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2552
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Magn Reson Med. 2021;85:2552–2567.
wileyonlinelibrary.com/journal/mrm
Received: 20 May 2020
|
Revised: 23 October 2020
|
Accepted: 25 October 2020
DOI: 10.1002/mrm.28602
FULL PAPER
3D Free-breathing multichannel absolute
B+
1
Mapping in the
human body at 7T
SebastianDietrich1
|
Christoph S.Aigner1
|
ChristophKolbitsch1
|
JohannesMayer1
|
JulianeLudwig1
|
SimonSchmidt2
|
TobiasSchaeffter1,3
|
SebastianSchmitter1,2,4
1Physikalisch-Technische Bundesanstalt (PTB), Braunschweig and Berlin, Germany
2Medical Physics in Radiology, German Cancer Research Center (DKFZ), Heidelberg, Germany
3Department of Medical Engineering, Technische Universität Berlin, Berlin, Germany
4Center for Magnetic Resonance Research, University of Minnesota, Minneapolis, Minnesota, USA
This is an open access article under the terms of the Creative Commons Attribution License, which permits use, distribution and reproduction in any medium, provided the original
work is properly cited.
© 2020 Physikalisch‐Technische Bundesanstalt. Magnetic Resonance in Medicine published by Wiley Periodicals LLC on behalf of International Society for Magnetic Resonance
in Medicine.
Parts of this work have been presented at the 2020 virtual annual meeting and the 2019 annual meeting of the International Society for Magnetic Resonance
in Medicine in Montreal, Canada.
Correspondence
Sebastian Schmitter, Physikalisch-
Technische Bundesanstalt (PTB),
Brauschweig and Berlin, Abbestraße 2–12,
10587 Berlin, Germany.
Email: sebastian.sc[email protected]
Funding informationThe German
Research Foundation (Grant Nos. SCHM
2677/2-1 and GRK2260-BIOQIC)
Purpose: To introduce and investigate a method for free-breathing three-dimensional
(3D)
B+
1
mapping of the human body at ultrahigh field (UHF), which can be used to
generate homogenous flip angle (FA) distributions in the human body at UHF.
Methods: A 3D relative B
+
1
mapping sequence with a radial phase-encoding (RPE)
k-space trajectory was developed and applied in 11 healthy subjects at 7T. An RPE-
based actual flip angle mapping method was applied with a dedicated
B+
1
shim setting
to calibrate the relative
B+
1
maps yielding absolute
B+
1
maps of the individual transmit
channels. The method was evaluated in a motion phantom and by multidimensional
in vivo measurements. Additionally, 3D gradient echo scans with and without static
phase-only
B+
1
shims were used to qualitatively validate
B+
1
shim predictions.
Results: The phantom validation revealed good agreement for
B+
1
maps between dynamic
measurement and static reference acquisition. The proposed 3D method was successfully
validated in vivo by comparing magnitude and phase distributions with a 2D Cartesian
reference. 3D
B+
1
maps free from visible motion artifacts were successfully acquired for
11 subjects with body mass indexes ranging from 19 kg/m2 to 34 kg/m2. 3D respiration-
resolved absolute
B+
1
maps indicated FA differences between inhalation and exhalation
up to 15% for one channel and up to 24% for combined channels for shallow breathing.
Conclusion: The proposed method provides respiration-resolved absolute 3D
B+
1
maps
of the human body at UHF, which enables the investigation and development of 3D
B+
1
shimming and parallel transmission methods to further enhance body imaging at UHF.
KEYWORDS
7 Tesla, actual flip angle imaging, body imaging,
B+
1
mapping, radial phase encoding, ultrahigh field
|
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DIETRICH ET al.
1
|
INTRODUCTION
One of the major challenges at ultrahigh magnetic field (UHF)
is the heterogeneous amplitude of the transmit (Tx) magnetic
field (
B+
1
). It yields a spatially heterogeneous flip angle (FA)
and, therefore, spatially varying contrast-limiting UHF ap-
plications. This issue has been addressed, among other solu-
tions, using multitransmit coils in combination with several
different parallel transmission (pTx) strategies.1,2
Several studies have also demonstrated that pTx methods
strongly improve signal-to-noise and contrast-to-noise ratios
when targeting the human body.3 Moreover, FA variations
in the human body are typically more pronounced than in
the human brain because of the larger and patient-depen-
dent body size. Static pTx methods are often sufficient to
achieve acceptable FA homogeneity for small two-dimen-
sional (2D) regions of interest (ROIs) such as used for the
prostate,4 whereas the complexity of the pTx strategies needs
to increase when large organs such as the liver5 are targeted.
Therefore, it is expected that an extension of the ROI to cover
a larger 3D volume in the human body also requires more
complex FA optimization strategies.
The knowledge of the underlying
B+
1
maps of each single
Tx channel is a prerequisite for any pTx optimization, inde-
pendent of the target region. Such maps are either acquired
for each subject at the beginning of each session or in case of
universal pulses6 during a set of training subjects, which then
eliminates the need for subject-specific
B+
1
mapping. Several
different techniques for acquiring (absolute)
B+
1
maps exist,
among them are methods based on the Bloch-Siegert shift,7
presaturation turbo fast low-angle shot,8 dual refocusing
echo acquisition (DREAM),9,10 and actual flip angle imag-
ing (AFI).11 A straightforward approach to obtain individual
B+
1
maps of each Tx channel consists of combining a sin-
gle absolute
B+
1
map obtained by one of the aforementioned
methods7–9,11 with a set of multiple Tx gradient echo (GRE)
acquisitions, where, for example, only a single Tx channel is
active per acquisition.12 Similarly, absolute 2D
B+
1
maps can
also be acquired based on the B1TIAMO method.13
For obtaining absolute
B+
1
maps in the human body at
7T of multiple 2D slices or 3D volumes, however, the afore-
mentioned methods are often limited by radiofrequency (RF)
power demand, motion, and/or blood flow and their acqui-
sition times might be too long for a breath-hold. Instead, a
method for generating relative
B+
1
estimations14 of the
NTx
Tx channels, initially introduced for the human brain, can be
acquired without high power requirements based on GRE im-
ages with low FAs. This allows the acquisition of
B+
1
maps of
a single slice within a few seconds, that is, within a breath-
hold when applied to the body. However, a limitation of this
method is that the amplitude of the maps is relative, if no ad-
ditional absolute
B+
1
map is acquired to calibrate these maps,
and that the maps are biased by the square-root of the proton
density.14 Despite these limitations, the relative
B+
1
mapping
method has proven to be highly suitable for FA optimization
in the human body at 7T, not only for
B+
1
shimming pur-
poses,15,16 but also for slice-selective pTx methods such as
spokes RF pulses.5,17,18 For pTx techniques that optimize the
FA within 3D volumes such as the whole liver or heart, how-
ever,
B+
1
maps covering the 3D volume are required. Although
the relative
B+
1
mapping technique mentioned previously has
been extended to multiple slices,17,19 full 3D mapping cannot
be achieved within a single breath-hold.
Therefore, we investigate in this work a method to acquire
free-breathing, 3D absolute
B+
1
maps in the human body at
7T.20 This technique is based on a fast, low-power–demanding
relative B
+
1
mapping method,14 as well as a radial phase-
encoding (RPE) acquisition scheme,21–23 which enables ret-
rospective binning in respiratory motion states. This flexible
3D acquisition scheme is applied for two acquisitions: (1) a
modified AFI calibration scan to acquire absolute
B+
1
maps
in a combined, single Tx configuration24; and (2) small FA
spoiled GRE acquisitions. We first use the GRE scans to
derive relative
B+
1
maps of the human body that are subse-
quently calibrated using the AFI scan resulting in estimated,
absolute
B+
1
maps of each channel and different respiratory
motion states. We quantitatively evaluate the method first in
a motion phantom scan and subsequently show the suitability
of this technique to compute
B+
1
maps in 11 volunteers. 2D
Cartesian cine GRE scans, as well as 3D GRE scans with and
without static phase-only
B+
1
shims are used to qualitatively
validate the resulting
B+
1
maps.
2
|
METHODS
Our method to compute multichannel 3D respiration-re-
solved, absolute
B+
1
estimations relies on two separate acquisi-
tions: an RPE-based, 3D respiration-resolved AFI (RPE-AFI)
acquisition,24 and a Tx channel-wise RPE-based23,25 3D
respiration-resolved-gradient echo (Tx-RPE-GRE) acquisi-
tion. An overview of the workflow is depicted in Figure 1A.
Relative
B+
1
maps are calculated from Tx-RPE-GRE and cali-
brated with RPE-AFI to obtain absolute
B+
1
map estimations.
To obtain a sufficiently high FA for the calibration, the RPE-
AFI is measured with a
B+
1
shim setting Φeff, providing a lo-
cally high Tx efficiency in the frontal region of the heart.
Calibration is then performed for ROIs in the heart with a FA
> 40° (see the white ROI in Figure 1A).
2.1
|
Radial phase encoding and
self-navigation
Using an RPE-based acquisition, respiration-resolved im-
ages are obtained by acquiring the 3D k-space as illustrated
2554
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DIETRICH ET al.
in Figure 1B. k-Space is sampled on a Cartesian grid along
the readout direction (kx), whereas the two phase-encoding
(PE) directions (ky, kz) are sampled on a non-Cartesian, radial
grid. After completing the
NPE
readouts along one radial line,
the next radial line is rotated by a golden-angle increment of
111.24°. The resulting RPE trajectory covered the PE plane ho-
mogenously and allowed retrospective binning of the data into
p=(1,...,Np)
respiratory motion states using self-navigation.
The
Np
undersampled respiration-resolved k-space data sets
are subsequently reconstructed using nonuniform fast Fourier
transformation (NUFFT) based on an iterative sensitivity-
encoding (SENSE) algorithm.23,26,27 A respiratory motion sur-
rogate for self-navigation is retrieved from the 1D projection
in the head–feet direction for ky = kz = 0.23 The temporal sam-
pling intervals for self-navigation and the number of readouts
per radial line are listed in Table 1 for the different acquisitions.
FIGURE 1 A, Workflow of proposed
mapping approach that uses acquired AFI data (with white lines as coil locations) (1) and calculated
magnitude of the sum (MOS) of estimated channel-wise
B+
1
for efficient shim setting (2,3) to calculate calibration factorλ (4) resulting in channel-
wise estimated, absolute
B+
1
maps (5). RPE trajectory with Cartesian data acquisition along readout direction kx on a radial grid in the phase-
encoding plane ky – kz. Between successive radial lines, the angle is increased by the golden angle, k-space data are retrospectively binned into
different respiratory motion states with the help of a motion surrogate. B, Each bin is reconstructed and RPE-GRE data for different respiratory
motion states can be acquired. Data are taken from subject S4. RPE-AFI, radial phase-encoding-actual flip angle imaging; Tx-RPE-GRE, transmit-
radial phase-encoding-gradient echo
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|
2555
DIETRICH ET al.
2.2
|
3D Multichannel Tx-RPE-GRE and
reconstruction of relative
B+
1
map
In the Tx-RPE-GRE acquisition, the RPE scheme is ap-
plied to a small tip angle GRE sequence, which consisted of
NTx +2
repetitive RPE-GRE scans. During each scan, only
a single Tx channel is active while the signal is received
through all
m=(1,...,NRx )
Rx elements.14,28 The ac-
tive Tx channel number
k=(1,...,NTx )
was incremented
through the scans. In addition to the
NTx
scans, we acquired
another scan with all
NTx
channels and one with no Tx chan-
nel active for transmission. The latter scan is used for noise
decorrelation of the receive channels used during image
reconstruction.
The magnitude of the resulting 3D respiration-resolved
image
|Ik,m,p(
r
)|
of Tx channel
k
, Rx channel
m,
and respira-
tory phase
p
for each spatial position
r
was then given by the
simplified signal equation:
with the complex receive profile
B
m,p(
r
)
, the complex
transmission profile
B+
k,p
(r
)
, the proton density
𝜌0(
r
)
, and
the spatially independent, complex scaling factor
c
includ-
ing RF pulse, hardware, and reconstruction-related scaling
factors. Consequently, the sum of all magnitude images
N
Rx
m
N
Tx
k
I
k,m,p
(r)
can be computed across Tx and Rx
(1)
|||
Ik,m,p(r)
|
|
|
=c
|
|
|
B
m,p(r)
|
|
|
|
|
|
B+
k,p(r)
|
|
|
𝜌0(r)
,
In vivo RPE–AFI Tx-RPE–GRE RPE–GRE Tx-GRE
TR1 [TR2] (ms) 10 [50] 5 3.7 5.2
TE (ms) 2.02 2.02 1.75 2.87
TA (min) 6.10 3.40 5.55 0.25
Oversampling 2.91
Nominal FA (degree) 90 20 25 15
Reference voltage (V) 170
FOV (mm3) 250 × 312–350 × 312–350 384 × 384 × 4
Voxel size (mm3) 4 × 4 × 4 4 × 4 × 4 1.4 × 1.4 × 1.4 4 × 4 × 4
BW (kHz) 25.54 25.54 172.48 48.96
Self-navigation TS
(ms)
480 80 162
Radial lines 384 256 2048
Slices 3
Slice gap (mm) 20
Phantom RPE–AFI Tx-RPE–GRE
TR1 [TR2] (ms) 10 [50] 4.26
TE (ms) 1.9 1.9
TA (min) 6.10 4.36
Oversampling 0.5
Nominal FA (degree) 60 5
Reference voltage (V) 30 30
FOV (mm3) 200 × 150 ×
150
200 × 150 × 150
Voxel size (mm3) 3.1 × 3.1 ×
3.1
3.1 × 3.1 × 3.1
BW (kHz) 25.54 25.54
Self-navigation TS
(ms)
480 140
Radial lines 384 192
AFI, actual flip angle imaging; BW, readout bandwidth; FA, flip angle; FOV, field of view; GRE, gradient
echo; RPE, radial phase-encoding; TA, acquisition time; TE, echo time; TR, repetition time; TS, sampling
time; Tx, transmit.
TABLE 1 Sequence parameter of
in vivo and phantom studies with TR,
TE, TA, oversampling as a percentage of
phase-encoding points additionally acquired
compared with a fully sampled Cartesian
scan with the same FOV and voxel size,
FA, BW, and self-navigation TS of k-space
center
2556
|
DIETRICH ET al.
channels. Assuming that the sum of magnitudes (SOM) of all
Tx channels is equal to the SOM of all Rx channels, which
has been observed to fit well for transceiver coils,14 then:
This assumption introduced a bias into the result-
ing
B+
1
maps that is proportional to the square root of the
proton density and symbolized by the hat. Estimated,
relative
B+
1
magnitude maps were obtained for each Tx
channel
k
and respiratory motion state
p
by weighting the
expression (Equation 2) by relative image signal intensities
R
k,p=
N
Rx
mIk,m,p(r)
N
Rx
m
N
Tx
k
Ik,m,p(r)
14:
Relative Tx phase maps
𝜙k,p(
r
)
of Tx channel
k
relative to
one fixed Tx channel
kf
are calculated based on the complex
images by:
Here,
denotes the complex conjugate and arg is the
phase of the complex value. The resulting estimated
B+
1
maps
were finally given by:
Two different sets of
B+
k,p
(r
)
maps are derived. A set of
nonrespiration-resolved
B+
1
maps
B+
k
(r
)
with
Np=1
is recon-
structed during the actual imaging session using all acquired
data without any binning. A second set of respiration-resolved
B+
1
estimations
B+
k,p
(r
)
with
Np=3
is obtained retrospectively
after each session because of reconstruction time limitations.
The nonrespiration-resolved
B+
k
(r
)
maps obtained with
the default transmission phase
ϕ0
are used to compute two
static phase-only
B+
1
shim settings based on multiple manu-
ally drawn ROIs covering multiple slices. The first shim set-
ting is computed for three transversal-slice ROIs covering the
human heart, by following the heart–lung border, to obtain
a homogenous
B+
1
shim
ϕhom
. The central slice Sc is placed
in the center of the heart, along the head–feet direction; the
other two slices are placed above and below Sc. The second
shim setting is computed for a ROIeff with a diameter of 2 to
3 cm, in a single transversal slice to obtain an RF-efficient
B+
1
shim
ϕeff
in an ROI covering the anterior section of the heart
and myocardium. The efficiency is computed as in Metzger
et al4 and aimed to maximize the constructive interference of
the
NTx
B+
1
maps in the target region.
The phase setting
ϕhom
is applied to achieve homoge-
neous FAs in subsequently acquired high-resolution 3D
RPE-GRE acquisitions, as well as 2D Cartesian cine GRE
scans;
ϕeff
is applied for an RPE-AFI mapping as outlined
below.
2.3
|
3D RPE-AFI
An RPE-GRE acquisition scheme was modified to acquire
two interleaved repetition times (TRs; RPE-AFI),24 which
enables the computation of absolute, respiration-resolved
B+
1
maps according to the AFI approach.11 The RPE-AFI data is
acquired in vivo with
B+
1
shim setting
ϕef
. Each readout of
the RPE k-space trajectory is consecutively sampled twice
with different TR (ratio
n=TR2TR1
) in an interleaved
fashion, which results in
Np
respiration-resolved image data
sets
A1,p(
r
)
and
A2,p(
r
)
. The FA
𝛼p(
r
)
for each motion state
p
and spatial position
r
is computed following Yarnykh.11
Before the FA calculation,
A1,p(
r
)
and
A2,p(
r
)
is median
filtered within a three-voxel neighborhood. The obtained FA
𝛼p(
r
)
in degree is converted to absolute
B+
1
maps, termed
B+AFI
p
(r)
,
which are given in μT or related to unit power in
μT/
kW.
2.4
|
Reconstruction of unbiased and
estimated, absolute multichannel
B+
1
maps
Unbiased, respiration-resolved, absolute
B+
1
maps of each Tx
channel are obtained following Van de Moortele et al,12 and
by combining the RPE-AFI with the images obtained by the
Tx-RPE-GRE acquisition. This approach, however, requires
a sufficiently high FA of >20° throughout the body, which
is not achievable in the center of the body with our set-up
because of RF power limitations.
To circumvent this, a different approach is used as already
suggested in Van de Moortele et al14 for brain applications by
calculating a single scaling factor
λ
based on the RPE-AFI
acquisition in a manually drawn
ROIλ
with a sufficiently high
FA of >40° (as shown in Figure 1A).
(2)
NTx
k
B+
k,p(r)
=
NRx
m
NTx
k
Ik,m,p(r)
c𝜌0(r)
(3)
B+
k,p(r)
=Rk,p(r)
NTx
k
B+
k,p(r)
=
N
Rx
mIk,m,p(r)
c𝜌0(r)
NRx
m
NTx
k
Ik,m,p(r)
𝜙
k,p(r)=arg
[N
Rx
m
Ik,m,p(r)I
kf,m,p(r)
].
(4)
B+
k,p(r)=
|||
B
+
k,p(r)
|||
ei𝜙k,p(r)
.
(5)
𝜆
=
B+AFI
p
NTx
k
B+
k,p(r)ei𝜙eff
ROIλ
.
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|
2557
DIETRICH ET al.
To achieve this calibration, the same efficient
B+
1
shim
ϕeff
is retrospectively applied to the
B+
k,p
(r
)
maps. The scaling
factor
𝜆
is subsequently applied to the relative
B+
k,p
(r
)
maps
(Equation 4) to yield estimated, absolute B
+
1
maps, termed
B+,𝜆
k,p
(r
)
in the following:
2.5
|
Experiments
All scans were obtained on a 7T system (Magnetom 7T;
Siemens Healthcare, Erlangen, Germany) equipped with
a pTx system and 8 × 1 kW amplifiers (Stolberg AG,
Stolberg, Germany). The phantom study was performed
using an in-house-built 8 Tx/8 Rx transceiver head coil.
The in vivo measurements were performed with a com-
mercial body coil array (MRI Tools, Berlin, Germany),
consisting of 32 transceiver elements (eight dipoles and 24
loops) that are driven in an 8 Tx/32 Rx channel mode. The
body coil has a field-of-view (FOV)/excitation of approxi-
mately 240 mm along the head–feet direction. It consists of
an anterior and a posterior half with fixed, solid housings.
The anterior and posterior halves have four parallel aligned
blocks with four elements each: one dipole and three loop
elements. The three loop elements are distributed along the
head–feet direction and each loop element covers approxi-
mately one-third of the dipole. In the Tx case, the blocks
of one dipole and three loop elements are combined with
fixed phases to form one Tx channel, respectively. The
body coil was certified by a notified body to comply with
the local specific absorption rate (SAR) limits in a first-
level controlled mode of 20 W/kg (IEC 60601-2-33) that
were ensured by limiting the RF power for each Tx channel
to the body coil. The conversion from RF power to a local
SAR was estimated based on electromagnetic simulations
with 10 million random-phase settings. The worst-case
shim combined with an RF power of 11.8 W applied to
each Tx channel resulted in a 10 g-averaged peak spatial
SAR of 11.1 W/kg; thus, it left an additional safety margin
factor of 1.8. Although such limits are then independent of
the applied Tx phase setting, the approach provides rather
conservative power limits.
To validate the respiration-resolved
B+
1
maps, a cylindrical
motion phantom (radius = 12.5 cm, height = 12.0 cm) filled
with water, 2% agaroses, 0.2% sodium chloride, 0.1% copper
sulfate (which corresponds to σphantom
=0.4sm
,
εr=80
) was
used that performed a translational sinusoidal-like motion
along the bore axis with a peak-to-peak amplitude of 3 cm and
a frequency of 0.1 Hz. To perform this motion, the phantom
was placed on a wagon driven by a piston that was attached to
a flywheel. The rotation of the flywheel was controlled by a
cord that was wound on the flywheel. This cord was unwound
by a step motor located in the operator room. Tx-RPE-GRE
and RPE-AFI data sets were acquired during the translational
movement of the phantom, referred to as “dynamic” acquisi-
tion. For comparison, RPE-AFI and Tx-RPE-GRE data were
acquired for two extreme motion states of the phantom mim-
icking an inhale and exhale state, referred to as “static” acqui-
sition. Sequence parameters are listed in Table 1. RPE-AFI
and Tx-RPE-GRE data were low-pass-filtered before NUFFT
reconstruction. For binning, self-navigation was used to re-
construct data into three different respiratory motion states
using an iterative SENSE algorithm.23,26,27
2.6
|
In vivo acquisitions
Eleven healthy subjects (age:
30 ±5
years, max/min: 35/21
years; body mass index [BMI]:
24 ±5
kg/m2, max/min:
34/19 kg/m2; male/female: 8/3) were included in this study
after obtaining ethical approval of the study by PTB's local
institutional review board and after the subjects provided
written informed consent.
All subjects underwent the same protocol, which started
with the coil's default
B+
1
shim and no additional Tx phase. The
default shim is set by the manufacturer as a trade-off between
B+
1
power efficiency, SAR efficiency, and avoidance of voids
throughout the entire ascending aorta, descending aorta, and
the heart. This default shim is used only as a starting point
because the maps and the resulting excitation pattern vary
within subjects. In this configuration, a Tx-RPE-GRE was ac-
quired under free-breathing; for comparison, a Cartesian mul-
tichannel GRE sequence (2D-Tx-GRE) was applied to obtain
three equidistant transversal 2D slices without cardiac gating
within one breath-hold. Additionally, a high-resolution RPE-
based anatomical GRE acquisition (RPE-GRE) was obtained
with default
B+
1
shim. After calculation of
ϕeff
based on the
B+
k,p
(r
)
maps generated by Tx-RPE-GRE reconstructed with
Np=1
and Equation 4, an RPE-AFI was obtained with phase
setting
ϕeff
. Here,
TR1=10ms, TR2=50ms
was used, thus at
the lower end within the range recommended by Yarnykh,11
which reduced acquisition time and allowed a k-space
center update of 0.48 seconds needed for sampling the self-
navigation signal. Additionally,
B+
1
mapping was performed
under a deep-breathing condition in two healthy subjects (S10
and S11). In these two cases, the data were reconstructed into
Np=
5 respiratory motion states, which required doubling the
acquisition time of both RPE-AFI and Tx-RPE-GRE with
otherwise same acquisition parameters. Thus,
p=5
indicates
the end inspiration motion state for deep breathing;
p=3
in-
dicates the end inspiration motion state for shallow breath-
ing. All RPE acquisitions were carried out with a rectangular
pulse with a pulse duration of 0.5 ms.
(6)
B+,𝜆
k,p
(r)=𝜆
B
+
k,p
(r)
.
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|
DIETRICH ET al.
To demonstrate the shimming results, an additional 3D
RPE-GRE sequence, with 1.4 mm isotropic resolution was
acquired under free-breathing without cardiac triggering or
gating, and 2D cine GRE acquisitions in transverse orientation
were performed with a homogenous
B+
1
shim setting
ϕhom
, with
acquisition parameters listed in Table 1. In one subject (S8), ad-
ditional 2D cine GRE scans were acquired during breath-holds
in four-chamber, long-axis and short-axis views with a TR =
4.4 ms, echo time = 2.0 ms, acquisition time = 17 s, FOV =
384 × 272 × 5 mm3, voxel-size = 1.3 × 1.3 × 5 mm3, nominal
FA = 50°, generalized autocalibrating partially parallel acquisi-
tions (GRAPPA) factor = 2 with 24 reference lines, retrospec-
tive electrocardiogram (ECG) gating, and 30-ms reconstructed
temporal resolution yielding 46 cardiac phases.
2.7
|
Image reconstruction
The Tx-RPE-GRE–based data were reconstructed during the
in vivo data acquisition on a separate computer (12 central
processing units (CPUs) with 2.1 GHz and 128 GB of ran-
dom-access memory (RAM)). To reconstruct nonrespiration-
resolved RPE-GRE data, which have been used to calculate
the
B+
k
(r
)
maps used online during the same scanning ses-
sion, a NUFFT was applied.26 The respiration-resolved data
sets were reconstructed after the scanning session on a sep-
arate reconstruction computer with 24 CPUs and 256 GB
RAM. Undersampled RPE data was reconstructed for
Np=3
respiration-resolved states using a NUFFT iterative SENSE
reconstruction.27 Respiratory binning was performed using a
self-navigation approach23 and a sliding window with a 20%
overlap for each side of the bin. The 3D RPE-GRE sequence
obtained for the shimming demonstration was reconstructed
using iterative SENSE reconstruction, resulting in 3D im-
ages in four different respiratory motion states. Subsequently,
a single set of high-resolution, respiratory motion-corrected
RPE-GRE images IMOCO was obtained based on the motion
fields from the four states. To this end, nonrigid motion fields
were estimated using NIFTYREG29 according to an algorithm
outlined in Cruz et al21 and Kolbitsch et al.30
2.8
|
Data analysis
Quantitative evaluation was performed for an ROI comput-
ing the mean efficiency and the coefficient of variation (CV)
following Metzger et al4 and Schmitter et al.28 The mean ef-
ficiency was defined as
with
NTx
k
B+,𝜆
k,p(r)ei𝜙k
being the magnitude of the sum
(MOS) and
NTx
k
B+,𝜆
k,p(r)
being the SOM of a channel-wise
Tx field
B+
k,p
(r
)
for a phase settings
ϕk
.31 The CV was defined as
with standard deviation SD. Moreover, pixel-wise differences
between
B+
1
estimations were computed to quantify the ac-
curacy of the proposed
B+
1
mapping approach. The differ-
ence between static and dynamic phantom experiments was
analyzed based on the MOS of static and dynamic
B+,𝜆
k
(r
)
maps in a Bland-Altman plot along with ±1.96 SD and mean
values over the whole volume. Nonrespiration-resolved and
respiration-resolved mean
B+,𝜆
k,p
distributions were analyzed
in line plots, showing a mean value of three slices. The mag-
nitude and phase differences between 2D
B+,𝜆
k
(r
)
maps and
3D
B
+,𝜆
k
(r
)
maps were analyzed in line plots for a matched
slice and manually selected lines.
All results were masked based on thresholded low-pass
filtered magnitude images with a kernel size of five voxels.
For Tx-RPE-GRE, the SOM of Tx channels was used.
3
|
RESULTS
In this section we show the acquisition and reconstruction of
respiration-resolved 3D absolute
B+
1
estimations of a moving
phantom and the human thorax at 7T for 11 subjects with a
wide range of BMIs and body dimensions. The figures il-
lustrate the changes compared with previous techniques ap-
plied in the body: the extension from 2D to 3D, the ability
to provide respiration-resolved maps, and the change from
relative to absolute maps. The validity of the
B+
1
maps is
shown by a static phase-only
B+
1
shim and successive cine
GRE acquisitions.
Figure 2 depicts motion-resolved
B+,λ
k,p
maps, each ob-
tained from threefold undersampled k-space data in the
motion phantom experiments from the dynamic scan in com-
parison with the static reference obtained from fully sampled
data. Depicted are the magnitude and phase images of two
motion states (mimicking inhale and exhale) for qualitative
comparison and a Bland-Altman plot of the difference be-
tween the static and dynamic MOS of
B+,λ
k,p
maps for inhale
and exhale, respectively. Bland-Altman plots show the vox-
el-wise difference for the whole phantom between the MOS
of the dynamic
B
+,𝜆
k,p
(r
)
and static
B+,𝜆
k
(r
)
measurement with
absolute (relative) mean difference
B+
1
=1.2 ±
2.6
μT/
kW
η=
mean
NTx
k
B+,𝜆
k,p(r)ei𝜙k
NTx
k
B+,𝜆
k,p(r)
ROI
,
CV
=
SD
NTx
k
B+,𝜆
k,p(r)ei𝜙k
mean
NTx
k
B+,𝜆
k,p(r)ei𝜙k
ROI
,
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|
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DIETRICH ET al.
(3% ± 6%) in the exhale position and a mean difference
B+
1
=1.2 ±
2.9
μT/
kW
(3 ± 6%) in the inhale position.
Background voxels were excluded from the analysis.
Figure 3 illustrates the measured
B+,λ
k,p
of subject S4
in transversal and sagittal view for Tx channel
k=1, 3
(Figure 3A,B). Differences in
B+,λ
k,p
between exhale and inhale
of up to 15% and between nonrespiration-resolved and exhale/
inhale of up to 10% and 7%,respectively, have been found.
Figure 4A shows channel-wise 2D Cartesian and 3D non-
respiration-resolved RPE
B+,λ
k,p
maps for each Tx channel in
a transversal view of the isocenter in subject S8. All 3D
B+,λ
k,p
maps show comparable results in terms of magnitude and
phase distribution compared with the 2D reference images;
none of the channels shows motion artifacts. Consistent re-
sults were found for all subjects. Quantitative comparison
was carried out with line plots for cross sections through
the heart of the central slice for Tx channel
k=2, 4
(Figure
4B,C) for phase and magnitude of 2D Cartesian (red) and 3D
RPE-based
B+,λ
k,p
(black).
Figure 5 illustrates the 3D absolute
B+,AFI
p
maps (RPE-AFI)
acquired with shim setting Φeff and absolute
B+,λ
k
maps acquired
with shim setting Φ0 that have been combined retrospectively
with shim setting Φeff. There is an excellent match between
both maps in the heart region, thus demonstrating the validity
of combining relative and absolute maps and to calculate the
calibration factor
λ
used to compute the absolute
B+,λ
k
maps.
The resulting CV and efficiency η for all volunteers are
shown in Figure 6, with values calculated in a region covering
three transversal slices of the heart as can be seen in Figure 7.
Across all subjects, the variations in the ROI could be substan-
tially reduced from CVpre = 42.5%
±
6.2% to CVpost = 20.6%
±
2.0% for Φhom. The mean efficiency was improved from ηpre =
53.2%
±
11.2% to ηpost = 90.0%
±
7.5% for Φeff.
Figure 7 depicts a qualitative comparison of
B+,λ
k
maps
and RPE-GRE images obtained for two different shim set-
tings for one coronal and three transversal views. Figure 7A
shows the results using the default phase setting Φ0, resulting
in signal cancellations in the human heart. Importantly, the
B+
1
maps qualitatively match the experimental RPE-GRE im-
ages and predict the underlying
B+
1
artifacts denoted by the
yellow arrows. Figure 7B shows the results after the homo-
geneous
B+
1
shim setting Φhom. The CV in these acquisitions
FIGURE 2 A, Comparison of magnitude and phase of estimated, absolute
B+
1
maps with dynamic data with exhale and inhale motion states,
and data from static acquisition for exhale and inhale position, with inhale position indicated by a solid line and exhale position by a dashed line,
with a peak-to-peak difference of 30 mm. All transmit channels of the dynamic scan show qualitatively comparable results in terms of magnitude
and phase distribution compared with the static reference images. B, Bland-Altman plot from MOS of B
+
1
maps with voxel-wise difference of static
to dynamic acquisition as a function of voxel-wise estimated mean
map is shown for inhale and exhale position of the phantom. With a mean
difference of
= 1.1 ± 2.9 μT/√kW for inhale position and mean difference
B+
1
= 1.3 ± 2.6 μT/√kW for exhale position. MOS, magnitude of
the sum
+1.96 SD: 6.2
mean: 1.3
-1.96 SD:-3.7
+1.96 SD: 6.7
mean: 1.1
-1.96 SD:-4.5
(A)
(B)
2560
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DIETRICH ET al.
was reduced from 42.9% (Φ0) to 19.8% (Φhom), which is re-
flected by the improved homogeneity.
In Figure 8, respiration-resolved 3D RPE-GRE images in
comparison with the MOS of respiration-resolved 3D
B+,λ
p
maps are featured. The mean free-breathing difference of the
position of the right hemidiaphragm between inhales and ex-
hales was 12.1 ± 2.6 mm. Cross sections of the
B+
1
distribu-
tion through the heart in coronal view (indicated by the solid
white line) for exhale and inhale show
B+,𝜆
p,
a (relative) dif-
ference up to 0.21 μT (24%), with mean
B+,𝜆
p
of 0.74 μT, be-
tween both respiratory motion states. The respiration-resolved
B+,𝜆
p
maps are free from motion artifacts and match the ex-
perimental data.
B+,λ
k
magnitude and phase distribution for Tx channel
1 are illustrated in three different respiratory motion states
p for exhale, inhale, and the nonrespiration-resolved re-
construction for subject S10 performing deep breathing in
Figure 9. No motion artifacts are visible in the respiration-
resolved images, and a clear difference between the
B+,λ
k
distributions for the exhale and inhale motion state is visi-
ble. The spatial displacement of the right hemidiaphragm be-
tween inhale and exhale was approximately 30 mm.
4
|
DISCUSSION
The presented respiration-resolved, absolute multichan-
nel
B+
1
mapping method is based on a relative
B+
1
mapping
method,14 providing proton-density relative 2D
B+
1
maps in
the human head. That technique has also been applied to the
human body at UHF to derive relative 2D
B+
1
maps within a
single slice or multiple slices to calculate B
+
1
shimming solu-
tions. Furthermore, it has been applied to optimize the
B+
1
field at 7T within various organs in the human body includ-
ing the aorta and the kidney, as well as the human heart.15,16,32
In addition, such a method has been used to calculate and
apply multispoke pulses in the human heart and the liver at
7T,5,33,34 as well as at 10.5T.18 The present method extends
this aforementioned technique with three modifications: (1)
by extending it to 3D mapping in the human body, (2) by
providing respiration-resolved maps, and (3) by enhancing
the 3D AFI technique to allow for generating respiration-
resolved absolute
B+
1
maps in the human body. The latter is
used for calibration of the channel-wise maps. 3D mapping
of the Tx field is highly beneficial when
B+
1
shim solutions or
pTx pulses are calculated for an entire 3D volume of interest.
The resulting maps provide the flexibility to either optimize
the B
+
1
field for a single slice or multiple slices in (multislice)
2D imaging or for the entire volume in 3D imaging.1
The ability to obtain
B+
1
maps with calibrated magnitude
is important for many applications, as it allows us to adjust
the absolute FA value. In this work, relative maps obtained
from the Tx-RPE-GRE acquisition were combined with an
RPE-AFI acquisition that provides respiration-resolved
B+
1
maps. Because this method uses GRE acquisitions obtained
in the linear, small FA regime to calculate the relative
B+
1
maps, the technique does not require high power levels.
FIGURE 3 Comparison of magnitude distribution between nonrespiration-resolved (NRR) exhale (P = 1) and inhale (P = 3) motion states for
transmit channel k = 1 (A) and k = 3 (B) in transversal and sagittal view for subject S4. Positions of one-dimensional
B+
1
curves for NRR exhale
(
P=1
) and inhale (
P=3
) motion states are indicated by solid white lines for left–right (i), anterior–posterior (ii), and head–feet (iii) direction
(A)
(B)
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|
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DIETRICH ET al.
FIGURE 4 A, Magnitude and phase from two-dimensional (2D) and three-dimensional (3D) estimated, absolute
B+
1
maps for each transmit
channel k for isocenter of subject S8. Vertical (dotted) and horizontal (solid) profiles of phase and magnitude for 2D
B+
1
estimation (red) and 3D
B+
1
estimation (black) of transmit channel k = 2 (B) and k = 4 (C). The solid and dashed line in phase images indicates the position of line plot data
(A)
(B)
(C)
FIGURE 5 In vivo RPE-AFI maps
B+,AFI
1
measured with Φeff of the human body with an estimation of
B+
1
shim setting Φeff from MOS of
estimated, absolute
B+
1
for subject S4. Images correspond to three cross sections taken from a three-dimensional volume. MOS, magnitude of the
sum; RPE-AFI, radial phase-encoding-actual flip angle imaging
2562
|
DIETRICH ET al.
Nevertheless, the gradient of the individual
B+
1
profiles
from the body surface towards the center of the body makes
it challenging to stay within the linear regime close to the
coil, while still providing sufficient SNR in the more dis-
tant body regions. On the one hand, we observed decreased
SNR in the posterior part of the heart in some volunteers.
On the other hand, T1 effects may bias the resulting
B+
1
maps close to the skin, particularly when short TR values
are being used to keep the scan time short as is the case
for cardiac
B+
1
mapping. This effect can be seen in the 3D
map in Figure 4 that used a 25% higher FA compared with
the 2D acquisition. Although the T1 bias could be avoided
by replacing the method with alternative techniques35,36
that provide T1 unbiased relative
B+
1
maps, such methods
may increase the acquisition time multifold. Furthermore,
a general challenge that is associated with the body size is
FIGURE 6 Calculated values of
the coefficient of variation (CV) and
the efficiency η of the
B+
1
shim settings
evaluated in hand-drawn regions of interest
around the hearts of all 11 subjects
FIGURE 7 A, Estimated, absolute
distribution for default shim setting Φ0 in coronal view, with the positions of transversal slices indicated
by white lines. The white circles mark the regions of interest (ROIs) for homogenous shim calculation as well as coefficient of variation (CV).
Corresponding respiratory motion-compensated high-resolution RPE-GRE images IMOCO are illustrated below for matching coronal and transversal
slices obtained with Φ0, with CV = 42.7% and η = 42.6%. B, Estimation of
B+
1
shim setting Φhom optimized for homogeneity in the ROI with CV =
19.5% and efficiency η = 51.7% for coronal and transversal view, as well as corresponding slices for IMOCO obtained with Φhom. The yellow arrows
denote the signal dropouts in (A) obtained with Φ0 that are recovered in (B) by applying Φhom. Corresponding images are taken from subject S3.
MOS, magnitude of the sum; RPE-GRE, radial phase-encoding-gradient echo
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|
2563
DIETRICH ET al.
given by regions of low FA in which the AFI method does
not yield reliable values. This problem could be avoided by
acquiring two or more AFI maps with different shim set-
tings so that at each point within the region at least one of
the shim settings provides sufficient
B+
1
. The latter option
is also exploited by the TIAMO method.37 In principle, a
similar technique, as proposed in Van de Moortele et al12
for the human brain, could be extended to the human body.
In that approach, individual, absolute
B+
1
maps of each
Tx channel are obtained by weighting relative Tx channel
maps with one absolute
B+
1
map acquired by transmitting
with all Tx channels combined with a given
B+
1
shim set-
ting. This approach, however, was not feasible because it
requires a well-defined FA map, ie, a FA > 20° throughout
the entire 3D volume. In this work, a well-defined FA map
could not be achieved in the center of the body given the
power limits of the coil. We refrained from extending the
duration of the rectangular RF pulse length because initial
investigations showed altered FA values.
The AFI method could be replaced by other, less SAR
demanding methods such as the DREAM technique,9 which
can provide very fast (ie, subsecond) absolute
B+
1
maps of
the human body. Furthermore, the method has also been used
to show respiration-induced
B+
1
changes in the human body
at 3T.38 However, the DREAM mapping technique is very
sensitive to flow-induced artifacts,39 which has also been
confirmed by our own experiments, and it has a limited dy-
namic FA range.9 Less flow/motion sensitivity has been ob-
served by AFI. To estimate a potential FA error in the left
ventricle and the myocardium (ie, within the ROI used for FA
calibration), separate experiments in a flow phantom were
conducted at 7T with a constant flow velocity of 55 cm/s,
which corresponds to maximum intracardiac blood veloci-
ties40 (see Supporting Information Figure S1). An average FA
error of 15% was found compared with flow-off conditions.
Furthermore, the impact of myocardial motion was estimated
by extended phase graphs, which resulted in an estimated
maximum error of <5%. The Bloch-Siegert shift mapping
FIGURE 8 RPE–GRE image in two respiratory states (exhale/inhale) with the corresponding MOS of estimated, absolute
B+
1
distribution for
subject S4 for
B+
1
shim setting Φhom. The yellow dashed line indicates the exhale position of the right hemidiaphragm. The solid white line denotes
the position of line plot data. Images correspond to two cross sections taken from a three-dimensional volume. MOS, magnitude of the sum; RPE-
GRE, radial phase-encoding-gradient echo
2564
|
DIETRICH ET al.
technique, however, could not be applied in the body at 7T
because of SAR constraints.
Therefore, the approach of calibrating the relative respira-
tion-resolved
B+
1
maps was employed in this work. Despite its
benefits, the method has limitations. As reflected in Equation 2,
the resulting maps are biased by the square-root of the proton
density. The impact on the resulting
B+
1
maps, however, is
expected to induce only minor tissue-dependent variations of
a few percentages. Furthermore, the technique is biased by
the assumption used to obtain Equation 2 (ie, the SOM of
the receive signal equals the SOM of the Tx signal), which
is particularly the case for the present coil, where groups of
four coil elements were combined to a single Tx element.
According to electromagnetic field simulations obtained with
the Duke model41 and performed for the applied 8 Tx/32 Rx
body-coil configuration, this assumption introduces an error
of 10% on average throughout the thorax. The error could be
reduced to 5% if the RF coil could be driven with 32 Tx and 32
Rx channels, as shown in Supporting Information Figure S2.
Nevertheless, despite those limitations, a high similarity was
found in qualitative comparisons between the excitation pat-
terns and the resulting images, which is in strong agreement
with previous work where the same technique was applied
slice selectively in 2D.5,15
The resulting absolute
B+
1
maps obtained by the proposed
method enabled successful
B+
1
phase shimming as has been
shown for cardiac imaging. The shim allowed for acquir-
ing 2D cine GRE acquisitions (see Supporting Information
Figure S3) in different views under breath-hold, as well as
3D respiration-resolved and respiration-corrected RPE-
based GRE imaging of the entire thorax obtained under
free-breathing, and in both cases, the image quality was
substantially improved compared with the default shim.
The magnitude of the absolute
B+
1
maps was used to cali-
brate the nominal FA for each sequence accordingly. In two
subjects, we observed phase singularities in the
B+
1
maps
of Tx channel 1, which was used as the reference chan-
nel. Based on calculating the phase difference, this effect
translated into the
B+
1
maps of the other seven Tx channels.
Although this effect could be avoided by changing the ref-
erence channel (see Supporting Information Figure S4), we
did not observe any impact on the resulting
B+
1
phase shim
magnitude; therefore, channel 1 was set as the reference Tx
channel.
Our initial results found that a shim setting calculated
on
B+
1
maps in a respiratory state
p
performs better than a
shim calculated on nonrespiration-resolved
B+
1
maps for im-
ages acquired in motion state
p
. Furthermore, it seems that
for shallow breathing patterns such differences in
B+
1
and CV
are rather minor42; thus, a respiration-state–resolved recon-
struction may not be needed. However, our initial results in-
dicate that the effect becomes stronger when deep breathing
FIGURE 9 Estimated, absolute
B+
1
magnitude and phase distribution for
transmit channel 1 in three different
respiratory motion states: p for exhale (P =
1), inhale (P = 5), and the nonrespiration-
resolved reconstruction for subject S10.
No motion artifacts are visible in the
respiration-resolved images, and a clear
difference between the
B+,λ
k
distribution
for exhale and inhale motion state is
visible. The deep-breathing difference of
the position of the right hemidiaphragm
between inhale and exhale was
approximately 30 mm
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|
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DIETRICH ET al.
is performed, which is in agreement with a recent study.43
However, more-detailed investigations are necessary to be
able to provide a more comprehensive description of the un-
derlying effects.
Both the scan time and the reconstruction time are lim-
itations of this method. Although a nonrespiration-resolved
reconstruction requires 2 minutes, and therefore has been
used during the session, the retrospective binning and recon-
struction of three phases requires 15 minutes to reconstruct,
which would have further prolonged the MR examination.
Importantly, image acquisition times of 3 minutes and 24 sec-
onds for the Tx-RPE-GRE and 6 minutes for the RPE-AFI
are too long, particularly for future patient studies. With an
increasing number of Tx channels,44,45 the image acquisition
needs to be accelerated to reduce total scan times, which could
be achieved by decreasing the spatial resolution, increasing the
undersampling factor, or introducing different k-space sam-
pling strategies.46 However, such steps might affect the quality
of respiration-resolved
B+
1
maps. For brain studies, so-called
universal pulses have been proposed,6 which calculate pTx RF
pulses based on training data sets in several subjects that are
successfully applied to subjects not included in the training
pool. Although applying this approach straightforwardly to the
body might be challenging because of variable body sizes, this
step, combined with a minimum of information on
B+
1
and/
or anatomy of the present subject,47 may achieve the desired
result at a minimum of additional calibration time.
Therefore, the present work provides a valuable basis for
performing 3D imaging of the human body without signal
dropouts and lays the foundation for future developments
necessary for body imaging at 7T.
ACKNOWLEDGMENTS
The authors gratefully acknowledge funding from the German
Research Foundation (SCHM 2677/2-1 and GRK2260-
BIOQIC). We thank André Kühne and Helmar Waiczies
(MRITools GmbH, Berlin, Germany) for discussions and
support with respect to the RF coil.
ORCID
Sebastian Dietrich https://orcid.
org/0000-0002-1610-909X
Christoph S. Aigner https://orcid.
org/0000-0003-3618-9610
Christoph Kolbitsch https://orcid.
org/0000-0002-4355-8368
Johannes Mayer https://orcid.org/0000-0002-2500-445X
Juliane Ludwig https://orcid.org/0000-0003-4042-8071
Simon Schmidt https://orcid.org/0000-0003-1835-4002
Tobias Schaeffter https://orcid.
org/0000-0003-1310-2631
Sebastian Schmitter https://orcid.
org/0000-0003-4410-6790
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SUPPORTING INFORMATION
Additional Supporting Information may be found online in
the Supporting Information section.
FIGURE S1 low impact on flip angle measurement in a
strongly varying B
+
1
field is investigated with an agarose flow
phantom. This phantom contains a pipe which is connected
to a flow pump (CardioFlow 5000 MR, Shelley Medical
Imaging Technologies, London, Canada) and driven with a
constant flow. Shown are the resulting flip angle for a flow
velocity of 0 cm/s (A) and 55 cm/s (B) and the correspond-
ing gradient echo image (C), with a blue arrow indicating the
flow direction. The relative difference RD of both measure-
ments a and b is shown in (D), for an ROI covering the pipe
a mean RD of 15% is calculated
FIGURE S2 Three different sets of
B+
1
/
B
1
maps were
available for analysis: 32 Tx and 32 Rx fields, 8 Tx and
32 Rx fields, which corresponds to the setup used in this
work, and 8 Tx and 8 Rx channels. For each of the three
cases, we compared the sum of magnitude (SOM) across
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|
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DIETRICH ET al.
the channels of simulated
B
1
, simulated
B+
1
and estimated
B+
1
, which was obtained using Equation 3, with synthetic
small flip angle gradient-echo-like images
I
k,m=B
+
k
B
m
.
The latter corresponds to the Tx-RPE-GRE acquisition.
Deviations between the simulated
B+
1
and the resulting es-
timated
B+
1
maps are quantified by the relative difference
RD of both(see Supporting Information Material M1).
Resulting maps are shown for the central transversal slice
through the heart are shown. Qualitatively the largest dif-
ference between the SOM is visible for the 8Tx/8Rx setup,
with an increasing number of Tx and Rx channels the SOM
maps become more homogenous which is also reflected in
the relative difference
B+
1
maps. Quantitively the standard
deviation of RD
B+
1
calculated for the whole thorax is 18%
for 8Tx/8Rx, 10% for 8Tx/32Rx, and 5% for 32Tx/32Rx.
Thus, using a 32Tx/32Rx coil setup would be ideal, but our
system has only 8 Tx channels, and therefore, we had to use
an 8Tx/32Rx configuration that still provides better results
as compared to an 8Tx/8Rx set
FIGURE S3 Two representative time frames in three dif-
ferent views capturing systole and diastole of subject S8 are
shown. Note that the same static phase-only shim Φhom, has
been used for all three different orientations. Despite using
the same shim setting, all three views are free of major
B+
1
re-
lated artifacts allowing to capture the whole heart movement.
The corresponding cine GRE can be found in Supporting
Information Video S1
FIGURE S4 Estimated, absolute
B+
1
maps of all the 11 sub-
jects for all Tx channels are shown. Despite large variations
in BMI (max/min: 34/19 kg/m²) the estimated, absolute
B+
1
maps are plausible and show no visible motion artifacts.
Transversal slice of all 11 subjects with magnitude and phase
for all eight Tx channels
VIDEO S1 Cine GRE of subject S8 in three different orien-
tations with 1.3 mm in-plane resolution, 5mm slab, and 46
cardiac phases
How to cite this article: Dietrich S, Aigner CS,
Kolbitsch C, et al. 3D Free-breathing multichannel
absolute
B+
1
Mapping in the human body at 7T. Magn
Reson Med. 2021;85:2552–2567. https://doi.
org/10.1002/mrm.28602