is currently a fourth-year Resident in Radiology at the
University of Michigan, Ann Arbor, MI. He completed his PhD under
Paul Lauterbur, PhD, at the University of Illinois,
Urbana-Champaign, and received his MD from the same institution
in 2000. He recently completed a year of Postdoctoral Fellowship
(2003-2004) at the Institute for Physics at the University of
Würzburg, Germany. Following residency, he will stay at the
University of Michigan for a Fellowship in Body MRI.
Higher field strengths have become attractive in nearly all
aspects of magnetic resonance due to the potential for increased
signal-to-noise ratios (SNR) and diminished acquisition times. Even
leaving aside the world of high-resolution, high-field protein
spectroscopy and focusing only on imaging, it becomes clear that
the definition of "high field" is different for different people.
In animal imaging and localized spectroscopy, for example, at least
2 institutions have at their disposal the services of 17.6T (750
MHz proton resonance) vertical bore spectrometers equipped with
imaging gradients. In the realm of human imaging for research, 4T
(170 MHz) systems have long been employed, 7T (298 MHz) and 8T (341
MHz) systems are functional, and a 9.4T (400 MHz) system is
operational and can soon be used for human imaging (Personal
comminication, Dr. X. Joe Zhou, University of Illinois, Chicago,
IL, July 2004).
The definition of high field in the clinic has also been
shifting. In the early 1980s, 0.3T was considered the upper limit
of practical imaging.
At this time, design considerations for radiofrequency (RF) coils
and difficulties with RF penetration seemed to limit the
application of high-field imaging.
Even as a high-field clinical standard of 1.5T (64 MHz) was being
established, questions were raised regarding what may be an ideal
field strength for imaging, given issues of power deposition,
relaxation times, chemical shift artifacts, and imaging times
required at higher fields.
It is interesting to note that many of these same concerns again
dominate the discussion nearly 20 years later as we move from a
1.5T (now standard) to 3T (128 MHz) as the new clinical high field.
As this review aims to offer the clinical scientist a summary of
high-field imaging and spectroscopy, this last definition of
high-field imaging will be used here.
Rationale: Why high field?
Put simply, the lure of higher field strengths is in the gain in
SNR that is possible with higher resonance frequencies. It is
worthwhile to take a closer look at how this gain in SNR comes
about. The SNR in magnetic resonance imaging (MRI) can be broken
down into 2 components- the detected signal and the noise.
Evaluation of each of these 2 components of the SNR reveals the
theoretical advantage of higher field MRI.
In nuclear magnetic resonance (NMR), and by extension MRI, the
magnitude of the detectable signal available for manipulation using
RF and gradient pulses is proportional to the difference in
population between spins or nuclei in 2 energy levels created when
the sample (or patient) is placed within the magnetic field.
To a first approximation, this difference in spin population
between the parallel and antiparallel configurations (
) can be written as follows
(See Equation 1.)
is the total number of protons in the body, γ the gyromagnetic
is Plank's constant,
is the Boltzmann constant, and
is the temperature. It is evident from equation 1 that this
difference in populations, which also represents the magnitude of
the detectable NMR signal, is directly proportional to the main
magnetic field strength,
A second contribution to an increase in detectable signal at
high field strengths comes from the physics of the signal detection
Signal detection in an RF coil can be seen as an application of
Faraday's law, where current is generated in a wire loop due to a
magnetic fiux enclosed by this loop of wire. An electromotive
is induced across the loop as follows:
(See Equation 2.)
represents the magnetic fiux in the coil. This magnetic fiux is
proportional to the precession frequency, ω
, which, in turn, is proportional to the main magnetic field
. Thus, the higher the field strength, the greater the induced
voltage in the RF coil, independent of the larger number of protons
available for detection (Equation 1). Putting the 2 factors
together, one sees that the signal strength in NMR is proportional
to the square of ω
However, the contribution of noise to the SNR must also be taken
into account. The noise, in terms of voltage in the detected
signal, can be written as:
(See Equation 3.)
are again the Boltzmann constant and temperature, respectively,
is the coil resistance,
is the effective patient resistance, and ∆
is the signal bandwidth. At higher fields (including 1.5T), when
large coils are used, the noise is dominated by the patient
resistance, which is proportional to the square of the field
strength. It becomes evident that the noise is linearly
notwithstanding adjustments to the bandwidth.
Combining the relationships of signal strength and noise to the
field strength, the SNR in MRI can be expected to increase
approximately linearly with field strength (for
The direct relationship between signal strength and
has been shown to hold true between 0.12T and 1.5T
and between 4T and 7T.
However, this analysis assumes that a constant bandwidth is used
for imaging. Depending on the imaging needs at hand, the receiver
bandwidth may also need to be increased (as, for example, with
echoplanar imaging [EPI], or when chemical shift artifacts are of
concern), further increasing the noise in the system and decreasing
the achievable SNR gain.
Some complicating factors
If SNR were the only factor to be considered, then this
discussion would imply that there is always a benefit from going to
ever-higher field strengths. However, high-field imaging also poses
some challenges that complicate such a straightforward analysis.
Relaxation effects, for instance, pose a distinct disadvantage at
higher fields. Various authors have examined the effects of field
strength on relaxation properties from 0.0002T up to 7T.
A tabulation of T1, T2, and T2* relaxation times for field
strengths ranging from 1.5T to 7T
and theoretical discussions about the dependence of relaxation
times on resonance frequency
can be found elsewhere. To summarize, however, T1 lengthens with
increasing field strength, while T2 and T2* both shorten.
These field-dependent effects have a pronounced impact on
imaging. The lengthening of T1, for example, means that for
non-single-shot sequences, a longer repetition time (TR) is
required for equivalent T1 weighting, or else an SNR or
contrast-to-noise ratio (CNR) penalty must be paid, partially
offsetting the benefits of field strength. T2 shortening is less
pronounced than the T1 lengthening. Nevertheless, this means that
for an equivalent echo time (TE), there will be a larger percentage
signal loss in high-field strength imaging than for an image with
the same parameters acquired at a lower field strength. The
shortening of T2* results in accentuation of susceptibility
gradients, both within tissue and at interfaces with vast
differences in magnetic susceptibility. Fortuitously, this produces
an increase in the blood-oxygenation-level-dependent (BOLD)
contrast used for most functional MRI (fMRI) experiments (see
below). However, it also results in worsening of susceptibility
artifacts in many rapid imaging sequences such as EPI
and spiral scanning,
where the echo-train length is limited by T2*.
While the nonionizing RF electromagnetic radiation used for the
in MRI is nonionizing, it does have the potential to cause adverse
effects. This is primarily in form of tissue heating from absorbed
RF radiation, mainly due to magnetic induction.
The specific absorption rate (SAR), given in units of joules per
kilogram, is a complex calculation measuring the rate of RF power
coupling with tissue, normalized by exposed mass.
MR parameters infiuencing the SAR include the duty cycle, D
(infiuenced by the number and spacing of RF pulses), the power of
the transmitted pulses (and therefore fiip angle, α), and the
resonant frequency, ω
. For a sphere with radius
and conductivity σ, for example:
(See Equation 4.)
From this, it is evident that adverse SAR effects are
exponentially worsened at high fields. It is also clear that
sequences, such as fast spin echo (FSE), that employ a large number
of closely spaced RF pulses with large fiip angles, are most likely
to cause increased power deposition at higher fields, and thus face
most limitations due to regulations on allowable SAR levels.
Unfortunately, because they rely on T2-weighted rather than
T2*-weighted echo trains, FSE sequences are also those least
sensitive to susceptibility induced image distortions, and,
therefore, potentially extremely useful at higher magnetic fields.
Furthermore, if the SAR problems of the affected sequences are
partially alleviated by employing longer repetition times (TR),
then not only does the imaging time lengthen, but also the T1
weighting of the sequence decreases, which is not necessarily
desirable. SAR effects, therefore, further complicate the choice of
sequences and parameters for high-field imaging and, potentially,
partially offset the benefits of the field strength.
Dielectric resonance and field focusing
Signal intensity in high-field images often have a bright center
and a relatively hypointense periphery when volume coils are used,
and an inhomogeneous appearance when surface coils are used.
This phenomenon is caused by the distortion of the transmitted
field by the patient or subject.
As the field strength/ Larmor frequency increases and, therefore,
wavelength decreases, the body dimensions become significant in
comparison to the wavelength of the transmitted radiation,
particularly when the dielectric constant of the imaged subject is
large (as is the case with tissue), as this also serves to decrease
the wavelength. This allows constructive and destructive
interference effects in the transmitted
to become significant, giving rise to the described inhomogeneities
in high-field images.
This phenomenon is sometimes called "dielectric resonance" in the
literature when the central image brightening at high fields is
explained. However, because the high conductivity of tissue serves
to dampen the standing waves required for dielectric resonance,
some authors prefer to use the term "field focusing" to explain the
origin of the inho-mogeneities.
Imaging strategies for high fields
While the promise of higher SNR provides the attraction to
high-field imaging, preserving this hard-won SNR is a more complex
task. The penalty for T1 lengthening, for example, is paid with
either an increase in imaging time due to longer TRs for equivalent
T1-weighting, loss in some of the potential gain in SNR by
maintaining an equivalent TR, or by using some of the increased SNR
to lengthen the echo train in multiple echo rapid sequences.
Multiple strategies are available to counter the worsened SAR
effects in FSE sequences. The most obvious solutions are to
decrease the duty cycle by increasing the echo spacing and
decreasing the echo-train length, or to employ smaller
and nonconstant fiip angles, α.
However, both these methods involve tradeoffs-the former entails
increased imaging time, and the latter, an SNR penalty with
concomitant lengthening of imaging time or loss of contrast. The
result is suboptimal scan efficiency; the CNR gained per unit time
is not maximal.
An interesting method that is free of these limitations is the
use of "hyper-echoes." Hennig and Scheffier
describe a family of sequences that employs 2
+ 1 equally spaced refocusing pulses after an initial 90° pulse, as
90° - (α°)
- 180° - (-α°)
This leads to the formation of a hyperecho or "echo of echoes,"
which contains, for tissues with T1>>T2, greater signal
intensity than the echo would otherwise have from conventional
Carr-Purcell-Meiboom-Gill (CPMG) experiments, due to contributions
from stimulated echo pathways. By using the hyperecho as the center
k-space line for the image, the net result is greater SNR than that
seen in comparable conventional FSE sequences, with marked
reduction of SAR. This advantage of the technique is presented in
Figure 1, which shows that compared with the fully refocused
sequence (Figure 1A), use of the hyperecho technique (Figure 1C)
results in SAR reduction of 76%, while tissue SNR is increased. The
re-duced αacquisition (Figure 1B), in contrast, results in
noticeable loss of SNR.
Partially parallel acquisition (PPA), also commonly called
parallel imaging, is an area of MRI that also has great potential
to help with several difficulties of higher field imaging. These
techniques employ coil sensitivity information from an RF coil
array to partially encode the MRI image.
This information is commonly used to accelerate the imaging process
or improve the spatial resolution per unit acquisition time.
The application of PPA has several positive implications for
high-field imaging. In EPI applications, such as diffusion imaging,
PPA offers the ability to reduce the echo-train length and
therefore geometric distortion due to susceptibility artifacts, and
improve spatial resolution due to a decreased point-spread
function, again due to decreased T2* decay in the shortened
The potential for artifact reduction by application of PPA is
illustrated in Figure 2,
which depicts diffusion-weighted EPI images acquired without PPA
(Figure 2A) and with PPA, using acceleration factors of 2 to 4
(Figure 2B through 2D). SAR effects can also potentially be reduced
with the application of PPA due to a decrease in the number of
pulses required to obtain a given image with a set resolution,
although this does require an increase in dead time.
High-field imaging is the ideal platform for the use of PPA because
it helps to offset what is considered a drawback of parallel
imaging-SNR. The intrinsic increase in SNR offered at higher
magnetic field strengths can be used to apply PPA for improvement
in imaging time or efficiency in situations where SNR
considerations would normally limit its application.
Applications of high-field imaging
Functional MR imaging
Functional MRI was perhaps the application earliest affected by
questions regarding moving to fields higher than 1.5T. Most
commonly, brain function is imaged by using BOLD contrast.
Deoxyhemoglobin is paramagnetic, while oxyhemoglobin is
diamagnetic. When a region of the brain is activated, the blood
fiow to that region increases out of proportion to the change in
oxygenation extraction of the tissue, resulting in signal changes
dependent on fiow, volume, and oxygenation status. These changes
can be used to indirectly detect neuronal function.
This experiment is affected at multiple levels by the field
strength. The BOLD fMRI signal change is dependent approximately
linearly upon field strength.
Moreover, as the field strength increases, the BOLD fMRI signal
contains increasing contributions from tissue as venous signal
diminishes due to T2* shortening. At low field strengths, a greater
proportion (if not the majority) of the fMRI signal comes from
larger vessels that are not as well localized for imaging function.
Higher resolution studies, constrained at lower fields due to SNR,
are also possible, allowing better localization of function.
However, the physiological noise, particularly that due to
respiration, is accentuated at high fields due to worsening of the
susceptibility effect, although multiple options exist to minimize
or remove such problems. Thus, overall the BOLD experiment is
greatly aided by higher field strengths, and fMRI has been one of
the driving forces behind moving to high fields for human imaging,
and is pushing the development of 7T to 9.4T magnets for use in
Perfusion-based methods for imaging brain function, such as
arterial spin labeling (ASL),
are also improved at high magnetic fields, as they are inherently
SNR poor. Moreover, the prolonged T1 with increased field strength
means that a labeled spin tag persists for longer, improving the
sensitivity of the perfusion measurement, although the results are
more mixed in terms of advantages for the actual ASL fMRI
A fast growing application of MRI in recent years has been MR
angiography (MRA), which benefits greatly from moving to higher
fields. The most obvious reason is the increased inherent signal
available at higher fields. More importantly, however, the inherent
lengthening of T1 at higher fields means that the "tag" in
time-of-fiight (TOF) angiography persists for a longer time,
improving the available SNR and also aiding in background
suppression. The result is better small vessel delineation and
fewer artifacts in display of larger vessels.
Magnetization transfer contrast (MTC) is often used to aid in back-
ground suppression in angiography at 1.5T. While SAR considerations
do raise questions about the feasibility of MTC at very high fields
in humans, at least at 3T it has been shown to be applicable.
The advantage of higher field strength for MRA applications is
clearly seen in Figure 3,
which presents cerebral MRAs obtained at 1.5T (Figure 3A) and 3T
(Figure 3B). Vessel detail and background suppression are superior
in the 3T image. The future possibilities of high-field MRA can
also be seen from Figure 4, which depicts an image from a pulmonary
angiogram obtained at 3T. Again, the vessel detail obtainable is
exquisite, and the image shows how the gains described earlier
could help expand the role of MRA in diagnostic imaging.
MR spectroscopy and spectroscopic imaging
Single-voxel localized spectroscopy and spectroscopic imaging
are increasingly important tools in the radiologist's hands,
finding application in a wide variety of diseases, including, but
not limited to, epilepsy, stroke, multiple sclerosis, human
immunodeficiency virus, and dementias. Due to inherent SNR
limitations and, therefore, signal averaging/imaging time
requirements, these are also areas that have much to gain at high
fields. The SNR considerations resulting in doubling of SNR with a
concomitant increase in field strength have largely been discussed
earlier and hold true for MR spectroscopy (MRS). In addition, the
doubling of field strength from 1.5T to 3T results in doubling of
the chemical shift between metabolites. Theoretically, therefore,
MRS at 3T should show double the SNR, improved baseline definition,
and improved spectral resolution, allowing the detection of less
concentrated metabolites. In practice, however, the improvement in
SNR in at least 1 direct comparison of 3T versus 1.5T was
considerably less than the predicted doubling of the SNR.
This is thought to be largely due to line broadening from T2
relaxation and field inhomogeneities. Nevertheless, MRS does
improve substantially at higher field strengths and has found
clinical application at 3T, especially with the availability of
automated algorithms that ease implementation.
The higher achievable SNR also means that it may be possible in the
future to perform MRS more easily with nuclei such as
Imaging the heart presents a special challenge in MRI in
general. Since the heart is a moving target, cardiac imaging
requires high temporal resolution for imaging, while preserving
high spatial resolution for applications such as coronary artery
In this regard, high fields seem to be ideal for this area of MRI,
as they provide the promise of increased SNR that can be used to
increase imaging speed or resolution. However, in practice, the SNR
gain has been less than the theoretically predicted linear increase
with field strength.
This may be partially due to the fact that cardiac high-field
imaging is made difficult by shortened T2* times (susceptibility
problems are already heightened in the chest cavity), field
focusing/dielectric resonance effects, and the need for rapid
imaging pulse sequences that require shortly separated RF pulses
that pose SAR problems. Nevertheless, high fields have opened new
doors in cardiac imaging, improving the contrast and resolution
available for short- and long-axis cine imaging,
and coronary artery imaging.
Parallel imaging is especially applicable in cardiac imaging, where
speed is at a premium. Some of the possibilities of cardiac imaging
at high fields are illustrated by the anatomical detail seen in the
papillary muscles and the coronary arteries in Figures 5 and 6.
The promise of increased SNR means that nearly all aspects of
clinical MRI can theoretically be improved by moving to higher
field strength. As has been the case with many advances in MRI,
neurological MRI has particularly benefited from the move to higher
fields. In addition to the advances in angiography, fMRI, MRS, and
MR spectroscopic imaging (MRSI), all of which directly affect
neurological imaging, routine structural imaging of the brain and
spine shows improved resolution due to SNR gains. Therefore,
lesions are more conspicuous both in contrast enhanced and
These observations have included, but are not limited to, tumors,
demyelinating lesions, hemangiomas, and ischemic lesions. Diffusion
tensor imaging (DTI) is also constrained by SNR/resolution
tradeoffs. Thus, the results at high fields can be remarkable, as
illustrated by the fiber tracking experiment on a patient with a
glioma in Figure 7.
Body MR poses problems related to low SNR, high temporal
resolution re-quirements due to motion, large susceptibility
gradients, and often wide bandwidth requirements to avoid chemical
shift artifacts. However, body and musculoskeletal structural
imaging also benefit from the SNR gains (see sections on cardiac
imaging and angiography). This is illustrated by the pulmonary
angiogram in Figure 4, the high-resolution anatomic image of the
normal hand in Figure 8A, and the image of the dysplastic left hip
in Figure 8B. Much needed clinical experience in areas such as
liver MR, renal imaging, and high-resolution musculoskeletal
imaging is now being gained and will no doubt contribute to the
wider use of high-field clinical magnets.
Generally, in MRI there is a complex tradeoff between SNR,
imaging time, and resolution. A fundamental hardware im-provement,
such as moving to a higher field (a shifting target which presently
means 3T for the clinician), allows one to bypass this tradeoff and
gain SNR without concomitant penalties in temporal or spatial
resolution, potentially allowing clinical examinations and studies
previously impossible. This promise of SNR is the lure for the
scientist and clinician alike toward high-field magnets, despite
their monetary costs and siting requirements. A straightforward
analysis shows that SNR can be expected to increase at most
approximately linearly with field strength. However, high-field
imaging poses some unique challenges, including relaxation
properties, power deposition, and RF in-homogeneities. Thus, the
practically realizable SNR gain is often considerably less than the
maximal predictions. Nevertheless, by adjusting the imaging
sequences and parameters, and by employing strategies specifically
targeted at high-field imaging, the problems can be at least
partially overcome. This allows the preservation of some of the
hard-won SNR to greatly improve MR image quality in a wide variety
of applications, particularly those that are SNR-starved at
The author would like to thank Mr. Robin Heidemann of the
University of Würzburg for his invaluable help in obtaining figures
and for discussions pertaining to parallel imaging at high fields.
Dr. Ioannis Panagiotelis of Siemens Medical Solutions graciously
provided images for several of the figures depicted here. Many
thanks also to Mr. Peter Schmitt, Dr. Mark Griswold, Dr. Xavier
Helluy, Dr. Peter Jakob (all at the University of Würzburg), Dr.
Andrew G. Webb (University of Illinois at Urbana-Champaign and
University of Würzburg), Mr. Alberto Vazquez, and Dr. Hero Hussain
(University of Michigan) for stimulating discussions and help with
proofreading or preparation of the manuscript. Funding for the
author's past year of work was provided through the Wolfgang Paul
Preis by the Alexander von Humboldt Stiftung.