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