Fast imaging techniques for body magnetic resonance (MR) imaging have evolved from faster pulse sequences to the current development of phased-array coils, parallel imaging, and 3.0 T magnets. This report will address the technical challenges of body MR imaging, the historical developments of fast imaging techniques, and the current advances and clinical applications of fast imaging techniques for body MR imaging.
is a 4th-year Radiology Resident at the Medical Center of
Delaware/Christiana Care Health System, Newark, DE. He received his
MD from Louisiana State University Medical Center in Shreveport,
LA. After completing his residency, Dr. Uppot will begin a
fellowship in Abdominal Imaging and Interventional Radiology at
Massachusetts General Hospital in Boston, MA in 2003.
is the Chairman of the Department of Radiology at the Medical
Center of Delaware/Christiana Care Health System, Newark, DE.
Fast imaging techniques for body magnetic resonance (MR)
imaging have evolved from faster pulse sequences to the current
development of phased-array coils, parallel imaging, and 3.0 T
magnets. This report will address the technical challenges of
body MR imaging, the historical developments of fast imaging
techniques, and the current advances and clinical applications of
fast imaging techniques for body MR imaging.
Currently, in most clinical practices, the role of magnetic
resonance (MR) in body imaging is to characterize solid organ
lesions identified on computed tomography, perform magnetic
resonance cholangiopancreatography (MRCP), and evaluate fetal
anomalies identified by ultrasound. Although MR imaging
historically provides superior tissue contrast, offers multiplanar
capabilities, has no ionizing radiation, and seems a natural choice
for body imaging, its role in body imaging had been limited because
of the slow image acquisition speed.
Technical challenges of body MR imaging
MR body imaging differs from neurologic or musculoskeletal
imaging because of the need for high temporal resolution.
High temporal resolution is needed to overcome cardiac and
respiratory motion, vascular pulsations, and bowel peristalsis. The
benefits of high temporal resolution with fast imaging techniques
include decreased motion-related artifacts, improved patient
acceptance, and increased patient throughput, reducing per patient
The tradeoffs of faster imaging, however, are decreased
signal-to-noise ratio (SNR) and decreased spatial resolution. The
challenge in fast MR imaging, therefore, is to decrease scan time,
while maximizing SNR and spatial resolution. Fast MR imaging
techniques are a result of advancements in pulse sequence design
and k-space acquisition techniques, advancements in gradient
technology, development of phased-array coils and parallel imaging,
and increases in magnet strength.
The first body MR imaging was performed on July 2, 1977 on an MR
scanner named "Indomitable" built by Dr. Raymond Damadian.
It was an axial proton density image of the thorax and took 4 hours
and 45 minutes to acquire. Since the "Indomitable," there has been
a continuous trend toward faster MR imaging (Table 1).
What determines the speed of MR image
In general, the speed of MR image acquisition is based on many
factors and scan parameters. Historically, attempts to reduce scan
times were made by reducing the repetition time (TR), the number of
phase-encoding steps, and the number of excitations (NEX) based on
the basic formula
TR is the time between successive radiofrequency (RF) excitations.
Decreasing the TR reduces the scan time.
The phase-encoding gradient localizes the signal along the short
axis of the anatomy based on the phase of the received MR signal.
Scan time is dependent on the phase-encoding gradient and how fast
it fills k-space
; k-space is a matrix that serves as a temporary computer data
storage space (Figure 1). Each data point in k-space holds
information for the entire image (Figure 2A). Data in the center of
k-space contains information regarding the image contrast (Figure
2B), and data in the periphery of k-space contain information
regarding the spatial resolution of the image (Figure 2C). In
conventional imaging, each phase-encoding step fills one line in
k-space (Figure 3).
: Increasing the NEX in each phase-encoding step allows more data
to be collected in each line of k-space. A lower NEX decreases scan
time, but also decreases SNR. In a conventional spin-echo imaging
sequence based on a TR of 2 seconds (500 to 4000 ms), 256
phase-encoding steps (in a 256 * 256 matrix), and 1 NEX, the total
scan time is 2000 * 256 * 1, or 8.5 minutes.
This prolonged acquisition time limits successful body imaging
because body physiology defines maximum breath-hold as
approximately 20 seconds, normal respiratory cycle as approximately
3 seconds, and cardiac-cycle period as approximately 1 second.
In order to avoid motion artifact, image acquisition times must be
shorter than these normal physiologic processes. Strategies to
reduce scan times involve reducing the value of each of the scan
time parameters (TR, phase-encoding steps, and NEX).
Fast imaging techniques
Faster pulse sequences
The conventional spin-echo sequence was introduced in the mid
It is the standard MR imaging pulse sequence with the application
of a 90° RF pulse, a 180° refocusing pulse, the slice selection,
and frequency-encoding and phase-encoding gradients (Figure 4A).
Each RF excitation results in 1 phase-encoding step that fills 1
line of k-space. Because of the long acquisition times (4 to 8
minutes) with conventional spin-echo sequences, it is not suitable
for body imaging. TR may be reduced in conventional spin-echo
sequences; however, this affects tissue contrast and results in
images that are T1-weighted.
In 1986, gradient-recalled echo (GRE) sequences were introduced.
These sequences modify the conventional spin echo by using smaller
(<90°) flip angles and by using gradients, rather than a 180°
rephasing pulse (Figure 4B). Gradient-recalled echo allows
T1-weighted, T1- and T2*-weighted, T2*-weighted, and proton density
images to be obtained with very short TRs. Image times were reduced
to 200 to 300 msec, for a 128 matrix, fast enough to obtain
single-slice breath- hold images of the abdomen.
Since the development of spin-echo and GRE sequences, numerous
pulse sequences have been developed (Table 2).
These new pulse sequences are the result of changes to the basic
spin-echo and GRE sequences combined with variations in how
phase-encoding steps are performed and how k-space is acquired.
These pulse sequences can be placed into 1 of 3 major categories:
multiple spin-echo, GRE, and echoplanar imaging (EPI).
Multiple spin-echo sequences
Multiple spin echo is also known as rapid acquisition relaxation
enhanced (RARE) sequence or fast spin echo on systems from GE
Medical Systems (Milwaukee, WI) and turbo spin echo on systems from
Phillips (Best, The Netherlands) and Siemens Medical Systems
(Iselin, NJ). Multiple spin echo, as described by Henning in 1986,
modifies the conventional spin-echo sequence by application of a
series of 180° refocusing pulses and obtaining echoes after each
180° pulse (Figure 4C). With each excitation, multiple lines of
k-space are filled. The number of echoes acquired per 90° RF pulse
is known as the echo train length (ETL) and imaging scan time is
decreased in proportion to the value of the ETL. For example, if
the ETL is 4, the imaging acquisition speed is increased by a
factor of 4. The echo time (TE) is only an "effective TE" and can
be used to weigh the image: for T1-weighted images, a short
effective TE is used; for T2-weighted images, a long effective TE
The advantage of a multiple spin-echo sequence is that k-space
is filled much more rapidly and image acquisition time is reduced.
Multiple spin-echo sequences are insensitive to magnetic
There are many disadvantages of multiple spin-echo sequences,
such as blurring caused by T2 decay
--with each refocusing echo, the signal intensity drops, resulting
in more distortion. Also, power deposition to the patient from
multiple 180° RF pulses is a limiting factor. The gradient and
spin-echo sequences have been developed to overcome this effect.
Another disadvantage of multiple spin-echo sequences is that fat is
bright on T2-weighted images because of no signal loss due to
J-coupling (carbon-proton interaction) effects.
Bright fat can interfere with evaluation of bright signal from
pathology (hyperintense lesion on T2-weighted images, or
gadolinium-enhancing lesion on T1-weighted images). Fat suppression
techniques are therefore important in fast MR imaging.
In GRE sequences, image contrast is determined by TR, TE, and
The rapid acquisition times in GRE allow for
electrocardiogram-gated cine images of the heart and true 3D
The disadvantage of GRE imaging is that it is susceptible to
magnetic field inhomogeneities, which degrade image quality. Also,
if TR is less than T2 decay time, echoes from one excitation will
remain present to interfere with subsequent echoes, also known as
transverse magnetization coherence. Pulse sequences have been
developed to eliminate this interference ("spoiled" GRE), or
capitalize on it and create constructive interference ("unspoiled"
Spoiled GRE sequences use RF pulse phase changes to prohibit
transverse magnetization coherence. Examples of spoiled GRE include
spoiled fast low-angle shot (FLASH) and spoiled gradient-recalled
acquisition in steady state (GRASS). Spoiled FLASH and spoiled
GRASS sequences are T1-weighting, with T1 contrast controlled by
the flip angle.
Sequences that encourage the coherence do so by refocusing the
transverse magnetization. Such pulse sequences include fast imaging
with steady-state precession (FISP), fast acquisition in steady
state (FAST), contrast-enhanced fast acquisition in steady state
(CE-FAST), steady-state free precession (SSFP), and reverse fast
acquisition in steady state imaging with steady-state precession
(PSIF). Both FISP and FAST images are based on the free induction
decay after each pulse and are T2-/T1-weighted, while CE-FAST and
PSIF are at the point of the subsequent RF pulse (Hahn echo) and
are therefore more T2-weighted.
In FISP, only the phase-encoding gradients are refocused, while in
true-FISP (described later in this article) all three gradients
(slice selection, frequency encoding, and phase encoding) are
In addition, ultra-fast GRE images using very low-flip angles
(5° to 10°) and very short TRs (<5 ms) have been developed that
provide a 'snap-shot' of tissue magnetization (snapshot FLASH and
magnetization-prepared rapid acquisition with gradient echo
These are proton densityweighted images with very little tissue
contrast. Ultra-fast snapshot FLASH can be used to obtain real-time
single-slice imaging of the beating heart without cardiac gating.
Echoplanar imaging was described by Mansfield in 1977
and was first used clinically in 1983.
In EPI, all of k-space is filled after a single excitation (Figure
4D [compare also with Figure 5D]). Echoplanar imaging can be
accomplished by applying a single 180° refocusing pulse after the
initial excitation pulse (spin-echo EPI) or via rapid gradient
switching (gradient-echo EPI). It can be single shot (all of
k-space in a single shot) or, if increased spatial resolution is
required, it can be interleaved (k-space is only partially filled
after each excitation). Because of the need for rapid on/off
switching and eddy-current free environments, EPI places great
demands on gradient coils.
Recent advances in gradient technology have made EPI a reality.
Ultra-fast scan times (total scan time <100 ms) in EPI have
allowed for morphological and functional cardiac imaging
and hepatic imaging.
The disadvantages of EPI are that it is limited because of magnetic
susceptibility, chemical shift artifact, and lower SNR; and it can
cause painful muscle stimulations if the gradients oscillate too
Modified k-space acquisition
Scan times may be decreased by modifying how k-space is
Rectangular field of view
In conventional imaging, k-space has a square field of view
(FOV) and typically 128 to 256 phase-encoding steps are needed to
fill k-space (Figure 5A). An alternative method is to change the
FOV to a rectangular space.
With the phase-encoding gradient oriented along the short axis of
the rectangle, less phase-encoding steps need to be performed to
fill k-space (Figure 5B). In the abdomen and pelvis, where the
cross section is more rectangular and no image information is
present in the area anterior to the abdomen, this method can be
used to reduce scan times.
Acquire only a part of k-space--half-Fourier imaging
As defined by the Hermitian conjugacy,
k-space is mathematically symmetric. Therefore, only part of
k-space may be filled and the remainder can be calculated. Two
techniques include half-Fourier imaging and fractional echo
imaging. Half-Fourier imaging makes use of the fact that k-space is
symmetric along the phase-encoding gradient. Therefore, if slightly
more than 50% of k-space is filled, the remainder can be calculated
mathematically (Figure 5C).
Fractional echo imaging, in contrast, uses the fact that k-space
is symmetric along the frequency gradient. Therefore, only a
fraction of the signal from the frequency-encoding gradient is
collected and the remaining is calculated (Figure 5D).
Fill multiple lines of k-space with each excitation
As described previously, multiple spin-echo and EPI sequences
reduce scan times by filling k-space faster (Figure 5E).
In spiral imaging, data in k-space are acquired in a series of
spiral trajectories (Figure 5F).
Spiral scanning decreases the switching demand on the gradient
coils when compared with gradient EPI. Acquiring data in a spiral
fashion is a more efficient use of gradient power. A spiral
trajectory can cover a given area of k-space with fewer repetitions
than linear acquisition of k-space. This method can reduce scanning
time and still maintain spatial resolution. Spiral MR has less
motion artifact than EPI because the spiral trajectory samples the
center views of k-space more than EPI.
Spiral imaging also has less misregistration artifacts (artifacts
related to obliquely moving spins relative to phase- and
frequency-encoding axes, appearing misplaced on the images). These
qualities make spiral imaging effective for the evaluation of flow.
Spiral MR imaging requires special reconstruction algorithms.
Because the center of k-space contains contrast information and
the periphery of k-space contains spatial resolution information,
in a dynamic study such as movement of a contrast bolus, only the
central portion of k-space needs to be updated.
This method saves time by only changing the central portion of
k-space and leaving the peripheral portion unchanged (Figure
Current advances and clinical applications
Recent advances in fast imaging techniques for body MR imaging
include advances in software: single-shot RARE, true FISP, and
volumetric interpolated breath-hold examination (VIBE); and
hardware: advances in gradient technology, phased-array coils and
parallel imaging, 3.0 T magnets, and moving table
Single-shot half-Fourier RARE is a type of multiple spin-echo
sequence in which 1 excitation fills half of k-space. The
advantages of this sequence include rapid acquisition times of 600
ms per image; a wide contrast range with T1-weighted, T2-weighted,
and hydrographic imaging; resistance to magnetic susceptibility
artifacts; and reduced tissue-inflow artifacts. The disadvantages
include high RF power deposition, image blurring as a result of T2
decay, cross talk, and hydrographic artifacts.
Its uses include MRCP,
fetal MR imaging,
and bowel imaging.
True FISP (Siemens Medical Systems), FIESTA (GE Medical Systems)
or balanced FFE (Phillips) are unspoiled GRE sequences in which
transverse magnetization is preserved and signal is acquired in a
steady-state free precession mode.
Its uses include cine imaging of the heart,
thoracic aortic dissections,
and small- and large-bowel imaging.
Volumetric interpolated breath-hold examination is a 3D GRE
sequence. Three-dimensional GRE may provide thinner sections, fat
saturation, higher SNR, and comparable image contrast than 2D GRE;
however, its use has been limited because of poor resolution and
limited anatomic coverage. An advanced technique, VIBE allows a
large volume of acquisitions with ¾2 mm pixel size within the time
frame of a single breath-hold.
It allows for image acquisition in 1 plane and reformatting in any
other plane used for the assessment of liver lesions and
high-resolution 3D MR cholangiography.
Advances in gradient technology have allowed fast MR pulse
sequences, such as EPI to become reality. Gradients are
electromagnetic coils situated in a 3D plane around the bore of the
magnet that perform the slice selection, frequency, and phase
encoding. The physical act of switching each of the gradients on
and off, or rise time (time to peak gradient strength), usually
takes 600 µsec.
Advances in gradient technology aim to increase the strength of the
gradient coils and decrease the rise times. Current gradient coil
peak strengths are 20 to 40 milliTeslas per meter (mT/m) and rise
times are 150 to 300 µsec.
High gradient switching rates are limited because they could induce
currents that may stimulate the heart and peripheral nerves.
Phased array coils and parallel imaging
Surface coils improve SNR by their proximity to the body.
However, surface coils can only image a limited area over which the
coil is placed. Phased-array coils are multiple surface coils
distributed throughout the area of interest that are decoupled
magnetically. The advantages of phased-array surface coils are a
higher SNR due to the proximity of the coils to the body surface
and the stronger signal that is received. Current phased-array
coils contain four to six elements.
Parallel imaging makes use of each of these elements to provide
spatial information. This spatial information acquired reduces the
number of phase-encoding steps and, therefore, allows for a 2- to
5-fold reduction in scan times.
Theoretically, all phase-encoding steps can be reduced with a very
large number of receivers.
However, most current efforts are directed toward partial parallel
techniques in which a few phased-array elements are used to reduce
the number of phase-encoding steps. Two partial parallel imaging
sequences recently introduced include simultaneous acquisition of
spatial harmonics (SMASH) and sensitivity encoding (SENSE).
First described in 1997, SMASH is a partial parallel imaging
In SMASH, multiple receiver coils acquire signal simultaneously
with different spatial sensitivities. These differences in spatial
sensitivity are used to reconstruct the missing data in k-space
from the reduced phase-encoding steps. Fewer phase-encoding steps
allow for faster scan times. A 2- to 4-times savings in image
acquisition time is possible with SMASH.
When the number of data points sampled in k-space is reduced,
the acquisition time is reduced. This less dense k-space preserves
spatial resolution, but reduces the FOV and introduces an aliasing,
or fold over, artifact. In SENSE, multiple receiver coils are used
and each receiver coil receives a different signal voltage from the
imaged object based on the position of the receiver coil relative
to the object.
Computational algorithms can use these differences in signal
voltage to subtract the aliasing effect and create a whole image.
The result of SENSE imaging is the acquisition of an unaliased
image with only half the k-space data and, therefore, half the scan
time. The disadvantage of SENSE is a decrease in SNR because of
fewer k-space data points, and the raw computational power
necessary for the complex reconstruction algorithms.
Increased magnet strengths
Currently, most clinical practices have 1.0 and 1.5 T magnets.
The FDA has approved 3.0 T magnets for clinical use. The main
advantage of increasing gradients is increased SNR. MR imaging
scanners with increased SNR supplement the SNR that is lost in
faster pulse sequences.
Moving table configurations
Historically, conventional MR imaging is restricted to a single
body region, ie, thorax, abdomen, or pelvis. With the development
of real-time gradient-echo imaging with true FISP, imaging of
multiple body regions is now feasible. Using a rolling table
platform moving at a speed of 5 cm/sec and real-time true FISP
imaging, MR imaging covering 150 cm, with 8-mm sections, and 110%
overlap was achieved in 30 seconds and could detect 8-mm hepatic
and pulmonary lesions.
This feasibility study opens up the possibility of whole-body MR
Fast imaging techniques for body MR imaging are in constant
evolution. The choice of pulse sequence and hardware will depend on
the clinical situation, the available hardware and software, and an
understanding of the tradeoffs in spatial resolution, SNR, FOV,
tissue contrast, and scanning time.
Current advances in fast techniques for body MR imaging offer
the potential of moving from anatomic imaging to physiologic
imaging. They offer the potential of moving from static imaging to
real-time anatomic evaluation, as in MR fluoroscopy, and in
real-time manipulation, as in MR intervention.
Future advances in fast MR imaging will involve continued increases
in the number of elements in phased-array coils and increases in