Microscopy analysis

For optical analysis, threads were put in a
homemade thin glass-chamber filled with fib-
rillogenesis buffer to prevent dehydration dur-
ing observation. Polarized light microscopy
was performed using a transmission Nikon
Eclipse E600 Pol, equipped with crossed po-
larizers and a Nikon DXM 1200CCD cam-
era. For electron microscopy analysis, all sam-
ples were fixed using 3.63% glutaraldehyde in
a cacodylate/saccharose buffer (0.05 M/0.3
M, pH 7.4) for 1 hour at 4°C. For scan-
ning electron microscopy (SEM), the sam-
ples were then washed three times in a ca-
codylate/saccharose buffer (0.05 M/0.3 M,
pH 7.4) and dehydrated through successive
increasing concentration ethanol baths from
70% to 100% alcohol. Thereafter, sam-
ples were dried at the carbon dioxide crit-
ical point and gold sputtered (20 nm) for
analysis. Samples were observed at 200M,
1kM, 5kM and 30kM with a Hitachi S-3400N
SEM operating at 8 kV with a probe cur-
rent around 95 µA. For transmission elec-
tron microscopy (TEM), samples were post-
fixed with 2% osmium tetroxide in cacody-
late/saccharose buffer solution, washed in ca-
codylate/saccharose solution, dehydrated in
successive ethanol baths (from 50° to 100°),
and finally embedded in araldite. Embed-
ded samples were sectioned on an Ultracut
Reichert Jung as already described [37]. 500
µm thin sections were stained with toluïdine
blue and observed on a Nikon Eclipse E600Pol
microscope equipped with a 40x/0.65 NA ob-
jective and a DXM 1200 CCD camera. 70 nm
ultrathin sections were contrasted with uranyl
acetate and observed on a transmission elec-
tron microscope FEI Tecnai spirit G2 operat-
ing at 120 kV. Images were recorded on a CCD

Camera (Orius Gatan 832 digital) at 4.4kM
and 15kM.

Mechanical tests

Mechanical tests were performed using a
BOSE® ElectroForce® 3200 Series II Test
Instrument. Specimens were attached to the
stainless steel crossheads by gluing each ex-
tremity with cyanoacrylate adhesive drop.
Sample and crossheads were immersed in
a glass Petri dish filled with fibrillogenesis
buffer at room temperature. The threads
diameter, D was determined by optical mi-
croscopy assuming full cylindrical symmetry.
Mechanical loading of the thread was achieved
by driving one crosshead at a constant veloc-
ity of 0.1 mm per second with a stepping mo-
tor (0.1-micron steps). The loading force F
was measured with a 220 g load cell. We de-
fined the nominal tensile stress as σ = $\frac{4F}{\pi D^2}$.
The nominal strain is defined as e = (L–L0)/
L0, where L is the distance between the glue
points and L0 the initial length. All measure-
ments were performed up to the specimens
rupture. The resulting stress-strain curves
were plotted and analyzed. At least 4 dif-
ferent threads were tested for each condition.

3 Results

Since the goal of our work was to find the
best conditions to synthesize dense colla-
gen threads that would be stable over time,
present anisotropic organization of collagen
fibrils and have the highest mechanical prop-
erties, we have, all along our work, narrowed
down the conditions to determine those giving
threads of optimized characteristics.
## Extrusion and Maturation

All threads look similar at the time of the extrusion in all buffers (Figure 2; Figure SI-1). Their surfaces present a "shark-skin" roughness, characteristic of so-called viscoelastic fluids (Larson, Denn). The roughness amplitude is of the order of 20 µm for a thread diameter of 710 µm (2.8%) as measured on Figure 2A. Noticeably, within experimental error and after averaging the shark-skin roughness, the swelling ratio Q, measured just downstream of the die exit (the so called die swell, of mere viscoelastic origin ??) remains close to 1. Figure 2 illustrates the threads synthesized

using buffer 1, 7 and 8 as visualized during extrusion (Fig. 2A,B) thanks to the optical system, and after the two weeks of maturation under polarized light (C) and in between crossed polarizers (D). We see that, depending on the buffer, the thread diameters evolve differently. No changes are observed in condition 7, while swelling and shrinkage are respectively observed in conditions 1 and 8. In Figure 2D, threads are positioned at nearly 45° to both polarizers. They appear homogeneously bright along the thread axis. In the case of condition 1, the bright signal concerns only the central part of the thread.
<figure><img src="image_113.png" /></figure>

Figure 2: 60 mg/ml collagen thread extrusion and maturation. Figure 2A,B shows the threads extrusions in extrusion buffers 1, 7 and 8. Figure 2A corresponds to the extrusion through a 650 µm inner diameter needle in condition 1, while Figure 2B corresponds to extrusion in the three buffers through needle of 390 µm inner diameter. A sharkskin pattern is observed on all threads, although it less pronounced in condition 8. Figure 2C shows the threads after the two weeks of maturation in the respective extrusion buffers. We observe the increase and shrinkage in diameters of threads extruded in condition 1 and 8 respectively, while the diameter of the threads in condition 7 does not evolve. The thread extruded and matured in buffer 1 shows a core shell profile, the core being denser than the shell. Figure 2C, shows the matured threads as seen in between cross-polarizers (crossed arrowed) and at around 45° from both polarizers. We see that in all the three conditions, the threads appear very bright indicating an overall alignment of the molecules. For the thread made in buffer 1, only the core shows birefringence. The shell, delineated as seen in the inset appears dark and remains invisible. This suggests a loss of molecular orientation along the thread axis.
Figure 3A shows the evolution of the threads diameter over time during the maturation process for all buffers and Figure 3B shows the plot of the long term swelling ratio $Q$, here of osmotic nature, in function of the ionic strength. The former show that diameters evolve over 10 days and then reach a plateau. The latter reveals three groups of behaviors.

For the first one, consisting in a cluster of conditions 1 to 6, we observe a swelling of the threads and find a $Q > 1$. For the second one, consisting in buffer 7, the diameters remain constant over time, with a ratio $Q = 1$. Finally for the third one, buffer 8, we observe a shrinkage of the collagen thread diameter ($Q < 1$). Data are summarized in Table 1.

<figure><img src="image_114.png" /><figcaption>Figure 3: Evolution of thread diameters upon maturation for the different buffers. After extrusion, the threads are left into the buffers for two weeks for a maturation process. The diameters of the thread have been measured regularly along the two weeks. We see that depending on the buffer some thread see their diameter increasing some decreasing. In condition 7 the diameter remains relatively stable throughout the two weeks period.</figcaption></figure>

At this stage, we chose to focalize only on conditions 1, as representative of the cluster of conditions 1 to 6, as well as condition 7 and 8.

Structural characterization

Figure 4 shows the external morphology and ultrastructures of threads obtained with fibrillogenesis buffers 1, 7 and 8 as seen by SEM. At low magnifications (200M, 1kM, 5 kM, top three lines), we observe the overall characteristics of the threads. In condition 1 (A),

threads present an outskirt of large "hairs" while, in condition 8 (C), threads present a smooth shell with hills and valleys with, locally some cracks revealing the internal ultrastructure. In condition 7 (B), the threads are rather smooth with a few "hairs" sticking here and there. At higher magnifications (bottom line), in condition 1, we see large (≈ 560 nm) linear fibers or bunches of compact fibrils (≈150 nm) on top of more tortuous fibrils (≈ 70 nm). The fibers present a large striation of ≈180 nm and sub-striations of approximately 67 nm. In condition 8, we ob-
serve a smooth top shell under which linear
fibers and fibrils (approximatively 150 nm di-
ameter can be seen. The fibers present large
striations of varied length and a sub striation
of 67 nm. Clear helical fibers are also seen. In
condition 7, densely packed wavy fibrils with
diameters of 100 nm in average are seen with

the 67 nm cross-striation. The fibrils are fused
in a more or less wavy pattern. Some larger
striations can also be seen. The measurement
of the diameters are indicated in Table SI-1
together with a scheme of the threads profile
for each group.

<figure><img src="image_115.png" /><figcaption>Figure 4: Ultrastructure of threads after maturation in buffers 1, 7 and 8 as seen by SEM. Figure 4A, B, C correspond respectively to threads synthesized in condition 1,7 and 8. The magnification increases from top to bottom (200M, 1kM, 5kM, 30kM). The cross-striated appearances of fibrils and fibers are clearly seen at 30kM in all three conditions. In condition 1, a 67nm striation in further integrated in a wider one of 180 nm. In condition 7, the 67 nm striations dominate although curly or wavy undulated fibrils, fused or not, are seen locally, which wavelength or pitch give yet other repeated distances. In condition 8, the 67 nm striation is seen both on straight and on helical fibrils.</figcaption></figure>
<figure><img src="image_116.png" /><figcaption>Figure 5: Ultrastructure of threads after maturation in buffers 1, 7 and 8 as seen by TEM (transverse sections). Figure 5A, B, C correspond respectively to threads synthesized in condition 1, 7 and 8. Top: view at low magnification from the outskirt (when present) to the inner part of the thread. Bottom: view at higher magnification from the outskirt to the inner part of the thread. Collagen fibrils appear cut in all directions from longitudinal to transverse, in which case they appear as dots. In A, the thread presents three different structural characteristics when going from the outskirt of the thread to the inner part. In B and C the threads present more homogeneous appearances. In B two types of fibrils are observed, some with greater diameters than others while in C the fibrils appear homogenous in diameters but with a reduced diameter.</figcaption></figure>

Figure 5 shows the distribution of the fibrils as well as their morphology as seen by transmission electron microscopy (Table 2).

Figure 5A, 5B and 5C, are cross-sections of threads respectively made in buffer 1, 7 and 8, at low magnification on the top part and
high magnification on the lower part. The
pictures show view from the external (Ex)
part to the internal part (In) of the threads
as indicated by the side arrows. In 5A, we
see that the thread present very large fibrillar
entities (≈ 500 nm) on the outmost edge of
the thread. The shell is constituted of large
(≈ 250 nm) individualized fibers that do not
present any specific alignement. In the core
part, thin fibrils (≈ 40 nm) are more tightly
packed, most of them being cut perpendicular
to their axis, giving the dotted features on the
photo. Fewer appear cut along their longitu-
dinal axis. Figure 5B shows ultra structure of
a thread made in buffer condition 7. We see a
cohabitation of well defined thick (≈ 80 nm)
fibrils presenting the 67 nm cross-striation, to-
gether with packages or aggregates of thin fib-
rils (≈ 20 nm). In the external part, the fibrils
are seen in majority aligned to the longitudi-
nal axis of the thread and thus appear as dots.
In Figure 5C, we observe a section of a thread
made in buffer condition 8. Here, we observe
dense packages (≈ 200 nm) of small fibrils (≈
25 nm) in majority aligned with the axis of
the thread in the external part and more ran-

domly orientated especially in the core. The
measurement of the diameters are indicated
in Table SI-1 together with a scheme of the
threads profile for each group

Mechanical properties

Figure 6 shows the results of the mechanical
tests and the threads after breakage. Figure
6A shows typical stress curves in function of
the strain obtained for the 3 selected condi-
tions 1, 7 and 8. Figure 6B shows typical pro-
file of thread extremities after breakage. Fig-
ure 6C shows Young modulus and Ultimate
Tensile Strength obtained for the three con-
ditions as measured over more than 4 threads
for each of them. We see that the highest me-
chanical properties are obtained for threads
extruded and maturated in PBS5X that ex-
hibit a Young's modulus of 1.66 ± 0.45 MPa
and an ultimate tensile strength of 0.33 ±
0.08 MPa, while threads synthesized in two
other conditions yield values of 0.30 ± 0.15
MPa and 0.13 ± 0.01 MPa respectively for
the Young modulus and UTS in condition 1
and of 0.88 ± 0.14 MPa and 0.16 ± 0.02 MPa
respectively for condition 8.
<figure><img src="image_117.png" /><figcaption>Figure 6: *Mechanical characterization.* Figure 6A shows the typical profiles obtained for the stress-strain curves in the three conditions 1, 7, and 8. Figure 6B shows the three threads at the point of breakage. Note that the breakage end is rather sharp in condition 1, intermediate in condition 7 and elongated in condition 8. Figure 6C gives Young Modulus and Ultimate Tensile Strength for the three conditions 1, 7 and 8 as measured over 4 threads.</figcaption></figure>

# 4 Discussion

The main objective of this work was to optimize the synthesis protocol in order to reproducibly obtain, homogeneous, high-density, anisotropic collagen threads with the best mechanical properties, that is high Young modulus and high UTS. As criteria we have selected: 1) the stability of the thread upon storage in the fibrillogenesis buffer to avoid swelling or contraction, 2) the ultrastructure of the threads, 3) their mechanical properties.

Another figure of merit is the linearity of the thread. This one can be affected by several gross instabilities [9] [9]. It is controlled both by the extrusion velocity and the nee-

dle diameter. In the present work, we chose the extrusion speed according to the analysis previously made on alginate samples [30] so as to limit instabilities. Since threads extruded with D0= 650 µm or 390 µm diameter with adapted extrusion speeds had similar structural and mechanical properties, we did most of the present study with a needle of 390 µm diameter to limit the quantity of collagen used. Noteworthy, diameters around 390 µm make threads easier for post-production processes (such as weaving), as compared to collagen threads with similar properties produced in the other studies, which diameters range from 50 to 200 µm sys-
tem [11],[5],[17],[7].

Remarkably, upon extrusion, the threads presented the same diameter as the inner diameter of the needle with no die swelling effect, often observed in the case of highly viscous solutions and that can reach up to 2 times the internal diameter of the needle [24]. This observation is puzzling since, even for a newtonian fluid, die swelling Q = 1.1 [35] is observed. Gross slip would suppress wall shear and normal stresses responsible for die swelling. It is however unlikely since the presence of shark-skin is indicative of a strong adhesion of the solution at the wall. At this stage, one can tentatively suggest that the strong anisotropy of the collagen solution could be one factor precluding die swelling.

Regarding the stability of the thread diameter upon fibrillogenesis and storage, Pins et al. (1997)[31] have shown the importance of fibril alignment on the mechanical properties obtained, starting from a 10 mg/ml collagen thread, and using a multistep procedure with two air-drying-rehydration stages, interspersed by a stretching step of different spans to align and condense the collagen fibrils, and followed or not by chemical crosslinking. The authors showed that on the final products, synthesized in the absence of crosslinking, the linear Young modulus increased from 1.8 MPa at 0% stretch to 46 MPa at 50% stretch, and UTS from 0.9 MPa to 7 MPa respectively. They made no ultrastructural study but showed the overall increased orientation by the increase of birefringence values. They concluded that collagen fibrils alignment, within the thread, increased its mechanical properties. Therefore the stability of the thread diameter after extrusion, that is limited swelling or contraction, stood as a

criteria on to select the best conditions. This stability indeed provided evidence of maintenance over time of collagen density as well as limited reorganization of collagen fibrils, as a consequence of the conservation of the shear induced alignment made by the extrusion.

Now, depending on the buffer we used, we observed swelling, contraction or stability of the thread. This variation in diameter and different structures obtained with different fibrillogenesis buffers most likely reflect the involvement of hydrogen and electrostatic bonding in collagen supramolecular organizations in our systems. Several studies on the effect of ionic force and pH in diluted collagen solutions have already been reported. The intermolecular space between triple helices has indeed been negatively correlated to the osmotic pressure applied [20]. Wood et al. also reported that increasing ionic strength between 130 and 310 mM increased fibril width [44]. This increase of fibril diameter was also observed in the case of collagen solutions at high concentration (100 mg/mL)[14]. Amongst all our conditions, the threads made in conditions 7 were the least affected upon storage. The overall alignment of the threads was confirmed by polarized light microscopy for the three conditions 1, 7 and 8, but was limited to the core of the thread in condition 1. This showed that, in our case, the swelling had disturbed the alignment of the fibrils at the surface of the thread. The ultra-structural characterization of the different threads confirmed the core shell nature of the thread synthesized in buffer 1. It revealed larger fibrils and wider interspaces in the external part than the internal. The wide interspace could potentially be more favorable to cell invasion. However, the inhomogeneity
of the thread with the very large aggregated fibers at its outmost edge, together with the pronounced difference from external to internal part, was opposite to the criteria that we had set. Threads made in condition 7 and 8, both presented ultra-structure more homogeneous than in condition 1. However, in condition 8, the thread showed a diameter contraction together with the presence of a crusty and smooth surface. We believe that this crust would be unfavorable for cell attachment and colonization. It may result from an artifact of SEM preparation although TEM images also reveal a rather continuous limit of very small fibrils. In fact, the small diameter of those fibrils, smaller than those found in living tissues, exception made of cornea, may potentially be problematic. Thus, although we observed two types of fibrils by TEM and SEM in the case of condition 7, the otherwise overall homogeneity of ultra-structure seemed to be the most favorable for future development.

The effectiveness of phosphate anion in changing the rate of fibrillogenesis has been reported for long [43],[15]. Using pepsin-treated collagen (4 mg/mL) Hayashi has shown that, at pH 6.8 with equal concentrations of monovalent and divalent ions, the time length to reach opacity was at a peak for a concentration around 150 mM phosphate. In our experiments, the samples in phosphate buffer 1 were the slowest to opacify taking up to 1 hour to generate stable and easy to handle threads. Roughly, the highest was the ionic strength the fastest the opacification occurred. Compared to buffer conditions 2 to 6, the relatively reduced swelling in buffer 1, in spite of a lower ionic strength, is in agreement with the model proposed by Mertz et al. In this model, the divalent phosphate ions

form salt bridges between positively charged residues on collagen molecules, especially at low ionic strength below the physiological values[26]. The opacification occurred more rapidly in buffer 8 with higher ionic strength creating a dense and continuous shell of a network of thin fibrils on the periphery of the thread. The intermediate situation is observed in the case of buffer 7, which presents intermediate ionic strength. The fact that the shark-skin structure observed after extrusion in all different buffers seems to be preserved only in buffer 8 may indicate that the fibrillation occurs more rapidly in this condition. In the case of buffers 1 and 7, the larger fibrils observed inside the core reflect molecular reorganization and thus a slower fibrillogenesis process than in buffer 8 which only gives rise to thin fibrils. Still, at pH 6, the precipitation was first retarded from 130 to 230 mM then accelerated again at 310 mM. Our study highlights phenomena at play in dense collagen solutions moreover constrained in cylindrical geometries, and the possible role of kinetics formation in the final structures obtained. Furthermore, the probable specific binding effect of phosphate anions, already observed at low collagen concentrations, seems to be confirmed at higher concentrations in our system.

Uquillas et al. (2011) showed that the variation of PBS concentration, in the buffers used to trigger fibrillogenesis, induced an increase of UTS and Young's modulus for PBS 1X, 5X and 10X compared to PBS 0.5X [39]. This compares pretty well with our results except for PBS10X for which we find a decrease of mechanical properties. They also reported an effect of the time left in the fibrillogenesis buffer on the mechanical properties of their scaffolds: after 12 hours incubation in
PBS 1X, the UTS and Young's modulus were around 0.4-0.5 MPa for both, and reached 1.2 MPa for UTS and 6 MPa for Young's modulus after 96h. Beside, they showed using SAXS that the D-bands characteristic of fibrillary collagen was present after 12 hours incubation in PBS 1X but not in PBS 10X, and that this pattern disappeared after 96h incubation in PBS 1X. In our case, although the thread diameters evolved over four days before reaching stabilization, after two weeks of maturation, the fibrils constituting the threads still showed cross-striations indicating good and stable fibrillogenesis in our three selected conditions (1, 7 and 8).

Concerning the mechanical properties in our systems, the highest Young's modulus (1.66 ± 0.45 MPa) was obtained with condition 7 compared to conditions 1 and 8 (0.30 ± 0.15 MPa and 0.90 ± 0.17 MPa respectively). This value may seem rather low compared to other reported scaffolds. For instance, Cornwell et al. (2007) measured Young's modulus around 4 MPa and UTS around 1.5 MPa [7]. However, these values were obtained on 10 mg/mL collagen threads that were air-dried and then soaked again in PBS[34]. Kato et al. (1989)[18] and Dunn et al. (1993)[11] had already obtained threads made with the same initial collagen concentration and in a similar fibrillogenesis buffer as Cornwell and coworkers. A modulus of 13.2 MPa on threads that had been rinsed in alcohol, air-dried then rewet in PBS was reported[18]. The measured UTS ranged between 16 MPa and 2.4 approximately, with a decrease for larger thread diameters [11] [18], in accordance with Cornwell et al. results [7]. The good mechanical values they report are likely resulting from enhanced fibril alignment and density due to

thread diameter contraction induced by the dehydration processes. However, the authors did not report any ultrastructural study to assess the alignment of the fibers. Using a similar process as Pins et al. (1997)[31] but starting with a 5 mg/mL collagen solution extracted in an initial polyethylene glycol solution (20% PEG MW 8000, in 94 mM sodium phosphate dibasic and 24 mM sodium phosphate monobasic), followed by dehydration in 70% isopropanol and drying, Cavallaro et al. (1994)[5] measured UTS around 1.2 MPa, with well aligned collagen fibers. By comparison, in our present work, the use of concentrated collagen solutions enables reaching moduli in the same MPa range in the absence of any crosslinking, air-drying and dehydration. Thus, our process still leaves possibilities to further reinforce the mechanical properties by applying further treatments as air-drying or chemical cross-linking or further bottom-up post-processing like braiding Finally, the stress-strain curve shapes and the different profiles adopted by the three types of threads after rupture, also indicate a better behavior of thread synthetized in condition 7. Indeed, in case of buffer 1, the thread breaks with a sharp fracture area, whereas in case of buffer 8, the crack propagates through a sliding mode, with fiber-like tips being visible at both ends after breakage. This mode of ductile rupture corresponds to the plateau observed in the stress-strain curve with constant stress over a large range of strain. In the case of condition 7, although no clear behavior arises from the stress-strain curve, the tip profile with several sharp breaks observed at different positions along the longitudinal axis, would be in favor with a multistep rupture. Thread in condition 7 thus present more
favorable behaviors, in between threads syn-
thetized in condition 1 and 8.

**5 Conclusion**

The home-made set up build to extrude dense
and highly viscous collagen solutions enabled
us to extrude regular threads over meters for
collagen at 60 mg/mL. Through the analysis
of different buffers, we have selected a condi-
tion (PBS 5X) in which the diameter of the
extruded solution was preserved during fibril-
logenesis and storage. "67 nm" cross-striated
collagen fibrils with two diameters (80 and 20

nm) were generated, the majority of which
were aligned parallel to the thread axis. How-
ever, some fibrils appear misaligned to this
axis and up to a 90° deviation. This some-
how disorganization may contribute to the
mechanical strength of the threads by block-
ing the preferential sliding direction along the
thread axis. The threads synthesized in this
condition present a Young modulus of 1.66 ±
0.45 MPa and an ultimate tensile strength of
0.33 ± 0.08 MPa without the use of any cross-
linkers or any drying step. We suggest those
threads to be good candidate for tendon re-
pair trials.
<table><caption>Table 2: S-I: Diameters and striations of fibrillar entities in threads synthesized in condition 1, 7 and 8.</caption><caption>TABLE 2</caption><thead><tr><th>CONDITION</th><th colspan="3">1</th><th colspan="2">7</th><th colspan="2">8</th></tr><tr><th>Scheme of thread</th><th colspan="3">Outskirt Extern Interne</th><th colspan="2">Extern Interne</th><th colspan="2">Extern Interne</th></tr><tr><th></th><th>Outskirt</th><th>Extern</th><th>Intern</th><th>Extern</th><th>Intern</th><th>Extern</th><th>Intern</th></tr></thead><tbody><tr><th>Diameters SEM (nm)</th><td>560/150</td><td>70</td><td></td><td>100</td><td></td><td>150</td><td></td></tr><tr><th>Diameters TEM (nm)</th><td>500</td><td>250</td><td>40</td><td>80/20</td><td>80/20</td><td>25</td><td>25</td></tr><tr><th>Striations SEM (nm)</th><td colspan="3">180 at 3kV<br>67 at 8kV</td><td colspan="2">67 at 8kV</td><td colspan="2">190 at 3kV</td></tr><tr><th>Striations TEM (nm)</th><td colspan="3">67</td><td colspan="2">67</td><td colspan="2"></td></tr></tbody></table>

## Supplementary information

<figure><img src="image_118.png" /><figcaption>Figure 7: Extrusion of collagen solutions through a 650 µm inner diameter needle in the first 6 buffer conditions as listed in Table 1. Bottom: higher magnification. The threads present a shark-skin like appearance and their diameter is equivalent to that of the inner diameter of the extruding needle.</figcaption></figure>
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## 5.4 Pure collagen vs collagen/alginate mixtures threads study

From the previous study, the threads with the best results are extruded in PBS5X. This condition is then our gold standard and we can explore other parameters as the collagen concentration and the ageing time. In this study, collagen concentration varies from 15 to 60 mg/mL. The threads are analyzed from characteristics at extrusion to their mechanical properties. Then, we produce hybrid threads composed of alginate and collagen and do the analysis to compare the two systems.

**Extrusion parameters choice:** in the study on the effect of buffers on fibrillogenesis, the output velocity was: $v̄=50$ mm/s. This speed was chosen from alginate extrusion experiments. Indeed, as several unintended instabilities were observed in the presence of cross-linking ions at low throughput, we decided to transpose those results to the collagen and work at high throughput to avoid threads bulk distorsions. We thus produced regular collagen threads and studied the influence of the ionic strength. In the current study, it was quite difficult to extrude the lowest concentration at this velocity range. Indeed, extrusion of such solution resulted in pieces of irregular "threads" of few millimeters. Therefore, the output velocity was reduced to $v̄= 3.4$ mm/s for which all concentrations lead to useable threads. Of course, we checked that this decrease did not induce huge morphology or structural changes on collagen threads at 60 mg/mL before beginning the concentration impact study. As collagen solutions at 60 mg/mL extruded in PBS 5X at 50mm/s and 3.4mm/s presented the same stability and structural morphologies (SEM, TEM, SHG), we use the new velocity $v̄= 3.4$ mm/s and kept the same previous needle diameter (D=390µm) for the whole following works.

### 5.4.1 Pure collagen threads of different concentrations

Three different concentrations were tested. In this part, A, B and C designate collagen threads at 15, 30 and 60 mg/mL respectively.

#### 5.4.1.1 Extrusion and maturation

The collagen threads were observed from the beginning of their formation and during their ageing in the fibrillogenesis buffer. The results obtained with different techniques are presented in Figure 5.11.

During the extrusion, the collagen solution flow is steady and homogeneous (Fig. 5.11 Top). However, at 30 (B) and 60 (C) mg/mL, a surface irregularity is observed. By optical view (Fig. 5.11 middle), it is clear that the threads have different diameters: A > B > C. This will be quantified in the following part. They exhibit regular shapes even if some roughness appears for B and C. The thread at 15 mg/mL (A) is quite transparent and smooth. At 30 mg/mL, the thread is transparent and presents a periodic roughness of almost 25 µm depth. At 60 mg/mL, the thread is more opaque and its surface is textured by a pattern of 10 µm height. This surface pattern is due to the sharkskin instability already observed at 60mg/mL in the previous study (5.3 p 157). It is not observed at 15 mg/mL as to reach the shear-shinning region of this condition, the extrusion velocity must be considerably increased. In fact, the imposed shear rate is about $4.5 s^{-1}$ in our experiments. If we take a look on the viscosity plot as a function of the shear rates given in Gobeaux et al. work,
([8] 3.6 p 94) at ·γ = 4.5 s⁻¹, collagen solutions at 30 and 60 mg/mL are in the strong shear-thinning region whereas the solution at 15 mg/mL is almost at the end of the newtonian region.

During observation by polarized light microscopy (Fig. 5.11 bottom), threads are positioned at nearly 45° to both polarizers. They appear homogeneously bright along the thread axis. The birefringence signal is higher while increasing collagen concentration. For the lowest one, the thread is almost not visible (A). For threads B and C, a central part is observed, although this might be due to the cylindrical geometry. Their inner structure seems to be aligned along the thread axis.

The threads A, B and C exhibit both regular shapes even if some roughness appears for B and C. Their inner structure seem to be aligned along the thread axis for B and C.

<figure><img src="image_119.png" /><figcaption>Figure 5.11: Collagen threads observed by different techniques at different concentrations. On the left, collagen thread at 15mg/mL (A). During the extrusion, the collagen flow is homogeneous and the thread seems to be smooth. It is confirmed by optical microscopy. At 30 and 60mg/mL (B and C), a pattern at the thread surface is observed during the extrusion. It is also visible by optical view where a regular roughness appears. Bottom: Polarized light microscopy technique: birefringence signal is visible for 30 and 60 mg/mL. The inner collagen organization seem to be aligned along the thread axis. Scale bars are 100µm</figcaption></figure>
5.4.1.1.1 Diameter stability

The collagen threads diameter was measured over 2 weeks by optical microscopy to assess their ageing evolution in the fibrillogenesis buffer (PBS 5X) chosen from the previous study. The results are presented in Figure 5.12 showing the plot of the ratio D/D0.

<figure><img src="image_120.png" /><figcaption>Figure 5.12: Evolution of collagen threads diameter over time. The ratio of the measured thread diameter D and the inner diameter of the needle was plotted as a function of the time (day). Dash-lines represent the diameter evolution trend. Optical microscopy images of collagen threads after 14 days of ageing in PBS 5X, at A) 15mg/mL; B) 30mg/mL and C) 60mg/mL</figcaption></figure>

For collagen threads A and B, the ratio D/D0 increases almost from 1 to 1.3 during the first two days. From the third day, D/D0 remains constant in both cases. For collagen thread C, the diameter ratio is equal to 1 and remains constant even after 2 weeks.

5.4.1.1.2 Surface analysis

The threads topography was observed by scanning electron microscopy (SEM) over time. The results are shown in Figure 5.13.

At day 1, SEM shows that the thread A surface is quite smooth while B exhibits more pleated patterns aligned in the thread axis. For C, we observe the same pattern, which now appears constituted by "ball of wool" made of collagen fibrils (Fig. 5.13 C). After two weeks in PBS 5X, B and C surface morphology do not differ much from day 1. For the lowest concentration, fibrillar structures are revealed. Those are smaller than those constituting the "ball of wool" of thread C.

These observations are in good agreement with the results obtained with optical techniques (Fig. 5.11 p 181). Thread A is quite smooth compared to B and C. The sharkskin surface instability seems to lead to those "ball of wool" of collagen fibrils for thread C. This perturbation is lower for threads B and A resulting only in more or less pronounced pleated. The fibrillar structures are larger for C than A that suggests a different fibrillogenesis kinetic.
<figure><img src="image_121.png" /><figcaption>Figure 5.13: SEM images of threads surface at A) 15mg/mL, B) 30mg/mL, C) 60mg/mL at day 1 and 14. A appears quite smooth compared to B and C where a more pronounced roughness is observed.</figcaption></figure>

### 5.4.1.2 Structural characterization

#### 5.4.1.2.1 Fibrils characterization

Figure 5.14 shows the distribution of the fibrils as well as their morphology as seen by transmission electron microscopy (TEM) after 2 weeks in PBS 5X. Figure 5.14 A, B and C, are cross-sections of threads respectively made of collagen at 15, 30, 60 mg/mL at low magnification on the top part and at high magnification on the lower part. The pictures show view of threads center areas.

In all threads, fibrils presenting the 67 nm cross-striation are observed and their diameter is about 40-60 nm. Their length is greater for A and C and can reach almost 1 µm and seem smaller in the thread B. For thread A (low magnification), the fibrils population is smaller per unit area than threads B and C. Indeed, as A has less collagen material (low concentration) and swells during ageing, collagen fibrils are more disconnected from each other. On top images, for threads A and B, most of the fibrils are cut perpendicular to their axis, giving the dotted features and fewer appear cut along their longitudinal axis. For thread C, both populations are observed.

The concentration effect is observed by TEM on threads centers, where the fibrils population appears smaller while decreasing the concentration. Fibrils are well aligned in the thread axis, which suggests that extrusion process tends to align them along the longitudinal thread axis. For the highest concentration, a transverse fibrils population is also observed.
<figure><img src="image_122.png" /><figcaption>Figure 5.14: Ultrastructure of threads after 2 weeks of maturation in PBS 5X as seen by TEM (transverse sections). Figure A, B, C correspond respectively to collagen threads at concentrations 15, 30 and 60mg/mL. Top: view at low magnification of a center threads region. Bottom: view at higher magnification from near region. Collagen fibrils appear cut in all directions from longitudinal to transverse, in which case they appear as dots.</figcaption></figure>

#### 5.4.1.2.2 Collagen fibrils organization and orientation

Second Harmonic Generation microscopy (SHG) and Polarization-resolved SHG (P-SHG) were performed on the collagen threads at different concentrations as a function of time. We first present the results obtained after 2 weeks of ageing in PBS 5X (Fig. 5.15) and detail those obtained at previous times if differences are revealed.

At lower magnification, SHG intensity is homogeneous in every direction for thread A. In the case of thread B, a texture (black striations) is observed within a distance of approximately 50µm from the surface. This shell is radially organized as it is observed on both "sides". The thread center is homogeneous. The same structure is observed for thread C, with more pronounced striking features, and a size of about 100µm. All those observations suggest that more concentrated collagen threads have a shell-core structure and that the one with the lowest concentration is homogeneous. This is in good agreement with polarized light microscopy previous observations, for which some doubts about the origin of this core structure arose.

P-SHG enables to calculate the anisotropy paramter ρ and the orientation φ of collagen in the image plane as previously explained, and thus to assess the micrometer scale disorder in the collagen fibril distribution. In the anisotropy "map" (Fig. 5.15), ρ is constant over the thread because the thread appears as a uniform medium for all conditions. The values of the anisotropy parameter is almost the same for the three concentrations and is
<figure><img src="image_123.png" /><figcaption>Figure 5.15: TOP: SHG imaging of collagen threads at a.15 mg/mL, b. 30 mg/mL and c. 60mg/mL in the center transverse thread planes. The threads are placed vertically and their borders are visible on each side (left and right) of the images. Scale bar: 100µm. P-SHG imaging a.1., b.1., c.1: zoomed-in images taken in the yellow squared areas respectively in a., b., c. BOTTOM: images of P-SHG analysis of the anisotropy parameter (a.2, b.2, c.2) and the orientation φ of collagen in the image plane (a.3, b.3, c.3) Scale bar: 10µm.</figcaption></figure>

equal in the superficial area to 1.34, 1.56 and 1.45 for respectively A, B and C. The same experiments were performed on native rat-tail tendons by Gusachenko et al. [9]. They found anisotropy parameters of about 1.40 ± 0.03 on the surface. Our results are almost in the same range, which may indicate that the collagen fibrils organization in the threads is close to those of tendons. Moreover, in our case, we observe that this parameter decreases, while going from the thread surface to its center (1.27, 1.52 and 1.40 are the ρ values at the center for A, B and C). This decrease is presumably an artefact due to diat-
tenuation.

The collagen fibrils orientation was also obtained by P-SHG measurements on regions of interest. Figure 5.15 (bottom) displays the resulting color maps obtained in the threads center transverse planes in the region close to the surface ("left side" in this example), as shown by the yellow squares on the top images. For the thread A, the fibrils orientation is the same in all directions, which means that it is isotropic. For threads B and C, φ varies abruptly from the textured shell to the core: from -60° to +50° for B and -70° to +70° for C. The same measurements were done at the thread "center-center" and at the other surface side for B. The results are shown in Figure 5.16.

<figure><img src="image_124.png" /><figcaption>Figure 5.16: SHG imaging and P-SHG areas (colored) of surfaces and middle regions at the threads center transverse sections for B: 30mg/mL and C: 60mg/mL. Schemes of collagen fibril orientation φ are given to help to visualize the radial organization. Scale bar: 50µm.</figcaption></figure>

Sharp angles variations are observed from the shell-core interfaces to the thread center. The angles changes are lower for thread B than C. The collagen fibrils orientation has radial symmetry. The resulting angles maps look like "ear of corn" patterns at interfaces. The mean orientation follows the thread longitudinal axis. Those complex organization may come from the sharkskin instability. Indeed, only threads B and C exhibit this pattern during extrusion. It means that strong strains are applied at the singular die exit and may lead to this collagen fibrils orientation difference between the shell (high strains) and the core. Moreover, as collagen fibrillogenesis at the thread surface occurs more rapidly for higher concentration, the matrix is "frozen" in the high strains state and can not relax. It might explain the φ angle amplitude, which is higher for C. Other parameters as the swelling of B can also be considered to understand those observations.

**Remark 1: Ageing time influence**

At day 1, collagen threads A and C have the same structure as presented at day 14. Only some collagen fibrillar structures at the surface were larger at day 14, these results being similar with those as observed by SEM (5.13 p 183). However, at day 1, the shell-core structure was not observed for thread B as shown on figure 5.17.

The sharp striated patterns of the shell are not visible by SHG. Surprisingly, while looking at the P-SHG measurements of φ angle at day 1, the interface between shell and core, and visible at day 14, matches an angle orientation difference (blue and red, the interface being represented by a yellow dashed line). It indicates that collagen is already oriented and during fibrillogenesis/ageing phase, this orientation becomes visible at a higher scale.
<figure><img src="image_125.png" /></figure>

Figure 5.17: SHG and P-SHG imaging of collagen thread at 30mg/mL (B) at day 1. It is already possible to see the different orientations in the core/shell regions even if it is not visible on SHG image as in the case of 60mg/mL (C). Yellow dashed line is an eye guideline to show the core/shell interface. Scale bar: 10µm.

**Remark 2: "Ball of wool" and 3D imaging**

By SHG technique, it is also possible to do 3D imaging of the threads. This kind of plane-by-plane acquisition (each 5µm) takes typically 1 to 2 hours. We give here an example of 3D imaging of a collagen thread at 60 mg/mL after 2 weeks in PBS 5X. At the outer surface, the famous "ball of wool" of collagen fibrils is clearly seen (Fig. 5.18, top left 1) as already observed by SEM (5.13 p 183).

<figure><img src="image_126.png" /></figure>

Figure 5.18: SHG images of collagen threads at 60 mg/mL after 2 weeks in PBS 5X. Images 1 and 2 correspond to the depths indicated by the arrowheads. The cross section was obtained by 3D reconstruction.

In this example, it is noticed that several concentric "layers" are revealed at the thread center depth (Fig. 5.18, top right 2). With the 3D reconstruction, we can observe the cross-section thread organization (Fig. 5.18). The main shell-core interface is visible and there, the SHG signal seems to be lower. If we come back to the scheme of thread C (Fig.
5.16 p 186, right), blue arrows were put on this area of about 10 µm, corresponding to the fact that φ is equal to 40-50 °between the shell and the core. From these two facts, we assumed that this small area may be a depletion area or a region where collagen molecules are mainly "antiparallel" leading to decreased SHG signal. Further TEM investigations must be done to assess the fibrils size and repartition between the surface, the shell and the core of the threads.

**Remark 3: discarded condition**

Unfortunately, as collagen thread at 15 mg/mL was difficult to handle, we decided to remove it for the mechanical characterization as it is obviously too weak for tendon repair applications.

5.4.2 Collagen/alginate threads

Hybrid collagen-alginate threads were made in order to combine the two systems and take benefit from both of them. In fact, in chapter 1 we reported the advantages of alginate (1.2.4 p 27), which can be degraded and thus created some porosity, and which also has non adhesive properties, both interesting characteristics in the view of tendon repair. On the other hand, as we already have presented it, collagen is the main component of connective tissues. The extrusions were performed in chloride calcium 1M to induce the gelation of alginate. The pH of this bath solution is about 6-7.

5.4.2.1 Ratio choice

This kind of mixtures is not easy to achieve due to proteins and polysaccharides in-
teractions, phase separation and precipitation that may occur leading to heterogeneous
systems. In order to ease the task, we produce collagen-alginate threads with effective
total material quantity similar to pure collagen of 30 milligrams per milliliter and with ex-
actly same extrusion parameters. Several ratios (2.2.3.2 p 84) are tested and the resulting
flow extrusion and threads are presented in Figure 5.19.

During extrusion, at ratios 1:1 (coll:alg) and 2:1, tortuous opaque domains are visible while the rest of the mixture solution is transparent. At ratio 5:1, the whole extrudate seems to be more homogeneous. Regarding the thread diameter uniformity, ratio 2:1 is clearly not straight and material is not well distributed inside the thread. By polarized light microscopy, the birefringence differences are more obvious. At ratio 1:1, very bright and dark regions alternates sharply, which means that some regions are rich and other poor in collagen. The material distribution is not homogeneous. For the thread at ratio 2:1, bright domain is homogeneous over a larger period. At ratio 5:1, birefringence signal is less intense but more homogeneous over the whole thread. Finally, while changing the extrusion buffer (CaCl₂ 1M) for few minutes by PBS 1X, threads at ratios 1:1 and 2:1 begin to disintegrate and ratio 5:1 remain quite stable.

According to these results, we assume that the best condition for the further study is the
collagen:alginate ratio 5:1.
<figure><img src="image_127.png" /><figcaption>Figure 5.19: Collagen:alginate threads at ratios: 1:1, 2:1, 5:1, observed by different techniques. On the left, real-time observations during the extrusion. Right: bright field views. Bottom: Polarized light microscopy images between two crossed polarizers at 45°. Scale bar are 250µm.</figcaption></figure>

### 5.4.2.2 Structural characterization

Some plugs appear once in a while and disturb the solution flow extrusion. Therefore, we use needle diameter of **D=460µm** and increase the piston velocity in order to have output velocity of $\bar{v}= 3.4 mm/s$ and thus the same shear rate as for collagen threads at 30 mg/mL.

#### Surface analysis

The topography of collagen-alginate (5:1) thread and pure ALG4 thread as control were observed by SEM. In Figure 5.20, ALG4 thread surface appears smooth. For the mixed thread, the surface also appears smooth, but some small aggregated structures are observed on it. A higher magnification, we can see the difference between the smooth (right) surface and those aggregates. As those structures are not observed for pure alginate, it is assumed that they might be made of collagen.
<figure><img src="image_128.png" /><figcaption>Figure 5.20: SEM imaging of pure ALG4 (control) and collagen:alginate (5:1) threads. Both thread surface are smooth, but some aggregates are observed on the hybrid thread.</figcaption></figure>

Microstructural organization

As for pure collagen at 30 mg/mL, microstructural organization was assessed by TEM and SHG techniques. In Figure 5.21, TEM images of two hybrid thread areas are shown. On the **left**, straight and abrupt transitions between two populations are observed. A population (1) composed of a network of small strands (length about 60-70nm) is enclosed between two similar areas (2), which are very concentrated and compact. At high magnification, this second population seems to be composed of thin "hair"-like structures (width of 5-6 nm) that are quite aligned with the interface axis. This phase (2) seems also to be composed of several parallel layers (bottom right corner).

TEM micrographs of a second representative area of the sample are shown on the figure 5.21 right. The two populations (1) and (2) described in the first area are found again. Here, the interface is not abrupt as in the other region. Moreover, a third population (3) seems to be located at the interface between the void-strands and the compact "hairs" region. Several concentrated tortuous bundles are observed between these populations. They have diameters about 70nm and length ranging from 500nm to 1µm. At high magnification, few of those larger shapes seem to present striations (red arrow) akin D-spacing of collagen fibrils.
<figure><img src="image_129.png" /><figcaption>Figure 5.21: TEM longitudinal sections images of two characteristic areas of collagen:alginate (5:1) thread in CaCl₂ 1M.</figcaption></figure>

Based on structural studies by other authors, we attempt to identify those different populations. The figure 5.22 shows TEM images of a cross-linked alginate gel at 2 % (w|w) [11] and a fibrillated collagen gel at 100-150 mg/mL (pH 13) [12].

<figure><img src="image_130.png" /><figcaption>Figure 5.22: TEM micrographs obtained from studies of a cross-linked alginate gel at 2 % (w|w) [11] and a fibrillated collagen gel at 100-150 mg/mL (pH 13) [12]. Scale bar: 200 nm.</figcaption></figure>
a) Schuster et al. [11] observed alginate gel networks. This network consists in polymer strands (dark) and voids (grey space), which are similar to the population (1) observed in our collagen:alginate thread.

b) Gobeaux et al. [12] obtained disordered fibrillar aggregates at 100-150 mg/mL at pH 13, that do not exhibit the typical cross-striation for collagen. They explained that these aggregates are similar to those obtained for lower pH of 6.2-12.2. Moreover, they measured the width of the nanofibrils located in the aggregates and found a width of 6 *nm*. These observations are in good agreement with our results for the population (2). In our case, the initial concentration is not comparable to gel at 100 mg/mL. However, as we observe phase separation, a local collagen material concentration may occur that would lead to structures similar to those of highest concentration as seen in Figure 5.22.

## SHG observations

To further investigate the microstructural collagen of hybrid thread, SHG observations were performed by Adèle Mauroux (PhD student under the supervision of Laurent Muller). In Figure 5.23, section view and their projections are given. There is a diffuse SHG signal in almost the whole thread. However, some areas are more intense. It means that fibrillar collagen is located in certain regions. In fact, we observe the same structures as previously observed by TEM: tortuous collagen phase (3) pointed by the red arrow. These "steamy" bright areas are assumed to be collagen fibrillated with a lower SHG intensity than in pure collagen threads. Those structures are located almost at the thread surface. The less intense SHG areas may be those of nanofibrils corresponding to the population (2). Finally, the darker regions may consist in alginate population (1). They seem to be preferentially located deeper inside the threads than population 3.

<figure><img src="image_131.png" /><figcaption>Figure 5.23: SHG images of collagen:alginate (5:1) thread in CaCl₂ section and the respective projections as given by the yellow lines. Scale bar: 100µm. SHG signal is observed in almost the whole thread but is more or less intense in certain areas. Red arrow points to a tortuous structure.</figcaption></figure>
### 5.4.3 Mechanical properties comparison

Stress-strain plots representative of each condition are presented in Figure 5.24. For collagen thread at 30mg/mL, a linear evolution before the rupture is observed. For the coll:alg thread, the material failure seems to be prolonged in term of strain compared to the pure collagen one. However, a better loading extent is observed for collagen 30mg/mL. Finally, at 60mg/mL, the stress roughly increases until reaching a plateau corresponding to a plastic behavior, whereas collagen 30 and its hybrid equivalent rather exhibit elastic behavior.

<figure><img src="image_132.png" /><figcaption>Figure 5.24: Tensile testing until rupture for collagen threads at 30 and 60mg/mL and coll:alg (5:1) thread in wet conditions and at room temperature. The characteristic stress-strain curves are given for the three conditions.</figcaption></figure>

<table><caption>Table 5.1: Tensile testing results: Young modulus E, Ultimate Tensile Strength UTS for collagen at 30 and 60 mg/mL and collagen:alginate at 5:1 ratio threads. N is the number of experiments performed.</caption><thead><tr><td>Condition</td><td>E (MPa)</td><td>UTS (MPa)</td><td>N</td></tr></thead><tbody><tr><td>Coll 30 mg/mL</td><td>0.20 ± 0.06</td><td>0.081 ± 0.016</td><td>10</td></tr><tr><td>Coll:Alg (5:1)</td><td>0.84 ± 0.17</td><td>0.11 ± 0.06</td><td>5</td></tr><tr><td>Coll 60 mg/mL</td><td>1.66 ± 0.45</td><td>0.31 ± 0.08</td><td>33</td></tr></tbody></table>

The summary results of tensile testing are shown in table 5.1. At 30 mg/mL, E of about 0.20 MPa is reached. This value is almost 8 times larger at 60mg/mL. Collagen threads at 30 mg/mL have UTS 4times smaller than at 60 mg/mL. Here, we clearly notice the collagen concentration effect: the more material in the thread, the best mechanical properties. While comparing pure collagen and hybrid collagen:alginate, we observe that the addition of alginate improve the mechanical properties of the threads, for which E (4-fold) and UTS (1.4-fold) are higher. But, the mixture thread can support smaller strains.
## 5.5 Preliminary study to cell culture

Before seeding the cells on threads, it is necessary to know if the collagen-based threads are stable in the cell culture medium. In fact, this medium is composed of 90% Dulbecco's Modified Eagle's Medium (DMEM: mixture of salts, vitamins etc), 10% Foetal Bovine Serum (FBS) and a small amount of antibiotics. The ionic strength of this medium is about 0.13M [13] and is set to be at physiological pH (7.4).

PBS 5X and CaCl₂ 1M have respective ionic strengths of 0.825 and 1M. The ionic strength is thus decreased while putting threads in culture with cells and this may lead to a structural change. Collagen based threads were put in DMEM in an incubated chamber maintained at 37°C and 5% CO₂ for 3 weeks. In this part, we observe the influence of complete cell culture medium, abbreviated DMEM, on these threads.

### 5.5.1 General morphology

#### 5.5.1.1 Optical observations

By optical microscopy, we observe the threads morphology evolution over time. The measurements are presented in the next section (Fig. 5.27 p 195). The pure collagen thread overall aspect is the same at day 0 as at day 14 (Fig. 5.25). No difference is observed by optical microscopy and we come to the same conclusion for collagen 60mg/mL.

On the contrary, hybrid thread morphology evolves over time. First, it clearly exhibits a larger diameter after 2 weeks in DMEM. As a consequence, the thread becomes more transparent as it is less concentrated per unit area. It is assumed that the ionic strength decreases from 1M to 0.13M leading to the swelling of the thread.

<figure><img src="image_133.png" /><figcaption>Figure 5.25: Bright field microscopy images of collagen 30 mg/mL and coll:alg (5:1) threads incubated during 0 and 14 days in DMEM at 37°C. Scale bar: 500µm</figcaption></figure>

#### 5.5.1.2 Topography analysis

The topography of pure collagen threads was not changed while incubating them in DMEM, on the contrary to the collagen:alginate mixture threads (Fig. 5.26). By comparison with the threads kept in CaCl₂ 1M (Fig. 5.20 p 190), the topography is rough and fibers cover the entire thread surface. At higher magnifications, those fibers appear quite thin
(< 1µm width) and do not have clear orientation. Finally, no typical fibrillar striations are seen.

<figure><img src="image_134.png" /><figcaption>Figure 5.26: SEM images of collagen:alginate (5:1) thread surface after 3 weeks in DMEM at 37°C. Thin fibers cover the entire thread surface and have no clear orientation.</figcaption></figure>

5.5.2 Diameter stability

The threads diameter D were measured on optical microscopy images before and dur-
ing the incubation in DMEM at 37°C. The figure 5.27 displays the results of these measure-
ments.

<figure><img src="image_135.png" /><figcaption>Figure 5.27: Evolution of collagen based thread diameters upon maturation for the different buffers: 1) PBS 5X for collagen 30 mg/mL, CaCl₂ 1M for coll:alg (5:1) and 2) cell culture medium DMEM at 37°C. The diameters of the threads have been measured regularly along the two phases.</figcaption></figure>

In the phase 1, threads remain in their extrusion solution. As previously presented, col-
lagen thread at 30 mg/mL swells in PBS 5X and its diameter is stable after 3 days. On the
contrary, hybrid thread diameter decrease by 10%, which indicates a shrinkage caused by
the high ionic strength of the extrusion solution.

At "day 0", both types of threads are put in DMEM at 37°C. In the phase 2, the ratio D/D0 decreases from 1.22 to 1.12 for pure collagen thread 30mg/mL over 3 days and then is stable. In the case of collagen:alginate mixture, in only few hours, the diameter ratio increases from 0.88 to 1.24, which is equivalent to a swelling of 36%. Then, the ratio slightly decreases and reaches 1.18 value closer to that of pure collagen. The ionic strength difference between CaCl₂ 1M and DMEM is assumed to be responsible for the strong and fast swelling of the collagen:alginate (5:1) thread. In the case of pure collagen thread, the small diameter decrease is not intuitive as one would have thought that it would also swell. The temperature is higher in the phase 2 which might also impact the thread stability.

**5.5.3 Structural characterization**

**5.5.3.1 Pure collagen threads**

Collagen threads at 30 and 60 mg/mL were observed by TEM while remaining in DMEM during 2 weeks. The following figure 5.28 displays the images comparison between the microstructure in PBS 5X and in DMEM.

<figure><img src="image_136.png" /><figcaption>Figure 5.28: TEM cross-sections images of collagen threads at 30 mg/mL (B) and 60 mg/mL (C) either matured in PBS5X or incubated in DMEM at 37°C during 2 weeks.</figcaption></figure>
For both collagen concentrations, cross-sections are very similar between the two buffers. For thread B, most of the collagen fibrils are aligned along the thread axis (dots) and very few fibrils are shifted in both buffers. The fibrils seem to be larger in DMEM than in PBS 5X. Moreover, at low magnification, we observe that there are more fibrils per unit area in DMEM. This is probably due to the small diameter decrease leading to the local collagen concentration. At 60mg/mL, the two conditions give very similar ultrastructure and no effect of the buffer change is revealed. For pure collagen threads, the passage from fibrillogenesis buffer to the cell culture medium does not change the thread structure, which remains quite stable.

### 5.5.3.2 Collagen:alginate threads

TEM observations were also performed on collagen:alginate threads incubated in DMEM during 3 weeks as shown in Figure 5.29:

<figure><img src="image_137.png" /><figcaption>Figure 5.29: TEM sections along the thread axis (left //) and perpendicular (right ⊥) of collagen:alginate (5:1) threads incubated in DMEM at 37°C during 3 weeks.</figcaption></figure>

Longitudinal (left) and perpendicular (right) threads sections images are presented in order to assess the main orientation of structures in the thread. At low magnification, fibrils
with 67 nm period striations are clearly observed and highly contrasted. The three differ-
ent populations are still present: 1) alginate rich phase, 2) nanofibrils collagen phase and
3) alginate and collagen fibrils coexistence region at the interface between 1 and 2. After
3 weeks in DMEM, there are more collagen fibrils and they are larger than those found for
CaCl₂ 1M (Fig. 5.21 p 191), resulting in an overlap with phase 1. Regarding the phase 2,
the main orientation of the nanofibrils bundles might be along the thread axis. Indeed,
for perpendicular sections, those bundles appear as huge areas of dots (red arrow), which
may correspond to transverse cut of collagen nanofibrils. For parallel sections, nanofib-
rils bundles are observed along their whole length. Finally, these bundles are separated by
voids areas, which may be here again alginate phase with typical polymer chains (higher
magnification).

In cell culture medium, fibrillogenesis of collagen molecules clearly continue to occur
compare to CaCl₂ 1M condition and leads to large striated fibrils formation in the "cosol-
uble" phase.

**Remark:** SHG observations were done on the hybrid threads (5:1) after 3 weeks of incu-
bation in DMEM at 37°C (Fig. 5.30).

<figure><img src="image_138.png" /><figcaption>Figure 5.30: SHG image of collagen:alginate (5:1) thread section incubated 3 weeks in DMEM at 37°C and the respective projections as shown by the yellow lines. Scale bar: 100µm.</figcaption></figure>

Looking overall, SHG signal intensity is high and well-defined structures are visible, in
contrast with hybrid thread left in CaCl₂ 1M (Fig. 5.23 p 192). Dark areas are still observed
and must correspond to alginate rich phase. On the transverse view (bottom), SHG sig-
nal appears more intense on the thread surface. This was already noticed for this thread
condition in CaCl₂ 1M. On the longitudinal view (right projection), we observe also the
signal enhancement over 50µm from the surface. Moreover, some light and darker layers
are alternating in this outlying region. It reminds the parallel organization seen by TEM,
where supposed alginate phase and collagen one are stacked on top of each other (Fig.
5.29 p 197, bottom left). Finally, straight structures parallel to the thread axis are pointed out by red arrows, suggesting an alignment in such area of collagen molecules in a overall more disordered and "tortuous" matrix.

5.5.3.3 Collagen stability: thermal analysis

The impact of cell culture medium on the collagen stability was investigated by DSC measurements on wet threads (pH 7.4) at a heating rate of 5°C/min. As shown in figure 5.31, pure collagen sample (red) has a denaturation temperature of 54°C, corresponding to the breaking of protein triple helices. This is in good agreement with values found in previous studies [14]. For pure alginate thread at 4 % (black), a non significant peak at 65°C due to thread dehydration is observed. While mixing collagen and alginate (green), a peak at 41 °C appears.

<figure><img src="image_139.png" /><figcaption>Figure 5.31: DSC measurements: denaturation of collagen 30mg/mL, coll:alg (5:1) and ALG4 threads kept either in their extrusion solutions (PBS 5X or CaCl₂ 1M respectively), or in DMEM during 2 weeks. Heating rate 5°C/min. Endotherms are down.</figcaption></figure>

After 3 weeks in DMEM at 37°C, a slight temperature shift from 54 to 52°C is observed
for pure collagen at 30mg/mL. In the case of hybrid thread, denaturation temperature
rises from 41 to 50 °C suggesting the presence of a more stable protein conformation. In
fact, we previously observed by TEM (Fig. 5.29 p 197) that fibrillar collagen population
was evolving in DMEM and presented typical 67 m striations whereas in CaCl₂ 1M rare
striated fibrils were seen (Fig. 5.21 p 191). As the quarter-staggered configuration is more
stable, a higher temperature is needed to denature those structures.
### 5.5.4 Mechanical properties

Tensile testing results performed in DMEM for the three thread conditions to be seeded with cells are presented in table 5.2. According to the results obtained in extrusion buffer (table 5.1 p 193), pure collagen threads exhibit the same mechanical properties. This is consistent with results from microscopic and thermal analysis. For coll:alg (5:1), E and UTS respectively decrease from 0.84 to 0.31 and 0.11 to 0.032. Young Modulus stays higher than pure collagen thread at 30 mg/mL.

<table><caption>Table 5.2: Tensile testing results: Young modulus E, Ultimate Tensile Strength UTS for threads made of collagen at 30 and 60 mg/mL and collagen:alginate at 5:1 ratio threads incubated in DMEM at 37°C during 3 weeks. N is the number of experiments performed.</caption><thead><tr><td>Condition</td><td>E (MPa)</td><td>UTS (MPa)</td><td>N</td></tr></thead><tbody><tr><td>Coll 30 mg/mL</td><td>0.18 ± 0.03</td><td>0.066 ± 0.005</td><td>14</td></tr><tr><td>Coll:Alg (5:1)</td><td>0.31 ± 0.10</td><td>0.032 ± 0.08</td><td>5</td></tr><tr><td>Coll 60 mg/mL</td><td>1.23 ± 0.12</td><td>0.31 ± 0.01</td><td>5</td></tr></tbody></table>

In DMEM, hybrid threads swell resulting in a lower material concentration per unit area. From tensile testing, threads exhibit a more alginate-like behavior in CaCl₂ 1M and a more collagen-like one in DMEM. In fact, calcium ions content is much lower in DMEM leading to a looser alginate network. This may be amplified and correlated with the thread swelling. Regarding collagen evolution, typical striations appear on collagen fibrils as seen by TEM (5.28) similar to the ones in pure collagen threads. We suggest that Young modulus is still higher in DMEM (than for pure collagen threads at 30mg/mL) because of collagen areas not aligned to the thread axis.

### 5.5.5 Discussion

While changing the thread bath condition from its fibrillogenesis buffer to cell culture medium, several structural changes were observed. In the case of pure collagen condition, threads at 30 and 60 mg/mL remain stable and only a slight diameter swelling was observed for the lowest concentration.

However, for mixed threads, several structural modifications occur leading to new properties. In fact, as presented in Chapter 2 (2.2.3.1 p 83), proteins (collagen) and polysaccharides (alginate) could interact in different way according to pH, ionic strength, temperature etc. We attempt here to explain those interactions when varying the bath condition parameters.

1) **Before extrusion**, we observed that collagen:alginate (5:1) mixed solution was opaque and white (Chapter 2, 2.33 p 84), suggesting a first interaction between the two systems. In fact, the mixed solution then in acetic acid (500mM) has a pH of about 3-4. At pH < pI (isoelectric point), collagen is positively charged [15]. Moreover, at pH > pKa (=3.5) of alginate [16], the polysaccharide is deprotonated. In acidic solvent, collagen and alginate interaction is thus attractive and may lead to an associative phase separation ("2-phase" of precipitation) as a white solution is observed.

2) **While extruding** this mixture in CaCl₂ 1M, three main changes occur: (i) pH increase to physiological value of 7.4 and (ii) alginate binding ions addition, (iii) high ionic strength.
With the pH increase, collagen fibrillogenesis begins. Moreover, collagen is still positively charged and alginate completely deprotonated. Indeed, as the ionic strength is high (1M), collagen pI value is shifted to 9 and so, pH = 7.4 < pI = 9, as shown by Freudenberg et al. [17]. Alginate is completely deprotonated and free G-blocks interact then with divalent ions through the gelation process and form a network of "egg-box" structures. At sight of our results (5.29), we obtained a first phase from the alginate and collagen precipitation still interacting; a second one rich in collagen for which fibrillogenesis begins, and a last one rich in gelled alginate.

3) **In cell culture medium**, the pH is the same but the ionic strength is lower than $CaCl_2$ 1M. Here again, three main changes occur compared to previous situation "2": (i) ionic strength decrease, (ii) calcium ions content decrease, (iii) temperature increase. As the ionic strength decreases, pH ≈ pI for collagen [17], which becomes then almost neutral. As the calcium ions content is lower, the disassembly of alginate network begins. At the same time, ionic strength decreases, leading to thread swelling, which reveals the weakening of the network. Regarding the temperature, the conformation of the "free" chains of alginate or collagen molecules may change and turn to the unfolded state [18], increasing the number of reactive sites available for each species. From our results, we observed the growth of fibrils exhibiting the typical D-spacing of 67nm in the mixed phase of collagen and alginate. It is assumed that in culture medium, the alginate network, in which collagen was entrapped, swells and comes undone (5.29, 5.30). This gives enough space for the entrapped collagen molecules to assemble. The collagen rich phase, secondary to the repulsive interaction with alginate rich phase, can not benefit from this network expansion and the nanofibrils areas stay confined.

Thus, alginate may disturb the collagen fibrillogenesis, but when put in cell culture medium, the mixed thread are composed of collagen phases as stable as pure collagen ones and displaying better mechanical properties.

## 5.6 Conclusion

In this work, several physico-chemical parameters were varied in order to select the condition for which thread topography, structural organization and mechanical properties were improved.

First, we showed that the extrusion process aligns collagen inside the pure collagen threads. This effect was lower for collagen:alginate mixed threads.

Then, collagen threads extruded in PBS 5X were revealed to be more stable and present the best Young modulus and Ultimate Tensile Strength compared to those obtained for lower or higher ionic strengths. Then, we have shown that the higher the collagen concentration, the better the threads properties. Finally, the collagen:alginate mixture (5:1) equivalent to collagen at 30 mg/mL exhibited several promising properties and must be improved to reach those obtained for pure collagen at 60 mg/mL.
**5.7 References**

[1] Francesco S Pavone and Paul J Campagnola. *Second harmonic generation imaging*. CRC Press, 2013. 146

[2] A Ustione and DW Piston. A simple introduction to multiphoton microscopy. *Journal of microscopy*, 243(3):221-226, 2011. 146

[3] Mathias Strupler. *Imagerie du collagène par microscopie multiphotonique*. PhD thesis, Ecole Polytechnique X, 2008. 147

[4] Stéphane Bancelin. *Imagerie quantitative du collagène par génération de seconde harmonique*. PhD thesis, Ecole Polytechnique X, 2013. 148

[5] Claire Teulon. *Imagerie quantitative de biopolymères par génération de second harmonic résolue en polarisation*. PhD thesis, Université Paris-Saclay, 2016. 148, 149

[6] DA Kleinman. Nonlinear dielectric polarization in optical media. *Physical Review*, 126(6):1977, 1962. 150

[7] Frédéric Gobeaux. *PHASES DENSES DE COLLAGÈNE DE TYPE I: TRANSITION ISOTROPE/CHOLESTÉRIQUE, FIBRILLOGENÈSE ET MINÉRALISATION*. PhD thesis, Université Pierre et Marie Curie-Paris VI, 2007. 157

[8] Frédéric Gobeaux, Emmanuel Belamie, Gervaise Mosser, Patrick Davidson, and Sophie Asnacios. Power law rheology and strain-induced yielding in acidic solutions of type i-collagen. *Soft Matter*, 6(16):3769-3777, 2010. 181

[9] Ivan Gusachenko, Gaël Latour, and Marie-Claire Schanne-Klein. Polarization-resolved second harmonic microscopy in anisotropic thick tissues. *Optics express*, 18(18):19339-19352, 2010. 185

[10] Ivan Gusachenko, Viet Tran, Yannick Goulam Houssen, Jean-Marc Allain, and Marie-Claire Schanne-Klein. Polarization-resolved second-harmonic generation in tendon upon mechanical stretching. *Biophysical journal*, 102(9):2220-2229, 2012.

[11] Erich Schuster, Johanna Eckardt, Anne-Marie Hermansson, Anette Larsson, Niklas Lorén, Annika Altskär, and Anna Ström. Microstructural, mechanical and mass transport properties of isotropic and capillary alginate gels. *Soft Matter*, 10(2):357-366, 2014. 191, 192

[12] Frédéric Gobeaux, Gervaise Mosser, Anny Anglo, Pierre Panine, Patrick Davidson, M-M Giraud-Guille, and Emmanuel Belamie. Fibrillogenesis in dense collagen solutions: a physicochemical study. *Journal of molecular biology*, 376(5):1509-1522, 2008. 191, 192

[13] Benoît GC Maisonneuve, Denis CD Roux, Peter Thorn, and Justin J Cooper-White. Effects of synthetic biomacromolecule addition on the flow behavior of concentrated mesenchymal cell suspensions. *Biomacromolecules*, 16(1):275-283, 2014. 194

[14] Maria Victoria Tuttolomondo, Juan Manuel Galdopórpora, Lea Trichet, Hugo Voisin, Thibaud Coradin, and Martin Federico Desimone. Dye-collagen interactions. mechanism, kinetic and thermodynamic analysis. *RSC Advances*, 5(71):57395-57405, 2015. 199
[15] John H Highberger. The isoelectric point of collagen. *Journal of the American Chemical Society*, 61(9):2302–2303, 1939. 200

[16] ARNE Haug and O Smidsrød. The effect of divalent metals on the properties of alginate solutions. *Acta Chem. Scand*, 19(2):341–351, 1965. 200

[17] Uwe Freudenberg, Sven H Behrens, Petra B Welzel, Martin Müller, Milauscha Grimmer, Katrin Salchert, Tilman Taeger, Kati Schmidt, Wolfgang Pompe, and Carsten Werner. Electrostatic interactions modulate the conformation of collagen i. *Bio-physical journal*, 92(6):2108–2119, 2007. 201

[18] Amit K Ghosh and Prasun Bandyopadhyay. Polysaccharide-protein interactions and their relevance in food colloids. In *The complex world of polysaccharides*. InTech, 2012. 201
# Chapter 6

# Collagen scaffolds and cell culture

"Even miracles take a little time."

Fairy Godmother in *Cinderella*

## Abstract

In this chapter, the mesenchymal stem cells (C3H10 T1/2) behavior was studied as a function of several parameters. First, we investigated the collagen concentration influence on cell morphologies and ability to colonize the threads. Preliminary in situ hybridization results of the target genes were given in order to assess qualitatively their expressions. Those first observations were promising. As cellular penetration was enhanced on collagen threads at 30 mg/mL, further quantitative analysis was performed enabling the comparison with corresponding collagen-alginate based threads. With this second study we tested the influence of: thread material, cell culture time, and mechanical loading during culture. We used Real time qPCR to quantify the expression of genes of interest (Col1a1, Tnmd and Scx) in order to characterize cell behavior. Characteristic tendon gene markers were expressed, which reflect the differentiation of C3H10 T1/2 into tendon-like cells.

## Contents

<table><tr><td><strong>6.1</strong></td><td><strong>Biological material and techniques</strong></td><td><strong>207</strong></td></tr><tr><td>6.1.1</td><td>Cell culture</td><td>207</td></tr><tr><td>6.1.2</td><td>Cell culture on threads under tension</td><td>207</td></tr><tr><td>6.1.3</td><td>Confocal Fluorescence Microscopy</td><td>209</td></tr><tr><td>6.1.4</td><td>Histology and Hybridization in situ</td><td>210</td></tr><tr><td>6.1.5</td><td>Real time q-PCR</td><td>212</td></tr><tr><td>6.1.5.1</td><td>Principle</td><td>212</td></tr><tr><td>6.1.5.2</td><td>Protocol</td><td>213</td></tr><tr><td><strong>6.2</strong></td><td><strong>Cell colonization study</strong></td><td><strong>217</strong></td></tr><tr><td>6.2.1</td><td>Mechanical properties of threads cultured with cells</td><td>217</td></tr><tr><td>6.2.2</td><td>Cell migration within the threads at different collagen concentrations</td><td>218</td></tr><tr><td>6.2.3</td><td>Preliminary In situ Hybridization results</td><td>221</td></tr><tr><td>6.2.4</td><td>Conclusion</td><td>222</td></tr><tr><td><strong>6.3</strong></td><td><strong>Behavior of stem-cells cultured on collagen based threads</strong></td><td><strong>223</strong></td></tr><tr><td>6.3.1</td><td>Load influence of threads morphologies</td><td>223</td></tr><tr><td>6.3.2</td><td>Cell colonization over culture time and load condition impact analysis</td><td>224</td></tr><tr><td>6.3.3</td><td>Genes of interest expression analysis</td><td>226</td></tr></table>
<table><tr><td>6.3.3.1 Qualitative analysis</td><td>226</td></tr><tr><td>6.3.3.2 Quantitative analysis: Real time qPCR</td><td>227</td></tr><tr><td>6.3.4 Discussion</td><td>230</td></tr><tr><td><b>6.4 Conclusion</b></td><td><b>231</b></td></tr><tr><td><b>6.5 References</b></td><td><b>232</b></td></tr></table>
**6.1 Biological material and techniques**

Checking behavior of cells in vitro in contact with putative implantable scaffolds is a prerequisite before studying in vivo models. In the present work, we chose the cell line C3H/10T1/2 to check cells behavior, such as migration, adopted morphology and eventual differentiation when cultured on dense collagen threads.

**6.1.1 Cell culture**

**Cell type: C3H/10T1/2**

The C3H/10T1/2 cell line was established by CA. Reznikoff, D.Branchow and C.Heidelberg in 1973 from C3H mouse embryos. According to their studies, these cells have similar properties and stability as murine fibroblasts after prolonged culture. The cells are very sensitive to post confluence inhibition of cell division. C3H/10T1/2 are functionally similar to mesenchymal stem cells or multipotent stromal cells that means that they can differentiate into different cell types (bone, cartilage, adipose cells etc). This cell line was chosen to be cultured on dense collagen threads in order to observe their morphologies, migration and behaviors on them.

<figure><img src="image_140.png" /><figcaption>Figure 6.1: C3H/10T1/2 cells morphology at: A) low and B) high densities. Images from ATCC, scale bars: 100 µm.</figcaption></figure>

**Cell culture**

C3H10T1/2 cells were cultured on culture flask of 75 or 150cm² with Dulbecco's Modified Eagle's Medium (DMEM low glucose, pyruvate) supplemented with 10% fetal bovine serum (FBS), and penicillin/streptomycin (100 U/100 µg per mL). Cells are put in an incubated chamber maintained at 37°C and 5% CO₂. Passages 18 to 23 were used during experiments.

**6.1.2 Cell culture on threads under tension**

The protocol from Kapacee et al. [1] is adapted in order to put threads under static tension during the cell culture. Figure 6.2 presents the protocol used in our study. Threads are cut into segments of about 2 – 3 cm length and put under tension on already prepared SYLGARD silicon gel-covered wells (Dow Chemical, Midland, MI, USA), in which two 8 mm sutures (Ethican, Sommerville, NJ, USA) are pinned 20 mm apart as shown in Figure 6.2. Each collagen thread is seeded with 2 mL of cells in culture medium with 350000 cells and placed to incubation conditions. Control collagen threads without cells are also followed. Culture medium is renewed every two days. After 1, 7, 14 and 21 days,
<figure><img src="image_141.png" /><figcaption>Figure 6.2: Protocol of cell culture under tension. 1) 6-well plate bases were covered with silicon gel (SYLGARD) and silk sutures are pinned with needles. 2) Threads are put under tension between the two sutures and 3) C3H10 T1/2 are added. 4) After 3 weeks of culture, 5) threads extremities are cut and observed by different techniques.</figcaption></figure>

the threads in presence of cells and the controls were retrieved by cutting the two extremities. For microscopic analysis, they are fixed in a 4% of paraformaldehyde (PFA) in PBS solution. Cell cultures were performed both at the LCMCP and in collaboration with Delphine Duprez and Marie-Ange Bonnin from the Laboratoire de Biologie du Développement.

## Static load range estimation

The explanation given below provides a reliable underestimate of the force applied. This estimation relies on the fact that the only force applied on the threads under tension is gravitational. Still, the threads extremities were also pinned apart in order to apply an additional tangent force, but the reproducibility was not easy and there was no possibility to measure these potential additional forces accurately.

A collagen thread pinned 20 mm (= 2L) apart can be considered as a flexible chain suspended between two points at the same height (Fig. 6.3). The thread is assumed to be almost cylindrical about its lowest point which is located at a height a from the anchored point altitude. The horizontal tension $T_h$ is estimated from the catenary curve equation which relates this constant tension with the respective lengths L, a, the apparent linear density $\mu_{app}$ of the thread, and the gravitational force g [2]:

$$T_h = \frac{L^2 - a^2}{2a} \mu_{app} g$$

<figure><img src="image_142.png" /><figcaption>Figure 6.3: Scheme of collagen thread cell culture configuration under tension. The horizontal tension $T_h$ is estimated by assuming that this situation is alike flexible chain suspended between two anchoring points.</figcaption></figure>
*Numerical application:* The different values of the physical quantities are: 2L = 20 mm, *a* ≈ 1 mm, *g* = 10 m/s². The apparent linear density μ<sub>app</sub> is the linear density of the thread adjusted by the buoyancy: μ<sub>app</sub> = μ<sub>thread</sub> - μ<sub>water</sub> = 0.12 g/m with μ<sub>thread</sub> = 0.25 g/m and μ<sub>water</sub> = πR²ρ<sub>water</sub>. We found an estimated horizontal tension of about: T<sub>h</sub> = 62.10<sup>-3</sup> mN. This corresponds to a strain of: l/l<sub>0</sub> = σ/E = T/EπR<sup>2</sup> where E is the Young modulus of the thread and is about 0.3 MPa. We found a static strain Δl/l<sub>0</sub>=0.16%. This is small compared to strain order of magnitude (about 1%) in the study of Kuo et al.[3] while using cycling load-
ing.

**6.1.3 Confocal Fluorescence Microscopy**

**Principle**

Confocal fluorescence microscopy allows 3D high resolution imaging (lateral:140 nm; ax-
ial: 1 µm) of labeled biological samples with little background noise. In a confocal fluo-
orescence microscope (Fig 6.4), the labeled sample is excited by a laser beam emitting
generally in the field of visible or ultraviolet. The fluorescence is excited over almost the
entire path of the beam. Thus, the fluorescence arising from out-of-focus positions rep-
resents an undesirable signal. In order to eliminate the majority of this spurious fluores-
cence, a pinhole is placed before the detector in the image focal plane of the microscope
objective. This hole allows to obtain a very good axial resolution but also to improve the
lateral resolutions. It is thus possible to scan the sample plane by plane at various depths
and finally reconstruct 3D images.

<figure><img src="image_143.png" /><figcaption>Figure 6.4: Scheme of confocal microscope: the laser beam illuminates a tiny region (through a first pinhole) of the sample in the focal plane. This light passes through the dichroic mirror and the second pinhole to a detector and a computer. The out-of-focus light is thus blocked. The pinhole is scanned sideways to form an image of the entire focal plane. The pinhole can then be scanned up and down to gather images from different focal planes and finally reconstruct 3D images. From "2-photons microscopy vs confocal microscopy" https://cogsci.stackexchange.com</figcaption></figure>

The observations were performed with an inverted Leica SP5 or Leica SP2 multipho-
ton laser scanning confocal microscope by Adèle Mauroux (PhD student under the direc-
tion of Laurent Müller) at the Collège de France. Images analysis and projections were
done with ImageJ software.

Fluorescent sample staining

Two fluorochromes are used to image the cells cultured on the collagen threads: DAPI (4',6-diamidino-2-phenylindole) which binds to DNA (nucleus) and Alexa Fluor 488 Phalloidin which binds to cells actin filaments (cytoskeleton). DAPI and Alexa Fluor excitation wavelengths are respectively 358 and 495 nm and their emission wavelengths 461 nm and 518 nm (Fig 6.5). There is an overlap region, which can lead to receive some DAPI signal in the Alexa channel.

<figure><img src="image_144.png" /><figcaption>Figure 6.5: Fluorescence spectra of DAPI (blue) and Alexa Fluor 488 Phalloidin (Green). Dashed lines are the excitation plots. Emission plots are represented by colored area. From "Fluorescent SpectraViewer" https://www.thermofisher.com.</figcaption></figure>

**Sample preparation:** Threads cultured with cells are fixed in solution of 4% PFA (diluted in PBS). Cells are then permeabilized by using 0.2% Triton X-100 for 15 min. This step allows the different fluorochromes to access their respective targets inside the cells. After rinsing with PBS solution, threads with cells are immersed in Alexa Fluor 488 Phalloidin (dilution: 1:200 in PBS) left in the dark at 4°C during 1 hour. Several PBS rinsing are needed to remove Alexa excess. Finally, the DAPI solution (dilution: 1:50000 in PBS) is added during 15 min and rinsed before the microscope observations. The same procedure is performed for SHG observations coupled with fluorescent microscopy.

**6.1.4 Histology and Hybridization in situ**

To examine the morphology, architecture and composition of the cells potentially
present in threads, several techniques are available such as histology and in situ hybridiza-
tion (ISH). ISH was also performed in collaboration with Delphine Duprez and Marie-
Ange Bonnin from the Laboratoire de Biologie du Développement. Threads with cells are
first removed and fixed at day 1, 7, 14 and 21 in a 4% of paraformaldehyde diluted in PBS
solution. Then, they are dehydrated by successive graded ethanol baths, xylene solutions,
and finally liquid paraffin. The successive baths gradually expose the sample to changes
in hydrophobicity, and avoid damaging cells. Threads in liquid paraffin are placed into a
mold with more paraffin. The wax finally cools down and forms a block that can be held
in a microtome and sliced. The transverse threads sections of 8µm are mounted on a slide in preparation for component identification via staining or in situ hybridization.

Commonly used staining of **Hematoxylin and Eosin** is performed on those threads slices. Hematoxylin colors the cell nucleus in dark blue and Eosin colors the remaining cell components in pink. This staining allows us to know if there are cells or not in the threads.

**In situ hybridization** allows localizing the expression of a gene of interest in its cellular environment. First, successive pre-treatments (deparaffinization, enzymatic digestion, permeabilization...) are performed on threads sections mounted on slides in order to avoid non specific hybridization and to access the cell genetic material.

<figure><img src="image_145.png" /><figcaption>Figure 6.6: In situ hybridization main steps: A) Membrane permeabilization, denaturation and addition of the probe. Hybridization occurs when the probe combine to the targeted RNA single strand. B) Addition of the antibody which binds to the probe. C) Colorless compound addition which becomes purple if the gene of interest was present. Adapted from [4].</figcaption></figure>

Then, the mRNA contained in cells and probes are denaturated (double to simple stranded RNA) by increasing the temperature. The probe is RNA strand labeled with a dioxygenin molecule (DIG). This probe will specifically hybridize (combine, base-pair to the corresponding mRNA single strand) to the mRNA of interest (Fig 6.6, A). An alkaline phosphatase conjugated with anti-DIG antibody is designed to detect the dioxygenin on the probe (Fig 6.6, B). Finally, a colorless compound that becomes purple when the phosphatase is removed, is added (Fig 6.6, C). Washing steps are necessary between the different probe and antibody additions to remove non-reagents excess.

**Choice of genes of interest:** Based on the literature [5], four mouse genes were chosen for the in situ hybridization:

- **Myogenic Differentiation1 (MyoD1):** this gene encodes a nuclear protein which plays a major role in the regulation of muscle differentiation. It was chosen as a control: if mesenchymal stem cells (MSC) differentiate into muscle-like cells, ISH will reveal this character. On the contrary, if MSC differentiate into another cell type, the expression will be lower and so its signal by ISH too. [5]

- **Collagen, type I, alpha 1 (Col1a1):** this gene encodes the alpha-1 subunit of the fibril-forming type I collagen which is the most abundant protein in most of connective tissues such as tendon. [5]

- **Scleraxis (Scx):** this gene encodes a protein (transcription factor) involved in mesoderm and expressed during the embryonic development of tendons and ligaments. [6]

- **Tenomodulin (Tnmd):** this gene encodes a type II transmembrane glycoprotein which is expressed at high levels in tenocytes and ligamentocytes. [7]
### 6.1.5 Real time q-PCR

#### 6.1.5.1 Principle

The Real-time quantitative Polymerase Chain Reaction allows to quantify the amplification of a targeted DNA molecule in real-time. It is divided in three main steps as shown in Figure 6.9.

<figure><img src="image_146.png" /><figcaption>Figure 6.7: Scheme of real time qPCR main steps: I) RNA extraction of cells cultured on collagen threads, II) Reverse transcription to synthetize cDNA from RNA strands, III) Polymerase Chain Reaction, which allows the amplification of cDNA material and thus quantification of gene expression.</figcaption></figure>

- I) **RNA extraction**: the collagen threads cultured with cells are removed from the 6-well plates and cell lysis is performed with Trizol in order to collect all the genetical material. Then, the RNAs are extracted through several phase separation, purification and centrifugation steps. All this procedure must be performed in very clean environment as RNase components and contaminants are present everywhere. Finally, the total RNAs concentrations are measured by spectrophotometry.

- II) **Reverse transcription (RT)**: this step consists in synthetizing a complementary double-stranded DNA from total RNA. First, a short single-stranded random primer anneals to the RNAs, which provides an initiation point for the reverse transcriptase (RTase) enzyme. In the presence of deoxynucleotide triphosphates dNTPs (dA, dT, dG, dC), the reverse transcriptase extends the primer to make a complementary DNA strand. Then, the RNA template is digested by ribonuclease (RNase) and the complementary strand of first cDNA is synthetized.

- III) **Polymerase Chain Reaction (PCR)**: this last step amplifies the DNA fragments of interest in order to be detected. First the temperature is increased in order to denature double-stranded into single-stranded cDNA. Then, specific chosen primers
anneal to their cDNA of interest. In the presence of dNTPs elongation of the com-
plementary strand occurs. Then, SYBR green which can bind to all types of double-
stranded nucleic acids, is intercalated while the strand is elongated. Once fixed, it
becomes a very good fluorophore and can be detected under excitation. All this
corresponds to one cycle. After one cycle, we obtained 2ⁿ copies of labeled cDNA
where n is the initial number of double-stranded cDNA. At the end of each cycle, the
fluorescent signal is measured allowing the quantification of the gene expression.

**6.1.5.2 Protocol**

RT-qPCR was performed to assess the mechanical load and time influence on the gene expression. For this purpose, collagen-based threads seeded with cells under static tension or not were removed at day 1, 7, 14, 21 to extract cells genetical material. For each condition, biological samples (threads with cells) sixplicat were analyzed by RT-qPCR.

**Extraction and purification of total RNAs**

The collagen-based threads cultured with cells are first immersed in a guanidium thiocyanate-phenol-chloroform solution, called TRIzol, which contains phenol (pH = 6). This solution allows lysing the cells and extracting the RNAs. The guanidium thiocyanate forms a complex with RNAs and protect them. In order to be sure to extract all the RNAs and to lyse the maximum number of cells that may have penetrated the matrix, the threads are scraped with cell scrapers. The supernatant is then collected in tubes and placed for storage at -80 °C.

After thawing and centrifugal steps, it is possible to eliminate all aggregates of col-
lagens which would have been dragged in TRIzol. The supernatants are then collected
and chloroform solution is added. After stirring, an emulsion is formed with RNAs will
moving into the choloroform phase. Centrifugation allows separating the two phases: the
phenol containing the cell debris and the genomic DNA at the bottom, and, the chloro-
form containing the RNAs at the top (and between them a phase composed of proteins).
The RNAs are then purified with an RNeasy kit (Qiagen, France). Briefly, the supernatants
(chloroform + total RNA) are removed and 70% ethanol is added. The samples are then
poured onto extraction columns. Total RNAs greater than 200 nucleotides in size bind to
the column membrane and the other elements are removed by centrifugation. The col-
umn is then washed with a solution containing guanidine. A solution containing a high
concentration of NaCl is added and allows the elution of the proteins still present. This is
repeated a second time and the tubes are then eluted with RNase-free distilled water. This
operation is repeated and the final aliquot is constituted by the total RNAs considered to
be purified. The purified samples are stored at -80 °C.

**Determination of total RNAs concentration**

Before carrying out the reverse transcription, the total RNAs concentrations are first mea-
sured by spectrophotometry. The absorption spectrum of the nucleotides shows a maxi-
mum peak at 260 nm, that of the aromatic amino acids at 280 nm. Thus, a double reading
at 260 and 280 nm and the ratio of optical densities (OD) OD260/OD280 determines the
degree of contamination by proteins. A ratio lower than 2 is a sign of high protein con-
tamination while ratios between 1.8 and 2.2 are considered acceptable.

**Reverse transcription (RT)**
The aim of the reverse transcription is to transform the extracted RNA into complementary DNA (cDNA) in order to carry out the polymerase chain reaction (PCR). To do so, non-specific primers capable of binding randomly to the RNAs are added to the RNAs aliquots. A first step at 65°C allows the denaturation of the secondary structures of the RNAs and fixing the primers. The tubes are then placed in ice and several reactants are added: 5X reaction buffer, DTT (Diothiothreitol), a reducing agent which destabilizes the RNases, dNTPs and the necessary enzyme (reverse transcriptase of the Moloney virus of murine leukemia). A step at 37°C allows the action of the enzyme and it is followed by a passage at 70°C which stops the reaction. The products of the RT are stable cDNAs which can be stored at -20°C.

## Choice of PCR primers

In order to prepare the necessary primers for qPCR, the sequences of the mRNAs of the genes were searched in the NCBI database (www.ncbi.nlm.nih.gov).

<table><caption>Table 6.1: Table of the primers</caption><thead><tr><td>Name</td><td>Direction (F or R)</td><td>Sequence</td></tr></thead><tbody><tr><td>36B4</td><td>F</td><td>ACCTCCTTCTTCCAGGCTTT</td></tr><tr><td>36B4</td><td>R</td><td>CTCCCACCTTGTCTCCAGTC</td></tr><tr><td>Collal</td><td>F</td><td>TGGAGAGAGCATGACCGATG</td></tr><tr><td>Collal</td><td>R</td><td>GAGCCCTCGCTTCCGTACT</td></tr><tr><td>Scx</td><td>F</td><td>CCTTCTGCCTCAGCAACCAG</td></tr><tr><td>Scx</td><td>R</td><td>GGTCCAAAGTGGGGCTCTCCGTGACT</td></tr><tr><td>Tnmd</td><td>F</td><td>AACACTTCTGGCCCGAGGTAT</td></tr><tr><td>Tnmd</td><td>R</td><td>AAGTGTGCTCCATGTCATAGGTTTT</td></tr></tbody></table>

The primers are in freeze-dried form and must be reconstituted in 1 mL of sterile water and then diluted in order to reduce their concentration to 7 µM. The primers pairs (Reverse R and Forward F) are then tested in conventional PCR under the conditions that will be used for the PCRs. For each pair, two mixtures are made: one contains a standard cDNA load, which is a mixture of 1 µL of all the samples, and makes it possible to verify that the primers amplify the target in a specific manner; an other mix contains no cDNA, but sterile water and ensures that the primers do not form dimers after 40 cycles. Once these two steps have been validated, the primers can be used in qPCR.

## Quantitative Polymerase Chain Reaction (qPCR)

Genes of interest have been studied in qPCR in order to compare their expression under different conditions and as a function of time. This technique consists in measuring the quantity of amplicons produced in each cycle thanks to a fluorescent intercalator, here SYBER Green. The intensity of the fluorescence increases with each cycle until it exceeds the detection threshold defining the Ct (cycle threshold), ie the number of cycles after which the number of amplicons formed exceeds the background noise. The smaller the Ct, the higher the initial target cDNA amount. It is thus possible to compare the number of initial copies between the different samples because at Ct, each tube has the same number of amplicons.

Initially, the pairs of primers are tested under the conditions of qPCR using standard samples. On a 96-well plate, for each tested pair of primers, a 5-point standard range, which
are dilutions of the standard (the same as for testing the primers in conventional PCR), are carried out at dilutions: 1/5, 1/25, 1/125, 1/625, 1/3125 with two negative controls (sterile distilled water) per range. Each range point is made in duplicate. The mixtures are prepared using the LightCycler 480 SYBR green I master kit containing the SYBR Green intercalant which attaches to the double-stranded DNA. The implemented temperature program is identical to that used in conventional PCR (but where the cycle times are reduced to a few seconds, here 10s-15s-15s) and the intensity is measured at the end of each cycle by a Light Cycler 480 system (Roche). The results are then analyzed using the LightCycler 480 software and allows us to determine the efficiency of the PCR.

First, the specificity of the amplification is determined through the calculation of the melting curve (Fig. 6.9). The melting curve corresponds to the fluorescence variation as a function of the temperature. When the temperature reaches the Tm (melting temperature) of the amplicon, it degrades and the fluorescence falls. Thus, when only one amplicon is present, a single fluorescence drop is recorded, which proves the specificity of the pair of primers (R and F). This step allows us to verify that the Tm of the amplicon corresponds to the one expected during the preparation of the primers.

<figure><img src="image_147.png" /><figcaption>Figure 6.8: Schemes of melting curves for the calibration points (top) and its respective derivative curve (bottom) with the melting temperature Tm.</figcaption></figure>

Then, the efficiency (curve slope) of each pair is calculated by introducing the cycle threshold Ct (Fig. 6.9) and by assigning arbitrary values to the (1/5, 1/25, 1/125, 1/625, 1/3125) dilutions. If the efficiency is greater than 1.85 then the pair of primers may be used for qPCR under these conditions.

For each gene tested, a 384-well plate was used. Each sample (thus each biological triplicate for each condition for each time) is made in triplicate, for the gene of interest and for the 36B4 reference gene [8]. Moreover, on each plate, two reference ranges (for the gene of
<figure><img src="image_148.png" /><figcaption>Figure 6.9: Scheme of the amplification curves with the cycle threshold Ct definition and the standard curve for the efficiency calculation</figcaption></figure>

interest and for the reference gene) are produced identical to those made for primers testing. The relative quantification technique used consists in transferring the Ct obtained on the standard points realized at the same time as the samples, which gives arbitrary values for each Ct measured. The arbitrary values of the samples of the target gene are then compared with the arbitrary values of the samples of the reference gene, which gives a ratio Target Gene/Reference Gene for each sample of the plate. It is this ratio, given by the software and that takes into account the efficiency, which we then use to compare the different conditions. For this purpose, a sample is chosen as the calibrator point and a second ratio of Studied condition ratio/Reference condition ratio is calculated.

Here the reference condition used is one of the biological triplicatas: **Coll 30mg/mL D7**
**under static load** which therefore has a ratio of 1. The ratios of the other conditions are
then compared to each other and presented in the form of histograms.

**Statistical analysis**

The gene expression results of real time qPCR are presented as the mean value ± SD (standard deviation). The statistical significance was assessed using one way analysis of variance (ANOVA) followed by Tukey posthoc test (comparing all pairs of groups) while comparing 3 data sets. We used this test for evaluating the influence of cell culture time as 3 different times. The significance level in all statistical analyses was set at a probability of P < 0.05. Example: collagen 30 mg/mL under static load at 1w, 2w and 3w were compared together and if there was a significant difference, it also referred to the time = 1 week. If significance was revealed between 2 and 3 weeks, it was signaled. For the comparison of two relevant data sets (ex: coll 30 mg/mL and col:alg both under static load at 1w), unpaired t test was used. Here again, the significance level was set at a probability of P < 0.05. We performed all data analysis with the software Prism (Graphpad).
**6.2 Cell colonization study**

In this first study, the cell colonization is studied as a function of collagen threads con-
centration (30 and 60 mg/mL) produced as previously described (Chapter 5 5.4.1 p 180).
The collagen threads with cells are put under static load during 3 weeks. First, we will
present the tensile testing results of the threads seeded with cells. Then, we will observe
the cell adhesion and migration by SHG microscopy coupled with fluorescence. Finally,
four genes will be tested to assess qualitatively the mesenchymal stem cells character after
being cultured 3 weeks on collagen threads.

**6.2.1 Mechanical properties of threads cultured with cells**

Tensile testing were performed after 3 weeks of cell culture on collagen threads under
static load. Figure 6.10 shows the results obtained for such threads and the respective
control either in fibrillogenesis buffer or DMEM as previously analyzed.

<figure><img src="image_149.png" /><figcaption>Figure 6.10: Tensile testing results of collagen threads at 30 mg/mL (blue) and 60 mg/mL (purple) seeded with cells and cultured over 3 weeks under static load and their respective controls. Top: Young's modulus and Bottom: Ultimate Tensile Stress</figcaption></figure>

Collagen threads at 60 mg/mL seeded with cells exhibit the best mechanical properties with E = 1.11 ± 0.30MPa and UTS = 0.14 ± 0.04MPa compared to collagen threads with
cells at 30 mg/mL. Young modulus is almost 10-fold and UTS 3-fold higher when the col-
lagen thread concentration is 60 mg/mL. We can also notice that there is a slight decrease
of both E and UTS for threads with cells compared to their controls in PBS 5X and in
DMEM without cells for the two concentrations. UTS for collagen threads at 60 mg/mL
with cells has a higher decrease compared to the controls. These decreases might be ex-
plained by the cell sheath around the threads which sometimes breaks away at the clamp-
ing jaw level leading to a few sliding of the thread during the tensile testing. Some set-up
improvements will be performed to avoid this kind of artifacts.

Limited literature was found about the mechanical properties of collagen scaffolds seeded
with cells *in vitro*. In most of the studies, cells were directly put inside the collagen based
construct.

Gentleman et al. performed tensile testing until failure of collagen threads at 10 mg/mL [9]. Post extrusion process of dehydration and cross-linking improved their mechanical properties as explained in the previous Chapter 5. However, relative analysis can be done from their collagen constructs study with and without cells. These collagen fibers (N=50) were embedded in a cell laden collagen gel (2.77 mg/mL) and cultured under static load during 25 days. The tangent moduli were compared with the corresponding scaffold without cells. They found a significant difference between the two constructs: the moduli of cell laden construct is almost 1.7-fold higher than the one without cells. In this case, cells improve mechanical properties of the constructs.

Kim et al. used 3D bioprinting to fabricate collagen based construct laden with cells and prior cross-linked with tannic acid [10]. In the case of pure cell laden collagen scaffold (40 mg/mL), they obtained Young's modulus E of 0.04 MPa. In our case, with collagen threads at 30 mg/mL, E is 3-fold higher which is quite promising.

Chen et al. produced cell laden collagen rings at various collagen concentrations (0.5 to 2 mg/mL) [11]. No post processing was performed. They showed that the modulus and UTS decreases while increasing the collagen concentration of their cell constructs. Its maximum (about 80 kPa and 35 kPa for modulus and UTS respectively) is reached for the collagen concentration of 0.5mg/mL. In our case, the opposite trend is observed and our results are significantly higher. But it should be kept in mind that cells may act differently if they are put from the outside or directly inside the collagen scaffold.

6.2.2 Cell migration within the threads at different collagen concentra-
tions

Cell migration within the threads was followed by SHG microscopy coupled with flu-
orescence at 3 and 4 weeks of cell culture for both collagen concentrations. Figure 6.11
displays the results obtained for collagen signal (Green) and cells fluorescence (Red and
Blue).

**Collagen concentration effect:** at 30 mg/mL and 60 mg/mL, C3H10 T1/2 cells covered the entire threads surface resulting in a cell multilayer of about 50µm representing around 5 cellular layers. In the case of collagen thread at 30 mg/mL at 4 weeks, we suspect the cell sheath to be ripped of the thread during handling. In all cases, cells adhere on the threads and proliferate a lot. The most notable concentration effect is the cell penetration. At
<figure><img src="image_150.png" /><figcaption>Figure 6.11: SHG images coupled with fluorescence of collagen threads seeded with C3H10 T1/2 cells during 3 weeks at 30mg/mL and 60mg/mL, under static load. Their respective cross sections projection are also given (bottom). Green: SHG signal; Red: Alexa Fluor 488 Phalloidin; Blue: DAPI. Scale bar = 50 µm.</figcaption></figure>

30 mg/mL, several holes are observed in the collagen threads where SHG signal is zero. Cells clearly invade those holes which are assumed to be created by cells through collagen matrix digestion as shown by Helary et al. [12]. In fact, without cells, such straight holes
were not observed (Chapter 5 5.4.1 p 180) on collagen threads. Thus, cells penetrate over 30µm at 30 mg/mL whereas at 60 mg/mL, they stay at the threads surface. This might be explained by the collagen density which is higher for 60 mg/mL and so the collagen "network" does not allow the cells to enter easily. Moreover, the stiffness is not the same which can also explain this observation.

**Cell culture time effect:** for cell cultures time of 3 or 4 weeks, the cell sheath thickness is almost the same for both conditions. No clear effect on the cell penetration is observed at 30 mg/mL. We can notice that cells do not seem to suffer from the high confluence and stay alive event after 4 weeks of culture.

**Axis effect:** at 30 mg/mL, straight holes in the collagen thread are observed. Each hole is equivalent to the size of one or two elongated cells. Their shape is aligned along the static load axis. Moreover, it has been shown that collagen thread surface is aligned along its longitudinal axis which can also induce cell orientation. Figure 6.12 shows the cell sheath surfaces. The cell actin filaments give the main orientation of the cells which is underlined by orange arrows. They follow the axis of the thread which is the same as the load axis. Thus it cannot be yet distinguished which mechanism is responsible for this cell direction. Observations of collagen threads seeded with cells without tension would be performed to answer it.

<figure><img src="image_151.png" /><figcaption>Figure 6.12: SHG images coupled with fluorescence of cell sheath surfaces on collagen threads at A)30mg/mL and B)60mg/mL, under static load. Red: Alexa Fluor 488 Phalloidin; Blue: DAPI. Orange arrows are eye guidelines to see the cell actin filaments main orientation. Scale bar = 50 µm.</figcaption></figure>
6.2.3 Preliminary In situ Hybridization results

In situ hybridization was performed on collagen threads at 30 and 60 mg/mL after 3 weeks of cell culture with C3H10 T1/2. As previously introduced ISH allows us to assess qualitatively the gene expression of cells. These mesenchymal stem cells may differen- tiate under appropriate conditions (transcription factors, growth factors or mechanical stimulation) [13]. Figure 6.13 presents the results obtained with the four genes of interest. First, we can notice that cells penetrate more in the collagen thread at 30 mg/mL than collagen 60 mg/mL as previously observed by SHG with fluorescence. It is clearly visible on Tnmd images (Fig. 6.13 bottom left).

<figure><img src="image_152.png" /><figcaption>Figure 6.13: In situ hybridization on collagen threads cross sections at 30 and 60 mg/mL seeded with C3H10 T1/2 cells during 3 weeks. Four genes were labeled: MyoD, Col1a1, Scx and Tnmd. Tnmd and Col1a1 exhibit stronger color level labeling in both conditions. Scale bar = 100 µm.</figcaption></figure>
As explained in the previous section, ISH reveals the expression of a gene if the dye reacts. In other words, if a cell is colored for the tested gene, it means that the gene is expressed. If it is not, the gene is low or not expressed.

ISH experiments on threads cross-sections reveals that for both collagen concentrations, myoD is not expressed as the cell sheath is not "colored". Considering a purple level, Tnmd is more expressed than Col1a1 and Scx. However, compared to myoD, it is assumed that Col1a1, Scx and Tnmd are expressed which suggests that mesenchymal stem cells C3H10 T1/2 do not exhibit muscle character and may differentiate into tendon-like cells on our collagen threads whether at 30 or 60 mg/mL. Further quantitative analysis are performed in the second study in order to assess more precisely the gene expression.

### 6.2.4 Conclusion

In this first study, we showed that mesenchymal stem cells adhered on pure collagen threads and could even penetrate over a depth of 30µm, leading to straight holes at the collagen thread surface as shown by SHG (6.11 p 219). The cellular sheath of thickness about 50µm was quite aligned along the thread axis. We assumed that this cell orientation may come from the static tension or the thread topography at which striated patterns were observed without celle by SEM (Chapter 5 5.4.1 p 180). Mechanical properties were not affected by the presence of the cells. However, measurements improvements must be done in order to avoid cellular sheath gliding during the tensile testing. Finally, the first qualitative analysis of the gene expression of the mesenchymal stem cells cultured on threads were very promising and indicated that these cells may differentiate into tendon-like ones. A further study of cell behavior on threads was thus needed to quantify their gene expression level.
## 6.3 Behavior of stem-cells cultured on collagen based threads

In this second study, collagen thread at 30 mg/mL is chosen for quantitative analysis of stem cells behaviors as cells adhere and penetrate significantly into it. We also decided to compare this condition with the best corresponding collagen-alginate ratio thread (Chapter 5 5.4.2 p 188). The idea is to evaluate the effect of alginate addition on cell behavior.

### 6.3.1 Load influence of threads morphologies

Collagen based threads morphologies were observed by phase contrast microscopy over the cell culture time and as a function of the load condition. Figure 6.14 shows the morphologies of such threads at 3 weeks. Under static load (A), threads are straight. Coll:Alg (5:1) thread is opaque as previously noticed (Chapter 5 5.4.2 p 188). Without static load (B), Coll:Alg (5:3) is also quite straight whereas collagen thread at 30 mg/mL collapses and looks like a ball of whool. The thread contour is discernable through the cell sheath. This phenomenon due to the contraction of the thread by the cells begins between the first and the second week of culture.

<figure><img src="image_153.png" /><figcaption>Figure 6.14: Morphologies observed by phase contrast microscopy of collagen thread at 30 mg/mL (left) and collagen:alginate (5:1) thread (right) after being seeded with cells and cultured during 3 weeks A) under static load and B) without load. Scale bar = 500 µm.</figcaption></figure>

Chen (M-Y) et al. investigated the contraction of cell-laden collagen gels either as a function of the cell or collagen initial concentrations [14]. The collagen matrix contraction can be divided into three phases: a lag phase, a linear rapid contraction phase, and a slow contraction phase [15]. While increasing collagen concentration, the lag phase
is longer. Moreover, while increasing the initial number of cells inside the gel, this first
phase is shortened. They found a lag phase of 14 days for collagen cell-laden gel at 2
mg/mL. Following this affirmation, lag phase for our collagen threads must be longer as
concentration is 10-fold higher. However, in our case, cells number increases over time as
cells have enough space to multiply and are not confined in a gel. This is why contraction
occurs more rapidly. We also removed during 1 day the static load on a collagen thread at
30 mg/mL after 3 weeks of cell culture. The cells contract the threads as if no tension had
been applied during 3 weeks. This is in good agreement with Chen et al. results [14] who
found a lag phase of 0 to 2 days for initial cell concentration of 1 to 0.25 ×10⁶cells/mL.

For coll:alg condition, threads keep straight with and without tension. A small inflec-
tion is visible for some threads (Fig 6.14, right B)). Regarding the mechanical properties of
such thread, it may be too stiff to be contracted by the cells (6.2.1 p 217). Moreover, coll:alg
threads diameter is larger than collagen one and the contraction cell force per area unit is
thus lower. The lag phase might be too long to be observed within our culture time scale.

**Remark:** in Figure 6.14, we can distinguish cells on the well plate which suggests that cells migrate and proliferate rapidly as threads were put in a new 6-well plate every 2 days.

**6.3.2 Cell colonization over culture time and load condition**
**impact analysis**

Cell colonization was assessed over time by confocal microscopy. After 3 days of cell
culture under static load, several cell-deserted areas are observed.

At day 7, these holes disappear for collagen 30 mg/mL threads and remain for coll:alg thread. In fact, the thread surface is smaller than for the mixed one. Thus, cells cover more rapidly this surface.

At day 21, the two types of threads are entirely covered by the cells and several cell layers
can be distinguished. Here again we see the thread contraction due to the cells at 30
mg/mL (Fig. 6.15, bottom left). In all cases, the outer layer of cells seems to be aligned
along the longitudinal thread axis. However, for condition coll:alg no effect of the tension
is visible. For pure collagen, some cell layers follow this thread axis orientation and other
are perpendicular to it (Fig. 6.15 arrows). Without tension, on the thread part located on
the left, which is still straight, cells are aligned along the thread axis (Fig. 6.15 arrows). Cell
morphologies may be induced to the thread topography "orientation".

From those observations, it is assumed that the static load does not induce the cell
orientation but this phenomena is due to the thread surface topology. Kishore et al. stud-
ied the effect of electrochemically collagen threads topography on cell morphology [16].
They fabricated either random or aligned collagen threads and put cells on it. Cells cul-
tured on random threads exhibit a polygonal phenotype, whereas they are well aligned
and have spindle shape on aligned collagen threads.
<figure><img src="image_154.png" /><figcaption>Figure 6.15: Confocal microscopy images of collagen thread at 30 mg/mL (left) and mixture coll:alg (5:1) (right) for different times: days 3, 7, 21 and without tension at 21 days. Alexa Phalloidin 488 and DAPI labeling. White arrows are eye guides for the cell sheet orientation on the thread surface. Scale bar = 250 µm.</figcaption></figure>

**Remark:** as collagen threads at 30 and 60 mg/mL, SHG coupled with fluorescence was perfomed on coll:alg (5:1) thread in order to observe how cells interact with the matrix.
The figure 6.16 displays the image of collagen (green) and cellular actin filaments (red) signals for a mix thread with cells cultured without tension.

<figure><img src="image_155.png" /><figcaption>Figure 6.16: SHG image coupled with fluorescence of coll:alg (5:1) thread at 2 weeks of culture with cells without static load. Cross sectional projection is given on the bottom. Green: collagen SHG signal, Red: Alexa Fluor 488 Phalloidin. Scale bar = 200 µm.</figcaption></figure>

Cells interpenetrate the surface of the collagen-based thread. Several "SHG signal" holes are present in the thread structure meaning that those areas are mainly composed of alginate. Cells do not colonize those areas. The alginate phase may degrade over time leading to new porosity inside the thread giving enough space for cells to invade.

### 6.3.3 Genes of interest expression analysis

The expression of the genes of interest were analyzed first qualitatively by ISH and then by Real-time qPCR.

#### 6.3.3.1 Qualitative analysis

ISH allows us to estimate gene expression prior Real-time qPCR. ISH reveals the expression of Scx in both conditions: collagen 30 mg/mL and coll:alg (5:1) thread. Col1a1 seems to be more expressed on coll:alg with brighter color level. Tnmd is lower expressed than Scx and Col1a1 for both threads.

The myoD gene is not expressed attesting that mesenchymal stem cells C3H10 T1/2 do not exhibit muscle character. As tendon markers are more or less expressed for each type
<figure><img src="image_156.png" /><figcaption>Figure 6.17: In situ hybridization on collagen threads at 30 mg/mL (left) and coll:alg (5:1) thread (right) with C3H10 T1/2 cells at 3 weeks, under static load. Four genes were labeled: MyoD, Col1a1, Scx and Tnmd. Scale bar = 200 µm.</figcaption></figure>

of threads, it is assumed that C3H10 T1/2 may differentiate into tendon like cells. This effect seems to be more pronounced for threads with alginate than pure collagen one.

**Remark:** no significant difference was noticed in the case of no load condition.

### 6.3.3.2 Quantitative analysis: Real time qPCR

The gene expression of Col1a1, Tnmd and Scx was analyzed quantitatively with real time qPCR and by using statistical tests to reveal the significant differences. For each gene, we will describe the results by varying one parameter at a time.
6.3.3.2.1 Col1a1 gene expression analysis

**Cell culture time effect:** For collagen 30 mg/mL threads, col1a1 gene expression increases over time with and without tension. Significant difference is noticed between 1 week and 3 weeks of culture. For coll:alg, col1a1 gene expression tends to decrease over time but no significant differences are observed.

**Material condition effect:** Significant differences are observed between pure collagen and mixed alginate threads for the first two weeks of cell culture. The gene expression is higher for cells cultured on coll:alg threads. At 3 weeks, col1a1 expression is almost the same for all conditions.

**Static load influence:** No clear effect of the tension is observed. Only one significant difference is noticed for mix threads at 3 weeks.

<figure><img src="image_157.png" /><figcaption>Figure 6.18: Collagen 1a1 gene: effect of cell culture time, load and material (pure collagen or mixed with alginate threads) conditions on tenogenic differentiation of mouse mesenchymal stem cells C3H10T1/2. (*P* < 0.05, **P* < 0.01, ***P* < 0.001).</figcaption></figure>

Conclusion

Col1a1 expression increases over time on pure collagen threads and decreases on mixed ones, on which the initial expression is higher than with pure collagen. However, after 3 weeks, expression tends to the same level in case; the expression of the cells is different at the beginning depending on the conditions but tends to the same value after three weeks. The static tension does not seem to present any stimulating effect and rather decreases the stimulating effect of the presence of alginate. Indeed, at 3 weeks, the expression in the non-loaded mix is higher than the loaded mix.
### 6.3.3.2.2 Tnmd gene expression analysis

**Cell culture time effect:** At 1 week, Tnmd is overexpressed for all conditions. The gene expression decreases significantly over the time in all cases.

**Material condition effect:** Tnmd expression is low for collagen 30 mg/mL (static load) at all times. It is significantly lower compared to coll:alg threads.

**Static load influence:** Tension has a significant increasing effect at 2 weeks of cell culture for both material conditions but at other time points, there is no difference.

<figure><img src="image_158.png" /><figcaption>Figure 6.19: Tenomodulin gene: effect of cell culture time, load and material (pure collagen or mixed with alginate threads) conditions on tenogenic differentiation of mouse mesenchymal stem cells C3H10T1/2. (*P < 0.05, **P < 0.01, ***P < 0.001)</figcaption></figure>

**Conclusion**

Tnmd gene is expressed at the beginning for all conditions but decreases over time. The material has an effect on its expression as it is significantly less expressed for pure collagen threads. Here again, it seems that the static load has no impact or a slight negative one.

### 6.3.3.2.3 Scx gene expression analysis

**Cell culture time effect:** Scx expression decreases significantly over time for pure collagen. It is quite constant over time for mixed threads.

**Material condition effect:** Scx is less expressed for pure collagen compared to coll:alg threads which exhibit very significant increase compared to pure collagen.
**Static load influence:** Static load has a positive effect at 1 week for pure collagen and a negative impact at 3 weeks for alginate mixtures.

<figure><img src="image_159.png" /><figcaption>Figure 6.20: Scleraxis gene: effect of cell culture time, load and material (pure collagen or mixed with alginate threads) conditions on tenogenic differentiation of mouse mesenchymal stem cells C3H10T1/2. (*P* < 0.05, **P* < 0.01, ***P* < 0.001)</figcaption></figure>

**Conclusion**

The thread material has clearly an influence on Scx expression as it is around 10-fold higher on coll:alg threads. Moreover, Scx expression is constant and remains at high level over the three weeks with or without load, whereas the expression decreases for pure collagen condition and the load has a negative effect at one week.

### 6.3.4 Discussion

From this gene expression analysis, the fact that the mesenchymal stem cells on the coll:alg mix threads express all tendon markers stands out. It is assumed that the cells behave as tenocytes while cultured on mix. The differentiation seems to be lower for pure collagen threads. This is in good agreement with the qualitative observations made from ISH images.

Kishore et al. studied the gene expression of Scx and Tnmd on random and aligned collagen threads. Scx expression is higher at day 3 on aligned collagen threads [16]. For Tnmd, they showed that its expression is 3-fold higher at day 7, and 10-fold higher when compared to random surfaces. The same trend is observed in our study. Based on their results, we can hypothesise that our coll:alg threads exhibit more aligned surface than the
pure collagen. This would favor tenogenic differentiation. Further investigations of surface characterization must be performed to verify this assumption.

Regarding the loading effect, Kuo et al. studied bone marrow mesenchymal stem cells gene expression [17]. Cells are put inside collagen gels at 3 mg/mL. They performed either static or cycling loading on those gels and cell culture was done over 7 days. They did not notice any difference while increasing the collagen gel concentration. Scx expression decreases over time for static loading, whereas Kuo et al. showed that static loading maintains Scx expression.

**In our study**, several hypotheses can be considered: the static load of about 125 mN is may not be sufficient enough to induce a difference with no tension condition. From this, we can conclude that thanks to its higher stiffness, coll:alg threads provide a better environment for cell differentiation. Based on Kuo et al. work [17], it will be interesting to perform the same study with cycling loading.

## 6.4 Conclusion

Mesenchymal stem cells were cultured on collagen-based threads over 3 weeks. From the studies of the cell colonization and behaviors on the threads, we showed that the cells could proliferate on the threads (pure or mixed condition) leading to a cellular sheath of about a width of 50µm. For pure collagen at 30 mg/mL, cells even digested collagen material and penetrated over a depth of 30µm from the thread surface. The cell sheath covered the whole thread surface available and seemed to be aligned along its axis. In Chapter 5, pure collagen threads exhibited striated patterns along the thread axis as shown by SEM. This oriented topography may guide and induce cell morphologies [18]. Moreover, the larger collagen structures (ball of wool) observed for collagen thread at 60 mg/mL could improve the cell adhesion. After 3 weeks, cells did not seem to suffer from the high confluence and preliminary ISH results proved that they could express specific genes of tenocytes.

The expression of these tendon specific genes were quantified by RT q-PCR for cells either cultured on pure collagen threads at 30mg/mL or its coll:alginate mixed equivalent ones. This was done as a function of time and mechanical loading. The ISH results were confirmed by this technique. Quite surprisingly, gene expression was higher in the case of collagen:alginate mixed thread than for pure collagen for Tnmd and scx. For Col1a1 gene, gene expression was almost the same after 3 weeks. Finally, the static mechanical loading was assumed to not have any effect on gene expression. This last parameter may be further investigated.
**6.5 References**

[1] Zoher Kapacee, Susan H Richardson, Yinhui Lu, Tobias Starborg, David F Holmes, Keith Baar, and Karl E Kadler. Tension is required for fibripositor formation. *Matrix Biology*, 27(4):371–375, 2008. 207

[2] *Differential equations*. D. C. Heath and Company, 1952. 208

[3] Catherine K. Kuo, Joseph E. Marturano, and Rocky S. Tuan. Novel strategies in tendon and ligament tissue engineering: advanced biomaterials and regeneration motifs. *BMC Sports Science, Medicine and Rehabilitation*, 2(1):20, 2010. 209

[4] *Developmental Biology*. Sinauer Associates, 2010. 211

[5] Ava E Brent and Clifford J Tabin. Developmental regulation of somite derivatives: muscle, cartilage and tendon. *Current opinion in genetics & development*, 12(5):548–557, 2002. 211

[6] Chisa Shukunami, Aki Takimoto, Miwa Oro, and Yuji Hiraki. Scleraxis positively regulates the expression of tenomodulin, a differentiation marker of tenocytes. *Developmental biology*, 298(1):234–247, 2006. 211

[7] Chisa Shukunami, Yuki Yoshimoto, Aki Takimoto, Hiroshi Yamashita, and Yuji Hiraki. Molecular characterization and function of tenomodulin, a marker of tendons and ligaments that integrate musculoskeletal components. *Japanese Dental Science Review*, 52(4):84–92, 2016. 211

[8] Jorge Laborda. 36b4 cdna used as an estradiol-independent mrna control is the cdna for human acidic ribosomal phosphoprotein po. *Nucleic acids research*, 19(14):3998, 1991. 215

[9] Eileen Gentleman, Andrea N Lay, Darryl A Dickerson, Eric A Nauman, Glen A Livesay, and Kay C Dee. Mechanical characterization of collagen fibers and scaffolds for tissue engineering. *Biomaterials*, 24(21):3805–3813, 2003. 218

[10] Won Jin Kim, Hui-Suk Yun, and Geun Hyung Kim. An innovative cell-laden α-tcp/collagen scaffold fabricated using a two-step printing process for potential application in regenerating hard tissues. *Scientific Reports*, 7(1):3181, 2017. 218

[11] Xiao Chen, Zi Yin, Jia-lin Chen, Wei-liang Shen, Huan-huan Liu, Qiao-mei Tang, Zhi Fang, Lin-rong Lu, Junfeng Ji, and Hong-wei Ouyang. Force and scleraxis synergistically promote the commitment of human es cells derived mscs to tenocytes. *Scientific reports*, 2:977, 2012. 218

[12] Christophe Helary, Ludmila Ovtracht, Bernard Coulomb, Gaston Godeau, and Marie Madeleine Giraud-Guille. Dense fibrillar collagen matrices: a model to study myofibroblast behaviour during wound healing. *Biomaterials*, 27(25):4443–4452, 2006. 219

[13] Giselle Chamberlain, James Fox, Brian Ashton, and Jim Middleton. Concise review: mesenchymal stem cells: their phenotype, differentiation capacity, immunological features, and potential for homing. *Stem cells*, 25(11):2739–2749, 2007. 221
[14] Meng-Yi Chen, Yulong Sun, Chunfeng Zhao, Mark E Zobitz, Kai-Nan An, Steven L Moran, and Peter C Amadio. Factors related to contraction and mechanical strength of collagen gels seeded with canine endotenon cells. *Journal of Biomedical Materials Research Part B: Applied Biomaterials*, 82(2):519-525, 2007. 223, 224

[15] Toshio Nishiyama, Naoki Tominaga, Keisuke Nakajima, and Toshihiko Hayashi. Quantitative evaluation of the factors affecting the process of fibroblast-mediated collagen gel contraction by separating the process into three phases. *Collagen and related research*, 8(3):259-273, 1988. 223

[16] Vipuil Kishore, Whitney Bullock, Xuanhao Sun, William Scott Van Dyke, and Ozan Akkus. Tenogenic differentiation of human mscs induced by the topography of electrochemically aligned collagen threads. *Biomaterials*, 33(7):2137-2144, 2012. 224, 230

[17] Catherine K Kuo and Rocky S Tuan. Mechanoactive tenogenic differentiation of human mesenchymal stem cells. *Tissue Engineering Part A*, 14(10):1615-1627, 2008. 231

[18] Giulio Abagnale, Antonio Sechi, Michael Steger, Qihui Zhou, Chao-Chung Kuo, Gülcan Aydin, Carmen Schalla, Gerhard Müller-Newen, Martin Zenke, Ivan G Costa, et al. Surface topography guides morphology and spatial patterning of induced pluripotent stem cell colonies. *Stem Cell Reports*, 2017. 231
CONCLUSION

The main objective of this thesis was to produce dense collagen threads mimicking
the native tendon structure to repair it in the future. For this multidisciplinary project, it
was thus necessary to combine both physico-chemical skills, due to the complexity of its
structure and biological ones, due to the final TE goal. As a consequence, a large range of
parameters from those different fields needed to be explored in order to fully understand
their influence and continuously improve the collagen threads properties until reaching
the best condition.

For this purpose, extrusion technique was chosen as it is quite convenient to implement
in a laboratory and is the first/intuitive method used to fabricate threads. We developed
a home-made set-up in order to control a large range of parameters and also to be able
to observe in real-time the thread production. Then, we decided to use a system model
similar to collagen that we could extrude in a vast amount to harness our experimental
device and first explore the influence of each parameter. Finally, we produced dense col-
lagen threads in a reproducible manner and in vitro studies were performed to assess the
behaviors of the cells cultured on our scaffolds.

First, we studied the extrusion of sodium alginate in the presence or not of binding ions (Chapter 4). In the first case, two characteristic instabilities of polymer melt extrusion were observed and quantified thanks to our experimental set-up configuration, which allows the visualization of the flow at the die exit and even inside the capillary. By varying die dimensions, we showed that the bulk undulation of the extrudate vanished for ratio of length over radius close to zero, indicating that the instability affected the whole pipe flow. Moreover, the shear-thinning level of such fluid solution may control the nature of the instability. It was the first time that such instabilities were reported in the case of polymer solutions. In the presence of binding ions, we reported several surprising instabilities such as the striated patterns on the thread surface or the gel clogging at the die exit. Such phenomena have never been previously described and further investigations were necessary to fully understand the mechanisms of alginate and calcium ions interaction during the extrusion.

Then, the new expertise on the complex fluid extrusion was used for the production of
collagen threads. We extruded for the first time collagen solutions at concentrations up
to 60 mg/mL (10-fold higher than in other studies) and succeeded to produce regular
threads in a reproducible manner, while controlling their quality in real time. The preparation and extrusion process allowed to obtain in only a few hours several meters of collagen threads about 500µm in diameter. We proved also that this method tends to align collagen molecules along the thread axis. Moreover, the studied alginate instability was successfully transposed on collagen systems, leading to the production of collagen threads exhibiting helical morphology, similar to the tendon "crimp" structure. This was obtained
at the end of the thesis and is promising compared to the previous studies on crimp replication (Chapter 4).

Hybrid threads were made of both collagen and alginate in order to take benefit from their respective properties (Chapter 5). After selecting the best collagen:alginate ratio, we showed that the addition of even a small amount of alginate improved the mechanical properties of the threads compared to the pure collagen ones.

Finally, mesenchymal stem cells were cultured on the collagen-based threads (Chapter 6). The cells colonized the threads surface and penetrated over a distance of 50 µm from the surface. They expressed the different genes characteristic of tendon-like cells suggesting that they differentiate into tenocytes when seeded on our threads. No difference was observed between cells cultured on threads with or without static load. However, the gene expression was improved on hybrid threads.

Although our threads exhibited promising properties, there is still a long way to repair the tendon and have a scaffold ready for clinical applications. This is why several further studies must be done on this subject and we will here detail many perspectives raised by our results.

## Perspectives

### Extrusion parameters investigation

Regarding the knowledge of the extrusion technique, several parameters still need to be further investigated. As we changed one parameter at a time, it was difficult to test each of them over a wide range. In fact, we began studying the effect of the die material and some corresponding results for alginate were presented. As the condition at the die wall seems to be critical to obtain some instability, it would be interesting to use PTFE tubing to remove stick effect on the wall and study the impact on instabilities. Then, we showed that the die exit angle can be shaped at a desired angle. From preliminary results, instabilities might be amplified by using sharp exit angles and this may be helpful to produce more pronounced crimped threads. Finally, the extrusion velocity range needs to be further explored and in particular for collagen solutions. Our experimental set-up also offers the possibility to impose the temperature during the extrusion. This parameter was not varied during this project. Such complex fluids may exhibit different behaviors while increasing or decreasing the temperature. In the case of collagen:alginate mixed solutions, it would be interesting to extrude them at lower temperatures which might enable to delay the two systems interactions so as to obtain more homogeneous threads.

### Threads properties improvement

#### *Composition*

In this thesis, no cross-linking agent or post-processing was used. In fact, those may decrease the biocompatibility of the scaffolds or lead to the denaturation of the collagen respectively. However, some natural cross-linking agents, present in the tendon, may be used to reinforce our threads. To do so, it would be interesting to introduce proteoglycans which bridge collagen molecules to each other. As shown in the Chapter 5, the threads produced from the higher collagen concentration exhibit the best properties. It may be possible to extrude collagen solution at concentration higher than 60mg/mL. This would lead to highly dense threads as shrinkage phenomenon might increase the local concen-
tration. Moreover, the strategy of mixing alginate and collagen seems to be very promising. By exploring the interactions between these proteins and polysaccharides, we would assess the best conditions to obtain quite cosoluble systems at high concentrations.

## Structure

Regarding its shape, the "crimp" like morphology was obtained at the end of this thesis project. It deserves to be further studied as it is easy to implement and does not need plenty of chemical treatments as in molding strategy. The die material may have an influence on this bulk "defect", which needs to be explored. Moreover, we did not have enough time to test the production of a "multiple"-based thread: the idea is to insert a metallic grid either upstream or downstream (to be chosen) in order to separate the flow into several parts that can be put back together downwards. This way we expect the alignment of collagen to be improved also in the core thread. A last approach would be to incorporate the cells directly inside the solution and extrude threads with cells inside. The cells may reorganize the collagen matrix at their length scale and improve the collagen structure.

## Scaffold assembly

Plenty of strategies may be applied to construct elaborated scaffolds from our threads. The first idea would be to braid several threads together as presented in Chapter 1 and form a cord-like construct. It would also be possible to assemble those cords and form a scaffold at a higher level. In our work, we observed the forces developed by the cells cultured on threads without static load. Some collagen threads were completely folded. It could be possible to assemble several threads pinned at their extremities and close enough to each other to allow the cell adhesion between threads. They would bind the threads together and produce their own extracellular matrix between them. In a similar approach, several threads "layers" could be aligned in a mold, which could also be filled with a collagen solution at a lower concentration and already containing cells.

## Cells behaviors

Mesenchymal stem cells expressed tendon-like cells gene while cultured on our collagen-based threads. Cell cultures on higher collagen concentration threads could be performed to assess the influence of this parameter on cell behaviors. The static load did not impact the cell behaviors. However, it would be interesting to test cell cultures on threads under cyclic load. This would be closer to the physiological situation. Some devices as Flexcell™ are available to perform such testing under cell culture.
List of abbreviations

– 36B4 : reference gene

– AA : acetic acid

– ALG: sodium alginate

– C3H10T1/2 : mesenchymal stem cell from mouse embryos

– C : concentration

– CaCl₂ : calcium chloride

– cDNA : complementary strand DNA

– χ : compressibility

– Col1a1: collagen type 1 alpha 1 gene

– coll : collagen

– C<sub>t</sub> : cycle threshold

– DAPI 4',6-diamidino-2-phenylindole dihydrochloride

– D<sub>p</sub>: piston diameter

– DSC : Differential Scanning Calorimetry

– DMEM Dulbecco's Modified Eagle's Medium

– DNA : Deoxyribonucleic acid

– E: Young Modulus

– ECM Extracellular matrix

– EDC : carbodiimide

– η : viscosity

– η₀ : plateau viscosity

– η* : complex viscosity

– γ̇ : shear rate

– γ̇₀ : zero shear rate

– γ : strain

– f : frequency

– F : force

– FACITs : Fibril Associated Collagen with Interrupted Triplehelix

– Ḡ : complex modulus

– G' : storage modulus

– G" : loss modulus

– Gly : Glycine

– I : ionic strength

– ISH : In sity Hybridization

– KCl : potassium chloride

– L : capillary length

– λ : relaxation time

– λ<sub>r</sub> : retardation time

– M : torque

– MPa : Mega Pascal (Pressure unit)
- mRNA: Messenger RNA

- myoD : myogenic differentiation gene

– N: Newton (Force unit)

– N₁, N₂ : normal stress differences

– ∇ : upper-convected derivative

– NaCl : sodium chloride

– Na₂HPO₄ : sodium phosphate

– nDNA Compacted DNA

– ω : angular frequency

– Ω : angular velocity

– P: pressure

– P (length): persistence length

– PBS : Phosphate Buffer Saline

– PCR Polymerase chain reaction

– Pe : Peclet number

– PFA : paraformaldehyde

– φ : orientation in the image plane (SHG)

– pI : isoelectric point

– ψ₁, ψ₂ : normal stress difference coefficients

– P-SHG: Polarization-resolved SHG

– PTFE: Polytetrafluoroethylene, Teflon

– Q: throughput

– q-PCR : quantitative PCR

– R: capillary radius

– Re : Reynolds number

– R<sub>H</sub> : hydraulic resistance

– ρ : bulk density

– ρ (SHG) : anisotropy parameter

– RNA : Ribonucleic acid

– RNase : Ribonuclease

– RT : Reverse Transcription

– σ : shear stress

– Scx: scleraxis gene

– SEM : Scanning Electron Microscopy

– SHG : Second Harmonic Generation

– S<sub>p</sub>: piston surface

– t: time

– τ : stress tensor

– TE: Tissue Engineering

– TEM : Transmission Electron Microscopy

– Tnmd: tenomodulin gene

– T<sub>m</sub> : melting temperature

– UTS: Ultimate Tensile Strength

– V: volume

– v<sub>p</sub>: piston velocity

– $\bar{v}$ : output velocity

– Wi : Weissenberg number
# Appendix
REVIEW ARTICLE

# From Tendon Injury to Collagen-Based Tendon Regeneration: Overview and Recent Advances

Clément Rieu<sup>a,†</sup>, Lise Picaut<sup>a,b,†</sup>, Gervaise Mosser<sup>a</sup> and Léa Trichet<sup>a,*</sup>

<sup>a</sup>Sorbonne Universités, UPMC Université Paris 06, CNRS, Collège de France, Laboratoire de Chimie de la Matière Condensée de Paris, F-75005, France; <sup>b</sup>Sorbonne Universités, UPMC Université Paris 06, CNRS UMR 7588, Institut des NanoSciences de Paris, F-75005, Paris, France

ARTICLE HISTORY

Received: March 24, 2017

Accepted: May 5, 2017

DOI:

10.2174/1381612823666170516130515

**Keywords:** Tendon, injury, repair, biomaterial, graft, collagen, scaffold.

**Abstract:** Tendon injury is a clinical, societal and economical issue. Moreover, tendon repair represents an important clinical challenge, partly due to the mechanical constraints that occur at the junctions with muscle and bone. Several strategies have been developed for tendon repair. In this review, we first assess the importance of tendon injuries from different sites and their causes. After a short overview of tendon three-dimensional organization, the complexity of the perfect repair quest is presented ranging from current clinical procedures to new engineering scaffolds. We then sum up tendon engineering requirements and focus on new collagen-based scaffolds, which raise promising prospects to mimic and repair tendon. In particular, we survey quantitatively a large panel of techniques to produce these scaffolds, detailing their principle and recent improvements.

# 1. INTRODUCTION: CLINICAL CHALLENGES

Tendons can present acute or chronic injuries caused by extrinsic or intrinsic factors, either alone or in combination (Fig. 1) [1]. Tendinopathies refer to the various chronic conditions that can affect tendons. They can remain silent over a large period of time, cause chronic pain and, potentially, lead to tendon tears and ruptures. These can also occur following lacerations by a sharp object or a tensile overload, in which case, it is now believed that an underlying tendinopathy is also involved [2], [3]. The term tendinitis, previously used, referred to tendon micro-tears associated with acute injuries due to an overload, causing inflammation, while tendinosis accounted for chronic wound following tendon overuse. The term tendinopathies is now preferred because there is no assumption about the underlying pathology, especially since the role of inflammation is still under debate [4], [5]. They are associated with a change in remodeling activity of tendon matrix.

External causes of tendinopathies include occupational, sporting activities, and prescription drugs [6]. Intrinsic factors include age and biomechanical imbalance [1]. The etiology of tendinopathies often appears to be multifactorial [2], [3], [7].

A UK study, performed in an acute Orthopedic Unit treating a well-defined catchment population of about 535, 000 and prospectively recording the demographic details over 5 years, provided results about the incidence of musculoskeletal tissue trauma [8]. One of the findings is that about two thirds of tendon injuries relate to the hand. Among them, forearm/hand extensor tendon injuries occur with an annual incidence of 17.87/100, 000, and mallet finger involving extensor tendon affecting mobility of the last phalanx, at a rate of 9.89/100, 000 (Fig. 1). A 10-year population-based study conducted in the US, examining hand or wrist open wounds secondary to acute trauma and involving tendons, found an incidence rate

of 33.2 tendon injuries per 100, 000 person-years, with work-related injuries accounting for 24.9% of total [9].

Achilles tendon rupture occurs at an annual rate of 11.33/100, 000 according to Clayton et al. (2008) [8] and has already been reported with rates ranging from 5.5 to 18/100, 000 (Fig. 1) [10]. Most acute ruptures concern people in their thirties and forties during sports activities. Achilles tendinopathy has a 5.9% lifetime cumulative incidence among sedentary people as compared to 50% among endurance athletes [11]. Tennis elbow, also known as lateral elbow tendinopathy, is secondary to excessive use of wrist extenders and forearm supinators. It is usually reported to affect 1 to 3% of the population each year, however, a descriptive epidemiology study shows a decrease in the incidence with a rate of 3.4 per 1, 000 [12]. The injury incidence in tennis players has been reported in proportions between 9 and 35% [13]. The rotator cuff, which is composed of four muscle-tendon units that originate on the scapula and cover the top of the humeral head, can also often be subject to tendinopathy. In a prospective study, the sex and age-standardized incidence of shoulder pain was evaluated to be 9.5/1, 000 among the patients presenting to general practice, with a proportion of 85% related to rotator cuff tendinopathy [14]. The incidence of rotator cuff tears was reported to be 3.73/100, 000 [8].

Work-related musculo-skeletal disorders concerning tendons are evident in occupations involving prolonged periods of repetitive, static work [14]. The prevalence of work related upper limb musculoskeletal disorders in UK was estimated to be 730 per 100, 000 people in year 2014/2015 [15]. There is an average of 17.7 working days lost for each case. According to the 2005 data collection of Eurostat figures on recognized occupational diseases (EODS) of 12 member states, the most common musculo-skeletal occupational disease was lateral elbow tendinopathy (14, 155 cases) [14]. In the United States, shoulder injury has been found to be the second most common cause after back pain for time away from work in manual laborers [16]. 41.5% of occupational shoulder injuries required more than 31 days away from work in 2007. In Sweden in 1995, among the 9, 398 reported work-related soft tissue diseases, 217 cases involved rotator cuff and 538 lateral elbow [17].

<footnote>*Address correspondence to this author at the Laboratoire de Chimie de la Matière Condensée de Paris, UPMC, Tour 43-44, Et. 4, 4 place Jussieu, 75252 Paris Cedex 05, France; Tel/Fax: +33-1-44276553, +33-1-44278102; E-mail: lea.trichet@upmc.fr</footnote>

<footnote>†These authors contributed equally to this work.</footnote>
<figure><img src="image_160.png" /><figcaption>**Fig. (1).** Location, annual incidence and main cause of most frequent tendon injuries. Data are referenced in the text. Sources: [8], [10]-[12]. Body image, courteously provided by Servier Medical Art under license https://creativecommons.org/licenses/by/3.0/fr/, was modified for clarity purpose.</figcaption></figure>

Some estimates of the cost of work-related upper limb and neck musculo-skeletal disorders are between 0.5 and 2% of gross national product [17].

Due to poor vascularization and low cellularity of the tissue, there is limited spontaneous regeneration of tendon. Treatments can be conservative or surgical, or a combination of both with post-intervention rehabilitation programs [2], [5], [10], [16]. Currently there is a high rate of re-injury. In the case of small and large rotator cuff tears, surgical repairs of ruptured tendons have re-tear rates of up to 35 and 94%, respectively [18]. Clinicians assess an unsatisfactory clinical outcome after suture for 25% of patients with hand tendon lacerations, and 7.7% of these repairing re-rupture [19].

## 2. TENDON STRUCTURE AND REMAINING QUESTIONS

Tendons and ligaments belong to connective tissues that present a cellular part and an extra-cellular matrix (ECM). Cells, out of which tenocytes are predominant, are found immersed into the ECM. They synthesize the bio-components and thus play a major role in the setting and renewing of the ECM. This last one is responsible for the structural and mechanical properties, and maintains the integrity of the tissue [20]. ECM is composed of many different biochemical components such as collagens, non-collagen proteins, and proteoglycans [21][22].

Collagens are widespread and play major roles in organisms. They constitute a family of 28 members [23], all being formed by three polypeptides chains. Each chain has repeating amino-acid triplet units -(Glycine-XY)n-, that extend variably depending on the family member, and with X and Y positions frequently occupied by proline and hydroxyproline. The presence of those -(Glycine-XY)n-, repeats leads to polyproline-II segments that supercoil into a right-handed helix. Each member of the family has different functions

and members can be, or not, co-localized, depending on the connec-tive tissue considered. In tendon, collagens reach a concentration of 30-31.6% of the wet weight [24], and 60-85% of the dry weight of the tissue [22]. In the collagen family, collagen I is the major struc-turing bio-component of the tendon, where it is found in its fibril form and arranged into a hierarchical manner as we will see later.

Among the other components found in the ECM, some regulate collagen fibril diameters and/or fibrils interdistance and mechanical properties. For instance, polysaccharides form bridges between fibrils maintaining the equidistance between them and contribute in holding the three-dimensional network. Those bridges are formed by small carbohydrate chains of about 15 kDa [25]. *In vitro*, it was also reported that, in the synthesis of collagen scaffolds, the addition of decorin increased their mechanical properties.

Tendon structure has been studied for the past decades and by various techniques and under various states. Despite the information acquired, the exact organization of tendon in living tissues remains unclear. The fact that it may be adapted to each location within the body and depend on each species, does not help in getting a precise idea [26]. New developments in microscopy will hopefully bring some valuable keys for this understanding [27]-[29]. Anyhow, some general characteristics are common to all tendons and ligaments and the most important are discussed below.

<figure><img src="image_161.png" /><figcaption>**Fig. (2).** Schemes illustrating the major structural characteristics of extracellular matrix of tendon: (A) Hierarchical organization of tendon unit. (B) Helical and crimp organization of fascicles (resp. top and bottom).</figcaption></figure>

Tendons and ligaments are organized in a hierarchical manner that starts with collagen I molecules, which, associated by five into “microfibrils”, pack into fibrils [30]. The fibrils are further associated into fibers that are joined into bundles delineated by the endotenon, a fine ECM of another composition and organization. The endotenon is continuous with the surrounding connective tissues and facilitates sliding between adjacent structures. In particular, it is in continuity with the epitendon that further tights together bundles of fibers into fascicles. Depending on the tendon, fascicles may present another covering layer named “paratenon” (Figure 2A). Finally, several fascicles may be bound together to give larger units [31].
<figure><img src="image_162.png" /></figure>

**Fig. (3). Scheme conciliating helix crimp and linear organization:** A single stranded helix is schematized on top, a partial super-helix in the middle and a full super-helix on bottom. Boxed by the dotted line, their appearance when observed between crossed-polarizers. Crossed-arrows indicate directions of polarizers. (A) **Top:** The extended helical strand and beneath the projected collagen molecular orientations are represented by lines. All orientations are nearly aligned along the helix axis. **Middle:** The partial super-helix is represented and the molecular orientations being all nearly aligned along one polarizer, the crimp is hardly seen and the partial super-helix appears nearly dark. **Bottom:** the symmetry of the complete super-helix is cylindrical. The mean projection of the molecular directions sums up along the axis of the helix. The complete super-helix appears dark under crossed-polarizers. (B) **Top:** The molecular orientations vary along the axis of the helix and present a repeating unit of alternating orientations. **Middle:** The partial super-helix is represented and the molecular orientations varying as illustrated above, the crimp is clearly seen as banded pattern under crossed polarizers. **Bottom:** the symmetry of the complete super-helix remains cylindrical with the mean projection of the molecular directions summing up along the axis of the helix. The complete super-helix still appears dark under crossed-polarizers.

Collagen fibrils within endotenon, epitendon and paratenon are not necessarily collinear to the tendon and can be found perpendicularly oriented [32]. Those layers possibly maintain the integrity at each level from bundles to tendon. They also possibly take part in its mechanical properties. Blood and lymphatic vessels, as well as nerves, are found in the endotenon, epitendon and paratenon.

Cross-polarized optical microscopic observations of excised rat tail tendons reveal an overall alignment of the bundles along the tendon axis. They also reveal units, from 10 to 100 µm, with alternating oblique orientations compared to the tendon axis. This feature is referred to as “crimp” (Fig. 2B bottom). Evidence of this crimp was further observed by atomic force and electron microscopies (scanning and transmission) [33][34]. The crimp was suggested to be a natural shock-absorber and to play a role in elastic recoil [35].

Shah *et al.* (1982) reported that the crimp was only observed at day 16th during the embryonic development of chick tarsus metatarsus region. The appearance of the crimp correlated with a high increase of collagen content between day 14th and day 16th [36]. This suggests the crimp to be highly related to the presence of collagen fibrils. On the other hand, for Herchenhan *et al.*, (2012) tenocytes, cells found inside tendon, induce formation of crimps [37]. Their conclusion relies on the fact that, knocking down cells using polyoxyethylene (9-10) *p*-t-octyl phénol (TX-100) within a biomaterial construct presenting a crimp, induces its loss. However, the authors do not take into account the fact that TX-100 also interacts with collagen and even destabilizes fibrils, possibly contributing to the loss of the crimp [38][39].

In addition to the overall alignment and the crimp, Vidal et al. (2003) analyzed under crossed polarizers transversal sections of tendons and showed that fascicles presented a helical organization along the tendon axis [40] (Fig. 2B top). This raises the question on how the crimp structure can be compatible with both a cylindrical symmetry and a helical organization. Geometrically speaking, one way to explain the crimp would be as over-winding of a helical organization. *In vivo*, under tension, it would be extended and fascicles would appear quasi collinear with the direction of the tendon. No “crimp” would be detected (Fig. 3A). However, when tendons

are relaxed and/or excised, the linear tension would decrease and the over-winding would collapse on itself while still being constrained transversally by the 90° transverse collagen sheath. The crimp would thus appear (Fig. 3B).

Mechanical tests made on tendon clearly show the difference of mechanical behaviors during the elongation of the crimp and after [35]. Computational modeling best fits the mechanical experimental data when a helical arrangement is injected into the models [41]. Reese *et al.* (2010) further improved the fit by combining a wavy pattern with a super-helical arrangement [42]. This data comforts the idea of a super-helical structure.

It has been shown that super-twisted organizations could be spontaneously generated in liquid-crystals [43]. Knowing that collagen forms different liquid crystal phases [44][45][46] that are subject to physico-chemical parameters [47], it is not excluded, although there is no proof yet, that the over-winding helical organization of tendon may also result from simple physical-chemical laws. Most probably, cells and physico-chemistry work hand in hand.

Modifications upon aging, injury and recovery have been studied at different levels: biochemical, fibrils (diameter, orientation) and crimp (presence or not, period) [21][48]. Here, again, it seems that there is not a unique scheme of evolution with different observations being made depending on tissues and/or species. To further highlight the complexity of the living tissues, their studies and the different routes to achieve their healing, reconstruction or replacement, it should be noted that control tendons, which are generally contralateral tendons, also see their characteristics being modified, possibly due to stress and postural changes during healing process.

Tendon wound healing process is characterized by three different stages: inflammation, repair, and remodeling [5], [49]. After injury there is a high release in growth factors, cytokines, and an increase in the expression of tendon-associated molecules, which cellular origins are still difficult to trace back [50]. During the repair phase, collagen type III is mainly synthesized, leading to the formation of a scar tissue with increased cross section area. During matrix remodeling, which can last months to years, collagen type I synthesis is increased and the newly formed fibers are aligned along
the longitudinal axis of the tendon. The tissue obtained often displays reduced mechanical properties due to the presence of disorganized collagen, as well as improper non-collagenous molecules and exogenous cells recruited at the wound site. These cells can lead to the formation of fibrous peritendinous adhesions with the surrounding tissue, impairing the normal stress transmission and hence the restoration of normally aligned collagen fibers [1].

The ultimate tendon repair is based on several requirements. The prior physiological criteria include, among others, the initial strength recovery, the joint gliding, the re-rupture avoidance and the shorter return-time to activity. Early active mobilization of the repaired tendon is advocated to improve those properties [19]. It was also proved that this early post-surgery stimulation increases the collagen type I cell production and that the quick tension restoration induces the parallel alignment of this renewed collagen [51]. In the following part, the different strategies clinically used from sutures to decellularized matrices will be described.

### 3. CURRENT CLINICAL STRATEGIES

The route chosen to induce healing of the tendon must take into account the gravity of the injury. From chronic tendon injury to tendon tears, the first objective is to reduce the patient's pain due to the tissue inflammation with specific drugs and with rehabilitation exercise [52], [53]. Then, emerging treatments consist in delivering growth factors through autologous whole blood or platelet rich plasma (PRP). Although *in vitro* studies show that these injections impact favorably the collagen production and degradation mechanisms [54], their use in clinical cases are controversial. Indeed, several questions regarding the injection frequency, volume, concentration etc. still remain open.

For acute tendon injury, the treatment depends on the lesion severity. If the gap between the two tendon extremities is less than 3 mm, the self-healing mechanism is engaged [55]. For larger injuries, if the joint motion is altered or even completely lost then surgery is needed. The length of the gap will determine the approach to be adopted. If the gap is less than 3 cm, suture techniques are required to rejoin the tendon. In the case of Achilles tendon, the conciliation of its 2 ends requires a great force. This is why, if the rupture is in an easy access location, surgeons use some devices as bone clamp [56] or Integra™ Achillon® system [57], [58], to join the tendon and then suture it. It should be kept in mind that the more handled is the injured tendon the more severe are the tissue trauma and adhesion. Wang *et al.* (2006) have even shown in a murine model that the suture directly affects the cellular healing process. The acellular zone does not heal and present inflammatory signs [59].

For gaps larger than 3 cm, augmentation of the ruptured tendon with additional implants or grafts is recommended [56]. Several types of devices from metallic implants to synthetic grafts are commercially available and already used by clinicians (Table 1, [60]).

#### 3.1. Sutures

Three methods of tendon suturing have been developed and optimized: the non-grasping, the grasping and the locking anchors. The non-grasping method was first performed on finger flexor tendon by Bunnell in 1918 [19] (Fig. 4A). This technique consists in looping the anchor around the epitenon. As a consequence, no collagen fibers bundles are injured. On the contrary, the methods based on grasping, (opened loop, Fig. 4B) and locking anchors (closed loop, Fig. 4C) pinch the collagen fibers.

However, those two techniques bring better mechanical properties and gapping resistance to the tendon. They took the name of their authors Kessler for the grasping suture and Krakow for the locking one. The last one is assumed to be the more effective and is continuously being improved [61].

<figure><img src="image_163.png" /><figcaption>**Fig. (4).** Scheme of the three sutures commonly used to repair tendon (Achilles tendon example in back view). (A): the non-grasping suture (Bunnell) consisting in looping the anchor around the epitenon. (B): the grasping suture (Kessler) based on 4 opened loops, which pinch collagen fibers. (C): The locking suture (Krackow) made of several closed loops.</figcaption></figure>

Several fibers made of polyester (Ethibond®), nylon, polypropylene or a combination of ethylene glycol and terephthalic acid (Dacron®) are already clinically used for suture alone. They are non-resorbable as they have to maintain more than one year the end-to-end tendon extremities [19]. Other materials have been developed for suture augmentation, which is a different approach. Schliemman *et al.* (2015) compared the reconstruction of a tendon with a polydioxanone (PDS) or cable wire augmentation and with suture anchor repair alone. They showed that suture augmentation in both cases provides less elongation and higher loads (almost a 1.5 factor) than the simple suture [62]. A negative aspect of cable wires is that a second surgery is needed to remove them.

Alternative methods have been proposed to avoid the use of sutures. Hohendorff *et al.* (2008) compared post-surgery results of Achilles tendon rupture treated either with fibrin glue or sutures. They noted that both strategies reached almost the same results. Moreover, fewer complications were observed for the fibrin glue cohort [63]. Rose *et al.* (2014) also worked on the biological tissue gluing. They used solutions of silica nanoparticles (NPs) to glue mismatched tissue. They tested it *in vitro* on a piece of calf liver with a created gap of 2 cm. They glued the two ends and kept it under finger pressure during 30 s [64]. They observed that the two parts of the liver remained attached and that the glue did not affect the rigidity or the permeability of the tissue. This second approach is very promising as it is quite simple and fast to implement. It must be tested on tendon to know if the bonds created with the silica NPs are strong enough to support physiological loads as applied in Achilles tendon sites for months *in vivo*.

#### 3.2. Metallic implants

Some other researchers have developed metallic implants to achieve tendon repair with the best mechanical properties.

Erol *et al.* (2007) developed an implant made of stainless steel wires with different spiral-shaped structures on sheep cadaveric Achilles tendons [65]. Briefly, the prostheses are composed of a coiled wire head, which hooks the tendon, and a wire passing through it to join the head to the other tendon end or bone. They also compared their implant with classical sutures. In some cases, they observed the breaking of the tendon-implant holding point and the fibering of tendon's ends, raising the problem of the material elastic properties. If the material is too stiff compared to the tendon, the rupture of the anchor will occur. To have the best mechanical properties, it is essential to use materials with the same range of elastic modulus as tendon. In their preliminary study, they obtained good mechanical results compared to sutures alone. But authors advised to use this kind of device only for bone-tendon repair. Ağır *et al.* (2014) also investigated this kind of metallic implant. In their
<table><caption>**Table 1.** Commercially available grafts or implants for tendon repair. The devices written in bold italic are detailed in sections 3.2 and 3.3.</caption><thead><tr><th>Type</th><th>Name</th><th>Source</th><th>Company</th><th>Refs.</th></tr></thead><tbody><tr><td rowspan="10">Metallic implants</td><td>Teno Fix®</td><td>Stainless steel</td><td>Ortheon Medical, Winter Park (USA)</td><td>[Wolfe 2007]</td></tr><tr><td>Tenolig®</td><td>Stainless steel (with Dacron fibers)</td><td>FH Orthopedics, (FR)</td><td>[Lacoste 2014]</td></tr><tr><td>Bio-Blanket®</td><td>Bovine dermis</td><td>Kensey Nash Corp. (USA)</td><td rowspan="3">[Gomes 2015]</td></tr><tr><td>Conexa®</td><td>Porcine dermis</td><td>Tornier/Wright Medicals (TN)</td></tr><tr><td>CuffPatch®</td><td>Porcine SIS</td><td>Organogenesis (MA, USA)<br>Licensed to Arthrotek (IN, USA)</td></tr><tr><td>Inforce® Reinforcement matrix</td><td>Porcine dermis</td><td>IntegraLife Sciences (NJ USA)</td><td>[Hollawell 2015]</td></tr><tr><td>OrthADAPT®</td><td>Equine pericardium</td><td>Pegasus Biologic Inc. (USA)</td><td>[Gomes 2015]</td></tr><tr><td>Restore™</td><td>Porcine SIS</td><td>DePuy Orthopedics (IN, USA)</td><td>[Phipatanakul 2009]</td></tr><tr><td>Shelhigh No-React® Encuff Patch</td><td>Bovine or porcine pericardium</td><td>Shelhigh Inc. (NJ, USA)</td><td>[Gomes 2015]</td></tr><tr><td>TissueMend®</td><td>Bovine fetal dermis</td><td>TEI Biosciences (MA, USA) licensed to Stryker<br>Orthopedics (NJ, USA)</td><td>[Song 2010]</td></tr><tr><td>Trellis™ Collagen Ribbon/Bio-tape XM</td><td>Porcine dermis</td><td>Wright Medical Technology, Inc. (USA)</td><td>[Berlet 2014]/[Gomes 2015]</td></tr><tr><td>Zimmer® or Permacol™</td><td>Porcine dermis</td><td>Tissue Science Laboratories (MA, USA)<br>licensed to Zimmer (IN, USA)</td><td>[Gomes 2015]</td></tr><tr><td rowspan="4">Xenografts</td><td>Arthroflex®</td><td>Human dermis</td><td>Arthrex, LifeNet Health Inc. (USA)</td><td>[Gomes 2015]</td></tr><tr><td>GraftJacket®</td><td>Human dermis</td><td>Wright Medical Technology (USA)</td><td>[Song 2010, Lee 2008]</td></tr><tr><td>Matrix™ HD</td><td>Human dermis</td><td>RTI Surgical (USA)</td><td rowspan="2">[Gomes 2015]</td></tr><tr><td>TenSIX™</td><td>Human dermis</td><td>Solana, Wright Medical (USA)</td></tr><tr><td rowspan="10">Synthetic grafts</td><td>Artelon® and Sportmesh™</td><td>(Polyurethane urea) polymer</td><td>Artimplant AB (SWE)<br>Biomet Sports Medicine (IN, USA)</td><td>[Schoaib 2016]</td></tr><tr><td>Biomerix™ RCR Patch</td><td>Polycarbonate (polyurethane urea)<br>with polyester fibers</td><td>Biomerix (USA)</td><td>[Gomes 2015]</td></tr><tr><td>Dacron® mesh</td><td>Polyethylene terephthalate (PET)</td><td></td><td></td></tr><tr><td>Gore-Tex®</td><td>Polytetrafluoroethylene</td><td>WL Gore and Associates (USA)</td><td></td></tr><tr><td>LARS ligament</td><td>PET</td><td>LARS (FR)</td><td></td></tr><tr><td>Leeds-Keio® or Poly-tape®</td><td>Polyester ethylene terephthalate</td><td>Xiros plc, Neoligaments (UK)<br>Yufu Itonaga Co. Ltd (JPN)</td><td></td></tr><tr><td>Marlex® mesh</td><td>Polyethylene polymer</td><td>C R Bard (USA)</td><td>[Morrey 2016]</td></tr><tr><td>Orthocoupler™</td><td>PET</td><td>bioLOGIC Corp., Surgical Energetics LLC<br>(USA)</td><td>[Melvin 2010]</td></tr><tr><td>Ortho-Tape®</td><td>PET</td><td>Neoligaments (UK)</td><td>[Abdullah 2015]</td></tr><tr><td>Sericuff™</td><td>Silkworm silk</td><td>Serica, Allergan (USA)</td><td rowspan="2">[Gomes 2015]</td></tr><tr><td>STR GRAFT™</td><td>Poly-L-Lactide</td><td>Soft Tissue Regeneration Inc. (USA)</td></tr><tr><td>X-repair™ X-repair-SL</td><td>Poly-L-Lactide</td><td>Synthasome (USA)</td><td></td></tr></tbody></table>
study, they used stainless steel wires but designed it with several loops [66]. The wire is applied from outside and allows the transversal anchoring of metallic pins in it. Here again, some limitations of the device are explained: these implants cannot be used for hand injuries and are more suitable for larger body tendons as Achilles tendons.

Metallic devices such as the Teno Fix® [67] or Tenolig® [68] are already commercially available. They were tested on cadaveric tendons and on patients respectively. Teno Fix® was compared with conventional sutures and no biomechanical improvement was shown switching from one to the other. Tenolig® exhibits different advantages such as the use of percutaneous surgery for the device positioning with ultrasonography monitoring to target the two Achilles tendon ends, very small wound scars and complications occurring only for 1 over 75 patients. The return time to sport was estimated to be 9 months. This kind of metallic implant is not largely used in clinical practice, as studies evaluating their effective improvement compared to classic sutures are still rare. Moreover, several limitations such as tearing of the site repair have been highlighted as mentioned above. Extensive improvements of these techniques are needed before becoming a common and viable strategy.

If the tendon injury has been neglected or if the site of injury has a poor quality due to tendinopathy, sutures or implants are not sufficient anymore to repair it [69]. Furthermore, to avoid the re-rupture, a reinforcement of the site is needed. To meet these objectives, grafts can be used. The tendon augmentation with grafts consists in adding a matrix either around the defect and then suturing it (onlay surgery), or in the ruptured tendon and then closing it with suture inside (inlay surgery). The combination of the classic sutures, which join the two tendon extremities, with the graft, may improve tendon repair. The grafts can come from different origins: the patient himself (autograft), another human being (allograft), animals (xenograft) or from synthetic materials. These different grafts are successively described in the following part.

### 3.3. Grafts

#### Autografts.

Sophisticated surgical procedures have been elaborated to overcome large tendon defects with patient's own "material". Three kinds of procedures more or less invasive allow accessing the tendon: open, mid-open and percutaneous. Beskin *et al.* (1987) compared four "open" techniques on 42 patients suffering from Achilles tendon rupture: Bunnell or Kessler sutures, "three bundle", and augmentation, either with the plantaris or with the peroneus brevis tendon [70]. These last two methods consist in using local tendon transfer to reinforce the Achilles tendon. The four techniques show promising results as 85 to 92% of patients return to normal long-term activity after the surgery. However, in each case, a single long longitudinal excision is performed which is very invasive, urging some clinicians to develop less invasive techniques. Carmont *et al.* (2007) also used the peroneus brevis tendon to reconstruct Achilles tendon [71]. Contrary to Beskin *et al.*, they did two para-midline incisions. Skin wounds are clearly smaller and skin integrity is preserved. Moreover, this allows preserving the surrounded tissues (tendon sheath) and may improve tendon's healing. To follow through, other autologous sources for reconstruction can be recruited. Ebied *et al.* (2016) repaired Achilles tendons thanks to free semi-tendinous tendon grafts from upper muscles [72]. Shewy *et al.* (2009) used two intratendinous flaps from the proximal gastrocnemius-soleus complex [73]. Briefly, they created two flaps in the musculo-tendinous junction and passed them through the Achilles tendon gap. They fixed the flaps extremities to the calcaneus (bone). Few complications were observed as wound gapping or infection, and long pain after surgery. However, this technique showed good results as patients recovered their pre-injury sporting activity level.

The main advantage of these two surgeries is the use of highly vascularized tissues in a site close to the injured tendon, but less invasive incisions are recommended for the future. Autografts present several key strengths such as absence of foreign body response, conservation of native mechanical properties and vascularization continuity. The main inconvenience is the risk of damaging the donor site tissue and thus creating new lesions. Moreover, the neotendon reconstructed with surrounding tendons or muscles parts is thicker than the native one. This can lead to unsightly aspect repair and to patient discomfort on the repair site.

Other strategies to repair injured tendons have been developed to avoid creating more lesions in patients. Tendon augmentation with xeno- or allograft may be good candidates for this purpose.

#### Xenо- and allografts

Xenografts are tissues or organs coming from other species that are pre-decellularized. Porcine decellularized small intestine submucosa (SIS) is the most commonly used xenograft. Several clinical studies have been performed to assess their effectiveness for tendon repair augmentation. Phipatanakul *et al.* (2009) used SIS (Restore™) for rotator cuff repair [74]. As repairs were not convincing and as some complications (infections, skin reactions) were observed, authors did not recommend SIS for such repair augmentation. Sclamberg *et al.* (2004) came also to the same conclusion for this repair area as 5 over 11 patients presented worse results after the surgery [75]. Clinicians are quite unanimous on this issue. However, SIS may be adapted for tendons in other locations such as the foot or ankle as only 6% of the patients showed complications such as re-tear or patch failure.

Other sources of xenografts are also used. Berlet *et al.* (2014) tested the Trellis™ device (a decellularized, freeze dried and sterilized porcine dermis), in a ribbon form on human cadaveric feet [76]. They created tendon defects greater than 5 cm. The tensile testing showed that in most cases, the suture failed first. As the tendon ends are not joined anymore, the load is only borne by the graft. Thus, in comparison with suture alone, the Trellis™ augmentation enhanced strength of Achilles tendon repair.

Other strategies, based on allografts, have been developed to overcome some rejection or infection reactions. Allografts are ECMs mostly coming from human dermis or SIS. Song *et al.* (2010) compared two commercially available devices on a cadaveric sheep model [77]. The first one is a xenograft named TissueMend®, which is derived from fetal bovine dermis and the second one, an allograft named GraftJacket® from human skin. The authors remind that grafts of dermis origin have better mechanical properties than matrices derived from SIS that were previously described. TissueMend® and GrafJacket® presented almost the same strength. They both improved tendon repair compared to suture alone. Lee *et al.* (2008) also used GraftJacket® matrix to augment tendon repair [78]. Clinical tests on Achilles tendon rupture were performed. At 3 months, patients recovered full range of motion and no re-rupture happened. The return-time to activity was about 12 weeks, which is quite encouraging.

Rather than choosing between xenografts and allografts, Hollawell *et al.* (2015) decided to combine them for Achilles tendon repair [79]. They used the allograft to span the gap between the two tendon ends and the xenograft (Inforce® Reinforcement Matrix) to secure the repair site. This graft combination improved the mechanical properties and physiological function, and patient pain was reduced. Some further investigations of this new strategy must be performed to better quantify its benefits.

Those acellular matrices either from animal or human origin show great advantages, overcoming the host reaction concerns and leading to good mechanical properties. Herbert *et al.* (2015) evaluated different decellularization processes needed to obtain such materials [80]. They studied the effect of using acetone or chloroform as fat agent reduction and antibiotics or peracetic acid as
biodecontaminants. They showed that the tensile strength and Young modulus of acellular porcine super flexor tendons are in the same range as the native one before the decellularization. However, this process altered the matrices properties at the low strains in the toe region where the crimp should play its shock absorber role. A potential way to overcome this problem is to reproduce this wavy structure. Branch *et al.* (2011) adopted this strategy with allograft for Achilles tendon repair [81]. They cut a human dermal acellular mesh in a z-wave shape, which finally looked like an accordion. They performed inlay surgery with a “strip and shoelace” weave. No re-rupture occurred and only one over the 3 case reports presented adhesion with the surrounding tissues.

Regarding the loss of cellular material due to the decellularization process, cell-based therapy combined with allografts emerged. Gaspar *et al.* (2015) reviewed the different cell strategies [82]. Before being implanted on the tendon injury sites, grafts can be sewn with cell suspension to enhance tendon regeneration. They highlighted the fact that the cell type chosen is crucial. Indeed, tenocytes or dermal fibroblasts are already differentiated, but their available quantity is limited. Progenitor stem cells overcome this issue, but can lead to bone ectopic formation. Finally, induced pluripotent stem cells may also be used on grafts. However, they raised the safety and efficiency problems. Güngormüs *et al.* (2016) used tenocytes and seeded them on decellularized rat tendons [83]. Finally, they implanted the cell-seeded scaffold in a rat model. No immune reaction occurred and their other results suggested that this scaffold improve mechanical properties of the repair.

Even if xeno- and allografts present some advantages, the matrix alteration due to the decellularization process is a problematic issue as it alters mechanical properties and leads to antigenic response [84]. Since it is essential to augment tendon with grafts having similar range of strength, synthetic grafts have been developed and are currently used in clinical practice.

## Synthetic grafts

Two categories of synthetic grafts are used to augment tendon. The non-resorbable grafts persist in the body and act as a substitute to overcome the lack of tissue. On the contrary, resorbable grafts are degraded by the body and support the fibrous ingrowth [85].

Non-resorbable grafts are commonly made of synthetic polymers such as polyester (Dacron®), polytetrafluoroethylene (PTFE or Teflon®) [86], and silicone rubber [87]. Such materials are already clinically used. Morrey *et al.* (2016) evaluated non resorbable polypropylene mesh commercially available (Marlex®) on 8 quadriceps tendons ruptures (knee) [88]. No re-rupture occurrence and good range of motion were the main advantages of this synthetic augmentation. As they tested it on a small cohort, with no mechanical evaluation, further studies are needed. Abdullah *et al.* (2015) evaluated the use of polyethylene terapthalate (PET) mesh called Ortho-Tape® [89]. They explored this new device on two patients who needed hand surgery because of massive tendon loss. Promising physiological results were shown. Melvin *et al.* (2010) developed their own device Orthocoupler®, made of 32 bundles of thousand PET fibers [90]. They implanted it in a goat model. Sixty days after the surgery failure forces for the specimens with the synthetic graft were 2.5 times higher than for sutures in the control group. Sutures were too fragile compared to the Orthocoupler® which is anchored and sutured in the muscle reinforcing the musculo-tendinous junction. It allows reinforcing the musculo-tendinous junction.

Finally, another criteria is the prevention of the tendon sheath adhesion. Non-resorbable grafts made of Teflon seem to be the appropriate candidates. Indeed, Williams *et al.* (1964) tested them on a dog model and compared them with autologous grafts [86]. Adhesions with the surrounding tissues were clearly observed for autogenous grafts, whereas Teflon graft surfaces glided and allowed

free joint motion. However, they reported some surgery problems as the solid anchoring of the Teflon grafts on tendon.

Non-resorbable grafts present some advantages such as their biocompatibility, their strength and the absence of adhesion with the tendon site. Nevertheless, they cannot be degraded, and are generally complicated to implement.

Resorbable grafts may overcome some of those issues. They can be composed of carbon ([91], [92]), polyglactin [93], polyglycolic acid (Dexon®) [94], or polylactic acid [95]. In the 1970s, Jenkins *et al.* (1980) investigated flexible carbon implants for ligament or tendon injuries. Special affinity of the living tissue with carbon was observed [91]. Moreover, carbon based scaffolds enhance the mechanical properties of the tendon and improve the alignment of collagen neo matrix. However, fragmentation of this kind of material often occurs and may lead to some irritation of the joint. Finally, they should not to be used close to metallic implants as they can react together. To overcome these drawbacks, Alexander *et al.* (1983) tried to coat carbon implants with polylactic acid but could not reduce the fragmentation process [96].

In this way, Cao *et al.* (2001) used unwoven polyglycolic acid fibers, which they assembled in a cord shape and wrapped with acellular intestinal submucosa membrane [97]. Tenocytes were cultured on it. Then, they created a 2.5 cm long defect in the second digital flexor profundus tendon of hen and repaired it with the cell-scaffold. They harvested the specimens at 8, 12 and 14 weeks and performed histological observations. The longer the graft stayed in the body, the better were the biomechanical properties. The authors suggested that a minimum time was required for a good tissue re-modeling of the tendon. Histological results showed that the neotendon contained a large amount of tenocytes and presented its own structure, easy to distinguish from native tendon.

Synthetic polymer scaffolds can also be designed as a tissue mesh. Shoaib *et al.* (2016) performed surgical repair in 7 patients suffering from symptomatic chronic Achilles tendon rupture with a bio-absorbable polyurethane polymer device (Artelon®) [98]. Artelon® can be seen as the resorbable equivalent of the Ortho-Tape® previously described. They compared this synthetic tendon augmentation to the Krakow suture alone. Better results were shown with the Artelon® graft. The authors also suggested that this device could be used for defect larger than 5 cm. Finally, they recommended combining it with non-resorbable suture in order to secure the augmentation over the years.

We have seen that synthetic grafts show promising results for tendon repair and in particular for the mechanical properties. However, some concerns persist as they may lead to foreign body response. This is also the same for xeno-allografts, and in the case of autografts, new lesions are created that may lead to secondary surgery in case of local degradation. Clinical practices or devices to repair tendon get closer to the ultimate goal, but they often lack specific criteria as strength or body tolerance. This is why some new scaffolds based on biocomponents are currently being developed. Such engineering is not straight-forward and needs to meet several strict specifications that we will discuss below.

## 4. SCAFFOLD DESIGN SPECIFICATIONS FOR TENDON REGENERATION

Scaffolds for reconstruction must meet several requirements as tendons are such specific tissues, with multiscale order and high mechanical solicitations. These are dependent on the site aimed for reconstruction. Indeed, tendon size and shape can vary greatly, in function of the location in the body. Whereas rotator cuff and patellar tendons are sheet-like, length and width being in the same order and much greater than the thickness, Achilles tendons are rather unidimensional, in a rope-like manner shape [99]. Typical dimensions are given in Table 2. Mechanical properties can also vary according to sites and persons, with Young modulus ranging from