is a partner at Radio-logical Services of New York, Staten
is a partner at Fairfax Radiological Consultants, Fairfax
Contrast-enhanced magnetic resonance angiography (CEMRA) has
been rapidly introduced into widespread clinical use over the past
several years. Intravenous gadolinium chelates for MR angiography
minimize flow artifacts, saturation effects, and long imaging times
that degrade time-of-flight (TOF) MR angiography (MRA), and their
use in three-dimensional (3D) MRA has produced excellent results.
Contrast-enhanced MRA was originally developed and described for
use in high-field (1.5 tesla [T]) magnetic resonance imaging (MRI)
systems. In routine clinical practice, it is performed almost
exclusively at 1.0 to 1.5 T. Low-field MRI systems are limited by
slower pulse sequences and lower signal-to-noise ratio (SNR)
qualities. This article will describe our initial experience
performing CEMRA on a 0.5 T closed-bore system.
Gadolinium-enhanced MRA was performed on 8 patients (ages 25 to
78 years; 5 men, 3 women) referred for MRI and MRA evaluation of
the aorta (n = 4) or renal arteries (n = 4). Aortic studies were
performed to evaluate for aortic aneurysm, and renal studies were
performed to evaluate for renal artery patency and stenosis. All
examinations were performed on a 0.5 T closed-bore MRI system
(Contour, GE Medical Systems, Milwaukee, WI). The 0.5 T system has
a maximum gradient strength of 15 millitesla per meter and a
maximal slew rate of 17 tesla per meter per second.
Prior to the start of the examination, a 22-gauge intravenous
catheter was placed in a forearm vein. All patients were given
breath-holding instruction prior to the study and were able to hold
their breath for at least 20 seconds. A fast 3D vascular TOF
gradient-echo (3D FSPGR) pulse sequence (10 to 13 msec/1.8 to 4.2
msec [repetition time/echo time], 35š to 40š flip angle, bandwidth
31.2 kHz, 3 to 6 mm slice thickness, 36 to 45 * 27 to 33 cm field
of view, 256 * 128 matrix, 24 slices acquired in a 30- to 35-second
breath hold) was used for the CEMRA studies. The sequence
parameters were tailored individually for each case to balance the
need for complete anatomic coverage with minimizing overall
acquisition time. The 3D FSPGR was a commercially available
sequence that provides the low repetition and echo times necessary
for breath hold contrast-enhanced vascular imaging. For imaging the
aorta, the authors used the standard built-in body coil. For
imaging the renal arteries, a wrap-around abdominal flex coil (GE
Medical Systems, Milwaukee, WI) was employed for increased SNR
Following a fast gradient-echo axial scout series, the anatomic
area of interest was selected and a 3D slab was prescribed (in the
coronal plane for renal studies and in the sagittal plane for
aortic studies). The scan delay time varied from approximately 20
to 30 seconds based on the patient's circulation time as estimated
by the attending radiologist (SD). Two complete breath-held
sequential data sets were acquired following intravenous injection
of gadolinium-DTPA (Magnevist, Berlex Laboratories, Wayne, NJ) at
0.2 mmol/kg into an antecubital vein. Gadolinium-DTPA was injected
by hand at a rate of approximately 2 mL/sec followed by a 10-mL
saline flush. A power injector was not available. Scan time was
approximately 30 seconds for each acquisition. The second
acquisition was obtained 10 seconds after the first acquisition to
obtain an angiographic data set in the venous phase and to provide
a backup acquisition in case of unanticipated prolonged circulation
of contrast. After the 3D FSPGR data sets were acquired, they were
transferred to an independent workstation (Advantage Windows
version 3.1P, GE Medical Systems, Milwaukee, WI). The authors
performed multiplanar volume rendering (MPVR), maximum intensity
projection (MIP), and 3D volume rendering (3DVR) reformatting to
visualize the relevant vasculature.
Signal intensity (SI) measurements and standard deviations were
determined by drawing the largest possible user-defined regions of
interest (ROIs) over three separate areas of the arterial anatomy
(ascending aorta, aortic arch, and descending aorta for aortic
exams, and abdominal aorta and each renal artery for renal exams).
In addition, ROIs were placed over areas of fat and muscle for
image contrast calculations, as well as outside the body in the
phase-encoding direction to measure the mean and standard deviation
of the background noise. Signal-to-noise ratio of the aorta was
defined as the mean SI divided by the standard deviation of the
background noise. Aorta-fat contrast-to-noise (C/N) was defined as
the SI of the artery minus the SI of the brightest fat within the
imaging volume, divided by the standard deviation of the background
noise. Aorta-muscle C/N was defined as the SI of the aorta minus
the SI of skeletal muscle divided by the standard deviation of the
background noise. For each patient, data from the first of two
post-gadolinium acquisitions was used since the second acquisition
was technically inferior because it was obtained in the venous
The quantitative signal characteristics are listed in Table 1.
The SNR and contrast-to-noise ratio (CNR) values compared favorably
with those published in similar studies performed at 1.5 T.
Quantitative data was agreeable for both renal and aortic exams
(Table 1). No difference was noted between the mean aortic and
renal SNR and CNR values. Signal did tend to decrease at the edges
of the magnetic field (Figures 1 and 2). Because of the lower
matrix (128 phase encodes) and thicker partitions (3 to 6 mm), the
overall resolution was lower than CEMRAs performed on most systems
at 1.5 T. This was especially true when the 3D images were viewed
orthogonal to the slice selection direction, and when imaging the
renal arteries (Figure 3). Each of the eight studies yielded
acceptable angiograms and the reader's self-assessed diagnostic
confidence was considered high. There was no difficulty depicting
the type B aortic dissection (Figure 4) even with a slice thickness
of 6 mm, and complete anatomic coverage of the aorta was easily
obtained (Figure 5).
The advantages and disadvantages of low-field MRI systems have
been discussed in prior literature and texts.
The current literature describes a variety of applications of
low-field MRI, primarily for musculoskeletal and neuroradiology
To the authors' knowledge, however, there is no literature on the
use of low-field systems for contrast-enhanced MR angiography. This
is unfortunate since low-field MRA may be the only noninvasive
means of obtaining angiographic information on patients who may not
have access to a high-field system.
Although low-field systems obey the same physical laws as their
high-field counterparts, their lower field strength results in
noticeable differences between the two platforms. Decreased
gradient system requirements allow cost savings in gradient system
design at the expense of lower, slower receive bandwidths.
Decreased T1 relaxation times permit shorter TR and scan times.
Low-field systems have less chemical shift and susceptibility and
fewer flow artifacts than their high-field counterparts.
Additionally, the smaller fringe field permits smaller site
requirements and less-restraining "open bore" configurations.
Unfortunately, low-field MRI systems are also subject to numerous
disadvantages. Signal-to-noise ratio decreases proportionally with
field size. Susceptibility artifact also decreases, and while this
is at times advantageous there is decreased sensitivity for
detecting calcification, iron, and hemorrhage. There is also a
smaller selection of pulse sequences and imaging options because
there are fewer low-field systems in the marketplace.
The radiologist must be aware of the limitations imposed upon
scan quality by the low-field environment. By far, the most
important limitation of low-field systems is their inherently lower
SNR and resolution as compared with their high-field counterparts.
Selecting the appropriate scan parameters to maximize the SNR of
the 3D imaging volume is critical to the success of the low-field
MRA examination. In any MRI acquisition, slice thickness is
directly proportional to SNR and affects resolution as well as
extent of anatomic coverage. We wish to keep slice thickness as
large as possible to maximize signal and anatomic coverage, but we
must also keep it small to maintain anatomic resolution. When
imaging a tortuous aorta, this becomes a problem that must be
addressed by adjusting both slice thickness and the number of
locations per slab. We routinely use a 3- to 6-mm slice thickness
and adjust our number of locations per slab to cover the anatomy
accordingly. When imaging renal arteries, we try to use a slice
thickness closer to 3 mm to prevent loss of resolution, usually at
the expense of time, anatomic coverage, and SNR.
The relatively slower pulse sequences of our low-field system
can use only 20 to 24 locations per 3D slab to keep the acquisition
time around a reasonable breath-hold time of ¾ 30 seconds.
Frequently 30 seconds is the upper limit of time that many older
patients can hold their breath. For elderly or
respiratory-compromised patients, this may still be too long for a
breath-hold, and fewer locations must be chosen with sacrifice of
some anatomic coverage. Given a 3- to 6-mm slice thickness and 20
to 24 locations per 3D slab, we can cover a volume of 6 to 14 cm in
depth, which we feel is adequate coverage for the majority of
patients we encounter.
Low-field systems have less stringent gradient requirements than
their high-field counterparts and, as a result, have lower receive
bandwidths, which accounts in part for their longer scan times.
While this is advantageous for SNR, it imposes a limit on the
shortest scan time for a 3D acquisition, and therefore on the
shortest practical time for a breath-hold acquisition. In practice
we use the highest receive bandwidth available on the system (32
kHz) to obtain a scan prescription that will cover all relevant
anatomy in a single breath-hold.
There are several limitations and weaknesses in this report. We
do not have conventional angiographic correlation for our studies.
A direct comparison with conventional angiography would be very
helpful to determine the sensitivity and specificity of low-field
CEMRA. We did not prospectively estimate the patients' circulation
time by performing a "timing run," as is now commonly performed.
No automatic triggering or fluoroscopic triggering is available at
0.5 T. The "best guess" technique of determining the scan delay
time used here can result in a substantial decrease in image
quality if the circulation time is misjudged. Finally, the
resolution of these exams is clearly lower than at 1.5 T, and a
direct comparison of studies performed at both field strengths
would be of great interest.
The limited resolution characteristic of low-field imaging is
particularly important when imaging the renal arteries, where
vessel diameters of 6 mm are common. This limits low-field MR
angiographic imaging of these vessels to assessing vessel patency,
course, and congenital anomalies. Grading renal artery stenosis
with low-field MR angiography is admittedly difficult, and great
care must be taken to use thin slice thickness, proper slab
selection, and meticulous bolus contrast administration.
The authors were pleasantly surprised by the image quality of
our CEMRA studies performed at 0.5 T, despite the limitations
mentioned above. By carefully balancing the imaging parameters,
adequate breath-hold CEMRA studies of the aorta and renal arteries
can be performed at 0.5 T. Additional evaluation at this field
strength needs to be performed to
determine the accuracy of these studies. Institutions that have 0.5
T systems can consider using them to perform this technique when a
high-field system is not available and the patient's clinical
condition prohibits transfer to another facility.