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DOI:10.2214/AJR.05.0932
AJR 2006; 186:1524-1532
© American Roentgen Ray Society

Perspective

Abdominal MRI at 3.0 T: The Basics Revisited

Elmar M. Merkle1 and Brian M. Dale2

1 Department of Radiology, Duke University Medical Center, Duke North, Rm. 1417, Box 3808, Erwin Rd., Durham, NC 27710.
2 Siemens Medical Solutions, USA, Cary, NC 27519.

Received June 1, 2005; accepted after revision July 20, 2005.

CME

This article is available for 1 CME credit. See supplemental data for this article at www.ajronline.org or visit www.arrs.org for more information.

Address correspondence to E. M. Merkle (elmar.merkle{at}duke.edu).

Abstract

OBJECTIVE. The purpose of our article is to describe the underlying physics concepts of abdominal MRI at 3.0 T and their impact on signal-to-noise ratio, susceptibility artifacts, chemical shift artifacts, and dielectric effects.

CONCLUSION. Abdominal MR sequence protocols optimized for 1.5-T scanners should not be transferred to 3.0 T without substantial modification. In addition, specific patient groups—for example, large patients with ascites—are not well suited to undergo an abdominal MRI study at 3.0 T.

Keywords: abdominal imaging • field strength • MRI • MRI technique • physics

Over the past 2 years, ultrahigh-field-strength whole-body 3.0-T MR systems have been installed in numerous institutions and are being increasingly used in clinics. Besides market considerations—for example, strategic investment to make a market statement and to stay competitive—the main reason to purchase an ultrahigh-field MR system is the anticipated twofold MR signal-to-noise ratio (SNR) compared with a standard 1.5-T MR scanner. This gain in SNR can be kept or traded for either speed, spatial resolution, or both. Although the number of accessory receiving coils has been limited in the past, the spectrum of dedicated receiver coils offered by vendors has increased significantly over the past 18 months, which allows almost all standard MRI examinations to be performed on a 3.0-T whole-body MR system. Although ultrahigh-field MR systems have already been shown to be advantageous for various indications in the brain and musculoskeletal system compared with standard high-field 1.5-T MR systems, only a few scientific studies have been published describing the use of 3.0-T MR systems in the chest, abdomen, and pelvis [1-5]. Unfortunately, insights gained in musculoskeletal or neuroimaging research at 3.0 T cannot simply be transferred to body MRI because MR sequence protocols and object sizes differ significantly in abdominal imaging. In addition, some artifacts are unique to ultrahigh-field abdominal MRI and are not seen in other regions of the body. It is also not clear which patient groups will benefit from an ultrahigh-field abdominal MRI study and which patient groups should remain on a 1.5-T MR scanner. This article will illustrate the underlying physics concepts of abdominal MRI at 3.0 T and their impact on SNR, susceptibility artifacts, chemical shift artifacts, and dielectric effects. On the basis of these fundamental considerations, basic recommendations will be provided for which patient groups will likely benefit from an ultrahigh-field MRI study and which patient groups should undergo a standard 1.5-T abdominal MRI examination.

3.0 T Offers Twice the SNR: A Persistent Myth

The idea that twice the magnetic field will give twice the SNR is appealing, and at first it seems correct because the intrinsic SNR in MRI is approximately proportional to the main magnetic field strength, B0 (equations 1 and 2) [6]. The equation for spin-echo-based MRI sequences is

Formula(1)

where SNRSE = signal-to-noise ratio for a spinecho pulse sequence, B0 = main magnetic field strength, V = voxel volume, NPE = number of acquired phase encode lines, NPA = number of acquired partitions, NAV = number of signals averaged, BW = receiver bandwidth per pixel, T1 = longitudinal relaxation time, and T2 = transverse relaxation time.

The equation for gradient-echo-based MRI sequences is

Formula(2)
*

where SNRGRE = signal-to-noise ratio for a spoiled gradient-echo sequence and {theta} =flip angle.

Note that, in both equations 1 and 2, the term under the square root is simply the total time spent acquiring data. Therefore, SNR is proportional to the main magnetic field strength, the voxel volume, the square root of the total sampling time, and some sequence-specific contrast-related terms. Some of these factors, such as the longitudinal relaxation time (T1), receiver bandwidth, and specific absorption rate limitations, can affect the SNR in a somewhat complicated manner by impacting other sequence-specific parameters (e.g., TR, flip angle).

The longitudinal relaxation time, T1, increases at a higher magnetic field strength, which causes a decrease in SNR [1, 7] (see equations 1 and 2) (Fig. 1A). The transverse relaxation time, T2, on the other hand, seems to be fairly independent of the main magnetic field strength [7]. However, one recently published study by de Bazelaire et al. [1] suggests a marked decrease of the transverse relaxation time (T2) at higher magnetic field strengths, which would further reduce the gain in SNR at ultrahigh-field-strength MRI (Fig. 1B) for long TE protocols. Given the optimistic assumption that the transverse relaxation time (T2) is independent of the main magnetic field strength and assuming only an increase of the longitudinal relaxation time (T1), equations 1 and 2 can be used to determine the theoretic maximum relative gain in SNR during MRI of the liver. For turbo spin-echo-based T2-weighted sequences with sequential acquisition such as HASTE sequences, an increase by a factor of approximately 1.8 in SNR can be obtained. For gradient-echo-based T1-weighted sequences such as in- and opposed-phase and VIBE (volume interpolated breath-hold examination) sequences, an increase by a factor of approximately 1.6-1.7 in SNR can be obtained. Thus, the theoretic twofold increase in SNR at 3.0 T compared with 1.5 T will not generally be obtained without sequence parameter optimization.


Figure 1
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  		Fig. 1A —Diagrammatic illustrations of relationships of relaxation times of liver tissue and main magnetic field strength (B0). Note that data of de Bazelaire et al. [1] are within confidence interval obtained by Bottomley et al. [7] even though 3.0-T data point is outside scope of article by Bottomley et al. Graph shows relationship of longitudinal relaxation time, T1, of liver tissue and main magnetic field strength. Data are shown from meta-analysis based on more than 800 study samples performed by Bottomley et al. [7] ({blacksquare}), together with data based on six volunteers acquired by de Bazelaire et al. [1] ({blacktriangleup}). Note nonlinear increase of T1 of liver tissue with main magnetic field strength. Also shown are nonlinear regression (solid line) and 95% confidence interval for data (dashed lines) as described by Bottomley et al.

 

Figure 2
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  		Fig. 1B —Diagrammatic illustrations of relationships of relaxation times of liver tissue and main magnetic field strength (B0). Note that data of de Bazelaire et al. [1] are within confidence interval obtained by Bottomley et al. [7] even though 3.0-T data point is outside scope of article by Bottomley et al. Graph shows relationship of transverse relaxation time, T2, of liver tissue and main magnetic field strength. Data are shown from meta-analysis based on more than 250 study samples performed by Bottomley et al. [7] ({blacktriangleup}), together with data based on six volunteers acquired by de Bazelaire et al. [1] ({blacksquare}). Note that no obvious relationship is seen between T2 of liver tissue and main magnetic field strength. Also shown are mean (solid line) and 95% confidence interval (dashed lines) for data as described by Bottomley et al.

 
In addition to the absolute change in the T1 relaxation time as a function of magnetic field strength, there are also relative changes for which the T1 relaxation time for one tissue increases at a different rate from the T1 relaxation time of another tissue. For example, according to Bottomley et al. [7], at 1.5 T the T1 relaxation time of the kidney is 32% greater than the T1 relaxation time of the liver (652 msec for kidney, 493 msec for liver), but at 3.0 T that difference shrinks to 21% (774 msec for kidney, 641 msec for liver). For other tissue pairs, the relative T1 dispersion may actually increase at ultrahigh field strength, rather than decreasing as shown here for kidney and liver. In any case, this example should illustrate why the contrast between various tissues on T1-weighted images at 3.0-T MRI cannot be identical to the contrast seen on standard 1.5-T T1-weighted MRI. Fortunately, on the basis of our current experience with more than 300 abdominal cases examined at ultrahigh-field MRI, this has not been a relevant clinical problem.

A second major factor with a negative impact on the gain in SNR is related to the specific absorption rate (SAR). When the main magnetic field strength is doubled, the SAR, a measure for energy deposition within the human body, increases by a factor of 4. Although the energy deposited at 3.0 T is still nonionizing, the increased SAR requires an increased concern for patient safety. Because body MRI at 3.0 T almost always runs at the upper limits of the allowed SAR deposition, patients are more likely to experience an uncomfortable sensation of warmth or heating. In addition, protocol adjustments are frequently necessary, such as an increase of the TR, a decrease in the number of slices, or a decrease of the flip angle. These adjustments are all undesirable because they increase scanning time, reduce anatomic coverage, alter contrast, or further reduce the gain in SNR at 3.0 T when compared with a standard 1.5-T MRI system. Finally, much of the radiofrequency transmitter and receiver technology at 1.5 T is relatively mature compared with the newer technology at 3.0 T.


Figure 3
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  		Fig. 2A —Chemical shift artifacts of the first and second kinds at various magnetic field strengths in same patient, 32-year-old woman. Axial gradient-echo in-phase image acquired at field strength of 1.5 T shows minimal chemical shift artifacts of the first kind along frequency-encoding axis.

 


Figure 4
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  		Fig. 2B —Chemical shift artifacts of the first and second kinds at various magnetic field strengths in same patient, 32-year-old woman. Axial gradient-echo in-phase image acquired at field strength of 3.0 T shows marked chemical shift artifacts of the first kind along frequency-encoding axis that appear bright toward higher part of readout gradient field (long arrow) and dark along lower part (short arrow).

 


Figure 5
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  		Fig. 2C —Chemical shift artifacts of the first and second kinds at various magnetic field strengths in same patient, 32-year-old woman. Axial gradient-echo opposed-phase image acquired at field strength of 1.5 T shows chemical shift artifact of the second kind in all pixels along fat-water interface.

 


Figure 6
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  		Fig. 2D —Chemical shift artifacts of the first and second kinds at various magnetic field strengths in same patient, 32-year-old woman. Axial gradient-echo opposed-phase image acquired at field strength of 3.0 T shows chemical shift artifact of the second kind, which is not significantly different from artifact seen at standard magnetic field strength of 1.5 T (compare withC).

 
All these reasons contribute to a gain in SNR that is less than the factor of 2.0 originally expected. This may help explain why Morakkabati-Spitz et al. [4] found that, "Visual signal to noise was rated equal for all 19 pelvic MRI studies" when applying a protocol with similar spatial and temporal resolution at both field strengths.

Chemical Shift Artifacts at 3.0 T: A Double-Edged Sword

The chemical shift artifact of the first kind is due to a difference in the resonant frequency between water and fat and is seen only along the frequency-encoding axis and the slice-selection dimension [8]. This difference in resonant frequency is directly proportional to the main magnetic field strength and has been measured as approximately 3.5 ppm, resulting in a difference of about 225 Hz at 1.5 T, or a difference of about 450 Hz at 3.0 T. This difference causes a chemical shift mis-registration, which is most easily seen around the kidneys (Figs. 2A, 2B, 2C and 2D). The chemical shift artifact of the first kind appears as a hypointense band, 1 to several pixels in width, toward the lower part of the readout gradient field, and as a hyperintense band toward the higher part of the readout gradient field. At a constant field of view, base resolution, and receiver bandwidth, the chemical shift artifact of the first kind will be twice as wide at 3.0 T as at standard 1.5-T imaging. Usually, this enlarged artifact does not cause substantial problems in clinical body MRI at 3.0 T. However, it may be problematic in selected cases such as the search for a subcapsular renal hematoma or an intramural aortic hematoma. In these cases, the receiver bandwidth can be increased to minimize the chemical shift artifact of the first kind. Unfortunately, this comes at the expense of SNR: doubling the receiver bandwidth will decrease the SNR by approximately 30% (see equations 1 and 2 Formula) (Figs. 3A, 3B and 3C). Another option is to repeat the MR pulse sequence with either chemical shift fat saturation, inversion nulling, or water excitation, which will eliminate chemical shift artifacts effectively, allow imaging at the lower bandwidth, and return the 30% loss in SNR.


Figure 7
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  		Fig. 3A —Effect of receiver bandwidth on size of chemical shift artifacts of the first kind at 3.0 T in 40-year-old man. Axial in-phase T1-weighted gradient-echo image through kidney shows significant chemical shift artifact (arrow) (receiver bandwidth, 210 Hz; pixel shift, 1.9; SNR [signal-to-noise ratio]LIVER, 75).

 

Figure 8
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  		Fig. 3B —Effect of receiver bandwidth on size of chemical shift artifacts of the first kind at 3.0 T in 40-year-old man. Axial in-phase T1-weighted gradient-echo image through kidney with twice receiver bandwidth as in A shows smaller chemical shift artifact (arrow) (receiver bandwidth, 415 Hz; pixel shift, 1.0; SNRLIVER, 56). However, SNR is also decreased by approximately 30%.

 

Figure 9
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  		Fig. 3C —Effect of receiver bandwidth on size of chemical shift artifacts of the first kind at 3.0 T in 40-year-old man. Axial in-phase T1-weighted gradient-echo image through kidney with four times receiver bandwidth as in A shows markedly smaller chemical shift artifact (arrow) (receiver bandwidth, 850 Hz; pixel shift, 0.5; SNRLIVER, 41). However, SNR is again substantially decreased by another 30%.

 
The chemical shift artifact of the second kind is not limited to the frequency-encoding axis but may be seen in all pixels along a fat-water interface because it is based on an intravoxel phase-cancellation effect in which fat and water exist in the same voxel [8]. The size of this artifact does not increase with the main magnetic field strength and is defined by the spatial resolution of the MR sequence (Figs. 2A, 2B, 2C and 2D). However, the TE needs to be adjusted because the frequency difference is twice as large as with the standard 1.5-T MR systems, as described in the section on chemical shift artifacts of the first kind. Using a 3.0-T MR system, both fat and water protons are in phase at 2.2, 4.4, 6.6 msec, and so on, and out of phase (also referred to as opposed phase) at 1.1, 3.3, 5.5 msec, and so on. Note that at 1.5 T, the fat and water are phase-opposed at 2.2 msec and in phase at 4.4 sec (nominal values). In short, by doubling the field strength we have halved the TEs for in-phase and opposed-phase imaging.

Fortunately, the increased difference in resonant frequency between water and fat at 3.0 T may also be advantageous because it allows a better separation of the fat and water peak during MR spectroscopy, and a better or faster fat suppression using other chemical shift techniques as well—for example, fat saturation and water excitation.

Susceptibility Artifacts: A Closer Look

Magnetic susceptibility is the extent to which a material becomes magnetized when placed in a magnetic field. Susceptibility artifacts occur as the result of microscopic gradients or variations in the magnetic field strength that occur near the interfaces of materials of different magnetic susceptibility. These artifacts are usually caused by metallic objects from previous surgical or interventional procedures near or in the imaging field of view because the susceptibility of metal is much higher than that of soft tissue. Susceptibility artifacts increase with the main magnetic field strength and are approximately twice as large in terms of volume at 3.0 T as at standard 1.5-T MRI [9] (Fig. 4). This may be advantageous in selected cases because metal-related susceptibility artifacts from surgical clips or surgical debris—for example, prior cholecystectomy or prior hepatic resection—may be better seen (Figs. 5A and 5B). However, it is possible that enlarged susceptibility artifacts may obscure important findings at 3.0-T MRI that may have been visualized at standard 1.5-T MRI. It must be clearly stated here that metal-containing devices that are considered MR safe at a field strength of 1.5 T are not necessarily safe at 3.0 T [10-15]. All these devices need to be rigorously tested at 3.0 T as well before affected patients can undergo an MRI examination at this field strength.


Figure 10
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  		Fig. 4 —Effect of magnetic field strength on size of susceptibility artifacts in various MR sequences in vitro.

A, Photograph shows phantom setup with water-filled straw (arrow) embedded in gelatin. Three pairs of surgical clips (arrowheads) are embedded at various distances from straw.

B, HASTE image acquired at 1.5 T (TR/TE, 1,010/128; field of view, 250 mm2; slice thickness, 2.8 mm; matrix, 256 x 205; bandwidth, 490 Hz) shows typical susceptibility artifacts caused by surgical clips.

C, HASTE image acquired at 3.0 T (1,010/128; field of view, 250 mm2; slice thickness, 2.8 mm; matrix, 256 x 205; bandwidth, 490 Hz) shows larger susceptibility artifacts caused by surgical clips when compared with B.

D, Gradient-echo image acquired at 1.5 T (118/2.4; field of view, 280 x 210 mm2; slice thickness, 3.0 mm; matrix, 256 x 154; bandwidth, 385 Hz) shows typical susceptibility artifacts caused by surgical clips.

E, Gradient-echo image acquired at 3.0 T (118/2.4; field of view, 280 x 210 mm2; slice thickness, 3.0 mm; matrix, 256 x 154; bandwidth, 385 Hz) shows larger susceptibility artifacts caused by surgical clips when compared with D. Note that artifact size increases by approximately 100% in terms of volume.

 

Figure 11
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  		Fig. 5A —Effect of magnetic field strength on size of metal-related susceptibility artifacts in vivo in 58-year-old man. Gradient-echo opposed-phase image acquired at 1.5 T (TR/TE, 200/2.2) shows typical susceptibility artifacts caused by surgical clips (arrows).

 

Figure 12
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  		Fig. 5B —Effect of magnetic field strength on size of metal-related susceptibility artifacts in vivo in 58-year-old man. Gradient-echo opposed-phase image acquired at 3.0 T (200/1.5) shows larger susceptibility artifacts (arrows) caused by surgical clips when compared with A despite shorter TE.

 


Figure 13
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  		Fig. 6A —Negative effect of magnetic field strength on size of gas-related susceptibility artifacts in vivo in 52-year-old man. Gradient-echo in-phase image acquired at 1.5 T (TR/TE, 200/4.4) shows minor susceptibility artifacts (arrows) in hepatic flexure and transverse colon caused by colonic gas.

 


Figure 14
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  		Fig. 6B —Negative effect of magnetic field strength on size of gas-related susceptibility artifacts in vivo in 52-year-old man. Gradient-echo in-phase image acquired at 3.0 T (200/4.4) shows larger susceptibility artifacts (arrows) in colon when compared with A. Note that these susceptibility artifacts obscure colonic wall.

 


Figure 15
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  		Fig. 7A —Positive effect of magnetic field strength on size of gas-related susceptibility artifacts in vivo in 65-year-old woman. Gradient-echo in-phase image acquired at 1.5 T (TR/TE, 200/4.4) shows minor susceptibility artifact (arrow) in left hepatic bile duct caused by pneumobilia that may be misinterpreted as a branch of portal venous system.

 


Figure 16
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  		Fig. 7B —Positive effect of magnetic field strength on size of gas-related susceptibility artifacts in vivo in 65-year-old woman. Gradient-echo in-phase image acquired at 3.0 T (200/4.4) shows markedly larger susceptibility artifact (arrow) in left hepatic bile duct caused by pneumobilia when compared with A. Enlarged susceptibility artifact makes misinterpretation less likely.

 


Figure 17
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  		Fig. 8 —Physical basis of radiofrequency shielding. In step A, rapidly varying magnetic field (black arrows) induces a circulating electric field (white arrows). In presence of a conductive medium (step B), circulating electric field leads to a circulating current (gray arrows). In step C, circulating current acts as electromagnet to produce magnetic field in opposite direction; and in step D, amplitude of overall magnetic field is reduced. Note that steps B and C require a conductive medium and that effect is stronger in more conductive medium. In imaging, this effect can be noticed in patients with ascites or fetal imaging, in which circulating currents can be established in relatively large regions of highly conductive fluid. Resulting artifacts are generally more visible in sequences that use a large number of radiofrequency pulses to generate contrast.

 


Figure 18
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  		Fig. 9A —Severe standing wave and conductivity artifact in 38-year-old woman with liver cirrhosis and ascites during ultrahigh-field-strength MRI at 3.0 T. Coronal HASTE image shows marked signal loss in center of image (arrows). Fluid accumulations in peritoneal cavity enlarge abdomen and increase electrical conductivity in field of view, causing severe artifacts.

 


Figure 19
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  		Fig. 9B —Severe standing wave and conductivity artifact in 38-year-old woman with liver cirrhosis and ascites during ultrahigh-field-strength MRI at 3.0 T. Axial HASTE image again shows marked signal loss in center of image (arrows) and represents severe standing wave and conductivity artifact.

 


Figure 20
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  		Fig. 9C —Severe standing wave and conductivity artifact in 38-year-old woman with liver cirrhosis and ascites during ultrahigh-field-strength MRI at 3.0 T. Contrast-enhanced axial gradient-echo T1-weighted image acquired at same level as B shows normal anatomy in center of field of view (arrows) and no evidence of susceptibility artifacts.

 
Susceptibility artifacts also occur next to gas-filled structures, such as the gas-filled bowel, because the susceptibility of gas is much less than that of soft tissue (Figs. 6A and 6B). Thus, bowel wall imaging in patients with inflammatory bowel disease or patients referred for MR colonography seems to be more challenging at 3.0 T. However, enlarged susceptibility artifacts due to a gas-soft-tissue interface may also be helpful in detecting gas—for example, intrahepatic pneumobilia or free intraperitoneal gas (Figs. 7A and 7B).


Figure 21
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  		Fig. 10 —Severe standing wave and conductivity artifact in 28-year-old pregnant woman during fetal ultrahigh-field-strength MRI at 3.0 T. Coronal HASTE image shows marked signal loss (long arrows). This artifact is caused by large amount of amniotic fluid and increased size of abdomen, which increase electrical conductivity in field of view. Short thin arrows mark placenta; thick arrows mark fetal torso. (Courtesy of Jones B, Cincinnati, OH)

 
Standing Wave Effects

In addition to the exacerbation of artifacts that are seen at 1.5 T, some new artifacts also begin to appear at 3.0 T. These artifacts are related to the higher frequency B1 transmit fields that are used at 3.0 T. The wavelength of the radiofrequency field at 128 MHz is 234 cm in free space, which is much larger than the field of view for clinical body imaging. However, water (and most body tissue) has a rather high dielectric constant, which reduces both the speed and wavelength of electromagnetic radiation. For visible light, this effect causes a straight stick entering water at an angle to appear bent. For MRI, this effect reduces the radiofrequency field wavelength from 234 cm in free space to about 30 cm in most human tissues—that is, water-containing tissues [16]. This size is approximately the size of the field of view for many body applications and can result in a so-called standing wave effect (often incorrectly called a "dielectric resonance" effect) [17]. As a result, strong signal variations across an image can be seen, especially brightening or dark "holes" in regions away from the receive coil caused by constructive or destructive interference from the standing waves. These artifacts become more pronounced the larger the region of interest is relative to the wavelength—that is, they are seen more in obese patients with a distended abdomen than in thin patients.

A rapidly changing magnetic field, like the radiofrequency transmit field, will induce a circulating electric field (Fig. 8). When this happens in a conductive medium, a circulating electric current is established. This current in turn acts like an electromagnet that opposes the changing magnetic field, reducing the amplitude and dissipating the energy of the radiofrequency field. The more conductive the medium, the stronger the opposing electromagnet and therefore the greater the attenuation of the radiofrequency field. In construction of the MRI suite, this principle is used by encasing the room in a copper conductor. Because of the high conductivity of copper, any incoming radiofrequency waves are almost completely attenuated and the magnet is shielded from external interference. To a lesser extent, large amounts of relatively highly conductive tissues can cause similar shielding effects, resulting in hypointense areas in the image where the radiofrequency field is partially attenuated [16].

These two effects combine to cause particularly strong artifacts for 3.0-T body MRI in pregnant patients and in patients with ascites (Figs. 9A, 9B, 9C and 10). In both cases, not only are the standing wave effects more pronounced because of the enlarged abdomen, but greater radiofrequency field attenuation is also present because of the increased amounts of highly conductive amniotic or ascitic fluid.

Summary and General Recommendations

Body MRI at 3.0 T is still in its infancy and will improve substantially over the next several years. However, radiologists need to know these several limitations based on the laws of physics:

First, overall, the gain in SNR at 3.0 T will be less than twofold compared with a standard 1.5-T MR system because of the inescapable increase of the longitudinal relaxation time T1. Also, the increased SAR deposition at ultrahigh field strength often requires protocol adjustments that can further reduce the anticipated gain. The gain in SNR will be higher in T2-weighted sequences than in T1-weighted sequences because longer TRs allow a more complete recovery of the longitudinal magnetization, and transverse relaxation times (T2) are fairly independent of the main magnetic field strength. Thus, patients referred for MR cholangiography may benefit from an ultrahigh-field-strength MR examination.

Second, chemical shift artifacts of the first kind are twice as large on ultrahigh-field MRI as on standard 1.5-T MRI. Chemical shift artifacts of the second kind, on the other hand, do not increase in size, although the timing is altered. Fortunately, the increased difference in resonant frequency between water and fat at 3.0 T is also advantageous because it allows a better separation of the fat and water peaks during MR spectroscopy and a better or faster fat suppression using chemical shift techniques.

Third, susceptibility artifacts are twice as big on 3.0-T MRI. Although patients referred for a colon study may be challenging, the search for gas—for example, free air or pneumobilia—should be easier on 3.0-T MRI. Patients with metal implants should undergo an MR examination at 3.0 T only if the metal-containing device has been proven to be MR safe for this field strength.

Fourth, standing wave and conductivity effects are usually not seen at a field strength of 1.5 T. At 3.0 T, these artifacts are most pronounced in pregnant women in the second and third trimesters because of the large amount of amniotic fluid and the increased size of the abdomen. Fetal MRI should therefore not be performed at 3.0 T because of these severe artifacts and the increased safety concerns. The same holds true for patients with a large amount of ascites, who are also not well suited for an ultrahigh-field MRI examination.

Finally, most patients can undergo an abdominal MRI study at 3.0 T with a reasonable outcome in terms of image quality.

Acknowledgments

We thank David E. Purdy and H. Cecil Charles for reviewing the manuscript and the subsequent stimulating discussions.

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