Investigation of Inter-Slice Magnetization Transfer Effects as a New Method for MTR Imaging of the Human Brain
Investigation of Inter-Slice Magnetization Transfer Effects as a New Method for MTR Imaging of the Human Brain
Jeffrey W. Barker 0 1 2
Paul Kyu Han 0 1 2
Seung Hong Choi 0 1 2
Kyongtae Ty Bae 0 1 2
Sung- Hong Park 0 1 2
0 1 Department of Radiology, University of Pittsburgh, Pittsburgh, Pennsylvania, United States of America, 2 Department of Bioengineering, University of Pittsburgh, Pittsburgh, Pennsylvania, United States of America, 3 Department of Bio and Brain Engineering, Korea Advanced Institute of Science and Technology , Daejeon , South Korea , 4 Department of Radiology, Seoul National University College of Medicine , Seoul , South Korea
1 Data Availability Statement: Data are openly available from Figshare at the following DOI: http://dx. doi.org/10.6084/m9.figshare.1281061. Data are also available from the D-Scholarship Institutional Repository at the University of Pittsburgh at the URL: http://d-scholarship.pitt.edu/23772/
2 Academic Editor: Su Lui, West China Hospital of Sichuan University , CHINA
We present a new method for magnetization transfer (MT) ratio imaging in the brain that requires no separate saturation pulse. Interslice MT effects that are inherent to multi-slice balanced steady-state free precession (bSSFP) imaging were controlled via an interslice delay time to generate MT-weighted (0 s delay) and reference images (5-8 s delay) for MT ratio (MTR) imaging of the brain. The effects of varying flip angle and phase encoding (PE) order were investigated experimentally in normal, healthy subjects. Values of up to *50% and *40% were observed for white and gray matter MTR. Centric PE showed larger MTR, higher SNR, and better contrast between white and gray matter than linear PE. Simulations of a two-pool model of MT agreed well with in vivo MTR values. Simulations were also used to investigate the effects of varying acquisition parameters, and the effects of varying flip angle, PE steps, and interslice delay are discussed. Lastly, we demonstrated reduced banding with a non-balanced SSFP-FID sequence and showed preliminary results of interslice MTR imaging of meningioma.
Funding: This study was supported by a grant from
National Research Foundation of Korea
(NRF2013R1A1A1061759) and a grant from Competitive
Medical Research Fund of the UPMC Health System.
The funders had no role in study design, data
collection and analysis, decision to publish, or
preparation of the manuscript.
Protons that are bound to macromolecules can exchange magnetization with free water
protons leading to magnetization transfer (MT) phenomena . Macromolecular protons cannot
be observed directly with magnetic resonance imaging (MRI) because of fast transverse
relaxation (T2 * 10 s); however, macromolecular protons can be preferentially saturated by
offresonance (with respect to free water) radio frequency (RF) irradiation, since the absorption
spectrum of macromolecular protons is much broader than that of free water. Exchange
between the two pools of protons can transfer an observable decrease in magnetization to the free
water pool. The percent signal decrease due to MT is called the magnetization transfer ratio
(MTR). Changes in MTR can often give information about the macromolecules involved in
Competing Interests: The authors have declared
that no competing interests exist.
generating the MT effects. One of the major applications of MTR imaging has been the
evaluation of white matter (WM) integrity in multiple sclerosis , in which myelin content has been
found to be significantly correlated with MTR . Other applications include imaging articular
cartilage of the knee , intervertebral disc degeneration , and characterization of brain
tumors [8, 9].
Interslice MT effects that are inherent to sequential multi-slice acquisitions  have
been used to generate contrast for MT asymmetry imaging without a separate saturation pulse
in the alternate ascending/descending directional navigation (ALADDIN) pulse sequence .
We hypothesize that this approach may be used for MTR imaging with an additional
acquisition of reference images without MT weighting. The purpose of this work was to demonstrate
the feasibility of MTR imaging in the brain using interslice MT effects to generate MT contrast
and to investigate the characteristics of interslice MTR image acquisition with multi-slice
balanced steady state free precession (bSSFP) imaging.
In this study, we validated the source of image contrast using a 10% agar phantom (high
MT effects) and saline phantom (no MT effects) and by comparing in vivo MTR images of the
brain acquired with the proposed interslice method to MTR images acquired with conventional
presaturation pulses. We compared in vivo MTR values for gray matter (GM) and WM at
varying flip angle and phase encoding (PE) order with predictions from numerical simulations of
the two-pool model using tissue parameters from the literature. Simulations were also used to
investigate the effects of varying the number of PE steps, flip angle, and interslice delay as well
as the accumulation of MT effects over multiple slices. Lastly, we demonstrate reduced banding
with non-balanced SSFP and applied the proposed interslice MTR imaging method to a
In 2D sequences, slice-selection is achieved by applying a linear gradient perpendicular to the
slice plane causing the Larmor precession frequency to vary as a function of position.
Excitation of a slice of interest is achieved by adjusting RF-pulse frequency to the Larmor frequency
of the desired plane. The slice of interest receives on-resonance excitation; however, the rest of
the volume receives off-resonance irradiation. During acquisition of one slice, excitation pulses
are effectively a series of off-resonance saturation pulses to future slices (Fig. 1). The
offresonance frequency received by neighboring slices is given by
where BW is the bandwidth of the RF-pulse, GAP is the interslice gap; THK is the slice
thickness; n is the slice index with positive indices indicating slices superior to the acquisition slice,
and negative indices indicating slices inferior to the acquisition slice; sign(GRAD) is the sign
of the gradient; ORD is +1 if ascending slice order and 1 if descending slice order. The
offresonance irradiation received by neighboring slices can saturate macromolecular protons
leading to interslice MT effects. This idea is illustrated in Fig. 1.
Interslice MT effects can be enhanced as a mechanism for generating contrast by the use of
bSSFP acquisition, in which high flip angle and short repetition time (TR) lead to high
saturation of the macromolecular pool. The interslice gap is set to a high value (e.g. 140% the slice
thickness) to avoid crosstalk caused by overlapping slice profiles and so that interleaving two
acquisitions gives a full set of images at typical gap size (e.g. 20% the slice thickness) . For
imaging in the brain, descending slice order is preferred to ascending slice order, in order to
Fig 1. Illustration of interslice MT effects. The application of a gradient varies the Larmor frequency f(z)
linearly in space (z). During excitation, the slice of interest (slice 0) receives on resonance excitation. With a
positive gradient polarity and descending slice order (shown above), the next slice to be acquired (slice 1)
receives off-resonance irradiation at a frequency offsets of 3840 Hz.
suppress signal contributions from blood perfusion. Because MT effects can accumulate over
multiple slices, a few extra dummy slices must be collected in the MT-weighted image sets to
ensure homogeneous MT contrast across slices. These slices can be positioned outside the
imaging volume of interest (e.g., above the head) and discarded during reconstruction. Reference
images without MT-weighting can be acquired by adding an interslice delay sufficient for T1
recovery. The MTR value, which measures the percent signal decrease, can be calculated pixel
by pixel as follows:
where IMT and IRef are the signal intensities of corresponding pixels in the MT-weighted and
reference images, respectively.
The modified two-pool model  can be used to quantitatively model MT effects
generated by sequential 2D bSSFP acquisitions. In this model, a free liquid pool (subscript f )
exchanges longitudinal magnetization with a restricted macromolecular pool (subscript r). Using
an RF-pulse given by
where RF is the angular frequency of the RF pulse, 0, f is the angular frequency at free water
resonace, and b1(t) is the magnitude of the RF pulse, the two pool model can be described
mathematically in a frame rotating at 0, f using the Bloch equations with additional terms for
exchange of longitudinal magnetization between the proton pools:
with z denoting the longitudinal component, and x and y denoting transverse components of
magnetization. The longitudinal and transverse magnetization decay constants are given by T1
and T2, respectively. The ratio of the fully-relaxed longitudinal magnetizations gives the ratio
of pool sizes (F = M0, r/M0, f), and kr is the pseudo-first-order exchange rate constant. The
saturation rate of the restricted pool is proportional to the square of the RF-pulse amplitude and to
the absorption lineshape, G():
where the term 0, r is the peak of the absorption spectrum of the macromolecular pool, which
can account for MT asymmetry when 0, r 6 0, f . For this study, MT asymmetry was not
considered (0, r = 0, f). On-resonance MT effects  were simulated for WM and GM by
setting G(0) = 1.4 105 s1 according to Gloor, Scheffler, and Bieri .
All imaging experiments were approved by the Institutional Review Boards at the University of
Pittsburgh and Seoul National University and written informed consent was obtained from
Instrumentation and Software
The experiments were performed on Siemens 3T Trio systems (Siemens Medical Solutions,
Erlangen, Germany), and a 12-element head matrix coil was used for reception with body coil
transmission for all data acquisitions. All simulations and analyses were performed with
Matlab (Mathworks, Natick, MA).
Raw data were reconstructed to images by the MR scanners. Dummy slices in the MT-weighted
images were discarded prior to MTR calculation. The number of MTR slices was equal to the
the number of reference image slices for all acquisitions described below. Calculation of MTR
images consisted of creating either a whole head or brain mask and calculating MTR pixel by
pixel inside the masked region according to Eq. 2. Brain masks were created from
segmentations generated using SPM8 software (Wellcome Trust Centre for Neuroimaging, London,
UK), whereas whole head masks were generated simply by thresholding based on intensity.
Computer simulations were performed using the parameters summarized in Table 1 for GM
and WM. We considered six prior slices of off-resonance saturation for MT-weighting (e.g.,
23040 Hz, 19200 Hz, 15360 Hz, 11520 Hz, 7680 Hz, 3840 Hz), unless otherwise indicated. For
reference image signal, the acquisition slice was simulated with no prior slices of off-resonance
saturation (i.e., full T1 recovery from prior slices), except when specifically investigating the
effects of varying interslice delay. We modeled excitation with a Gaussian windowed sinc pulse.
Other simulation parameters were taken to match acquisition parameters, such as flip angle,
number of PE steps, TR, and RF duration. The set of differential equations 47 were solved
using the 4th/5th order Runge-Kutta algorithm. For analysis, we calculated MTR values using
the magnitude of the transverse magnetization at the center line of k-space for MT-weighted
and reference signal simulations. In addition to comparing simulations with in vivo data, we
used simulations to investigate the dependence of MT contrast on the number of preceding
slices for varying flip angles and number of PE steps. Lastly, we investigated the effects of
varying the interslice delay from 08 s by simulating reference image acquisition with 6 prior slices
of off-resonance irradiation with a specified delay time between each slice.
A cylindrical 10% agar phantom with a diameter of 140 mm and height of 180 mm was imaged
for initial assessment of interslice MTR imaging. Additionally, a cylindrical saline phantom
with a diameter of 120 mm and height of 195 mm was imaged with the same acquisition
parameters as a negative control. The MTR images were reconstructed from MT-weighted and
reference bSSFP images according to Eq. 2. Centric PE order was used for acquisition. Other
acquisition parameters were as follows: slice order = descending; slice-select gradient polarity =
positive; readout gradient polarity = positive; TR/TE = 4.56/2.28 ms; matrix size = 256 256;
field of view = 256 256 mm2; flip angle = 50; slice thickness = 4 mm; interslice gap = 5.6 mm
(0.8 mm after interleaving); scan direction = axial; PE direction = anterior-posterior; dummy
PE steps = 30; phase oversampling = 50%; number of averages = 1; RF-pulse BW = 1600 Hz;
acquisition BW = 501 Hz/pixel; number of slices = 19 and 18 for each interleaved MT-weighted
*Measured in this study. All other values from Stanisz et. al. .
image set (including 6 dummy slices each; 12 total) and 25 for reference images; interslice
delay = 0 s for MT-weighted images and 6 s for reference images.
Comparison of Interslice and Presaturation MT Effects
For three normal, healthy subjects (age 2140), MTR images were acquired using the proposed
interslice method and using conventional presaturation pulses with an identical bSSFP readout
(single slice; no interslice MT effects), in order to confirm the source of image contrast in vivo.
Three different off-resonance irradiation frequencies were used for the presaturation pulses
corresponding to the offset frequencies of the first (3200 Hz), second (6400 Hz), and third
prior (9600 Hz) slices of the interslice method. The average RF-power off-resonance irradiation
of the two methods were equivalent (94 T). The following parameters of the bSSFP readout
were the same for both acquistion methods: TR/TE = 4.15/2.08 ms; matrix size = 128 128;
field of view = 220 220 mm2; flip angle = 60; slice thickness = 5 mm; phase partial Fourier =
6/8; scan direction = axial; PE direction = right-left; initial dummy PE steps = 30; centric PE
order; phase oversampling = 0%; RF-pulse BW = 1333 Hz; and acquisition BW = 592 Hz/pixel.
The interslice MTR images were acquired with descending slice order and positive
slice-select gradient. The rest of the imaging parameters for the proposed interslice method were as
follows: gap = 7 mm; number of average = 1; number of slices = 19 (including 6 dummy slices)
for MT-weighted images and 13 for reference images; interslice delay = 0 s for MT-weighted
images and 5 s for reference images; nominal scan time of 10 s for MT-weighted images and
scan time of *1.1 min for reference images. The off-resonance irradiation condition of the
interslice method was: pulse width = 1.2 ms; inter-pulse interval = 4.15 ms (*29% duty cycle);
and average RF power = 0.94 T.
For the images acquired with presaturation, both MT-weighted and reference images were
acquired with number of slices = 1, number of averages = 1, and a sufficient acquisition delay
time of 5 s to get rid of any residual signals prior to each measurement. A pulse train of 75
Gaussian pulses were used for off-resonance irradiation with the following saturation
condition: flip angle = 578.4; pulse width = 20 ms; inter-pulse interval = 40 ms (50% duty cycle);
total saturation duration = 3 s; average RF power = 0.94 T (equivalent to interslice method
above); off-resonance irradiation frequencies = +3200 Hz, +6400 Hz, and +9600 Hz.
Regions of interest (ROI) for the data were created manually for WM. Mean MTR values
were computed for the WM ROI for each subject, and the results were averaged across subjects.
Signal to noise ratio (SNR) was estimated from the difference image (IRef IMT) as the mean of
the signal in the WM ROI divided by the standard deviation of a large region in the difference
image containing only noise.
Effects of Varying Flip Angle and Phase Encoding Order
For six normal, healthy subjects (age 2439), MTR images were reconstructed from
MTweighted and reference bSSFP images acquired at varying flip angles from 15 to 90 in 15
intervals with descending slice order and positive slice-select gradient. Additional scan parameters
were as follows: TR/TE = 4.11/2.06 ms; matrix size = 128 128; field of view = 230 230 mm ;
slice thickness = 5 mm; interslice gap = 7 mm; scan direction = axial; PE direction =
anteriorposterior; initial dummy PE steps = 10/30 for linear/centric PE order; phase oversampling =
50%; number of averages = 1; RF-pulse BW = 1067 Hz; acquisition BW = 673 Hz/pixel; number
of slices = 15 for MT-weighted images (including 8 dummy slices) and 7 for reference images;
interslice delay = 0 s for MT-weighted images and 8 s for reference images. Scan time for
MTweighted images was 13 s and scan time for reference images was *1 min. Lastly, we measured
the observed T2 value of the center slice using a multi-contrast spin echo sequence with echo
times varying from 30 ms to 300 ms in 30 ms intervals.
Regions of interests for the data were created by automatic segmentation of GM and
WM via SPM8 software. Segmentation results were checked manually to ensure quality.
Mean MTR values were computed for the WM and GM ROIs as a function of flip angle for
each subject, and the results were averaged across subjects. Signal to noise ratio was estimated
from the difference image (IRef IMT) as the mean of the signal in the combined WM and GM
ROI divided by the standard deviation of a large region in the difference image containing
Comparison of bSSFP and SSFP-FID
For five normal, healthy subjects (ages 2449), MTR images were acquired for near full brain
coverage using a bSSFP sequence. Two of the subjects were also imaged with a SSFP-FID
sequence. Common acquisition parameters for both sequences were as follows: slice order =
descending; slice-select gradient polarity = positive; readout gradient polarity = positive; matrix
size = 256 256; field of view = 256 256 mm2; flip angle = 50; slice thickness = 3 mm;
interslice gap = 4.2 mm (0.6 mm after interleaving); scan direction = axial; PE direction =
anterior-posterior; dummy PE steps = 30; phase oversampling = 50%; number of averages = 1;
RF-pulse BW = 1600 Hz; acquisition BW = 501 Hz/pixel; number of slices = 19 and 18 for each
interleaved MT-weighted image set (including 6 dummy slices each; 12 total) and 25 for
reference images; interslice delay = 0 s for MT-weighted images and 6 s for reference images. For
the bSSFP sequence, TR/TE = 4.56/2.28 ms and total scan time = 4.36 min (1.17 min for
MT-weighted images and 3.18 min for reference images). For the SSFP-FID sequence, TR/TE
= 4.31/2.21 ms and total scan time = 4.24 min (1.10 min for MT-weighted images and 3.14 min
for reference images).
For each subject ROIs were created manually for WM. Mean MTR and SNR for the WM
ROI were estimated for each subject, and the results were averaged across subjects. The MTR
and SNR values were compared for the bSSFP and SSFP-FID sequences.
Interslice MTR Imaging of Meningioma
For a meningioma patient, MTR images were acquired with full brain coverage using
acquisition parameters: PE order = centric; slice order = descending; slice-select gradient polarity =
positive; readout gradient polarity = positive; TR/TE = 4.15/2.08 ms; matrix size = 128 128;
field of view = 220 220 mm2; flip angle = 60; slice thickness = 5 mm; interslice gap = 7 mm;
scan direction = axial; PE direction = right-left; initial dummy PE steps = 30; phase
oversampling = 0%; phase partial Fourier = 6/8; number of averages = 1; RF-pulse BW = 1333 Hz;
acquisition BW = 592 Hz/pixel; number of slices = 18 and 19 for each interleaved MT-weighted
image set (including 9 dummy slices each; 18 total) and 19 for reference images; interslice
delay = 0 s for MT-weighted images and 5 s for reference images. The total scan time was 2.1
min (0.4 min for MT-weighted images and 1.7 min for reference images).
For comparison with the proposed method, T2-weighted images were acquired using a Turbo
Spin Echo (TSE) sequence with imaging parameters: number of slices = 25; TR = 6000 ms; TE =
93 ms; echo train length = 18; matrix size = 640 520; field of view = 220 178 mm2; number of
acquisitions = 1; slice thickness = 5 mm; flip angle = 120; and scan time = *1.2 min.
Fig. 2 shows the resulting MTR images of the 10% agar phantom and saline phantom. The
MTR images of the agar phantom (left) showed relatively homogeneous MTR values with the
exception of dark spots caused by air bubbles trapped in the phantom. The saline phantom
(right), which was expected to have no MT effects, was free from extraneous signals (note the
scale difference between the two images). Together, the agar and control phantom images
strongly support MT effects as the source of image contrast.
Comparison of Interslice and Presaturation MT Effects
Fig. 3 shows the representative MTR images using the interslice method and using
conventional presaturation with a bSSFP readout for different offset irradiation frequencies corresponding
to the offset frequencies of the first, second, and third prior slices of the interslice method.
Visually, the signals in WM were higher than GM for all MTR images (Fig. 3a). The MTR images
acquired with presaturation showed decreasing MTR and SNR for increasing offset irradiation
frequencies. The average SNR and MTR values from the proposed interslice method were
similar to those with presaturation at an offset irradiation frequency corresponding to the first
prior slice of the interslice method (Fig. 3bc), indicating that contribution of the first prior
slice is dominant in the interslice MTR method and that the saturation efficiency of the
interslice method is comparable to conventional method.
Effects of Varying Flip Angle and Phase Encoding Order
Fig. 4 shows the center slice images for varying flip angle and for linear and centric PE from a
representative subject. Centric PE images showed better GM and WM contrast, suggesting that
relaxation effects influenced image contrast with linear PE.
Fig. 5a and 5b show the results of ROI analysis along with two-pool model simulations of
the acquisition protocol. Overall, MTR and SNR values increased with flip angle within the
tested range. Centric PE images showed higher MTR values and substantially higher SNR than
linear PE images (Fig. 5c). Simulated MTR values using tissue parameters from the literature
agreed well with the in vivo values. Substitution of the observed T2 value of 85 ms (from fitting
Fig 2. MTR images of a 10% agar phantom (left) and saline phantom (right).
Fig 3. Comparison of MTR images generated with interslice MT effects and with presaturation. Offset irradiation frequencies of the presaturation
pulses corresponded to the offeset frequencies of the first (+3200 Hz), second (+6400 Hz), and third (+9600 Hz) prior slices of the interslice method. Average
RF-power of saturation was equivalent in both methods. Baseline (MT-weighted) and MTR images (a) from a representative subject are shown. Both MTR (b)
and SNR (c) were calculated for white matter. Error bars show the 95% confidence interval of the group average.
the T2 map to a single-exponential function) for the literature value (69 ms ) produced
simulated MTR values in closer agreement with in vivo results.
Accumulation of MT Effects from Prior Slices
Fig. 6 shows simulations of the longitudinal magnetization (Mz, f) for WM and GM for varying
number of PE steps (RF-pulses) per slice, and two different flip angles. In agreement with the
results in Fig. 3, the majority of saturation was generated by the first prior slice; however, a few
Fig 4. Center slices of MTR images from a representative subject are shown for varying flip angles and for linear phase encoding (top) and centric
phase encoding (bottom).
Fig 5. Mean MTR values across subjects from regions of interest for white (a) and gray matter (b). Predicted values from simulating the two-pool model
(solid lines) with parameters from the literature show close agreement with the in vivo values. Centric phase encoding shows substantially better SNR (c)
than linear phase encoding. Error bars show the 95% confidence interval of the group average.
dummy slices are needed to account for the contributions of earlier (2nd, 3rd, etc.) prior slices
to reach a steady value of longitudinal magnetization across slices. More dummy slices are
needed for lower flip angles and for lower number of PE steps per slice. The value of
magnetization reached depended on the number of PE steps per slice and appeared to asymptotically
approach true steady-state MT effects (i.e., the state that would be achieved after an infinite
chain of saturation pulses at off-resonance frequency 1). For 128 PE steps, 56 dummy slices
should be included to reach steady MT effects. For 256 or more PE steps, 34 dummy slices
appeared to be sufficient. Including more dummy slices than needed would have minimal impact
on scan time (*13 s per dummy slice).
Effects of Varying Interslice Delay
Fig. 7 shows simulations with varying interslice delay time for acquisition of reference images.
In general, the MTR asymptotically increased with interslice delay time at rate that is
Fig 6. Saturation of the longitudinal magnetization accumulates over multiple prior slices with the majority of saturation due to the first prior slice.
For white (a) and gray (b) matter, simulations show the longitudinal magnetization as a function of the number of prior slices for varying number of phase
encoding steps per slice and for varying flip angles. One prior slice = saturation at 3840 Hz, two prior slices = saturation at 7680 Hz followed by 3840 Hz,
three prior slices = saturation at 11520 Hz, 7680 Hz, and 3840 Hz, etc.
Fig 7. Simulated MTR values for white (solid line) and gray matter (dashed line) for varying interslice
delay time for reference image acquisition. Simulations were performed using sequence parameters that
matched the bSSFP acquision for images in Fig. 8.
dependent on the T1 value of the tissue, with longer T1 values requiring a longer interslice
delay to recover. An interslice delay time of 34 s (rather than 8 s) for acquisition of reference
images would still maintain most of the MT contrast, indicating that the scan time for the
acquisition of reference images could be reduced accordingly (35 s per slice).
Comparison of bSSFP and SSFP-FID
Fig. 8 shows representative images using the interslice MTR imaging method with bSSFP and
SSFP-FID sequences. The SSFP-FID sequence is a non-balanced SSFP sequence that is less
Fig 8. Comparison of interslice MTR imaging with bSSFP and SSFP-FID sequences. The SSFP-FID sequence significantly reduced banding artifacts in
slices 35, but SNR was 22% lower than with bSSFP.
sensitive to banding artifacts in regions of high susceptibility. Regions near the sinuses in slices
35 were corrupted by banding artifacts with the bSSFP, but no artifacts were present in the
same slices acquired with the SSFP-FID sequence. The mean MTR values of WM were 32%
and 33% for bSSFP and SSFP-FID, respectively. The SNR in WM was 19.5 and 15.2 for bSSFP
and SSFP-FID respectively. Overall, the SSFP-FID sequence eliminated banding artifacts, but
reduced SNR by 22%.
Interslice MTR Imaging of Meningioma
Fig. 9 shows the MTR images acquired over the brain tumor region. The T2 images showed
higher signal in the tumor region compared to normal tissue. The MTR images from the
interslice method showed distinct signal characteristics in the brain tumor regions, different from
the T2 images.
In this study, we demonstrated the feasibility of using interslice MT effects to generate contrast
for MTR imaging. Furthermore, we validated the source of contrast as MT with phantoms and
by comparing interslice MT effects with contrast generated from conventional presaturation
pulses. We investigated effects of varying flip angle, number of PE steps, number of prior slices,
and interslice delay. We showed that banding artifacts could be reduced by using a
nonbalanced SSFP sequence with no modifications to the acquisition strategy. Finally, we
demonstrated the proposed method in a meningioma subject, in which the interslice MTR images
showed distinct contrast.
Interslice MTR Signal Characteristics
In the interslice method, data are acquired during the transient period of bSSFP acquisition.
The magnetization state at the start of acquisition of the imaging slice (partially saturated for
MT-weighted, relaxed for reference images) will move towards the steady state of the bSSFP
Fig 9. Interslice MTR images of a brain tumor (meningioma). Distinct signal characteristics in the MTR images were visible in the brain tumor regions.
readout, determined by the dynamics of bSSFP with contributions from on-resonance MT
effects. This explains why lower MTR values were seen with linear PE. With a sufficiently large
number of TRs before the acquisition of the center of k-space, no contrast is expected between
the MT-weighted and reference image acquisitions. Centric PE mitigates this issue by capturing
the MT contrast at the beginning of the acquisition in the low spatial frequencies. Despite this
limitation, we have validated the image contrast as MT effects in phantoms (Fig. 2), shown that
the contrast was predictable based on quantitative models of MT (Fig. 5), and generates distinct
contrast in preliminary imaging of a meningioma subject (Fig. 9).
The results in Fig. 3 showed that the contributions from the second and third prior slices
were minor compared to the first prior slice. The contributions are futher reduced in the
interslice method due to nominal delay times of *0.65 s and *1.3 s for the second and third prior
slices, respectively, in contrast to the first prior slice which has a nominal delay time of 0 s.
The results from imaging of the meningioma patient showed that MTR images can be acquired
in a relatively short period (e.g. *2.1 min) over the whole brain region using the proposed
interslice method. Assessment of brain tumor tissue is difficult, and the distinct MTR contrast
as shown in Fig. 9 may reflect unique metabolic information of the tumor region.
The interslice MTR method may offer some advantage in terms of SAR compared to other
methods, such as gradient echo methods, which rely on a separate pulse for saturation, or
onresonance MTR imaging [20, 21], which relies on a short RF-pulse duration to generate MT
effects. The SAR levels for the images acquired in Fig. 8 ranged from 3756% of the scanner limit
during acquisition of MT-weighted images with a 1 ms RF-pulse; however, we have shown that
high MTR values can be achieved with a longer 1.5 ms RF-pulse (Fig. 5). Since SAR levels have
a quadratic dependence on RF-pulse amplitude, the interslice MTR method may be optimized
for low SAR applications or high field applications by increasing the RF-pulse duration.
In the ALADDIN sequence, sequential multi-slice bSSFP acquisitions with alternating slice
order and slice-select gradient polarity, as well as alternating readout gradient polarity  are
used for interslice MT asymmetry  and perfusion imaging . The ALADDIN sequence
can be modified by adding acquisition of a set of reference images to calculate MTR images.
This would allow for simultaneous acquisition of four different image contrasts: baseline
bSSFP images (MT-weighted and reference), perfusion, MT asymmetry, and MTR images.
The MTR is considered a semi-quantitative measurement, since the value is dependent on
scan parameters. In quantitative MT imaging, the two-pool model is fit to multiple
MTweighted acquisitions, in which parameters such as bound pool fraction (F) and magnetization
exchange rate (kr) may offer better insight into tissue characteristics. Agreement of the data
and two-pool model simulations in this study suggest the possibility for developing
quantitative MT methods using interslice MT effects. Saturation power and off-resonance frequency
can be controlled simultaneously by changing excitation pulse duration, which is inversely
proportional to off-resonance frequency and inversely proportional to the square of saturation
power (W). The technique shows potential as a novel method for fast quantitative MT, since a
full set of MT-weighted images can be collected in less than 1 min.
Comparison with On-resonsance MTR Imaging with SSFP
Another SSFP based method for MTR has previously been developed using the difference in
on-resonance MT effects of long and short RF-pulse duration acquisitions [20, 21]. Thus, it is
worth briefly comparing the proposed method to the on-resonance MTR method. The
offresonance signal responses of bSSFP are periodic as a function of the off-resonance frequencies
and the TR of the bSSFP sequence. The on-resonance MTR method is based on the assumption
that both MT-weighted (short RF-pulse) and reference (long RF-pulse) images are acquired on
the pass-band region of the bSSFP off-resonance responses. This assumption seems reasonable
for normal brain regions; however, the assumption may not hold under pathological
conditions, such as brain tumors or hemorrhages that can cause high susceptibility effects. Because
of using different TR values between MT-weighted and MT free imaging, the on-resonance
MTR method also shows significantly reduced MTR in relatively short T2 components, as
demonstrated 40% reduction of MTR in 4% agar phantom with T1 = 1960 ms and T2 = 43 ms .
In contrast, the interslice MTR method uses the same scan parameters for acquisition of
MTweighted and references with only the addition of an interslice delay time, which does not affect
off-resonance responses of bSSFP and also provides high MTR values in relatively short T2
components as demonstrated in the 10% agar phantom with T1 = 1700 ms and T2 = 36 ms
(Fig. 2). The interslice method may potentially offer more reliable MTR measurements in
In terms of sensitivity, the 3D on-resonance MTR method can provide higher SNR because
of volumetric averaging effects. The proposed interslice method may be implemented in 3D as
multiple overlapping thin slab acquisition, which can improve spatial resolution and SNR and
requires further evaluation. Also, the availability of higher flip angle and longer data sampling
time due to no restrictions on RF duration and TR in the interslice MTR method can further
improves its MTR (as shown in Figs. 4 and 5) and SNR (lower acquisition BW), respectively.
These factors of the interslice method can partly compensate for the volumetric averaging
effects of the on-resonance MTR method.
Fig 10. Interslice MTR images with short interslice delay times. Matrix size = 128 128, FA = 60, TR/TE
= 4.15/2.08 ms, RF-pulse duration = 1.24 ms, slice thickness = 5 mm, number of slices = 24 (excluding 6
dummy slices), total scan time = 56 s (0.7 s delay) and 87 (2.0 s delay). Two scans (each with 12 number of
slices excluding 3 dummy slices) were spatially interleaved, in order to provide near whole brain coverage
with no gap.
In terms of scan time, both the interslice and on-resonance MTR methods require covering
the region of interest using bSSFP acquisitions twice (MT-weighting and MT-free). If the TR of
the interslice method (e.g. 4 ms) is the same as the average of the short and long TRs of the
onresonance method (e.g. 3 ms and 5 ms ), the time for actual data acquisition will be similar
between the two methods, except for the interslice delay required by the interslice method.
Data from a single subject (Fig. 10) showed preliminary evidence that good MTR maps can be
acquired with near whole brain coverage with very short interslice delay times (0.72.0 s),
which makes the total scan time of the interslice method (11.5 min) about 50% longer than
the on-resonance method, depending on in-plane matrix size. Note that while a short interslice
delay does not provide completely MT-free reference images, the long TR (5 ms) acquisition
for the on-resonance method also does not provide completely MT-free conditons. Scan time
and SNR are related to each other and thus should be systematically evaluated together, but
this is beyond the scope of the current study.
We demonstrated the feasibility of MTR imaging using interslice MT effects generated from
sequential multi-slice bSSFP acquisition. The technique provides a method for MTR imaging
without additional saturation pulses. Centric PE provided higher MTR values, higher SNR, and
better image contrast. Linear PE images showed image contrast influenced by relaxation effects.
Simulation of the two-pool model with parameters from the literature agreed well with in vivo
data and provided a useful tool for investigating the characteristics of interslice MT effects. The
new technique could provide MTR images covering the whole brain of a tumor patient within
a clinically feasible scan time of *2 min (or less with sequence optimization). Potential unique
applications include optimization for low SAR imaging and simultaneous MTR, MT
asymmetry, and perfusion imaging. Further work is needed to systematically compare the proposed
method with on-resonance MTR method and evaluate clinical usefulness of the
We would like to thank Dr. Chan-Hong Moon for providing the agar phantom.
Conceived and designed the experiments: JWB PKH SHC KTB SHP. Performed the
experiments: JWB PKH SHC SHP. Analyzed the data: JWB PKH SHP. Wrote the paper: JWB PKH
SHC KTB SHP.
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