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Global and local tension measurements in biomimetic skeletal muscle tissues reveals early mechanical homeostasis

Tension and mechanical properties of muscle tissue are tightly related to proper skeletal muscle function, which makes experimental access to the biomechanics of muscle tissue formation a key requirement to advance our understanding of muscle function and development. Recently developed elastic in vitro culture chambers allow for raising 3D muscle tissue under controlled conditions and to measure global tissue force generation. However, these chambers are inherently incompatible with high-resolution microscopy limiting their usability to global force measurements, and preventing the exploitation of modern fluorescence based investigation methods for live and dynamic measurements. Here, we present a new chamber design pairing global force measurements, quantified from post-deflection, with local tension measurements obtained from elastic hydrogel beads embedded in muscle tissue. High-resolution 3D video microscopy of engineered muscle formation, enabled by the new chamber, shows an early mechanical tissue homeostasis that remains stable in spite of continued myotube maturation.

Skeletal muscle is one of the most abundant tissues in the human body and is crucial for 2 essential functions such as limb movement, thermogenesis and maintaining 3 posture [9,26,38]. Skeletal muscle atrophy is debilitating and is a typical outcome of 4 diseases like muscular dystrophies but also aging. Characterizing the biomechanics of 5 muscle tissue is a key element to understand both, muscle development and muscle 6 degeneration. While this characterization has been largely achieved on the global tissue systems are of limited use in contraction studies due to randomly oriented myotubes 23 and the inability to maintain long-term cultures [14,22,39,49]. In contrast, modern 24 reconstituted 3D in vitro skeletal muscle systems are demonstrated to be an efficient 25 tool for rapid and reliable drug screening [2,54] and allow for testing of personalized 26 treatments on cells harvested from individual patients. Besides these medical 27 advantages, such functional muscle tissues were reported to successfully mimic native 28 muscle tissue with long-term structural integrity [23,27] and allow new insights and 29 fundamental knowledge of muscle tissue development, force generation during 30 contraction as well as the phases of disease onset and 31 progression [2,11,19,24,25,31,32,41,48,51]. In culture platforms allowing for in situ 32 force measurements, 3D skeletal muscle tissues self organize between mm sized posts 33 around which muscle precursor cells are seeded together with an extracellular matrix. 34 These chambers are commonly based on polydimethylsiloxane (PDMS) molds that can 35 be tuned in their elastic properties and are soft enough to allow measurement of muscle 36 contraction forces by recording the post deflection. Such non-invasive measurements of 37 force generation in reconstituted muscles are indeed based on the elastic properties of 38 PDMS [1,29]. A downside of PDMS is that the poor optical properties paired with 39 routinely used material thickness prohibits high numerical aperture objectives that are 40 necessary for high resolution microscopy, and hence make modern 3D microscopy 41 methods like confocal and spinning disk microscopy on living tissue impossible. Instead, 42 to study 3D tissues with high resolution requires fixation and removal from the culture 43 device, thus preventing research on questions related to dynamic processes. Hence, they 44 are of only limited use for the investigation of spatiotemporal research questions such as 45 dynamics of cell-cell interaction or myoblast fusion. Moreover, PDMS has an immense 46 capacity to absorb chemicals and proteins and is therefore unsuitable for serum-free 47 biocompatible oil microdroplets to evaluate cell-generated forces in living tissue for the 66 first time [10] and very recently, deformable PAA beads were used to determine local 67 tension on a cell scale within cancer spheroids, zebrafish embryos and during 68 phagocytosis [13,28,53,55]. Contrary to oil microdroplets, PAA beads are compressible 69 and are therefore able to reveal isotropic tissue pressure. 70 In this study, we open new ways to study the global and local mechanical properties 71 of biomimetic skeletal muscle tissue. We first present a novel technique to culture in 72 vitro skeletal muscle tissues in 3D that inverses the geometry, so that the tissue is 73 located closely to a microscopy glass slide, which allows real-time high resolution 74 imaging, while still self organizing around thin posts of polymethylmethacrylate 75 (PMMA) that are connected to the top of the chamber and thereby allowing both, 76 accurate quantification of global forces via post deflection and high resolution 3D 77 fluorescence microscopy. Next, we incorporate custom-made elastic PAA beads into the 78 biomimetic muscle tissues to trace local tension within the tissues during development. 79 both, the global and local forces match in magnitude, as expected. Interestingly, we 80 observe an immense increase in local tension inside muscle tissues in vitro within the 81 first week of development which remains at a similar level in the following week, 82 although myotube diameter continues to increase. Notably, global muscle tissue 83 pre-tension did not increase in the second week of differentiation, either. By enabling 84 real-time high resolution microscopy on 3D in vitro muscle tissues for the first time, we 85 envision our novel reconstituted muscle tissue device will be an advantage to the skeletal 86 muscle research community by enabling the dynamic study of (sub-)cellular events. posts. Prior to use, the molds were sterilized using 70 % ethanol and the wells were 97 coated with a Poloxamere solution in ddH 2 O over night at 4 • C (5 % Pluronic R F-127, 98 Sigma, St. Louis, USA) to render the surface non-adhesive. MyoTACTIC PDMS molds 99 were fabricated exactly as described in [1].

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2D and 3D skeletal muscle progenitor culture 101 The C2C12 mouse muscle progenitor cell line was obtained from ATCC. The AB1190 102 and AB1167 immortalized human muscle progenitor cell line [33] was obtained from 103 Vincent Mouly (Paris, France). Channelrhodopsin-2 was stably introduced into the 104 AB1190 cell line by lentiviral transduction as described on the [8] webpage. The Fubi-ChR2-GFP plasmid was a gift from Edward Boyden (Addgene plasmid 106 #22051) [7]. LifeAct-GFP and H2B-mCherry was stably introduced into the AB1167 107 cell line by lentiviral transduction. The pLenti6-H2B-mCherry plasmid was a gift from 108 Torsten Wittmann (Addgene plasmid #89766) [37]  To produce PAA beads a water-in-oil emulsion approach was used. Since in this study 136 predominantly 2.1 kPa (Young's modulus) beads were utilized, the fabrication and 137 characterization for this elasticity is described, representatively. To generate beads with 138 other elasticity, adapted concentrations of monomers and crosslinkers are required.

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The water phase was prepared by mixing two parts of acrylamide solution (40 % v/v, 140 Sigma, St. Louis, USA) with one part of N,N'-methylenebisacrylamide solution (2 % 141 v/v, Sigma, St. Louis, USA). This PAA solution was then diluted to 13.5 % (v/v) using 142 a 65 % (v/v) phosphate buffered saline (PBS, Sigma, St. Louis, USA) to obtain a 143 pre-bead-mix. The mechanical properties of the PAA beads can be tuned by changing 144 the dilution of the pre-bead-mix. The oil phase was composed of 3 % Span R 80 (Sigma, 145 St. Louis, USA) in n-hexane (Merck, Darmstadt, Germany). Shortly before 146 polymerization the pre-bead-mix as well as the oil phase were degassed for 10 min at 50 147 mbar. The free-radical cross-linking polymerization of the PAA solution was initiated by 148 adding 1.5 % (w/v) ammonium persulphate (APS, AppliChem, Darmstadt, Germany) 149 and the pH-value of the solution was neutralized using NaOH solution. The emulsion 150 was then generated by injecting the pre-bead-mix into the n-hexane with a 100 µl 151 Hamilton syringe (Hamilton, Reno, USA). Hereafter, the polymerization was catalysed 152 by adding 3 % (v/v) N,N,N',N'-tetramethylethylenediamine (TEMED, Sigma, St. Louis, 153 USA) and the emulsion was degassed again for 6 min. The supernatant was discarded 154 and the polymerization was kept at 85 • C for 10 min. To reach end of gelation the beads 155 were incubated at room temperature over night. Next day, the beads were washed five 156 times with n-hexane and transferred to 65 % PBS. Finally, the beads were labeled Mechanical characterization of PAA beads 162 To characterize the mechanical properties of the elastic PAA beads the shear modulus G 163 and bulk modulus K were determined. The shear modulus was measured via active 164 microrheology using a custom optical tweezers setup. For this purpose, smaller beads 165 with a diameter of 1 µm (Latex beads, Sigma, St. Louis, USA) were embedded into the 166 PAA beads. The complex shear modulus, G * = G + iG , was measured for varying 167 oscillation-frequency of the 1 µm beads [3]. The shear modulus G used for further 168 analysis was estimated by averaging shear storage moduli G for low frequencies from 1 169 Hz to 10 Hz. The bulk modulus K was measured via an osmotic pressure approach 170 using 2 MDa dextran molecules (Dextran T2000, Pharmacosmos, Wiesbaden, Germany) 171 which were dissolved in the PAA beads solution. Since the pore size of the PAA gel is 172 much smaller than the hydrodynamic radius of the dextran, an osmotic pressure 173 p = K ∆V V acts on the PAA beads and gives rise to a compression ∆V /V proportional to 174 the bulk modulus K. The exerted osmotic pressure depends on the concentration of 175 dextran and has been calibrated previously [35]. Here, three different concentrations (60 176 g/l, 85 g/l and 100 g/l) were used, corresponding to osmotic pressures of 6 kPa, 12 kPa 177 and 18 kPa, respectively. The distribution of PAA bead diameters in the three solutions 178 was measured and compared to the uncompressed situation by segmenting images of 179 PAA beads with the find-edges-function of Fiji and fitting ellipses to the edges. The  In order to reconstruct forces acting in the tissue, the deformation of embedded PAA 184 beads were quantified using a two step process implemented in a graphical user interface 185 developed in Matlab (Mathworks, US). First, the beads were segmented and fitted using 186 spherical harmonics. Then an analytical solution to the elastic problem was applied to 187 gain the forces on the beads. For bead segmentation, a small volume around each PAA 188 bead was cut out of the 3D image stack and if desired the effective pixel density was 189 interpolated by a factor of 2 for display purposes. In the cutout a threshold intensity 190 was determined according to segmentation parameters that were typically set to clipping 191 level (50 %) and peak intensity voxels (3 %) and the image was then transformed into a 192 binary image. Subsequently, the surface was determined by extracting voxels at the 193 edge of the binary volume and the algorithm returns the coordinates (x i , y i , z i ) of these 194 voxels that do form the surface of the PAA bead. In order to quantitatively capture the 195 deformation, spherical harmonics were fitted to the segmented bead surface using a  In detail, following this approach, the coordinates of the segmented surface (x i , y i , z i ) 205 were transformed into spherical coordinates r(θ, φ). The surface defined by the spherical 206 harmonics

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was then fitted to the segmented surface by rotating r sh (θ, φ) to align with the 208 major direction of deformation of r(θ, φ) and minimizing the residual (r sh − r) 2 by 209 varying c 00 and c 20 . Thus, a sphere of radius c 00 / √ 4π with uniaxial deformation 210 proportional to c 20 was fitted to the deformed PAA bead. This fitting procedure was 211 repeated five times with randomly picked starting values in each case. The parameter 212 combination with the lowest residual was taken as the result. The sought force dipole F 213 acting on the PAA bead was finally calculated by with the scaling factors N r and N θ as well as e z being the unit vector on the main 217 deformation axis. Equation 2 was derived following classical elasticity theory [30], 218 resulting in expression for the stress on the sphere that was then integrated over the 219 half sphere to get the force dipole acting on the PAA bead. The axis of the force dipole 220 relates to the axis of rotation when fitting the spherical harmonics. deformation in x is considered and y,z are assumed to be fixed. Hooke's law gives The constraint deformation enforces that only the component u xx is different from zero, 282 hence which means that the stresses are not independent, but that In first order, the tissue around the bead remains mostly undeformed for the small 287 deformations measured, hence the absolute stress σ yy acts on the beads. For a Poisson 288 ratio ν = 0.5 the tensional stress in x direction is equal in magnitude to the measured 289 stress in y and z direction. Therefore we can directly compare the measured local tension derived by the bead 291 deformation of 2.4 ± 0.9 kPa to the global tension of 1.8 ± 0.67 kPa, and confirm that 292 these match within the error. All results are presented as mean ± SD and statistical differences of experimental 295 groups were analysed by unpaired t-test using GraphPad Prism 6.0 software, where p < 296 0.05 was considered as significant. Significances were subdivided into three levels: (p = 297 0.05 -0.01), (p = 0.01 -0.001), (p < 0.001) The number of biomimetic muscle tissues 298 (N) and analysed PAA beads (n) is indicated in every figure legend. In Figure 4E 'n' 299 refers to the number of measured myotubes. systems [1,2,31]. To demonstrate the imaging quality and observation capacity of this 327 chamber design, C2C12 mouse myoblasts were seeded within a Geltrex R -Fibrin matrix 328 at defined ratios and cultured in growth medium for two days allowing for equilibration 329 to the 3D environment. The medium was then exchanged to a low serum formulation to 330 support differentiation (i.e. fusion to form multinucleated 'myotubes') for a period of up 331 to 2 weeks ( Figure 1C). During remodeling, the tissues anchored to each end of the 332 posts, which act as artificial tendons ( Figure 1B images of transverse 3D muscle tissue cryosections were prepared that displayed 341 myotubes which are consistently aligned close to each other and are evenly distributed 342 throughout the entirety of the tissue ( Figure 1E).

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To demonstrate the capacity of time-lapse high resolution microscopy on living 3D 344 skeletal muscle tissues, we generated a Lifeact-GFP and H2B-mCherry labeled human 345 myoblast cell line (AB1167-Lifeact-GFP-H2B-mCherry) and recorded a timeseries of 10 346 day old 3D skeletal muscle tissue dynamics using a spinning disk microscope. The  Figure 2B). These measurements are consistent with force measurements obtained in 383 PDMS based chambers [1]. Our post deflection analysis software is able to trace exerted 384 forces on the post over time from acquired videos with minimal background signal of ± 385 8.5 µN ( Figure 2C). For optogenetic twitches we observe a roughly one second 386 contraction phase before the in vitro muscle tissue returned to the relaxed state. By 387 comparison, a tetanus contraction induced by 2 mM ACh reached its peak of contractile 388 force more slowly and the time elapsed before return to the relaxed state was almost 25 389 seconds ( Figure 2D). 390 We further established a method to evaluate pre-tension during skeletal muscle   The bottom portion is affixed to a microscopy grade glass. B, Images of remodeled C2C12 muscle tissues at 7 days of differentiation anchored to the end of the posts captured to provide a view looking up into a well in which the top and bottom parts are fashioned together (left; 4X objective) or from the side looking at the top part to visualize 6 pairs of posts with tissues at the bottom. C, Schematic workflow used to raise 3D skeletal muscle tissues in vitro. D-E, Representative confocal microscopy longitudinal (D, whole mount, flattened stack, 40X water immersion objective) and transverse (E, cryosection, single snap, 40X water immersion objective) images of multinucleated myotubes within a 14 day old muscle tissue immunostained for sacomeric alpha-actinin (SAA, green) and counterstained with Hoechst 33342 to visualize nuclei (blue). The cross section of the tissue was measured to be 0.17 ± 0.03 mm 2 . F, Timeseries of nuclear motion withing developing 3D skeletal muscle tissue 10 days after differentiation, demonstrating the possibility of high resolution imaging during living tissue formation. Lifeact-GFP (green) and H2B-mCherry (red) was stably introduced into AB1167 cells. tissue development in vitro. Specifically, we treated C2C12 skeletal muscle tissues with 392 10 % SDS to dissolve the myotube membranes, thereby releasing the pre-tension To determine the local tension in the muscle tissue, we implemented elastic hydrogel 404 beads as tension sensors, an approach first used by others in different 405 systems [13,28,53,55]. We custom-made elastic PAA beads by a water in oil emulsion 406 approach ( Figure 3A). The use of PAA beads as tension sensors requires precise 407 mechanical characterization. Using optical tweezer based active microrheology on 1 µm 408 11/23 polystyrene particles embedded within the PAA beads, we determined a shear modulus 409 of G = 710 ± 270 Pa by averaging shear storage moduli G for low frequencies from 1 Hz 410 to 10 Hz ( Figure 3B). Additionally, the bulk modulus was determined to be K = 41 ± 6 411 kPa by an osmotic pressure approach ( Figure 3C). Here three different concentrations of 412 2 MDa dextran (60 g/l, 85 g/l and 100 g/l) were applied to the elastic beads, which 413 corresponds to osmotic pressures of 6 kPa, 12 kPa and 18 kPa, respectively. PAA bead 414 diameter in the three solutions was measured and compared to the corresponding 415 uncompressed situation. Consequently, the Poisson ratio ν = (3K − 2G)/(6K + 2G) was 416 determined as ν = 0.491 ± 0.005, which is consistent with previous reports of Poisson's 417 ratio in PAA [5]. For the force calculation, the Young's modulus E = 2G(1 + ν) of the 418 PAA beads used in this study was hence determined to be E = 2.1 ± 0.8 kPa. 419 We next used the mechanical properties of the PAA beads to determine the forces 420 exerted on the beads, from images where bead deformation was measured. Previous 421 studies by others used finite element methods for the analysis [53], which have the has an user friendly interface ( Figure 3E) and is available on GitHub [58].

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Local tension within in vitro skeletal muscle tissues during 435 development 436 We then conducted a quantitative study of local tension within 3D skeletal muscle  We first imaged the PAA beads contained within the living muscle microtissues via 451 spinning disc microscopy on the same day we seeded the C2C12 cells mixed within the 452 ECM matrix into the culture device. The culture chamber is crucial to obtain 453 fluorescence images of the beads that have sufficient quality for subsequent BDA. We 454 repeated this tension investigation in living tissues at 7 and 14 days of differentiation, 455 respectively. Whereas the beads were fairly spherical directly after seeding, we observed 456 significantly higher deformation, and therefore tension, within the tissues after one week 457 of differentiation ( Figure 4C). The average internal tissue tension at this time-point was 458

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A C D B E 5 µm Figure 3. PAA bead characterization and analysis software. A, Schematic workflow of bead fabrication. Using a syringe a water in oil emulsion is created and the acryl amide solution is polymerized to produce elastic PAA beads. B, Graph showing the determination of the shear modulus of elastic PAA beads, which was measured for varying oscillation-frequency of 1 µm beads. C, Graph depicting the determination of the bulk modulus of elastic PAA beads, which was measured by an osmotic pressure approach using dextran. D, Segmentation of a deformed PAA bead. The surface (red) was segmented and a linear combination of spherical harmonics of degree n = 0, 2 and order m = 0 (blue) were fitted to the surface. E, Screenshot of PAA bead deformation analysis software front-end. Providing a point spread function enables the automatized program to deconvolve the image and determine the edge position of an elastic bead (green) in a given cutout box. The bead's shape is then fitted by spherical harmonics, which are used to calculate the main axis of the force dipole (blue line) and the corresponding tension analytically.
slightly more than 2.4 ± 0.9 kPa, and maintained a similar level of tension following 2 459 weeks of differentiation. As a control experiment, we measured the tension of beads 460 embedded in ECM matrix only made of geltrex and fibrin, and without cells. We found 461 no significant difference between beads in ECM matrix without cells after seven days 462 when compared to beads measured directly after seeding. This demonstrates that the 463 observed bead deformation is indeed due to remodeling and differentiation of the cells in 464 the ECM matrix. To further control these measurements, we cultivated 3D muscle 465 tissues with very stiff 150 kPa beads integrated throughout. At this high stiffness, we 466 expect that the measured forces cannot deform the beads sufficiently for detection with 467 microscopy. Indeed, in this case we did not observe deformation of the PAA beads after 468 one week of differentiation as compared to the day of tissue seeding (Supplementary 469 Figure 2). Consistently, when 7 days old tissues with PAA beads embedded are 470 dissolved using 5% SDS, the beads return to their initial spherical shape ( Figure 4C).

471
To further validate our results acquired via spinning disk microscopy in PMMA/glass 472 molds, we produced C2C12-derived muscle microtissues in the previously described 473 MyoTACTIC PDMS mold [1] and imaged fixed tissues that were removed from the 474 mold and imaged in a FEP capillary using Multiview Light-Sheet microscopy to provide 475 a comparison. Multiview Light-Sheet microscopy has the advantage of isotropic 476 resolution, and hence checks potential problems due to spinning disk microscopy and 477 deconvolution. We first conducted studies and confirmed that fixation of in vitro 478 skeletal muscle tissues per se did not impair embedded PAA bead shape significantly by 479 comparing images of the same bead before and after fixation in our PMMA/glass molds 480 (Supplementary Figure 3). Consistent with bead data obtained using spinning disk 481 confocal microscopy in our PMMA/glass device, we observed a significant increase of 482 local tension exerted on the elastic beads after 7 days of skeletal muscle microtissue 483 differentiation of slightly more than 2.3 ± 0.9 kPa ( Figure 4D), but again, no significant 484 difference was detected after a second week of differentiation. Importantly, myotube 485 diameter increased significantly from 10.5 ± 1.4 µm at one week of differentiation to 486 13.6 ± 2.5 µm in the second week ( Figure 4E). This shows that while the initial phase 487 of tissue compaction myotube differentiation influences beads tension, a further increase 488 in myotube diameter does not result in more tension on embedded PAA beads.

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The emergent importance of mechanobiology has demonstrated that mechanical 491 interaction between cells as well as stiffness and tensile forces provide important 492 signalling elements for cell biology and cell fate. Muscle tissue is intrinsically exposed to 493 large forces and rapid changes in tension and stiffness. Hence, it is to be expected that 494 the mechanical properties of the environment are of particular relevance for the 495 homeostasis of muscle tissue. To study these interactions, not only the precise 496 observation of self organized 3D tissue is required, but also a non-invasive readout of 497 global and local forces is important.  tissue with PAA beads embedded (red) and immunostained for sacomeric alpha-actinin (green) and counterstained with Hoechst 33342 to visualize nuclei (blue). B, Representative confocal snap of a transverse section of a 14 day old C2C12 tissue. Color coding is identically equal to A. C, Bar graph of local tension within C2C12 muscle tissues at the day of seeding (N=7, n=28), and after one (N=4, n=28) and two weeks of differentiation (N=4, n=29). Beads were imaged in the context of living tissues raised in PMMA molds for BDA utilizing spinning disk microscopy. Also shown is negative controls in which tension was measured in the cell-free geltrex/fibrin matrix alone (N=3, n=20) as well as following the dissolution of tissues using 5 % SDS (N=3, n=10). D, Bar graph of local tension within C2C12 muscle tissues raised in the MyoTACTIC PDMS microtissue platform that were fixed and removed for imaging using Multiview Light-Sheet microscopy on the day of seeding (N=13, n=39), and after one (N=17, n=48) and two weeks of differentiation (N=19, n=46). E, Bar graph of myotube diameter analysis conducted on C2C12 muscle tissues at one (N=3, n=49) and two (N=3, n=57) weeks of differentiation. Each myotube diameter data point reflects the average of three measurements per myotube.

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prevents the protein sponge effect, while the geometry allows positioning a glass 509 coverslip in close proximity to the self-organized muscle tissue, thus enabling high 510 resolution microscopy.

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Changing the material and adapting the geometry to still enable global force 512 measurements while gaining high quality optical access is a key advancement. It now 513 allows time resolved high resolution measurements of in vitro muscle tissue development, 514 myotube formation and fusion, without the need of fixation and tissue removal as 515 necessary in the previous designs. Using fluorescent protein tagging it was directly 516 possible to gain high resolution 3D images of actin networks with a high time resolution. 517 Such fluorescent protein based life stains in combination with this approach will help 518 understanding the dynamic interactions between muscle cells during differentiation, 519 fusion, myotube formation and maturation. Therefore, we enable real-time high 520 resolution imaging during cultivation of in vitro muscle tissues for the first time which is 521 of great benefit for various future research issues such as myoblast fusion or myotube 522 maturation. A previous 2.5D myotube culture approach was indeed able to monitor first 523 myotube maturation and nuclear movement to the periphery of skeletal myotubes by 524 dynamic high resolution imaging [44]. However, this approach cannot draw conclusions 525 within a 3D system, is not capable to conduct functional or contraction studies and is 526 furthermore unable to quantify force generation of the tissue, which all can be combined 527 using our novel approach.

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The here presented 3D in vitro muscle tissues raised close to a glass coverslip 529 displayed myotube structures characteristic for progressing maturation e.g. sacomere 530 striations and multinucleation ( Figure 1) and indeed showed functional responses to 531 contractile stimuli (Figure 2, Supplementary Video 3). While the force measurement via 532 post deflection is a commonly used method [29], we now offer an advanced and reliable 533 readout for global force generation which is even more precise due to sharply milled post 534 edges and higher quality of imaging through glass. We therefore pave the way for future 535 contraction studies of diseased or individual patient related tissues. As for instance, by 536 shortening the height of the PMMA posts we can vary the spring constant of the posts 537 dramatically and therefore enable first isometric contraction investigations of skeletal 538 muscle tissues in vitro which is impossible using highly flexible PDMS.

539
To enable the study of potential correlations of contractile forces and tension 540 between cells, we exploit the post deformations to determine the overall tissue 541 contractility while the deformable beads are used to measure the local tension between 542 individual cells and tubes. PAA beads were just previously reported to function as 543 tension sensors within cancer spheroids in vitro, during phagocytosis as well as within 544 zebrafish embryos in vivo [13,28,53,55]. Combining local and global analysis is only 545 possible with the new chamber design, as the bead deformation analysis relies on high 546 resolution images. Furthermore, in contrast to the previous published tension analysis 547 programs, the here introduced approach focuses on the principal components of tissue 548 tension and directionality, which largely simplifies the analysis and the comparison 549 between different treatments, cell types and contraction trigger approaches. The access 550 to the global and the local tension also allows to independently crosscheck the values. we use t = f t /A to predict a local average tension of t = 1.8 ± 0.67 kPa, which is in 555 excellent agreement to the average value of 2.4 ± 0.9 kPa obtained by the elastic bead 556 analysis (Figure 4, see Methods and Material for details). This crosscheck suggests that 557 the post deflection measurements are already sufficient to determine average tension in 558 the tissue, but the local measurement can be used to then test the tension as function of 559 position in the tissue. 560

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Establishing the combined approach for the first time in developing 3D skeletal 561 muscle tissues in vitro, we observe a significant increase in local tension between 562 myotubes in C2C12 cells after one week of differentiation which does not further 563 increase in the following week ( Figure 4). Interestingly, we do not observe an increase of 564 global tissue pre-tension in the second week of differentiation, either ( Figure 2). 565 However, we indeed monitor a significant increase in myotube diameter from week one 566 to week two of differentiation as others reported before [1,2,31]. In addition, Afshar et 567 al. showed slight in vitro muscle tissue remodelling from week one to week two of 568 differentiation. Hence, neither myotube diameter nor progressing tissue remodelling has 569 a significant impact on local cellular tension within the 3D muscle tissues. Therefore, a 570 mechanical homeostasis may be reached on a cellular level among myoblast fusion, 571 myotube death and myotube progressing maturation after one week of in vitro muscle 572 tissue differentiation. Furthermore, we can speculate that myotubes may be protected 573 by a mechanical buffer layer composed of e.g. large glycocalyx biopolymers that are 574 known to have crucial mechanical functions for cells and cell aggregates [17]. Our 575 results suggest that cellular tension within in vitro muscle tissues is more connected to 576 tissue pre-tension.

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The novel culture mold is milled using PMMA and hence, easy and reliable to 578 fabricate. The chamber design also allows scaling to a large number of experiments in 579 parallel, which is a key element for screening approaches. It was designed for sufficient 580 gas exchange, easy medium exchange as well as drug delivery via holes in the lid of the 581 mold. Utilizing PMMA and microscopy glass for our molds, we overcome the issues of 582 immense chemical absorption by the material and poor optical properties that previous 583 approaches possessed [1,2,31]. Thus, we can additionally offer serum-free culturing of in 584 vitro skeletal muscle tissues and accurate drug dose response studies.

585
In conclusion, we provide a novel technique for functional 3D in vitro skeletal muscle 586 tissue cultivation that enables real-time high resolution microscopy of living 3D 587 biomimetic muscle tissue for the first time. We use the new approach for first global as 588 well as local cellular force investigations of developing in vitro muscle tissues. We 589 observe that cellular tension in C2C12 in vitro muscle tissues is closely related to global 590 tissue pre-tension and reaches a mechanical homeostatic phase after one week of 591 differentiation. However, in vitro skeletal muscle tissue maturation is still progressing in 592 the second week of differentiation, interestingly. Further, cultivation of muscle tissues 593 directly on glass enables whole new opportunities to study highly complex and dynamic 594 issues of myogenesis in 3D in vitro. In addition, an easy and reliable readout of 595 contractile forces makes the novel culture mold applicable for individualised drug 596 screening as well as diagnosis.