The myotendinous junction (MTJ) is the interface connecting skeletal muscle and tendon tissues. This specialized region represents the bridge that facilitates the transmission of contractile forces from muscle to tendon, and ultimately the skeletal system for the creation of movement. MTJs are, therefore, subject to high stress concentrations, rendering them susceptible to severe, life-altering injuries. Despite the scarcity of knowledge obtained from MTJ formation during embryogenesis, several attempts have been made to engineer this complex interfacial tissue. These attempts, however, fail to achieve the level of maturity and mechanical complexity required for in vivo transplantation. This review summarizes the strategies taken to engineer the MTJ, with an emphasis on how transitioning from static to mechanically inducive dynamic cultures may assist in achieving myotendinous maturity.

Musculoskeletal conditions continue to plague our community with approximately 1.71 × 109 people affected worldwide, making them the leading contributor to disability globally.1 Although only being responsible for 2% of total hospital discharges, musculoskeletal conditions reported 5.4% of the total hospital costs for children and adolescents aged 20 years or younger, correlating to a $7.6 billion economic burden. In fact, in 2011, there was an estimated $213 billion annual cost of direct treatment and lost wages in the United States alone.2 Arising from various diseases, injuries, and myopathies, musculoskeletal conditions are largely incurable or irreparable. Trauma to muscles and tendon alone contributes 20.6% of all work-related musculoskeletal injuries in Australia (Safe Work Australia, 2016), in which the MTJ represents the primary site of injury at 58.7%.3 Current surgical methods, such as suturing, allografts, and xenografts, result in residual scar tissue that disrupts the fragile biomechanics of the muscle–tendon interface, and hence do not achieve the therapeutic rehabilitation required for patients to return to pre-injury activities. Although the interdisciplinary field of tissue engineering has previously encompassed interfacial tissues, this domain remains relatively unexplored, particularly regarding the MTJ. Groups that have targeted the MTJ have observed interface-specific markers indicating MTJ development; however, current efforts have failed to recapitulate the maturity and mechanical complexity of native MTJs. Herein, we review the finite studies targeted at MTJ regeneration, revealing a common trend that may be extrapolated to other interfacial tissues. In particular, this review focuses on the mechanosensitive traits of skeletal muscle and tendon constructs in three-dimensional (3D) culture, in which the prospect of maturing MTJ constructs under uniaxial strain has been explored.

To rationally design a MTJ engineering strategy, we should consider the process by which it is originally formed. Herein, we begin by describing the development and mechanical characteristics of MTJ constituents in utero, followed by the biomechanics of fully developed MTJs in humans. Subsequently, through summarizing the finite studies targeted at MTJ engineering, the lack of dynamic culture methods has been identified as a limitation. Finally, this gap has been linked back to the established literature on individual tissue responses to uniaxial strain in dynamic cultures, revealing a targeted future perspective aimed at recapitulating the native environment more appropriately to be identified.

MTJ formation is unequivocally reliant on the interactions between developing tissues during embryogenesis. Originating from the myotome, myogenesis relies on a complex, yet not fully elucidated signaling network comprised of molecules, such as Wingless-related integration site proteins (Wnts), sonic hedgehog (Shh), and bone morphogenetic proteins (BMPs).4 These molecules, especially Wnt1 and Wnt3a, heavily influence the expression of myogenic regulatory factors (MRFs) responsible for myogenic lineage progression and differentiation.4,5 Contrarily, tenogenesis originates from the undifferentiated mesenchymal cells within the lateral plate mesoderm.6 Initially, Shh signals derived from the notochord induce paired box transcription factors (Pax) 1 and 9, which regulate mesenchymal differentiation. This process ultimately commits SRY-box transcription factor 9 (Sox-9)-expressing chondrogenic mesenchymal cells to their tendinous fate.7 Post mesenchymal differentiation, tendon progenitors express various transcription factors, including Scleraxis (Scx), Mohawk (Mkx), early growth response 1 (Egr1), and early growth response 2 (Egr2).8 

MTJ formation begins as tenocytes attach to skeletal muscle cells (myocytes). The chemical and mechanical signaling between these tissues during embryonic development is poorly understood; however, electron microscopy has been widely incorporated to analyze the origins of the MTJ. The earliest morphological modification observed at the MTJ is the formation of close associations between myogenic cells and tendon fibroblasts.9 Here, a dual identity can be seen as fibroblasts transdifferentiate by switching on a myogenic program, allowing fusion into myofibers.10 Simultaneously, extracellular material accumulates at the surface of muscle cells, representing the first appearance of the basement membrane.9,11 Myofibril production subsequently increases, allowing subsarcolemmal densities to appear at the intracellular surface of the MTJ.9 Finally, associations between myofibril thin filaments and subsarcolemmal densities occur, resulting in membrane folding, later depicted as invaginations.11 

To date, there has been limited reported data regarding the mechanical properties of embryonic skeletal muscle, particularly in mammals. Although myotendinous development is a generic process in vertebrates, Fig. 1 displays not only the variability between species but also the rapid formation in utero as evident in the increased stiffness and ultimate tensile strength (UTS). A study by McBride et al. focused on individual tendons adjacent to the femur and tibiotarsus of fertilized white leghorn eggs. At post-fertilization day (PFD) 14, a Young's modulus (E), UTS, and strain at failure (SAF) were determined to be 0.216 ± 0.060 MPa, 2.052 ± 1.112 MPa, and 12.77% ± 1.91%, respectively. At PFD 17, these values had increased to 1.02 ± 0.27 MPa, 21.411 ± 2.634 MPa, and 29.83 ± 5.33, indicating not only the rapid pace of embryonic tendon development but also the disparity between embryonic and adult tendon mechanical properties.12 This disparity may be seen by Nakagaki et al.'s elastic modulus of caged, 8-month-old calcaneal tendon of chickens (210.51 ± 46.01 MPa).13 It has been determined from Marturano et al. that the elastic modulus of the embryonic tendon increased nonlinearly as a function of embryonic stage at both the nano- and microscale.14 A further in-depth review of embryonic development may be found here.15 

FIG. 1.

Timeline of embryonic milestones and mechanical properties from conception to birth. (a) Ref. 12, (b) Ref. 16, (c) Ref. 17, (d) Ref. 18, (e) Ref. 19, (f) Ref. 20, (g) Ref. 14, (h) Refs. 21 and 22, (i) Ref. 23, (j) Ref. 24, and (k) Ref. 25. Created with BioRender.com.

FIG. 1.

Timeline of embryonic milestones and mechanical properties from conception to birth. (a) Ref. 12, (b) Ref. 16, (c) Ref. 17, (d) Ref. 18, (e) Ref. 19, (f) Ref. 20, (g) Ref. 14, (h) Refs. 21 and 22, (i) Ref. 23, (j) Ref. 24, and (k) Ref. 25. Created with BioRender.com.

Close modal

For the MTJ to develop into an integrated mechanical unit that can support and transmit great forces, many connective complexes must be considered.26 These connections must not only allow high tensile stresses to be endured but also act as the junction in which intracellular myofilament contractions transmit force to extracellular proteins found in the tendon.27 In total, five milestones of MTJ formation have been reported in the literature, including26,28–31

  • Actin microfilaments extend from the last Z-line of sarcomeres and terminate in the ridge-like protrusions.

  • Actin-binding proteins cross bind highly aligned actin filaments together.

  • Actin filament bundles insert into the sarcolemma plasma membrane via an electron-dense subsarcolemmal layer. The anchorage of these actin filaments may be perpendicular or oblique.

  • Transmembrane proteins that link cytoskeletal elements to basal lamina components. Basil lamina is seemingly thicker in these regions.

  • Proteins that link the basement membrane to collagen fibril-rich matrix in such a way that the collagen fibrils are parallel to the muscle myofilaments.

Two major and distinct transmembrane linkage systems have been described at the MTJ, which rely on the dystrophin-associated glycoproteins complex (DGC) and the binding and signaling protein “α7β1” integrin. Both systems constitute a structural link between cytoplasmic actin and tendinous extracellular matrix proteins. These links heavily rely on laminin 211, which is prevalent in both complexes, and represents the unique isoform that is found in adult human MTJs.26,27

The adult MTJ contains a complex morphology to combat stress concentrations and avoid injury. Advancements in scanning techniques, such as transmission and scanning electron microscopy (TEM and SEM, respectively), and focused ion beam (FIB) have steered scientists away from the idea of a simplistic planar divisional interface between skeletal muscles and tendons, revealing a more complex, interwoven design. Here, tendinous extracellular matrix (ECM) folds and protrudes into invaginations of the muscle cell membrane (sarcolemma) as represented in Fig. 2. In fact, while studying the human MTJ in 3D, Knudsen et al. more accurately described the connection between tissues as ridge-like protrusions30 that serve two primary mechanical purposes.

FIG. 2.

Ultrastructure of the MTJ. Created with BioRender.com.

FIG. 2.

Ultrastructure of the MTJ. Created with BioRender.com.

Close modal

First, the interwoven nature of the MTJ greatly increases the contact area between skeletal muscles and tendons, therefore significantly reducing stress concentrations.32 2D TEM images taken by Noonan et al. revealed a 10–20-fold increase in surface area compared to a smooth planar transition;33 however, recent 3D modeling by Knudsen et al. indicates an even greater increase.30 Second, the intertwining design positions the muscle cell membrane at acute angles relative to the applied force, compelling the sarcolemma to be predominately exposed to shear forces.34 Shear forces are optimal as cell membranes display greater strength of adhesion while subject to contractile shear stresses as opposed to tension loading.35 Thus, the ability of the MTJ membranes to transmit force is accentuated through invaginations which result in greater load-bearing capabilities. The ultrastructure of the MTJ may be seen below in Fig. 2.

Generally, collagen fibrils within the tendon are crimped until the onset of strain, resulting in an initial toe region of a stress–strain curve up until 2%–2.5% strain,36 in which tendonous strains primarily reside under physiological conditions.37 Once stretched beyond 4%, plastic deformation occurs, representing mild injury. If stretched over 10%, however, complete rupture will occur.38,39 The tendon is also described as a viscoelastic tissue, meaning compliance at low strain rates and resilience at high strain rates. This was elegantly displayed by Wren et al., who increased the rate of strain from 1 to 10 mm/s during tensile loading of the human Achilles tendon (AT).40 Here, a greater strain rate resulted in increased E, UTS, and SAF from 816 ± 218 to 822 ± 211 MPa, 71 ± 17 to 86 ± 24 MPa, and 7.5% ± 1.1% to 9.9% ± 1.9%, respectively.

To date, insufficient data have been acquired regarding the in vivo biomechanics of the MTJ due to the difficulty in attaining results during physical activity. Therefore, postmortem destructive tests are the primary source for mechanical properties of the MTJ. Several studies27,41 claim that the mechanical properties of the MTJ are that as displayed in Table I.

TABLE I.

Reported mechanical properties of the myotendinous junction and its constituents.

E (MPa) UTS (MPa) SAF (%)
Native MTJ  0.2789 ± 0.1509  0.1478 ± 0.016 31  122.4 ± 19.18 
Muscle  0.005–2.8  ⋯  ⋯ 
Tendon  500–1850  52 –120  5–16 
E (MPa) UTS (MPa) SAF (%)
Native MTJ  0.2789 ± 0.1509  0.1478 ± 0.016 31  122.4 ± 19.18 
Muscle  0.005–2.8  ⋯  ⋯ 
Tendon  500–1850  52 –120  5–16 

Upon further investigation, many of these references are obtained from studies conducted on not only different tissues but also different species. In fact, analysis of the original data referenced in these reviews provides greater insight into the variability of myotendinous tissue (see Table II). Here, * represents the corresponding value presumed to be identified in summary Table I.

TABLE II.

Analysis of references cited by several reviews elucidating the mechanical properties of adult human MTJ. * Indicates a possible origin source for the above-mentioned summary table (Table I).

Species Tissue E (units defined below) UTS (MPa) SAF (%) Reference
Human  Various wrist tendon  438.1 ± 93.7–726.1 ± 73.5 MPa  51.6 ± 9.3*–74.0 ± 13.5  11.4 ± 1.0–16.6 ± 1.7*  37  
Human  AT  816 ± 218–822 ± 211 MPa  71 ± 17–86 ± 24  7.5 ± 1.1–9.9 ± 1.9  40  
Rabbit  Patellar  549 ± 13–1390 ± 53 MPa  57.1 ± 2.5  5.3 ± 0.2*–14.1 ± 0.6  42  
Rabbit  Patellar  955 ± 97–1855 ± 77* MPa  ⋯  ⋯  43  
Human  Patellar  504 ± 222*–660 ± 266 MPa  53.6 ± 10.0–64.7 ± 15.0  14 ± 6–15 ± 5  44  
Rabbit  Achilles  281 ± 104.6–618 ± 87.0 MPa  23.9 ± 3.8–67.3 ± 4.2  15.7 ± 2.9*−16.3 ± 2.7*  45  
Pig diaphragm  MTJ  0.2789 ± 0.1509* MPa  0.1478 ± 0.016 31*  122.4 ± 19.18*  41  
Human  PT  643.1 ± 53.0 MPa  68.5 ± 6.0  13.5 ± 0.7  46  
Canine  PT  457 ± 98.0 MPa  122 ± 25.6  32.3 ± 7.8  47  
MDX  Muscle  18 ± 6 kPa  ⋯  ⋯  48  
C57  Muscle  12 ± 4 kPa  ⋯  ⋯  48  
Murine  Muscle  11.5 ± 1.3–45.3 ± 4.0 kPa  ⋯  ⋯  49  
Rabbit  TA muscle  1.75 ± 1.18–2.79 ± 0.67 kPa  ⋯  ⋯  50  
Human  Gracilis tendon  612.8 ± 40.6 MPa  111.5 ± 4.0  33.9 ± 1.5  51  
Human  Patellar  305.5 ± 59.0 MPa  58.3 ± 6.1  35.1 ± 4.4  51  
Human  Patellar  307 ± 17 MPa  43.7 ± 3.9  23.4 ± 1.4  52  
Species Tissue E (units defined below) UTS (MPa) SAF (%) Reference
Human  Various wrist tendon  438.1 ± 93.7–726.1 ± 73.5 MPa  51.6 ± 9.3*–74.0 ± 13.5  11.4 ± 1.0–16.6 ± 1.7*  37  
Human  AT  816 ± 218–822 ± 211 MPa  71 ± 17–86 ± 24  7.5 ± 1.1–9.9 ± 1.9  40  
Rabbit  Patellar  549 ± 13–1390 ± 53 MPa  57.1 ± 2.5  5.3 ± 0.2*–14.1 ± 0.6  42  
Rabbit  Patellar  955 ± 97–1855 ± 77* MPa  ⋯  ⋯  43  
Human  Patellar  504 ± 222*–660 ± 266 MPa  53.6 ± 10.0–64.7 ± 15.0  14 ± 6–15 ± 5  44  
Rabbit  Achilles  281 ± 104.6–618 ± 87.0 MPa  23.9 ± 3.8–67.3 ± 4.2  15.7 ± 2.9*−16.3 ± 2.7*  45  
Pig diaphragm  MTJ  0.2789 ± 0.1509* MPa  0.1478 ± 0.016 31*  122.4 ± 19.18*  41  
Human  PT  643.1 ± 53.0 MPa  68.5 ± 6.0  13.5 ± 0.7  46  
Canine  PT  457 ± 98.0 MPa  122 ± 25.6  32.3 ± 7.8  47  
MDX  Muscle  18 ± 6 kPa  ⋯  ⋯  48  
C57  Muscle  12 ± 4 kPa  ⋯  ⋯  48  
Murine  Muscle  11.5 ± 1.3–45.3 ± 4.0 kPa  ⋯  ⋯  49  
Rabbit  TA muscle  1.75 ± 1.18–2.79 ± 0.67 kPa  ⋯  ⋯  50  
Human  Gracilis tendon  612.8 ± 40.6 MPa  111.5 ± 4.0  33.9 ± 1.5  51  
Human  Patellar  305.5 ± 59.0 MPa  58.3 ± 6.1  35.1 ± 4.4  51  
Human  Patellar  307 ± 17 MPa  43.7 ± 3.9  23.4 ± 1.4  52  

Comparing the original data to the referenced values (Table I), although being within similar range, this table appears to be an over-simplification of MTJ properties as it does not reveal the extent to which biomechanics vary upon location or species. Considering postmortem mechanical properties of the human MTJ constituents, Loren and Lieber studied tendons of the human wrist, observing E (at maximal tetanic tension), UTS, and SAF of 438.1 ± 93.7 to 726.1 ± 73.5 MPa, 51.6 ± 9.3 to 74.0 ± 13.5 MPa, and 11.4% ± 1.0% to 16.6% ± 1.7%, respectively.37 Comparatively, Johnson et al. studied the tensile and viscoelastic properties of the human patellar tendon, revealing an E, UTS, and SAF of 504 ± 222–660 ± 266 MPa, 53.6 ± 10–64.7 ± 15 MPa, and 14% ± 6%–15% ± 5%, respectively.44 

Hanson et al. determined E, UTS, and SAF of the human AT to be 222.8 ± 84.6–316.8 ± 110 MPa, 21.9 ± 9.9–28.1 ± 9.8 MPa, and 13.8% ± 4.4%–16.3% ± 3.5%, respectively. Comparatively, the Iliopsoas tendon gave an E, UTS, and SAF of 63.5 ± 23.6–165.3 ± 67.3 MPa, 6.8 ± 2.1–22.5 ± 7.3 MPa, and 18.3% ± 3.5%–19.7% ± 5.2%, respectively.53 Mabe et al. found E, UTS, and SAF to be 201 ± 70 MPa, 16.0 ± 7.32 MPa, and 0.15% ± 0.07%, compared to that of the quadriceps tendon 153 ± 46 MPa, 19.1 ± 5.42 MPa, and 0.16% ± 0.02%, respectively.54 

As expected, there is a greater abundance of studies conducted on animals as opposed to humans; however, the study by Pollock et al. revealed that the elastic properties of tendons do not vary significantly from animals of different body mass.55 These findings may explain how different tendons in different species consistently have properties within or close to the range found within humans.31,42,43,45,55–66 Similarly, the data on the tensile properties of the skeletal muscles in humans is scarce compared to that of other species.48–50,67

Focusing on living human mechanical properties, Maganaris et al. stimulated the human tibialis anterior (TA) muscle using conductive aluminum pads, enabling ultrasonography and magnetic resonance imaging of the connecting AT.68 By stimulating isometric loads, a maximum stress, strain, and stiffness of 25 MPa, 2.5%, and 1.2 GPa were found. Conversely, Kongsgaard et al. used a calf raising apparatus to study the AT revealing a max stress, strain, and modulus of 29 ± 3MPa, 4.2% ± 1.1%, and 2.0 ± 0.4 GPa, respectively.69 Furthermore, the Achillies MTJ averaged a proximal displacement of 7.1 ± 0.9 mm. A summary of living human mechanical properties of MTJ constituents under load can be seen below in Table III.

TABLE III.

Mechanical properties of living, in vivo adult human myotendinous tissues.

Tissue Stimuli Analysis E (Units defined below) Max stress (MPa) Max strain (%) Reference
Free Achilles tendon  Isometric plantarflexion ramp contractions  Ultrasonography  2.0 ± 0.4–1.9 ± 0.5 GPa  29 ± 3–31 ± 4  4.2 ± 1.1–4.5 ± 1.4  69  
Ultrasonography, electromyography  788 ± 181 MPa  36.5 ± 4.6  8.0 ± 1.2  70  
Cyclic isometric contractions  MRI      2.8–4.7  71  
Achilles tendon  Isometric plantar flexion  Ultrasound      5.1 ± 1.1  72  
Rest  SUE and MyotonPRO  363.38 ± 54.11 kPa      73  
Surae aponeurosis and Achilles tendon  Isometric plantarflexion ramp contractions  Ultrasonography  1048–1474 MPa  41.6  4.4–5.6  74  
Gastrocnemius aponeurosis (distal)  Isometric plantarflexion ramp contractions  Ultrasonography, electromyography      1.4 ± 0.4  70  
Gastrocnemius tendon  Isometric plantarflexion contraction  Ultrasonography  1.16 ± 0.15 GPa  32.4 ± 2.3  4.9 ± 1  75  
Gastrocnemius tendon  Isometric plantarflexion contraction  Ultrasonography  1.16 ± 0.15 GPa  32.4 ± 2.3  4.9 ± 1  75  
Tibialis anterior tendon  Electrical stimulation  Ultrasonography, MRI  1·2 ± 0·15 GPa  25 ± 2·5  2·5 ± 0·4  68  
Patellar tendon  Isometric knee extension  Ultrasonography  1.5 ± 0.4–1.8 ± 0.6 GPa  34 ± 8–53 ± 13  5.3 ± 0.7–5.8 ± 1.0  76  
Gastrocnemius (medial)    SUE and MyotonPRO  22.59 ± 3.31 kPa    73  
Gastrocnemius (lateral)  23.56 ± 4.08 kPa  73  
Tibialis anterior  Rest and knee flexion  Ultrasound shear wave imaging  40.60 ± 1.00–258.10 ± 15.00 kPa    77  
Gastrocnemius (medial)  16.50 ± 1.00–225.40 ± 41.00 kPa  77  
Soleus  14.50 ± 2.00 to 55.00 ± 5.00 kPa  77  
Tissue Stimuli Analysis E (Units defined below) Max stress (MPa) Max strain (%) Reference
Free Achilles tendon  Isometric plantarflexion ramp contractions  Ultrasonography  2.0 ± 0.4–1.9 ± 0.5 GPa  29 ± 3–31 ± 4  4.2 ± 1.1–4.5 ± 1.4  69  
Ultrasonography, electromyography  788 ± 181 MPa  36.5 ± 4.6  8.0 ± 1.2  70  
Cyclic isometric contractions  MRI      2.8–4.7  71  
Achilles tendon  Isometric plantar flexion  Ultrasound      5.1 ± 1.1  72  
Rest  SUE and MyotonPRO  363.38 ± 54.11 kPa      73  
Surae aponeurosis and Achilles tendon  Isometric plantarflexion ramp contractions  Ultrasonography  1048–1474 MPa  41.6  4.4–5.6  74  
Gastrocnemius aponeurosis (distal)  Isometric plantarflexion ramp contractions  Ultrasonography, electromyography      1.4 ± 0.4  70  
Gastrocnemius tendon  Isometric plantarflexion contraction  Ultrasonography  1.16 ± 0.15 GPa  32.4 ± 2.3  4.9 ± 1  75  
Gastrocnemius tendon  Isometric plantarflexion contraction  Ultrasonography  1.16 ± 0.15 GPa  32.4 ± 2.3  4.9 ± 1  75  
Tibialis anterior tendon  Electrical stimulation  Ultrasonography, MRI  1·2 ± 0·15 GPa  25 ± 2·5  2·5 ± 0·4  68  
Patellar tendon  Isometric knee extension  Ultrasonography  1.5 ± 0.4–1.8 ± 0.6 GPa  34 ± 8–53 ± 13  5.3 ± 0.7–5.8 ± 1.0  76  
Gastrocnemius (medial)    SUE and MyotonPRO  22.59 ± 3.31 kPa    73  
Gastrocnemius (lateral)  23.56 ± 4.08 kPa  73  
Tibialis anterior  Rest and knee flexion  Ultrasound shear wave imaging  40.60 ± 1.00–258.10 ± 15.00 kPa    77  
Gastrocnemius (medial)  16.50 ± 1.00–225.40 ± 41.00 kPa  77  
Soleus  14.50 ± 2.00 to 55.00 ± 5.00 kPa  77  

Apart from stimuli type and analysis method, simulation protocol, subject age, and sex, the definition of free tendon plays an important role in comparable outcomes. For example, previous methods have estimated Achilles tendon mechanical properties from measurements of distal medial gastrocnemius MTJ displacement in relation to an external marker.69 This method, however, does not account for elongation or contraction of the tendon–aponeurosis structure and, therefore, the results vary. It is concluded that heterogeneous analysis methods and tissue types produce inconsistent outcomes while considering the biomechanics of myotendinous tissues.

Despite the adaptable and complex inner workings of the MTJ, injuries are still prevalent.86 Although most tears occur near the MTJ, as opposed to through it (indirect tears),78 the MRI reveals the resultant scar tissue is sufficient to disrupt interfacial biomechanics.79–81 Tears within close proximity to the MTJ generally occur between the cell membrane and lamina densa of the basement membrane.35 

These injuries mainly arise from a passively overstressed tissue, or fast eccentric contractions, where the skeletal muscle contracts while lengthening.82 In fact, eccentric forces greater than 20% of the average maximum isometric force are sufficient to induce rupture.29 

In skeletal muscles, it is widely accepted that either disrupted sarcomeres, or alternatively, damage to the excitation–contraction coupling system represents the immediate signs of injury.83 Furthermore, sarcomeres may be disturbed as a lack of homogeneity results in uneven energy absorption.84 Most commonly, muscles that are exposed to these contractions include, but are not limited to, the hamstring, rectus femoris, hip adductor, and calf muscles.82 Although a universal grading system is yet to be agreed upon, a classification system compromising four grades are identified by location (distal, middle, and proximal), severity, and symptoms. These grades are as follows:85–90 

  • Grade 0: edema or fluid adjacent to an intact tendon/aponeurosis/epimysium without myofibril detachment.

  • Grade 1: myofibril detachment without tendon/aponeurosis/epimysium change.

  • Grade 2: myofibril detachment with adjacent tendon/aponeurosis/epimysium increased signal, delamination, or defect but no retraction.

  • Grade 3: myofibril detachment with adjacent tendon retraction.

Grade 1 and 2 tears may be treated with non-steroidal anti-inflammatory drugs (NSAIDS), protection, rest, ice, compression, and elevation (PRICE) protocol,91 and physiotherapy where a full recovery is expected, whereas grade 3 tears require surgical intervention.92 Currently, suturing of allografts and xenografts are the only available clinical option. A review of current suturing approaches for the MTJ may be found here.93 Although this generally allows for the re-attachment of the torn muscle–tendon unit, significant scar tissue results in compromised biomechanics and increased likelihood of reinjury.93 In cases where the interface is completely severed, suture may not allow the possibility to re-join the MTJ. Since most myotendinous injuries are untreated, surgical outcomes are not well documented.92,94

Furthermore, if surgery is delayed, muscular atrophy may make it impossible to re-attach. Although experimental pharmacologic agents are still being explored,95 suture-based approaches remain the predominant surgical intervention for repair of a grade 3 MTJ injury. Suture-based approaches remain the predominant surgical intervention for repair of a grade 3 MTJ injury.

Tissue engineering (TE) approaches provide great hope to assist the natural repair of musculoskeletal diseases. Recently, a trend has been developing around the hypothesis that tissue-engineered scaffolds that mimic embryonic developmental forces may dictate the differentiation of embedded cells down a specific lineage more effectively. These approaches, however, remain undeveloped and insufficient due to the limited understanding of tissue interactions during embryogenesis. Extracellular forces that contribute to embryonic development may, however, be applied to tissue-engineered scaffolds in hopes of recreating developmental-specific microenvironments. For example, Kardon et al. revealed that in the avian hind limb, the initial morphogenetic events, formation of tendon primordia, and initial differentiation of myogenic precursors occur autonomously with respect to one another.96 This effect was further iterated by Kieny et al. who studied the development of an embryonic chick wing. Here, they concluded that tendons start to develop autonomously from the muscle bulks, but for their maintenance and further development they require connexion to a muscle belly.97 Although being imperative to interfacial tissue development, in utero relationships remain heavily unexplored, therefore, tissue engineering advancements are predominantly governed by the understanding of individual tissues as opposed to the relationship between the two. Thus, the pillars of tissue engineering (cells, growth factors, and scaffolds) must be adapted to house complementary environments for both skeletal muscles and tendons prior to in vivo transplantation. Few attempts have been made to engineer the MTJ, some of which utilize scaffolds to provide initial mechanical support and disparity between skeletal muscles and tendons. Not only do scaffolds provide mechanical support but they also pre-define tissue geometry and replicate the ECM, thus allowing for provisional cellular adherence, migration,98 proliferation, and differentiation.99 To achieve desired outcomes, the following five parameters should be considered and manipulated in such a way that provokes intended biological responses:100 

  1. Biodegradability

  2. Biocompatibility

  3. Engineered scaffold architecture

  4. Surface properties

  5. Mechanical properties

Further increasing the complexity, these factors are all time-dependent, making the ideal scaffold challenging to fabricate. Scaffolds for individual tissues will not be reviewed here, as they have been elsewhere.101 Instead, heterogeneous structures able to mimic dissimilar tissue properties will be analyzed.

Although planar homogeneous scaffolds are most often reported in the literature, they are incapable of producing spatially altered mechanical properties.102 Despite this, architectural adaptations, such as the transition from linear to crimped fibers, may be introduced to make tissue-specific scaffolds. For example, Hochleitner et al. used melt electro-writing below the critical translation speed to obtain sinusoidal polymer fibers.103 Under tensile loading, these fibers, although homogeneous, exhibited a toe region resembling the uncrimping of collagen fibrils in tendinous tissue. Producing a similar structure, Wu et al. fabricated a tendon-bioinspired wavy scaffold via electrospinning, displaying a tensile toe region and UTS consistent with that of the native tendon. Expectedly, however, the tensile modulus within this toe region fell below that of the native tendon, indicating optimization is required.104 In general, electrospun fibers maintain high surface area–volume ratios that mimic the natural ECM nanostructure. One disadvantage of electrospun fibers, however, are their ability to maintain sufficient mechanical properties, particularly for tendons, and, therefore, will only be discussed in hybrid scaffold situations. For example, Sahoo et al. electrospun poly (lactic-co-glycolic acid) (PLGA) nanofibers onto the surface of a knitted PLGA scaffold,105 resulting in increased proliferation, cellular function, failure load, elastic stiffness, and toe stiffness, whereas cell attachment remained comparable.

Heterogeneous scaffolds are a promising approach for mimicking the mechanical impedance seen at interfacial tissues. Melt electro-writing (MEW)-controlled deposition style presents a unique approach to achieving these mechanical and architectural variances. As practically all transplantable tissues require a 3D geometry, a cylindrical collector is often used for target tissues, such as heart valves, tendon, ligament, and skeletal muscle. For example, Saidy et al. printed polycaprolactone (PCL) onto a 22 mm diameter cylinder for heart valve tissue-engineered scaffolds,106 creating discrete and controlled architectural changes as diamond-shaped pores turned rectangular [Fig. 3(b)]. Although not having undergone appropriate tensile testing, compression testing revealed that, as expected, variations in tubular morphology had a direct impact on mechanical properties.

FIG. 3.

Heterogeneous scaffolds in tissue engineering (a) Fabrication of different patterns by changes in collector speed at and below CTS reproduced with permission from Hochleitner et al., Bionanomaterials 17(3), 159–171 (2016). Copyright 2016 Authors, licensed under a creative commons attribution (CC BY) license.109 (b) Spatially heterogeneous tubular scaffolds for in situ heart valve tissue engineering using melt electro-writing, reproduced with permissions from Saidy et al., Adv. Funct. Mater. 32(21), 2110716 (2022). Copyright 2022 John Wiley and Sons License.106 (c) Dimension-based multiphasic scaffold, reproduced with permission from Hrynevich et al., Small 14(22), e1800232 (2018). Copyright 2018 John Wiley and Sons.110 (d) Quadrant intersection between varying fiber diameter and pore size, reproduced with permission from Xie et al., “Mater. Des. 181, 108092 (2019). Copyright 2019 Authors, licensed under a creative commons attribution (CC BY) license.107 (e) HUVECs in the scaffold with pore sizes of 100 and 200 μm on Day7, reproduced from reproduced with permission from Xie et al., “Mater. Des. 181, 108092 (2019). Copyright 2019 Authors, licensed under a creative commons attribution (CC BY) license.107 (f) Cells' reaction to various grid sizes, reproduced with permission from Nguyen et al., Mater. Sci. Eng., C 103, 109785 (2019). Copyright 2019 Elsevier.108 

FIG. 3.

Heterogeneous scaffolds in tissue engineering (a) Fabrication of different patterns by changes in collector speed at and below CTS reproduced with permission from Hochleitner et al., Bionanomaterials 17(3), 159–171 (2016). Copyright 2016 Authors, licensed under a creative commons attribution (CC BY) license.109 (b) Spatially heterogeneous tubular scaffolds for in situ heart valve tissue engineering using melt electro-writing, reproduced with permissions from Saidy et al., Adv. Funct. Mater. 32(21), 2110716 (2022). Copyright 2022 John Wiley and Sons License.106 (c) Dimension-based multiphasic scaffold, reproduced with permission from Hrynevich et al., Small 14(22), e1800232 (2018). Copyright 2018 John Wiley and Sons.110 (d) Quadrant intersection between varying fiber diameter and pore size, reproduced with permission from Xie et al., “Mater. Des. 181, 108092 (2019). Copyright 2019 Authors, licensed under a creative commons attribution (CC BY) license.107 (e) HUVECs in the scaffold with pore sizes of 100 and 200 μm on Day7, reproduced from reproduced with permission from Xie et al., “Mater. Des. 181, 108092 (2019). Copyright 2019 Authors, licensed under a creative commons attribution (CC BY) license.107 (f) Cells' reaction to various grid sizes, reproduced with permission from Nguyen et al., Mater. Sci. Eng., C 103, 109785 (2019). Copyright 2019 Elsevier.108 

Close modal

Steering away from biomechanics, changes in pore sizes have been shown to dictate cell attachment, alignment, and growth rate. For example, Xie et al's Melt electro-wrote high-resolution PCL scaffolds with four distinct zones segmented by pore size.107 Here, the proliferation rate of bone marrow-derived stem cells (BMSCs) and human umbilical vein endothelial cells (HUVECs) in small pores was determined to be three times faster than in larger pores, with changes in fiber diameter dictating the spread of cells, attachment, and alignment [Fig. 3(e)]. Similarly, Nguyen et al. seeded NIH-3T3 cells on MEW PCL scaffolds [Fig. 3(f)], confirming that cell growth is considerably influenced by lattice structure grid size, enabling the control of spatially attached cell populations.108 

To date, advances in tissue engineering the MTJ rely extensively on the knowledge gained from individual tissue types. Engineering strategies can be segregated into scaffold-based and scaffold-free approaches, depending on their intended function. Scaffold-based approaches provide greater promise in mimicking the translational stiffness between skeletal muscles and tendons. Here, individual scaffolds are engineered for skeletal muscles and tendons, where an overlap defines the interface (Fig. 4).

FIG. 4.

Common structural design of engineered myotendinous junction. Here, cells and scaffolds are spatially deposited, in which the 10% interfacial region signifies the main area for analysis of specific MTJ markers. Created with BioRender.com.

FIG. 4.

Common structural design of engineered myotendinous junction. Here, cells and scaffolds are spatially deposited, in which the 10% interfacial region signifies the main area for analysis of specific MTJ markers. Created with BioRender.com.

Close modal

For example, Ladd et al. used a 10% overlap of co-electrospun PCL and poly(L-lactide) (PLLA) fibers, creating a mechanical disparity between muscle and tendon, respectively.41 This unique fabrication technique produced a continuous structure, which subsequently underwent tensile loading. As targeted, the PCL side (representing SM) produced the lowest stiffness and UTS of 4.490 ± 1.604 and 1.069 ± 0.2713 MPa, whereas PLLA (representing tendon) had the highest moduli and UTS of 27.62 ± 6.063 and 3.741 ± 0.8486 MPa, respectively. The central region (representing the MTJ) gave an E and UTS of 20.06 ± 7.773 and 2.384 ± 0.5987 MPa. Conclusively, the whole composite structure produced an E and UTS of 7.339 ± 2.131 and 0.5058 ± 0.2130 MPa, respectively, and could last up to 100 cycles during cyclic loading. Subsequent embedding of C2C12 and NIH/3T3 cells into evenly distributed bovine collagen I provided no reduction in cell viability and accommodated attachment, survival, and differentiation of myoblasts into myotubes. These traits, however, did not progress into evidence of cell re-organization at the interface.

Similarly, Merceron et al. also used spatial deposition of synthetic polymers with a 10% overlap to create an interfacial region.111 Here, bioprinting was utilized to deposit polyurethane (PU) with C2C12 cells (representing muscle), and PCL with NIH/3T3 cells (representing tendon), in which the deposition of cells had a negligible effect on viability. Mechanical characterization of the composite structure displayed a yield strain of 300%, which was expectedly driven by the superior elasticity of PU. Elastic modulus of PU, PCL, and the interface was 0.39 ± 0.005, 46.67 ± 2.67, and 1.03 ± 0.14 MPa, respectively, which, compared to Ladd et al., were mechanically inferior. On the muscle side, C2C12 cells expressed both desmin and myosin heavy chain (MHC), aligned along a unilateral plane, and began to show multinucleation, indicating their differentiation into myotubes. On the tendon side, cells began to secrete collagen I, producing a distinct interfacial region (Fig. 5). Focusing on the junction, upregulation of MTJ-associated genes, such as pax, talin 1 (tln1), vinculin, integrin β1, laminin α1, and laminin α2, were prevalent.

FIG. 5.

Current attempts at engineering the myotendinous junction (a) Co-electrospun triphasic scaffold compromising PCL and PLLA synthetic polymers, reproduced with permissions from Ladd et al., Biomaterials 32(6), 1549–1559 (2011). Copyright 2011 Elsevier.41 (b) 3D bio-printed scaffold with PCL and PU, reproduced with permission from Merceron et al., Biofabrication 7(3), 035003 (2015). Copyright 2015 IOP Publishing Ltd., through Clearance Center.111 (c) A bio-printed complex tissue model for myotendinous junction with biochemical and biophysical cues, reproduced with permission from W. J. Kim and G. H. Kim, Bioeng. Transl. Med. 7(3), e10321 (2022). Copyright 2022 Authors, licensed under Creative Commons Attribution (CC BY) license.113 (d) A standard dumbbell-shaped muscle model, reproduced with permissions from Laternser et al., SLAS Technol. 23(6), 599–613 (2018). Copyright 2018 Elsevier.112 (e) Macroscopic and schematic images of procedures used in MTJ complete rupture model preparation and sheet-pellet transplantation, reproduced with permission from Hashimoto et al., PeerJ 4, e2231 (2016). Copyright 2016 Authors, licensed under a Creative Commons Attribution (CC BY) license.116 (f) In vivo culture of a 3D bio-printed MTJ construct, reproduced with permission from Merceron et al., Biofabrication 7(3), 035003 (2015). Copyright 2015 IOP Publishing Ltd,, through Clearance Center.111 

FIG. 5.

Current attempts at engineering the myotendinous junction (a) Co-electrospun triphasic scaffold compromising PCL and PLLA synthetic polymers, reproduced with permissions from Ladd et al., Biomaterials 32(6), 1549–1559 (2011). Copyright 2011 Elsevier.41 (b) 3D bio-printed scaffold with PCL and PU, reproduced with permission from Merceron et al., Biofabrication 7(3), 035003 (2015). Copyright 2015 IOP Publishing Ltd., through Clearance Center.111 (c) A bio-printed complex tissue model for myotendinous junction with biochemical and biophysical cues, reproduced with permission from W. J. Kim and G. H. Kim, Bioeng. Transl. Med. 7(3), e10321 (2022). Copyright 2022 Authors, licensed under Creative Commons Attribution (CC BY) license.113 (d) A standard dumbbell-shaped muscle model, reproduced with permissions from Laternser et al., SLAS Technol. 23(6), 599–613 (2018). Copyright 2018 Elsevier.112 (e) Macroscopic and schematic images of procedures used in MTJ complete rupture model preparation and sheet-pellet transplantation, reproduced with permission from Hashimoto et al., PeerJ 4, e2231 (2016). Copyright 2016 Authors, licensed under a Creative Commons Attribution (CC BY) license.116 (f) In vivo culture of a 3D bio-printed MTJ construct, reproduced with permission from Merceron et al., Biofabrication 7(3), 035003 (2015). Copyright 2015 IOP Publishing Ltd,, through Clearance Center.111 

Close modal

Yet another example of bioprinting's efficacy in creating spatially deposited co-cultures was described by Laternser et al.112 Here, a dumbbell-shaped construct was deposited in between two posts, where primary human skeletal muscle-derived cells were printed between five successive layers of gelatin methacrylate (GelMA), and tenocytes between a GelMA/polyethylene glycol dimethacrylate (PEGDMA)-based ink. Interestingly, tenocytes were printed around posts, whereas skeletal muscle-derived cells were printed in between, with a gap of 3 mm to create a clear border, maintaining a >95% cell viability. Myoblasts formed aligned areas and were able to attach to the tendinous tissue; however, after 1–2 days, the structure tore at one of the bioink boarders. The successful attachment of myoblasts to tenocytes indicated the formation of MTJ-like structures, possibly resultant from the developed tension between posts.

Another form of mechanical stimulation for engineering the MTJ was investigated by Kim et al., who used a combination of biochemical and biophysical cues.113 One culture analyzed cell-laden human adipose-derived stem cells (hASC) in bioinks including collagen, muscle-derived ECM (mECM), and tendon-derived ECM (tECM)), in which flow-induced alignment of cells in bioinks was achieved by controlling (i) the collagen concentration in the core and (ii) the flow rates in the core and sheath. Using the idealized flow rates for optimized cellular alignment, 14 days post culture, paxillin, integrin, fibronectin, and talin were found in the MTJ region, suggesting MTJ formation. Furthermore, upregulated expression of integrin β1- and MTJ-associated genes [Pax, Tln1, thrombospondin 1 (Thbs1), collagen type 1 (Col1a1), laminin alpha 1 (Lama1), myosin heavy chain 2 (Myh2), and Scx] was identified, leading to the conclusion that MTJ formation can be clearly affected by the physical interfacing shape between muscle and tendon cells.

Scaffold-free approaches hold promise in circumventing issues associated with scaffold fabrication, degradation, and reduced biologically active volume. Larkin et al. evaluated the co-culture of skeletal muscle satellite cells and fibroblasts obtained from pregnant Fischer 344 rats based on their ability to create MTJs.114 One week of monolayer culture enabled the muscle and tendon cells to roll up into 3-D constructs around tendon tails, in which a highly aligned interface of collagen fibrils and myotubes formed. Furthermore, although less concentrated than found in adult samples, both paxillin and talin were localized at the interface. Proceeding to tensile testing, an average E of 37.2 ± 10.3 kPa (for muscle construct) was revealed, being comparable to one quarter the passive stiffness reported for young adult rat soleus muscle. Conclusively, Larkin et al. successfully engineered constructs resembling neonatal MTJs using a cell approach, greatly expanding the potential to control the phenotype of skeletal muscles in culture.85 

Continuing from this, Kostrominova et al. adopted the same method to engineer self-organizing tendon (SOT) constructs.115 Briefly, once myocytes fused to form spontaneously contracting multinucleated myotubes, SOT constructs were pinned onto the muscle cell monolayer, allowing the construct to roll up around the anchors. At the early stages of culture, MTJ-like structures were represented by sub-sarcolemma densities, which later developed into well-defined folding of the plasma membrane. These plasma membranes were surrounded by type I collagen with the characteristic striation pattern.

Several other studies regarding engineered MTJ constructs have been summarized in Table IV and may be visualized in Fig. 5 below.

TABLE IV.

Current biomaterial and biomaterial-free approaches to engineering the myotendinous junction in three-dimensions.

Scaffold Materials Fabrication Cells Location Findings Reference
Tri-phasic  Collagen/PCL/PLLA  Electrospinning  C2C12/NIH3T3  Vitro  • E: 4.49 (1.604), 27.62 (6.063), and 20.06 (7.773) MPa for muscle, tendon, and junction, respectively.
• No indication of MTJ formation. 
41  
Tri-phasic  Hydrogel/PU/PCL  Bio-printing  C2C12/NIH3T3  Vitro  • E = 0.39 (0.05), 46.67 (2.67), and 1.03 (0.14) MPa for muscle, tendon, and junction, respectively.
•   Upregulation MTJ-associated genes paxillin, talin, vinculin, integrin β1, laminin α1, and laminin α2. 
111  
Mono-phasic  Polystyrene  Bio-printing  C2C12/ C3H10T1/2  Vitro  • Functionalizing oriented sub-micron fibers with printed GFs provides instructive cues to spatially control cell fate and alignment to mimic native tissue organization.  117  
5 layers  GelMA/PEGDMA  Bio-printing  Rat tenocytes/human myoblasts  Vitro  • Attachment of myoblasts to tenocytes.  112  
Tri-phasic  Porcine m-dECMs + fibronectin/porcine t-dECMs/ collagen I (throughout)  Bio-printing  hASCs  Vitro  • E = 40 ± 5 kPa.
• Upregulated gene expression integrin β1 and MTJ-associated genes.
• Paxillin, integrin, fibronectin, and talin expression.
• Attachment of aligned actin filaments and collagen fibers. 
113  
Triphasic  Female dog dECM derived from gastrocnemius muscle and AT  ⋯  ⋯  Vivo  • Contractile function reached 48.4% compared to that of uninjured MTJ.
• Responsive, vascularized, and innervated MTJ within 6 months of implantation. 
118  
⋯  SOT constructs  ⋯  Fibroblasts/satellite cells  Vitro  • 14–17 days of cocultured resulted in an interface of collagen fibrils and myotubes orientated along the longitudinal axis.
• Interface resembles neonatal MTJ.
• Paxillin and talin concentration at junction.
• E = 37.2 ± 10.3 kPa.
• Failure occurred on muscle side of interface. 
114  
⋯  SOT constructs  ⋯  Fibroblasts/satellite cells  Vitro  • Tapered myofibers integrated into collagen-rich matrix at the interface.
• Interface resembles neonatal MTJ.
• Some interfaces revealed folding of the plasma membrane with surrounding extracellular type 1 collagen and characteristic striation patterns. 
115  
Bi/triphasic  Collagen, agarose, and hydroxyapatite  ⋯  Dermal fibroblasts and Sket.4u cells  Vitro  • Increased expression of tenomodulin and αSMA.
• Aligned tendon and muscle cells.
• Stiffness gradient achieved from muscle hydrogel (20 kPa) to tendon (140 kPa). 
119  
Biphasic  Type 1 collagen  ⋯  Skeletal muscle cells  Vitro  • Muscle fiber ends showed invaginations of the sarcolemma.
• Actin filaments from the last sarcomere terminated at the end of muscle fibers.
• Basal lamina observed.
• Highly contractile multinucleated muscle fibers. 
120  
Monophasic  Skeletal muscle-derived mesenchymal stem cells (SM-d-SCs) sheet pellets  ⋯  ⋯  Vivo  • Favorable results in the reconstruction and/or reconnection of ruptured muscles and tendons.
• Engrafted green fluorescent protein (GFP)+ fibroblast-like cells migrated to bridge skeletal muscles and tendons.
• 36% recovery in the short-term group, and 49% in the long-term group. 
116  
⋯  Type 1 collagen and tECM  ⋯  C2C12  Vitro  • Loading upregulated Myh1, Myh2, and Myh4 in tECM.
• Pax and collagen XXII expression increased in tECM and accentuated by mechanical stimulation. 
121  
tECM and mECM  ⋯  C2C12 and tendon fibroblasts  Vitro  • 50%–70% paxillin increase in ECM hydrogels compared to type 1 collagen hydrogel.
• ECM-conditioned media does not promote paxillin.
• Collagen XXII present in tissue-derived ECM, but to a lesser extent. 
122  
Scaffold Materials Fabrication Cells Location Findings Reference
Tri-phasic  Collagen/PCL/PLLA  Electrospinning  C2C12/NIH3T3  Vitro  • E: 4.49 (1.604), 27.62 (6.063), and 20.06 (7.773) MPa for muscle, tendon, and junction, respectively.
• No indication of MTJ formation. 
41  
Tri-phasic  Hydrogel/PU/PCL  Bio-printing  C2C12/NIH3T3  Vitro  • E = 0.39 (0.05), 46.67 (2.67), and 1.03 (0.14) MPa for muscle, tendon, and junction, respectively.
•   Upregulation MTJ-associated genes paxillin, talin, vinculin, integrin β1, laminin α1, and laminin α2. 
111  
Mono-phasic  Polystyrene  Bio-printing  C2C12/ C3H10T1/2  Vitro  • Functionalizing oriented sub-micron fibers with printed GFs provides instructive cues to spatially control cell fate and alignment to mimic native tissue organization.  117  
5 layers  GelMA/PEGDMA  Bio-printing  Rat tenocytes/human myoblasts  Vitro  • Attachment of myoblasts to tenocytes.  112  
Tri-phasic  Porcine m-dECMs + fibronectin/porcine t-dECMs/ collagen I (throughout)  Bio-printing  hASCs  Vitro  • E = 40 ± 5 kPa.
• Upregulated gene expression integrin β1 and MTJ-associated genes.
• Paxillin, integrin, fibronectin, and talin expression.
• Attachment of aligned actin filaments and collagen fibers. 
113  
Triphasic  Female dog dECM derived from gastrocnemius muscle and AT  ⋯  ⋯  Vivo  • Contractile function reached 48.4% compared to that of uninjured MTJ.
• Responsive, vascularized, and innervated MTJ within 6 months of implantation. 
118  
⋯  SOT constructs  ⋯  Fibroblasts/satellite cells  Vitro  • 14–17 days of cocultured resulted in an interface of collagen fibrils and myotubes orientated along the longitudinal axis.
• Interface resembles neonatal MTJ.
• Paxillin and talin concentration at junction.
• E = 37.2 ± 10.3 kPa.
• Failure occurred on muscle side of interface. 
114  
⋯  SOT constructs  ⋯  Fibroblasts/satellite cells  Vitro  • Tapered myofibers integrated into collagen-rich matrix at the interface.
• Interface resembles neonatal MTJ.
• Some interfaces revealed folding of the plasma membrane with surrounding extracellular type 1 collagen and characteristic striation patterns. 
115  
Bi/triphasic  Collagen, agarose, and hydroxyapatite  ⋯  Dermal fibroblasts and Sket.4u cells  Vitro  • Increased expression of tenomodulin and αSMA.
• Aligned tendon and muscle cells.
• Stiffness gradient achieved from muscle hydrogel (20 kPa) to tendon (140 kPa). 
119  
Biphasic  Type 1 collagen  ⋯  Skeletal muscle cells  Vitro  • Muscle fiber ends showed invaginations of the sarcolemma.
• Actin filaments from the last sarcomere terminated at the end of muscle fibers.
• Basal lamina observed.
• Highly contractile multinucleated muscle fibers. 
120  
Monophasic  Skeletal muscle-derived mesenchymal stem cells (SM-d-SCs) sheet pellets  ⋯  ⋯  Vivo  • Favorable results in the reconstruction and/or reconnection of ruptured muscles and tendons.
• Engrafted green fluorescent protein (GFP)+ fibroblast-like cells migrated to bridge skeletal muscles and tendons.
• 36% recovery in the short-term group, and 49% in the long-term group. 
116  
⋯  Type 1 collagen and tECM  ⋯  C2C12  Vitro  • Loading upregulated Myh1, Myh2, and Myh4 in tECM.
• Pax and collagen XXII expression increased in tECM and accentuated by mechanical stimulation. 
121  
tECM and mECM  ⋯  C2C12 and tendon fibroblasts  Vitro  • 50%–70% paxillin increase in ECM hydrogels compared to type 1 collagen hydrogel.
• ECM-conditioned media does not promote paxillin.
• Collagen XXII present in tissue-derived ECM, but to a lesser extent. 
122  

Currently, the task of engineering interfacial tissues, especially that of the MTJ, reveals the lack of understanding in several other domains. For example, knowledge about how tissue-specific cells interact with the microenvironment and ECM of opposing tissues is scarce. Gaffney et al. decellularized the porcine AT and gastrocnemius muscle to form tissue-specific prehydrogel digests. Seeding of C2C12 myoblasts and tendon fibroblasts in their opposing ECM hydrogel revealed a 50%–70% increase in paxillin expression, and collagen XXII to a lesser extent.122 Most recently, Gaffney et al's continuation of this research represents the only mechanically stimulated study focused on engineering a myotendinous construct.121 Here, C2C12 cells were seeded in either type 1 collagen or tendon-derived extracellular matrix (tECM) and conditioned in a customed designed bioreactor. Cyclic tensile exposure of 10% strain, for 10 800 cycles at 1 Hz (3 h per day) resulted in the upregulation of myosin heavy chains (Myh1, Myh2, and Myh4), Pax, and type XXII collagen. Interestingly, tECM consistently produced greater upregulation of these factors compared to that of type 1 collagen, possibly due to the presence of tendon-specific ECM factors. Furthermore, transmission of uniaxial loads from scaffolds to cells is heavily influenced by hydrogel stiffness, as is actin–myosin striation. In fact, hydrogel mechanical properties have become increasingly influential in dynamic cultures and must, therefore, be considered an imperative aspect of interfacial tissue engineering. As previously stated, different hydrogels are commonly used in MTJ constructs; however, evolving to dynamic cultures involving uniaxial strain introduces more variables, such as tensile modulus, flexibility, and fatigue. Engler et al. found an optimal gel modulus of 12 kPa for skeletal muscle constructs,48 within the range of native skeletal muscle. Decellularized MTJ (D-MTJ) has only been reported once in the literature by Zhao et al., who seeded muscle satellite cells into D-MTJ derived from porcine AT MTJ. This study failed to investigate the coculture of tenocyte and myoblasts; however, it successfully characterized the mechanical properties of decellularized MTJ.62 Considering human cells, Tsuchiya et al. cocultured semitendinosus and gracilis tendon–muscle tissues, revealing three main findings: (i) human-derived tenocytes do not enhance proliferation of myoblasts, (ii) human-derived tenocytes enhance myotube formation, and (iii) human-derived myoblasts and tenocytes release factors important for myotube formation.123 

In essence, the foundation of myotendinous engineering is fragile due to the limited knowledge of cellular interactions in their opposing ECMs. Therefore, it is difficult to identify the exact mechanisms responsible for the expression of interfacial characteristics, especially since this is not fully understood during in vivo development. The majority of data indicating the dependency of skeletal muscles and tendons during development has arisen from NULL studies, which strategically eliminate inputs, such as muscle contraction, to investigate the effect on tendon development.

As discussed, all techniques used to engineer the MTJ have employed heterogeneous structures in combination with cocultured skeletal muscle and tendon cells. Here, the junction is represented in a slight overlap of hydrogels, cells, and scaffolds (where relevant). Although the exact mechanisms behind MTJ development are for the most part unknown, tissue engineers commonly analyze MTJ-specific gene and protein expression to determine the success of their study, such as collagen XXII. Furthermore, the existence of invaginations directly correlates with mechanical competency and is, therefore, commonly seen as an indicator of MTJ maturation. Although there has been some investigation into the response of tendon and SM cells in their opposing dECMs, this information is scarce, making it increasingly difficult to predict in vitro outcomes and create a targeted, standardized approach. The trend of further maturing MTJ constructs via mechanical stimuli has been identified as future steps by many authors;41,111,124 however, few have investigated the effect of mechanotransduction on engineered MTJ constructs.121 The transition to dynamic cultures must, however, be related back to the vast body of knowledge gained from individual tissues, such as skeletal muscles and tendons, respectively. As the progression from 2D to 3D cultures often produces conflicting results, it is important to only refer to 3D cultures for implantable tissue constructs.

To facilitate the transition to dynamic cultures, mechanically inductive bioreactors are required. Hermetically sealed, contaminant-free bioreactors provide an environment that can mimic complex and dynamic in vivo environments.125 Bioreactors can serve three main purposes, including (i) systematic investigations to be undertaken in a controlled and reproducible manner; (ii) pre-conditioning and maturing engineered constructs prior to in vivo implantation; and (iii) the expansion of cells for in vivo transplantation.125–127 Pre-conditioning is of great importance as loss of cellular function is currently one of the main limitations in tissue engineering.128 Thus, to replicate in vivo interactions, bioreactors aim to control physico-chemical stimuli parameters to elicit specific cellular responses. Several bioreactor categories have been developed since the early 1980s, including rocking bed, batch stirred tank, rotating wall vessels, perfusion, and isolated expansion automated systems.126,127,129 Considering the myotendinous junction, stretch bioreactors are favorable; however, only few are commercially available. General bioreactor considerations revolve around the following five domains:130 

  1. Physical design/material selection

  2. Mass transfer (nutrient)

  3. Mechanical stimulation

  4. Electrical stimulation

  5. Feedback control system

Here, only two imperative domains will be described, including mass transfer and mechanical stimulation, as they have the most relevance to MTJ engineering.

In vivo, cells are generally within 100 μm from capillary networks that provide gas and nutrient supply and waste disposal.131 Bioreactors must, therefore, emulate the rich and extensive vascular network of the human body via the circulation of nutrient-rich culture media. Three forms of mass transfer may occur to accommodate this in a bioreactor, including (i) convection, (ii) diffusion, and (iii) perfusion. Mass transfer in static cultures solely rely on diffusion to achieve adequate mass transfer, which, due to the diffusive penetration threshold of ∼100–200 μm, commonly results in heterogeneous cell distributions.132 Lack of adequate perfusion is commonly visible in tissue constructs by the development of a central necrotic core surrounded by a dense layer of viable cells.126 As practically all transplantable tissues will exceed these dimensions and lack vasculature, mass transfer drastically restricts scalability.125,133 Conversely, dynamic cultures further incorporate convection via continuous media flow, and perfusion, via mass flow throughout the scaffold itself. Dynamic environments are most commonly facilitated by pumps and rotary motion; however, they may also be generated with heat exchangers, humidifiers, bubble traps, and oxygenators.

Suitable mass transfer, thus, allows culture media constituents, such as nutrients, glucose, proteins, vitamins, oxygen, and pH, to be homogeneously delivered to engineered tissue. The concentration of these constituents, however, varies with tissue growth and type and must continuously be replenished. This may be achieved via peristaltic pump systems that induce the constant supply and circulation of fresh medium. In essence, the addition of fluid flow within bioreactors drastically increases the complexity of analysis and design considerations. For example, pumping flow rate may be set to induce laminar or turbulent flow, which is further influenced by a bioreactor's architecture.134 Ideally, mass transfer from supplying arteries and arterioles of the perimysium at the myotendinous junction must be emulated to avoid necrosis and enable reliable tissue growth.

Regarding the maturation of myotendinous structures, the implementation of mechanical forces in vitro appears to increase tissue maturity. To date, there is convincing evidence that physical forces are imperative for tissue development during embryogenesis.126 In fact, a study by Sungsoo et al. displayed that stress-induced signal transduction is at least 40 times faster than growth factor-induced signal transduction while considering the human airway smooth muscle (HASM) cells.135 Forces assumed to contribute to this consist of hydrodynamic125 and hydrostatic pressure, fluid dynamics, mechanical stresses and strains, and electrical cues.133 To replicate one aspect of these, a dimension-based mechanical stimulation is often embedded into bioreactors.136 

In three dimensional tissue culture, forces may be applied via three mechanisms, including (i) fluid flow, (ii) mechanical vibration, and (iii) mechanical stimulation.137,138 Focusing on mechanical stimulation, forces may be applied via uniaxial or biaxial tension, compression, and/or torsion depending on the native tissue environment that is being emulated. Tensile strain is the most commonly applied biophysical stimulus used to mimic the native muscle environment. While each individual muscle belly shortens during movement, the contraction of individual myofibers against two ends that are tethered to the rigid bone induces tensile strains in the tissues as they are contracting.139 In reported stretch bioreactors, one end is often clamped into a stationary position, whereas the other is clamped to an oscillation source, commonly connected to a stepper motor.140 Here, strain concentrations increase in magnitude approaching the clamps, leading to altered cellular responses which may be able to mimic interfacial tissues, such as the MTJ.

Studies demonstrate that, dependent on the exposure routine, mechanical input can (i) stimulate ECM production,141 (ii) improve cell/tissue organization,136 (iii) direct cell differentiation142 and alignment,143 and (iv) enhance targeted tissue functions.144 For example, by analyzing cells more susceptible to stretching, such as fibroblasts, cellular alignment along the axis of strain can be observed after just three hours of exposure.145 This is expected as fixation of a scaffold between two anchor points creates predictable lines of isometric strain which cells sense and respond to accordingly. This mechanically mediated internal tension is sufficient in promoting cellular alignment along the principal axis of strain,140,146,147 and as previously described, was used by both Laternser et al.112 and Gaffney and Laternser121 to engineer the MTJ.

Considering uniaxial strain exposure, little is known about which specific mechanical force(s) or regimes of application (magnitude, frequency, and duty cycle) induce specific cellular responses.131 Brown et al. showed that tissue engineering programs should be customized for axial vs limb tendons while studying Scx-GFP mice cells under a cyclic uniaxial strain.148 Significant consideration must, therefore, be given to determine the appropriate application of biophysical stimuli, such as static vs cyclic, continuous vs intermittent, low vs high amplitude, among others. For example, a frequency of 1 Hz is similar to the natural stride frequency, mimicking the strain cycle of muscles in locomotion.139 Referring back to embryonic cues however, Kodama et al. found that muscle-driven movements began at E14 in mice, creating a timepoint benchmark for the transition from static culture to mechanically inducive dynamic cultures.24 These findings have further implications to the cues dictating tendon development, suggesting tendon progenitor cells (TPCs) begin to differentiate as a function of their particular anatomical microenvironments prior to E14. This was further consolidated by Lipp et al., who found muscular contraction to begin at E13.5.149 

Adhering to this embryonic phenomenon, Foolen et al. applied a cyclic stretch after a static culture to study the subtle balance between contact guidance and stress concentrations in 3D.147 Here, cells in the core displayed alignment along the principal axis of the strain; however, cells became increasingly oblique toward the exterior due to decreasing contact guidance toward the exterior of the construct. Similar results can be seen from Boerboom et al.150 and Rubbens et al.151 Conversely, Chen et al. restrained cell deformation between two posts for 24 h before applying a cyclic uniaxial strain, resulting in homogeneously aligned cells from surface to core.146 It has been suggested that the earliest forces experienced by MTJ constituents is via the elongation of the connecting bone,16 in which the increasing cell-mediated tension between bioreactor posts may have mimicked this function and enhanced homogeneous cellular alignment.146 

Cyclic loading, however, produces inconsistent results on signaling and myogenic marker expression, making it difficult to determine optimal cyclic parameters, whereas static loading largely produces consistent results. Although the exact science behind mechanotransduction is yet to be elucidated, it has been hypothesized that the strain effect on myogenesis may occur by signaling intracellular transcription factors. This further iterates that the expression of myogenic markers is a direct indicator of the progression of immature constructs to functional implants. There have been many studies conducted on the effect of mechanical strain on skeletal muscles and tendons individually; however, there have not yet been many studies specifically investigating the effects of mechanical strain at the MTJ (see Tables V and VI).

TABLE V.

3D tendon constructs under uniaxial strain.

Cell Type Type Substrate/scaffold Mechanical parameters Effects Reference
Tendon-derived stem cells (TDSCs)  Cyclic  Confluent monolayer with ECM deposition (rolled up)  6% strain, 0.25 Hz, 8 h/d, 6 days  • Uniaxial > biaxial for tenogenic differentiation markers.
• Compact, aligned, and connected F-actin.
• Increased expression of Scleraxis, Mohawk, TNMD, and COL1A1.
• Osteogenic markers also relevant. 
152  
hMSCs  Cyclic  Electrospun polycaprolactone fibers  Intermittent 12% strain, 0.01in/s, for 21 days  • Increased collagen and ECM production
• Fiber arrangement has no discernible pattern.
• Addition of cells increased elastic moduli from 30.33 ± 8.37 to 216.53 ± 55.71 MPa 
141  
BM-dMSCs  Cyclic  Decellularized tendon scaffolds  0%, 3%, or 5% strain at 0.33 Hz for up to 1 h daily for 11 days  • 3% strain gave best results
• UTS of 17.7 ± 3.8 MPa
• Elastic moduli of 119 ± 44 MPa, within 25% of native tendon value (98 ± 25 MPa)
• Altered ECM composition.
• Deep cellular integration
• Tendon-like gene expression 
153  
BM-dMSCs  Cyclic  Raw silk knits  12h/day, 0.1 Hz  • Collagen synthesis increased.
• Improved cellular elongation and ECM deposition along the direction of the mechanical strain.
• Upregulated gene expression (collagen 1, tenascin-C, and TNMD).
• Minimal differences in tensile properties 
154  
Fibroblasts  Cyclic  Elastomeric polyurethane (Tecoflex SG-80A)  10% strain, 0.25 Hz, 8 h/day  • Increased proliferation
• Presence of ascorbic acid results in increased elastic modulus in the high strain region
• Increased fibroblast DNA content
• Ascorbic acid increases collagen 1 and fibronectin matrix accumulation.
• Homogeneous distribution of ECM and cells. 
155  
TDSCs  Cyclic  Poly(L-lactide-co-ε-caprolactone)/collagen (P(LLA-CL)/Col) scaffold  2%, 4%, and 8% strain at 0.3, 0.5, and 1.0 Hz  • No effect on cell viability
• 0.5 Hz and 4% strain gave produced greatest increase in proliferation
• Increased collagen 1, tenascin-C expression, TNMD, and SCX 
156  
MSCs  Cyclic  Collagen fiber scaffold  10% strain, 1 Hz, 3 h on/3 h off, 14 days  • Increased SCX
• Increased ECM deposition (collagen 1, collagen 3, and tenascin-C) 
157  
Fibroblasts  Cyclic  Collagen disk fiber  0.1 Hz, 0%–0.7% strain, 7 days  • Increased SCX and TNMD
• Fiber nuclei deformed to more or a spindle (elongated) morphology, aligned to the long axis of the fiber.
• Increased peak stress, Young's modulus, and toughness.
• No change in failure strain or toe-in strain 
158  
Tenocytes  Cyclic  Collagen sponge  3% preload followed by 10% strain at 0.5, 1, and 2 Hz.  • Scaffold moduli of 17.5 kPa, and stiffness of 0.022 N/mm
• Tenocytes and scaffold fibers promoted homogeneous alignment 
159  
hMSCs  Cyclic  PCL electrospun yarns  5% strain, 1 Hz, for 1 h per day  • More textured and rounded cells
• No evidence of cell infiltration into the structure
• Upregulation of Col1a1, Col1a2. Col3a1, tenascin-C, elastin, and fibronectin.
• Increased Young's modulus and UTS. 
160  
BMSC's  Cyclic  Decellularized tendon slices  3% strain, 0,2 Hz, 12 h/day, 7 days  • Upregulated decorin and TNMD
• No changes in UTS or stiffness
• Enhanced cell infiltration deeper into the tissue
• Improved ECM protein expression
• Alignment along principal axis of strain 
161  
MSCs  Cyclic  Decellularized tendon matrix scaffolds  2% strain at 1 Hz (various regimes)  • Strain decreased cellular alignment.
• Short stretching may promote cell integration.
• Downregulated collagen 1A2 expression
• Upregulated collagen 1A2 and 3A1, and SCX 
162  
MSCs  Cyclic  Electrospun PCL nanofiber yarns  4% strain at 0.5 Hz, 2 h/day for 12 days  • Promotion of TNMD and COL1
• Increased total collagen content.
• Increased SCX, TNC, COL1, COL3, TNMD, and VEGFA genes 
163  
MSCs  Cyclic  Collagen 1 sponges  Stimulated once every 5 min for 8 h/d to a peak strain of 2.4% for 2 weeks.  • Increased collagen 1 and 3
• No significant increase in decorin, fibronectin, and GAPDH gene expression
• Linear modulus increased from 0.0005 ± 0.002 to 0.02 ± 0.01 MPa
• Maximum stiffness increased from 0.003 ± 0.001 to 0.005 ± 0.002 MPa 
164  
Fibroblasts  Cyclic  Bioartificial tendons  1 h/day at 1% elongation and 1 Hz  • Increased nuclei elongation and cytoplasmic extensions
• Modulus increased from 0.49 to 1.8 MPa
• UTS strength increased 2.9 fold after 7 days 
165  
AT  Cyclic  Achilles tendon  0.25 Hz for 8 h/day, 0%–9% for 6 days  • 3% strain results in slightly disrupted ECM structure and significant cellular morphology changes.
• 6% strain circumvented the previously mentioned issues, and also had no type 3 collagen expression.
• 9% strain resulted in partial tears and 45% apoptosis. 
166  
MSCs  Static  GelMA hydrogel yarn  15% for 7 days, with an additional 15% from day 7–14  • Higher fraction of elongated cells and spreading.
• Cell alignment up to 50% in the direction of stretch. 
167  
MSCs  Cyclic  Human umbilical cord  2% strain, 0.5 h/day, 0.5 cycle/min, and 14 days  • Decreased cell infiltration.
• Some alignment in direction of stretch, but ECM still somewhat disorganized/wavy.
• Increased UTS.
• Day 7 revealed upregulated collagen type 1 and 3, and tenomodulin.
• Day 14 revealed downregulated collagen type 1 and 3 while elastin and SCX were upregulated. 
168  
TDSCs  Cyclic  P(LLA-CL)/Col  4% strain, 0.5 Hz, 2 h/day, and 14 days.  • Increase in cell number.
• More uniform distribution of cells.
• Col 1 and 3, decorin, tenascin C, and biglycan upregulated.
• Downregulated Runx2, Col 2, and aggrecan
• UTS increased from 43.18% ± 6.58% to 59.58% ± 7.81% of normal patella tendon in rabbits.
• Young's modulus increased from 23.30% ± 3.83% to 51.99% ± 7.16% of normal patella tendon. 
169  
BM-MSCs  Cyclic  Human umbilical cords  2% strain, 1 h/day, 1 cycle/min, calibrated 1% strain during rest periods (several other variations)  • Increase in cell number.
• UTS ranging from 1.06 (0.34) to 1.58 (0.35) MPa
• Increased fiber alignment, although quite far from ideal scenario.
• Upregulation of biglycan, tenascin C, Col 1, and Col 3.
• Slow frequencies produce significant decorin downregulation. 
170  
Fibroblasts  Cyclic  Human flexor and extensor tendon  0.625–2.5 N, up to 8 days  • Conditioning duration has significant effect on material properties, the load magnitude does not.
• Increased UTS and elastic modulus for all loading regimes. 
171  
MSCs  Cyclic  Braided hyaluronate elastic band  1 Hz at 0.1 N (≈10% strain)  • Increased collagen type 1 and 3, decorin, SCX, and tenascin-C.
• Biochemical and mechanical stimuli was used in tandem. 
172  
MSCs  Cyclic  Collagen type 1 sponge  0%–2.4% strain, 1 Hz for 20 s followed by 100 s rest, for 5 h/day  • Increased type 1 collagen.
• Increased linear stiffness. 
173  
MSCs  Cyclic  Poly(ethylene glycol)-based hydrogel  10% strain, 1 Hz, 3 h of strain followed by 3 h without  • No change in cell number or viability
• Increased Col1, Col2, and tenascin C. 
174  
MSCs  Cyclic  Decellularized tendon scaffold  3% strain, 0.33 Hz, 1 h/day  • Increased SCX, Col1, Col3
• Decreased TNMD
• Increased mechanical properties. 
175  
Tenocytes  Cyclic  Decellularized tendon  1.25N stretch, 1 cycle/min, in alternating 1 h periods of load and rest, over 5 days.  • UTS increased from 35.69 ± 5.62 to 71.17 ± 14.15 N.
• Elastic modulus increased from 632 ± 86 to 1091 ± 169 MPa (no significant difference from freshly harvested tissue) 
176  
MSCs  Cyclic  Collagen sponge  2.4%, at 1 Hz, 8 h/day, with either 100 or 3000 cycles/day.  • Stiffness decreases with increase cycles/day.
• Col 1, 3, and Decorin was maximal at 3000cycles/day, with a minimal value at 100cycles/day.
• Fibronectin expression was maximal at 0cycles/day, and minimal at 100cycles/day. 
177  
BMSCs  Cyclic  Collagen sponge  15% strain, 1 Hz, and 3 days  • Increased SCX and TNC
• Decreased Col1a1 expression.
• Failure strain increased from 98.58% ± 8.92% to 116.40% ± 14.51%.
• Young's modulus decreased from 14.17 ± 0.87 to 7.80 ± 1.72 kPa.
• Increased collagen scaffold porosity. 
178  
MSCs  Cyclic  Collagen type 1  1% strain, 1 Hz, 30 min/day, and 7 days  • 50% increase in collagen production.
• SCX initially decreased, but subsequently increased.
• Wnt5a and Wnt14 downregulated.
• Wnt16, Col1,3, XII, and elastin upregulated. 
179  
Fibroblasts  Cyclic  Collagen type 1  10% strain for 15 min, 0% strain for 15 min, and 15 min rest for 24 h  • Cells aligned parallel to force from 110° to 180° (not seen in static cultures)
• Upregulated collagen I, collagen III, TNMD, and TGF‐β 
143  
Cell Type Type Substrate/scaffold Mechanical parameters Effects Reference
Tendon-derived stem cells (TDSCs)  Cyclic  Confluent monolayer with ECM deposition (rolled up)  6% strain, 0.25 Hz, 8 h/d, 6 days  • Uniaxial > biaxial for tenogenic differentiation markers.
• Compact, aligned, and connected F-actin.
• Increased expression of Scleraxis, Mohawk, TNMD, and COL1A1.
• Osteogenic markers also relevant. 
152  
hMSCs  Cyclic  Electrospun polycaprolactone fibers  Intermittent 12% strain, 0.01in/s, for 21 days  • Increased collagen and ECM production
• Fiber arrangement has no discernible pattern.
• Addition of cells increased elastic moduli from 30.33 ± 8.37 to 216.53 ± 55.71 MPa 
141  
BM-dMSCs  Cyclic  Decellularized tendon scaffolds  0%, 3%, or 5% strain at 0.33 Hz for up to 1 h daily for 11 days  • 3% strain gave best results
• UTS of 17.7 ± 3.8 MPa
• Elastic moduli of 119 ± 44 MPa, within 25% of native tendon value (98 ± 25 MPa)
• Altered ECM composition.
• Deep cellular integration
• Tendon-like gene expression 
153  
BM-dMSCs  Cyclic  Raw silk knits  12h/day, 0.1 Hz  • Collagen synthesis increased.
• Improved cellular elongation and ECM deposition along the direction of the mechanical strain.
• Upregulated gene expression (collagen 1, tenascin-C, and TNMD).
• Minimal differences in tensile properties 
154  
Fibroblasts  Cyclic  Elastomeric polyurethane (Tecoflex SG-80A)  10% strain, 0.25 Hz, 8 h/day  • Increased proliferation
• Presence of ascorbic acid results in increased elastic modulus in the high strain region
• Increased fibroblast DNA content
• Ascorbic acid increases collagen 1 and fibronectin matrix accumulation.
• Homogeneous distribution of ECM and cells. 
155  
TDSCs  Cyclic  Poly(L-lactide-co-ε-caprolactone)/collagen (P(LLA-CL)/Col) scaffold  2%, 4%, and 8% strain at 0.3, 0.5, and 1.0 Hz  • No effect on cell viability
• 0.5 Hz and 4% strain gave produced greatest increase in proliferation
• Increased collagen 1, tenascin-C expression, TNMD, and SCX 
156  
MSCs  Cyclic  Collagen fiber scaffold  10% strain, 1 Hz, 3 h on/3 h off, 14 days  • Increased SCX
• Increased ECM deposition (collagen 1, collagen 3, and tenascin-C) 
157  
Fibroblasts  Cyclic  Collagen disk fiber  0.1 Hz, 0%–0.7% strain, 7 days  • Increased SCX and TNMD
• Fiber nuclei deformed to more or a spindle (elongated) morphology, aligned to the long axis of the fiber.
• Increased peak stress, Young's modulus, and toughness.
• No change in failure strain or toe-in strain 
158  
Tenocytes  Cyclic  Collagen sponge  3% preload followed by 10% strain at 0.5, 1, and 2 Hz.  • Scaffold moduli of 17.5 kPa, and stiffness of 0.022 N/mm
• Tenocytes and scaffold fibers promoted homogeneous alignment 
159  
hMSCs  Cyclic  PCL electrospun yarns  5% strain, 1 Hz, for 1 h per day  • More textured and rounded cells
• No evidence of cell infiltration into the structure
• Upregulation of Col1a1, Col1a2. Col3a1, tenascin-C, elastin, and fibronectin.
• Increased Young's modulus and UTS. 
160  
BMSC's  Cyclic  Decellularized tendon slices  3% strain, 0,2 Hz, 12 h/day, 7 days  • Upregulated decorin and TNMD
• No changes in UTS or stiffness
• Enhanced cell infiltration deeper into the tissue
• Improved ECM protein expression
• Alignment along principal axis of strain 
161  
MSCs  Cyclic  Decellularized tendon matrix scaffolds  2% strain at 1 Hz (various regimes)  • Strain decreased cellular alignment.
• Short stretching may promote cell integration.
• Downregulated collagen 1A2 expression
• Upregulated collagen 1A2 and 3A1, and SCX 
162  
MSCs  Cyclic  Electrospun PCL nanofiber yarns  4% strain at 0.5 Hz, 2 h/day for 12 days  • Promotion of TNMD and COL1
• Increased total collagen content.
• Increased SCX, TNC, COL1, COL3, TNMD, and VEGFA genes 
163  
MSCs  Cyclic  Collagen 1 sponges  Stimulated once every 5 min for 8 h/d to a peak strain of 2.4% for 2 weeks.  • Increased collagen 1 and 3
• No significant increase in decorin, fibronectin, and GAPDH gene expression
• Linear modulus increased from 0.0005 ± 0.002 to 0.02 ± 0.01 MPa
• Maximum stiffness increased from 0.003 ± 0.001 to 0.005 ± 0.002 MPa 
164  
Fibroblasts  Cyclic  Bioartificial tendons  1 h/day at 1% elongation and 1 Hz  • Increased nuclei elongation and cytoplasmic extensions
• Modulus increased from 0.49 to 1.8 MPa
• UTS strength increased 2.9 fold after 7 days 
165  
AT  Cyclic  Achilles tendon  0.25 Hz for 8 h/day, 0%–9% for 6 days  • 3% strain results in slightly disrupted ECM structure and significant cellular morphology changes.
• 6% strain circumvented the previously mentioned issues, and also had no type 3 collagen expression.
• 9% strain resulted in partial tears and 45% apoptosis. 
166  
MSCs  Static  GelMA hydrogel yarn  15% for 7 days, with an additional 15% from day 7–14  • Higher fraction of elongated cells and spreading.
• Cell alignment up to 50% in the direction of stretch. 
167  
MSCs  Cyclic  Human umbilical cord  2% strain, 0.5 h/day, 0.5 cycle/min, and 14 days  • Decreased cell infiltration.
• Some alignment in direction of stretch, but ECM still somewhat disorganized/wavy.
• Increased UTS.
• Day 7 revealed upregulated collagen type 1 and 3, and tenomodulin.
• Day 14 revealed downregulated collagen type 1 and 3 while elastin and SCX were upregulated. 
168  
TDSCs  Cyclic  P(LLA-CL)/Col  4% strain, 0.5 Hz, 2 h/day, and 14 days.  • Increase in cell number.
• More uniform distribution of cells.
• Col 1 and 3, decorin, tenascin C, and biglycan upregulated.
• Downregulated Runx2, Col 2, and aggrecan
• UTS increased from 43.18% ± 6.58% to 59.58% ± 7.81% of normal patella tendon in rabbits.
• Young's modulus increased from 23.30% ± 3.83% to 51.99% ± 7.16% of normal patella tendon. 
169  
BM-MSCs  Cyclic  Human umbilical cords  2% strain, 1 h/day, 1 cycle/min, calibrated 1% strain during rest periods (several other variations)  • Increase in cell number.
• UTS ranging from 1.06 (0.34) to 1.58 (0.35) MPa
• Increased fiber alignment, although quite far from ideal scenario.
• Upregulation of biglycan, tenascin C, Col 1, and Col 3.
• Slow frequencies produce significant decorin downregulation. 
170  
Fibroblasts  Cyclic  Human flexor and extensor tendon  0.625–2.5 N, up to 8 days  • Conditioning duration has significant effect on material properties, the load magnitude does not.
• Increased UTS and elastic modulus for all loading regimes. 
171  
MSCs  Cyclic  Braided hyaluronate elastic band  1 Hz at 0.1 N (≈10% strain)  • Increased collagen type 1 and 3, decorin, SCX, and tenascin-C.
• Biochemical and mechanical stimuli was used in tandem. 
172  
MSCs  Cyclic  Collagen type 1 sponge  0%–2.4% strain, 1 Hz for 20 s followed by 100 s rest, for 5 h/day  • Increased type 1 collagen.
• Increased linear stiffness. 
173  
MSCs  Cyclic  Poly(ethylene glycol)-based hydrogel  10% strain, 1 Hz, 3 h of strain followed by 3 h without  • No change in cell number or viability
• Increased Col1, Col2, and tenascin C. 
174  
MSCs  Cyclic  Decellularized tendon scaffold  3% strain, 0.33 Hz, 1 h/day  • Increased SCX, Col1, Col3
• Decreased TNMD
• Increased mechanical properties. 
175  
Tenocytes  Cyclic  Decellularized tendon  1.25N stretch, 1 cycle/min, in alternating 1 h periods of load and rest, over 5 days.  • UTS increased from 35.69 ± 5.62 to 71.17 ± 14.15 N.
• Elastic modulus increased from 632 ± 86 to 1091 ± 169 MPa (no significant difference from freshly harvested tissue) 
176  
MSCs  Cyclic  Collagen sponge  2.4%, at 1 Hz, 8 h/day, with either 100 or 3000 cycles/day.  • Stiffness decreases with increase cycles/day.
• Col 1, 3, and Decorin was maximal at 3000cycles/day, with a minimal value at 100cycles/day.
• Fibronectin expression was maximal at 0cycles/day, and minimal at 100cycles/day. 
177  
BMSCs  Cyclic  Collagen sponge  15% strain, 1 Hz, and 3 days  • Increased SCX and TNC
• Decreased Col1a1 expression.
• Failure strain increased from 98.58% ± 8.92% to 116.40% ± 14.51%.
• Young's modulus decreased from 14.17 ± 0.87 to 7.80 ± 1.72 kPa.
• Increased collagen scaffold porosity. 
178  
MSCs  Cyclic  Collagen type 1  1% strain, 1 Hz, 30 min/day, and 7 days  • 50% increase in collagen production.
• SCX initially decreased, but subsequently increased.
• Wnt5a and Wnt14 downregulated.
• Wnt16, Col1,3, XII, and elastin upregulated. 
179  
Fibroblasts  Cyclic  Collagen type 1  10% strain for 15 min, 0% strain for 15 min, and 15 min rest for 24 h  • Cells aligned parallel to force from 110° to 180° (not seen in static cultures)
• Upregulated collagen I, collagen III, TNMD, and TGF‐β 
143  
TABLE VI.

3D skeletal muscle constructs under uniaxial strain.

Cell Type Type Substrate/scaffold Mechanical parameters Major findings Reference
C2C12  Static  Fibrin  10% static strain for 6 h, followed by 18-h rest at 3% static strain. Repeated for 6 days.  • Higher degree of alignment (>85% in direction of strain, residual 15 showed a 15%–30% deviation) and preservation.
• Thicker (≈ 20%) and longer (≈30%) myotubes.
• Greater sarcomeric patterning. 
136  
Myoblasts  Cyclic  Collagen  10% stretch 3 times/minute for the first 5 min of every h, duration ranging from 5 days to 3 weeks.  • Uni-directional alignment achieved within 5 days and maintained in vivo.
• Depolarization with KCI resulted in measurable contractile responses after 3 weeks; however, dissipated within 1 week in vivo
180  
Myoblasts  Static  Collagen  Sustained uniaxial tension (unspecified)  • Sustained uniaxial tension promotes myoblast fusion to form aligned multi-nucleate myotubes.
• Myotube alignment parallel to strain.
• Surge in both IGF-1 Ea and MGF expression on day 3 of the developing construct 
181  
Myoblasts  Static/cyclic  Collagen sponge  7.5%–15% static strain for 6 h.  • Increase MMP-2 expression and matrix remodeling.
• Upregulation depends on amount and type of strain. 
182  
C2C12 myoblasts  Static  Fibrin gel  25% strain  • Cells aligned parallel to strain direction.
• Cell proliferation was identical to that cell alignment.
• Adjacent cells were not in contact, due to fibrin bundles restricting cellular activity to the space in between.
• Increased proliferation compared to control, and 50% strain produced enhanced proliferation compared to 25% strain. 
183  
Muscle-derived cells (MDCs)  Static  Collagen constructs  Isometric tension (no strain)  • Aligned myotubes.
• Tension increased as structure contracted, causing concave center. 
184  
C2C12 myoblasts  Cyclic/static  Collagen matrix  Rotation of strain increase, strain maintenance, strain release, and rest for 6 days  • IGF-IEa constitutively expressed in myoblasts and myotubes under endogenous tension.
• Single ramp stretch and cyclic loading increased mechano-growth factor (MGF), whereas static strain did not. 
185  
C2C12 myoblasts  Cyclic  Aliphatic diisocyanate-based polyurethane (PU) fibers  (i) 5% strain at 1 Hz, (ii) 10% strain at 1 Hz, and (iii) 5% strain at 1 Hz with constant 5% pre-stretch. 1 h on, 5 h rest.  • Myotubes exhibited evidence of overstretching at both 5% and 10% strains.
• Very few myoblasts fused to form myotubes on the scaffold, and those that did were short and lacked sarcomeric structures.
• Overstretching resulted in a proliferation effect, as seen in response to injury.
• Produced increased fraction of striated myotubes from 75% to 85%. 
186  
Muscle precursor cells (MPCs)  Cyclic  Acellular collagen matrices  10% strain stretched 3 times per minute for the first 5 mins of every hour for periods ranging from 5 days to 3 weeks.  • Unidirectional orientation within 5 days.
• Generation of contractile response after 3 weeks.
• Conditions tissue produced contractile responses with 1% and 10% specific force in vivo, whereas non-conditioned tissue does not. 
180  
MPCs  Cyclic  Porcine bladder acellular matrix scaffold  10% strain, three times per minute for the first 5 min of every hour, for one week.  • In vivo implantation resulted in decreased recovery time. This, however, also incorporated electrical stimulation.  187  
MDCs)  Cyclic  Bladder acellular matrix scaffold  10% strain, three times per minute for the first 5 min of every hour, for 5–7 days  • Cells exhibited an elongated and aligned morphology.
• TEMR-1SPD produced a significant reduction in nuclei, whereas TEMR-2SPD produced a significant increase.
• Strain produced higher levels of multinucleated cells. 
188  
Fibroblasts  Cyclic  Collagen gel  480 dynes/h, 15-min intervals of load, rest, unload, rest for 16 h. Initially loaded with 120 Dyne force.  • High aspect ratio configurations produced dramatic alignment parallel with the axis of strain, shape control of cells, elongated, and approximately bipolar, whereas low aspect ratio configurations did not (after 16 hours).
• Alignment is predominately “nose to tail” configuration.
• No alignment of collagen fibrils was evident. 
189  
Skeletal muscle cells  Cyclic  Collagen/MATRIGEL  5%, 10%, and 15% strain. 5 stretch and relaxation cycle. Repeated 3 times with rest in between. Total duration of 30 minutes.  • Internal longitudinal tensions align cells into parallel arrays which then fused into aligned multinucleated myofiber.
• Mechanical conditioning of HBAMs improved HBAM myofiber diameter and area percentage. 
190  
Fibroblasts  Cyclic  Polyethylene terephthalate (PET)  5%, 0.5 Hz, 1 h daily, 14 days  • GAG per DNA was statistically lower in the bioreactor as compared to the static culture.
• Stretched and non-stretched studies produced a sevenfold increase in cell numbers compared to Petri-dish cultures.
• Little differences seen between stretch and non-stretched specimens. 
191  
Cardiac progenitor cells (CPCs)  Cyclic  Decellularized human skin (d-HuSk)  10% strain, 1 Hz, for 7 days.  • Dynamic culture promoted hCPC migration toward the inner layers of the scaffolds.
• Upregulated cardiac alpha actin.
• CD117 and TBX3 significantly down regulated. 
192  
MSCs  Cyclic  Hyaluronate band with 3D fibrin hydrogel  10% strain, 1 Hz for 72 h.  • Upregulation of SCX, type 1 collagen, decorin, and tenascin-C.  193  
Mesenchymal stromal cells  Cyclic  Non-crosslinked porcine-derived acellular dermal matrix  1%, 5%, and 10% strains, 0.2, 0.33, and 0.5 Hz.  • Sustained SCX and tenascin C increase.
• Wnt 16 and tenomodulin time-dependent increase. 
194  
hADSCs  Cyclic  Spider silk with type 1 collagen  2% strain for 8 h daily at 1 Hz. 16 h rest between every stressing. 21 days duration  • Cells and silk fibers orientated and aligned toward the direction of mechanical stimulation.
• Cell density decrease. 
195  
C2C12 Myoblasts  Cyclic  Collagen gel  3% strain, 1 Hz, 4 h/day, for 2 days  • Cell diameter decreased.
• Elongated and aligned along direction of strain.
• Increased differentiation and myotube size. 
142  
C2C12 Myoblasts  Cyclic  DegraPol microfibrous membranes  Phase C: 1 mm, 0.5 Hz (30 s rest between each pulse), 28 min rest  • Eightfold increase in myosin content.  196  
C2C12 Myoblasts  Static  GelMA  0%–45% strain for 5 days  • Improved cell spreading.
• Improved elongation.
• Improved alignment along direction of strain. 
197  
C2C12, MPCs  Cyclic  Fibrin  2-day uniaxial ramp stretch (0%–2%), followed by 2%–6% dynamic stretch (3 h on, 3 h off)  • Orientation of myotubes in direction of attachment between anchor points.
• Increased cross striations in MPC constructs. 
144  
Cell Type Type Substrate/scaffold Mechanical parameters Major findings Reference
C2C12  Static  Fibrin  10% static strain for 6 h, followed by 18-h rest at 3% static strain. Repeated for 6 days.  • Higher degree of alignment (>85% in direction of strain, residual 15 showed a 15%–30% deviation) and preservation.
• Thicker (≈ 20%) and longer (≈30%) myotubes.
• Greater sarcomeric patterning. 
136  
Myoblasts  Cyclic  Collagen  10% stretch 3 times/minute for the first 5 min of every h, duration ranging from 5 days to 3 weeks.  • Uni-directional alignment achieved within 5 days and maintained in vivo.
• Depolarization with KCI resulted in measurable contractile responses after 3 weeks; however, dissipated within 1 week in vivo
180  
Myoblasts  Static  Collagen  Sustained uniaxial tension (unspecified)  • Sustained uniaxial tension promotes myoblast fusion to form aligned multi-nucleate myotubes.
• Myotube alignment parallel to strain.
• Surge in both IGF-1 Ea and MGF expression on day 3 of the developing construct 
181  
Myoblasts  Static/cyclic  Collagen sponge  7.5%–15% static strain for 6 h.  • Increase MMP-2 expression and matrix remodeling.
• Upregulation depends on amount and type of strain. 
182  
C2C12 myoblasts  Static  Fibrin gel  25% strain  • Cells aligned parallel to strain direction.
• Cell proliferation was identical to that cell alignment.
• Adjacent cells were not in contact, due to fibrin bundles restricting cellular activity to the space in between.
• Increased proliferation compared to control, and 50% strain produced enhanced proliferation compared to 25% strain. 
183  
Muscle-derived cells (MDCs)  Static  Collagen constructs  Isometric tension (no strain)  • Aligned myotubes.
• Tension increased as structure contracted, causing concave center. 
184  
C2C12 myoblasts  Cyclic/static  Collagen matrix  Rotation of strain increase, strain maintenance, strain release, and rest for 6 days  • IGF-IEa constitutively expressed in myoblasts and myotubes under endogenous tension.
• Single ramp stretch and cyclic loading increased mechano-growth factor (MGF), whereas static strain did not. 
185  
C2C12 myoblasts  Cyclic  Aliphatic diisocyanate-based polyurethane (PU) fibers  (i) 5% strain at 1 Hz, (ii) 10% strain at 1 Hz, and (iii) 5% strain at 1 Hz with constant 5% pre-stretch. 1 h on, 5 h rest.  • Myotubes exhibited evidence of overstretching at both 5% and 10% strains.
• Very few myoblasts fused to form myotubes on the scaffold, and those that did were short and lacked sarcomeric structures.
• Overstretching resulted in a proliferation effect, as seen in response to injury.
• Produced increased fraction of striated myotubes from 75% to 85%. 
186  
Muscle precursor cells (MPCs)  Cyclic  Acellular collagen matrices  10% strain stretched 3 times per minute for the first 5 mins of every hour for periods ranging from 5 days to 3 weeks.  • Unidirectional orientation within 5 days.
• Generation of contractile response after 3 weeks.
• Conditions tissue produced contractile responses with 1% and 10% specific force in vivo, whereas non-conditioned tissue does not. 
180  
MPCs  Cyclic  Porcine bladder acellular matrix scaffold  10% strain, three times per minute for the first 5 min of every hour, for one week.  • In vivo implantation resulted in decreased recovery time. This, however, also incorporated electrical stimulation.  187  
MDCs)  Cyclic  Bladder acellular matrix scaffold  10% strain, three times per minute for the first 5 min of every hour, for 5–7 days  • Cells exhibited an elongated and aligned morphology.
• TEMR-1SPD produced a significant reduction in nuclei, whereas TEMR-2SPD produced a significant increase.
• Strain produced higher levels of multinucleated cells. 
188  
Fibroblasts  Cyclic  Collagen gel  480 dynes/h, 15-min intervals of load, rest, unload, rest for 16 h. Initially loaded with 120 Dyne force.  • High aspect ratio configurations produced dramatic alignment parallel with the axis of strain, shape control of cells, elongated, and approximately bipolar, whereas low aspect ratio configurations did not (after 16 hours).
• Alignment is predominately “nose to tail” configuration.
• No alignment of collagen fibrils was evident. 
189  
Skeletal muscle cells  Cyclic  Collagen/MATRIGEL  5%, 10%, and 15% strain. 5 stretch and relaxation cycle. Repeated 3 times with rest in between. Total duration of 30 minutes.  • Internal longitudinal tensions align cells into parallel arrays which then fused into aligned multinucleated myofiber.
• Mechanical conditioning of HBAMs improved HBAM myofiber diameter and area percentage. 
190  
Fibroblasts  Cyclic  Polyethylene terephthalate (PET)  5%, 0.5 Hz, 1 h daily, 14 days  • GAG per DNA was statistically lower in the bioreactor as compared to the static culture.
• Stretched and non-stretched studies produced a sevenfold increase in cell numbers compared to Petri-dish cultures.
• Little differences seen between stretch and non-stretched specimens. 
191  
Cardiac progenitor cells (CPCs)  Cyclic  Decellularized human skin (d-HuSk)  10% strain, 1 Hz, for 7 days.  • Dynamic culture promoted hCPC migration toward the inner layers of the scaffolds.
• Upregulated cardiac alpha actin.
• CD117 and TBX3 significantly down regulated. 
192  
MSCs  Cyclic  Hyaluronate band with 3D fibrin hydrogel  10% strain, 1 Hz for 72 h.  • Upregulation of SCX, type 1 collagen, decorin, and tenascin-C.  193  
Mesenchymal stromal cells  Cyclic  Non-crosslinked porcine-derived acellular dermal matrix  1%, 5%, and 10% strains, 0.2, 0.33, and 0.5 Hz.  • Sustained SCX and tenascin C increase.
• Wnt 16 and tenomodulin time-dependent increase. 
194  
hADSCs  Cyclic  Spider silk with type 1 collagen  2% strain for 8 h daily at 1 Hz. 16 h rest between every stressing. 21 days duration  • Cells and silk fibers orientated and aligned toward the direction of mechanical stimulation.
• Cell density decrease. 
195  
C2C12 Myoblasts  Cyclic  Collagen gel  3% strain, 1 Hz, 4 h/day, for 2 days  • Cell diameter decreased.
• Elongated and aligned along direction of strain.
• Increased differentiation and myotube size. 
142  
C2C12 Myoblasts  Cyclic  DegraPol microfibrous membranes  Phase C: 1 mm, 0.5 Hz (30 s rest between each pulse), 28 min rest  • Eightfold increase in myosin content.  196  
C2C12 Myoblasts  Static  GelMA  0%–45% strain for 5 days  • Improved cell spreading.
• Improved elongation.
• Improved alignment along direction of strain. 
197  
C2C12, MPCs  Cyclic  Fibrin  2-day uniaxial ramp stretch (0%–2%), followed by 2%–6% dynamic stretch (3 h on, 3 h off)  • Orientation of myotubes in direction of attachment between anchor points.
• Increased cross striations in MPC constructs. 
144  

Overall, the mechanical stimulation of skeletal muscle cells promotes a higher degree of alignment in the direction of the strain,136 thicker and longer myotubes,136 greater sarcomeric patterning,136 promoted myoblast fusion,181 significantly higher MMP-2 expression,182 enhanced proliferation,183 unidirectional nose to tail orientation,189 generation of contractile force,180 elongated and aligned morphology,188 higher level of multinucleated cells,190 promoted hCPC migration,192 and upregulation of scleraxis, collagen I, decorin, and tenascin C193 while maintaining comparable cell survival.198 Similarly, engineered tendon constructs promoted alignment,143,161,167 ECM production,157,161,179 maturation,160,169,178 and expression of tenogenic markers.143,175,178 As depicted above, a uniaxial strain has a profound effect on various cellular activities. Further studies are reviewed by Somers et al.139 To induce this strain, several strain-inducing bioreactors have been commercialized and are available from a number of companies (Fig. 6) including Strex/Strexcell, Electroforce, Flexcell, Biodynamic, Tissue growth technologies, Ebers, Cellscale, and Con Whitley scientific.

FIG. 6.

Example of commercialized bioreactors capable of applying a uniaxial strain to 3D tissue-engineered constructs. (a) Ebers TC-3 bioreactor (ceeiaragon.es). (b) Biodynamic 5270* bioreactor (docplayer.net). (c) Cellscale MCT6 bioreactor (cellscale.com).

FIG. 6.

Example of commercialized bioreactors capable of applying a uniaxial strain to 3D tissue-engineered constructs. (a) Ebers TC-3 bioreactor (ceeiaragon.es). (b) Biodynamic 5270* bioreactor (docplayer.net). (c) Cellscale MCT6 bioreactor (cellscale.com).

Close modal

The vast majority of strain bioreactors have been custom designed and are primarily research focused, such as those seen in Table VII. Lab-specific bioreactors commonly incorporate a 3D-printed frame, transparent Perspex lid, stepper motor linear actuator, and a wide range of clamps to constrain the engineered tissue, similar to commercialized devices. Here, a wide range of materials, including Perspex, silicon, aluminum, polycarbonate, resin, polyurethane, Teflon, and stainless steel, are utilized, which all achieve self-proclaimed biocompatibility.189–191,199–201 Given this large range of materials, however, various cleaning protocols and surface treatments are mentioned. For example, to improve biocompatibility and hydrophobicity, sylgard-184 has previously been coated on the bottom of a resin-based bioreactor chamber.202 Similarly, basic features, such as gas vents, are common throughout, where only the most intricate designs incorporate mechanical feedback (via in line or bending beam load cells) and automated media exchange systems. Finally, there are a finite number of designs that incorporate electrical stimuli with a uniaxial strain;158,186,200,203–205 however, they will not be summarized here as they are not the focus of this review.

TABLE VII.

Custom-made bioreactors capable of providing uniaxial strain on three-dimensional tissue-engineered constructs. ✓ indicates transparent at top and bottom. ✓✓ indicates transparent at top and bottom.

Stretch source Separate chambers? No of samples Metallic free media? Sensory feedback (mechanical)? Automated media exchange Media exchange ports? Gas exchange filter? Transparent? Reference
Stepper motor driver  ✗  Unspecified  ✗  ✗  ✗  ✗  ✗  ✓✓  199  
Lead screw micro-stepping motor  ✓  ✗  ✓  ✗  ✗  ✓  ✓✓  189  
Linear actuator stepper motor  ✓  ✓  ✓  ✗  ✓  ✓  ✓  190  
Linear actuator stepper motor  ✓  ✗  ✓  ✗  ✓  ✓  ✓✓  206  
Linear actuator  ✗  12  ✗  ✗  ✓  ✓  ✗  ✓✓  191  
Micro-stepper motor  ✓  ✗  ✗  ✗  ✗  ✓  ✓✓  207  
Stepper motor  ✓  ✓  ✗  ✗  ✓  ✓  ✓  192  
Linear motor  ✓  ✓  ✗  ✗  ✗  ✗  ✓✓  193  
Micro-gear motor  ✓  ✓  ✓  ✗  ✗  ✓  ✓  208  
Motor (unspecified)  ✓  ✗  ✓  ✓  ✓  ✓  ✓  194  
Driving motor (unspecified)  ✓  ✗  ✓  ✓  ✓  ✗  ✓✓  209  
Drive motor (unspecified)  ✗  ✗  ✗  ✗  ✗  ✓  ✓✓  195  
Electric servomotors  ✓  ✗  ✓  ✓  ✓  ✗  ✓✓  210  
Linear motor  ✗  10  ✓  ✗  ✗  ✗  ✗  ✓✓  180  
Linear activator stepper motor  ✓  12  ✗  ✗  ✗  ✗  ✗  ✓  211  
Stepper motor actuator  ✓  ✗  ✗  ✗  ✗  ✗  ✓✓  212  
Linear actuator  ✓  ✓  ✗  ✗  ✗  ✓  ✓  202  
Motor-ball screw actuator  ✓  ✗  ✗  ✗  ✗  ✗  ✓  166  
Linear actuator stepper motor  ✗  ✓  ✓  ✗  ✗  ✗  ✓  141  
Linear actuator stepper motor  ✓  ✗  ✓  ✗  ✗  ✓  ✓  213  
Stretch source Separate chambers? No of samples Metallic free media? Sensory feedback (mechanical)? Automated media exchange Media exchange ports? Gas exchange filter? Transparent? Reference
Stepper motor driver  ✗  Unspecified  ✗  ✗  ✗  ✗  ✗  ✓✓  199  
Lead screw micro-stepping motor  ✓  ✗  ✓  ✗  ✗  ✓  ✓✓  189  
Linear actuator stepper motor  ✓  ✓  ✓  ✗  ✓  ✓  ✓  190  
Linear actuator stepper motor  ✓  ✗  ✓  ✗  ✓  ✓  ✓✓  206  
Linear actuator  ✗  12  ✗  ✗  ✓  ✓  ✗  ✓✓  191  
Micro-stepper motor  ✓  ✗  ✗  ✗  ✗  ✓  ✓✓  207  
Stepper motor  ✓  ✓  ✗  ✗  ✓  ✓  ✓  192  
Linear motor  ✓  ✓  ✗  ✗  ✗  ✗  ✓✓  193  
Micro-gear motor  ✓  ✓  ✓  ✗  ✗  ✓  ✓  208  
Motor (unspecified)  ✓  ✗  ✓  ✓  ✓  ✓  ✓  194  
Driving motor (unspecified)  ✓  ✗  ✓  ✓  ✓  ✗  ✓✓  209  
Drive motor (unspecified)  ✗  ✗  ✗  ✗  ✗  ✓  ✓✓  195  
Electric servomotors  ✓  ✗  ✓  ✓  ✓  ✗  ✓✓  210  
Linear motor  ✗  10  ✓  ✗  ✗  ✗  ✗  ✓✓  180  
Linear activator stepper motor  ✓  12  ✗  ✗  ✗  ✗  ✗  ✓  211  
Stepper motor actuator  ✓  ✗  ✗  ✗  ✗  ✗  ✓✓  212  
Linear actuator  ✓  ✓  ✗  ✗  ✗  ✓  ✓  202  
Motor-ball screw actuator  ✓  ✗  ✗  ✗  ✗  ✗  ✓  166  
Linear actuator stepper motor  ✗  ✓  ✓  ✗  ✗  ✗  ✓  141  
Linear actuator stepper motor  ✓  ✗  ✓  ✗  ✗  ✓  ✓  213  

It is widely hypothesized that engineered tissues that mimic embryonic (as opposed to adult) mechanical properties will produce greater biological responses. Although producing promising results, the paucity of knowledge surrounding MTJ's embryonic development, particularly that of skeletal muscle, makes outcomes difficult to quantify. Currently, attempts to engineer the MTJ rely on heterogeneous scaffolds aimed to provide the appropriate environment for skeletal muscle and tendon tissues simultaneously. In doing so, the interfacial region may be analyzed for interface-specific genes, proteins, and architectural characteristics. To date, although favorable results have been achieved, the mechanical discrepancy and architecture complexity between skeletal muscles and tendons is yet to be mimicked. To overcome this issue, many groups identify the possibility to mature myotendinous constructs in a dynamic, strain-inducive bioreactor as a means of tissue maturation;41,111,124 however, few have advanced to this stage.121 The translation to dynamic cultures must be reliant on the established body of knowledge known about individual tissues, such as skeletal muscles and tendons separately. Here, the significant progress of engineering tissues may be merged to develop targeted strategies focused on interfacial regions such as the MTJ. Future studies are advised to direct their attention to the translation from static to strain-inducive dynamic bioreactors, where a cyclic regime following an initial static strain period shows the greatest results.

The authors would like to acknowledge the National Health and Medical Research Council (Ideas Grant 2002723) and FSHD Global Research Foundation for supporting this work. F.S. was supported by an RMIT STEM Scholarship.

The authors have no conflicts to disclose.

Ethics approval is not required.

Finn Edward Snow: Conceptualization (lead); Visualization (lead); Writing – original draft (lead); Writing – review & editing (lead). Cathal O'connell: Writing – review & editing (supporting). Peiqi Yang: Writing – review & editing (supporting). Magda Kita: Conceptualization (supporting). Elena Pirogova: Supervision (equal); Writing – review & editing (equal). Richard Williams: Writing – review & editing (supporting). Robert Kapsa: Conceptualization (supporting); Supervision (equal); Writing – review & editing (supporting). Anita Quigley: Supervision (equal); Writing – review & editing (equal).

Data sharing is not applicable to this article as no new data were created or analyzed in this study.

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