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JOURNAL OF MAGNETIC RESONANCE IMAGING 28:1273–1279 (2008)
Technical Note
Inflow Quantification in Three-Dimensional
Cardiovascular MR Imaging
Reza Nezafat, PhD,1* Daniel Herzka, MD,3* Christian Stehning, PhD,4*
Dana C. Peters, PhD,1 Kay Nehrke, MD,4 and Warren J. Manning, MD1,2
Purpose: To investigate blood inflow enhancement (or lack
thereof) in three-dimensional (3D) cardiovascular MR for both
single phase whole-heart and cine biventricular functions.
Materials and Methods: A 3D imaging sequence is proposed in which radiofrequency excitation gradient is
changed without modifying image acquisition or phase/
slice encoding. This imaging sequence enables direct inflow
measurement while retaining static voxel signal-to-noise
ratio. Inflow measurements were performed for both
spoiled gradient-echo (GRE) imaging and balanced steadystate free precession (SSFP) in 18 healthy subjects.
Results: For single phase imaging, increasing slab thickness from 3 to 10 cm lead to 73% and 59% reductions in
contrast-to-noise ratio (CNR) with GRE and SSFP, respectively. For cine acquisitions, systolic CNR was reduced by
85% and 50% for the GRE and SSFP acquisitions, respectively, while diastolic CNR was reduced by 64% and 42%.
Conclusion: There is significant loss of CNR between blood
and myocardium when using larger 3D slabs due to saturation of inflowing spins. The loss of contrast is less pronounced for SSFP than for GRE, though both acquisition
techniques suffer.
Key Words: 3D cardiac imaging, 3D whole heart imaging,
inflow quantification
J. Magn. Reson. Imaging 2008;28:1273–1279.
© 2008 Wiley-Liss, Inc.
IN TWO-DIMENSIONAL (2D) balanced steady-state free
precession (SSFP) and spoiled gradient echo (GRE) acquisitions, fresh and unsaturated spins are continuously flowing into the imaging slice. These inflow spins
1
Departments of Medicine (Cardiovascular Division) and 2Radiology,
Beth Israel Deaconess Medical Center and Harvard Medical School,
Beth Israel Deaconess Medical Center, Boston, Massachusetts.
3
Clinical Sites Research Program, Philips Research North America,
Briarcliff Manor, NY.
4
Tomographic Imaging, Philips Research Europe, Hamburg, Germany.
The first three authors contributed equally to this work.
*Address reprint requests to: R.N., Beth Israel Deaconess Medical Center, 330 Brookline Avenue, Boston, MA, 02215.
E-mail: [email protected]
Received February 19, 2008; Accepted June 6, 2008.
Grant sponsor: American Heart Association; Grant number AHA SDG0730339N
DOI 10.1002/jmri.21493
Published online in Wiley InterScience (www.interscience.wiley.com).
© 2008 Wiley-Liss, Inc.
tend to improve the blood–myocardium contrast-tonoise ratio (CNR) as they yield higher blood signal to
noise ratio (SNR), particularly for SSFP sequences that
use high imaging flip angles. The 3D cardiovascular MR
(CMR) imaging is appealing due to its potentially higher
SNR compared with 2D imaging (1). The 3D CMR imaging has recently gained importance due to the availability of phased arrays with a higher number of coil elements that enable the use of 2D parallel imaging to
reduce total acquisition time (2– 6). Whole heart 3D
coronary imaging, which applies a single large slab over
the entire heart in an axial orientation obviates the use
of multiple thin slab 3D acquisitions targeted to the
right and left coronary arteries (6 – 8). Similarly, single
breath-hold (9) or self-navigated 3D cine acquisitions
are now becoming feasible. In addition to SNR gain in
3D cine, slice misregistration is typically improved, permitting more accurate evaluation of cardiac function.
Both SSFP and GRE sequences are used in 3D CMR,
with SSFP imaging more commonly used at 1.5 Tesla (T)
due to its superior SNR and CNR (10,11). GRE imaging
is more robust at higher magnetic fields (e.g., 3.0T),
which display increased field inhomogeneity. Imaging
of coronary vein anatomy for evaluation of patients undergoing cardiac resynchronization therapy has also
been shown to be more robust with GRE than SSFP
even at 1.5T, mainly due to increased B0 field inhomogeneity surrounding the veins (12).
The transition from targeted thin 3D slab acquisitions to thick 3D slab whole heart should be advantageous with both GRE and SSFP due to higher spatial
coverage and potential SNR increase of 公Nz, where Nz is
the number of slice partitions or encoding steps (1). The
effects of slab size on inflow-based image contrast are
yet to be explored in a systematic manner.
In 2D and small slab 3D imaging, fresh inflow results
in blood signal enhancement thus blood–myocardium
CNR improvement. With large 3D slab acquisitions, inflowing spins experience a higher number of radiofrequency (RF) pulses within the imaging volume, driving
the magnetization closer to its steady-state magnitude
(13). Therefore, the effect of inflowing spins on image
contrast is expected to be greatly reduced (13). In the
absence of inflow enhancement, GRE produces a T1
weighting, leading to a reduction of blood–myocardium
CNR because both tissues have very similar T1s (1). In
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Nezafat et al.
Figure 1. Image acquisition set-up: a large 3D axial 10-cm slab (100 slices of 1 mm thickness) is prescribed to cover the entire
heart. As usual, frequency (x), phase (y), and slice (z) encoding are performed. In the reference experiment, the imaging RF pulse
excites the entire 10 cm volume, as shown, with an extra 20% slice encoding to prevent fold-over from RF imperfections. In
subsequent experiments, b, c, and d, the excitation volume is reduced by changing the magnitude of the RF encoding gradient,
reducing the effective slab width to 6.0, 4.0, and 3.0 cm, respectively. The frequency, slice and phase encoding gradients are left
unchanged and cover the entire volume. This experiment maintains voxel size and scan time constant, and, therefore, changes
in SNR and CNR can be attributed to inflowing spins.
contrast, SSFP yields images with T2/T1 contrast,
which in the absence of any inflow enhancement, might
be more advantageous in terms of blood–myocardium
CNR (1). Also, because the innate contrast produced by
these sequences may not be as good as that observed
with thin slab acquisitions, prepulses such as T2 magnetization preparation (14,15) or magnetization transfer (MT) may be used to create additional contrast
(12,16). However, during the long transit time of the
blood through a thick 3D-slab, considerable accumulative phase errors may evolve due to flow encoding of
the phase encoding gradients (17).
In this study, we sought to investigate inflow saturation in 3D single phase and cine CMR imaging by quantifying the contribution of inflowing spins to SNR and
CNR with varying slab thickness. An imaging sequence
is designed such that RF excitation slab thickness is
varied in different acquisitions, thereby enabling direct
inflow measurement while retaining static voxel SNR.
MATERIALS AND METHODS
To study the effect of inflowing blood in a large slab
prescription, a series of imaging studies were performed
in each subject. The RF excitation to image encoding
ratio, that is, excitation slab thickness relative to image
slab size, was changed between different image acquisitions by changing the amplitude of the slab selection
gradient of Gz (i.e., reducing the excitation size) between
different acquisitions, as shown in Figure 1. A slice
selective 3D acquisition is used with excited RF profile
covering the entire imaging volume. The excitation slab
thickness was reduced manually in consecutive acquisitions without modifying the image acquisition volume, that is, frequency and phase encoding steps. By
keeping the imaging parameters of slice, phase and
frequency encoding the same between all acquisitions,
the SNR and CNR changes between different acquisitions can be associated with the saturation of inflowing
spins. The flowing spins into the imaging plane will
experience different RF histories based on the excitation profile.
MR imaging studies were performed on two 1.5T Philips Achieva (Philips Medical Systems, Best, NL). One
system was equipped with a 16-channel receiver and a
5-element cardiac phased-array receiver coil for single
phase whole heart coronary acquisitions. The second
system used a 32-channel receiver and a 32-element
cardiac phased array for 3D cine scans.
In Vivo Studies
Three groups of healthy adult subjects were imaged. In
each group, images were acquired in 6 subjects (totaling 18 subjects, 10 females, mean age 27 years). For 3D
cine imaging, 6 healthy subjects were studied using a
32-channel coil with both 3D SSFP and GRE cine. For
single phase imaging, the other 12 subjects were studied using a 5-element coil using either 3D single phase
GRE (in 6 subjects) or 3D single phase SSFP (in 6 other
subjects). Written informed consent was obtained from
all participants and the protocols were approved by
both Institutional Review Boards.
3D GRE Single Phase Whole Heart CMR
An electrocardiogram triggered, segmented 2-chamber
SSFP cine (repetition time [TR] ⫽ 3 ms, echo time [TE] ⫽
1.4 ms, ␣ ⫽ 60°, temporal resolution of 29 ms, spatial
resolution of 1.6 ⫻ 2.0 mm2 reconstructed to 1.21 ⫻
1.21 mm2) dataset was acquired at the level of the
mid-ventricle to identify the onset of the diastolic rest
period. This was followed by four whole heart CMR
imaging studies that differed only in RF slab encoding
gradient amplitude. A single phase sequence similar to
one used for coronary artery and vein imaging was used
(7,12). To improve blood–myocardium contrast, an MT
pulse train consisting of eight 600° Sinc-Gaussian RF
Inflow Quantification in 3D Cardiovascular MRI
pulses, each with duration of 15 ms and frequency
off-set of 500 Hz was used for all MT preparations (12).
A spectrally selective fat saturation sequence was also
used to suppress the fat signal. The navigator beam was
positioned at the dome of the right hemidiaphragm with
an acceptance window of 7 mm, automatic respiratory
drift adaptation, and RF excitation angle of 25°. An
axial volume covering the entire heart was prescribed
off the two-chamber images and initial scout images.
For each 3D dataset, a 10-cm-thick slab was acquired,
using 50 slice encoding partitions reconstructed to 100
one-mm-thick slices. For imaging, 20 –25 RF excitations (asymmetric Gaussian-weighted-Sinc with four
and two lobes and duration of 672 ␮s) with spoiled GRE
readouts: TE ⫽ 1.1 ms, TR ⫽ 3.6 ms, ␣ ⫽ 30°, BW ⫽ 383
Hz/Pixel, acquisition time of ⬃ 4:40 min for heart rate
of 60 bpm. Partial echo (62.5%) was used in the acquisition. A field of view (FOV) of 270 ⫻ 270 mm2 was
imaged with a scan matrix of 140 ⫻ 140 yielding a voxel
size of 2 ⫻ 2 ⫻ 2 mm3 reconstructed to a 0.9 ⫻ 0.9 ⫻ 1
mm3. A total of four studies with slab thickness of 10
cm, 6 cm, 4 cm, and 3 cm were acquired. Due to the
long duration of these studies, the images were prescribed with lower spatial resolution to ensure the successful completion of all 4 studies in all subjects. No
parallel imaging was used to enable absolute SNR and
CNR comparisons.
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tions and imaging artifacts with SSFP from increased
field inhomogeneity. However, for consistency, in this
study, 3D GRE cine images were acquired at 1.5T as
well. The imaging parameters were: TR ⫽ 3.6 ms, TE ⫽
1.3 ms, BW ⫽ 1562 Hz/Pixel, and ␣ ⫽ 20°. The scan
geometry and matrix size were identical to those used
for the SSFP acquisition with the same acceleration rate
and coil. Both 3D GRE and SSFP cine images were
acquired in one imaging session in one group of healthy
subjects.
3D SSFP Single Phase Whole Heart CMR
For 3D SSFP inflow quantification, a similar study to
the one used for 3D GRE (previous section) was performed on six subjects. The following imaging parameters, if different from above, were used: TE ⫽ 1.8 ms,
TR ⫽ 3.7 ms, ␣ ⫽ 90°, BW ⫽ 1470 Hz/Pixel. An excitation Gaussian-weighted Sinc RF pulse with three lobes
and duration of 1.49 ms was used for SSFP imaging.
3D SSFP Cine CMR
A 3D cardiac cine acquisition was prescribed to image
the heart throughout the cardiac cycle. The imaging
parameters were as follows: TR ⫽ 3.6 ms, TE ⫽ 1.8 ms,
BW ⫽ 1562 Hz/Pixel, ␣⫽ 60°. A 10-cm 3D slab consisting of 25 slices (interpolated to 50 slices) was acquired
using a FOV of 270 ⫻ 220 mm2, with a scan matrix of
128 ⫻ 102, yielding a voxel size of 2.1 ⫻ 2.1 ⫻ 4 mm3.
A 2D acceleration of rate 2 ⫻ 2 in the phase and slice
encoding directions and a partial Fourier acquisition
(reduction factor 0.625) were used to reduce the acquisition time to fit within a single breath-hold (30 s).
Although respiratory navigators can be used to reduce
the time constraint on image acquisition, transition to
steady-state artifacts that could present a complicating
influence on the measurement of inflow effects. A 32element phased-array coil was used in all 3D cine experiments to enable highly accelerated imaging of rate
2 ⫻ 2.
3D GRE Cine CMR
Although 3D cine GRE is not currently being used at
1.5T due to insufficient blood–myocardium CNR, it is
commonly used in CMR at 3.0T because of SAR limita-
Figure 2. a– h: Mid-diastolic 3D SSFP and GRE images acquired with different slab widths: 10 cm (a,e), 6.0 cm (b,f), 4.0
cm (c,g), and 3.0 cm (d,h). The images demonstrate the increased contrast found at thinner slab widths. Because all
experiments are acquired with exactly the same parameters,
the loss in contrast is known to be a direct result of saturation
of inflowing spins. Note that the contrast in stationary tissues
(e.g., the chest wall) is the same for all acquisitions.
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Data Analysis
In whole heart 3D SSFP and GRE acquisitions, the SNR
of arterial blood and myocardium were measured by
drawing a region of interest (ROI) in the left ventricle
and myocardium in the septum in the middle slice. The
middle slice was chosen in all analyses to remove the
effects of slice imperfection. Complex image data were
not available due to use of partial Fourier acquisitions.
Thus, the SNR analysis was based on the magnitude
images. The standard deviation of the noise was measured using an ROI in the air across the chest wall. No
corrections to the noise statistics were applied to account for the use of multi-channel magnitude images in
the measurement process. SNR was calculated as the
ratio of the mean signal to the standard deviation of the
noise. CNR between the blood and myocardium was
measured as the mean signal difference divided by the
standard deviation of the noise. Regression analysis
was used to calculate slope of SNR and CNR (i.e., rate of
decline in SNR and CNR) in arbitrary unit. Percent
change of SNR or CNR per cm increase in slab thickness
was also calculated by dividing the calculated slope of
each acquisition, calculated from the regression analysis, to the mean SNR or CNR of different subjects at 3
cm acquisition. Parallel imaging (SENSE) was used for
3D cine acquisitions making an absolute measure of
SNR unfeasible. However, there was no difference in
imaging parameters and identical coil sensitivity maps
were used in all eight 3D cine acquisitions on each
volunteer, keeping g-factors relatively constant. Therefore, “relative” SNR and CNR (i.e., SNR and CNR scaled
with g-factor loss) were measured in 3D cine acquisitions and were directly comparable. The relative SNR
was normalized to the maximum value observed in the
Nezafat et al.
respective measurement series (i.e., using the smallest
slab thickness). Error bars were calculated as the standard deviation of the SNR measurements across different volunteers.
RESULTS
Figure 2 shows an example set of middle slices acquired
using the 3D single phase whole heart SSFP and GRE
sequences with RF excitation slab thicknesses of 10
cm, 6 cm, 4 cm, and 3 cm. These results show an
improvement in contrast between ventricular blood and
myocardium as the slab thickness decreases, which
can only be associated with enhancement from inflowing spins. It is clear that the thicker slab thicknesses
lead to saturation of spins flowing into the imaging
volume. Compared to SSFP, there is more inflow contrast enhancement in the GRE images.
Figure 3 shows the SNR and CNR measurements for
single phase whole heart acquisitions. SNR measurements for 3D GRE acquisitions show that SNR is lost
with a slope of 4.4 (5.6% loss per cm) and 1.5 (2.8% loss
per cm) for arterial blood and myocardium, respectively
(Fig. 3a). For SSFP acquisitions the slopes are 4.0 (5.3%
loss per cm) and 1.2 (3.0% loss per cm) for arterial blood
and myocardium, respectively (Fig. 3b). There is slight
decline in myocardium SNR that could be associated
with through plane motion of the myocardium. Blood–
myocardium CNR shows losses associated with saturation of inflowing spins for both sequences with slopes of
2.4 (9.9% loss per cm) and 2.7 (8.2% loss per cm) for
GRE and SSFP, respectively, with overall lower CNR in
GRE (Fig. 3c,d). Blood–myocardium CNR was reduced
by 73% and 59% by increasing the excitation slab
Figure 3. a,b: SNRs for GRE
(a) and SSFP (b) from 3D single
phase whole heart acquisitions. Blood SNR declines with
saturation for both sequences.
c,d: Blood myocardium CNR
measurements for GRE (c) and
SSFP (d) imaging sequences
show losses associated with
saturation of inflowing spins
for both sequences with overall
lower CNR in the GRE. Error
bars represent standard deviation.
Inflow Quantification in 3D Cardiovascular MRI
thickness from 3 cm to 10 cm in GRE and SSFP acquisitions, respectively.
Figure 4 shows sample 3D cine images of middle slice
acquired with GRE (top two rows) and SSFP (bottom two
rows) in mid systole and end diastole. As with the single
phase whole heart acquisitions, four slab sizes are
shown: 10 cm, 6 cm, 4 cm, and 3 cm. An improvement
in CNR with decreasing slab thickness can be seen in
both acquisitions with a more pronounced change in
images acquired with GRE. Comparing images acquired
in mid-systole versus diastole shows that there are considerable changes in blood signal in mid-systolic phase,
in which there is significant inflow, compared with the
diastolic phase. These images suggest that the variability of inflow throughout the cardiac cycle has an impact
on CNR and SNR.
Figure 5 shows the results from relative SNR and
CNR measurements made on cine acquisitions, both in
mid-systole and late diastole. Systolic blood SNR in
GRE images decreased with a slope of 27.3 (11.4% loss
per cm), while diastolic blood SNR in GRE images decreased with a slope of 7.0 (7.0% loss per cm). A lower
diastolic decline rate could be associated with a reduction in fresh inflowing blood during the diastolic phase
(Fig. 5a). Myocardial SNR did not show considerable
change with GRE (Fig. 5a). In SSFP images, both systolic and diastolic blood SNRs decreased with a slope of
19.2 (6.8% loss per cm) and 13.1 (5.4% loss per cm),
respectively (Fig. 5b). Similar to the results observed
with GRE, myocardial SNR in SSFP images showed
little change (Fig. 5b). Overall, SSFP showed higher SNR
for both diastolic and systolic phases compared with
GRE. There was insufficient blood–myocardium CNR in
Figure 4. Example 3D cine images in systole
and diastole for GRE (top rows) and SSFP (bottoms rows) acquired with different slab sizes.
The 3D SSFP images have higher CNR compared with GRE images. The CNR of SSFP images is improved with thinner slabs.
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the GRE cine acquisitions, with slopes of 27.2 (12.2%
loss per cm) and 6.9 (8.4% loss per cm) for systolic and
diastolic phases (Fig. 5c), respectively. The relative CNR
of GRE acquisitions decreased when the excitation slab
thickness was increased from 3 cm to 10 cm for both
systolic (85%) and diastolic (64%) cardiac phases. The
3D SSFP cine images showed higher CNR with a decline
slope of 18.9 (7.0% loss per cm) and 12.8 (5.6% loss per
cm) for systole and diastole, respectively (Fig. 5d). The
relative CNR of SSFP acquisitions decreased when the
excitation slab thickness was increased from 3 cm to 10
cm for both systolic (50%) and diastolic (42%) cardiac
phases.
DISCUSSION
In this study, we sought to investigate the saturation of
inflowing spins in both 3D single phase whole heart and
cine imaging. 3D cine is appealing due to ease of prescription, faster acquisition speed, and the reduction of slice
misregistration which could result in miscalculation of
left ventricle (LV) end diastolic and end systolic volumes.
Although recent advances in coil technology have already
enabled the acquisition of 3D cine images, image quality
has not been as good as that observed with 2D cine imaging in all slices mainly due to loss of contrast (e.g., such
as reduced CNR in LV apex (5,6)). Our results suggest that
blood signal saturation may be one of the main sources of
insufficient image quality in 3D cine. Other sources of
poor image quality could be imperfect 3D slice profile,
field inhomogeneity, mixing of blood components with a
different excitation history profile, or phase errors that
lead to flow artifacts due to undesired flow encoding by
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phase encoding gradients. Further study is required to
quantify the CNR loss resulting from a transition from 2D,
the gold-standard in CMR LV function evaluation, to thick
slab 3D imaging, as used here. The losses in SNR and
CNR quantified in this work reduce the effective gains
expected when increasing the number of z-partitions in
3D CMR. New approaches should be tested and found to
counteract the loss of contrast due to inflowing spins.
These may include the use of contrast agents as well as
other volumetric approaches such as multi-slab acquisitions.
At 1.5T, 3D whole heart coronary artery imaging using the SSFP sequence is an alternative to targeted
small slab acquisition using either SSFP or GRE. However, GRE acquisitions are more robust in the presence
of artifacts from spins flowing through field inhomogeneities (18) and pericardial fluid. Nevertheless, whole
heart imaging using GRE at 1.5T can be problematic
due to insufficient contrast. At 3.0T, the GRE sequence
has emerged as the choice for coronary imaging due to
its robustness to field inhomogeneity and lower specific
absorption rate. The methodology for inflow quantification proposed in this study could be used at 3.0T, with
the caveat that additional signal loss would result from
artifacts caused by flowing blood transiting through an
inhomogeneous field. Additional experiments are required to quantify inflow effects in CMR at 3.0 T.
Considering single phase whole heart imaging, we
observed 73% and 59% losses in blood SNR in GRE and
SSFP acquisitions, respectively, when comparing the
thinnest slabs (3 cm) and the thickest (10 cm) slabs.
The methodology used in this work was designed not to
affect voxel SNR in the absence of inflow, keeping it
Nezafat et al.
constant relative to the acquisition with a 10 cm slab.
Normally, the change from a 3 cm slab to a 10 cm slab
(assuming 2 mm z-partitions) should yield an increase
in SNR of 82% (⫽公(10/3) without considering blood
saturation due to a decrease in the noise. Hence, the
transition from thin slab to thick slab might still be
advantageous from an SNR perspective, albeit with
much reduced benefits. This SNR increase would also
directly translate into a CNR increase of the same magnitude. This work shows that this theoretical CNR improvement is reduced by considering inflow saturation.
Hence, alternative approaches should be considered to
restore contrast and image quality. Preparatory pulses
(e.g., T2 Prep or MT) or use of exogenous contrast agents
might further assist in maintaining contrast.
Study Limitations
In the 3D cine studies, parallel imaging was used to reduce the acquisition time to one single breath-hold induration. Nevertheless, the breath-hold duration was significant (30 s), and not appropriate for use in a clinical
examination. However, this study was designed to minimize confounding factors and imaging artifacts that could
arise from nonbreath hold acquisitions, permitting accurate assessment of inflowing spin saturation. For example, respiratory navigators could interrupt the steadystate, while averaging could introduce blurring.
The loss in blood SNR observed in this work is
primarily due to saturation of the signal of inflowing
spins. However, other second order effects could have
also contributed to the loss. Though care was taken
to minimize the differences between the 4 different
Figure 5. a,b: Relative SNR for GRE
(a) and SSFP (b) measured from 3D
cine images for systolic blood, diastolic blood and myocardium. Relative CNRs for GRE (c) and SSFP (d)
imaging sequences show that both
systolic and diastolic CNRs decrease
with increasing slab thickness. Error
bars represent standard deviation.
Inflow Quantification in 3D Cardiovascular MRI
acquisitions (with four different slab thicknesses),
the method of quantification used in this work does
lead to one difference: the amplitude of the RF slab
encoding gradient. Although the change in amplitude
had negligible effects on the timing of the acquisition,
it could have affected signal intensity. Intra-voxel
dephasing is most prominent for stronger gradients,
suggesting that experiments carried out with thinner
slabs were more prone to signal loss in the blood pool.
Hence, this mechanism could have biased the results
by decreasing the SNR of blood observer in thinner
slab acquisitions. Although the experiment could
have been carried out using first moment nulling
gradients which minimize intra-voxel dephasing,
these gradients are too time consuming as they extend TE and TR and are not regularly used in cardiac
imaging. Therefore, they were not used in this work.
Although this might have introduced a bias into the
resulting measurements, it is unlikely to be highly
significant for the SNRs reported in this work, which
were measured in the center of the blood pool.
An accurate computer-controlled flow phantom or
numerical simulation of inflow in the SSFP or GRE
imaging sequences study could further validate this
experimental imaging study. However, it is difficult to
mimic in vivo conditions due to the complicated flow
patterns observed in the cardiovascular system, especially in the ventricles. There are also additional in
vivo conditions such as transition of blood flow
through the inhomogeneous magnetic field (e.g.,
lung) or variations and position of the heart within
the chest cavity that could further complicate such
theoretical analysis. Although no flow phantom or
numerical study is presented, this in vivo study is
sufficient to demonstrate the loss of LV blood and
myocardial CNR and SNR in 3D CMR. Furthermore,
the static phantom study confirmed the accuracy of
the imaging sequence by validating the slice profile
and measuring the slice thickness.
The SNR and CNR measurements were performed only
in the middle slice to remove the effects of imperfections in
the slice profile. Inflow measurements would differ if performed closer to base or apex of the heart, although it is
likely the trends would have been similar to those observed in this work. Additionally, the measurements of
inflow saturation were performed in a relatively young
cohort of healthy subjects with relatively higher velocity of
the LV blood flow (i.e., lower inflow saturation) compared
with patients with cardiovascular disease.
The changes in the SNR and CNR values presented
in this study may differ for different imaging flip angles. For example, a higher imaging flip angle could
increase saturation, yielding lower overall CNR. The
flip angles chosen in this work reflect those used both
in the literature and in our clinical practice, for both
GRE and SSFP. Therefore, the results obtained within
this work are likely relevant to the majority of cardiac
MR scans performed using 3D imaging. Nevertheless,
the effect of imaging flip angle on inflow saturation in
the context of slab thickness needs further investigation.
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CONCLUSION
In conclusion, in this study, we quantify the inflow
enhancement in 3D single phase and 3D cine cardiac
imaging. The results show there is significant loss of
CNR between blood and myocardium when using larger
3D slabs due to saturation of inflowing spins. The loss
of contrast is less pronounced for SSFP than for GRE,
though both acquisition techniques suffer.
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