Exploration of New Contrasts, Targets, and MR Imaging and Spectroscopy Techniques for Neuromuscular Disease – A Workshop Report of Working Group 3 of the Biomedicine and Molecular Biosciences COST Action BM1304 MYO-MRI

Abstract

Neuromuscular diseases are characterized by progressive muscle degeneration and muscle weakness resulting in functional disabilities. While each of these diseases is individually rare, they are common as a group, and a large majority lacks effective treatment with fully market approved drugs. Magnetic resonance imaging and spectroscopy techniques (MRI and MRS) are showing increasing promise as an outcome measure in clinical trials for these diseases. In 2013, the European Union funded the COST (co-operation in science and technology) action BM1304 called MYO-MRI (www.myo-mri.eu), with the overall aim to advance novel MRI and MRS techniques for both diagnosis and quantitative monitoring of neuromuscular diseases through sharing of expertise and data, joint development of protocols, opportunities for young researchers and creation of an online atlas of muscle MRI and MRS. In this report, the topics that were discussed in the framework of working group 3, which had the objective to: Explore new contrasts, new targets and new imaging techniques for NMD are described. The report is written by the scientists who attended the meetings and presented their data. An overview is given on the different contrasts that MRI can generate and their application, clinical needs and desired readouts, and emerging methods.

Full text

Journal of Neuromuscular Diseases 6 (2019) 1–30
DOI 10.3233/JND-180333
IOS Press
1
Review
Exploration of New Contrasts, Targets,
and MR Imaging and Spectroscopy
Techniques for Neuromuscular Disease –
A Workshop Report of Working
Group 3 of the Biomedicine and
Molecular Biosciences COST
Action BM1304 MYO-MRI
Gustav J. Strijkersa,∗, Ericky C.A. Araujob, Noura Azzaboub, David Bendahanc,
Andrew Blamired, Jedrek Burakiewicze, Pierre G. Carlierb, Bruce Damonf,
Xeni Deligiannig, Martijn Froelingh, Arend Heerschapi, Kieren G. Hollingsworthd,
Melissa T. Hooijmansa, Dimitrios C. Karampinosj, George Loudosk, Guillaume Madelinl,
Benjamin Martyb, Armin M. Nagelm, Aart J. Nederveena, Jules L. Nelissena,
Francesco Santinig, Olivier Scheideggern, Fritz Schicko, Christopher Sinclairp,
Ralph Sinkusq, Paulo L. de Sousar, Volker Straubd, Glenn Waltersand Hermien E. Kane
aAmsterdam UMC, University of Amsterdam, Amsterdam, The Netherlands
bNMR Laboratory, Neuromuscular Investigation Center, Institute of Myology & NMR Laboratory,
CEA/DRF/IBFJ/MIRCen, Paris, France
cAix Marseille University, Marseille, France
dInstitute of Cellular Medicine, Newcastle University, Newcastle-upon-Tyne, UK
eDepartment of Radiology, Leiden University Medical Center, Leiden, the Netherlands
fVanderbilt University Medical Center, Nashville, USA
gDepartment of Radiology, Division of Radiological Physics, University Hospital Basel, Basel,
Switzerland & Department of Biomedical Engineering, University of Basel, Basel, Switzerland
hUniversity Medical Center Utrecht, Utrecht, the Netherlands
iRadboud University Medical Center, Nijmegen, the Netherlands
jTechnical University of Munich, Munich, Germany
kBioEimmision Technology Solutions, Athens, Greece
lNew York University School of Medicine, New York, USA
∗Correspondence to: Gustav J. Strijkers, Amsterdam UMC,
University of Amsterdam, Amsterdam, the Netherlands. E-mail:
g.j.strijkers@amc.uva.nl.
ISSN 2214-3599/19/$35.00 © 2019 – IOS Press and the authors. All rights reserved
This article is published online with Open Access and distributed under the terms of the Creative Commons Attribution Non-Commercial License (CC BY-NC 4.0).

2G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
mInstitute of Radiology, University Hospital Erlangen, Friedrich-Alexander-Universit¨at
Erlangen-N¨urnberg (FAU), Erlangen, Germany & Division of Medical Physics in Radiology,
German Cancer Research Center (DKFZ), Heidelberg, Germany
nDepartment of Neurology, Inselspital, Bern University Hospital, University of Bern, Switzerland
oUniversity of T¨ubingen, Section on Experimental Radiology, T¨ubingen, Germany
pUCL Institute of Neurology, London, UK
qKing’s College London, London, UK
rUniversit´e de Strasbourg, Strasbourg, France
sUniversity of Florida, Gainesville, USA
Abstract.Neuromuscular diseases are characterized by progressive muscle degeneration and muscle weakness resulting in
functional disabilities. While each of these diseases is individually rare, they are common as a group, and a large majority
lacks effective treatment with fully market approved drugs. Magnetic resonance imaging and spectroscopy techniques (MRI
and MRS) are showing increasing promise as an outcome measure in clinical trials for these diseases. In 2013, the European
Union funded the COST (co-operation in science and technology) action BM1304 called MYO-MRI (www.myo-mri.eu), with
the overall aim to advance novel MRI and MRS techniques for both diagnosis and quantitative monitoring of neuromuscular
diseases through sharing of expertise and data, joint development of protocols, opportunities for young researchers and
creation of an online atlas of muscle MRI and MRS. In this report, the topics that were discussed in the framework of working
group 3, which had the objective to: Explore new contrasts, new targets and new imaging techniques for NMD are described.
The report is written by the scientists who attended the meetings and presented their data. An overview is given on the
different contrasts that MRI can generate and their application, clinical needs and desired readouts, and emerging methods.
Keywords: Neuromuscular disease, muscle, MRI, MRS, biomarker, outcome measure, myo-mri
ABBREVIATIONS
ADC apparent diffusion coefficient
ADP adenosine diphosphate
ATP adenosine triphosphate
bSSFP balanced steady state free precession
cCSA contractile cross sectional area
CK creatine kinase
COST cooperation in science and technology
CS corticosteroid/compressed sensing
(in the context of MR image
reconstruction methods)
CSA cross sectional area
CT computed tomography
CV coefficient of variation
DENSE displacement encoding with
stimulated echoes
DESS double echo steady state
DMD Duchenne muscular dystrophy
DWI diffusion weighted imaging
DTI diffusion tensor imaging
EPG extended phase graph
18F-FDG 18 F-fluorodeoxyglucose
FH foot-head
FSHD facioscapulohumeral dystrophy
GCL gastrocnemius lateral head
GCM gastrocnemius medial head
Gr gracilis
GRAPPA generalized autocalibrating partial
parallel acquisition
GRMD golden retriever muscular dystrophy
IDEAL iterative decomposition of water and
fat with echo asymmetry and
least-squares estimation
IR inversion recovery
ISIS image-selected in vivo spectroscopy
NMD neuromuscular diseases
MIRACLE motion-insensitive rapid configuration
relaxometry
MOLLI modified look-locker inversion
recovery
MR magnetic resonance
MRE magnetic resonance elastography
MRI magnetic resonance imaging
MRS magnetic resonance spectroscopy
MSE multi-spin-echo
MT magnetization transfer
MTR magnetization transfer ratio
NBD nemo-binding-domain
PC phase contrast
PCr phosphocreatine
PDE phosphodiesters
PDFF proton density fat fraction
PER peroneus

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 3
PET positron emission tomography
Pi inorganic phosphate
PME phosphomonoesters
PS pulsed saturation
PSR pool size ratio
pSSFP partially spoiled steady state free
precession
qMT quantitative magnetization transfer
RF rectus femoris/radio frequency (in the
context of MRI excitation pulses and
signal acquisition)
SIR selective inversion recovery
sLASER semi-localized by adiabatic selective
refocusing
SMA spinal muscular atrophy
SNR signal-to-noise ratio
SOL soleus
Sr Sartorius
SSFP steady state free precession
STEAM stimulated echo acquisition mode
STIR short-TI inversion recovery
TA tibialis anterior
TESS triple echo steady state
TIRM turbo inversion recovery magnitude
TP tibialis posterior
TSE turbo spin echo
UTE ultra-short echo time
VI vastus intermedius
VL vastus lateralis
VM vastus medialis.
INTRODUCTION
Neuromuscular diseases (NMD) form a large
group of individually rare diseases that are present
in all populations and affect people of all ages. They
are often, particularly at chronic stages, character-
ized by progressive muscle degeneration and muscle
weakness resulting in functional disabilities. Many of
them result in chronic disability, which poses a signif-
icant healthcare burden for society, and most of them
lack an effective therapy. Even though NMDs have
very different causes and pathogenic mechanisms,
fibrosis, edema, and fat replacement are frequently
observed histological features. Diagnosis and ther-
apy development for NMD has rapidly expanded in
recent years [1–3] and there is an urgent need to
develop objective, non-invasive outcome measures
to monitor disease progression and treatment effect
[4, 5]. Muscle biopsies have been used extensively
to classify NMD and to gain a better understand-
ing of their underlying pathomechanisms. However,
because they are invasive, it is undesirable to repeat
them often. Also, they only assess a small sam-
ple in a single muscle, leading to non-representative
results and making them less suitable as outcome
measures for clinical trials. Functional measures are
often used in clinical trials as the primary outcome,
but most of them heavily rely on patient cooperation
and motivation, and are therefore inherently variable
and subjective.
The use of magnetic resonance imaging and spec-
troscopy techniques (MRI and MRS) applied to NMD
is showing increasing promise as an outcome measure
in clinical trials [6, 7]. Unfortunately, progress has
been hindered by the rarity of individual NMDs and
lack of options for pooling data from different groups.
Therefore, in 2013, the European Union funded the
cooperation in science and technology (COST) action
BM1304 called MYO-MRI (www.myo-mri.eu). The
overall aim of MYO-MRI was to advance novel MRI
and MRS techniques for both diagnosis and quanti-
tative monitoring of neuromuscular diseases through
sharing of expertise and data, joint development of
protocols, opportunities for young researchers and
creation of an online atlas of muscle MRI and MRS.
There were four working groups in the action, which
all held two working group meetings a year, where
data were shared and discussed in an open and infor-
mal atmosphere. In this report, the topics that were
discussed in the framework of working group 3,
which had the objective to explore new contrasts,
new targets and new imaging techniques for NMD
are described.
This report is written by scientists who attended the
meetings and presented their data. The report starts
with the different contrasts that MRI can generate,
and describes their recent developments as applied to
skeletal muscle. Clinically,the two most used contrast
mechanisms are those that determine fat infiltration
and from that muscular fraction (area), in the form
of T1-weighted imaging, and those that use T2relax-
ation to characterize a combination of various states
of water in muscles (T2-weighted imaging with fat
suppression, or inversion-recovery based short-tau
inversion recovery (STIR) or turbo inversion recovery
magnitude (TIRM) sequences). As the focus of the
working group 3 meetings was on new contrasts, T1-
and T2-weighted imaging will not be discussed, and
the report will focus on the more experimental imag-
ing sequences that can generate quantitative image
contrasts. This report will describe the different

4G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
contrasts and their applications, followed by a section
on the most pressing clinical needs, with an overview
of applied and desired readouts for NMD, as well
as combinations of MR techniques. In addition, an
overview is given of emerging methods, including
new developments in post-processing technology and
acceleration techniques. The report ends with sum-
marizing perspectives.
CONTRASTS
Dixon imaging
The increased number of potential treatments for
NMDs results in a higher need for a reliable and
sensitive biomarkers which would allow objective
quantification of disease progression [8–13]. Since
one of the main characteristics of chronic NMDs is
fatty infiltration of muscle tissue, a method quantify-
ing fat content would be a natural choice for such
a biomarker. A technique commonly employed in
MRI to measure fat content is the Dixon method
[14–16]. It utilizes the difference in resonance fre-
quencies of water and fat to produce separate maps
of water and fat distribution. While solely water or
fat images remain qualitative, the combination of the
two becomes quantitative by giving the relative con-
tent of the two species, generally defined through a
fat fraction FF = f/(w+f), where wand fare water and
fat signals, respectively.
The Dixon technique is well-established, but many
variants exist, and one of the objectives of working
group 3 was to reach a consensus on the optimal
approach to apply it in the muscles. In particular a
number of confounding factors needs to be addressed
to make the method independent of the protocol
and system where the data is acquired, and some
trade-offs need to be addressed. First, to properly
compensate for main field inhomogeneities, it is rec-
ommended that at least 3 echoes (Dixon points) are
acquired [15, 16]. The sequence parameters need
to be chosen so that the sequence is minimally
T1-weighted, since water and fat tissue have very dif-
ferent T1values. This can be avoided by using long
repetition times and low flip angles [17, 18]. Also, the
complexity of the fat spectrum needs attention – basic
implementation of Dixon reconstruction assumes a
single spectral peak, however around 30% of the fat
protons resonate at frequencies different from that
of the main peak [19–21]. T2* relaxation effects
need to be accounted for. This can be included in
the reconstruction model, but often requires acquisi-
tion of additional echoes to be stable [22]. Moreover,
noise bias from magnitude correction can influence
the measurements, so a post-processing correction
method is advised, particularly if the signal-to-ratio
(SNR) of the images is low [17]. Finally, phase data
from the scanner needs to be reliable – this can be
confounded by bipolar readouts, eddy currents, and
gradient delays. Using monopolar readout sequences
helps to mitigate some of these, otherwise modeling
is needed [23–25].
If these confounders are accounted for, the
Dixon technique is considered to produce a setup-
independent measurement of fat fraction, generally
referred to as proton density fat fraction (PDFF), and
working group 3 published a consensus paper detail-
ing what is necessary to achieve this goal, as well as
listing important NMD studies to date utilizing the
Dixon technique [26].
T2Relaxation time mapping
The transverse relaxation time (T2) is one of the
main variables that determine the MR signal inten-
sity. It describes the nuclear spin-spin interactions and
its value highly depends on the molecular tumbling
rate. T2relaxation time mapping can be performed
for the T2of the water signal (commonly referred
to as water-T2or MRS-T2), the fat signals or of all
signals combined (commonly referred to global-T2
or MRI-T2). Muscle water T2variations provide rel-
evant information about disease activity and muscle
physiological status but are highly nonspecific. They
may indeed reflect inflammatory processes, myocyte
swelling, sarcoplasmic leakiness, cell necrosis, den-
ervation, or simply hydrostatic edema. More about
T2as a biomarker for inflammation in NMD can be
found in section 3.1.
The standard way to measure T2consists of
adjusting an exponential model to the signal decay
measured with a multi-spin-echo (MSE) sequence.
Quantitative relaxometry in fat infiltrated muscles
should be interpreted with care though since water
and fat protons T2largely differ. Because of the long
T2of fat compared to the T2of normal muscle tissue,
a single-component analysis of non-fat-suppressed
signal results in a global T2that primarily reflects
the fat content in tissues and may hide underlying
alterations of the muscle tissue itself [27]. Different
approaches were proposed to decompose the global
signal decay into a sum of fat and water contribu-
tions. A bi-exponential fit was introduced by Yao

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 5
et al. and combined with a Dixon acquisition to deter-
mine water T2[28]. More recently, a tri-exponential
model was introduced, where the fat signal decay was
approximated by a bi-exponential (based on the anal-
ysis of subcutaneous fat) and the muscle tissue decay
by a single exponential and resulted in the simul-
taneous estimation of water T2and the fat fraction
[29]. However, in practice, the signal resulting from
an MSE sequence rarely displays a pure T2decay. It is
sensitive to instrumental imperfections like non-ideal
slice pulse profiles, B1+inhomogeneities, and insuf-
ficient crushing schemes. This leads to the generation
of T1-weighted stimulated echoes that contaminate
the signal decay curves [30]. This can be resolved
applying more complex models based either on Bloch
equations [31] or extended phase graph (EPG) anal-
ysis [30]. Recently, a bi-component EPG approach
was proposed to simultaneously quantify the mus-
cle water T2and fat fraction from a standard MSE
acquisition (Fig. 1) [32].
Alternatively, a method that exploits the chemi-
cal shift between fat and water [33, 34] has been
proposed, combining an IDEAL (iterative decom-
position of water and fat with echo asymmetry and
A
120 ms
0
B
C
Fig. 1. Bi-component extended phase graph (EPG) approach to simultaneously quantify the muscle water T2and fat fraction. (A) Water T2
map. (B) Fat image. (C) Water image. Figure adapted from Marty et al. with permission [32].

6G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
least-squares estimation) decomposition of water and
fat signals with an MSE sequence was also specifi-
cally designed for simultaneous estimation of water
T2and fat fraction.
Several rapid T2-mapping sequences based on
2D- or 3D-SSFP sequences have been developed for
musculoskeletal imaging, including bSSFP [35, 36],
pSSFP [37, 38], DESS [39], TESS [40]. These meth-
ods are generally sensitive to B1+inhomogeneities
and flip angle profiles, and some of them show, in
addition, some sensitivity to motion. Recently a rapid
B1+,B
0, and motion-insensitive 3D-SSFP relaxom-
etry approach, MIRACLE [41], was presented. It
allows accurate T1and T2mapping within one sin-
gle scan, but was to date only applied in the brain.
To mitigate the effects of B1+inhomogeneities on
muscle water T2quantification T2-prepared 3D-TSE
sequences employing an adiabatic T2preparation
have been also proposed [42, 43].
Although rarely mentioned, the T2-relaxation of
muscle water is multi-exponential [44], as revealed by
spectroscopic studies. At least three T2components
can be observed: a short one (T2< 5 ms) attributed
to hydration water, and two others around 30 and
150 ms, with respective relative fractions of about
90% and 5%, reflecting the water compartmenta-
tion into intracellular, interstitial, and vascular spaces
[45]. On standard single-exponent T2-mapping meth-
ods the signal is dominated by the longer T2
components. Multicomponent T2decay analysis has
the potential of determining compartment-specific
T2values, relative volumes, and compartmental
exchange rates.
Diffusion-weighted imaging
Diffusion-weighted imaging (DWI) has been
widely applied in various medical and non-medical
fields, ranging from solid material, phantom and
animal experiments to human tissue studies. DWI
uses diffusion-sensitizing pulsed-field gradients to
assess displacement of water protons in vitro and
in vivo [46]. This has motivated its use to probe
microstructural information beyond the resolution of
conventional MR imaging, i.e. in the range of several
10’s of micrometers. Depending on the pulsed-field
diffusion gradient scheme and strategy of analysis of
the acquired DWI data, different parameters can be
estimated from the investigated object: the apparent
diffusion coefficient (ADC) and the directional dif-
fusivity [47], the displacement probability [48], the
sizes of the diffusion-restricting boundaries, e.g. pore
or cell diameter [49], and the transition probability
between compartments delimited by semi-permeable
boundaries [50–52].
When applying DWI in a clinical setting, espe-
cially the analysis of the directional diffusivity and
the time-dependent analysis of DWI parameters have
shown to be a promising tool for the non-invasive
investigation of human skeletal muscle architec-
ture [53, 54]. However, its use can be hampered
by the time-consuming acquisitions, as measure-
ment time is directly proportional to the number
of diffusion-sensitizing gradient directions, the steps
of diffusion weighting (the number of b-values),
and for specific sequences also the diffusion time
(time between diffusion gradients). The ADC can be
estimated easily and fast using one image without
diffusion weighting and three images with orthog-
onal gradient directions in order to get the trace
image. Prior studies in pathological human skele-
tal muscle have demonstrated a change in ADC
due to muscle denervation (24% increase) [55], in
inflammatory myopathies (17% increase) [56], and
deformation-induced injury (between 16% increase
and 10% decrease, depending on time after injury)
[57] – hence the ADC could serve as a biomarker
in muscular disorders. However, analysis of ADC
in resting healthy human muscles has yielded a
broad range of normal values between 1.26 ×10–3
and 1.99 ×10–3mm2/s [54–56, 58–61], i.e. a 58%
difference between the lowest and the highest val-
ues. Additionally, changes in muscle ADC are rather
unspecific. Changes in ADC and directional diffusiv-
ity in healthy human muscles have previously been
reported to be dependent on exercise [62, 63], train-
ing condition [64], active muscle contraction [65],
and passive joint position [66, 67]. Recent systematic
evaluations of DWI in healthy human skeletal mus-
cle have furthermore described the DWI parameters
to be dependent on different technical issues dur-
ing acquisition and post-processing, such as spatial
resolution, diffusion-encoding parameters, signal-to-
noise ratio, and phase-sensitive variations due to
microcirculation [68–71]. Fatty infiltration of mus-
cle tissue is a hallmark of NMD and confounds DWI
measurements [72, 73]. Hence new acquisition meth-
ods for robust fat suppression need to be used in such
cases [74, 75]. Careful planning of MR examinations
can take the DWI parameter dependences mentioned
above into account, but the variability of values in
resting healthy human muscle means that interpret-
ing DWI changes in pathological conditions remains
challenging.

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 7
Magnetization transfer
Magnetization transfer (MT) MRI is sensitive to
the relative sizes of, and exchange between, the
free water and solid-like macromolecular proton
pools in tissues. One mechanism for this exchange
is a through-space interaction between magnetic
dipoles, commonly termed cross-relaxation. Another
mechanism is the chemical exchange of protons in
free water protons and certain functional groups on
macromolecules, thought to occur by way of the
intermediate pool of the macromolecule’s interfacial
water. The MT’s sensitivity to the relative propor-
tions of free water and macromolecules may make
this technique sensitive to common aspects of muscle
pathology such as inflammation and fibrosis. Because
the rate of chemical exchange between water and
functional groups such as amides is sensitive to pH
[76], MT methods may potentially afford insight into
a tissue’s metabolic status as well.
Two common MT methods are the pulsed satu-
ration (PS-MT [77]) and selective inversion recovery
(SIR-MT [78, 79]) approaches. Both approaches take
advantage of the broad linewidth of the macromolec-
ular proton signal, relative to the narrow linewidth of
the water proton signal. In PS-MT, the macromolec-
ular signal is saturated with an RF pulse centered on
a resonance frequency sufficiently different from the
water signal. Typically, the saturation pulse is applied
several kHz off-resonance to water. MT takes place
between the solid and free water proton pools, and
as the saturated magnetization enters the free water
pool, the water signal is reduced in amplitude. In SIR-
MT, a narrow-bandwidth RF pulse is used to invert
the water signal selectively. MT between the solid
and free water proton pools causes the water signal
to recover as a bi-exponential function.
Semi-quantitative methods based on PS-MT meth-
ods express the reduction in water signal amplitude
following off-resonance saturation relative to the
signal observed in a control condition (typically
achieved by centering the saturation pulse ∼50 kHz
off-resonance to water, beyond the reasonable lim-
its of the macromolecular peak). This quotient is
called the MT-ratio (MTR). A disadvantage of this
approach is that the MTR is sensitive to the satura-
tion pulse power and offset frequency. Consequently,
if studies have employed different RF saturation pulse
parameters, then strict quantitative comparisons of
the results are not possible. In truly quantitative MT
(qMT) methods, data are collected at a variety of sat-
uration offsets (for the PS-MT method) or inversion
times (for SIR-MT). These data are then fitted to a
biophysical model, allowing estimation of the ratio
of macromolecular to free water protons (the pool
size ratio, PSR), the relaxation rates of these pools,
and the rates of exchange between them.
MT has been developed and applied in healthy
muscles [80–90]. Some of these works have pro-
vided fundamental quantitative or semi-quantitative
descriptions of the MT process in healthy muscle tis-
sue. For example, Harrison et al. observed that the
amount of signal loss due to off-resonance MT satu-
ration pulses is greater in the T2signal component
associated with the intracellular space than in the
T2component associated with the extracellular space
[82]. Normative values for exchange rate, PSR, and
other important parameters have been reported [83,
85, 86, 88, 89], as a function of variables such as age
[85, 88] and sex [88]. Other studies have advanced
the understanding of the MT process in muscle or
implemented technical advances. For example, Louie
et al. studied the effect of intracellular pH on the
MT rate and observed a direct, linear dependence of
the MT rate on pH [84], which they postulated may
reflect base-catalyzed amide-water proton exchange
[76]. Other advances have included the correction
of MT parametric maps for RF inhomogeneity [91,
92], the effects of fat signal contamination [90],
the reduction of data acquisition time by adopting
reduced parameter models [90], and establishment
of the reproducibility of the technique [88, 89].
MT has also been used to characterize human
NMD and animal models of human disease. Quan-
titative MT has been used to study small animal
models of muscle inflammation, wherein it has been
shown that inflammation decreases the PSR [93].
Inasmuch as MT reflects both the macromolecular
and free water proton pools, the PSR may also be
influenced by fibrosis [94, 95]. Also consistent with
inflammation, Sinclair et al. have observed a reduced
MTR in patients with peripheral neuropathies [87].
They further observed that the MTR was correlated
with clinical severity and was even reduced in oth-
erwise normally appearing images. In a study of
patients with Charcot-Marie-Tooth disease and inclu-
sion body myositis, Sinclair et al. likewise observed
reductions in MTR [91].
MR Elastography
Palpation is frequently used in the physical exam-
ination of patients with NMD. Pathological features
related to NMD such as fat infiltration, fibrosis, and

8G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
edema affect the muscle biomechanical properties
and can be detected using palpation by the sense of
touch. Although very practical in the clinical setting,
assessment of skeletal muscle biomechanical proper-
ties using manual palpation is of a qualitative nature
and therefore less useful to objectively assess disease
progression and for treatment evaluation. A quantifi-
cation of the biomechanical properties of tissue can
be done with a specialized MRI technique called MR
Elastography (MRE). Since the invention of this pal-
pation by imaging technique, it is increasingly used
to quantify the biomechanical properties of skeletal
muscle [96, 97].
MRE is based on the imaging of shear waves,
most often introduced in the tissue by an external
vibrating actuator mechanically coupled to the tis-
sue, usually the skin. During vibration, the tissue
displacements, i.e. the shear waves, are imaged with
a motion-encoded MRI sequence synchronized to the
motion [98]. By inversion of these shear wave images,
viscoelastic biomechanical properties, such as the
complex shear modulus G* can be subsequently esti-
mated [99].
Current applications of skeletal muscle MRE can
be roughly divided in four categories. A first appli-
cation involves skeletal muscle rheology. Skeletal
muscle exhibits nonlinear viscoelastic anisotropic
material behavior, which makes accurate determi-
nation of the biomechanical properties challenging.
Several research groups focused on sophisticated
experimental and analysis techniques to approxi-
mate the true mechanical properties. Most effort was
put into including the anisotropic behavior of skele-
tal muscle. Some proposed a linear (visco-) elastic
transverse isotropic material approach as intermedi-
ate solution [101–106]. Other relevant rheological
features for skeletal muscle, including nonlinear
viscoelasticity, frequency dependence (power-law)
behavior, incompressibility, tissue fluid content,
tension-compression nonlinearity, and inhomogene-
ity are currently often neglected, however, could
be worthwhile to investigate further [107]. Sec-
ondly, the study of the biomechanical properties of
healthy skeletal muscle is an active field of research,
since it is believed that there are relevant differ-
ences in stiffness between muscles and because
thorough understanding of muscle’s baseline biome-
chanical properties is considered essential to study
muscle injury or disease. Most studied are the indi-
vidual skeletal muscles in the lower extremities,
upper extremities, and shoulder [108–117]. Further-
more, MRE-derived biomechanical properties during
skeletal muscle contraction and relaxation can be
used as readout for skeletal muscle function. Several
groups have studied the relationship between mus-
cle biomechanical properties and load [115, 116],
as well as the effects of skeletal muscle exercise
on muscle biomechanical properties, and the differ-
ence in relaxed versus contracted skeletal muscle
[100, 105, 109, 113, 118–120]. The fourth appli-
cation is MRE applied for studying skeletal muscle
pathology. MRE has been applied to study mechan-
ical changes due to neuromuscular dysfunction,
myositis, deformation-induced damage, testosterone
treatment, aging, disuse, hyperthyroidism, myofas-
cial pain, and Duchenne muscular dystrophy (DMD)
[100, 110, 120–127]. For example, in boys with
DMD, Bensamoun et al. have observed elevated vas-
tus medialis muscle stiffness at rest and decreased
muscle stiffness in contracted state compared to
healthy controls (Fig. 2) [100].
31P and 13 C MR Spectroscopy
Following initial animal experiments, MR spec-
troscopy (MRS) was first applied to human subjects
in the 1980s, using the 31P nucleus to monitor the
levels and fate of high-energy phosphates in skele-
tal muscle [128]. The technique was initially applied
in healthy individuals at rest and during exercise,
and rapidly also in NMD patients [129, 130]. 31P
MRS offers a unique non-invasive window on some
key high-energy phosphate metabolites such as ATP,
phosphocreatine (PCr) and inorganic phosphate (Pi),
which are present at sufficient tissue levels to gener-
ate resonances with good SNR. Other compounds that
may be estimated from usually less intense signals in
31P MR spectra of muscles are total NAD(H), phos-
phomonoesters (PME) and phosphodiesters (PDE)
[131]. In addition, intracellular pH can be derived
from the chemical shift of the Pi resonance [132],
free Mg++ from shifts of the ␣-ATP and ␥-ATP peaks
and a measure of free cytosolic ADP may be derived
from the creatine kinase (CK) equilibrium assuming
its substrates to be free in solution [133]. The enzyme
kinetics of the CK reaction and of ATPases may be
estimated from saturation transfer experiments [134,
135]. During in-magnet exercise bioenergetics data
involving PCr, ATP, Pi, and tissue pH can also be
acquired [136], Recent studies have identified so-
called 31P-31 P nuclear Overhauser effects for ATP
due to its transient binding to large molecular struc-
tures like mitochondria [137].

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 9
Fig. 2. Biomechanical characterization of skeletal muscle in DMD children. (A) MR Elastography derived sheer stiffness color maps projected
on parts of the legs of a healthy subject and a child with DMD. VM = vastus medialis, Sr= sartorius, Gr =gracilis. (B) Quantification of the
shear stiffness in the vastus medialis and the subcutaneous adipose tissue (*= P<0.1). Figure reproduced with permission from Ref. [100].
To be able to excite and acquire 31P spectra, or
any other nucleus apart from 1H, a dedicated coil is
needed. 31P MR spectroscopy of skeletal muscle is
traditionally done by only using a 31P surface coil
(for transmit and receive) close to the muscles of
interest, sampling mostly signal from tissue adjacent
to the coil. However, spectra acquired in this way
are not very well localized and do not arise purely
from one muscle which can complicate interpreta-
tions, especially in NMDs where muscles of clinically
asymptomatic patients can already show fatty infil-
tration in some muscles, while others still appear to
be normal [138]. This has been addressed in sev-
eral ways, including localization methods like ISIS
or sLASER to select single voxels within a mus-
cle [139, 140] or spectroscopic imaging [141–143].
As these methods usually sacrifice some of the total
information present in 31P MR spectra of muscles
an educated selection of a method for a specific
purpose is required. It was recently discovered that
high-energy phosphate metabolism and fatty infil-
tration may be very different within a muscle and
therefore localization is also required to capture this
heterogeneity [144–147].
A major advantage of using 31P MRS in NMD is
that the metabolites sampled are hardly present in
fat and are therefore representative of muscle and to
a lesser extent interstitial tissue rather than the fat.
This is potentially very useful in therapy develop-
ment, as drugs are commonly designed to preserve
or improve muscle tissue. It is generally thought that
fat replacement of muscle tissue is irreversible, while
other processes may be at least partially reversible.
Typical findings in resting 31P spectra of skeletal mus-
cle affected by NMD are a low total P-compound level
due to atrophy and fatty infiltration, a low PCr/ATP
ratio indicating a loss in contractile elements or unbal-
anced energy metabolism, high Pi/PCr ratio which
may indicate increased ADP and thus unbalanced
ATP production versus consumption, changes in PDE

10 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
Fig. 3. (top row) Representative reconstructed water images of the right lower leg and (bottom row) 31P spectra of the TP muscle of a DMD
patient at (A) baseline, (B) 12-months and (C) 24-months. PDE/ATP ratios in TP muscles are shown in the graph and %fat for the all analyzed
muscle for all three time points are as follows. Baseline: GCL =5.6%; GCM =7.3%; SOL =7.1%; PER =14,4%; TA =6.24%; TP =4.2%;
12-months: GCL =6,6%; GCM =8.8%; SOL =5.2%; PER =20.6%; TA =5.71%; TP =4.3%; 24-months: GCL =10.1%; GCM =11.3%
SOL =5.9%; PER =24.7%; TA =7.3%; TP =4.3%. GCL =gastrocnemius lateral head, GCM=gastrocnemius medial head, SOL =soleus,
PER =peroneus, TA =tibialis anterior, TP =tibialis posterior. Figure reproduced with permission from Hooijmans et al. [157].
levels and an increased tissue pH [148–150]. How-
ever, not all NMDs show the same changes, and even
if the metabolic alterations are similar, they likely
arise from different mechanisms for each pathology
[150, 151]. As a result, it is unlikely that metabolic
alterations as shown by 31P MRS are specific markers
for disease in NMD.
The technique has also been used to study patho-
physiology and timing of events in the disease
cascade. Recently, it was postulated that the increased
tissue pH in patients with DMD is due to a splitting
of the Pi peak, with an interstitial Pi associated with
damaged dystrophic myocytes and/or expanded inter-
stitial space related to fibrosis [152]. An additional
Pi peak has also been assigned to a mitochondrial
compartment in healthy muscles [153]. Using chem-
ical shift imaging, it was shown that in patients with
facioscapulohumeral dystrophy (FSHD), metabolic
changes were only present in fat infiltrated mus-
cles [154], while in both Becker and Duchenne
muscular dystrophy, the PDE peak was increased
already in muscles without fat infiltration [155,
156]. PDE is often assigned to membrane break-
down products and its signal may be used as a
biomarker for treatments [6]. Longitudinal stud-
ies which followed 31P MRS changes over time
in NMD patients are scarce [9, 152, 157]. In the
forearms of patients with DMD, it was shown that the
ratio of the alkaline Pi signal over Pi and Pi/PCr were
increased over one year in non-ambulant patients only
[9, 152]. In the legs of patients with DMD (an exam-
ple is demonstrated in Fig. 3), no changes in 31PMR
indices were observed in a 2-year follow up, but in
this study no distinction was made between ambulant
and non-ambulant patients [157].
Studies with 31P MRS during and after exercise
to assess the dynamics of high-energy phosphate
metabolism have only rarely been performed in
patients with muscular dystrophies, but have been
done in metabolic and inflammatory myopathies. In
general, changes in tissue pH and PCr depletion dur-
ing exercise and differences in metabolite and tissue
pH recovery after exercise were noted in a number of
diseases [149, 158–160]. More recent studies on this
topic are scarce [161–163], and maybe more practi-
cal to assess the kinetics of CK and ATPase is the
application of saturation transfer experiments at rest.
It has been performed in muscles of patients such as
diabetics type II, but not yet in NMD [164].
As 13C MRS is a relatively insensitive method
and 13C nuclei only occur at 1% natural abun-
dance it is only feasible to detect highly concentrated
compounds at natural abundance such as lipids and
glycogen [165]. As a result, application of this

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 11
Fig. 4. 1H and 23 Na MR images of a patient with a muscular channelopathy (paramyotonia congenita). The right lower leg (*) was cooled
with an ice bag, which results in an increase of the intracellular Na+concentration and in severe muscle weakness. Two reference tubes
were used for signal normalization (1:51.3 mmol/L NaCl and 5% agarose gel, 2:51.3 mmol/L NaCl solution). (A) T1-weighted and (B)
T2-weighted 1H MRI revealed no pathological signal differences between the cooled (*) and the non-cooled muscle. (C) The cooled leg
shows a slightly increased total Na+concentration. (D) 23Na inversion recovery (IR) MRI revealed a distinct increase of the signal in the
cooled leg and, thus, visualizes the increase of the intracellular Na+content. Note, reference tube 2 (pure saline solution) shows no signal
intensity in 23Na IR MRI. Images adapted from Ref. [172] with permission.
method in NMD is limited to a study in adult
onset acid maltase deficiency [166]. As an alter-
native, 13C labeled compounds can be used, such
as 13C1-glucose, as this allows following metabolic
conversions. As 13C labeled material is costly and
the method requires special hardware it is not much
applied and only in a rare case to a neuromuscu-
lar disease [167]. A much higher sensitivity can be
achieved with so-called hyperpolarized 13C, but this
new approach is still in development and has not been
applied yet to neuromuscular diseases [168].
Sodium imaging
Sodium ions (Na+) play an important role
in the ion homeostasis of skeletal muscle tis-
sue. In healthy tissue, the intracellular con-
centration ([Na+]in = 10–15 mmol/L) is about 10-
fold lower than the extracellular concentration
([Na+]ex = 145 mmol/L). This concentration gradi-
ent across the cell membrane is maintained by the
Na+/K+-ATPase pump and is of utmost importance
for the excitation and inhibition of muscle cells.
Sodium MRI enables non-invasive imaging of the
stable and naturally occurring (100% natural abun-
dance) isotope (23Na). However, its quadrupolar
nature leads to a fast transverse relaxation and, thus,
requires acquisition techniques that enable ultra-short
echo times [169]. In addition, dedicated hardware
such as transmit/receive coils that are tuned to the
Larmor frequency of 23Na are required [170].
During the last decades, 23Na MRI has evolved
into a versatile tool in biomedical research [171]. It
enables non-invasive determination of the total tissue
Na+concentration (Fig. 4C) and relaxation-weighted
measurements (Fig. 4D). The latter allows at least
a partial separation between different Na+compart-
ments. For example, it was shown that 23Na inversion
recovery MRI can visualize changes of the intracel-
lular Na+content in muscular channelopathies [172]
that are not visible in conventional 1H MRI (Fig. 4).
23Na inversion recovery MRI and conventional 23Na
MRI in combination with sophisticated modelling
have also been applied to provide a quantitative esti-
mation of the intracellular Na+concentration in brain
tissue [173].
Although 23Na MRI largely benefits from the
increased signal-to-noise ratio at ultra-high magnetic
field strength (e.g. 7 T), 23Na MRI of skeletal mus-
cle can also reliably be performed at clinical field
strengths (3 T).
In NMD, it has been shown that muscle tissue of
patients with myotonic dystrophy [174] and DMD
[175] have increased sodium concentrations. Also in

12 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
the muscles of mdx-mice (a model for DMD), abnor-
mally high intracellular concentrations have been
reported [176]. In DMD, the absence of dystrophin
modifies the gating properties and expression level of
the voltage-gated Na+channel (Nav1.4). This results
in an increased Na+concentration under the sar-
colemma [177]. It was hypothesized that the elevation
of the intracellular Na+concentration contributes to
an elevated rate of cell death in dystrophic muscle.
In vitro studies demonstrated that both effects can
be reversed by a specific Nav1.4 channel blocker
(tetrodotoxin) [177]. In DMD patients, it has been
shown, that the K+-sparing diuretic agent eplerenone
reduces muscular Na+overload and edema [178].
However, long-term studies with larger patient num-
bers are required to investigate treatment effects of
eplerenone in more detail. In future, 23Na MRI might
be used in clinical studies that investigate disease pro-
gression or therapy effects in neuromuscular diseases.
CLINICAL NEEDS
Inflammation & disease activity
Multiple MR imaging strategies have been shown
to be sensitive to muscle inflammation and muscle
damage in the NMDs. For instance, STIR sequences
are commonly used to enhance the contrast gener-
ated by in areas of muscle inflammation and edema.
In many NMDs it is believed that areas that are
hyperintense on STIR images precede the eventual
deposition by fibrofatty tissues. For instance, Marden
et al. [179] observed regions of increased signal inten-
sity on STIR images in the muscles of young DMD
boys in the absence of fatty infiltration. This finding
is consistent with the conjecture that inflammation,
as well as necrosis and edema, occurs early in DMD,
prior to the loss in contractile tissue and accumulation
of fatty infiltration. The importance of inflammation
in DMD is supported by the 100-fold higher serum
levels of TNF-␣in boys with DMD compared to con-
trols (27.8 vs. 0.27 ng/L) [180]. Changes in the serum
levels of TNF-␣with age are also consistent with the
observations that muscle expression of TNF-␣and
IL-6 decreases with age [181]. In another type of mus-
cular dystrophy, FSHD, a direct relationship between
inflammatory markers (serum and tissue) and hyper-
intensity observed on STIR images has also been
documented [182]. Due to these early inflammatory
changes a wide range of therapeutics are in clini-
cal trials for NMD targeting specific inflammatory
pathways [183].
Changes in muscle T2, MT, and diffusion have all
been commonly used to look at acute muscle dam-
age and inflammation. Damage and inflammation
has been induced in animal models using myotoxins
[184, 185], a local inflammatory agent [93, 186–188],
and in humans [189–191] and preclinical models
[192–195] following eccentric muscle contractions.
Due to the close association between muscle dam-
age (loss of sarcoplasm integrity), inflammation, and
edema, it is often difficult to separate these individual
components from net changes in MRI signal. Despite
this, careful animal studies along with detailed histo-
logical analysis at specific time points [93, 186] and
multimodal MR acquisitions are providing insight
into sources of the different contrast mechanisms [57,
93, 184, 186]. Often in these analyses it is necessary to
take into account the multicomponent nature of relax-
ation or diffusion that are occurring on the subpixel
scale [186]. Due to the fact that the primary prob-
lem in many of the muscular dystrophies is related to
a mutation or absence of a structural protein, many
modern therapeutics aim at replacing or mitigating
these defective or missing proteins which leads to
bouts of muscle damage and inflammation [183].
Agents that are known to decrease inflammation
have also been shown to lower muscle T2. Cur-
rently the only intervention that has been shown
to be successful in mitigating some of the effects
of dystrophy are glucocorticosteroids. Studies have
shown that corticosteroid treatment can significantly
improve muscle strength and function, reduce the
incidence of scoliosis, maintain the respiratory and
cardiac function, as well as prolong ambulation and
survival in boys with DMD [196–198]. Although
there is increasing evidence of long-term benefits
of corticosteroids in DMD, the exact mechanism of
action of these drugs in dystrophic muscle is unclear
but a primary mechanism is believed to reduce muscle
inflammation possibly by modulating NF-κ␤activa-
tion. Arpan et al. measured the muscle water T2in 5
to 7 year-old DMD boys whom were using corticos-
teroids and age matched corticosteroid na¨
ıve subjects
and found that muscle water T2measured by 1H-
MRS was significantly reduced in both the vastus
lateralis and soleus muscles of the boys on corticos-
teroids [199], consistent with the anti-inflammatory
effect of corticosteroids (Fig. 5). At the low levels of
fatty tissue deposition in these young DMD patients,
they also found lower T2values in both thigh and
lower leg muscles of boys on corticosteroid treatment.
Of the muscle groups studied, the gracilis muscle was
the only exception. Where, the gracilis is known to

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 13
Fig. 5. Cross-sectional comparisons of water T2measured by MRS between corticosteroid- treated (CS) and corticosteroid-na¨
ıve (CS-
na¨
ıve) boys with Duchenne muscular dystrophy. (A) Spectroscopic relaxometry measurement using 1H-MRS STEAM to quantify water T2
in individual muscles. (B) T2values for both the soleus (Sol) and vastus lateralis (VL) muscles were lower in CS boys compared to the
CS-na¨
ıve boys, indicating less damage in the muscles of boys on corticosteroid treatment. Figure reproduced with permission from [199].
be one of the most preserved muscles in DMD and
hence might not reveal treatment response at such a
young age. Both MRI and MRS measures detected
the beneficial effects of corticosteroids on the skele-
tal muscles as early as 3 months after drug initiation.
Additional evidence of ability of water T2to be a
readout for inflammation was shown by Marty et al.
[32] in which they found an almost complete recovery
to normal water T2values after 3 months of steroid
treatment in patients with juvenile dermatomyositis.
Whereas the clinical benefits of steroids in DMD
are clear, their use also comes with significant side
effects and great effort has been spent to identify
new anti-inflammatory agents for the NMDs [183].
Preclinical studies in the golden retriever muscu-
lar dystrophy (GRMD) model revealed that MRI
water T2changes following the treatment with a
nemo-binding-domain (NBD) peptide, which is a
specific inhibitor of NF-κ␤[200]. Post NBD treat-
ment, GRMD dogs had normalized postural changes
and a trend towards lower tissue injury on T2
images. Unfortunately, despite phenotypic improve-
ment in the dystrophic dogs, NBD administration
over time led to infusion reactions and an immune
response in both treated GRMD and wild type
dogs. Other drugs that target inflammation that have
shown great success in preclinical models such as
Vamorolone (VBP-15; NCT03439670), Edasalonex-
ent (CAT-1004; NCT02439216), Histone deacetylase
2 inhibitor (Givinostat; NCT02851797) are now in
clinical trials [183] and either plan to or include MRI
and MRS measures as secondary outcome measures.
23Na imaging – see also section 2.7 – showed that
areas of hyperintensity on STIR images from skeletal
muscle in DMD subjects are directly related to muscle
edema [201]. In a pilot study, Glemser et al. per-
formed a longitudinal assessment of edema in 2 DMD
patients treated with eplerenone or steroid using
a combination of 23Na, Dixon, and STIR imaging
[178]. This study demonstrated that longitudinal
changes in sodium content, and presumably tissue
edema, can be noninvasively monitored throughout
therapeutic intervention.
Taken together, these findings indicate that mul-
tiple MR sequences may be sensitive to early
inflammation in the NMDs and emphasize the
potential of MRI and MRS as biomarkers for the
quantification of early and subtle muscle changes
caused by the disease process and evaluation of ther-
apeutic interventions in NMD.
Fat infiltration
Muscle fat infiltration has been extensively inves-
tigated in NMDs with single-voxel 1H MRS and 2D
and 3D imaging. Over the past decade, fat infiltration
has been evaluated with respect to functional mea-
sures, the rate of disease progression, in clinical trials,
and to assess differences between and within indi-
vidual muscles. A listing of important NMD studies
to date assessing muscle fat fraction can be found
in the recently published article from our COST
action [26]. All these findings together showed that
muscle fat infiltration differs between muscles and
diseases, is highly associated with disease progres-
sion and that functional and strength measures are
correlated to the fat fraction in NMD. A recent large
longitudinal study in DMD reinforced this finding,
with correlations to functional endpoints and sen-
tinel events ranging from 0.59 to 0.78 depending
on the muscle studied [202]. Besides all the possi-
ble readouts used for fat fraction itself — i.e. whole

14 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
muscle fat fraction, specific volume fat fraction for an
individual muscle or muscle groups, delta %fat etc.
– other closely related measures, such as contractile
cross-sectional-area (cCSA = %fat ×total CSA) or
muscle volume have also been widely investigated
and it was shown that specific strength is reduced
in a number of NMD [10, 203–205]. Recently, this
analysis was extended by obtaining physiological
cross-sectional areas in patients with FSHD, showing
that specific strength is also reduced in these patients
[206]. Besides the relation with function, some recent
studies also focused on the methodological aspects
of fat infiltration as outcome measure. Different MR
fat quantification methods have been assessed and
benchmarked against each other and other methods
for fat quantification [13, 26, 207–210]. Muscle fat
fraction measurements showed to be highly repro-
ducible from day-to-day and across centers, with CV
values ranging between 1.8–7.3% [10, 88, 211]. How-
ever, in situations where muscles inevitably change
in between measurements, for instance in a maturing
study population (for instance in patients with DMD
and patients with spinal muscular atrophy, SMA),
repositioning according to internal muscle references
and bony landmarks becomes more challenging and
could affect the reproducibility.
It is important to consider the non-uniform fat
distribution patterns along the proximodistal mus-
cle axis in various NMDs [145–147, 212, 213].
This non-uniform shape emphasizes the need for
accurate spatial localization between measurements,
as a slight shift of the imaging stack or volume
can result in over- or underestimation of fat frac-
tion between measurement points. Subsequently,
the choice of intrinsic parameters of most imag-
ing and spectroscopy approaches, i.e. slice gaps,
slice thickness, restricted voxel size, results in lim-
ited and location specific information, which could
amplify the effect of inaccurate spatial localization.
3D acquisitions with full limb coverage would allow
accurate offline matching of datasets in a standard-
ized way as well as the possibility to retrospectively
decide which readout is most suitable for the study
set-up.
Fibrosis
Besides fatty infiltration of skeletal musculature,
patients with NMD often show excessive accumu-
lation of extracellular matrix proteins as collagen
(fibrosis) in their skeletal muscles. The process
of formation of proteins in the extracellular space
between remaining myocytes and resulting scars
is physiologically active during repair of damaged
musculature (e.g. after injuries). This repair involves
several cell types acting in response to various local
and systemic signals, but those activities are not
persisting in healthy subjects. In contrast, during
chronic tissue damage in muscular dystrophies, an
inflammatory cell infiltration persists, leading to
activation of fibroblasts and their transformation
to myofibroblasts, which are continuously building
up excess amounts of connective tissue, while the
reparative capacity of muscle stem cells is more and
more attenuated.
Volume share and composition of the extracellu-
lar space in musculature involved in NMD strongly
depends on the intensity and duration of inflamma-
tory activity (Fig. 6). Their non-invasive detection is
of high clinical interest in order to better stage the dis-
ease and monitor therapeutic interventions. In early
involvement of musculature, inflammatory activity
leads to increasing amounts of interstitial fluid with-
out fibers. Thus, an increased amount of free water
is detectable by prolonged T2values in MRI, and
also other MR detectable features of musculature
(e.g. proton density, diffusion,magnetization transfer,
23Na content) undergo changes due to inflamma-
tion. Persistent inflammatory activity often leads to
an irreversible and increasing accumulation of pro-
tein fibers in the extracellular space (fibrosis). It is
evident that mechanical properties (stiffness) and sig-
nal characteristics of musculature in MRI depend
on the composition of the extracellular compartment
containing variable amounts of fluid and increasing
volume share of protein fibers.
However, since fibrotic areas and remaining mus-
cular elements are distributed microscopically, it is
impossible to selectively record MR signals from the
interstitial compartment only. We have to deal with a
superposition of signal contributions from remain-
ing functional musculature, from the components
of the interstitial space (mainly water and protein
fibers, but also cellular elements), and in some cases
from fat in cases with fatty degeneration. So, valid
and unambiguous interpretation of results from MR
imaging with respect to fibrosis of musculature is
very challenging. The gold standard for assessment
of fibrosis (in any type of tissue) is biopsy. Unfor-
tunately, this procedure is invasive, and results are
often not representative for extended tissue areas,
since fibrosis is often not disseminated homoge-
nously, neither microscopically nor macroscopically.

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 15
Fig. 6. (A) Histological sections of healthy muscle, inflamed muscle, and heavily fat infiltrated and fibrotic muscle tissue. (B) Schematic
of changes in volume share and composition of the interstitial space due to inflammation and as a consequence of fibrosis and fat
infiltration.
Therefore, non-invasive imaging techniques for
reliable assessment of the degree of fibrosis are highly
desired.
The most common MR approach for examina-
tion of muscle stiffness is MR elastography [101,
123]. This technique works with mechanical stimula-
tion and assessment of mechanical wave propagation
inside musculature and is established for assessment
of fibrosis in other organs like liver and breast. Appli-
cations in musculature are more difficult because of
the variable muscle tone and geometric anisotropy of
muscle tissue (see section 2.5).
Other approaches to non-invasively measure fibro-
sis by MRI are currently investigated. Changes in
water diffusion, magnetization transfer, prolonged
T1, prolonged T1-rho, short T2components, and late
enhancement after administration of a contrast agent
were correlated to fibrosis in myocardium, liver, and
bone marrow [94, 214–217]. However, only a few
approaches have been discussed regarding applica-
tions in NMD [6, 93]. Pathological events concurrent
to fibrosis in muscular dystrophy, such as inflamma-
tion and muscle damage, may affect the contrast in the
same direction or in opposite direction than fibrosis,
which poses a challenge to disentangle the contribu-
tion of fibrosis. Nevertheless, a promising technique
that deserves further attention in the context of NMD
is ultra-short echo time (UTE) imaging, which was
already applied to characterize age-related difference
in collagenous tissue in muscle [218].
Muscle cell morphology/architecture
Routine clinical MRI investigation in patients with
NMD continues to rely largely on T1-weighted,
T2-weighted and STIR methods to obtain high
definition anatomical information. These methods
primarily distinguish between apparently normal
appearing muscle fibers and muscle tissue with strong
fatty infiltration, which represents the end stage of the
disease process. In the context of clinical patient man-
agement and treatment there is an important need to
visualize the early and ongoing pathological changes,
which are occurring in the remaining muscle fibers.
For clinical trials, such information is vital to under-
stand drug-target interaction and to identify early
treatment response. In the muscular dystrophies, dis-
ease progression is characterized at the cellular level
by cycles of degeneration and regeneration of mus-
cle fibers. Over time, the ability to regenerate fibers
is lost and fibers are eventually replaced by fibrotic
tissue and fat. Histological measurement of muscle
in animal models shows that the distribution of mus-
cle fiber sizes changes in dystrophic muscle over and
above differences driven by maturation or healthy
ageing [219]. Specifically, an increase in the propor-

16 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
tion of fibers with much larger or smaller diameter
than normal muscle is observed, represented by a shift
and broadening of the muscle fiber size histogram
[220]. In addition, muscle sarcolemma permeability
to large tracer molecules is increased in dys-
trophic muscle, driven by damage to the membrane
[221, 222].
These two basic morphological properties (fiber
size and permeability) are fundamental determinants
of the effective (apparent) diffusivity of water within
the tissue microstructure. Indeed, the earliest in vivo
studies of self-diffusion characteristics of water in
tissue by magnetic resonance methods were made in
excised muscle specimens [223]. In that early work
Tanner demonstrated that measurement of water dif-
fusivity at extremely long diffusion time (hundreds
of milliseconds to seconds, achieved through the use
of stimulated echoes) showed evidence of restricted
behavior. The magnitude of the effect increased
with diffusion time as each water molecule has an
increased probability of interacting with major bar-
riers in the tissue such as the sarcolemma itself.
More recently such measurements have been revis-
ited as imaging experiments in living tissue, both
in animal models and in human subjects [50, 220,
224–226]. In the context of NMD, studies in wild
type mice have shown that clear maturation related
changes can be seen in diffusion restriction [220],
while in the mdx-mouse model significant differences
are reported, particularly at long diffusion times with
reduced restriction compared to wild type animals
[227]. Applications of the DWI technique in humans
are also described in section 2.3.
To quantify the changes seen using diffusion meth-
ods, Tanner et al. applied a simple biophysical model
of regularly spaced semi-permeable membranes to
estimate muscle fiber size and permeability [223].
It is however clear from the progressively complex
models which have subsequently been proposed to
quantify muscle water diffusion behavior [224, 226,
228], that sarcolemmal permeability and muscle fiber
scale are inherently opposing factors – increasing
permeability reduces restriction whatever the change
in muscle fiber size. In diseased muscle, such as in
the muscular dystrophies where significant changes
are expected in muscle fiber permeability, it seems
unlikely that a unique solution is possible and that the
effects of altered permeability will dominate the abil-
ity to assess muscle fiber sizes [229]. Nevertheless,
the potential to differentiate normal from diseased
muscle using long diffusion times and to detect
changes with time [227], as suggested by recent
animal data suggest that further development is
warranted.
EMERGING IMAGING TECHNOLOGY
Muscle segmentation
Segmentation of individual muscles is of high
interest in NMD, as fat infiltration can differ greatly
even between adjacent muscles. This issue is also of
interest if one aims at monitoring changes resulting
from training and/or pharmaceutical interventions.
The task of segmenting MR images into meaning-
ful compartments has been recognized as challenging
for a variety of reasons. In addition to the high muscle
shape variability among subjects, fatty infiltration
and muscle atrophy can change the visibility and
location of borders between muscles and therefore
greatly complicate the segmentation task. So far,
manual segmentation of anatomical structures is the
gold standard and has been used in multiple studies.
However, this approach is widely acknowledged as
very time-consuming and can be operator-dependent
[230]. For example, accurate segmentation of the
quadriceps femoris compartment in controls may
take a few hours.
More recently, several semi-automated and auto-
mated methods have been tested on MR images
of healthy subjects. An example by Ogier et al.
is shown in (Fig. 7) [231]. They implemented an
algorithm allowing a semi-automatic transverse prop-
agation of a number of manually drawn masks,
resulting in an 85% reduction in segmentation time.
A random walk algorithm based on graphs has
been reported by Baudin et al. [232, 233], while
Gilles et al. [234] proposed a method based on
mesh deformable registration models. In order to
take into account the large inter-individual vari-
ability, Prescott et al. [235] used a semi-automatic
segmentation method based on the pre-selection
of appropriate templates selected from a database.
Using atlas-based registration, Ahmad et al. proposed
a semi-automatic segmentation tool for quadriceps
muscles [236]. At the whole body-level and using
a multi-atlas based method, Karlsson et al. reported
a 3% volume error for the quadriceps femoris com-
partment [237]. Using a similar approach, Le Troter
et al. further confirmed a 3% volume error for
the quadriceps femoris compartment but reported
much larger errors when considering individual
muscles [238]. The corresponding muscle volume

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 17
A
B
C
Fig. 7. Semi-automatic segmentation of the upper leg muscles on axial slices. (A) Manual segmentations (blue colored ROIs) of vastuslateralis
(VL), rectus femoris (RF), vastus medialis (VM) and vastus intermedius (VI). Sagittal and axials views of the automatic segmentations of
intermediate slices (yellow-orange colored ROIs) based on (B) 2 and (C) 4 initial manual segmentations. Figure reproduced from Ref. [231]
with permission.
errors measurements ranged from 4 (vastus medialis
muscle) to 17% (rectus femoris and vastus lateralis
muscles).
Overall, segmentation methods have been devel-
oped and tested so far mostly in healthy subjects so
that the issues of fat infiltration and muscle atrophy
as confounding factors still need to be systematically
addressed.
Accelerated imaging
The acquisition time of quantitative MRI is a bar-
rier to the expanded use of these methods in NMD,
particularly for DMD and SMA, where restricting the
duration of the protocol is critical to ensuring child
compliance and avoiding motion artifacts.
The acquisition time of an MRI acquisition is con-
ventionally limited by the need to acquire a matrix of
raw data (k-space) of an equivalent size to the even-
tual desired imaging matrix (the Shannon-Nyquist
criterion). Whereas the read dimension is acquired
rapidly, the other dimensions using phase encoding
require multiple repetition times. If the amount of k-
space data points to be acquired can be reduced by
the use of additional information, then the number of
repetition time delays and hence the acquisition time
can be reduced.
Parallel imaging, now available on all modern
scanners, achieves this by exploiting the spatial sen-
sitivity of multi-channel array coils. More recently,
research effort has exploited a different source of
additional information, the sparsity of the image
under a mathematical transform, such as discrete
wavelets or the calculation of total variation. If the k-
space is undersampled quasi-randomly, such that no
coherent aliasing is produced under inverse Fourier
transformation, then the sparsity in spatial and/or
temporal domains can be exploited to permit a high

18 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
Fig. 8. Reconstructions of the left lower leg of a subject showing (A) the calculated fat fraction maps and (B) the total water and fat signal
map for a fully sampled acquisition (top row) using CS and traditional parallel imaging (GRAPPA) with different undersampling ratios
(subsequent rows). Figure reproduced with permission from Hollingsworth et al. [239].
quality iterative reconstruction. This method has been
coined compressed sensing.
Research into the use of compressed sensing in
MRI has proliferated [240], with applications in
skeletal and cardiac muscle relevant to NMD. It has
been demonstrated that combined compressed sens-
ing and parallel imaging could be applied to a 3D
IDEAL-type fat fraction measurement in skeletal
muscle at acceleration factors of up to 5×without
significant loss of image quality or impairment of fat
fraction fidelity (Fig. 8) [22, 239, 240].
The measurement of cardiac function is of increas-
ing interest in NMD. In preclinical cardiac imaging,
a Cartesian spatiotemporal undersampling scheme
has been used to allow an acceleration of up to 3×
[241] while producing comparable results for end
systolic and end diastolic volumes and the early to
late filling ratio in diastole. Subsequently, a 2D multi-
slice golden-angle radial acquisition was applied to a
sequence with ultra-short echo time, which reduces
artifacts due to flow and susceptibility [242]. Golden-
angle radial schemes permit post-hoc decisions to
be made about the number of spokes to be recon-
structed per cardiac phase and the number of cardiac
phases resolved in the spatiotemporal scheme. This
permitted accelerations of 2, 4 and 5×to be assessed
for end-systolic-volume, end-diastolic-volume, ejec-
tion fraction and cardiac output, using Bland-Altman
analysis.
The assessment of myocardial extra-cellular vol-
ume by pre- and post- contrast gadolinium T1
measurement has found an increasing role in car-
diac studies, with potential applications to monitoring
the neuromuscular disease process [243]. Marty et
al. have demonstrated T1measurement using com-
pressed sensing reconstruction with a 2D multi-slice
radial MOLLI sequence which permits accurate
reconstruction of T1maps in a breath hold of 5 heart
beats, a substantial time saving over the 12–17 heart-
beats conventionally required [244]. This has recently
been extended to rapid T1-mapping of skeletal muscle
in Becker muscular dystrophy [245].
Measuring muscle contraction & strain
The MRI techniques used for dynamic muscle
contraction and strain imaging can be categorized
in three main groups: MR tagging, phase contrast
(PC), and displacement encoding with stimulated
echoes (DENSE). Tagging has been used to moni-
tor tissue displacement and deformation during active
isometric contraction in humans [246, 247] as well as
for muscle indentation in rats [248] and in humans
[249, 250]. The PC method, which enables tissue
velocity measurements in three directions, has been
applied to quantify muscle inertial forces [251], as
well as changes in fascicle length [252] and strain
rate [253, 254]. DENSE is conceptually similar to
PC, as it encodes displacement on the phase of the
signal, but it is based on the acquisition of stimulated
echoes. One of the advantages of DENSE over PC
is that it allows to encode motion over longer time

G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease 19
Fig. 9. Accelerated 3D dynamic muscle contraction imaging of the lower leg during active plantarflexion and dorsiflexion. (A) Two velocity-
imaging frames from a 12-frame movie with varying degree of image acceleration. The reference image required 10:52min of scan time and
326 repetitions of the motion task, whereas the 6.41 times accelerated scan was only 2:46 min and 83 repetitions. (B) Velocities in foot-head
(FH) direction for 4 muscles of the lower leg during the motion cycle for various degree of image acceleration. Figure was reproduced with
permission from Mazzoli et al. [259].
intervals, which is especially beneficial for skele-
tal muscles, that are characterized by low velocities
[255].
Dedicated MRI technology has been developed
to visualize and quantify muscle deformation dur-
ing active and passive contraction [256]. Dynamic
muscle imaging can be performed in real time [257]
or in a triggered fashion by mechanically restricting
muscle motion to repeatable patterns, or by using
electrical muscle stimulation [254, 258]. Because
of imaging speed restrictions, real-time imaging is
commonly limited to a 2D single slice approach, or
3D with coarse spatiotemporal resolution [250, 257].
When the motion task can be consistently repeated,
displacements or velocities can be acquired in a seg-
mented way, allowing for volumetric acquisition and
3D strain measurements. Mazzoli et al. [259] have
recently introduced a 4D PC method for quantifica-
tion of velocities and strain rate of the muscles in the
lower leg during active plantarflexion and dorsiflex-
ion, using the latest compressed sensing acceleration
technology requiring only 2:46 min of scan time
(Fig. 9).
Dynamic imaging offers great opportunities to
study musculoskeletal healthy function and dysfunc-
tion from (neuromuscular) disease, particularly when
combined with structural information and model-
ing of muscle function. PC MRI in skeletal muscles
showed non-uniform strain values in the biceps
brachii during elbow flexion, as well as in the
soleus aponeurosis-tendon complex [260, 261]. A
later study combined such information with finite ele-
ment modeling and confirmed that a different fascicle
length and curvature within the muscle are the main
cause for non-uniform strains [262]. Non-uniform
strain values were also reported in the tibialis ante-
rior [263], as well as changed strain rate values in the
gastrocnemius muscle [264].
Sinha et al. observed significant differences in
strain rates between young and elderly subjects in
the medial gastrocnemius, as well as smaller angles
between the principal strain rate and the fiber direc-
tion in the older cohort [253]. Since acquisition
was limited to a single slice, the out-of-plane com-
ponent of the strain rate could not be measured,
and was estimated under the assumption of vol-
ume incompressibility [265, 266]. Non-collinearity
of the strain directions and fiber directions and
non-uniform strains observed are believed to be
caused by architectural heterogeneity in terms of fiber
lengths and pennation angles [263]. Diffusion tensor
imaging (DTI) combined with appropriate post-
processing, allows for determination of fiber lengths
[267, 268], fiber curvatures [68, 269, 270], and pen-
nation angles [268]. We foresee that combining these
architectural parameters with local 3D strain mea-
surement will assist in a better understanding of
mechanisms of altered muscle force production and
lateral force transmission as a consequence of healthy
aging and NMD.

20 G.J. Strijkers et al. / MR Imaging and Spectroscopy Techniques for Neuromuscular Disease
MRI/PET
MRI/PET is a relatively new multimodal imag-
ing technique, which combines superior anatomical,
structural, and functional information provided by
MRI, with the greater sensitivity of PET for providing
molecular information via the detection of radiola-
beled molecular tracers. The evolution of MRI/PET
has been rapid over the past decade and has success-
fully addressed the technical challenges, which are
related to the design of PET detectors that operate in
the presence of the strong magnetic field in the MR
scanner. Initially this resulted in high-performance
preclinical prototype and research systems, but by
now both preclinical and clinical PET/MRI systems
are available from several vendors [271].
There may be added value to PET/MRI in the field
of NMD, considering that MRI/PET is practically
the combination of a standard MRI, which means
that all the MRI techniques described in this paper
may be enriched with molecular information from
PET. Currently, because of its young age, there are
only few examples of application of MRI/PET in the
neuromuscular field. Behera et al. [272] observed
increased 18F-FDG uptake in the affected nerve of
animals with neuropathic pain. Based on these find-
ings Lee et al. showed increased 18F-FDG uptake in
the ipsilateral trapezius muscle in a limited number
of patients with varying degree of severity of spinal
accessory neuropathy [273]. Priola et al. showed
that the combination of PET and MRI can pro-
vide additional information to the standard computed
tomography (CT) for the identification of myasthenia
gravis, the detection and type of the thymic abnor-
mality, and preoperative planning [274]. Haddock
et al. employed hybrid MRI/PET to determine the
relationship between relative 18F-FDG uptake and
MRI T2changes in skeletal muscles following resis-
tance exercise. They found a high correlation between
18F-FDG uptake and changes in muscle T2with phys-
ical exercise, leading to an improved insight into the
metabolic changes that occur with muscle activation
[275]. Since the family of neuromuscular disorders is
associated with a large number of muscle pathologies
the role of this technology is expected to increase as
new protocols and PET tracers become available.
SUMMARIZING PERSPECTIVES
For the seven working group three meetings of
COST action BM1304, research teams from a large
number of universities have contributed unpublished
preclinical and clinical data with the intention to dis-
cuss the potential and pitfalls of a large range of
advanced imaging techniques. The workshops were
attended by a core group of attendees and a selected
group of invited experts per topic. In this workshop
report, an overview of the topics that were discussed
was presented.
Of the techniques that were discussed, fat fraction
by MRS or Dixon imaging as well as T2relax-
ation time mapping are already used as an outcome
measure in a number of clinical trials as a pri-
mary or secondary endpoint (e.g. NCT02851797,
NCT02439216) [276]. In the research realm, a signif-
icant number of publications have appeared on DTI
and 31P; for DTI several groups have been working
together on consensus protocols for application in
skeletal muscle, whereas the commercial availabil-
ity of higher field human MR systems has resulted
in renewed interest in phosphorous MRS and MRI
by several groups. In terms of practical implementa-
tion, both (semi-) automatic segmentation and scan
acceleration are areas of great interest, and a number
of groups are actively pursuing these goals. Imaging
of fibrosis and/or other aspects of muscle tissue apart
from fatty infiltration is being regarded as a great
clinical need, but the optimal technique to assess this
remains a topic for further study.
The atmosphere during the workshops was open
and informal, and the frequency of the meetings
created common ground for new collaborations and
strengthening of existing ones. The action brought the
community together and had several tangible outputs
in the form of shared publications. The final activity
of MYO-MRI consisted of the first International Con-
ference on Imaging in Neuromuscular Disease [277].
The conference was very successful, with 200 partici-
pants and almost equal contribution from clinical and
academic partners, with more than 10% industry par-
ticipation. This illustrates the large interest in imaging
in NMD and shows that collaboration between disci-
plines in this field is very active.
ACKNOWLEDGMENTS
This work was support by The European Coop-
eration in Science and Technology (COST) action
BM1403 (MYO-MRI).
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