Download Fetal magnetic resonance imaging: jumping from 1.5 to 3 tesla

Pediatr Radiol (2014) 44:376–386
DOI 10.1007/s00247-013-2857-0
REVIEW
Fetal magnetic resonance imaging: jumping from 1.5 to 3 tesla
(preliminary experience)
Teresa Victoria & Diego Jaramillo & Timothy Paul Leslie Roberts &
Deborah Zarnow & Ann Michelle Johnson & Jorge Delgado &
Erika Rubesova & Arastoo Vossough
Received: 12 July 2013 / Revised: 20 October 2013 / Accepted: 11 December 2013
# Springer-Verlag Berlin Heidelberg 2014
Abstract Several attempts have been made at imaging
the fetus at 3 T as part of the continuous search for
increased image signal and better anatomical delineation
of the developing fetus. Until very recently, imaging of
the fetus at 3 T has been disappointing, with numerous
artifacts impeding image analysis. Better magnets and
coils and improved technology now allow imaging of
the fetus at greater magnetic strength, some hurdles in
the shape of imaging artifacts notwithstanding. In this
paper we present the preliminary experience of evaluating the developing fetus at 3 T and discuss several
artifacts encountered and techniques to decrease them,
as well as safety concerns associated with scanning the
fetus at higher magnetic strength.
Keywords Magnetic resonance imaging . Fetus . 3.0 tesla .
Safety . Artifacts
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T. Victoria (*)
Radiology Department, Center for Fetal Diagnosis and Treatment,
The Children’s Hospital of Philadelphia,
34th Street and Civic Center Boulevard,
Philadelphia, PA 10104, USA
e-mail: [email protected]
D. Jaramillo : T. P. L. Roberts : D. Zarnow : A. M. Johnson :
J. Delgado : A. Vossough
Radiology Department, The Children’s Hospital of Philadelphia,
Philadelphia, PA, USA
E. Rubesova
Department of Radiology, Lucile Packard Children’s Hospital,
Stanford University, Stanford, CA, USA
Introduction
MR imaging of the gravid patient was first reported in 1983
[1], with the initial applications primarily related to maternal
and placental abnormalities [2]. In the 1990s ultrafast MR
sequences evolved, decreasing fetal motion artifact and
allowing the evaluation of the highly mobile fetus. Fetal MR
imaging was soon shown to be complementary to US in the
evaluation of multiple fetal pathologies. In parallel, improved
MR imaging technology and stronger magnets allowed the
transition from the 0.6-T machines first available for imaging
to the currently utilized 1.5-T machines. In 2002, the U.S.
Food and Drug Administration approved the use of 3-T magnets for human use. Since then numerous papers have praised
its use for imaging of the brain [3], musculoskeletal system [4]
and abdomen [5] and for various indications [6].
The search for better anatomical delineation and evaluation
of the fetus warrants evaluation of the feasibility of imaging at
3 T. Another reason to consider fetal imaging at 3 T is that
some imaging centers only have one magnet available for all
of their imaging needs, that magnet being a 3-T system.
Anecdotal presentations have been made discussing imaging
of the fetus at 3 T, mostly coming from Europe and more often
with disheartening results. The goal of this paper is to present
the preliminary experience at the Children’s Hospital of Philadelphia with fetal imaging at 3 T, with a focus on one of the
main impediments: the numerous artifacts encountered at 3-T
MRI. These are particularly prominent in the case of the
moving fetus within an environment of amniotic fluid.
Gain in signal-to-noise ratio: advantages of 3-T imaging
The signal created at MR imaging is the result of an excess of
unpaired protons aligned in the direction of the magnetic field.
The number of aligned protons is directly proportional to the
Pediatr Radiol (2014) 44:376–386
strength of the magnetic field, such that at higher strengths,
more protons are available to create the MR signal with
resultant increased net magnetization. The signal depends on
the square of the magnetic field strength, quadrupling at 3 T;
yet the noise also doubles such that the overall theoretical gain
of signal-to-noise ratio (SNR) at 3 T doubles in magnitude.
Other parameters also affect the overall gain of signal, including longitudinal relaxation time (T1), transverse relaxation
time (T2) and sequence parameters. In practice, and as published in multiple prior papers [5–7], the overall gain of signalto-noise ratio at 3 T is more likely to be 1.7–1.8 times the
signal at 1.5 T (Fig. 1). This gain in SNR can then be exploited
to achieve higher image quality or traded for increased spatial
resolution, increased temporal resolution (decreased acquisition time), or a combination of both. In the case of fetal
imaging, where the target is usually very small and highly
mobile, this gain can theoretically represent a tremendous
advantage. The gain in signal-to-noise ratio can also allow
more streamlined implementation of SNR-hungry applications such as use of parallel imaging with multichannel coils
to speed up single-shot MRI protocols, to reduce echo time on
long echo train sequences to mitigate susceptibility, and to
decrease radiofrequency heating by minimizing the number of
pulses.
3-T coils and gain in signal-to-noise ratio
One crucial contributor to the SNR performance is the
radiofrequency coil. Phased-array technique combines
multiple coil elements to provide a large field of view
while maintaining signal. Our 3-T phased-array coils
have 32 available elements, of which, depending on
the field of view, 8 are usually used for imaging at a
given time during fetal scanning. This is in contrast to
the typical 6-channel phased-array coil configuration
used at 1.5-T fetal scanning. In addition to the
phased-array coil, a surface coil is placed anterior to
the mother’s abdomen; for the 3-T magnet the surface
coil is equipped with 18 elements, whereas for the 1.5T magnet it contains only 6 elements. An important
advantage of utilizing phased-array coils is the potential
to employ parallel imaging. In this technique, the number of samples in k-space is reduced by a factor of two
or greater, thereby significantly shortening the acquisition time [8], a critical advantage in fetal imaging, when
it is difficult for the mother to hold her breath for long
periods of time and when the fetus is frequently moving
(Fig. 2). Relative loss of signal-to-noise ratio is a limitation to the use of parallel imaging. However the
overall gain of signal while imaging at 3 T is such that
this small loss does not detract from the gain of using
this technique.
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3-T protocol modification
Changes in T1 imaging
As the magnetic field strength increases, so does tissue T1
relaxation time. For the soft tissues of the abdomen and pelvis
in the adult patient T1 relaxation time has been reported to be
approximately 20–40% longer at 3 T than at 1.5 T [9] and for
adult musculoskeletal tissues, approximately 15–22% greater
[4]. Although no data detail exact relaxation times in fetal tissue,
it is fair to assume that they are also longer at 3 T than at 1.5 T. If
the same signal strength is to be maintained, repetition time
should be increased. However an increase in repetition time
results in unwanted increased acquisition time and potentially
worsening motion artifact of a highly mobile fetus. Our current
3-T protocols (Tables 1 and 2) reflect an increase in repetition
time at 3 T, while the matrix is held constant and the field of
view is decreased, with additional signal gain coming from the
increased number of coil elements at 3 T (see above). The
overall acquisition time remains essentially the same for both
magnet strengths (about 1 s per slice), such that image quality is
not compromised. The use of flip-angle characteristics can also
affect the T1 properties of the evaluated soft tissues, with
increasing angles resulting in increased T1 contrast (Fig. 3).
Practically speaking, there are some differences in the expected T1 signal of the fetal soft tissues at higher magnet
strength, the most conspicuous being the iso- to slightly hyperintense liver when compared to the lungs (Fig. 3). This
might be caused by the shorter T1 time and characteristics of
the liver compared to the lungs. At 1.5 T, the inherent T1
hyperintense signal of the liver is a helpful feature in certain
pathologies, including congenital diaphragmatic hernia. In these
fetuses, it is crucial to determine whether there is herniated
intrathoracic liver, because a herniated intrathoracic liver is a
negative prognostic indicator of postnatal outcome [10–12].
MR can be very helpful in evaluating liver position because the
Fig. 1 Comparison between fetuses of equal gestational ages imaged at„
1.5 T and 3 T MRI. a, b Two fetuses at 23 weeks of gestational age.
Sagittal steady-state free precession (SSFP) images at (a) 1.5 T and (b)
3 T show increased soft-tissue contrast at 3 T, with increased detail of the
hepatic and renal parenchyma. c, d Two different fetuses, shown at
27 weeks of gestational age, who have left congenital diaphragmatic
hernia and herniated liver (arrow in d). In these sagittal half-acquisition
single-shot fast spin-echo (SSFSE) images at 1.5 T (c) and 3 T (d), the
overall conspicuity of the fetal soft tissues is increased at 3 T. e, f Axial
SSFP images of the same fetuses as in (c) and (d). At 3 T (f) there is more
detail of the herniated bowel than at 1.5 T (e), and the border between the
liver and bowel is better delineated. g, h SSFSE images of two fetuses at
29 weeks of gestational age. Axial images at 1.5 T (g) and at 3 T (h) show
the migrational bands at 3 T (arrows in h), whereas at 1.5 T they are not
visualized. i, j Same fetuses as in (g) and (h). Sagittal images at 1.5 T (i)
and at 3 T (j) show at 3 T a clear visualization of the aqueduct and optic
chiasm (arrow and arrowhead, respectively, in j). k Axial image shows the
fetal pancreas (arrows) in a different 28-week fetus can be seen at 3 T in
this SSFSE image; in the experience of the authors, the pancreas is never
seen at 1.5 T at this gestational age. B bowel, L liver, S stomach
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379
Fig. 2 Twenty-four-week fetus
with omphalocele. a Sagittal 3-T
SSFSE MR image of the fetus
without parallel imaging. b The
same fetus with parallel imaging.
Note the increased conspicuity of
the soft tissues and increased
signal-to-noise ratio in (b)
manifested as increased sharpness
of the vertebral bodies and cardiac
contour. SSFSE half-acquisition
single-shot fast spin echo
echogenicity of liver and lung at US is similar, limiting
evaluation [13]. At 3 T, where the fetal liver is iso- to slightly
hyperintense with respect to lung, it can be more difficult than
at 1.5 T to unequivocally define liver position. However, when
analyzed in conjunction with other available fluidsensitive sequences, including echoplanar sequences
and single-shot fast spin-echo (SSFSE) and steady-state
free precession (SSFP) sequences, liver position should
be easily discernible.
Another difference when imaging at T1 is the improved
contrast of the osseous structures, which remain dark against
the intermediate signal of the surrounding soft tissues. At 3 T
Table 1 Scanning parameters for fetal body MRI
Plane
Sequence
FS
TR/TE (ms)
ETL
FOV (mm)
Matrix
Slice/gap
NEX
Other imaging parameters
3-plane localizers
Sagittal
Coronal
T2 SSFSE
T1 SGE
T1 VIGRE
No
No
No
1,100/80
180/4.76
3.49/1.37
192
-
250–300
250–300
250–300
265×205
256×205
192×192
3/0
4/0
2/0
1
1
1
3 planes
3 planes
T2 SSFP
T2 SSFSE
No
No
4.35/1.86
1,100/78
205
250–300
250–300
256×256
256×205
4/0
3/0
1
1
FA-180, Pat-2, PPF-4/8, PO-63%
FA-75, Pat-2, PO-25%
Breath hold
PPF-6/8, SPF-6/8
PO-40%, FA-10%
PO-25%, FA-70
Pat-2
PO-63%, FA-180, PPF-4/8
ETL echo train length, FA flip angle, FOV field of view, FS fat saturation, Pat parallel acquisition technique, PO phase oversampling, PPF partial phase
fourier, SGE spoiled gradient echo, SPF slice partial fourier, SSFP steady-state free precession, SSFSE half-acquisition single-shot fast spin echo, VIGRE
volume interpolated gradient echo
Table 2 Scanning parameters, neuro-fetal MRI
Plane
Sequence
FS
TR/TE (ms)
ETL
FOV (mm)
Matrix
Slice/gap
NEX
Other imaging parameters
3 planes
3 planes
Axial, coronal
T2 SSFSE
T2 SSFSE
EPI
No
No
Yes
1,100/62
1,100/78
5,030/40
205
205
-
250–300
250–300
250–300
256×205
256×205
192×192
3/0
3/0
3/0
1
1
1
FA-150, PO-50%, Pat-2
FA-150, PO-50%, Pat-2
PPF-7/8, FA-90
Axial, coronal
EPI
Yes
4,430/29
-
250–300
192×192
3/0
1
Axial
T1 SGE
No
204/4.76
-
250–300
256×166
4/0
1
EPI factor – 192, PO-25%, Pat-2
PPF-6/8, FA-90, Pat-2
EPI factor– 192, PO 25%
FA-60, PPF-7/8, PO-25%, Pat-2
EPI echoplanar imaging, ETL echo train length, FA flip angle, FOV field of view, FS fat saturation, Pat parallel acquisition technique, PO phase
oversampling, PPF partial phase fourier, SGE spoiled gradient echo, SSFSE half-acquisition single-shot fast spin echo
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Fig. 3 Coronal 3-T T1-weighted spoiled gradient echo (SGE) MR images
of a 30-week fetus at different flip angles. a 30°, (b) 60°, (c) 75°. Note
increased T1 hyperintensity of the liver (L) and meconium (arrow) with
increasing flip angles. d, e Coronal T1-W SGE images in a 1.5-T magnet
with flip angles of 30° (d) and 75° (e). Note the increased T1 character of
both liver and meconium at 1.5 T when compared to a 3-T magnet
increased sharpness and spatial resolution of the osseous
structures can be achieved (Fig. 4), including those of the
notoriously difficult-to-analyze hands and feet. This might
be helpful when evaluating the fetus with a presumed diagnosis of skeletal dysplasia, because US has limited sensitivity in
evaluating such entities and MRI at 1.5 T provides limited
bony detail [14]. Future studies could determine whether 3-T
fetal imaging helps in the prenatal evaluation of the fetus with
a diagnosis of skeletal dysplasia.
Also improved is visualization of the subcutaneous fat in
the developing fetus, which can be seen as early as the mid- to
late second trimester, whereas at 1.5 T this is delayed until the
third trimester (Fig. 5).
The remaining structures that are bright at 1.5 T remain so at
3 T, namely the thyroid and meconium [15]. It is worth mentioning that in our limited experience with the second- and
third-trimester fetus (the youngest fetus had a gestational age of
20 weeks), the dynamics of meconium signal, well described in
prior papers [16, 17], appear unchanged between 1.5 T and 3 T.
Fig. 4 Evaluation of bony structures at 3 T MRI. a Sagittal 3-T T1-W SGE
image of a 24-week fetus clearly demonstrates the five metacarpals (the hands
and feet are notoriously difficult to evaluate at 1.5 T). b T1-W SGE image
shows sharp scoliosis in this 24-week fetus. c Sagittal T1-W SGE image
demonstrates sharp visualization of the vertebral bodies in this 30-week fetus.
Note the relatively hypointense liver (arrow). SGE spoiled gradient echo
Changes in T2 imaging
T2 relaxation time is essentially unchanged or slightly decreased with increasing magnetic field strengths, possibly because some mechanisms of T2 relaxation are prolonged, whereas others become more efficient [6]. T2* is the observed or
Pediatr Radiol (2014) 44:376–386
381
Fig. 5 Coronal 3-T T1-W SGE
MRI in this 24-week fetus
demonstrates the rim of T1
hyperintense fat along the surface of
the fetus; this is not usually
identified until the third trimester
using 1.5 T MRI. SGE spoiled
gradient echo
also called “conductivity effect,” manifests as hypointense
areas in the regions of radiofrequency inhomogeneity [7,
18]. In fetal imaging, these two effects, an enlarged abdomen
coupled with the presence of amniotic fluid, combine to
produce particularly undesired artifacts that result in areas of
blackout centered in the field of view, where the fetus is
positioned. This artifact is particularly accentuated on the
single-shot fast spin-echo sequences, the workhorse of fetal
imaging.
Several approaches can be used to minimize standing wave
artifacts or so-called dielectric resonance artifacts. One feasible method to improve field homogeneity is the use of dielectric pads or radiofrequency cushions, given the method’s
simplicity and non-invasiveness. The cushions contain a gel
encapsulated in synthetic material. The gel is mixed with a
gadolinium- or manganese-based medium to eliminate the
signal from the gel itself. The conducting material has a high
dielectric constant and shorter wavelength than the surrounding soft tissues, altering the magnetic field inhomogeneity by
practically changing the geometry of the imaged subject [19].
There is one impediment of its use in the gravid patient: the
weight of the commercially available pad, 4–6 k, cannot be
tolerated when placed on the anterior abdominal wall of a
gravid patient. One way to bypass this is to create alternative,
lighter dielectric pads. In 2007 Kataoka et al. [20] showed in a
preliminary study that pads filled with saline solution can be
as effective as the ones filled with gel. We designed a saline
cushion that acts as a dielectric pad. This pad contains four
150-cc bags of saline solution placed in a pillow, which is
placed flat onto the maternal abdomen. This saline cushion,
which weighs about 2 lb in total, is easily tolerated by the
pregnant patient and results in partial dissipation of the dielectric artifact (Figs. 6 and 7).
Alternative ways to decrease the radiofrequency field inhomogeneity artifacts include use of multichannel transmission body coils [21] and active radiofrequency shimming
coupled with parallel imaging [22], although these techniques
are not widely available on all MR systems. As reported in the
literature, use of multichannel transmission body coils can
provide a more uniform radiofrequency field resulting in
enhanced field homogeneity, reduced dielectric shading and
improved consistency of the images [23].
In steady-state free precession sequences, field inhomogeneity can result in off-resonance effects, potentially aggravating banding artifacts, particularly at the edge of the field of
view. Changing the orientation of scanning or the field of
view, or modifying the radiofrequency frequency or bandwidth might help to displace the artifact away from the area
of interest, without eliminating the artifact from the image.
When interpreting the images, the practical effect of field
inhomogeneity results in a relatively greater time spent manipulating window and level settings, even if almost invariably each individual fetal structure can be better seen at 3 T.
effective T2 value, a product of the intrinsic T2 value mixed
with the local field inhomogeneity. The effect of T2* is more
pronounced at the greater magnet strength and results in greater
magnetic susceptibility than at 1.5 T, which is discussed below.
Artifacts encountered at 3 T
Magnetic field heterogeneity and standing wave artifact
At 3 T it is more difficult to maintain a homogeneous field,
resulting in worsening of a number of artifacts. Radiofrequency field inhomogeneity is one of the main challenges of fetal
imaging at 3 T, particularly for echoplanar, spin-echo and
steady-state free precession sequences. The increase in magnet strength translates into an increase in frequency and a
decrease in wavelength. In water and amniotic fluid, the
decreased radiofrequency wavelength at 3 T might approximate the size of the field of view. This causes the generation of
standing waves and constructive and destructive interference
patterns that result in heterogeneous signal, manifested as
areas of hypointensity or blackout mixed with areas of brightening. The larger the field of view in comparison with the
wavelength, the worse the artifact becomes, making the imaging of obese and pregnant patients particularly challenging
at 3 T [6, 7].
There is an additional hurdle in imaging fetuses at 3 T: the
presence of amniotic fluid. At 3 T the rapidly changing
magnetic field induces a circulating electrical current field
through the highly conductive amniotic fluid. The current acts
as an electromagnet that opposes the changing magnetic field,
reducing the amplitude and dissipating the energy of the
radiofrequency field. This radiofrequency shielding artifact,
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Fig. 6 Use of saline pads. Sagittal 3-T SSFSE MRI of a 25-week fetus
with myelomeningocele. The scan was done with (a) and without (b)
saline pads placed over the maternal abdomen. The dark band or standing
wave artifact secondary to dielectric artifact seen along the heart in (b) is
dissipated in (a). c The same imaging parameters as in (a) with the
exception of decreased slice thickness (3 mm in this image versus
5 mm in a). Note the increased resolution of the soft-tissue structures
with the higher slice thickness. SSFSE half-acquisition single-shot fast
spin echo
Magnetic susceptibility artifacts
times (TE). In brain imaging, use of a lower echo times
at 3-T imaging is often necessary compared to 1.5-T
imaging in order to decrease geometric distortion of the
head and accentuated susceptibility of intracranial vessels seen on echoplanar imaging. The normal susceptibility and hypointensity of deoxyhemoglobin within the
intracranial venous structures is accentuated at 3 T compared to 1.5 T, which can interfere with a reliable
diagnosis of intracranial hemorrhage. Reduction of echo
time is generally necessary at 3-T imaging to offset this
accentuated hypointensity and conspicuity of the vessels. Use of a wider readout bandwidth or use of smaller voxel sizes also have a similar effect in decreasing
susceptibility, but all of these modifications need to be
balanced with the resultant decrease in signal-to-noise
ratio.
Another artifact commonly encountered at 3 T fetal imaging is
magnetic susceptibility. This refers to the extent a material
becomes magnetized in a field and occurs at the interface of
substances with different magnetic susceptibility. Local field
inhomogeneity can result in intravoxel dephasing, resulting in
signal attenuation, image geometric distortion and localized
regions of high or low signal intensity [18]. Susceptibility
artifacts worsen with increasing magnetic strengths, being
approximately twice as large at 3 T as at 1.5 T [7]. This is
particularly the case with echoplanar sequences. This artifact
is more conspicuous at the interface of the gas-filled colon in
echoplanar imaging sequences (EPI), more so in the older,
vertex-positioned fetus, where the head is in close apposition
to the maternal rectum. Yet there are potential advantages to
this increased susceptibility, including increasing conspicuity
of intracranial or adrenal bleeds on echoplanar imaging sequences (Fig. 8).
A few technical modifications can minimize the degree of
distortion of the image, including changing the readout direction to alter the location of the artifact away from the field of
view, implementing parallel imaging and using shorter echo
Chemical shift artifact at 3 T
In general chemical shift artifacts are accentuated at 3 T
compared to 1.5-T imaging. Yet this is not an issue in fetal
imaging because the two sources of fetal fat (subcutaneous
Fig. 7 Dielectric pads. Axial 3-T EPI MRI of a 23-week fetus at the level of the kidneys shows the fetus with (a) and without (b) dielectric pads over the
maternal abdomen. Note the increased field homogeneity and decreased distortion of the images with the dielectric pad. EPI echoplanar imaging
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Fig. 8 Axial spin-echo EPI
sequence (a) performed at 3 T with
a technique typically used at 1.5 T
(TR/TE=4,930/94 ms, no parallel
imaging) shows marked geometric
distortion and susceptibility. b
Decreasing the TE to 29 ms and
enabling parallel acceleration
factor of 2 markedly decreases the
geometric distortion and excessive
susceptibility artifacts. c, d Axial
spin-echo EPI in a 25-week fetus
with myelomeningocele
demonstrates increased conspicuity
of normal veins and germinal
zones of the brain at higher echo
time (TE 42 ms in image c; TE
29 ms in image d). Care must be
exercised not to confuse these
normal structures with abnormal
foci of hemorrhage in the brain.
EPI echoplanar imaging, TE echo
time, TR repetition time
and intra-abdominal fat) appear relatively late in pregnancy,
thus not contributing to artifact formation.
diagnostic imaging, (2) the lower magnetic fields (16 T vs.
3 T) and (3) the fact that we are dealing with fetuses as
opposed to embryos.
Magnetic resonance safety
Gradient switching
While the benefits of signal-to-noise ratio and concomitant
advantages of spatial or temporal resolution are compelling
for fetal MRI at 3 T, several considerations should be
evaluated to ensure the safety of mother and child. A
premise of this discussion is the concept that, for clinically
indicated patients, the benefits of MRI outweigh any procedure risks at 1.5 T. Two to three decades of practice
generally support this widely held belief [24–26]. The
discussion below outlines key safety considerations and
their impact at 3 T MRI, in direct comparison to those at
1.5 T. Issues to be addressed are:
In conventional MRI, the process of image formation requires
the application of exogenous linear magnetic field gradients
(in x- or y- or z- directions) at various short time periods, to
encode spatial position across the body. Most current MRI
devices limit the maximum magnetic field gradient to approximately 40–50 mT/m. Given this relatively low value in
contrast to the 3 T (3,000 mT) of a static field it is not likely
that the mere presence of the gradient field is contributory to
risk. However because these gradient fields are pulsed on and
off at key times as dictated by the imaging pulse sequence,
attention has to be paid to the changing magnetic field experienced during onset and offset of a magnetic field gradient
pulse. Especially in rapid imaging techniques, such as
echoplanar imaging, where gradient pulses are applied and
reversed extremely rapidly in order to obtain snapshot images,
it is possible that high rates of change of magnetic field are
transiently experienced. By Faraday’s law of electromagnetic
induction, a conducting material experiencing such a changing magnetic field will experience an electrical current. In
practice this means that exposure to magnetic field gradient
switching carries with it the risk of involuntary peripheral
nerve stimulation. Most contemporary MR hardware limits
the magnetic field gradient switching rate (slew rate) to ~200
mT/m/ms. Given the length of magnetic field gradient coils
(~1 m), this implies a changing magnetic field of up to 100
(1) static field exposure,
(2) gradient field switching, and
(3) radiofrequency power deposition.
Static field exposure
There is little direct evidence of harmful effects with shortterm exposure to high static magnetic fields. Although one
study of frog embryos exposed to long durations of 16-T
magnetic fields demonstrated a mechanistically plausible interaction [27], it is difficult to extrapolate from this paper
given (1) the much shorter exposure times in humans during
384
mT/ms (or 100 T/s), 50-cm removed from the center of the
imaging field of view. U.S. Food and Drug Administration
safety limits, in fact, prevent exposure to changing magnetic
fields (dB/dt) of more than 60 T/s, with scanner failsafe
software limiting these exposures. Furthermore, because the
fetus is generally at or near the center of the imaging field of
view during imaging, fetal exposure changes are much lower
(e.g., 1 cm from the imaging FOV center the dB/dt is only 2 T/
s even at maximum slew rates). Of further note, maximum
slew rates are similar at both 1.5 T and 3 T and thus incremental risks at 3 T vs. 1.5 T are essentially negligible. In fact,
the FDA limits of exposure to changing magnetic fields are
independent of magnetic field strength.
Radiofrequency power deposition
The excitation and refocusing radiofrequency pulses of
the MR sequences deposit energy into the subject being
imaged. The rate of energy deposition is defined by the
specific absorption rate (SAR), which is reported in
W/kg and is based on approximate modeling of the
subject body.
The rate of energy deposition is dependent on the amplitude of the applied radiofrequency fields, which tends to
increase with increasing magnetic field strength and in fact
quadruples in transitioning from 1.5 T to 3 T (all other factors
being kept constant). Optimizing radiofrequency pulse design
might be expected to mitigate this phenomenon, for example
decreasing the number of slices, lowering the flip angles,
increasing repetition time, the use of parallel imaging techniques and shorter echo train lengths (in fast spin-echo sequences) can all decrease the specific absorption rate. Although in general, energy deposited can lead to undesired
subject heating, a variety of fluid and flow mechanisms might
mitigate this in practice. However these factors are not considered in determining worst-case-scenario limits to specific
absorption rate. The FDA safety limits for radiofrequency
exposure are set at 4 W/kg for the maternal whole body
(independent of static magnetic field strength), with scanner
failsafe controls prohibiting exposure beyond that. Of practical note, typical fetal MR pulse sequences are shown in
Table 3, along with their specific absorption rate loads
at 1.5 T and 3 T for approximately equivalent image
contrasts. The values were obtained for a phantom fetus
(Fig. 9), constructed with bottles of different consistencies (oil, water, cream-based salad dressing) that could
simulate the soft-tissue consistencies of the fetus. This
phantom fetus was wrapped in saline bags imitating the
amniotic fluid and then placed inside an adult body-type
phantom emulating the mother. Their specific absorption
rates were then obtained at 1.5 T and 3 T. Values were
also obtained for real fetuses at 1.5 T and 3 T, all of
Pediatr Radiol (2014) 44:376–386
Table 3 Specific absorption rate loads for MR imaging at 1.5 T and 3 T.
Note that values are always kept below the 4 W/kg required by the U.S.
Food and Drug Administration
1.5 T
1.5 T
3T
3T
Fetus
Phantom
Fetus
Phantom
SSFSE
SSFP
T1
EPI
≤2.1
≤0.8
≤2.8
1.6
≤2.1
0.7
1.4
1.4
0.6
≤0.8
1.6
1.1
≤0.7
<0.7
0.2
0.2
EPI echoplanar imaging, SSFP steady-state free precession, SSFSE halfacquisition single-shot fast spin echo
them kept well below the whole-body exposure limit of
4 W/kg required by the FDA.
Aside from the maternal whole-body specific absorption
rate limitations stipulated by the FDA, there is an additional
consideration to the specific absorption rate received by the
fetus: its inhomogeneous distribution among the imaged fetal
soft tissues [28]. Radiofrequency field inhomogeneity and
dielectric and standing wave effects may produce focal SAR
hot spots that can be projected around or within the fetus. This
complication is addressed by recent sophisticated modeling of
the fetal environment by Hand et al. [28, 29], who suggest that
the maximum energy deposited to the fetus is a fraction of that
received by the mother. In their calculations, the highest local
specific absorption rate is in the mother, with the fetus being
exposed to a peak of 50–70% of the maternal maximum SAR
at 3 T. They also showed that in MRI studies done under
normal-mode operating conditions set forth by the International Electrotechnical Commission in 2008, where the maternal whole-body average specific absorption rate is ≤2 W/kg
and the fetal whole-body specific absorption rate is 1.24 W/kg
and 1.14 W/kg at typical 1.5-T and 3-T configurations, respectively [28]. They also stipulated that the average temperature
of the fetus under these conditions remains below 38OC.
Care must be placed to limit the amount of time imaging the fetus, however, because the maximum temperature predicted in the fetus can exceed 38OC following
continuous exposures of about 7.5 min or longer. In
Fig. 9 Fetal phantom containing fluids of different consistencies is
wrapped in saline solution bags, created to measure the specific absorption rates. Refer to Table 3 for further correlation
Pediatr Radiol (2014) 44:376–386
practice, our fetal imaging protocols do not use long
continuous exposures, with most sequences falling well
below the 1-min scanning time.
Alternative methods to decrease specific absorption
rate aim at minimizing dielectric and standing wave artifacts by decreasing radiofrequency field inhomogeneity, a
task that can be accomplished by the use of multichannel
phased-array and parallel transmission, as discussed above
[30, 31].
Taken together, although there are very real concerns about
the safety of the MRI procedure to the fetus at 3 T, the various
mechanisms that could theoretically have a biologically plausible impact are limited to the levels associated with ongoing
scanning at 1.5 T.
IRB approval and informed consent prior to scanning fetuses
at 3 T
The rationale for scanning at 3 T is to take advantage of the
additional signal intensity and resolution. The goal is to obtain
the best and most detailed anatomical information of the fetus
to make the best informed decisions about management. With
this in mind, after thorough discussion with the Institutional
Review Board (IRB) department, and armed with extensive
information about the safety of fetal scanning at 3 T, we
proceeded to scan. The reason to do so was to optimize a
clinically indicated study of the fetus, done to potentially
improve the visualization of existing abnormalities. Because
this was not research but a clinically indicated examination,
IRB approval was not required, as agreed by the head of the
IRB department.
Our institution does not require written informed consent
prior to MR scanning at 1.5 T or 3 T for any anatomical part.
This includes fetal imaging at 1.5 T. We extrapolated this
thinking for fetal 3-T imaging.
We also recognize that these guidelines might not apply at
all institutions, and that directors at each center have to make
policies and rules as they see fit.
Conclusion
We have presented our preliminary experience in imaging the
fetus at 3 T. As expected in other fields of 3-T imaging, the
signal-to-noise ratio is superior to that at 1.5 T when technique
is optimized. Yet numerous hurdles are presented along the
way, mainly in the form of artifacts. These can be minimized
with techniques discussed in this paper. Future work will be
done to further improve the available sequences with the aim
of obtaining the highest-quality images to improve diagnostic
accuracy when evaluating the developing fetus.
Conflicts of interest None
385
References
1. Smith FW, Adam AH, Phillips WD (1983) NMR imaging in pregnancy. Lancet 1:61–62
2. Stark DD, McCarthy SM, Filly RA et al (1985) Pelvimetry by
magnetic resonance imaging. AJR Am J Roentgenol 144:947–950
3. DeLano MC, Fisher C (2006) 3 T MR imaging of the brain. Magn
Reson Imaging Clin N Am 14:77–88
4. Kuo R, Panchal M, Tanenbaum L et al (2007) 3.0 Tesla imaging of
the musculoskeletal system. J Magn Reson Imaging 25:245–261
5. Chang KJ, Kamel IR, Macura KJ et al (2008) 3.0-T MR imaging of the abdomen: comparison with 1.5 T. Radiographics 28:
1983–1998
6. Barth MM, Smith MP, Pedrosa I et al (2007) Body MR imaging at
3.0 T: understanding the opportunities and challenges. Radiographics
27:1445–1462, discussion 1462–1464
7. Erturk SM, Alberich-Bayarri A, Herrmann KA et al (2009) Use of
3.0-T MR imaging for evaluation of the abdomen. Radiographics 29:
1547–1563
8. Glockner JF, Hu HH, Stanley DW et al (2005) Parallel MR imaging:
a user’s guide. Radiographics 25:1279–1297
9. De Bazelaire CM, Duhamel GD, Rofsky NM et al (2004) MR
imaging relaxation times of abdominal and pelvic tissues measured
in vivo at 3.0 T: preliminary results. Radiology 230:652–659
10. Hedrick HL, Danzer E, Merchant A et al (2007) Liver position and
lung-to-head ratio for prediction of extracorporeal membrane oxygenation and survival in isolated left congenital diaphragmatic hernia.
Am J Obstet Gynecol 197:e421–e424
11. Victoria T, Bebbington MW, Danzer E et al (2012) Use of magnetic
resonance imaging in prenatal prognosis of the fetus with isolated left
congenital diaphragmatic hernia. Prenat Diagn 32:715–723
12. Jani J, Cannie M, Sonigo P et al (2008) Value of prenatal magnetic
resonance imaging in the prediction of postnatal outcome in fetuses
with diaphragmatic hernia. Ultrasound Obstet Gynecol 32:793–799
13. Victoria T, Danzer E, Adzick NS (2013) Use of ultrasound and MRI
for evaluation of lung volumes in fetuses with isolated left congenital
diaphragmatic hernia. Semin Pediatr Surg 22:30–36
14. Victoria T, Epelman M, Coleman BG et al (2013) Low-dose fetal CT
in the prenatal evaluation of skeletal dysplasias and other severe
skeletal abnormalities. AJR Am J Roentgenol 200:989–1000
15. Zizka J, Elias P, Hodik K et al (2006) Liver, meconium, haemorrhage:
the value of T1-weighted images in fetal MRI. Pediatr Radiol 36:
792–801
16. Saguintaah M, Couture A, Veyrac C et al (2002) MRI of the fetal
gastrointestinal tract. Pediatr Radiol 32:395–404
17. Rubesova E, Vance CJ, Ringertz HG et al (2009) Three-dimensional
MRI volumetric measurements of the normal fetal colon. AJR Am J
Roentgenol 192:761–765
18. Soher BJ, Dale BM, Merkle EM (2007) A review of MR physics: 3 T
versus 1.5 T. Magn Reson Imaging Clin N Am 15:277–290
19. Franklin KM, Dale BM, Merkle EM (2008) Improvement in B1inhomogeneity artifacts in the abdomen at 3 T MR imaging using a
radiofrequency cushion. J Magn Reson Imaging 27:1443–1447
20. Kataoka M, Isoda H, Maetani Y et al (2007) MR imaging of the
female pelvis at 3 tesla: evaluation of image homogeneity using
different dielectric pads. J Magn Reson Imaging 26:1572–1577
21. Vernickel P, Roschmann P, Findeklee C et al (2007) Eightchannel transmit/receive body MRI coil at 3 T. Magn Reson
Med 58:381–389
22. Ullmann P, Junge S, Wick M et al (2005) Experimental analysis
of parallel excitation using dedicated coil setups and simultaneous RF transmission on multiple channels. Magn Reson Med
54:994–1001
23. Willinek WA, Gieseke J, Kukuk GM et al (2010) Dual-source parallel
radiofrequency excitation body MR imaging compared with standard
386
24.
25.
26.
27.
Pediatr Radiol (2014) 44:376–386
MR imaging at 3.0 T: initial clinical experience. Radiology 256:966–
975
Baker PN, Johnson IR, Harvey PR et al (1994) A three-year followup of children imaged in utero with echo-planar magnetic resonance.
Am J Obstet Gynecol 170:32–33
Kanal E, Gillen J, Evans JA et al (1993) Survey of reproductive
health among female MR workers. Radiology 187:395–399
Myers C, Duncan KR, Gowland PA et al (1998) Failure to detect
intrauterine growth restriction following in utero exposure to MRI. Br
J Radiol 71:549–551
Denegre JM, Valles JM Jr, Lin K et al (1998) Cleavage planes in frog
eggs are altered by strong magnetic fields. Proc Natl Acad Sci U S A
95:14729–14732
28. Hand JW, Li Y, Hajnal JV (2010) Numerical study of RF exposure
and the resulting temperature rise in the foetus during a magnetic
resonance procedure. Phys Med Biol 55:913–930
29. Hand JW, Li Y, Thomas EL et al (2006) Prediction of specific
absorption rate in mother and fetus associated with MRI examinations during pregnancy. Magn Reson Med 55:883–893
30. Van den Berg CA, van den Bergen B, Van de Kamer JB et al (2007)
Simultaneous B1+homogenization and specific absorption rate
hotspot suppression using a magnetic resonance phased array transmit coil. Magn Reson Med 57:577–586
31. Homann H, Graesslin I, Eggers H et al (2012) Local SAR management by RF shimming: a simulation study with multiple human body
models. MAGMA 25:193–204