A MRI-Compatible Combined Mechanical Loading and MR Elastography Setup to Study Deformation-Induced Skeletal Muscle Damage in Rats

Deformation of skeletal muscle in the proximity of bony structures may lead to deep tissue injury category of pressure ulcers. Changes in mechanical properties have been proposed as a risk factor in the development of deep tissue injury and may be useful as a diagnostic tool for early detection. MRE allows for the estimation of mechanical properties of soft tissue through analysis of shear wave data. The shear waves originate from vibrations induced by an external actuator placed on the tissue surface. In this study a combined Magnetic Resonance (MR) compatible indentation and MR Elastography (MRE) setup is presented to study mechanical properties associated with deep tissue injury in rats. The proposed setup allows for MRE investigations combined with damage-inducing large strain indentation of the Tibialis Anterior muscle in the rat hind leg inside a small animal MR scanner. An alginate cast allowed proper fixation of the animal leg with anatomical perfect fit, provided boundary condition information for FEA and provided good susceptibility matching. MR Elastography data could be recorded for the Tibialis Anterior muscle prior to, during, and after indentation. A decaying shear wave with an average amplitude of approximately 2 μm propagated in the whole muscle. MRE elastograms representing local tissue shear storage modulus Gd showed significant increased mean values due to damage-inducing indentation (from 4.2 ± 0.1 kPa before to 5.1 ± 0.6 kPa after, p<0.05). The proposed setup enables controlled deformation under MRI-guidance, monitoring of the wound development by MRI, and quantification of tissue mechanical properties by MRE. We expect that improved knowledge of changes in soft tissue mechanical properties due to deep tissue injury, will provide new insights in the etiology of deep tissue injuries, skeletal muscle damage and other related muscle pathologies.

Introduction Sustained mechanical loading and deformations of skeletal muscle in the proximity of bony structures may lead to damage in the deep soft tissue layers. This category of injury is recognized as a special type of pressure ulcer by the National and European Pressure Ulcer Advisory Panel (NPUAP/EPUAP) and is commonly referred to as deep tissue injury [1][2][3][4][5]. The pressure ulcer may initially stay invisible and remain undetected for days, until the injury becomes visible as a purple or maroon discolored skin spot, thereafter rapidly developing into a severe and difficult to heal Stage III or Stage IV pressure ulcer [6,7]. The populations at high risk for developing deep tissue injury include intensive-and acute-care patients, hospice patients, as well as people that require long-term care, for example after a spinal cord injury or stroke [8]. Common risk areas are the coccyx, sacrum, buttocks, knee, and heel [9,10]. Deep tissue injury is related to increased morbidity and mortality [7,[11][12][13]. Cost-of-illness studies for pressure ulcers report a significant cost burden for society. In particular for deep tissue injuries; with wound closing times between 127 up to 155 days, typical treatment costs are 10,000-15,000 GBP per ulcer [14][15][16].
Clinicians and experienced wound nurses have noticed differences in mechanical stiffness between healthy and injured muscle in early stages of deep tissue injury development by means of palpation [17,18]. Finite element analysis (FEA) studies have shown that changes in mechanical properties of the affected tissue may lead to larger deformations and aggravated wounds [19,20]. Thus, the changes in mechanical properties of soft tissues have been proposed as one of the risk factors in the development of deep tissue injury, and analysis of tissue stiffness may be useful as a diagnostic tool for early detection [21]. However, the spatial-temporal changes in muscle mechanical properties from deformation-induced injury have never been systematically studied and quantified in a controlled setting.
A rat model for deep tissue injury has been studied extensively before, using MRI and FEA methods [22][23][24][25][26][27][28][29][30][31][32]. This well-established animal model of deep tissue injury involves the compression of the rat Tibialis Anterior (TA) muscle with a custom loading device that can be placed in an MRI scanner for imaging during the development of the wound [24]. Maps of the global transversal relaxation time T 2 were used as a readout of muscle edema as a consequence of damage and inflammation after load release [23,26,31,33]. T 2 is generally accepted as quantitative damage marker in musculoskeletal MRI [34][35][36][37][38][39]. Animal-specific FEA models, where the geometry and loading were derived from MRI, supplied estimations of the local tissue deformations. This combined experimental-numerical approach resulted in new understandings of the mechanical boundary conditions that contribute to the development of deep tissue injury [32,40].
The animal model raised new questions about the role of the muscle tissue mechanical properties in the etiology of wound development. Particularly, it is currently not known to which extent muscle stiffness alters under compression during the development of the injury and in the early healing stage, nor how stiffness changes contribute to the development of severe wounds. Since stiffness changes have been reported during tissue injury development, assessment of changes in tissue stiffness could serve a role in diagnosis of deep tissue injury [7,19,[41][42][43][44].
Magnetic Resonance Elastography (MRE) is an MRI technique for measuring mechanical properties of soft tissue by measuring mechanically induced shear waves [45,46]. The small strain harmonic shear waves, induced by an external MRI-compatible actuator, are imaged using a motion-encoded phase contrast MRI sequence. By inversion of the wave image data, Hookean viscoelastic mechanical properties, such as the shear storage modulus G d , can be quantified [47,48]. In liver diseases MRE proves to be a technique of great diagnostic and etiologic value [49][50][51][52], and is also increasingly utilized for use in the kidney, spleen, brain, breast, heart, lungs, prostate, and skeletal muscle [53][54][55][56][57][58][59][60][61]. MRE studies of skeletal muscle have primarily focused on healthy skeletal muscle to better understand the complex anisotropic viscoelastic mechanical properties and to assess changes in muscle shear stiffness during active and passive contraction [62][63][64][65][66][67][68][69][70]. In a number of studies MRE has been applied to characterize muscle disease and has proven to have added value in addition to the more conventional MRI methods; MRE is proposed as objective evaluation method to monitor the efficacy of treatments for pathologic or injured muscle [71][72][73][74][75][76][77][78].
The application of MRE to assess deep tissue injury was proposed in several papers, but it has, to the best of our knowledge not been performed yet [19,20,[41][42][43]79]. In this paper, we present an MRI-compatible setup for studying the mechanical properties of rat TA muscle during the development of deep tissue injury. The setup allows for controlled deformation under MRI-guidance, monitoring of the wound development by MRI, and quantification of tissue mechanical properties and potential changes therein by the MRE technique.

Methods Setup
A schematic illustration of the setup is shown in Fig 1. The setup is fixed in (A) a fiberglass tube with an outer diameter of 86 mm and a wall thickness of 2 mm. Inside the tube the following parts are mounted: (B) a 3D-printed water-circulated heating blanket for maintaining the rat body temperature, (C) an anesthesia mask with supply and exhaust of anesthesia gas for sedating the rat, (D) a fixation block to secure the setup in the MRI magnet, (E) the indentor, and (F) the MRE actuator. (G) A base plate with (H) electromagnetic shaker (LDS V201, Brüel and Kjaer, Royston, UK) is fixed to the wall of the Faraday cage at a safe distance from the stray field of the 7.0 T MRI magnet. The MRE actuator is brought into motion via (I) a splined transmission rod (type S518, Sullivan Products, Baltimore, US). The shaker is powered with an amplifier (QUAD 50E, Huntington, UK) and waveform generator (Agilent 33220A, Santa Clara, US) placed outside the Faraday cage. The shaker is air cooled with constant airflow. Synchronization of the MRE sequence with the shaker motion is performed via transistor-transistor logic (TTL) triggering. The TTL trigger signal is a 3V voltage level signal which is pulled to ground in case of a trigger. Signals from TTL triggering and waveform generator were visualized on an oscilloscope (Agilent InfiniVision 2000X oscilloscope, Santa Clara, US) to verify correct synchronization. The MRE actuator is roughly based on the design of Sinkus et al. that was used for MRE of colon tumors in mice and for studying an mdx mouse model of muscular dystrophy [71,80,81]. A photograph of a rat positioned in the setup can be found in the online supplement (S1 Fig). The indentation and MRE components (E and F) are shown in greater magnification and detail in Fig 2. The rat can be conveniently positioned in the setup using a right-left direction adjustable u-shaped profile with a cutout for the rat's groin, which enables positioning of the right leg of the rat under the indentor (Fig 2 indentor). The indentor can be positioned on the rat TA muscle using a movable indentor holder and rotatable half arch. The indentor is a cylindrical rod (3 mm in diameter, 36.5 mm long) with a rounded head and is composed of two, screwed-together parts. The indentor assembly contains a hollow compartment, which is filled with an aqueous solution of 1 g/L CuSO 4 for MRI visualization, inside a rigid solid rod. The MRE actuator is mounted on a dovetail profile allowing for adjustment in the axial direction. The height and depth of the indentation can be controlled by removing spacer plates or by altering the position of the MRE piston (Fig 2 MRE piston). If desired the MRE piston can be replaced by pistons of a different shape or size. The MRE piston is brought into motion via the drive rod attached to the electromagnetic shaker ( Fig 1H) and cantilever. Detailed overview of all above described parts of the indentation and MRE components can be found in the online supplement (S2 Fig).
Indentation of the rat TA muscle is achieved by positioning the indentor above the TA muscle using the rotatable half-arch, by manually pushing the rod into the muscle, and by fixating it at the desired indentation depth and orientation. The MRE piston is softly coupled to the tendon at the distal side of the TA muscle. All parts are made of PET-P (polyethylene terephthalate polyester), or from PEEK (polyether ether ketone). Both materials are MRI-compatible, and have high mechanical strength, stiffness and creep resistance.

Animal model
To investigate the feasibility of MRE based shear stiffness changes occurring during indentation induced damage 20 female Sprague-Dawley (SD) rats (11 to 13-week-old, ± 250 g, Charles River, Paris, France) were used. To optimize the settings of the experimental setup and the MR sequences 13 rats were used (results not shown). MRE measurements pre and post 2 h indentation was performed in 7 rats. In 1 of these rats MRE measurements were also performed during indentation. A duration of 2 h indentation was chosen based on previous studies, in which different durations of deformation, ischemia and reperfusion were studied [40]. The results of these experiments showed that deformation, ischemia and reperfusion all contribute to the damage process but that deformation induced damage is dominant for periods less than 2 h. Animals were housed under standard laboratory conditions with a 12 h light/dark cycle and were maintained on a standard diet and with access to water ad libitum. Rats were anesthetized with isoflurane (4.0 vol% for induction, 1.0-2.0 vol% for maintenance) in 0.6 L/min medical air. Buprenorphine (0.05 mg/kg s.c.) was administered for analgesia. Eye ointment was applied to prevent eye dehydration. The rat was placed in supine position in the MR-compatible setup. Body temperature was maintained at 35-37˚C with the heating blanket and monitored with a rectal temperature sensor. Respiration was monitored with a balloon pressure sensor placed on the abdomen and maintained in a physiological range by adjusting the anesthesia. The right leg of the rat was shaved and positioned in the u-shaped profile filled with alginate molding substance for firm fixation and susceptibility matching. To assess the TA muscle during indentation the top of the alginate cast was removed making the TA muscle accessible to the indentor. After indentation the alginate cast was closed again. Indentation of the TA muscle for 2 h took place inside the MRI scanner. The MRE actuator was coupled to the tendon at the distal side of the TA muscle. After the measurements the animals were sacrificed by means of exsanguination from the inferior vena cava. This procedure was performed under anesthesia and after administration of analgesia.

Data analysis
The MRE acquisitions provide phase data for the harmonic vibrations which are proportional to the harmonic displacements. By using an inversion algorithm the displacement data can be converted to elastograms, i.e. images representing the local tissue linear elastic shear storage modulus G d (real part of the complex shear modulus G Ã ). The inversion process assumes linear (visco-) elasticity, isotropy, and local homogeneity for all tissues, and aims to solve the following partial differential equation: Where G Ã is the complex shear modulus, ρ is the density, q is the curl of the complex displacement vector (q(x) = r × u(x)) derived from the MRE phase data, and ω is the known angular frequency. In case of a continuous monochromatic wave the displacement is defined as: with A the amplitude, and k the complex wave number. Combination of Eqs 1 and 2 produces: Combining Eq 3 with the fact that k = β + iα, the complex shear modulus G Ã can be split in terms of shear storage modulus G d and shear loss modulus G i . Full details of the inversion algorithm implemented in the ROOT data analysis framework (ROOT 5.34/17, CERN, Meyrin, Switzerland) were previously described by Sinkus et al [83][84][85].
Quantitative T 2 maps were obtained by pixel-wise fitting of the MR signal to S ¼ S 0 e À TE=T 2 (Mathematica 10, Wolfram Research, Champaign, USA). Pixels with goodness of fit R 2 < 0.7 and SNR < 4 were excluded from further analysis.
Region-of-interest (ROI) based analysis of the elastograms and T 2 maps of the 6 rats measured before and after 2 h of indention was performed (Matlab R2016a, The Mathworks, Inc., Natick, Massachusetts, USA). An ROI was defined by manually outlining the TA muscle in the center slice. Mean values of G d , displacement amplitude A tot and T 2 were determined in this ROI before and after indentation. The percentages of significantly elevated G d and T 2 pixels were calculated by thresholding of pixels with values increased by 2 standard deviations as compared to the mean values pre-indentation. Mean, standard deviation (sd), and coefficient of variation (CV) were calculated to quantify the group average and variation at baseline, during and after indentation. Mean values of G d and T 2 were also determined in a circular ROI of 4 x the diameter of the indentor, positioned at the center of indentation. Paired sampled t-tests were conducted to test for significant differences between before and after indentation (Origin 2015, OriginLab Corporation, Northampton, USA).

Results
The experimental methods presented here combine soft tissue indentation with MRE to assess the skeletal muscle shear storage modulus G d before, during and after deep tissue injury development. Both the indentation device as well as the MRE actuator are designed to allow flexible positioning on top of the TA muscle in the hind leg of a rat and can be used before, during, and after indentation inside a small animal MRI scanner. The combination of an RF surface receiver coil and volume transmission coil resulted in high-quality MRI and MRE images of the whole TA muscle before, during, and after 2 h indentation. Fixation with alginate assures a good leg fixation with a perfect anatomical fit. In addition, it improves susceptibility matching enabling fast imaging with EPI-based sequences.

MRI
The manually set indentation depth can accurately be determined since the indentor is visible in the MR images ( Fig 3B). Indentation depth was varied between rats resulting in different degrees of deep tissue injury in the TA muscles. Fig 3 shows representative axial anatomical T 1 -weighted images of a rat leg (A) pre, (B) during, and (C) post indentation together with the corresponding quantitative (D-F) T 2 maps of a slice through the center of the region of deformation. The indentor pressing in the TA muscle is clearly visible in the T 1 -weighted images (arrow). The tibia bone, fibula bone, and skin contour provide anatomical landmarks from which indentation depth can be estimated. Compression of the TA muscle between the indentor and tibia bone is seen in (B). Elevated T 2 values compared to pre indentation are seen in the T 2 map post indentation (Fig 3F). These increased T 2 values are indicative for the formation of edema and skeletal muscle damage as validated before [25,26,40].
Several axial slices of the 3D ultra-short echo time images of the alginate cast around the leg before, during, and after indentation are shown in Fig 4. Due to its short T 2 relaxation time, the alginate is only visible on the ultra-short echo time scans and not on the regular T 1 -weighted and T 2 -weighted scans. The presence of alginate therefore does not interfere with the anatomical scans and the MRE measurements, and improves image quality by providing good susceptibility matching between the leg and its surroundings. During indentation (Fig 4 during) the top of the alginate cast was removed. Newly applied alginate is brighter on the post-indention ultrashort echo time scan (Fig 4 post). The cast closely follows the anatomy of the leg and little displacement of the leg between the three conditions is seen. The contours and the rigid fixation by the cast provide essential boundary conditions for FEA calculations of deformation related strains, as previously demonstrated [86].  Table 1. The displacement amplitude was in the range of 0.5-10 μm (mean 2 μm).

Magnetic resonance elastography
In Fig 6 eight offsets of coronal SE-MRE magnitude images (mag) and wave images (θ 1θ 8 ) are shown before, during, and after indentation. Distinct wave propagation was observed during the equally distributed time offsets θ 1 -θ 8 of the 900 Hz wave with a time resolution of 0.1389 ms. In Fig 6B the dark spot in the TA corresponds to the location of the indentor. The wave pattern in (Fig 6B1-6B8) during indentation was distinctly different compared to pre indentation (Fig 6A1-6A8). The changes in the wave patterns relate to changes in tissue elasticity and geometry due to indentation, and also by reflection of shear waves at the indentor surface. Post indentation (Fig 6C1-6C8) the waves show a different pattern compared to pre indentation (Fig 6A1-6A8). A distinct change in the wave pattern was observed due to altered tissue stiffness of the wound caused by indentation.
In Fig 8, representative MRE elastograms (Fig 8A) before, as well as (Fig 8B) after 2 h of indentation are shown together with the corresponding T 2 -maps (Fig 8C and 8D). Increased shear storage modulus G d was observed in the TA in the elastogram that was measured following 2 h of indentation. T 2 -maps after indentation revealed a large region of elevated T 2 values in the TA muscle, indicative for edema and muscle injury. The region of increased shear storage modulus G d was less diffuse and smaller compared to the region of elevated T 2 . The region of high shear storage modulus G d highlights a central focus with severe muscle damage, whereas T 2 elevation occurs in a larger region with edema surrounding the injury. The MRE elastograms of the other analyzed animals show the same consistent results.
For all 6 animals measured before and after end of indentation, the shear storage modulus G d maps of the TA muscle were reconstructed at four time points: before, 30 min after, 60 min after, and 90 min after end of indentation. Two animals had missing data at the 90 min after time point. The T 2 maps were calculated at two time points: before and 45 min after end of indentation. T 2 and G d cannot be measured at the same time, particularly because of the long,  Tables 2 and 3. G d Standard deviation (0.1 kPa) and coefficient of variation (3%) were low at baseline in non-damaged TA muscle.  Averaged across all animals, the mean T 2 of the TA muscle at 45 min after end of indentation was significantly higher than before indentation (p < 0.01). A trend of increased mean shear storage modulus G d after indentation was apparent. Mean shear modulus G d at 30 min after indentation was significantly higher than before indentation (p < 0.05).
The percentage significantly elevated T 2 and G d pixels after end of indentation was significantly higher compared to before indentation (T 2 ; p < 0.01, G d ; p < 0.05 at 30 min after). The percentage elevated T 2 pixels after indentation was significantly higher compared to the percentage elevated G d pixels after indentation at all three time points (p < 0.01, p < 0.01, p < 0.05). In both T 2 and G d maps before indentation a comparable percentage of significantly elevated pixels was found (3.6 ± 0.6% for G d and 4.0 ± 0.7% for T 2 ).

Discussion
In this paper we have presented a combined Magnetic Resonance (MR) compatible indentation setup and MR Elastography (MRE) setup to study the spatial-temporal changes in mechanical properties that result from deformation-induced injury. The setup was designed in such way that tissue indentation and the MRE methods can either be used separately or simultaneously. The setup was successfully tested in a deep tissue injury rat model by damageinducing indentation and MR Elastography measurements of the rat TA muscle. Results show that it is possible to quantitatively measure muscle mechanical properties over time. The elastograms after 2 h of indentation revealed a local increase in shear storage modulus G d after indentation. The spatial extent of change in stiffness was also evaluated by quantifying the percentage pixels with significantly enhanced G d values with respect to baseline. Quantifica  These results show the high potential of the setup to measure non-invasively the quantitative mechanical properties related to deep tissue injury in vivo.
There are still options to further improve the indentation setup. Firstly, the setup does not allow for an indentation controlled from outside the scanner; i.e. currently the setup has to be taken out of the scanner bore to apply, adjust or remove the indentor. This manual operation, especially if multiple indentation depths are of interest, is time consuming and repositioning of the setup following indentor adjustment may cause misalignment of image data sets. Image registration can be employed to correct for such misalignment. However, a controlled indentation device would be preferred. Also, the indentor has no force sensor incorporated in the current implementation. Although indentation control and force sensing can be achieved inside an MRI scanner (e.g. see [90] for a clinical scanner set-up) these were not implemented in the current study due to the space challenges in a small animal MR scanner. Knowledge of the indentation force boundary conditions, combined with MRI data of the undeformed and deformed configuration would enable inverse FEA based investigation of hyperelastic tissue mechanical properties [91]. The MRE part was designed to be flexible in feet-head and anterior-posterior direction, but is limited in left-right direction due to the fixed position of the electromagnetic shaker driving the MRE part. In the positioning of the leg of the animal this limitation was taken into account, which resulted in successful coupling of the MRE actuator piston in all animals. Although only MRE measurements of 900 Hz were presented here, higher and lower frequencies (100 Hz-  1500 Hz) were also evaluated. Frequencies between 1000 and 1500 Hz did not always result in sufficient displacement (< 0.5 μm displacement) to produce waves through the whole TA muscle. At 900 Hz a decaying wave with displacement in the range of 0.5-10 μm (mean 2 μm) was induced. This displacement amplitude is comparable to other MRE studies [87,92]. We observed that the MRE actuation system behaves frequency dependent; less power was needed with lower frequencies to generate sufficient displacement at the MRE piston. In addition, tests with a prototype of the current MRE actuator showed eigenfrequencies with different designs of cantilevers resulting in higher MRE piston amplitudes [80]. Altogether, measurements at each frequency in the range 100 Hz-1500 Hz are possible but further optimization to obtain the correct amplitude at the MRE piston is required. T 2 elev., percentage of pixels with elevated T 2 (> mean T 2 before + 2 x sd); Ind.,circular ROI of 4x indentor's diameter *** p < 0.01 with paired sampled t-test. doi:10.1371/journal.pone.0169864.t003 The alginate cast allowed proper fixation of the animal leg with anatomical perfect fit, provided boundary condition information for FEA and provided good susceptibility matching. The use of alginate has advantages over plaster [93][94][95]. Alginate does not clot, is easy to prepare, can be molded anatomically around the rat leg and is very easy to remove without any remnants. In addition, the surrounding alginate provided appropriate susceptibility matching beneficial for MRI, i.e. it reduced artefacts and improved image quality. For this reason, the MRE measurements during indentation, where the top of the alginate cast was removed, was performed with SE-MRE instead of SE-EPI-MRE. SE-MRI is less sensitive to field inhomogeneities but more time consuming. A typical SE-MRE scan with three encoding directions plus reference will take approximately 1h. To capture dynamic changes in tissue mechanical properties SE-EPI-MRI is therefore preferred.
After application of the indentor to the TA muscle inside the MRI scanner, increased signal intensity in the TA muscle was observed on T 2 maps as a result of edema, resulting from the induced damage and associated inflammatory response [22][23][24][25][26]. ROI analysis showed a significant increase in T 2 after end of indentation. A significant increase in the percentage of significantly elevated T 2 pixels was observed after indentation. The area of pixels with elevated T 2 at 45 min after indentation was significantly larger than those with elevated G d at 30, 60 and 90 min after indentation. Using T 2 alone as specific damage area indicator might result in overestimation of the size of the damaged area, as in the initial injury phase fluids might diffuse between the fibers proximally and distally from the wound. The size of the affected region in T 2 in comparison to G d may provide insights in the causes and consequences of deformation induced muscle damage.
In the current study deep tissue injury and skeletal muscle tissue were of interest. Shear moduli data for the more superficial adipose tissue were not derivable in the current study because fat suppression techniques were employed for the MR acquisitions, enabling minimization of chemical shift artefacts. If such data is however desired non-EPI MRE sequences can be employed without the use of fat suppression.
The employed MRE inversion algorithm assumes Hooke's law of linear (visco) elasticity (as is valid for the small strain vibrational motions) and assumes isotropic tissue behavior. However, as is the case for muscle tissue, the epi-, peri-and endomysium structure causes the tissue to be anisotropic. Furthermore, damage and deformation may independently cause changes in elastic symmetry and therefore induce anisotropy. Since the inversion algorithm assumes isotropy the reconstructed shear moduli information may potentially be affected by any underlying anisotropy. Exactly how anisotropy is reflected in analyses based on isotropic inversion methods is currently unknown, and may also relate to the directions wave happen to locally follow. However, when combined with MRI-compatible force sensing, diffusion tensor-based fiber architecture assessment, and inverse FEA, the presented methods do offer a means to assess such effects in the future [86,90,[96][97][98]. This is because the initial anisotropy and non-linear elastic behavior would be accurately reflected in FEA based on such measurements, and elastic tensor data for any state of deformation can be directly compared to MRE findings. In principle such analysis can be coupled with FEA incorporating constitutive formulations of damage and swelling.
In conclusion, we have presented an MRI-compatible setup for studying the mechanical properties of rat TA muscle. The setup allows for controlled deformation under MRI-guidance, monitoring of the wound development by MRI, and quantification of tissue mechanical properties by MRE. We expect that improved knowledge of changes in soft tissue mechanical properties due to deep tissue injury, will provide new insights in the etiology of deep tissue injuries, skeletal muscle damage and other related muscle pathologies. Future studies will focus on the pathophysiological processes behind the changes in mechanical properties of damaged skeletal muscle and evaluation of the changing mechanical properties over time.

Ethical approval
All animal experiments were approved by the Animal Care and Use Committee of Maastricht University, Maastricht, The Netherlands (protocol 2013-047, Maastricht University, Maastricht, The Netherlands) and performed in accordance with the Directive 2010/63/EU for animal experiments of the European Union.