrained by SAR limitations. Certain pulse sequences that require repeated excitation with extremely short repetition time, such as balanced SSFP sequences (true fast imaging with steady-state precession), and those that demand rapid high flip angle refocusing pulses, such as single-shot FSE and other RARE-type sequences, pose particular challenges in translation to higher field strengths.

Two of the simplest solutions are to reduce the flip angle or to increase RF pulse duration. Reducing flip angles will cause changes in image contrast and decreased SNR. For example, lower refocusing flip angles from 180° to, say, 140° for RARE imaging results in decreased image contrast owing to an introduction of T1 dependence of the signal (26). If these changes in image contrast and the loss of SNR are not tolerable, the desired flip angle can be maintained by lengthening the pulse duration in order to lower the power deposition of a sequence. This inevitably causes a less efficient study and longer acquisition times. Similarly, other strategies to increase the "down time" between RF pulses, such as increased echo train spacing or lengthened repetition time, can reduce SAR, but these changes also increase acquisition times. Longer acquisition times can be particularly undesirable in chest and abdominal applications where acquisitions are performed during suspended respiration.

Several newer strategies are being pursued to reduce RF power deposition without diminishing image quality or lengthening acquisition times.

Variable-rate selective excitation (VERSE) methods, first described almost 2 decades ago, are seeing a resurgence at higher field strengths (27,28). VERSE techniques combine a time-varying gradient waveform with a modified RF waveform to provide the same excitation profile with less RF deposition (Fig 9). The time transformation of a frequency-selective excitation pulse maintains the amplitude integral of the pulse but enables the lowering of the high amplitude values and increasing of the low amplitude values. Consequently, pulse power is considerably reduced. The scaling of the time intervals is given by the slope of the slice-select gradient. Flip angle and section profile are maintained. The excitation becomes slightly susceptible to chemical shift–related problems, depending on the lower gradient values applied during excitation. Nevertheless, these pulses are an attractive approach to achieving high flip angles without exceeding SAR limits (Fig 10). VERSE transformations also apply to adiabatic excitation pulses, which can achieve B₁-insensitive selective or nonselective inversion of magnetization. Adiabatic pulses are generally energy demanding. The VERSE transformation keeps their energy contribution low, but retains their valuable off-resonance features.

View larger version (12K): [in this window] [in a new window] [Download PPT slide]   Figure 9: Example of (left) an unfiltered frequency-selective sinc pulse and (right) its variable-rate transformed version. Dashed lines represent the gradient, which is modulated during the RF pulse in the VERSE implementation. The amplitude values of the central main pulse lobe are much reduced and cause the considerable power reduction.

View larger version (146K): [in this window] [in a new window] [Download PPT slide]   Figure 10: Sagittal T2-weighted FSE MR images obtained (left) without VERSE (4000/115; 150° flip angle; echo train length, 25) and (right) with VERSE (4000/116; 180° flip angle; echo train length, 25). The VERSE-transformed RF pulse generates comparable image quality with a 44% power reduction. This case illustrates the benefit of increased refocusing flip angle without exceeding SAR limits by use of VERSE.

A complementary approach that can be used with FSE-type sequences, particularly those using variable flip angle methods, is to transfer the transverse magnetization back to the longitudinal direction prior to the next RF excitation pulse. This method is particularly useful at 3 T, where T1 relaxation times are observed to be prolonged. For example, fluids with long relaxation times appear less saturated on T2-weighted FSE images by using this technique. Restoring the magnetization is also helpful for heavily T1-weighted imaging with short echo trains in order to drive the magnetization to equilibrium within the selected repetition time.

Compatible with these new strategies is the use of parallel imaging methods. By reducing the necessary number of phase-encoding steps, parallel imaging techniques can help decrease RF exposure without compromising image quality (see below).

Signal-to-Noise Ratio
One of the major advantages of high-field-strength imaging is the potential increase in SNR, which increases linearly with field strength because of increased magnetization of nuclear hydrogen spins and higher resonance frequency. Increasing field strength from 1.5 T to 3 T leads to a theoretical twofold increase in signal intensity while noise remains relatively unaffected by field strength. This theory is verified by findings of phantom studies at 3 T and 4 T (33,34) and is supported by results of human abdominal imaging studies at 3 T (16,35) and 4 T (36). In practice, studies comparing SNR at 1.5 T and 3 T are difficult to design, since comparison of imaging protocols by using identical protocols may not be realistic given differences in T1 and T2 relaxation times and SAR constraints, as well as differences in performance characteristics of phased-array coils (16).

In human subjects, the theoretical increase in SNR may not be completely realized because of several factors. At high field strength, T1 relaxation times, chemical shift, and susceptibility effects increase, while T2 relaxation times decrease. To optimize tissue contrast, higher bandwidths are often required, thereby decreasing the overall SNR. SAR limitations may reduce imaging efficiency or image quality. For example, use of reduced flip angles for RF-intensive sequences such as balanced SSFP sequences or single-shot FSE sequences decreases image contrast and SNR. Again, parallel imaging strategies can potentially compensate for the limitations of SAR and maximize the potential gains in SNR at higher field strengths.

Parallel Imaging at 3 T
Parallel imaging techniques such as simultaneous acquisition of spatial harmonics (known as SMASH) (37) and sensitivity encoding (or SENSE) (38) use the spatial signal variation across elements of a coil array to provide spatial encoding information, thereby reducing the number of phase-encoding steps needed to generate an image. The improved efficiency can be used to reduce acquisition time or to increase spatial resolution or both.

To perform parallel imaging, signal must be recorded from individual receiver coil elements, necessitating a separate receiver channel for each element. The factor by which parallel imaging can reduce the number of phase-encoding steps, sometimes referred to as the R factor, generally can be no more than the number of independent coil elements in the phase-encoding direction. For three-dimensional applications, parallel imaging can theoretically be implemented in both phase-encoding directions, allowing for an even more substantial reduction in acquisition time.

With parallel imaging methods, SNR is reduced by a factor called g, the geometric factor, which reflects the lack of total spatial independence of coil profiles. The geometric relationships between individual coils result in g values greater than 1. SNR in parallel imaging, SNR_par, is reduced by a factor that is proportional to g and also to the square root of the R factor: SNR_par = SNR/(gR).

Since g varies spatially within the image, routine SNR measurements with parallel imaging techniques can be challenging. Reeder et al (39) have proposed two approaches to measuring SNR and demonstrated their comparability in a study using a 32-channel coil for cardiac imaging at 1.5 T. The multiple-acquisition method requires repeated acquisitions with the same imaging parameters. For a given pixel, the mean signal intensity across acquisitions is divided by the standard deviation across acquisitions to derive pixel-by-pixel SNR measurements. Alternatively, the difference method proposed by Reeder et al uses only acquisition of two identical images. SNR for a given region of interest is derived from the ratio between the average of the signal intensities for the two acquisitions and their difference. Surrogate measurements for SNR that are also used include the ratio between the mean signal intensity and the standard deviation of signal intensity within a reasonably large region of interest that contains a homogeneous sample of single tissue type.

The use of parallel imaging at high field strength is particularly beneficial for two reasons: (a) The increased SNR at high field strength compensates for the loss of SNR usually associated with parallel imaging and (b) parallel MR imaging can help solve the spectrum of problems associated with high-field-strength imaging described above, such as increased susceptibility artifacts, increased SAR, T2* decay in echo-planar imaging, and T2 blurring in FSE imaging (40).

As illustrated in Figure 11, by reducing the number of phase-encoding steps needed to generate an image, parallel imaging can be used to decrease susceptibility artifacts in echo-planar imaging by shortening TE. For SAR-intensive sequences, such as single-shot FSE methods, parallel imaging shortens the required echo train length for imaging. Consequently, full 180° refocusing pulses can be maintained without exceeding SAR limits (Fig 12), thereby preserving image contrast. Even with the same TE, the reduction in echo train length with parallel imaging results in visibly reduced T2 blurring, as shown in Figure 13.

View larger version (42K): [in this window] [in a new window] [Download PPT slide]   Figure 11: Transverse echo-planar MR images obtained through the abdomen illustrate of the use of parallel imaging to reduce susceptibility artifacts. Echo-planar methods are used to obtain diffusion-weighted images in the abdomen. Left: Image acquired with standard echo-planar sequence (2000/51; 90° flip angle) with effective TE of 51 msec shows substantial distortion of the kidneys (arrows). Middle: Image acquired with same echo-planar sequence (2000/23; 90° flip angle), combined with threefold parallel acceleration, reduces effective TE to 23 msec and thus greatly reduces distortion artifacts. Right: Half-Fourier RARE image shown as a reference of the anatomy (800/96; 132° flip angle).

View larger version (64K): [in this window] [in a new window] [Download PPT slide]   Figure 12: Transverse half-Fourier RARE MR images obtained through the abdomen (800/96) illustrate use of parallel imaging to solve SAR problems. Left: To acquire 25 sections of transverse half-Fourier RARE images within a breath hold (20 seconds), the flip angle of the refocusing RF pulses had to be reduced to 132° to avoid exceeding SAR limitations. Right: By using parallel imaging with threefold acceleration, the number of phase encoding steps (and thus the number of refocusing pulses needed) was dramatically reduced, allowing use of regular 180° refocusing pulses. Shorter echo train length with parallel imaging reduces T2 blurring in liver (arrowheads) and improves corticomedullary differentiation in the kidney (arrow).

View larger version (67K): [in this window] [in a new window] [Download PPT slide]   Figure 13: Transverse abdominal MR images show reduction of T2 blurring by using parallel imaging. Both (left) standard turbo spin-echo image (2200/84; 180° flip angle) and (right) image obtained with parallel acquisition of twofold acceleration (2200/84; 180° flip angle) have the same effective TE of 84 msec. However, the reduction in echo train length from 21 to 13 by using parallel imaging results in slightly reduced T2 blurring (arrow).

	IMPLICATIONS FOR COMMONLY USED MR PULSE SEQUENCES

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
Having reviewed some of the relevant principles of MR physics at 3 T, this article turns now to the implications of these principles on several of the most commonly used body and cardiovascular MR imaging sequences: T1-weighted gradient echo, T2-weighted FSE, single-shot FSE, three-dimensional gradient echo, and balanced SSFP. A brief discussion of gadoliniumenhanced imaging then follows.

T1-weighted Gradient-Echo Imaging
For most routine abdominal and pelvic applications, T1-weighted gradient-echo imaging with in-phase and opposed-phase imaging is desirable. The addition of opposed-phase imaging is useful for the detection of hepatic steatosis, characterization of adrenal adenomas, and identification of fat-containing tumors such as angiomyolipomas and myelolipomas.

Ideally, in-phase and opposed-phase images are acquired by using a dual-echo gradient-echo sequence. By imaging both phases in one acquisition, the technique is not only efficient but also guarantees image registration for optimal comparison. The greater chemical shift between fat and water at 3 T means that TEs at which fat and water protons are in-phase and opposed-phase are more closely spaced at 3 T (1.14 msec). At 1.5 T and even more so at 3 T, high receiver bandwidths may be necessary to allow for sampling of two separate echoes following each RF excitation.

When configuring the in-phase and opposed-phase gradient echo sequence, it is preferable for the opposed-phase acquisition to be performed at a shorter TE than the in-phase so that signal loss on the opposed-phase image can be attributable to fat-water signal cancellation, rather than to T2* decay. Ensuring that opposed-phase imaging is occurring at shorter TE is particularly important to differentiate steatosis from iron deposition. Theoretically, potentially desirable combinations of TEs would include (a) 1.14 and 2.3 msec, (b) 1.14 and 4.6 msec, (c) 3.4 and 4.6 msec, and (d) 3.4 and 6.9 msec, for opposed-phase and in-phase acquisitions, respectively. Actual combinations of dual-echo gradient-echo TEs that are feasible with commercial systems tend to vary from these values depending on constraints on gradient strength and signal sampling times (Fig 14). Choosing between options depends on balancing the trade-offs between the improvements in SNR and lower susceptibility effects associated with shorter TEs against the higher receiver bandwidths and subsequent degradation in SNR necessary to achieve these TEs.

View larger version (131K): [in this window] [in a new window] [Download PPT slide]   Figure 14: Transverse abdominal T1 opposed-phase and in-phase MR images from dual-echo gradient-echo sequences (repetition time, 237 msec; flip angle, 90°) with different pairs of TEs (and receiver bandwidths): A, TEs of 1.6 msec (930 Hz/voxel) and 2.9 msec (975 Hz/voxel); B, 1.6 msec (930 Hz/voxel) and 4.9 msec (400 Hz/voxel); C, 3.7 msec (1220 Hz/voxel) and 4.9 msec (1220 Hz/voxel); and D, 3.7 msec (610 Hz/voxel) and 7.4 msec (610 Hz/voxel). Depending on system specifications, the trade-offs between higher receiver bandwidths (and consequently lower SNR) must be balanced against the greater susceptibility and signal loss from longer TE.

T2-weighted FSE
For routine abdominal imaging, T2-weighted FSE sequences are typically performed with suppression of fat signal to enhance lesion conspicuity, particularly in the detection of lesions in a steatotic liver. Options include short-tau inversion recovery imaging, with inversion time set to null fat, and frequency-selective fat-suppressed FSE imaging.

For inversion recovery methods, the inversion time at 3 T should be adjusted slightly, although the change in T1 relaxation time for fat at 3 T is modest. In our experience, frequency-selective FSE imaging tends to be more robust at 3 T given the greater spectral separation between fat and water (Fig 15). However, successful fat suppression requires high B₀ uniformity. To address SAR limitations at 3 T with FSE imaging, implementation of parallel imaging helps to reduce the necessary echo train length, thereby also reducing T2 blurring. Variable refocusing flip angle sequences and VERSE RF excitation, if implemented, also help reduce SAR and allow high quality images at 3 T (32).

View larger version (58K): [in this window] [in a new window] [Download PPT slide]   Figure 15: Transverse fat-suppressed T2-weighted MR images of the breast obtained at (left) short-tau inversion-recovery FSE imaging (6290/79; 120° refocusing; inversion time, 200 msec; acquisition time, 5 minutes 28 seconds for 33 sections) and (right) frequency-selective fat suppressed FSE imaging (6720/111; 150° refocusing; and acquisition time, 3 minutes 22 seconds for 31 sections) in a 25-year-old woman with inflammatory breast carcinoma. Substantial time savings can be realized by using frequency-selective fat suppression when magnetic field homogeneity can be achieved.

Three-dimensional Gradient-Echo Imaging
Volumetric imaging of the abdomen and pelvis typically uses a fat-suppressed interpolated three-dimensional gradient-echo sequence that is performed both prior to and following administration of intravenous gadolinium-based contrast material. A similar version of the sequence performed without fat suppression is used for gadolinium-enhanced MR angiographic applications.

Parallel imaging can be implemented in both phase-encoding directions for three-dimensional pulse sequences. The additional SNR at 3 T affords the potential for larger reductions in phase-encoding steps with parallel imaging "squared" (Fig 16). For example, a reduction by a factor of two in each of two directions results in an overall reduction by a factor of four. The ongoing development of MR systems with greater numbers of receiver channels and the design of new coils optimized for parallel imaging are likely to show further improvements in contrast material–enhanced imaging at higher field strengths.

View larger version (60K): [in this window] [in a new window] [Download PPT slide]   Figure 16: Oblique sagittal maximum intensity projection of a contrast-enhanced three-dimensional MR angiogram of the thoracic aorta obtained by using parallel imaging in both phase-encoding directions for a total parallel factor of six (3.8/1.2; 21° flip angle; imaging matrix, 438 x 640; field of view, 32 x 40 cm; interpolated section thickness, 1.75 mm; number of interpolated partitions, 56; and acquisition time, 10.2 seconds). Reduced acquisition time comes at the expense of decreased SNR.

Susceptibility and flow-related artifacts as shown in Figures 2–4 can be reduced with focused shimming on the region of interest, and frequency adjustments may be necessary to minimize the artifact in the region of interest. Despite these adjustments, artifacts relating to flow and motion may persist (Fig 4). When free of artifacts, SSFP sequences provide improved signal intensity at 3 T compared with 1.5 T and comparable measures of left ventricular function (43–45). Because of SAR limitations, SSFP sequences are frequently performed at 3 T by using longer repetition times and lower flip angles than at 1.5 T, which in turn can result in worsened off-resonance artifacts or diminished SNR and image contrast. In view of these limitations, spoiled-gradient-echo techniques such as fast spoiled gradient-recalled imaging or fast low-angle shot imaging, which benefit from the increased SNR at 3 T, may be preferable alternative methods for assessing cardiovascular function (46).

Gadolinium-enhanced T1-weighted Gradient-Echo Imaging
In this section, special consideration is given to conventional contrast-enhanced imaging in body and cardiovascular MR at 3 T and to the potential for improved contrast-to-noise ratios at higher field strengths. Relaxivities, which are inversely related to relaxation times, have been measured for the most commonly used gadolinium chelates at 1.5 T and 3 T. Rohrer et al (47) reported relaxivities for gadopentetate dimeglumine (Magnevist; Berlex Laboratories, Wayne, NJ) as follows: r1 = 4.1 L·mmol^–1·sec^–1 and r2 = 4.6 L·mmol^–1·sec^–1 at 1.5 T compared with 3.7 L·mmol^–1·sec^–1 and 5.2 L·mmol^–1·sec^–1 at 3 T, respectively. The lower r1/r2 ratio at 3 T suggests that at high concentrations of contrast material, the increased signal intensity at T1-weighted imaging may be offset by the loss of signal intensity due to T2* effects. Consequently, for first-pass applications such as contrast-enhanced MR angiography at 3 T, greater signal intensity and image contrast may be achieved by aiming for lower concentrations of gadolinium chelate. Leiner et al (48) compared contrast-enhanced peripheral MR angiograms in two healthy volunteers at 1.5 T and 3 T and found increased signal intensity and contrast in the lower legs at 3 T. In their protocol, they empirically lowered the injection rate for this study from their standard maximum rate of 1.8 mL/sec to a lower rate of 1.2 mL/min in order to avoid the T2 losses associated with higher concentrations of gadolinium-based contrast material at 3 T.

The SNR benefit of 3-T imaging is particularly great on contrast-enhanced images, both in the abdomen (16) and brain (49), leading investigators to propose that half doses of contrast material may be sufficient for some contrast-enhanced imaging applications (49). For contrast-enhanced perfusion imaging of tissues such as the breast and heart, early results are also showing improved image SNR and contrast-to-noise ratio at 3 T by using similar imaging protocols (50,51). Additional work determining optimum sequence parameters and doses of gadolinium-based contrast material for different applications at 3 T is needed (52). Whether gains in SNR and contrast-to-noise ratio translate into measurable improvements in diagnosis remains to be determined.

	FUTURE DIRECTIONS FOR 3-T BODY AND CARDIOVASCULAR IMAGING

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
Continued advances in pulse sequence design developments at 3 T will be driven by the desire to realize SNR gains in the setting of SAR constraints and imaging artifacts. The higher SNR at higher field strengths is particularly attractive to investigators seeking to push the limits of parallel imaging. At both 1.5 T and 3 T, vendors are developing phased-array coils and systems with multiple independent channels with ever-increasing complexity. Recent published reports describe implementation of new 32-channel phased-array coils for brain (53) and body or cardiac imaging (54–56) at 3 T, with the latter offering the potential for imaging of whole-heart coronary arteries in a single breath hold; research systems with 128 receiver channels at 3 T are becoming available. One potential use of high-channel-array coils is performance of single-echo acquisition MR imaging, in which an independent full image is acquired with every echo by eliminating phase encoding and using spatial information obtained from the array of coils (41). Furthermore, new ideas about implementing parallel concepts with transmitting arrays are garnering substantial attention in academic and industry engineering laboratories (17,42,57).

Clinical MR spectroscopy may finally see greater widespread clinical utilization at 3 T. The improved SNR and spectral separation at 3 T facilitate new explorations in thoracic and abdominal applications (13) and may obviate endorectal coils in prostate MR imaging (58) and spectroscopy. The potential for MR spectroscopy to improve the diagnostic accuracy of breast MR for malignancy has been supported by findings of an early study at 4 T (59). In the transition to 3 T, optimization of pulse sequence timing needs to take into consideration the changes in T1 and T2 relaxation times of relevant metabolites at 3 T (60). Imaging at higher field strengths may also help to advance the use of MR with other nuclei, such as sodium 23, in the clinical setting (61).

Patient comfort will also drive new developments in high-field-strength MR imaging. The 3-T systems will evolve to become as short and workflow-optimized as systems at lower field strengths, without compromises in field homogeneity. Overall system costs are likely to decline.

The future will see the wider introduction of higher field strength into routine clinical practice. The recent proliferation of clinical 3-T whole-body MR systems is being echoed by a parallel spread of research 7-T whole-body systems. Whether the even more substantial challenges of magnetic field inhomogeneity and SAR limitations at 7 T can be overcome to realize the greater SNR and spectroscopic benefits at this higher field strength will be the subject of intense research during the next decade.

	CONCLUSION

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 
The superb quality of today's state-of-the-art 1.5-T imaging represents the product of over 20 years of research and development. By comparison, higher-field-strength whole-body MR systems are still in their infancy. As discussed in this article, 3-T MR imaging cannot be performed by simple adaptation from 1.5-T imaging. Key challenges at high field strength are increased SAR, increased B₀ and B₁ heterogeneity, and changes in T1 and T2* relaxation times, and all of these issues are providing the impetus for substantial research and development in MR hardware and software. At present, for pulse sequences and applications that are relatively free of artifacts, the greater SNR and better spectral separation at 3 T are starting to be realized in the clinical setting. The effects of these improvements on diagnostic accuracy need to be quantified.

In the long term, the intense research focus on MR imaging at high field strength will yield many creative solutions that will benefit MR imaging at all field strengths. Without a doubt, we will see in the next decade tremendous strides in clinical MR imaging and spectroscopy that will benefit patients in ways that we can only begin to imagine.

	ESSENTIALS

TOP ABSTRACT INTRODUCTION SAFETY OVERVIEW GENERAL PRINCIPLES OF MR... IMPLICATIONS FOR COMMONLY USED... FUTURE DIRECTIONS FOR 3-T... CONCLUSION ESSENTIALS References

 

• MR imaging at 3 T cannot be performed by simple adoption of 1.5-T pulse sequence parameters.
  
• Key challenges at high field strength are increased specific absorption rate (SAR), increased B₀ and B₁ heterogeneity, and changes in T1 and T2* relaxation times, and all of these issues are providing the impetus for substantial research and development in MR hardware and software.
  
• SAR limits favor the use of modified pulse sequence designs, including variable-rate selective excitation, hyperechoes, restore pulses, and variable flip angle signal evolution.
  
• Parallel imaging takes advantage of the higher SNR at 3 T to provide improved spatial or temporal resolution or both.
  
• The increased SNR and improved spectral separation favor the use of MR spectroscopy at 3 T.
  

	ACKNOWLEDGMENTS

 
We thank KellyAnne McGorty, RT (New York University, NY), for her assistance in collecting images and protocol parameters, Linda Moy, MD (New York University), and Nouha Salibi, PhD (Siemens Medical Solutions, Erlangen, Germany), for providing the breast MR spectroscopy images, and Bernd Stoeckel, PhD (Siemens Medical Solutions), for his insightful discussions.

	FOOTNOTES

 

Abbreviations: FSE = fast spin echo • RARE = rapid acquisition relaxation enhancement • RF = radiofrequency • SAR = specific absorption rate • SNR = signal-to-noise ratio • SSFP = steady-state free precession • TE = echo time • VERSE = variable-rate selective excitation

² Current address: Siemens Medical Solutions, Erlangen, Germany.

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