Synthesis of an artificial tendon

Lise Picaut

► To cite this version:

Lise Picaut. Synthesis of an artificial tendon. Physics [physics]. Université Pierre et Marie Curie, 2017. English. NNT: 2017PA066265 . tel-03990442

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Submitted on 15 Feb 2023

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Université Pierre et Marie Curie

École doctorale n° 397 : Physique et Chimie des Matériaux

Institut des Nanosciences de Paris

Laboratoire de Chimie et de La Matière Condensée de Paris

# Synthèse d'un tendon artificiel

présentée et soutenue publiquement par

Lise PICAUT

le 09 Octobre 2017

Directeur de thèse : **Tristan BAUMBERGER**

Co-directrice de thèse : **Gervaise MOSSER**

## Devant le jury composé de :

<table><tr><td>Cécile Ayako Dreiss,</td><td>Directrice de Recherche au King's College London, Royaume-Uni - IPS</td><td>Rapportrice</td></tr><tr><td>Charlotte Vendrely,</td><td>Maitre de Conférence HDR à l'Université de Cergy Pontoise, France ERRMECe</td><td>Rapportrice</td></tr><tr><td>Dominique Hourdet,</td><td>Professeur à l'Université Pierre et Marie Curie, France - SIMM</td><td>Examinateur</td></tr><tr><td>Sophie Asnacios,</td><td>Maitre de Conférence à l'Université Pierre et Marie Curie, France - MSC</td><td>Examinatrice</td></tr><tr><td>Fabien Le Grand,</td><td>Chargé de Recherche HDR CNRS, France - Institut de Myologie</td><td>Examinateur</td></tr></table>
# Contents

<table><thead><tr><th></th><th></th><th></th></tr></thead><tbody><tr><td>Contents</td><td></td><td>iii</td></tr><tr><td>INTRODUCTION</td><td></td><td>1</td></tr><tr><td>0.1 References</td><td></td><td>3</td></tr><tr><td>1 Literature review specific to tendon and its repair strategies</td><td></td><td>5</td></tr><tr><td>1.1 Tendon</td><td></td><td>6</td></tr><tr><td>1.1.1 Anatomy and structure</td><td></td><td>6</td></tr><tr><td>1.1.1.1 Human anatomy</td><td></td><td>6</td></tr><tr><td>1.1.1.2 Structure</td><td></td><td>6</td></tr><tr><td>1.1.2 Clinical situation</td><td></td><td>13</td></tr><tr><td>1.1.2.1 Pathologies and injuries description</td><td></td><td>13</td></tr><tr><td>1.1.2.2 Incidence</td><td></td><td>13</td></tr><tr><td>1.1.2.3 Current strategies</td><td></td><td>15</td></tr><tr><td>1.2 Tissue engineering for tendon repair</td><td></td><td>23</td></tr><tr><td>1.2.1 Scaffold design specifications</td><td></td><td>23</td></tr><tr><td>1.2.2 Presentation of the main TE techniques used for tendon repair</td><td></td><td>24</td></tr><tr><td>1.2.3 Synthetic polymer scaffolds</td><td></td><td>26</td></tr><tr><td>1.2.4 Scaffolds based on natural/organic components</td><td></td><td>27</td></tr><tr><td>1.3 Conclusion</td><td></td><td>31</td></tr><tr><td>1.4 References</td><td></td><td>33</td></tr><tr><td>2 Methods and materials</td><td></td><td>45</td></tr><tr><td>2.1 Methods</td><td></td><td>47</td></tr><tr><td>2.1.1 Extrusion set-up</td><td></td><td>47</td></tr><tr><td>2.1.1.1 Key constituent part: the needle</td><td></td><td>48</td></tr><tr><td>2.1.1.2 Set-up performances and limits</td><td></td><td>52</td></tr><tr><td>2.1.1.3 Conclusion</td><td></td><td>59</td></tr><tr><td>2.1.2 Rheological testing of polymer solutions</td><td></td><td>60</td></tr><tr><td>2.1.2.1 Physical basis</td><td></td><td>60</td></tr><tr><td>2.1.2.2 Rheometer geometries</td><td></td><td>67</td></tr><tr><td>2.1.3 Mechanical testing of extruded threads</td><td></td><td>69</td></tr><tr><td>2.2 Materials</td><td></td><td>71</td></tr><tr><td>2.2.1 Collagen solutions</td><td></td><td>71</td></tr><tr><td>2.2.1.1 Collagen type I and its fibrillogenesis</td><td></td><td>71</td></tr><tr><td>2.2.1.2 Chemical properties</td><td></td><td>73</td></tr><tr><td>2.2.1.3 Collagen solutions preparation</td><td></td><td>75</td></tr><tr><td>2.2.1.4 Collagen extraction and purification</td><td></td><td>75</td></tr><tr><td>2.2.1.5 Collagen solutions dosage</td><td></td><td>76</td></tr><tr><td>2.2.1.6 Collagen solutions purity</td><td></td><td>77</td></tr></tbody></table>
<table><tbody><tr><td>2.2.1.7</td><td>Collagen solutions concentration</td><td>78</td></tr><tr><td>2.2.2</td><td>Model system: sodium alginate</td><td>79</td></tr><tr><td>2.2.2.1</td><td>Structure</td><td>79</td></tr><tr><td>2.2.2.2</td><td>Chemical properties</td><td>79</td></tr><tr><td>2.2.2.3</td><td>Solution preparation</td><td>82</td></tr><tr><td>2.2.3</td><td>Collagen-alginate mixtures</td><td>83</td></tr><tr><td>2.2.3.1</td><td>Chemical properties</td><td>83</td></tr><tr><td>2.2.3.2</td><td>Solution preparation</td><td>84</td></tr><tr><td>2.3</td><td>References</td><td>85</td></tr><tr><td>3</td><td>Rheological properties of collagen/alginate solutions</td><td>87</td></tr><tr><td>3.1</td><td>Collagen solution rheological characterization</td><td>88</td></tr><tr><td>3.1.1</td><td>Oscillatory shear flow</td><td>88</td></tr><tr><td>3.1.1.1</td><td>Complex modulus measurement</td><td>88</td></tr><tr><td>3.1.1.2</td><td>Comparison with Gobeaux et al. work [1]</td><td>89</td></tr><tr><td>3.1.2</td><td>Steady shear flow</td><td>90</td></tr><tr><td>3.1.2.1</td><td>Viscosity</td><td>90</td></tr><tr><td>3.1.2.2</td><td>First normal stress difference</td><td>92</td></tr><tr><td>3.1.2.3</td><td>Typical physical quantities for other concentrations</td><td>93</td></tr><tr><td>3.2</td><td>Alginate solutions rheological characterization</td><td>94</td></tr><tr><td>3.2.1</td><td>Viscosity analysis</td><td>94</td></tr><tr><td>3.2.1.1</td><td>Alginate concentration choice</td><td>94</td></tr><tr><td>3.2.1.2</td><td>Shear-thinning nature</td><td>95</td></tr><tr><td>3.2.2</td><td>Complex modulus analysis</td><td>96</td></tr><tr><td>3.2.3</td><td>First normal stress difference and Weissenberg number</td><td>97</td></tr><tr><td>3.3</td><td>Conclusion</td><td>99</td></tr><tr><td>3.4</td><td>References</td><td>100</td></tr><tr><td>4</td><td>Extrusion instabilities of polymer solutions</td><td>101</td></tr><tr><td>4.1</td><td>Quick literature overview: polymer melt extrusion instabilities</td><td>103</td></tr><tr><td>4.1.1</td><td>Surface instability: the sharkskin</td><td>104</td></tr><tr><td>4.1.2</td><td>Spurt-flow defect</td><td>104</td></tr><tr><td>4.1.3</td><td>Bulk distorsion</td><td>104</td></tr><tr><td>4.1.3.1</td><td>Helical</td><td>104</td></tr><tr><td>4.1.3.2</td><td>Gross melt fracture</td><td>105</td></tr><tr><td>4.1.4</td><td>Conclusion</td><td>105</td></tr><tr><td>4.2</td><td>Extrusion in a bath of non-binding ions</td><td>105</td></tr><tr><td>4.2.1</td><td>Context and motivations</td><td>105</td></tr><tr><td>4.2.1.1</td><td>Conclusion</td><td>108</td></tr><tr><td>4.2.2</td><td>Results: Experimental evidence of helical, supercritical instability in pipe flow of shear thinning fluids</td><td>108</td></tr><tr><td>4.2.3</td><td>Implementation on other systems: In the "crimp" quest</td><td>122</td></tr><tr><td>4.2.3.1</td><td>Quick literature overview</td><td>122</td></tr><tr><td>4.2.3.2</td><td>Helical instability: from alginate to collagen</td><td>123</td></tr><tr><td>4.3</td><td>Extrusion in a bath of cross-linking ions</td><td>125</td></tr><tr><td>4.3.1</td><td>Overall observations</td><td>125</td></tr><tr><td>4.3.2</td><td>Smooth region</td><td>126</td></tr><tr><td>4.3.3</td><td>Die swelling</td><td>128</td></tr><tr><td>4.3.4</td><td>Wavy, striated patterns and smooth alternation</td><td>130</td></tr><tr><td>4.3.4.1</td><td>Volumetric waves and surface striations</td><td>130</td></tr></tbody></table>
<table><tbody><tr><td>4.3.4.2</td><td>Oscillating flow</td><td>134</td></tr><tr><td>4.3.5</td><td>Gel clogging</td><td>135</td></tr><tr><td>4.4</td><td>Conclusion</td><td>137</td></tr><tr><td>4.5</td><td>References</td><td>138</td></tr><tr><td><b>5</b></td><td><b>Preparation and characterization of collagen scaffolds</b></td><td><b>141</b></td></tr><tr><td>5.1</td><td>Threads structure characterization techniques</td><td>143</td></tr><tr><td>5.1.1</td><td>Photonic microscopies</td><td>143</td></tr><tr><td>5.1.1.1</td><td>Optical (bright field) and polarized microscopies</td><td>143</td></tr><tr><td>5.1.1.2</td><td>Second Harmonic Generation (SHG) Microscopy</td><td>146</td></tr><tr><td>5.1.2</td><td>Electron microscopies</td><td>152</td></tr><tr><td>5.1.2.1</td><td>Scanning Electron microscopy (SEM)</td><td>153</td></tr><tr><td>5.1.2.2</td><td>Transmission Electron microscopy (TEM)</td><td>154</td></tr><tr><td>5.1.3</td><td>Differential Scanning Calorimetry (DSC)</td><td>155</td></tr><tr><td>5.2</td><td>Collagen batch quality criterion</td><td>156</td></tr><tr><td>5.3</td><td>Fibrillogenesis buffer (Ionic strength) effect</td><td>157</td></tr><tr><td>5.4</td><td>Pure collagen vs collagen/alginate mixtures threads study</td><td>180</td></tr><tr><td>5.4.1</td><td>Pure collagen threads of different concentrations</td><td>180</td></tr><tr><td>5.4.1.1</td><td>Extrusion and maturation</td><td>180</td></tr><tr><td>5.4.1.2</td><td>Structural characterization</td><td>183</td></tr><tr><td>5.4.2</td><td>Collagen/alginate threads</td><td>188</td></tr><tr><td>5.4.2.1</td><td>Ratio choice</td><td>188</td></tr><tr><td>5.4.2.2</td><td>Structural characterization</td><td>189</td></tr><tr><td>5.4.3</td><td>Mechanical properties comparison</td><td>193</td></tr><tr><td>5.5</td><td>Preliminary study to cell culture</td><td>194</td></tr><tr><td>5.5.1</td><td>General morphology</td><td>194</td></tr><tr><td>5.5.1.1</td><td>Optical observations</td><td>194</td></tr><tr><td>5.5.1.2</td><td>Topography analysis</td><td>194</td></tr><tr><td>5.5.2</td><td>Diameter stability</td><td>195</td></tr><tr><td>5.5.3</td><td>Structural characterization</td><td>196</td></tr><tr><td>5.5.3.1</td><td>Pure collagen threads</td><td>196</td></tr><tr><td>5.5.3.2</td><td>Collagen:alginate threads</td><td>197</td></tr><tr><td>5.5.3.3</td><td>Collagen stability: thermal analysis</td><td>199</td></tr><tr><td>5.5.4</td><td>Mechanical properties</td><td>200</td></tr><tr><td>5.5.5</td><td>Discussion</td><td>200</td></tr><tr><td>5.6</td><td>Conclusion</td><td>201</td></tr><tr><td>5.7</td><td>References</td><td>202</td></tr><tr><td><b>6</b></td><td><b>Collagen scaffolds and cell culture</b></td><td><b>205</b></td></tr><tr><td>6.1</td><td>Biological material and techniques</td><td>207</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>6.2</td><td>Cell colonization study</td><td>217</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></tbody></table>
<table><tr><td>6.2.3 Preliminary In situ Hybridization results</td><td>221</td></tr><tr><td>6.2.4 Conclusion</td><td>222</td></tr><tr><td>6.3 Behavior of stem-cells cultured on collagen based threads</td><td>223</td></tr><tr><td>6.3.1 Load influence of threads morphologies</td><td>223</td></tr><tr><td>6.3.2 Cell colonization over culture time and load condition impact analysis</td><td>224</td></tr><tr><td>6.3.3 Genes of interest expression analysis</td><td>226</td></tr><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>6.4 Conclusion</td><td>231</td></tr><tr><td>6.5 References</td><td>232</td></tr><tr><td>CONCLUSION</td><td>235</td></tr><tr><td>List of Figures</td><td>XXIX</td></tr><tr><td>List of Tables</td><td>XXXV</td></tr></table>
# INTRODUCTION

Throughout the human body, about 4000 tendons have been counted [1]. These entities bind muscles to bones and stabilize the joints of the skeletal structure. Two-thirds of their content consists in water and they are known to be poorly vascularized. We use tendons in every daily motion, from the finger snap to the most arduous task. Even for passive movements and postures, our tendons work actively. During effort, they can support 17 times the average total body weight [2]! Of course, such sollicitations inevitably lead to some damages and tendinopathies, tendonitis, tendosynovitis, tendon ruptures are their sworn enemies [3]. Indeed, with injured tendons, motions become painful or even impossible causing joint immobilization. Those injuries are very frequent and so far standard therapies have a low success rate, mostly coming from post-surgery complications: surrounding tissue adherence, foreign body response, inflammation, re-rupture...

For this reason, it is essential to find innovative and sophisticated solutions to repair tendons and thus to turn towards tissue engineering. To claim such amazing mechanical properties, tendons cannot be so simple as springs between bone and muscle, made of water. In fact, tendons have very well organized structures whose major protein components are collagens. To mimic tendons, scaffolds must meet many requirements including reproducing the collagen hierarchical structures from the molecular to the tissue levels.

Several tendon-like scaffolds either made of synthetic polymers or more natural components have been tested to repair tendons. Many techniques are used to produce constructs in a great variety of shapes (threads, tubings, sponges, sheets, 3D constructs etc.) However, they are still not actually close to the "ideal" artificial tendon. Indeed, when trying to reach native tendon mechanical properties, other essential properties as biocompatibility and resorbability are neglected through the use of unsuitable materials or methods. Moreover, most of the techniques used are quite complicated to be implemented.

**In this general context, the present thesis aims to produce biomimetic threads which can be colonized by cells, by developing a bottom-up approach.** As collagen type I is the main component of native tendons, we logically decided to use it as raw material and at concentrations comparable to those of tendon. Such **dense states** have never been tested before. In order to remain close to the **physiological conditions**, no chemical, cross-linking, drying or other synthetic procedures have been used. Therefore, we choose **extrusion** technique as it is the easiest way to produce threads and to be implemented in a laboratory or industry.

The main objective was thus to determine the best conditions to form collagen threads with a tendon-like structure. For this purpose, extrusion process must be perfectly harnessed. A model system, sodium alginate, which has similar collagen rheological proper-
ties was chosen to test a large range of parameters. The results, our understanding and
new expertise was the transposed to the production of collagen threads. Different pa-
rameters, specific to collagen were then varied to evaluate their influence on the threads
formation, ageing, microstructure, mechanical properties and cell culture.

**A first chapter** will briefly describe the tendon functions and structure at the different scales. Based on the clinical context of the tendon repair, the main promising tissue engineering approaches available will be presented and, in particular, those based on collagen. This state of the art part was based on the article review co-authored by Clément Rieu [4] (Appendix).

**The second chapter** will be devoted to the presentation of the methods and materials used during the development of our methodology. A home-made experimental set-up of extrusion will allow us to control several parameters and to observe in real-time the flow at the die exit. First, we will describe this device and techniques used to characterize rheological and mechanical properties of, respectively, solutions used and resulting synthetized threads. In a second part, the complex fluid systems and their respective physico-chemical properties will be described.

**In the third chapter**, rheological characterization will first be performed on collagen solution. Our results will be analyzed and compared with the work of Gobeaux et al. [5] done over a large range of concentrations. Based on these first results, alginate concentrations will then be chosen to obtain solutions exhibiting similar rheological properties as dense collagen ones.

**The fourth chapter** will be focused on alginate extrusion in order to better harness this technique. First, we will briefly review the polymer melt extrusion studies. Then, we will present the results obtained for alginate extrusion either in a bath without or with cross-linking ions. Based on polymer melt expertise, we will attempt to analyze and understand the extrusion instabilities observed.

**The fifth chapter** will present three main studies on collagen based threads production and the respective techniques used to characterize them. Criteria as the stability over time, the microstructure and the mechanical properties will determine the best condition during those studies. First, the influence of ionic strength on collagen threads will be investigated. Then, the collagen concentration (15 to 60mg/mL) will be varied and hybrid system will be tested. A last study will be performed to assess the behavior of our threads in an in vitro cell culture medium in absence of cells.

**In the last chapter**, mesenchymal stem cells seeding of collagen based threads will be performed. Behaviors are studied as a function of several parameters. Cell morphologies, migration, and gene expression will be studied in order to characterize the cells behaviors on our threads.

Finally, in the conclusion, our main results and contribution to tissue engineering for ten-
don repair will be outlined and several project perspectives will be provided.
0.1 References

[1] Harrison Wein and Vickie Contie. Protect your tendons, preventing the pain of tendinitis. *NIH News in Health*, 2014. 1

[2] Marsha Rutland, Dennis O'Connell, Jean-Michel Brismée, Phil Sizer, Gail Apte, and Janelle O'Connell. Evidence-supported rehabilitation of patellar tendinopathy. *North American journal of sports physical therapy: NAJSPT*, 5(3):166, 2010. 1

[3] *Tendon Regeneration: Understanding Tissue Physiology and Development to Engineer Functional Substitutes*. Eds. Academic Press, 2015. 1

[4] Clément Rieu, Lise Picaut, Gervaise Mosser, and Lea Trichet. From tendon injury to collagen-based tendon regeneration: overview and recent advances. *Current pharmaceutical design*, 2017. 2

[5] Frédéric Gobeaux, Emmanuel Belamie, Gervaise Mosser, Patrick Davidson, and So-phie Asnacios. Power law rheology and strain-induced yielding in acidic solutions of type I-collagen. *Soft Matter*, 6(16):3769, 2010. 2
# Chapter 1

# Literature review specific to tendon and its repair strategies

"Sometimes the right path is not the easiest one."

Grandmother Willow in
Pocahontas

## Abstract

In this chapter, we will present the tendon functions and structure at the different length scales. After establishing the clinical situation of tendon injuries and evaluating their incidence, we will present the current strategies available to repair tendons. Then, the tissue engineering (TE) approaches will be described and the tendon scaffold specifications will be set. Finally, we will give an overview of promising TE scaffolds and in particular we will focus on the collagen-based ones obtained by extrusion process.

## Contents

<table><tr><td>1.1</td><td>Tendon</td><td>6</td></tr><tr><td>1.1.1</td><td>Anatomy and structure</td><td>6</td></tr><tr><td>1.1.1.1</td><td>Human anatomy</td><td>6</td></tr><tr><td>1.1.1.2</td><td>Structure</td><td>6</td></tr><tr><td>1.1.2</td><td>Clinical situation</td><td>13</td></tr><tr><td>1.1.2.1</td><td>Pathologies and injuries description</td><td>13</td></tr><tr><td>1.1.2.2</td><td>Incidence</td><td>13</td></tr><tr><td>1.1.2.3</td><td>Current strategies</td><td>15</td></tr><tr><td>1.2</td><td>Tissue engineering for tendon repair</td><td>23</td></tr><tr><td>1.2.1</td><td>Scaffold design specifications</td><td>23</td></tr><tr><td>1.2.2</td><td>Presentation of the main TE techniques used for tendon repair</td><td>24</td></tr><tr><td>1.2.3</td><td>Synthetic polymer scaffolds</td><td>26</td></tr><tr><td>1.2.4</td><td>Scaffolds based on natural/organic components</td><td>27</td></tr><tr><td>1.3</td><td>Conclusion</td><td>31</td></tr><tr><td>1.4</td><td>References</td><td>33</td></tr></table>
# 1.1 Tendon

## 1.1.1 Anatomy and structure

### 1.1.1.1 Human anatomy

Tendon is a dense connective tissue, poorly vascularized and located between muscle and bone (Fig. 6.1) whereas ligaments connect bone to bone. It allows transmitting muscle-contraction force to the skeleton in order to produce motion or simply keep a posture. The extremities of the tendon are called the entheses. They extend from one side to the muscle to form the myotendinous junction, and on the other side to the bone to form the osteotendinous junction. They serve to dissipate stress away from interfaces and are subject to wear and tear.

<figure><img src="image_2.png" /><figcaption>Figure 1.1: Scheme of a tendon which connects muscle to bone and allows motion [1]</figcaption></figure>

There are about 4000 tendons throughout the human body and they are required for any motion in our daily life: grasp an object with the hand, rotate the shoulder, bend the knee, walk etc [1]. The tendons are well known for their incredible mechanical properties. Indeed, they can support 17 times the average total body weight while subjecting to tensile or compression forces.

Like all connective tissues in the body, the tendon is composed of some cells, called in this case tenocytes, tenoblasts and fibroblasts (about 20% of the total volume) and an abundant extracellular matrix (about 80% of this volume). This matrix contains about 65% water and 35% of solid constituents synthesized and secreted by the cells: collagen (30%) and a small amount of non-collagenous glycoproteins (5%). The collagen content represents 60 to 85% of the tendon dry weight [2].

To fulfill its goals/functions, the tendon must be highly shaped and organized as we will see below.

### 1.1.1.2 Structure

Tendon structure has been studied for the past decades 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 species does not help in getting a precise idea [3]. New developments in microscopy will hopefully bring some valuable keys for this understanding [4–6]. Anyhow, some general characteristics are common to all tendons and the most important are discussed in this section.
1.1.1.2.1 Hierarchical organization

Tendons have a hierarchical fibrillar organization that can be compared with russian
dolls arrangement. In fact, each unit is divided into smaller units until the fundamental
one (Fig. 1.2). This architecture is uniaxial as the different units are aligned along its
longitudinal axis.

<figure><img src="image_3.png" /><figcaption>Figure 1.2: Scheme of a tendon hierarchical structure. Adapted from [1]</figcaption></figure>

Surrounding sheaths

The tendon is encircled by an outer connective sheath called the epitenon (Fig. 1.2). It is further ensheated in an another layer of loose connective tissue, the paratenon. Those sheaths give their white color/aspect to the tendon (Fig. 1.3)and contain blood vessels, nerves and cells . Their main functions are to protect the tendon fibers, to reduce friction with adjacent tissues and to facilitate tendon sliding where friction forces are very high. For tendons surrounded by other bones as in hands or feet, an other structure also helps the tendon gliding: the synovial bursae. They are enclosed cavities of various sizes which contain viscous synovial fluid. This whole system acts as a cushion to protect and lubricate the joint.

Fascicles/Bundles

The tendon largest unit is the fascicle which is composed of collagen fibers bound
together. As shown in Figure 1.3, their shape is irregular and their diameter can vary from
20 to 200 µm. Each fascicle is surrounded and delineated by an interfascicular matrix
called endotenon (white junctions) which contains cells.
<figure><img src="image_4.png" /><figcaption>Figure 1.3: Cross-sectional image of an equine superficial flexor tendon. The fascicles arrangement and the respective endotenon sheath (black arrows) can be seen. White scale bar: 1mm
Adapted from [2]</figcaption></figure>

Fibers and fibrils

A collagen fiber is composed of several collagen fibrils and has a diameter ranging
from 50 to 100 µm (Fig. 1.2 and Fig. 1.4, a).

<figure><img src="image_5.png" /><figcaption>Figure 1.4: Scheme and images (TEM [7]) of a) collagen fibers and b) fibrils organization. Different views of 1. cross-sections and 2. along the fibers axis. White scale bar: 1µm.</figcaption></figure>
Collagen fibrils are parallel to each other (Fig. 1.4, 1) and are aligned along the longitu-
dinal fiber axis (Fig. 1.4, 2). However, an axial offset called D is observed and results as
a periodic striation of 67 nm recognizable as an alternation of dark and light stripes (Fig.
1.4, bottom right) visible in transmission electron microscopy (Fig. 1.4, 2). Depending on
the tissues and the hydration of the fibrils, this value can vary between 64 and 68 nm. To
explain this periodic striation, Hodge and Petruska developed a theoretical model in 1963
[8]. The collagen molecules are packed in a quarter staggered orientation which means
that each molecule overlaps its neighbor by a quarter of its length (1.4, bottom right, c)).
The resulting "empty" space is due to the triple helix length of 4.47xD (=300 nm) which
is a noninteger value of the period D. The lighter space is called "gap" and measures 47
nm long while the denser stripe is called "overlap" and measures 20 nm long (1.4, bottom
right). Finally, the collagen triple helix is the smallest unit of the tendon [9].

Wavy structure: the "crimp"

Collagen fibrils within tendon fascicles sheaths (endotenon, epitenon and paratenon) are not necessarily collinear to the tendon and can be found perpendicularly oriented [10]. Indeed, cross-polarized optical microscopic observations of excised rat tail tendons reveal units, from 10 to 100 µm, with alternating oblique orientations compared to the tendon axis. This feature is referred to as "crimp" (Fig.1.5).

<figure><img src="image_6.png" /><figcaption>Figure 1.5: Helical (Top) and crimp organization (Bottom) of tendon fascicles. Adapted from [11]</figcaption></figure>

Evidence of this crimp was further observed by atomic force and electron microscopies (scanning and transmission) [12; 13]. The crimp was suggested to be a natural shock-absorber and to play a role in elastic recoil [14].

Shah et al. (1982) reported that the crimp was only observed at day 16th during the em-
bryonic 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 [15].
This suggests the crimp to be highly related to the presence of collagen fibrils. On the
other hand, for Herchenhan et al.(2012), tenocytes may induce formation of the crimps
[16] . Their conclusion relies on the fact that knocking down cells using polyoxyethylene
(9–10) p-t-octyl phenol (TX-100) within a biomaterial construct presenting a crimp in-
duces 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 [17; 18].

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 [19] (Fig. 1.5, 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 an 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. 1.6, A). 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. 1.6, B).

<figure><img src="image_7.png" /></figure>

Figure 1.6: 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. [11]

It has been shown that super-twisted organizations could be spontaneously generated in liquid-crystals [20]. Knowing that collagen forms different liquid crystal phases [21–23] that are subject to physico-chemical parameters [24], 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 physicochemi-stry work hand in hand.

1.1.1.2.2 Mechanical properties

Tendons have to respond and adapt themselves to the loading transmitted by the mus-
cles. Their entire structure is sollicited at every hierarchical level in order to ensure their
function. There are two different tendon functions: positional and energy storing. In the
first group, tendons transmit forces to bone in order to move joints and are quasi inex-
tensible. In the second one, they act as springs and improve the locomotion efficiency.
However, their intrinsic mechanical properties are lower to positional tendons. For in-
stance, Achilles tendons are energy storing. In the following section, we will describe the
characteristic mechanical properties of tendons, which are generally analyzed by traction
along the tendon axis and up to the tendon rupture *in vitro*. Those experiments are typ-
ically done on traction apparatus as described in section 2.1.3. The results of this kind of
tensile testing are given as stress-strain curves detailed below.

Typical stress-strain curve

<figure><img src="image_8.png" /><figcaption>Figure 1.7: Typical stress-strain curve and corresponding images at the microscale of a rat Achille tendon by polarized light microscopy Scale bar=100µm. In the toe region the macroscopic crimps disappear, in the heel region the molecular kinks in the gaps regions progressively straighten. Adapted from [14],[25]</figcaption></figure>

Briefly, the force (N) is measured during the tendon extension/elongation. The stiffness (N/m) of the material is defined as the ratio of the force and the extension extracted from the gradient of the curve. In order to have characteristic variables independent of sample size, the data need to be normalized. For this purpose, stress and strain are calculated as:

$$
\text{Stress (Pa)} = \frac{\text{Force (N)}}{\text{Cross-sectional area (m}^2\text{)}}
$$

$$
\text{Strain} = \frac{\text{Change in length (m)}}{\text{Original length (m)}}
$$

The failure stress, failure strain and the modulus (stress-strain slope) will be determined
from the tensile testing. Figure 1.7 shows the typical stress-strain curve obtained for such
test. The tendon response to the constant elongation is clearly not linear. In fact, four
different regions can be distinguished from this schematic curve:
1) **Toe region:** in this first region, the crimp previously described and shown in Figure 1.7 (1) Top right) by polarized light microscopy begins to strech. As a spring initially coiled, it unfolds while elongation is applied. The tendon fibers tend to align in the loading direction which generates little effort. Mechanical tests made on tendon clearly show the difference of mechanical behavior during the elongation of the crimp and after [14]. Computational modeling best fits the mechanical experimental data when a helical arrangement is injected into the models [26]. Reese et al. (2010) further improved the fit by combining a wavy pattern with a super-helical arrangement [27]. This data comforts the idea of a super-helical structure as previously described.

2) **Heel region:** the tendon stiffness begins to rise. This is due to reordering at the molecular level. (Fig. 1.7 (2))

3) **Linear region:** a linear zone is obtained between the stress and the deformation. During this phase, micro-tears will form and start damaging the fibers of the tendon, to the macroscopic rupture zone of the tendon. (Fig. 1.7 (3))

4) **Failure:** the microcracks become too large and lead to the macroscopic rupture of the tendon fibers. (Fig. 1.7 (4))

Tendons display viscoelastic behaviors where the "elastic" one comes from the collagen and the "viscous" part from the interfascicular matrix. The four regions represent the tendon behaviors of most of the tendon. However, the order of magnitude of the characteristics physical values vary according to sites and persons, with Young modulus ranging from 0.5 to 1.9 GPa, ultimate tensile strength from 50 to 120 MPa and strain at break between 5 and 20% (Table 1.1) [28]. Thus, the tendinous tissue is capable of modifying its structure, and its composition is a function of the mechanical stresses which it undergoes.

<table><caption>Table 1.1: Mechanical properties of human tendons [28]</caption><thead><tr><td>Tissue</td><td>Modulus (MPa)</td><td>Ultimate Tensile Strength (MPa)</td><td>Ultimate Strain (%)</td></tr></thead><tbody><tr><td>Gracilis Tendon</td><td>590-1734</td><td>111-112</td><td>7-19.4</td></tr><tr><td>Semitendinous Tendon</td><td>540-1081</td><td>89-124</td><td>8-23</td></tr><tr><td>Achilles Tendon</td><td>819 ± 208</td><td>79 ± 22</td><td>9 ± 2</td></tr></tbody></table>

1.1.1.2.3 Other components

As previously mentioned, tendons 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 [29]. The ECM is composed of many different biochemical components such as collagens, non-collagen proteins, and proteoglycans [30; 31].
1.1.2 Clinical situation

**1.1.2.1 Pathologies and injuries description**

Tendons can present acute or chronic injuries caused by extrinsic or intrinsic factors, either alone or in combination (Fig. 1.8) [32]. 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 [33; 34]. The term tendinitis 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 [35; 36]. They are associated with a change in remodeling activity of tendon matrix.

External causes of tendinopathies include occupational, sporting activities, and prescrip-
tion drugs [37]. Intrinsic factors include age and biomechanical imbalance [32]. The eti-
ology of tendinopathies often appears to be multifactorial [33; 34; 38].

1.1.2.2 Incidence

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 [39]. 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. 1.8). A 10-year population-based study led 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 [40].

Achilles tendon rupture occurs at an annual rate of 11.33/100,000 according to Clay-
ton et al. (2008) [39] and has already been reported with rates ranging from 5.5 to 18/100,000
(Fig. 1.8) [41]. 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 [42]. Tennis elbow, also
known as lateral elbow tendinopathy, is secondary to excessive use of wrist extensors 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 [43]. The injury incidence in tennis players has been reported in propor-
tions between 9 and 35% [44]. 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 inci-
dence 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 [45]. The
incidence of rotator cuff tears was reported to be 3.73/100,000 [39].
<figure><img src="image_9.png" /><figcaption>Figure 1.8: Location, annual incidence and main cause of most frequent tendon injuries. Data are referenced in the text. Sources: [39], [41], [42], [43]. 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>

Work-related musculo-skeletal disorders concerning tendons are evident in occupations involving prolonged periods of repetitive, static work [45]. The prevalence of work related upper limb musculoskeletal disorders in UK was estimated to be 730 per 100,000 people in year 2014/2015 [46]. 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) [45]. 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 [47]. 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 [48]. 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 [48].

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 [33], [36], [41], [47]. 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 [49]. Clinicians assess an unsatisfactory clinical outcome after suture for 25% of patients with hand tendon lacerations, and 7.7% of these repairs re-rupture [50].

### 1.1.2.3 Current strategies

The gravity of the injury will determine the strategie to be used to heal the tendon. From chronic tendon injury to tendon tears, it is essential to reduce the patient's pain due to the tissue inflammation with specific drugs and with rehabilitation exercises [51, 52]. Some new treatments based on growth factors delivery through autologous whole blood or platelet rich plasma (PRP) are emerging. Even if in vitro studies prove that these injections improve the collagen production and degradation mechanisms [53], this kind of strategy in clinical cases is still controversial. Indeed, a standard protocol with determined parameters as the injection frequency, volume, concentration etc. is not implemented.

For acute tendon injury, the lesion severity will impose the treatment. For lesions with a gap between the two tendon extremities less than 3 mm, the tendon is able to self-heal [54]. 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 [55] or Integra™ Achillon® system [56],

#### 1.1.2.3.1 Sutures

Three methods of suturing tendon 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 [50] (Fig. 1.9, A). This technique consists in looping the anchor around the epitenon. As a consequence, no collagen fibers bundles are injured. On the contrary, the grasping (Fig. 1.9,B) and locking anchors (Fig. 1.9, C) based respectively on an opened and a closed loop, 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 most effective and is constantly being improved [57].

Several fibers made of polyester (Ethibond®), nylon, polypropylene or a combination of ethylene glycol and terephthalic acid (Da-cron®) are already clinically used for suture alone. They are non-resorbable as they have to maintain almost more than 1 year the end-to-end tendon extremities [50]. 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 [58]. A negative aspect of cable wires is that a second surgery is needed to remove them.
<figure><img src="image_10.png" /><figcaption>Figure 1.9: 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>

Alternative methods have been proposed to avoid the use of suture. Hohendorff et al. (2008) compared post-surgery results of Achilles tendon rupture treated either with fibrin glue or suture. They note that both strategies reach almost the same results. Moreover, fewer complications are observed for the fibrin glue cohort [59]. 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 [60]. They observed that the two parts of the liver were attached together again and that the glue didn't 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.

1.1.2.3.2 Metallic implants

With the same approach as using cable wire sutures, some other researchers have de-
veloped metallic implants. Their aim is to obtain tendon repairs with the best mechanical
properties. Erol et al. (2007) developed an implant made of stainless steel wire with differ-
ent spiral-shaped structure on sheep cadaveric Achilles tendon [61] (Fig. 1.10, A). 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 com-
pared their implant with classic sutures. In some cases, they observed the breaking of
the tendon-implant holding point and the fibering of tendon's ends. Indeed, they raised
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 than the tendon. In
their preliminary study, they obtained good mechanical results compared to suture 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 (Fig. 1.10, B). In their
<figure><img src="image_11.png" /><figcaption>Figure 1.10: Images of different metallic implants: A: coiled spiral with a T-shaped wire stuck to the proximal end of the spiral [61]; B: stainless steel wires with several loops [62], C: nitinol device implemented on digital flexor tendon [63].</figcaption></figure>

study, they also used stainless steel wires but designing it with several loops [62]. 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 tendon. Reese et al. (2013) developed a device for digital flexor tendon repair (Fig. 1.10, C) [63]. It is made of superelastic nitinol (metal alloy of nickel and titanium) tubular parts and is designed specially for hand region (Fig. 1.10, B). They obtained promising results on cadaver model and showed that this kind of device is easy to use and may allow early mobilization of the injured finger post surgery.

Metallic devices as the Tenofix® [64] or Tenolig® [65] are already commercially available. They were tested on cadaveric tendons and on patients, respectively. Tenofix® was compared with conventional suture and no biomechanical improvements were shown passing from one to the other. Tenolig® exhibits different advantages such as the use of percutaneous surgery for the device positioning monitored with ultrasonography (noninvasive imagery technique) to target the two Achilles tendon ends, very small wound scars and complications occuring only for 1 over 75 patients and finally, the return time to sport was estimated to be 9 months. This kind of metallic implants are not largely used in clinical practice, as studies evaluating their effective improvement compared to classic sutures are still rare. Moreover, several limitations such as the site of repair tearing have been highlighted as mentioned above. Extensive improvements of these techniques are needed before becoming a common and viable strategy.
1.1.2.3.3 Grafts

If the site of the tendon injury is neglected or has a poor quality due to tendinopathy,
sutures or implants are not sufficient anymore to repair it [66]. Furthermore, to avoid the
re-rupture, a reinforcement of the site is needed. To meet these objectives, a source of
new material called grafts must be used. The tendon augmentation with grafts consists in
adding the 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 suture, which joins the two tendon extremities, and the graft may improve
the tendon repair. The grafts can come from different origins: the patient himself (auto-
graft), another human being (allograft), animals (xenograft) or it can be synthetic. These
grafts used for tendon repair will be successively described in the following part.

Autografts

Sophisticated surgical procedures have been elaborated to overcome large tendon de-
fects with patient's own "material". Three kinds of procedure more or less invasive allow
accessing the tendon: open, mid-open and percutaneous. Beskin et al. (1987) compare
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 [67]. Those 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. This
is why some clinicians have developed less invasive techniques. Carmont et al. (2007)
also used the peroneus brevis tendon to reconstruct Achilles tendon [68]. Contrary to Be-
skin et al., they did two para-midline incisions. Skin wounds are clearly smaller and skin
integrity is preserved. Moreover, this allows to preserve the surrounded tissues (tendon
sheath) and may improve its healing. To follow through, other autologous sources for the
reconstruction can be recruited. Ebied et al. (2016) repaired Achilles tendon thanks to free
semi-tendinous tendon graft from upper muscles [69]. Shewy et al. (2009) used two intra-
tendinous flaps from the proximal gastrocnemius-soleus complex [70] (Fig. 1.11). Briefly,
they create two flaps in the musculo-tendinous junction and passed them through the
Achilles tendon gap. They fix the flaps extremities to the calcaneus (bone). Few compli-
cations are observed as wound gapping or infection, and long pain after surgery. This
technique showed goods 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 close site to the injured tendon, but less invasive incisions are recommended for the
future. Autografts present several key strengths such as no foreign body response, con-
servation of native mechanical properties and vascularization continuity. The main in-
convenient is the risk of damaging the donor site tissue and thus to create new lesions.
Moreover, the neo tendon reconstructed with surrounding tendons or muscles parts is
thicker than the native one. This can lead to unsightly aspect repair and to a patient dis-
comfort on the repair site. Other strategies have been developed to avoid creating more
lesions in patients to repair injured tendon. Tendon augmentation with xeno-allograft
may be good candidates for this purpose.
<figure><img src="image_12.png" /><figcaption>Figure 1.11: Scheme of Achilles tendon rupture repair by using autologous tendon: two flaps (red arrows) are created in the musculo-tendinous junction and passed through the Achilles tendon gap. Then, the flaps extremities are fixed to the calcaneus (bone) and finally the whole is sutured. [70]</figcaption></figure>

Xeno- 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 xeno-graft. Several clinical studies have been performed to assess their effectiveness for tendon repair augmentation. Phipatanakul et al. (2009) used SIS for rotator cuff repair [71]. 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 pre-sented worse results after the surgery [72]. Clinicians are quite unanimous on this issue. However, SIS may be adapted for tendons in other locations such as 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 design on human cadaveric feet [73]. They created tendon defects greater than 5 cm. The tensile testing showed that in most cases, the suture failed first. As a consequence, the xenograft collagen ribbon insured alone the load with tendon ends not joined anymore. Thus, in comparison with suture alone, the Trellis™ augmentation enhanced strength of Achilles tendon repair. Other decellularized matrices from animal origin are commercialized for other applications (Restore™, OrthADAPT™ (Fig. 1.12, A), TissueMend®, Trellis™, Zimmer®, Cuffpatch® etc.) but are already used for 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 [74]. 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 (Fig. 1.12, B). The authors remind that grafts of dermis origin have better mechanical properties
<figure><img src="image_13.png" /><figcaption>Figure 1.12: Images of commercially available A: OrthADAPT™ xenograft and B: Grafjacket® allograft.</figcaption></figure>

than matrices derived from SIS, previously described. TissueMend® and GrafJacket® presented almost the same strength. They both improved tendon repair compared to structure alone. Lee et al. (2008) also used GraftJacket® matrix to augment tendon repair [75]. 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 [76]. They used the allograft to span the gap between the two tendon ends and the xenograft to secure the repair site. This grafts 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 [77]. 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, is to reproduce this wavy structure. Branch et al. (2011) adopted this strategy with allograft for Achilles tendon repair [78]. 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 [79]. Before being implanted on the tendon injury sites, grafts can be sown
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 differenti-
ated, 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 raise the safety and efficiency problems. Güngormüs et al.
(2016) used tenocytes and seeded them on decellularized rat tendons [80]. Finally, they
implanted the cell-seeded scaffold in a rat model. No immune reaction occurred and their
other results suggested that this scaffold improves mechanical properties of the repair.

Even if xeno- and allografts present some advantages, the matrix alteration due to the de-
cellularization process is a problematic issue as it alters mechanical properties and leads
to antigenic response [81]. It is essential to augment tendon with grafts having almost the
same range of strength. To do so, 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 [82].

Non-resorbable grafts are commonly made of synthetic polymers such as polyester (Da-
cron ®), polytetrafluoroethylene (PTFE or Teflon®) [83], and silicone rubber [84]. Such
materials are already clinically used. Morrey et al. (2016) evaluated non resorbable polypro-
pylene mesh commercially available (Marlex®) on 8 quadriceps tendons ruptures (knee)
[85]. 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, and as they did not per-
form mechanical evaluation, further studies are needed. Abdullah et al. (2015) evaluated
the use of polyethylene teraphtalate (PET) mesh called Orthotape® [86]. 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 [87]. They im-
planted 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. Su-
tures were too fragile compared to the Orthocoupler® which is anchored and sutured in
the muscle. 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 [83]. Adhesions with the surrounding tissues were clearly observed for autogenous grafts, whereas Teflon grafts surface glide and allow 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 their non-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 ([88], [89]), polyglactin [90], polyglycolic acid (Dexon®) [91], or polylactic acid [92]. In the 70s, Jenkins et al. (1980) investigated flexible carbon implants for ligament or tendon injuries. Special affinity of the living tissue with carbon was observed [88]. 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 leads 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 [93].

<figure><img src="image_14.png" /><figcaption>Figure 1.13: Image of commercially available synthetic graft: Artelon®. [94]</figcaption></figure>

Synthetic polymer grafts are also 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®) [94] (Fig. 1.13). Artelon® can be seen as the resorbable equivalent of the Orthotape® 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 to combine 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, the inherent risk of creating new lesions is too high and may lead to secondary surgery. So clinical practices or devices to repair tendon get closer to the ultimate goal but they often lack one specific criteria as strength or body tolerance. This is why, some new scaffolds based on synthetic or biocomponents are currently developed. Such engineering is not straight-forward and needs to meet several strict specifications that we will discuss below.
## 1.2 Tissue engineering for tendon repair

### 1.2.1 Scaffold design specifications

Tendon being such a specific tissue, with multiscale order and high mechanical so- licitations, scaffolds for reconstruction must meet several requirements. These are de- pendent 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 ten- dons 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 [95]. Typical dimensions are given in Table 1.14.

<table><caption>Figure 1.14: Requirements for an ideal scaffold for tendon repair with description and quantitative criteria when assessable. [11]</caption><thead><tr><th></th><th>Requirements</th><th>Descriptions</th><th colspan="2">Criteria</th><th>Refs.</th></tr></thead><tbody><tr><td rowspan="6">Tendon mimism</td><td rowspan="3">Size / Shape</td><td rowspan="3">Scaffold size and shape depend on the site of the tendon to be reconstructed</td><td>Rotator cuff tendon</td><td>3 x 21 x 17 mm</td><td rowspan="3">[Verdiyeva 2015]</td></tr><tr><td>Patellar tendon</td><td>5 x 33 x 59 mm</td></tr><tr><td>Achilles tendon</td><td>5 x 14 x 120 mm</td></tr><tr><td rowspan="3">Mechanical Strength</td><td rowspan="3">Can bear usual mechanical sollicitation throughout the healing process and promote cell differentiation</td><td>Young Modulus</td><td>0.5 - 1.9 GPa</td><td rowspan="3">[Ladd 2011, Jung 2009]</td></tr><tr><td>UTS</td><td>50 - 120 MPa</td></tr><tr><td>Strain at Break</td><td>5 - 20 %</td></tr><tr><td>Alignment</td><td>Alignment in the direction of the load to promote cell alignment and differentiation</td><td>Size threshold</td><td>100 nm</td><td>[Van Delft 2008]</td></tr><tr><td rowspan="4">Temporary biocompatible scaffold</td><td>Biotolerable</td><td>Absence of immune rejection from the scaffold and its degradation products in vivo (and non toxic)</td><td colspan="2">Non cytotoxic, genotoxic, carcinogenic ...</td><td></td></tr><tr><td>Bioactive</td><td>The scaffold promotes cell adhesion and survival</td><td colspan="2">Must be porous to allow exchange with medium<br>Good colonization and viability of the cells on the scaffold</td><td></td></tr><tr><td>Regenerative</td><td>The scaffold fosters the formation of a new tendon by inducing cell differentiation</td><td colspan="2">No observable difference between the new and old tendons</td><td></td></tr><tr><td>Biodegradable /resorbable</td><td>Degradation of the scaffold to leave room to the new tendon/and elimination of degradation products</td><td>Time range :</td><td>2 to 3 months</td><td>[Gilbert 2007, Bady-lak 1995]</td></tr><tr><td>Process</td><td>Tunability</td><td>The material can be shaped in a controled way to suit different location and achieve hierarchical structure</td><td colspan="2">Control of the structure from molecular to macroscopic scale, and tunable size and shape of the scaffold</td><td></td></tr></tbody></table>

Scaffolds must meet as much as possible the native mechanical properties to resist so- licitations and promote cell differentiation [96]. Even though tendons and ligaments are sometimes indiscriminately named in papers given their similarities, they do exhibit dif- ferences when it comes to mechanical properties. Ligaments are usually not as stiff (mod- ulus from 5 to 450 MPa) and strong (UTS between 1 and 50 MPa) as tendons and are more elastic (strain at break from 10 to 40 %) [28]. Another specificity that the scaffolds must mimic is the uniaxial structure of tendons, aligned along the direction of load. To obtain neo-tendons by promoting cell alignment and differentiation, scaffolds should mimic this anisotropy. The influence of scaffold anisotropy has been widely investigated, by using
microgrooves in the micrometer scale for instance [97][98]. Scaffold alignment may be
important down to 100 nm, based on the work of Van Delft and co-workers [99].

Scaffolds for reconstruction must also comply with general specifications for regenerative scaffolds. Biocompatibility is an obvious specification but its definition is not straightforward and has evolved with time. Now, it goes further than the old definition, i.e. the absence of immune rejection or toxicity, called "bio-tolerable" by Mertz [100], adding "the ability of a device material to perform with an appropriate host response in a specific situation" [101]. Scaffolds should not be inert to the body but bioactive, promoting cell adherence and proliferation. Scaffold porosity is here paramount to ensure exchanges and cell penetration.

Besides, the scaffold must be regenerative, enabling cells to differentiate and adopt the relevant phenotype to reconstruct the ECM. Many parameters are at stake here such as mechanical properties of the scaffold, its structure and alignment as well as chemical cues [96] of the ECM or growth factors. To allow good regeneration, the scaffold should at least be biodegradable or even better, bioresorbable, based on definitions given by Vert [102]. This includes the degradation and the elimination of the polymers and by-products. Time of degradation and the subsequent change in mechanical properties are also crucial for tendon reconstruction but very seldom investigated due to the complexity of such a study. To assess the question, works on small intestinal submucosa (SIS) in canine Achilles tendon repair, carried by Badylak et al. (1995) and then Gilbert et al. (2007) are valuable. Badylak et al. (1995) reported full use of the injured leg by the dogs after 9 to 13 weeks and recovery of injured tendon strength after 12 weeks as assessed by tensile testing [103].

Degradation rate of SIS was investigated by Gilbert et al. using carbon 14 measurements, showing 60% loss after 4 weeks and total degradation after 2 months [104]. Thus, we can hypothesize that time span aimed for degradation should be around two months and that, after degradation of the scaffold, recovery and strengthening continue in the month scale.

Last but not least is the processability of the material. This factor is key to meet the spec-
ifications listed above and remains one of the main challenges for biomaterials, as the
processing should not degrade their properties.

We will first present the tissue engineering techniques which are mainly used to pro-
duce scaffolds for tendon repair. Then, we will review the different materials that com-
posed suwh scaffold and we will focus on the collagen type I, which is our "raw material"
for our work.

1.2.2 Presentation of the main TE techniques used for tendon repair

A variety of TE engineering techniques is currently investigated in order to produce scaffolds that fulfill all tendon criteria. Here, we report the procedures mainly used and which lead to constructs tested *in vitro* and *in vivo* for tendon repair.

As tendons exhibit hierarchical structure made of several fibers, fibrils bundles at different
scales, the main idea is to produce threads and assemble them by braiding, gel embedding
etc process. The synthesis of threads can be performed by different techniques as the
three presented below:
<figure><img src="image_15.png" /><figcaption>Figure 1.15: Schematic of the techniques of: A: Extrusion protocol where solution held in the syringe is forced to pass through the tubing and arrives in the buffer which induces the thread formation; B: Electrospinning protocol where a high voltage is applied between the nozzle and the collector, forming a jet of the polymer solution that dries into fiber; C: 3D printing setup where a dispensing unit is monitored in the three directions of space by a computer to deposit bio-ink layer by layer. Adapted from [11]</figcaption></figure>

Extrusion

The extrusion process consists in forcing a material or a solution to pass through a die, giving its shape to the resulting extrudate. This method provides bars, tubes or fibers with a high production rate (few cm/s [105]). It is quite simple to implement as it only needs a syringe with a needle and a bath to induce the thread formation at the die exit (Fig. 1.15 A). The main advantages are its "technical simplicity" and its efficiency. Several parameters play a major role in such technique and may be well controlled in order to produce regular threads. Moreover, most of the authors used not only one extrusion bath and increase the set-up complexity by adding others baths to dehydrate the new forming thread [105]. Finally, co-extrusion procedure, directly derived from simple extrusion, is also widely used to mix directly two species or to produce hollow fibers.

Electrospinning

Electrospinning is a widespread technique for fibrous scaffold production (Fig. 1.15, B). This method, based on a simple and inexpensive set-up, uses an electrical field to cast various materials, such as polymer solutions or melts, into fiber, making it attractive for tissue engineering. Briefly, a high voltage is applied between a grounded target and a needle provided with the liquid of interest, leading to charge accumulation at its surface [106], [107]. For sufficiently high voltage, the surface tension is overcome, leading to the formation of a liquid jet that forms a fiber which dries and is collected on the target, resulting in a fibrous mat. To achieve electrospinning of a polymer solution, many requirements
must be met. In addition to the setup parameters such as voltage or solution feeding rate, and the ambient conditions such as temperature and humidity, intrinsic characteristic of the solution are paramount to electrospin it. Ideal solutions must have low surface tension and high conductivity [108], viscous enough to form a stable jet that does not break before drying [109] but not too much to ensure sufficient flowability [110] and dry fast before reaching the collector [111]. Usual solvents such as acetic acid lack the viscoelastic properties needed to form a stable jet and do not evaporate fast enough before reaching the collector [112].

### 3D printing

3D printing or additive manufacturing, has breached in most good manufacturing sectors in the last few years and tissue reconstruction is no exception (Fig. 1.15, C). Based on the knowledge acquired on scaffold manufacturing and facing new intrinsic problems, it gives high hopes for its capacity to design complex macroscopic shapes of scaffolds as seen in the previous paragraph, as well as enable precise cell deposition. It consists in a layer-by-layer computer aided deposition that enables the construction of 3D structure designed on the computer. This technique is currently used in research on skin, cartilage, bone, aortic valve [113]. Different methods of deposition have been developed such as inkjet, extrusion or laser-assisted bioprinting [113],[114],[115], as well as different hydrogel based inks, called bio-inks [116].

### 1.2.3 Synthetic polymer scaffolds

Synthetic polymer scaffolds can have a sheet-like shape for tendon augmentation repair, similar to the one of synthetic grafts or have a more tendon-like structure made of thread bundles [117]. The materials used are here based on poly(-hydroxy acids). In this category, we find the most commonly used polymer: PGA, PLA, PCL...[2]

<figure><img src="image_16.png" /><figcaption>Figure 1.16: PGA/PLA scaffold example: The scaffold was composed of an inner part of PGA unwoven fibers (A) and an outer part of a net knitted with PGA/PLA fibers (B). The outcome of assembled two parts (C). [118]</figcaption></figure>

In this way, Cao et al. (2001) used unwoven polyglycolic acid (PGA) fibers, which they assembled in a cord shape and wrapped with acellular intestinal submucosa membrane [119]. 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 remodeling of the tendon. Histological results showed that the neo tendon contained a large amount of tenocytes and presented its own structure, easy to distinguish from native tendon.
Other *in vivo* studies came to the same conclusions. Indeed, Ouyang et al. (2003) tested with their biodegradable knitted poly-lactide-co-glycolide (PLGA) scaffold for Achilles tendon repair on a rabbit leg model [117]. Bone marrow cells were seeded on it before implantation. This pre-implantation step allowed a higher rate of tissue formation and remodeling than the cell-free construct. With the same approach, Deng et al. (2014) also investigated this type of polymer scaffold [118]. PGA/PLA constructs were fabricated by combining PGA unwoven fibers longitudinally arranged to form a cord with an outside net scaffold knitted with PGA and PLA fibers. The composite scaffold resulted in a cord (Fig. 1.16). They showed that the *in vitro* culture of these scaffolds before *in vivo* implantation improves the new matrix formation thanks to the *in vitro* cells already present in the scaffold. Moreover, their constructs bridged large defects of about 3 cm and displayed quite good mechanical properties (E=50MPa).

Synthetic polymer scaffolds present promising *in vivo* results and there is no biocompatibility concerns as for synthetic grafts. Thus, the main strength of synthetic constructs relies on their mechanical properties. However, other approaches based on natural components close to the native ones arose to avoid the question of possible inflammation or rejection reactions of the patient.

### 1.2.4 Scaffolds based on natural/organic components

#### Silk based scaffolds

Silk is increasingly investigated as a potential material for tendon repair. This protein often extracted from silkworm displays unique mechanical properties when used as a construct material. For now, scaffolds made from silk fibers are tested for the anterior cruciate ligament repair. Altman et al. (2002) designed a sophisticated silk fiber wire-rope and tested it in *in vitro* conditions [120]. Their scaffolds exhibit benefits in terms of mechanical properties (close to the native ligament ones), biocompatibility and degradation rate, which is slow. In this way, Fan et al. (2008) used knitted silk mesh filled with silk sponges, similar construct as shown in the figure 1.16. Those scaffolds were first put in contact with cells and then implanted in rabbit model [121]. Here again, after 24 weeks, remarkable results were obtained, which are promising for clinical applications. However, some questions about the degradation rate still remain. In fact, as those matrices are slowly degraded by the body, this may lead to body foreign response, or even worse, fibrosis. Moreover, the purification of the silk must be perfectly executed: if some traces of sericin, which is a glue-like protein, are not removed, this may lead to increased adhesion of the scaffold with the surrounding tissues.

#### Polysaccharides

The polysaccharides represent an other source of materials for tendon TE. In this category of biopolymer, we can find: alginate, chitin, chitosan, agarose, pectin etc. In Nowotny et al. (2016) work, pure chitosan scaffolds are produced by wet electrospinning and braiding steps [122]. An example of such construct is presented in figure 1.17. The mechanical properties are improved with larger chitosan filaments and reach moduli of about 100 MPa. Cells adhered on the scaffolds in a comparable manner to the control (polystyrene 2D wells).

Other polysaccharides such as alginate may also be used. Namba et al. (2006) tested algi-
<figure><img src="image_17.png" /><figcaption>Figure 1.17: Braided chitosan scaffolds.[122]</figcaption></figure>

nate solution injection post-surgery on tendon repair site in rabbits [123]. They showed that the lesions coated with alginate presented less adhesions and that the joint exion was improved. This non-adhesive property is promising for the synthesis of new scaffolds as adhesion occurs on the scared tendon areas and decreases the entire recovery of the joint motion.

Silk and polysaccharides display both amazing properties and are good candidates in the
tendon gold standard quest. As some aspects need to be improved, several studies are
nowadays based on hybrid scaffolds such as: silk/PLGA

Collagen based scaffolds

Collagen type I present lots of advantages as it is the main component of tendons
and because it is biocompatible and degraded by enzym with low inflammatory response
[81], [124]. It promotes cell adhesion [125] and differentiation [96], and is one of the most
favorable biomaterials for cell survival compared to cheap and easy to use alginate for
instance, or to chitosan [116]. It is thus a promising material that fuels much innovative
research [105].

a) Extrusion

Collagen solution extrusion was first developed in the 1960s to create new biosutures for ophthalmic surgeries [126]. Since then, the potential of this molecule was highlighted leading to its widespread use in tissue engineering. Collagen solution extrusion is quite easy to implement requiring only a syringe, a needle or tubing, a collecting bath and a device to control the flow rate (syringe pump) (Fig. 1.15, A). Basically, to produce collagen threads the syringe is filled with a collagen solution and mounted in the syringe pump. Then, the solution will pass through the tubing and will go out in a bath containing the fibrillogenesis buffer. Finally, some other post-production processes can be performed [127-133]. Self-assembly of collagen may occur either with a pH transition when using an acidic collagen solution and a basic to neutral pH buffer, or through a temperature jump for cooled neutral collagen solution.
In most studies, the collagen is extracted from rat tail (RTT) or bovine Achilles tendons (BAT). Zeugolis et al. (2008) compared the properties of threads made of RTT or BAT acidic collagen [134]. By scanning electron microscopy, they observed no significant differences in the substructure of the collagen fibers in the two kinds of threads. Collagen concentration of solution commonly used ranges from 5 mg/mL to 30 mg/mL in acidic conditions (pH=2) [128], [130], [132; 133; 135]. For the fibrillogenesis buffer, the same basis is found in different studies: 135 mM of NaCl, 30 mM of TES (tris(hydroxylmethyl)methyl-2-aminoethane sulphonic acid) and 30 mM sodium phosphate dibasic with a resulting pH of 7.5 [129; 130; 133]. Other buffers as phosphate saline buffer (PBS) [131] or polyethylene glycol in sodium phosphate dibasic and monobasic are also used. Characteristic dimensions of collagen threads, obtained by this technique, range from 100 µm to 1 mm for wet diameter and around 20 µm to 100 µm after drying [111], [129; 130; 133]. The length depends on the syringe volume capacity and the inner diameter of the tubing. In general, the order of magnitude of the collagen thread length ranges between a few centimeters to meter. Mechanical properties of collagen strands have been measured routinely in dry conditions and, less frequently, in wet conditions. Cornwell et al. (2007) measured Young modulus of wet collagen fibers around 4 MPa [132]. Regarding the elastic modulus of native tendon, those fibers are between 2 and 3 orders of magnitude "weaker". For the ultimate tensile strength (UTS), Cavallaro et al. (1994) [105], Cornwell et al. (2007) [132] and Dunn et al. (1993) [130] obtained results ranging from 1.2 to 10 MPa. Dunn et al. (1993) also observed that UTS decreases when the diameter fiber increases. So, in wet conditions, which mimic the physiological environment, mechanical properties of collagen fibers are still far from that of native tendon (Fig. 1.14). This is why different cross-linking (CXL) agents are commonly used to improve those characteristics. Cross-linking process consists in creating new bonds between collagen molecules with chemical species or external excitation sources. The most commonly used on collagen threads are glutaraldehyde, carbodiimide (EDC), cyanamide, hydroxysuccinimide (NHS) and dehydrothermal treatment (DHT), simple drying process, or ultra violet (UV) exposition.

Kato et al. (1989) have shown that they can produce collagen threads cross-linked with glutaraldehyde having an UTS of 50 to 66 MPa in wet conditions [131]. Those results are in the same range than UTS of rat tail tendons (33 to 39 MPa). The UTS is divided by a factor 2 for severe dehydration and for EDC treatment. As for Young modulus, they reach 400-500 MPa with glutaraldehyde CXL process. Zeugolis et al. (2008) evaluated mechanical properties of collagen threads dried under their own weight [136]. Young modulus of those strands ranged from 1 to 3 GPa and UTS from 0.2 to 0.3 GPa. These results have similar values as native tendon (Fig. 1.14). However, we have to keep in mind that in this case, tensile testing is performed in dry conditions, which are not close to the physiological ones. As for Kato et al. (1989) [131], they lose on average 3 orders of magnitude in wet conditions. Finally, to have as many bonds as possible, some combinations of CXL processes can be used (dehydrothermal with cyanamide [130] etc).

The alignment of collagen fibrils inside the collagen threads is also a parameter to take into account. The extrusion technique has the main advantage to induce collagen fibrils alignment in the longitudinal axis, thanks to the high strain applied when collagen solution is forced to pass through the narrow die ([105], [129]). This phenomenon can be further increased by stretching the threads [127]. Lai et al. (2011) observed that, *in vitro*, human fibroblasts cultured during 1 or 2 days have an elongated shape following the longitudinal axis of the collagen fibrils [129]. This suggests that there is some con-
tact guidance of the cells on the collagen threads structure. Cornwell et al. (2007) high-
lighted the problem of CXL post-process on collagen threads. Indeed, they observed that
fibroblast migration rate decreased by using CXL agents [132]. A compromise between the
improvement of mechanical properties with CXL techniques and the upgrading of cell
proliferation must be found. To go further, Dunn et al. (1993) implanted their collagen
threads in subcutaneous space of rabbit during 4 weeks. The threads were surrounded by
a fibrous tissue and were colonized by inflammatory cells and fibroblasts [130]. Threads
crosslinked with DHT combined with cyanamide were degraded over 4 weeks. Cavallaro
et al. (1994) also had promising results in a canine ligament model and a rat abdominal
repair model that can be transposed for tendon repair [105]. They noticed neovascular-
ization and cell infiltration in the collagen threads. In both studies, no rejection of the
collagen scaffolds was observed.

b) *Electrospinning*

Collagen electrospinning was achieved using fluoro-alcohols[137]. However, such a
process was proven to denature collagen [112]. Indeed, resulting collagen is soluble in
water, does not exhibit characteristic D-banding under transmission electron microscopy
and has lower denaturation temperature. Further proofs were given thanks to circular
dichroism, electrophoresis and infrared spectroscopy. Attempts succeeded in electro-
spinning collagen in benign solvent such as Phosphate Buffer Saline and ethanol mix
[106] or acidic solution with poly(ethylene oxide) [138]. Bürck et al. (2013) demonstrated
that electrospun collagen fiber from this solvent did not preserve much better the native
structure of collagen and that part of the denatured collagen was intrinsically due to the
electrospinning process and not to the solvent [111].

Recently, Elamparithi et al. (2015) reported successful electrospinning of collagen in 97:3 (acetic acid:DMSO) mix with fibers exhibiting the characteristic D-banding of native collagen [139]. However, no further investigation on native collagen content was carried out. The cross-linking of the resulting fibrous mat after electrospinning with 1-ethyl-3(3-dimethylaminopropyl) carbodiimide hydrochloride does not allow for water solubility test and gives surprisingly low mechanical properties with modulus and tensile strength at maximum load of a few kPa. This may be suitable for non-load bearing tissue reconstruction such as cardiac tissue, aimed by the authors, but not for tendon reconstruction.

Thus, electrospinning of collagen for tendon reconstruction seems only relevant when collagen is co-electrospun with biodegradable synthetic polymers such as poly(l-lactide-co-ε-caprolactone) [140] or poly(L-lactic acid) [99], as it improves biocompatibility of synthetic polymers. For instance, an interesting work was published by Ladd et al. (2011) who electrospan two different polymers with collagen to reproduce the musculo-tendinous junction: the poly(L-lactic acid), being stiffer, for the tendon part and poly(ε-caprolactone) for the muscle part [141]. Whereas the synthetic polymers provide mechanical strength, the collagen promotes good cell viability close to 100% after 7 days of both C2C12 myoblasts and NIH3T3 fibroblasts seeded on each part.

c) 3D *printing*

The use of collagen as an ink for 3D printing raises an obvious interest and is inves-
tigated especially for bone [142] and cartilaginous tissue reconstruction [143]. However,
current research faces difficulties to obtain self-standing 3D printed structures. Thus, col-
lagen is often used to enhance the bioactivity of widespread bioinks such as alginate [144]
or agarose [115] that are cheap and easy to use but show lower bioactivity compared to
main components of the ECM. It is also used with calcium phosphate [145] for bone re-
construction or mixed with fibrinogen [146] that is easy to crosslink.

<figure><img src="image_18.png" /><figcaption>Figure 1.18: 3D plotted collagen scaffold (33 layers) in dry state [147]</figcaption></figure>

Other techniques to print collagen self-standing structures are the deposition of collagen on a cold plate to freeze the solution when deposed [144], or the use of sacrificial negative molds [148], [149]. An original process that is worth mentioning is the printing of a core-sheath collagen-alginate structure [150]. Here again, collagen is sometimes used without inducing fibrillogenesis. Besides, the protocols used can be too harsh to incorporate cells in the bio-ink, preventing the use of additive manufacturing to deposit cells in an accurate and homogeneous way. A few articles deal with pure collagen self-assembled during or after printing [151], [152]. However, for collagen [152] as well as for fibrinogen [146], this leads to tedious protocols where the scaffold is incubated, after each layer printing, either at 37°C or with thrombin respectively.

One currently adopted strategy is the printing side-by-side of an artificial biodegradable polymers and a bio-ink with incorporated cells. The artificial polymer confers mechanical integrity to the scaffold whereas the bio-ink allows exact cell deposition with suitable bioactivity [146], [153]. To our knowledge, despite the control on porosity, macroscopic structure and cell incorporation that this technique offers, no article about 3D printing of collagen for tendon reconstruction has been published yet. This can be explained by the difficulty of handling collagen and the poor mechanical properties obtained with collagen in other domains such as extrusion. As collagen is often used with other bio-inks for its bioactivity, some chose gelatin instead for its lower price. For instance, Merceron et al. (2005) published the co-printing of two synthetic polymers polyurethane (PU) and poly(ε-caprolactone) (PCL) together with a bio-ink based on hyaluronic acid, fibrinogen and gelatin, to reproduce the musculo-tendinous junction [154]. The bio-ink contained C2C12 myoblasts on the PU side while the bio-ink on the stiffer PCL side contained NIH/3T3. The construct exhibited good cell viability, high cell alignment and characteristic musculo-tendinous gene upregulation. They highlighted the ability offered by 3D printing to deposit different cell types in precise locations, paving the way to complex multicellular scaffolds for regeneration.

**1.3 Conclusion**

Current tendon repair methods are mainly based on sutures and grafts that have proven
to be effective but exhibit intrinsic drawbacks. Even though tendon structure is not com-
pletely understood, tendon regeneration based on bioengineered scaffold is promising.
The diversity of techniques and the increasing know-how in scaffold design allows tack-
ling new exciting aspects such as alignment, multi-scale organization, scaffold tailoring or
multi-cell deposition. Noteworthy mechanical stimulation appears also to be a key factor
for tendon healing, independently of the suture, graft or scaffold used. Collagen I stands
out from other biomolecules as main component of tendon and thanks to its suitable bio-
compatible properties. However, mechanical properties of collagen scaffolds remain an
issue. Cross-linkers such as carbodiimide are mainly used to improve them but are not
sufficient to bridge the gap and may introduce some drawbacks.

**In this work,** we thus chose collagen type I solutions to be shaped as tendon-like scaf-
fold, at concentration range almost 20-fold higher than in other studies. We have seen that
the first and easiest way used to produce threads is the extrusion technique. However, this
process must be fully controlled as complex fluids extrusion may lead to severe flow in-
stabilities, which disturb threads morphologies. Therefore, we designed a home-made
extrusion set-up, where a large range of pieces and parameters are adjustable and where
the exit flow is observed in real-time. A system model with similar rheological properties
is chosen to study those instabilities and control them while extruding dense collagen so-
lutions. Some hybrid threads will also be tested to combine these two systems exhibiting
interesting behaviors. Finally, the collagen based threads will be cultured with cells *in vitro* under different conditions to assess their interactions with biological systems.
1.4 References

[1] Geoffroy Nourissat, Francis Berenbaum, and Delphine Duprez. Tendon injury: from biology to tendon repair. *Nature reviews Rheumatology*, 11(4):223–233, 2015. 6, 7

[2] *Tendon Regeneration: Understanding Tissue Physiology and Development to Engineer Functional Substitutes*. Eds. Academic Press, 2015. 6, 8, 26

[3] Helen L. Birch. Tendon matrix composition and turnover in relation to functional requirements: *Tendon matrix composition and turnover. International Journal of Experimental Pathology*, 88(4):241–248, August 2007. 6

[4] Stéphane Bancelin, Charles-André Couture, Katherine Légaré, Maxime Pinsard, Maxime Rivard, Cameron Brown, and François Légaré. Fast interferometric second harmonic generation microscopy. *Biomedical Optics Express*, 7(2):399, February 2016. 6

[5] 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, May 2012.

[6] Alexandre Fouré. New Imaging Methods for Non-invasive Assessment of Mechanical, Structural, and Biochemical Properties of Human Achilles Tendon: A Mini Review. *Frontiers in Physiology*, 7, July 2016. 6

[7] Yufei Li, Gion Fessel, Marios Georgiadis, and Jess G Snedeker. Advanced glycation end-products diminish tendon collagen fiber sliding. *Matrix Biology*, 32(3):169–177, 2013. 8

[8] John A Petruska and Alan J Hodge. A subunit model for the tropocollagen macromolecule. *Proceedings of the National Academy of Sciences*, 51(5):871–876, 1964. 9

[9] Pramod B Voleti, Mark R Buckley, and Louis J Soslowsky. Tendon healing: repair and regeneration. *Annual review of biomedical engineering*, 14:47–71, 2012. 9

[10] P. Kannus. Structure of the tendon connective tissue. *Scandinavian journal of medicine & science in sports*, 10(6):312–320, 2000. 9

[11] Clément Rieu, Lise Picaut, Gervaise Mosser, and Lea Trichet. From tendon injury to collagen-based tendon regeneration: overview and recent advances. *Current pharmaceutical design*, 2017. 9, 10, 23, 25

[12] S. P. Nicholls, L. J. Gathercole, A. Keller, and J. S. Shah. Crimping in rat tail tendon collagen: morphology and transverse mechanical anisotropy. *Int. J. Biol. Macromol.*, 5:283–288, 1983. 9

[13] J Dlugosz, L. J. Gathercole, and A. Keller. TRANSMISSION ELECTRON MICROSCOPE STUDIES AND THEIR RELATION TO POLARIZING OPTICAL MICROSCOPY IN RAT TAIL TENDON. *Micron*, 9:71–82, 1978. 9

[14] James H.-C. Wang. Mechanobiology of tendon. *Journal of Biomechanics*, 39(9):1563–1582, January 2006. 9, 11, 12
[15] J. S. Shah, E. Palacios, and L. Palacios. Development of crimp morphology and cellular changes in chick tendons. *Developmental biology*, 94(2):499-504, 1982. 9

[16] Andreas Herchenhan, Nicholas S. Kalson, David F. Holmes, Patrick Hill, Karl E. Kadler, and Lee Margetts. Tenocyte contraction induces crimp formation in tendon-like tissue. *Biomechanics and Modeling in Mechanobiology*, 11(3-4):449-459, March 2012. 9

[17] Xiao Wen Hu, David Knight, P. and John Chapman, A. The effect of non-polar liquids and non-ionic detergents on the ultrastructure and assembly of rat tail tendon collagen fibrils in vitro. *Biochimica et Biophysica Acta*, 1334:327-337, 1997. 9

[18] Jeffrey S. Cartmell and Michael G. Dunn. Effect of chemical treatments on tendon cellularity and mechanical properties. *Journal of biomedical materials research*, 49(1):134-140, 2000. 9

[19] Benedicto de Campos Vidal. Image analysis of tendon helical superstructure using interference and polarized light microscopy. *Micron*, 34(8):423-432, December 2003. 10

[20] Françoise Livolant and Amélie Leforestier. Condensed phases of DNA: structures and phase transitions. *Progress in polymer science*, 21(6):1115-1164, 1996. 10

[21] N. S. Murthy. Liquid crystallinity in collagen solutions and magnetic orientation of collagen fibrils. *Biopolymers*, 23(7):1261-1267, 1984. 10

[22] Hideatsu Maeda. An Atomic Force Microscopy Study of Ordered Molecular Assemblies and Concentric Ring Patterns from Evaporating Droplets of Collagen Solutions. *Langmuir*, 15(24):8505-8513, November 1999.

[23] Marie-Madeleine Giraud-Guille. Liquid crystallinity in condensed type I collagen solutions: a clue to the packing of collagen in extracellular matrices. *Journal of molecular biology*, 224(3):861-873, 1992. 10

[24] Paulo De Sa Peixoto, Ariane Deniset-Besseau, Marie-Claire Schanne-Klein, and Gervaise Mosser. Quantitative assessment of collagen I liquid crystal organizations: role of ionic force and acidic solvent, and evidence of new phases. *Soft Matter*, 7(23):11203, 2011. 10

[25] Marco Franchi, Vittoria Ottani, Rita Stagni, and Alessandro Ruggeri. Tendon and ligament fibrillar crimps give rise to left-handed helices of collagen fibrils in both planar and helical crimps. *Journal of Anatomy*, 216(3):301-309, March 2010. 11

[26] Rafael Grytz and Günther Meschke. Constitutive modeling of crimped collagen fibrils in soft tissues. *Journal of the Mechanical Behavior of Biomedical Materials*, 2(5):522-533, October 2009. 12

[27] Shawn P. Reese, Steve A. Maas, and Jeffrey A. Weiss. Micromechanical models of helical superstructures in ligament and tendon fibers predict large Poisson's ratios. *Journal of Biomechanics*, 43(7):1394-1400, May 2010. 12

[28] Ho-Joong Jung, Matthew B Fisher, and Savio L-Y Woo. Role of biomechanics in the understanding of normal, injured, and healing ligaments and tendons. *BMC Sports Science, Medicine and Rehabilitation*, 1(1), December 2009. 12, 23
[29] Christopher A. Squier and W. H. Bausch. Three-dimensional organization of fibroblasts and collagen fibrils in rat tail tendon. *Cell and tissue research*, 238(2):319-327, 1984. 12

[30] G. Zhang, B. B. Young, Y. Ezura, M. Favata, L. J. Soslowsky, S. Chakravarti, and D. E. Birk. Development of tendon structure and function: regulation of collagen fibrillogenesis. *J Musculoskelet Neuronal Interact*, 5(1):5-21, 2005. 12

[31] Hazel R. C. Screen, David E. Berk, Karl E. Kadler, Francesco Ramirez, and Marian F. Young. Tendon Functional Extracellular Matrix: TENDON FUNCTIONAL EXTRA-CELLULAR MATRIX. *Journal of Orthopaedic Research*, 33(6):793-799, June 2015. 12

[32] D. S. Morais, J. Torres, R. M. Guedes, and M. A. Lopes. Current Approaches and Future Trends to Promote Tendon Repair. *Annals of Biomedical Engineering*, 43(9):2025-2035, September 2015. 13

[33] Stavros Thomopoulos, William C. Parks, Daniel B. Rifkin, and Kathleen A. Derwin. Mechanisms of tendon injury and repair: TENDON INJURY AND REPAIR. *Journal of Orthopaedic Research*, 33(6):832-839, June 2015. 13, 14

[34] Jonathan S. Young and N. Maffuli. Etiology and epidemiology of Achilles tendon problems. *The Achilles Tendon*, pages 39-49, 2007. 13

[35] Graham Riley. Tendinopathy—from basic science to treatment. *Nature Clinical Practice Rheumatology*, 4(2):82-89, February 2008. 13

[36] A.J. Lomas, C.N.M. Ryan, A. Sorushanova, N. Shologu, A.I. Sideri, V. Tsioli, G.C. Fthenakis, A. Tzora, I. Skoufos, L.R. Quinlan, G. O’Laighin, A.M. Mullen, J.L. Kelly, S. Kearns, M. Biggs, A. Pandit, and D.I. Zeugolis. The past, present and future in scaffold-based tendon treatments. *Advanced Drug Delivery Reviews*, 84:257-277, April 2015. 13, 14

[37] Trevor Lewis and Jill Cook. Fluoroquinolones and Tendinopathy: A Guide for Athletes and Sports Clinicians and a Systematic Review of the Literature. *Journal of Athletic Training*, 49(3):422-427, June 2014. 13

[38] Francesco Oliva, Alessio Giai Via, and Nicola Maffulli. Calcific tendinopathy of the rotator cuff tendons. *Sports medicine and arthroscopy review*, 19(3):237-243, 2011. 13

[39] Robert A.E. Clayton and Charles M. Court-Brown. The epidemiology of musculoskeletal tendinous and ligamentous injuries. *Injury*, 39(12):1338-1344, December 2008. 13, 14

[40] Johanna P. de Jong, Jesse T. Nguyen, Anne J. M. Sonnema, Emily C. Nguyen, Peter C. Amadio, and Steven L. Moran. The Incidence of Acute Traumatic Tendon Injuries in the Hand and Wrist: A 10-Year Population-based Study. *Clinics in Orthopedic Surgery*, 6(2):196, 2014. 13

[41] Cheng Sun, Qi Zhuo, Wei Chai, Jiying Chen, Wei Yang, Peifu Tang, and Yan Wang. Conservative interventions for treating Achilles tendon ruptures. In *The Cochrane Collaboration, editor, Cochrane Database of Systematic Reviews*. John Wiley & Sons, Ltd, Chichester, UK, October 2013. 13, 14
[42] Alexander Scott and Maureen C. Ashe. Common Tendinopathies in the Upper and Lower Extremities: *Current Sports Medicine Reports*, 5(5):233-241, October 2006. 13, 14

[43] T. L. Sanders, H. Maradit Kremers, A. J. Bryan, J. E. Ransom, J. Smith, and B. F. Morrey. The Epidemiology and Health Care Burden of Tennis Elbow: A Population-Based Study. *The American Journal of Sports Medicine*, 43(5):1066-1071, May 2015. 13, 14

[44] B M Pluim. Tennis injuries: occurrence, aetiology, and prevention. *British Journal of Sports Medicine*, 40(5):415-423, May 2006. 13

[45] European Agency for Safety and Health at Work. OSH in figures: Work-related musculoskeletal disorders in the EU - Facts and figures, 2010. 13, 14

[46] UK Health and Safety Executive. Work-related Musculoskeletal disorder (WRMSDs) Statistics, Great Britain, 2015, 2015. 14

[47] Agency for Healthcare Research and Quality. Comparative Effectiveness of the Non-operative and Operative Treatments for Rotator Cuff Tears, 2010. 14

[48] European Agency for Safety and Health at Work. Work-related neck and upper limb musculoskeletal disorders, 1999. 14

[49] Nelly Andarawis-Puri, Evan L. Flatow, and Louis J. Soslowsky. Tendon basic science: Development, repair, regeneration, and healing: TENDON DEVELOPMENT, INJURY, AND REPAIR. *Journal of Orthopaedic Research*, 33(6):780-784, June 2015. 15

[50] Shelley Rawson, Sarah Cartmell, and Jason Wong. Suture techniques for tendon repair; a comparative review. *Muscles, ligaments and tendons journal*, 3(3):220, 2013. 15

[51] Marc A. Childress and Anthony Beutler. Management of chronic tendon injuries. *Am Fam Physician*, 87(7):486-490, April 2013. 15

[52] K. M. Khan and A. Scott. Mechanotherapy: how physical therapists' prescription of exercise promotes tissue repair. *Br J Sports Med*, 43(4):247-252, April 2009. 15

[53] Emily A. Sundman, Brian J. Cole, and Lisa A. Fortier. Growth factor and catabolic cytokine concentrations are influenced by the cellular composition of platelet-rich plasma. *Am J Sports Med*, 39(10):2135-2140, October 2011. 15

[54] R. H. Gelberman, M. I. Boyer, M. D. Brodt, S. C. Winters, and M. J. Silva. The effect of gap formation at the repair site on the strength and excursion of intrasynovial flexor tendons. An experimental study on the early stages of tendon-healing in dogs. *J Bone Joint Surg Am*, 81(7):975-982, July 1999. 15

[55] Patrick S. Buckley and David I. Pedowitz. Bone Reduction Clamp to Gain Length in Repairing Chronic Achilles Tendon Ruptures. *Orthopedics*, August 2016. 15

[56] M. D. Zeev Feldbrin, M. D. David Hendel, M. D. Alexander Lipkin, M. D. Dan Zin, and M. D. Louis Schorr. achilles tendon rupture and our experience with the achillon device. September 2016. 15
[57] Javier Maquirriain. Achilles Tendon Rupture: Avoiding Tendon Lengthening during Surgical Repair and Rehabilitation. *Yale J Biol Med*, 84(3):289-300, September 2011.
15

[58] Benedikt Schliemann, Niklas Grüneweller, Daiwei Yao, Clemens Kösters, Simon Lenschow, Steffen B. Roßlenbroich, Michael J. Raschke, and Andre Weimann. Biomechanical evaluation of different surgical techniques for treating patellar tendon ruptures. *International Orthopaedics (SICOT)*, 40(8):1717-1723, October 2015.
15

[59] Bernd Hohendorff, Wolf Siepen, Liesbeth Spiering, Lukas Staub, Thilo Schmuck, and Andreas Boss. Long-term Results after Operatively Treated Achilles Tendon Rupture: Fibrin Glue versus Suture. *The Journal of Foot and Ankle Surgery*, 47(5):392-399, September 2008. 16

[60] Séverine Rose, Alexandre Prevoteau, Paul Elzière, Dominique Hourdet, Alba Marcellan, and Ludwik Leibler. Nanoparticle solutions as adhesives for gels and biological tissues. *Nature*, 505(7483):382-385, January 2014. 16

[61] Bülent Erol, Barış Kocaoğlu, and Tanıl Esemenli. Spiral-shaped Metallic Implant in the Treatment of Achilles Tendon Ruptures: An Experimental Study on the Achilles Tendon of Sheep. *The Journal of Foot and Ankle Surgery*, 46(3):155-161, May 2007.
16, 17

[62] İsmail Ağır, Mahmut Nedim Aytekin, Onur Başçı, Barış Çaypınar, and Bülent Erol. Tendon-Holding Capacities of Two Newly Designed Implants for Tendon Repair: An Experimental Study on the Flexor Digitorum Profundus Tendon of Sheep. *Open Orthop J*, 8:135-139, June 2014. 17

[63] Shawn P Reese and Erik N Kubiak. A nitinol based flexor tendon fixation device: Gapping and tensile strength measurements in cadaver flexor tendon. *Journal of biomechanical engineering*, 136(1):014501, 2014. 17

[64] Scott W. Wolfe, Andrew A. Willis, Deidre Campbell, Jonathan Clabeaux, and Timothy M. Wright. Biomechanic Comparison of the Teno Fix Tendon Repair Device With the Cruciate and Modified Kessler Techniques. *The Journal of Hand Surgery*, 32(3):356-366, March 2007. 17

[65] S. Lacoste, J.M. Féron, and B. Cherrier. Percutaneous Tenolig® repair under intraoperative ultrasonography guidance in acute Achilles tendon rupture. *Orthopaedics & Traumatology: Surgery & Research*, 100(8):925-930, December 2014. 17

[66] Robert A Magnussen, Richard R Glisson, and Claude T Moorman III. Augmentation of achilles tendon repair with extracellular matrix xenograft: a biomechanical analysis. *The American journal of sports medicine*, 39(7):1522-1527, 2011. 18

[67] James L. Beskin, Richard A. Sanders, Stephen C. Hunter, and Jack C. Hughston. Surgical repair of Achilles tendon ruptures. *Am J Sports Med*, 15(1):1-8, January 1987.
18

[68] Michael R Carmont and Nicola Maffulli. Less invasive Achilles tendon reconstruction. *BMC Musculoskelet Disord*, 8:100, October 2007. 18
[69] Ayman M. Ebied and Adel Abdel Azim Foda. Less Invasive Reconstruction of Chronic Achilles Tendon Rupture with free Semitendinosus Tendon Autograft - A case Series. *International Journal of Orthopaedics*, 3(2):539-543, April 2016. 18

[70] Mohamed Taha El Shewy, Hassan Magdy El Barbary, and Hisham Abdel-Ghani. Repair of Chronic Rupture of the Achilles Tendon Using 2 Intratendinous Flaps From the Proximal Gastrocnemius-Soleus Complex. *Am J Sports Med*, 37(8):1570-1577, January 2009. 18, 19

[71] Wesley P. Phipatanakul and Steve A. Petersen. Porcine small intestine submucosa xenograft augmentation in repair of massive rotator cuff tears. *Am J. Orthop.*, 38(11):572-575, November 2009. 19

[72] Steven G. Sclamberg, James E. Tibone, John M. Itamura, and Sina Kasraeian. Six-month magnetic resonance imaging follow-up of large and massive rotator cuff repairs reinforced with porcine small intestinal submucosa. *J Shoulder Elbow Surg*, 13(5):538-541, October 2004. 19

[73] Gregory C. Berlet, Christopher F. Hyer, Thomas H. Lee, and Barbara E. Blum. Collagen Ribbon Augmentation of Achilles Tendon Tears: A Biomechanical Evaluation. *The Journal of Foot and Ankle Surgery*, 53(3):298-302, 2014. 19

[74] Lin Song, Raymond E. Olsen, Jeffrey P. Spalazzi, and Twana Davisson. Biomechanical Evaluation of Acellular Collagen Matrix Augmented Achilles Tendon Repair in Sheep. *The Journal of Foot and Ankle Surgery*, 49(5):438-441, September 2010. 19

[75] Daniel K. Lee. A Preliminary Study on the Effects of Acellular Tissue Graft Augmentation in Acute Achilles Tendon Ruptures. *The Journal of Foot and Ankle Surgery*, 47(1):8-12, January 2008. 20

[76] Shane Hollawell and William Baione. Chronic Achilles Tendon Rupture Reconstructed With Achilles Tendon Allograft and Xenograft Combination. *The Journal of Foot and Ankle Surgery*, 54(6):1146-1150, November 2015. 20

[77] Anthony Herbert, Gemma L. Jones, Eileen Ingham, and John Fisher. A biomechanical characterisation of acellular porcine super flexor tendons for use in anterior cruciate ligament replacement: Investigation into the effects of fat reduction and bioburden reduction bioprocesses. *Journal of Biomechanics*, 48(1):22-29, January 2015. 20

[78] J. Palmer Branch. A Tendon Graft Weave Using an Acellular Dermal Matrix for Repair of the Achilles Tendon and Other Foot and Ankle Tendons. *The Journal of Foot and Ankle Surgery*, 50(2):257-265, March 2011. 20

[79] Diana Gaspar, Kyriakos Spanoudes, Carolyn Holladay, Abhay Pandit, and Dimitrios Zeugolis. Progress in cell-based therapies for tendon repair. *Advanced Drug Delivery Reviews*, 84:240-256, April 2015. 21

[80] Cansın Güngörmüş, Dürdane Kolankaya, and Erkin Aydin. Histopathological and biomechanical evaluation of tenocyte seeded allografts on rat Achilles tendon regeneration. *Biomaterials*, 51:108-118, 2015. 21
[81] S. J. Kew, J. H. Gwynne, D. Enea, M. Abu-Rub, A. Pandit, D. Zeugolis, R. A. Brooks, N. Rushton, S. M. Best, and R. E. Cameron. Regeneration and repair of tendon and ligament tissue using collagen fibre biomaterials. *Acta Biomater*, 7(9):3237-3247, September 2011. 21, 28

[82] Mickey D. Stapp. Implantable materials and grafts in tendon surgery. *Reconstructive surgery of the foot and the leg*, pages 289-292, 1993. 21

[83] Roger D. Williams and Spencer F. August. Experimental evaluation of a Teflon tendon prosthesis. *The American Journal of Surgery*, 107(6):913-916, 1964. 21

[84] M. Holtz, M. L. Midenberg, and S. E. Kirschenbaum. Utilization of a silastic sheet in tendon repair of the foot. *J Foot Surg*, 21(4):253-259, 1982. 21

[85] Matthew C. Morrey, Jonathan D. Barlow, Matthew P. Abdel, and Arlen D. Hanssen. Synthetic Mesh Augmentation of Acute and Subacute Quadriceps Tendon Repair. *Orthopedics*, 39(1):e9-e13, January 2016. 21

[86] Shalimar Abdullah. Usage of synthetic tendons in tendon reconstruction. *BMC Proceedings*, 9(Suppl 3):A68, 2015. 21

[87] Alan Melvin, Alan Litsky, Joel Mayerson, David Witte, David Melvin, and Natalia Juncosa-Melvin. An artificial tendon with durable muscle interface. *J. Orthop. Res.*, 28(2):218-224, 2010. 21

[88] D. H. Jenkins and B. McKibbin. The role of flexible carbon-fibre implants as tendon and ligament substitutes in clinical practice. A preliminary report. *J Bone Joint Surg Br*, 62-B(4):497-499, November 1980. 22

[89] D. P. M. Corey. *Tendon grafts and implants*. 1988. 22

[90] J. M. Roberts, G. L. Goldstrohm, T. D. Brown, and D. C. Mears. Comparison of unrepaired, primarily repaired, and polyglactin mesh-reinforced Achilles tendon lacerations in rabbits. *Clin. Orthop. Relat. Res.*, (181):244-249, December 1983. 22

[91] C. B. Howard, B. McKibbin, and Z. A. Ralis. The use of Dexon as a replacement for the calcaneal tendon in sheep. *Bone & Joint Journal*, 67-B(2):313-316, March 1985. 22

[92] M. D. Liem, H. G. Zegel, F. C. Balduini, M. L. Turner, J. M. Becker, and A. Caballero-Saez. Repair of Achilles tendon ruptures with a polylactic acid implant: assessment with MR imaging. *AJR Am J Roentgenol*, 156(4):769-773, April 1991. 22

[93] H. Alexander, A. B. Weiss, and J. R. Parsons. Absorbable polymer-filamentous carbon composites—a new class of tissue scaffolding materials. *Aktuelle Probl Chir Orthop*, 26:78-91, 1983. 22

[94] Ahmed Shoaib and Viren Mishra. Surgical repair of symptomatic chronic achilles tendon rupture using synthetic graft augmentation. *Foot and Ankle Surgery*, September 2016. 22

[95] Gunay Verdiyeva, Kiron Koshy, Natalia Glibbery, Haroon Mann, and Alexander M. Seifalian. Tendon Reconstruction with Tissue Engineering Approach—A Review. *Journal of Biomedical Nanotechnology*, 11(9):1495-1523, September 2015. 23
[96] Wing Yin Tong, Wei Shen, Connie W.F. Yeung, Ying Zhao, Shuk Han Cheng, Paul K. Chu, Danny Chan, Godfrey C.F. Chan, Kenneth M.C. Cheung, Kelvin W.K. Yeung, and Yun Wah Lam. Functional replication of the tendon tissue microenvironment by a bioimprinted substrate and the support of tenocytic differentiation of mesenchymal stem cells. *Biomaterials*, 33(31):7686-7698, November 2012. 23, 24, 28

[97] Su A Park, In Ae Kim, Yong Jae Lee, Ji Won Shin, Chong-Rak Kim, Jeong Koo Kim, Young-Il Yang, and Jung-Woog Shin. Biological responses of ligament fibroblasts and gene expression profiling on micropatterned silicone substrates subjected to mechanical stimuli. *Journal of Bioscience and Bioengineering*, 102(5):402-412, November 2006. 24

[98] James HC Wang, Fengyan Jia, Thomas W Gilbert, and Savio LY Woo. Cell orientation determines the alignment of cell-produced collagenous matrix. *Journal of biomechanics*, 36(1):97-102, 2003. 24

[99] F.C.M.J.M. van Delft, F.C. van den Heuvel, W.A. Loesberg, J. te Riet, P. Schön, C.G. Figdor, S. Speller, J.J.W.A. van Loon, X.F. Walboomers, and J.A. Jansen. Manufacturing substrate nano-grooves for studying cell alignment and adhesion. *Microelectronic Engineering*, 85(5-6):1362-1366, May 2008. 24

[100] Leslie Mertz. What Is Biocompatibility?: A New Definition Based on the Latest Technology. *IEEE Pulse*, 4(4):14-15, July 2013. 24

[101] Use of International Standard ISO 10993-1, "Biological evaluation of medical devices - Part 1: Evaluation and testing within a risk management process". Technical report, U.S. Department of Health and Human Services Food and Drug Administration Center for Devices and Radiological Health, June 2016. 24

[102] Dietmar W. Hutmacher. Scaffolds in tissue engineering bone and cartilage. *Bi materials*, 21(24):2529-2543, 2000. 24

[103] S Badylak, D Freytes, and T Gilbert. Extracellular matrix as a biological scaffold material: Structure and function. *Acta Biomaterialia*, 5(1):1-13, January 2009. 24

[104] Gilbert, Stewart-Akers, Simmons-Byrd, and Badylak. Degradation and remodeling of small intestinal submucosa in canine achilles tendon repair. *J Bone Joint Surg*, pages 621-630, 2007. 24

[105] J. F. Cavallaro, P. D. Kemp, and K. H. Kraus. Collagen fabrics as biomaterials. *Biotech-nol. Bioeng.*, 43(8):781-791, April 1994. 25, 28, 29, 30

[106] Bin Dong, Olivier Arnoult, Meghan E. Smith, and Gary E. Wnek. Electrospinning of Collagen Nanofiber Scaffolds from Benign Solvents. *Macromolecular Rapid Communications*, 30(7):539-542, April 2009. 25, 30

[107] Nuanchan Choktaweesap, Kunawan Arayanarakul, Duangdao Aht-Ong, Chidchanok Meechaisue, and Pitt Supaphol. Electrospun gelatin fibers: effect of solvent system on morphology and fiber diameters. *Polymer journal*, 39(6):622-631, 2007. 25

[108] Yu-Qin Wan, Ji-Huan He, Jian-Yong Yu, and Yue Wu. Electrospinning of high-molecule PEO solution. *Journal of Applied Polymer Science*, 103(6):3840-3843, March 2007. 26
[109] Jamil A. Matthews, Gary E. Wnek, David G. Simpson, and Gary L. Bowlin. Electro-spinning of Collagen Nanofibers. *Biomacromolecules*, 3(2):232–238, March 2002.
26

[110] Jamil A. Matthews, Eugene D. Boland, Gary E. Wnek, David G. Simpson, and Gary L. Bowlin. Electrospinning of collagen type II: a feasibility study. *Journal of bioactive and compatible polymers*, 18(2):125–134, 2003. 26

[111] Jochen Bürck, Stefan Heissler, Udo Geckle, Mohammad Fotouhi Ardakani, Reinhard Schneider, Anne S. Ulrich, and Murat Kazanci. Resemblance of Electrospun Collagen Nanofibers to Their Native Structure. *Langmuir*, 29(5):1562–1572, February 2013. 26, 29, 30

[112] Dimitrios I. Zeugolis, Shih T. Khew, Elijah S.Y. Yew, Andrew K. Ekaputra, Yen W. Tong, Lin-Yue L. Yung, Dietmar W. Hutmacher, Colin Sheppard, and Michael Raghunath. Electro-spinning of pure collagen nano-fibres – Just an expensive way to make gelatin? *Biomaterials*, 29(15):2293–2305, May 2008. 26, 30

[113] Anh-Vu Do, Behnoush Khorsand, Sean M. Geary, and Aliasger K. Salem. 3d Printing of Scaffolds for Tissue Regeneration Applications. *Advanced Healthcare Materials*, 4(12):1742–1762, August 2015. 26

[114] Jason A. Inzana, Diana Olvera, Seth M. Fuller, James P. Kelly, Olivia A. Graeve, Edward M. Schwarz, Stephen L. Kates, and Hani A. Awad. 3d printing of composite calcium phosphate and collagen scaffolds for bone regeneration. *Biomaterials*, 35(13):4026–4034, April 2014. 26

[115] Daniela Filipa Duarte Campos, Andreas Blaeser, Kate Buellesbach, Kshama Shree Sen, Weiwei Xun, Walter Tillmann, and Horst Fischer. Bioprinting Organotypic Hydrogels with Improved Mesenchymal Stem Cell Remodeling and Mineralization Properties for Bone Tissue Engineering. *Advanced Healthcare Materials*, 5(11):1336–1345, June 2016. 26, 31

[116] Sean V. Murphy, Aleksander Skardal, and Anthony Atala. Evaluation of hydrogels for bio-printing applications. *Journal of Biomedical Materials Research Part A*, 101A(1):272–284, January 2013. 26, 28

[117] Hong Wei Ouyang, James CH Goh, Ashvin Thambyah, Swee Hin Teoh, and Eng Hin Lee. Knitted poly-lactide-co-glycolide scaffold loaded with bone marrow stromal cells in repair and regeneration of rabbit achilles tendon. *Tissue engineering*, 9(3):431–439, 2003. 26, 27

[118] Dan Deng, Wenbo Wang, Bin Wang, Peihua Zhang, Guangdong Zhou, Wen Jie Zhang, Yilin Cao, and Wei Liu. Repair of achilles tendon defect with autologous ascs engineered tendon in a rabbit model. *Biomaterials*, 35(31):8801–8809, 2014. 26, 27

[119] Yilin Cao, Yongtao Liu, Wei Liu, Qingxin Shan, Samuel D. Buonocore, and Lei Cui. Bridging tendon defects using autologous tenocyte engineered tendon in a hen model. *Plast. Reconstr. Surg.*, 110(5):1280–1289, October 2002. 26

[120] Gregory H Altman, Rebecca L Horan, Helen H Lu, Jodie Moreau, Ivan Martin, John C Richmond, and David L Kaplan. Silk matrix for tissue engineered anterior cruciate ligaments. *Biomaterials*, 23(20):4131–4141, 2002. 27
[121] Hongbin Fan, Haifeng Liu, Eugene JW Wong, Siew L Toh, and James CH Goh. In vivo study of anterior cruciate ligament regeneration using mesenchymal stem cells and silk scaffold. *Biomaterials*, 29(23):3324-3337, 2008. 27

[122] J Nowotny, D Aibibu, J Farack, U Nimtschke, M Hild, M Gelinsky, P Kasten, and Ch Cherif. Novel fiber-based pure chitosan scaffold for tendon augmentation: biomechanical and cell biological evaluation. *Journal of Biomaterials Science, Polymer Edition*, 27(10):917-936, 2016. 27, 28

[123] Jiro Namba, Kozo Shimada, Masanobu Saito, Tsuyoshi Murase, Hideaki Yamada, and Hideki Yoshikawa. Modulation of peritendinous adhesion formation by alginate solution in a rabbit flexor tendon model. *Journal of Biomedical Materials Research Part B: Applied Biomaterials*, 80(1):273-279, 2007. 28

[124] Julie Glowacki and Shuichi Mizuno. Collagen scaffolds for tissue engineering. *Biopolymers*, 89(5):338-344, May 2008. 28

[125] Malgorzata Barczyk, Sergio Carracedo, and Donald Gullberg. Integrins. *Cell Tissue Res*, 339:269-280, 2010. 28

[126] Davide Enea, Frances Henson, Simon Kew, John Wardale, Alan Getgood, Roger Brooks, and Neil Rushton. Extruded collagen fibres for tissue engineering applications: effect of crosslinking method on mechanical and biological properties. *Journal of Materials Science: Materials in Medicine*, 22(6):1569-1578, June 2011. 28

[127] George D. Pins, David L. Christiansen, Raj Patel, and Frederick H. Silver. Self-assembly of collagen fibers. Influence of fibrillar alignment and decorin on mechanical properties. *Biophysical journal*, 73(4):2164, 1997. 28, 29

[128] M. C. Wang, G. D. Pins, and F. H. Silver. Collagen fibres with improved strength for the repair of soft tissue injuries. *Biomaterials*, 15(7):507-512, June 1994. 29

[129] Edwina S. Lai, Claire M. Anderson, and Gerald G. Fuller. Designing a tubular matrix of oriented collagen fibrils for tissue engineering. *Acta Biomaterialia*, 7(6):2448-2456, June 2011. 29

[130] M. G. Dunn, P. N. Avasarala, and J. P. Zawadsky. Optimization of extruded collagen fibers for ACL reconstruction. *J. Biomed. Mater. Res.*, 27(12):1545-1552, December 1993. 29, 30

[131] Y. P. Kato, D. L. Christiansen, R. A. Hahn, S. J. Shieh, J. D. Goldstein, and F. H. Silver. Mechanical properties of collagen fibres: a comparison of reconstituted and rat tail tendon fibres. *Biomaterials*, 10(1):38-42, January 1989. 29

[132] Kevin G. Cornwell and George D. Pins. Discrete crosslinked fibrin microthread scaffolds for tissue regeneration. *Journal of Biomedical Materials Research Part A*, 82A(1):104-112, July 2007. 29, 30

[133] Lei Huang, Karthik Nagapudi, Robert P.Apkarian, and Elliot L. Chaikof. Engineered collagen-PEO nanofibers and fabrics. *Journal of Biomaterials Science, Polymer Edition*, 12(9):979-993, January 2001. 28, 29
[134] Dimitrios I. Zeugolis, Gordon R. Paul, and Geoffrey Attenburrow. Cross-linking of extruded collagen fibers-A biomimetic three-dimensional scaffold for tissue engineering applications. *Journal of Biomedical Materials Research Part A*, 89A(4):895-908, June 2009. 29

[135] W E Teo and S Ramakrishna. A review on electrospinning design and nanofibre assemblies. *Nanotechnology*, 17(14):R89-R106, July 2006. 29

[136] D.I. Zeugolis, R.G. Paul, and G. Attenburrow. The influence of a natural cross-linking agent (Myrica rubra) on the properties of extruded collagen fibres for tissue engineering applications. *Materials Science and Engineering: C*, 30(1):190-195, January 2010. 29

[137] Kyong Su Rho, Lim Jeong, Gene Lee, Byoung-Moo Seo, Yoon Jeong Park, Seong-Doo Hong, Sangho Roh, Jae Jin Cho, Won Ho Park, and Byung-Moo Min. Electrospinning of collagen nanofibers: Effects on the behavior of normal human keratinocytes and early-stage wound healing. *Biomaterials*, 27(8):1452-1461, March 2006. 30

[138] A. Szentivanyi, U. Assmann, R. Schuster, and B. Glasmacher. Production of bio-hybrid protein/PEO scaffolds by electrospinning. *Materialwissenschaft und Werkstofftechnik*, 40(1-2):65-72, January 2009. 30

[139] Anuradha Elamparithi, Alan M. Punnoose, and Sarah Kuruvilla. Electrospun type 1 collagen matrices preserving native ultrastructure using benign binary solvent for cardiac tissue engineering. *Artificial Cells, Nanomedicine, and Biotechnology*, pages 1-8, May 2015. 30

[140] T Xu, C Gregory, P Molnar, X Cui, S Jalota, S Bhaduri, and T Boland. Viability and electrophysiology of neural cell structures generated by the inkjet printing method. *Biomaterials*, March 2006. 30

[141] Mitchell R. Ladd, Sang Jin Lee, Joel D. Stitzel, Anthony Atala, and James J. Yoo. Co-electrospun dual scaffolding system with potential for muscle-tendon junction tissue engineering. *Biomaterials*, 32(6):1549-1559, February 2011. 30

[142] Kai-Feng Lin, Shu He, Yue Song, Chun-Mei Wang, Yi Gao, Jun-Qin Li, Peng Tang, Zheng Wang, Long Bi, and Guo-Xian Pei. Low-Temperature Additive Manufacturing of Biomimic Three-Dimensional Hydroxyapatite/Collagen Scaffolds for Bone Regeneration. *ACS Applied Materials & Interfaces*, 8(11):6905-6916, March 2016. 30

[143] Hyeong Jin Lee, Yong Bok Kim, Seung Hyun Ahn, Ji-Seon Lee, Chul Ho Jang, Hyeon Yoon, Wook Chun, and Geun Hyung Kim. A New Approach for Fabricating Collagen/ECM-Based Bioinks Using Preosteoblasts and Human Adipose Stem Cells. *Advanced Healthcare Materials*, 4(9):1359-1368, June 2015. 30

[144] Kai Sun, Ruixin Li, Wenxue Jiang, Yufu Sun, and Hui Li. Comparison of three-dimensional printing and vacuum freeze-dried techniques for fabricating composite scaffolds. *Biochemical and Biophysical Research Communications*, 477(4):1085-1091, September 2016. 30, 31

[145] B. Derby. Printing and Prototyping of Tissues and Scaffolds. *Science*, 338(6109):921-926, November 2012. 31
[146] Tao Xu, Kyle W Binder, Mohammad Z Albanna, Dennis Dice, Weixin Zhao, James J Yoo, and Anthony Atala. Hybrid printing of mechanically and biologically improved constructs for cartilage tissue engineering applications. *Biofabrication*, 5(1):015001, March 2013. 31

[147] Anja Lode, Michael Meyer, Sophie Brüggemeier, Birgit Paul, Hagen Baltzer, Michaela Schröpfer, Claudia Winkelmann, Frank Sonntag, and Michael Gelin sky. Additive manufacturing of collagen scaffolds by three-dimensional plotting of highly viscous dispersions. *Biofabrication*, 8(1):015015, February 2016. 31

[148] C. Z. Liu, Z. D. Xia, Z. W. Han, P. A. Hulley, J. T. Triffitt, and J. T. Czernuszka. Novel 3d collagen scaffolds fabricated by indirect printing technique for tissue engineering. *Journal of Biomedical Materials Research Part B: Applied Biomaterials*, 85(2):519-528, 2008. 31

[149] SeungHyun Ahn, SuYeon Lee, Youngseok Cho, Wook Chun, and GeunHyung Kim. Fabrication of three-dimensional collagen scaffold using an inverse mould-leaching process. *Bioprocess and Biosystems Engineering*, 34(7):903-911, September 2011. 31

[150] MyungGu Yeo, Ji-Seon Lee, Wook Chun, and Geun Hyung Kim. An Innovative Collagen-Based Cell-Printing Method for Obtaining Human Adipose Stem Cell-Laden Structures Consisting of Core-Sheath Structures for Tissue Engineering. *Biomacromolecules*, 17(4):1365-1375, April 2016. 31

[151] Su A Park, In Ae Kim, Yong Jae Lee, Ji Won Shin, Chong-Rak Kim, Jeong Koo Kim, Young-Il Yang, and Jung-Woog Shin. Biological responses of ligament fibroblasts and gene expression profiling on micropatterned silicone substrates subjected to mechanical stimuli. *Journal of Bioscience and Bioengineering*, 102(5):402-412, November 2006. 31

[152] SangJun Moon, Syed K. Hasan, Young S. Song, Feng Xu, Hasan Onur Keles, Fahim Manzur, Sohan Mikkilineni, Jong Wook Hong, Jiro Nagatomi, Edward Haeggstrom, and others. Layer by layer three-dimensional tissue epitaxy by cell-laden hydrogel droplets. *Tissue Engineering Part C: Methods*, 16(1):157-166, 2009. 31

[153] Edward Kang, Gi Seok Jeong, Yoon Young Choi, Kwang Ho Lee, Ali Khademhosseini, and Sang-Hoon Lee. Digitally tunable physicochemical coding of material composition and topography in continuous microfibres. *Nature Materials*, 10(11):877-883, September 2011. 31

[154] Tyler K Merceron, Morgan Burt, Young-Joon Seol, Hyun-Wook Kang, Sang Jin Lee, James J Yoo, and Anthony Atala. A 3d bioprinted complex structure for engineering the muscle-tendon unit. *Biofabrication*, 7(3):035003, June 2015. 31
# Chapter 2

# Methods and materials

"Like so many things, it is not
what's outside, but what is inside
that counts."

The Merchant in *Aladdin*

## Abstract

In this chapter, the home-made experimental set-up of extrusion is first presented and its main parts are detailed. Moreover, we quantified the new device's performances and limits. Then, the physical bases of rheological and mechanical testing are explained and the experiments' parameters are given. In a second part, the materials used in the following studies are described: collagen type I, sodium alginate and collagen-alginate mixtures. The solutions' purification, preparation and their physico-chemical properties are reported here.

## Contents

<table><tbody><tr><td><strong>2.1 Methods</strong></td><td><strong>47</strong></td></tr><tr><td>2.1.1 Extrusion set-up</td><td>47</td></tr><tr><td>2.1.1.1 Key constituent part: the needle</td><td>48</td></tr><tr><td>2.1.1.2 Set-up performances and limits</td><td>52</td></tr><tr><td>2.1.1.3 Conclusion</td><td>59</td></tr><tr><td>2.1.2 Rheological testing of polymer solutions</td><td>60</td></tr><tr><td>2.1.2.1 Physical basis</td><td>60</td></tr><tr><td>2.1.2.2 Rheometer geometries</td><td>67</td></tr><tr><td>2.1.3 Mechanical testing of extruded threads</td><td>69</td></tr><tr><td><strong>2.2 Materials</strong></td><td><strong>71</strong></td></tr><tr><td>2.2.1 Collagen solutions</td><td>71</td></tr><tr><td>2.2.1.1 Collagen type I and its fibrillogenesis</td><td>71</td></tr><tr><td>2.2.1.2 Chemical properties</td><td>73</td></tr><tr><td>2.2.1.3 Collagen solutions preparation</td><td>75</td></tr><tr><td>2.2.1.4 Collagen extraction and purification</td><td>75</td></tr><tr><td>2.2.1.5 Collagen solutions dosage</td><td>76</td></tr><tr><td>2.2.1.6 Collagen solutions purity</td><td>77</td></tr><tr><td>2.2.1.7 Collagen solutions concentration</td><td>78</td></tr><tr><td>2.2.2 Model system: sodium alginate</td><td>79</td></tr><tr><td>2.2.2.1 Structure</td><td>79</td></tr><tr><td>2.2.2.2 Chemical properties</td><td>79</td></tr></tbody></table>
<table><tr><td>2.2.2.3 Solution preparation</td><td>82</td></tr><tr><td>2.2.3 Collagen-alginate mixtures</td><td>83</td></tr><tr><td>2.2.3.1 Chemical properties</td><td>83</td></tr><tr><td>2.2.3.2 Solution preparation</td><td>84</td></tr><tr><td><b>2.3 References</b></td><td><b>85</b></td></tr></table>
**2.1 Methods**

**2.1.1 Extrusion set-up**

As mentioned above in section 1.2.2, extrusion consists in forcing a material to pass through a die. This method allows either to produce threads as in the textile industry and in tissue engineering (3D printing) or to study instabilities as in the microfluidics field. To fulfill the present goal, we built an in-home extrusion apparatus. This tailored set-up is designed as follows:

<figure><img src="image_19.png" /><figcaption>Figure 2.1: Scheme of experimental set-up of extrusion</figcaption></figure>

A steel rod is clamped to a stepping motorized stage which imposes its velocity $v_p$ ranging from 0 to 2000 µm/s. A load cell LCM201 Series from OMEGA (Fig 2.1, dark blue rectangle) was calibrated in tension for compression ranging from 100 N to 500 N. It is connected to an on-line transducer in order to assess the force F applied on the syringe piston. During the extrusion of our complex fluids, forces of 10 to 50N are reached. A syringe (Terumo) of 1 mL contains the fluid to be extruded. It has an inner diameter equal to the piston diameter $D_p$ of 4.45 mm. Metallic needles, glass capillaries or PTFE tubings with characteristic length L and inner radius R are set at the syringe end.
A steel structure is fabricated to hold the syringe and its capillary vertically. It is composed of a cylindrical piece, which is shaped to fit tightly the syringe and a second piece to insert the optical glass cell (Hellma) containing the extrusion bath. This one can be removed so as to collect the extruded thread at any time during the process, while keeping the syringe in place. Holes are drilled on the last piece to create optical windows. The whole set-up is put inside a temperature controlled enclosure.

An optical mounting is added to observe the extrusion in real-time. A quasi-parallel beam is created with a light source (S) limited by a diaphragm placed at the focal length of a convergent lens. A polarizer (P) and an analyzer (A) are inserted so that the birefringence of the stressed fluid can be observed. The light beam illuminates the glass cell and the extrusion flow is followed thanks to a video camera equipped with a macro-objective and placed on the other side. Acquisition speeds of 600 frames/s can be reached. This optical system allows us to obtain extrusion images of a high contrast level as shown in Figure 2.2. It allows us to distinguish clearly small changes of the refractive index as well as other structural particularities.

<figure><img src="image_20.png" /><figcaption>Figure 2.2: Example of an extrusion in a solution bath at the needle exit. The black rectangle is the needle exit. It is possible to see the gelation front ("black" region) growing progressively from the flow outside. Scale bar = 200µm</figcaption></figure>

The motorized rod stage, the load cell signal and the camera are piloted by a program interface made by Olivier Ronsin to synchronize and control them easily.

### 2.1.1.1 Key constituent part: the needle

Our set-up allows us to control almost all the physical parameters of the extrusion such as the piston velocity, die diameter, temperature as well as chemical parameters by adjusting the collecting bath nature. Moreover, an other set-up part can be adapted to control more sensitive aspects of the extrusion as the die exit material. In fact, the die exit consists in an empty hose in which the fluid from the syringe will pass through. Along the die, this fluid will interact with the walls die. This interaction is governed by the hose material dimensions, composition and the fluid nature.
Thanks to our experimental set-up, the syringe exit can be changed and thus the material of the hose too. In the microfluidics field, typical set-up do not display such adaptability. Indeed, the polydimethylsiloxane (PDMS) is the major component used for the fabrication of devices to study flow physics. Even if some coating processings (gase-phase, plasma, sol-gel...) can be used to modify the PDMS surface nature, those strategies are not very convenient to implement [1]. In our case, hoses made of different materials (stainless steel, glass or PTFE) will be prepared and used as detailed in the following section.

### 2.1.1.1.1 Metallic needles

Sterile stainless steel needles with a normalized Luer nozzle were purchased from Sterican (BRAUN) and Fine-ject. Several gauges were available from 18G to 27G. A large gauge corresponds to a small diameter. Table 2.1 presents the needle characteristic dimensions supplied by the manufacturer.

<table><caption>Table 2.1: Characteristic dimensions of stainless steel needles</caption><thead><tr><td>Length (mm)</td><td>Outer diameter (μm)</td><td>Inner diameter (μm)</td><td>Gauge</td></tr></thead><tbody><tr><td>20</td><td>400</td><td>220</td><td>27G</td></tr><tr><td>25</td><td>450</td><td>240</td><td>26G</td></tr><tr><td>25</td><td>500</td><td>250</td><td>25G</td></tr><tr><td>30</td><td>600</td><td>390</td><td>23G</td></tr><tr><td>30</td><td>700</td><td>460</td><td>22G</td></tr><tr><td>50</td><td>800</td><td>500</td><td>21G</td></tr><tr><td>120</td><td>800</td><td>500</td><td>21G</td></tr><tr><td>40</td><td>900</td><td>650</td><td>20G</td></tr><tr><td>40</td><td>1100</td><td>780</td><td>19G</td></tr></tbody></table>

<figure><img src="image_21.png" /><figcaption>Figure 2.3: Example of a 23G needle tip, Scale bar = 250μm</figcaption></figure>

These hypodermic needles exit present almost 3 angles (2.3). The aim of this geometry is to ease penetration with less painful puncture for example during an injection or a blood sampling. In our case, this makes the extrusion study at the exit more complicated. This is why, needles were recustomized in the laboratory in order to remove the sharp lancet point and obtain a blunt tip.

For this, the needle is fixed to a support and mounted on the saw machine to cut the sharp lancet point with the diamond wire (Fig 2.4,1). The cut-needle is placed into a glass tube of the same height and sealed on its basis on a glass slide (Fig 2.4,2). The whole system is put in the heat chamber at 110 °C. A mixture of beeswax and rosin previously prepared and heated at the same temperature is filled into the glass tube (Fig 2.4,3). It is cooled at room temperature and the excess of resin is removed. The glass tube containing the needle can be handled and place in contact with a polishing lathe (Fig 2.4,4) coated with silicon carbide powder (abrasive element). The needle extremity is polished until
<figure><img src="image_22.png" /><figcaption>Figure 2.4: Protocol of blunt needles fabrication. 1) Cutting the sharp lancet point, 2) Needle installation for 3) Resine embedding, 4) Exit polishing, 5) Acetone cleaning to remove resin remains, 6) Needle with blunt tip</figcaption></figure>

its edges are smooth, which is controlled by optical microscopy. Finally, the glass tube is removed and the blunt needle is washed with acetone to bring out resin remains (Fig. 2.4,5). The needle obtained is finally ready for use (Fig 2.4,6).

Similar procedures will be adapted to create glass capillaries of different diameters and structures as shown in the following part.

### 2.1.1.1.2 Glass capillaries

Borosilicate glass capillaries were purchased from CM Scientific. The outer diameter was chosen in order to be compatible with the inner diameter of stainless steel needles. In fact, the plastic extremity (Luer nozzle) of the needle is essential to be mounted on the syringe. The idea is to keep this part and to insert the glass capillary inside. Therefore glass capillaries with inner diameter of 200, 300 and 500 µm, outer diameter of respectively 330, 400 and 700 µm, were chosen.

Needles are cut at 1 cm of the input extremity with diamond wire saw (Fig 2.5 1)). Capillaries are cut to the desired length with a diamond-point and put inside the needle. The cut extremity was inside the needle and the manufactured one outside to have a clean blunt exit. The connecting point is glued and sealed horizontally with araldite (Fig 2.5 2)). The whole system is dried at room temperature overnight.

To study the input effect during the extrusion, combination of glass capillaries were fabricated. As shown in Figure 2.5 a) and b), two capillaries of different diameters were affixed to create either a branch narrowing or expansion.

Other strategies were performed to study the effects of the capillary inner surface or of the exit edges. For the first one, glass capillaries were put close to a hydrofluoric acid (HF) solution at 1% during several hours. By capillary action, the HF solution goes up to a certain height, given by the Jurin's law and react with the borosilicate glass. The idea is to control the height of the treatment and to make the inner surface more porous in order to change the fluid/capillary interaction. However, no clear effects were observed and this approach was not followed up.

For the second one, to study the effect of the sharp edges of the capillary extremity, MF-
<figure><img src="image_23.png" /><figcaption>Figure 2.5: Protocol of glass needles fabrication. 1) Cutting the sharp lancet point, 2) Capillary installation and sealing with araldite, different configurations of a) narrowing and b) expansion capillaries</figcaption></figure>

microforge at the laboratory *Matière et Systèmes Complexes, Paris* 7 was used to fire-polish those singularities. Briefly, the glass capillary tip is heated close to the resistive electrical heating element and brought to a temperature where its surface melts. As a consequence, the glass tip becomes smooth. However, the heating seems to be too intense for those capillaries as it leads to a diameter narrowing. Some extrusions were tested with those new capillary but will not be detailed in the following chapter.

<figure><img src="image_24.png" /><figcaption>Figure 2.6: Microscope images of a) 200 µm capillary extremity without any treatment, b) Fire-polished 200 µm capillary tip. Scale bar = 100 µm</figcaption></figure>

To go further in the study of the fluid/material interaction, the question of the wetting of the surface arises. Indeed, stainless steel needles and borosilicate glass are both hydrophilic materials. The idea was to study the extrusion of the fluid against a hydrophobic surface in order to lower the wall frictions and for this, we chose Teflon.

### 2.1.1.1.3 PTFE tubing

Polytetrafluoroethylene (PTFE, Teflon) tubing purchased from Adtech Polymer Engineering were used. The same procedure as for glass capillaries (Fig. 2.5) was performed to obtain capillary of PTFE tubing. The PTFE tubing chosen has an inner diameter of 300 µm and an outer diameter of 760 µm. As the araldite does not stick to the Teflon, the connecting point was sealed with Parafilm M.

An unexpected property of the PTFE was a great surprise during the first test with this type of capillary. In fact, the contrast obtained with the optical system is clearly enhanced
when PTFE tubing is used, making details easier to distinguish the tubing. This is due to the refractive index of the PTFE. In fact, it is close to the one of the water. For the same wavelength of 589.29 nm and temperature (20 °C), the refractive index of PTFE is 1.35 – 1.38 and 1.33 for the water. Compared to the borosilicate refractive index (1.47), PTFE tubing thus ensures a better insight of what is happening in the capillary, and helps to better understand extrusion phenomena.

### 2.1.1.2 Set-up performances and limits

Our experimental set-up of extrusion allows us to impose the velocity of the piston, i.e. the throughput. As for a capillary rheometer, the force applied on the piston is measured, in our case, with the load cell presented previously. The steady state is achieved when the force signal (monitored in real-time) becomes constant. In the following part, the different conditions to reach a steady flow regime will be described.

#### 2.1.1.2.1 Friction behavior

During the extrusion experiments, force signal oscillations are observed for a restricted range of $v_p$ values. This was first attributed to extrusion instability akin to the spurt-flow instability which manifests itself by periodic relaxation oscillations of the force. However, no flow instability was discerned at the capillary exit. To assess the origin of those signal fluctuations, we performed extrusion with an empty syringe (air), or a syringe filled with either distilled water or Rhodorsil® Oils 47V1000 (polydimethylsiloxane oils) at different piston speeds and registered the respective force signal. No capillary was mounted on the syringe. The results obtained for piston velocities of 100 and 500 µm/s are presented in the following Figure 2.7:

<figure><img src="image_25.png" /><figcaption>Figure 2.7: Force signals registered during the extrusion of a) Air, b) Water and c) Oil, made without capillaries at piston velocity of 100 µm/s (dotted blue) and 500 µm/s (purple).</figcaption></figure>

For all conditions, the force signal is constant at about 0.5 N at $v_p=500$ µm/s. At 100 µm/s, periodic oscillations of 1N are clearly observed even with the empty syringe. Those fluctuations are composed of two phases: a first one characterised by a slow force relaxation increase and a second one by an instantaneous decline. This can only be due to the piston frictions against the walls of the syringe. In fact, the piston rod extremity (in black, Fig. 2.8, 1)) composed of a thermoplastic latex-free elastomer is ultra-flexible to perfectly fit the inner diameter of the syringe and to create a tight seal. It plays the role of a spring. [2]

It is known that stick-slip can be suppressed either by stiffening the system or by lubricating i.e by replacing solid friction (with a threshold) by a viscous friction [3]. The inner surface of the syringe is lined by silicone oil. Despite the presence of this lubricant,
<figure><img src="image_26.png" /><figcaption>Figure 2.8: Image of 1) Terumo piston and 2) Home-made Dural piston. Scale bar = 1 cm.</figcaption></figure>

stick of this piston extremity may occur. Thus, when the piston sticks to the syringe inner surface, the force signal increases and when it slips, the force signal decreases etc. This phenomenon happens for velocities ranging typically from 25 to almost 150 µm/s. As a consequence, extrusion study is not sufficiently reliable for this range of piston velocity. To assess those velocities and use the entire available velocity range, the second solution of stiffening the system is selected. To do so, a home-made Dural piston is produced (Fig. 2.8, 2)). This new piston is used for extrusion of alginate solutions of low concentration (≤ 4 % (w/w)). For higher concentrations of alginate or for collagen solutions, the commercial piston is kept because no stick-slip is observed and furthermore in the second case sterility must be ensured.

2.1.1.2.2 Steady flow regime conditions

In order to study the different flow instabilities at the die exit, a steady state regime must be reached. In the following part, the different conditions to reach this regime will be described.

<figure><img src="image_27.png" /><figcaption>Figure 2.9: Scheme of the { syringe + capillary} system and the characteristic variables. A syringe is filled with a volume V of a solution with a density ρ. During the time t and t + dt the piston (left dark gray) of section S_p advances at the velocity v_p and forces the solution at the pressure ΔP to pass through the capillary of section S and exit at the velocity v̄</figcaption></figure>

Figure 2.9 displays the variables of our system. A global characteristic of the exit flow rate in the capillary is the average velocity $\bar{v}$ defined from the volumic throughput Q and related by: $\bar{v} = Q/(\pi \times R^2)$. As previously mentioned, the velocity $v_p$ is imposed by the piston motion. However, due to the finite compressibility of the syringe and the fluid, $\bar{v}$ will be not imposed instantaneously when imposing $v_p$. Indeed, a transient regime occurs during a time $\tau$ before the steady state is reached. During the time $\tau$, a significant
portion of the fluid could be wasted. It is thus of practical importance to understand the parameters affecting $\tau$.

Between $t$ and $t+dt$, the mass entering in the capillary: $\rho v_p S_p dt$ is not equal to the mass emerging from the capillary: $\rho \bar{v} S dt$ as given by the continuity principle. The volume change of the fluid under pressure in the syringe must be taken into account and is written as $\frac{d(\rho V)}{dt}$. Moreover, as the capillary volume is small compared to the syringe volume, its compressibility will not be significant. All the parameters are defined as described in Figure 2.9. It can thus be written:

$$ \rho \times v_p \times S_p = \rho \times \bar{v} \times S + \frac{d(\rho V)}{dt} \quad (2.1) $$

We are seeking to express the velocity $\bar{v}$. The steady state solution is known and given by:

$$ \bar{v}(\infty) = v_p \times \frac{S_p}{S} $$

During $dt$, the energy due to the compression is stored either by creating volume or by compressing the fluid containing air bubbles. This is why, the syringe compliance $\chi_{syringe}$ and the fluid compressibility $\chi_{fluid}$ can be defined as:

$$ \begin{aligned} \chi_{syringe} &= \frac{1}{V} \times \frac{dV}{dP} \\ \chi_{fluid} &= \frac{1}{\rho} \times \frac{d\rho}{dP} \end{aligned} $$

Thus, $\rho V$ may vary with an increase of $\Delta P$ and we can write:

$$ \begin{aligned} \frac{d(\rho V)}{dt} &= \rho \frac{dV}{dt} + V \frac{d\rho}{dt} \\ &= \rho V \times (\chi_{syringe} + \chi_{fluid}) \times \frac{d(\Delta P)}{dt} \end{aligned} $$

An effective compressibility is defined as: $\chi_{eff} = \chi_{syringe} + \chi_{fluid}$ and the equation 2.1 becomes:

$$ v_p \times S_p = \bar{v} \times S + V \times \chi_{eff} \times \frac{d(\Delta P)}{dt} $$

By introducing $\bar{v}(\infty)$:

$$ \bar{v} + \frac{V \times \chi_{eff}}{S} \times \frac{d(\Delta P)}{dt} = \bar{v}(\infty) \quad (2.2) $$

However, the velocity $\bar{v}$ depends on the pressure $\Delta P$ and is related to it via the relationship:

$$ \Delta P = R_H \times \bar{v} \times S $$

where $R_H$ is the hydraulic impedance for newtonian fluids and is a function of the throughput $Q$ hence of $\bar{v}$.
In the case of a Newtonian fluid, R<sub>H</sub> is constant and equation 2.2 becomes:

$$
\bar{v} + V \times \chi_{eff} \times R_H \times \frac{d(\bar{v})}{dt} = \bar{v}(\infty) \quad (2.3)
$$

We can thus define the system capacitance C as V × χ<sub>eff</sub> and deduce the time τ via a "RC"
circuit analogy:

$$
\tau = R_H \times C
$$

The solution of equation 2.3 is given by:

$$
\bar{v}(t) = \bar{v}(\infty) \times (1 - e^{-t/\tau}) \quad (2.4)
$$

**At low throughput,** we measured the characteristic time τ by recording signal force dur-
ing an extrusion as presented in Figure 2.10 (left).

<figure><img src="image_28.png" /><figcaption>Figure 2.10: Force signals (black) and their respective piston velocities (red) over the time obtained for alginate solution extrusions at: low throughput (left) and high throughput (right). For low velocities, the force signal becomes constant after a characteristic time τ (left) whereas it is immediate (right) at high velocities.</figcaption></figure>

At $v_p$= 100 µm/s, the transient regime is clearly observed and the steady state is reached after a time $\tau$ about 100s. At low speed, it is also possible to calculate the hydraulic resistance $R_H$.

- Hydraulic resistance R<sub>H</sub>

The hydraulic resistance is determined by the Hagen-Poiseuille equation in the case of a laminar flow and a Newtonian fluid. In our case, the Reynolds number Re is defined as:

$$
Re = \frac{\rho \bar{v} R}{\eta}
$$

with ρ the fluid density, η the dynamic viscosity in Pa.s. Typical values for our system are:
ρ=1000 kg/m³; v̄= few cm/s, R of 250 µm and η around 30 Pa.s. Thus, Re is about 10⁻⁴ ≪ 1,
which means that the fluid flow is laminar. The Hagen-Poiseuille equation can be used:

$$
R_H = \frac{8\eta L}{\pi R^4} \tag{2.5}
$$
where L is the capillary length of 3cm, R about 250 µm and η around 30 Pa.s. In this case, the hydraulic resistance is: R<sub>H</sub> = 5.8 × 10<sup>14</sup> Pa.s.m<sup>-3</sup>.

From the characteristic time τ and the hydraulic resistance values, we can estimate the effective compressibility of the fluid in the Newtonian regime.

– The effective compressibility χ<sub>eff</sub>

The compressibility of liquids is of the order of 10<sup>-10</sup> Pa at room temperature. We can estimate the effective compressibility χ<sub>eff</sub> from the different physical parameters of our set-up by:

$$
\chi_{eff} = \frac{\tau}{V_{tot} R_{H}}
$$

where τ is about 100 s, V<sub>tot</sub> the total volume considered of 0.2 mL = 2×10<sup>-7</sup> m<sup>-3</sup> and the hydraulic resistance R<sub>H</sub> previously calculated: 5.8 × 10<sup>14</sup> Pa.s.m<sup>-3</sup>. The effective com- pressibility is χ<sub>eff</sub> = 10<sup>-6</sup>Pa<sup>-1</sup> ≫ χ<sub>liquids</sub> . In microfluidics, this can be explained by the die deformability. In our case, the syringe deformability is negligible. Another possible explanation could be the presence of air in the liquid. In fact, some air bubbles may be entrapped in our solutions (Fig. 2.11). This raises the question of the effect of those air bubbles on the fluid compressibility.

<figure><img src="image_29.png" /><figcaption>Figure 2.11: Image of extrusion with a glass capillary. Air bubbles are entrapped in the alginate solution. The volume V of interest is represented by the dotted line rectangle. Scale bar = 200 µm</figcaption></figure>

We estimated the air bubbles volume V<sub>air</sub> in a sodium alginate solution passing in a glass capillary of 100 µm radius by measuring the bubbles diameter (Fig. 2.11, red circles) in a delimited solution volume V (Fig. 2.11, dotted rectangle). We obtained:

$$
V_{air} = 0.03 \times V
$$

The fluid compressibility is related to the air volume fraction φ<sub>air</sub> and the pressure P by:

$$
\chi_{fluid} = \frac{\phi_{air}}{P}
$$

where φ<sub>air</sub> = 3 × 10<sup>-2</sup> and P is considered as equal to the atmospheric pressure of 10<sup>5</sup> Pa. We thus obtain a fluid compressibility of 3 × 10<sup>-7</sup>Pa<sup>-1</sup> ≫ χ<sub>water</sub>.

Thus, at low throughputs, characteristic physical quantities as the hydraulic resistance and the effective compressibility were calculated. χ<sub>eff</sub> values seem to be overestimated and other parameters might be taken into account to explain this order of magnitude. Typical time of about 100 s are observed before reaching steady state at low velocities.

**At higher throughputs,** the syringe would be empty before reaching a steady state. Fortu- nately, due to the shear thinning (viscosity decreases while increasing shear rate) behavior
of our solutions, the hydraulic resistance $R_H$ decreases markedly as the throughput $Q$ increases. The steady state regime is reached quasi-instantaneously as observed in Figure 2.10 (right).

To conclude, steady state regime is reached after waiting almost 100s at low throughput and few seconds for high throughput. While this signal remains constant and for laminar flows, the average velocity $\bar{v}$ will be calculated as:

$$ \bar{v} = v_p \times \left( \frac{D_p}{2 \times R} \right)^2 \qquad (2.6) $$

### 2.1.1.2.3 Pressure drop

The pressure drop in a capillary rheometer is the difference between the pressure values measured respectively upstream (reservoir) and downstream (extrudate):

$$ \Delta P = P_{up} - P_{down} $$

$$ \text{with: } P_{up} = \frac{F - F_{friction}}{\pi R_{piston}^2} $$

$$ \text{and: } P_{down} = p_0 $$

where $p_0$ is the atmospheric pressure and $F_{friction}$ estimated as the force on the piston which remains when the solution is extruded without any capillary. The pressure drop is then assumed to be negligible.

<figure><img src="image_30.png" /><figcaption>Figure 2.12: Scheme of a cylinder pipe of length $L$ and radius $R$. The flow velocity is represented by the arrows and the wall shear rate $\sigma_w$ acting on the lateral wall.</figcaption></figure>

For example, in the case of a newtonian fluid the Hagen-Poiseuille relationship holds between $\Delta P$ and the throughput $Q$. In the following, we will make use of the average shear rate $\bar{\gamma}$ (rather than $Q$) and the wall shear stress $\sigma_w$ (rather than $\Delta P$). These quantities are defined as:

$$ \bar{\gamma} = \frac{\bar{v}}{R} = \frac{Q}{\pi R^3} $$
and, $\sigma_w$:

$$ \sigma_w = \frac{\Delta P R}{2 L} $$

The latter results from the momentum balance in steady flow (Fig. 2.12):

$$ (\Delta P + P_0)\pi R^2 - P_0 \pi R^2 - \sigma_w \times 2\pi RL = 0 $$

with R and L respectively the radius and length of the pipe.

The figure 2.13 displays the experimental rheological characteristics $\sigma_w(\bar{\gamma})$ obtained for ALG4 through glass capillaries of different radii.

<figure><img src="image_31.png" /><figcaption>Figure 2.13: Wall shear stress as a function of shear rate for glass capillaries of different radii: R = 100 (blue), 150 (green) and 250 (red) µm. The blue curve has been computed from the Cross law [?] extrapolated from the range of shear rates accessible in the cone-plate rheometer (shadowed zone)</figcaption></figure>

Since the range of shear rates relevant to the capillary rheometer is much larger than the one accessible for the cone plate rheometer, one may wonder whether it is legitimate to extrapolate the Cross law (well fitting the rheometer data) to the wider range. We have therefore calculated $\sigma_w(\bar{\gamma})$ for the Cross law. This means integrating:

$$ \dot{\gamma} = \frac{dv}{dr} $$

with:

$$ \begin{cases} \sigma(r) = & \frac{\Delta P r}{2 L} \\ \sigma(r) = & \eta(\dot{\gamma}) \times \dot{\gamma} \\ v(r) = & 0 \text{ no slip hypothesis} \end{cases} \qquad (2.7) $$
The integration must be performed numerically for the Cross law. However, it is interesting to have a look at the analytical solution corresponding to a "power law" fluid with η(γ) ~ γ^m-1 (Fig. 2.14).

<figure><img src="image_32.png" /><figcaption>Figure 2.14: Velocity profiles for the flow of a fluid obeying a power law. Orange: m = 1, Blue: m=0.8</figcaption></figure>

In this case: v(r) = vmax(1 - (r/R)α) with α = (3-m)/(2-m). For a newtonian fluid (m = 1), one recovers the Poiseuille parabolic profile (for which σw = η₀γ₀(γ/γ₀)ⁿ. For lower m values, (Fig. 2.14, blue), the profile is flat over a large part of the capillary, therefore exhibiting a central plug-like flow (Fig. 2.14). When solving for v(r) using the full rheological law (including the newtonian plateau) one finds that the profile evolves from the parabolic one when γ̄ ≪ γ₀ to the power law profile when γ̄ ≫ γ₀. Thus, from the profile v(r) one can compute γ̄, γw the shear rate at the wall under no slip condition, hence σw = γw × η(

$$ \dot{\gamma}_w \approx 5 \frac{\bar{\nu}}{R} $$

The pressure drop is thus quite well estimated as it gives a correct indication of pressure magnitude. Moreover, the last definition will allow us to compute the Weissenberg number for our extrusion experiments.

### 2.1.1.3 Conclusion

Our experimental set-up of extrusion is designed either to study extrusion instabilities origin and morphologies or to produce threads in a reproducible manner thanks to the stepping motorized rod stage. Even if the depth of focus is not sufficient to perform velocimetry measurements, the optical set-up allows us to observe with a high contrast the extrusion directly through the bath.

The particular strengths are also its adaptability as the die material and the bath can be changed, even during extrusion for the second one and its performance. Indeed, the set-up allows us to produce threads very rapidly, around several meters in few minutes.

Few set-up limitations were identified and yet analyzed in order to overcome those drawbacks before the extrusion flow studies. In fact, piston friction was observed and a new metallic one was made to remove this parasitic phenomenon. Air bubbles entrapped in the solution will be useful passive tracers in Chapter 4. The steady state regime was seen to be reached after almost 100s at low throughput. However, thanks to shear-thinning properties of our solutions, this time is close to zero at high throughput and avoids to have empty syringe before reaching the steady state. Thus, those limits are turned into
advantages and may not impact further studies. Finally, the pressure drop was estimated, giving us a correct indication of pressure magnitude.

## 2.1.2 Rheological testing of polymer solutions

### 2.1.2.1 Physical basis

Rheology is the part of the mechanics that studies elasticity and flow characteristics of a material. For the flow characteristics of a fluid, two types of experiments can be performed: the shear flow and the oscillatory shear flow that we will describe first. To characterize a solid, elongation tests are performed to assess physical quantities such as the Young Modulus.

#### 2.1.2.1.1 Shear flow

A simple case to define the different physical quantities at play in a fluid between two parallel plates subjected to a shear stress $\sigma$ is depicted in Figure 2.15:

<figure><img src="image_33.png" /><figcaption>Figure 2.15: Velocity profile for a single shear flow of a newtonian liquid between two parallel plates. The upper plate is moving at a constant velocity V. The shear rate $\dot{\gamma}$ is constant along the y-axis, and is equal to: $\dot{\gamma}(y) = \frac{\partial v_x(y)}{\partial y}$. The tangential shear stress $\sigma_{xy}$ is applied along the x-axis on a fluid layer surface whose normal vector $\mathbf{n}$ is along the y-axis.</figcaption></figure>

The lower plate is fixed, and the stress $\sigma$ (Pa) is applied on the upper plate which then moves at a speed V. This configuration is called planar Couette flow. In laminar flow regime, the liquid can be modeled by a succession of infinitely thin layers parallel to each other and parallel to the plates. The fluid deformation is carried out by successive sliding of the layers on each other by friction without matter transport from one layer to the other but with momentum transport. The boundary conditions under the "no slip" hypothesis well verified for simple fluids are $v_x(0) = 0$ and $v_x(e) = V$. The strain rate is defined as:

$$ \dot{\gamma} = \frac{d\gamma}{dt} $$

In the case of planar Couette flow with no wall slip, it reads:

$$ \dot{\gamma} = \frac{V}{e} $$

For an infinitely thin liquid layer, the local shear rate is defined as:

$$ \dot{\gamma}(y) = \frac{\partial v_x}{\partial y} $$
The stress $\sigma_{xy}$ is the tangential force per unit area applied on the liquid layer by the surrounding fluid. One defines the shear viscosity $\eta$ as $\eta = \sigma_{xy}/\dot{\gamma}$. Generally, $\eta$ depends on the shear rate, so that:

$$ \sigma_{xy} = \eta(\dot{\gamma}) \dot{\gamma} $$

For a newtonian fluid, the viscosity $\eta$ does not depend on the shear rate $\dot{\gamma}$ and the stress $\sigma_{xy}$ is proportional to the shear rate $\dot{\gamma}$. In the case of a simple shear flow, $\sigma_{xy}$ is constant over the thickness $e$ by mechanical equilibrium between the different fluid layers. When, as in the case for many fluids, $\sigma_{xy}$ is a monotonous increasing function of $\dot{\gamma}$, the shear rate is also constant over the thickness $e$ in steady state regime.

## Stress tensor

For a general flow (different from simple shear) and for more complex materials (viscoelastic), it is necessary to introduce the stress tensor which reveals all the forces applied on an element of fluid by the surrounding medium.

<figure><img src="image_34.png" /><figcaption>Figure 2.16: On the left: Stress components $\sigma_{xy}$, $\sigma_{yy}$ and $\sigma_{zy}$ applied on a surface $dS$ with the normal $\mathbf{n}$ oriented along the y-axis. On the right: Stress force $df$ acting on a surface $dS$ with normal vector $\mathbf{n}$ presents no specific orientation.</figcaption></figure>

Considering a surface element $dS$ placed horizontally: the normal $\mathbf{n}$ to the surface $dS$ is along the y-axis (Fig 2.16). A stress $\sigma$ applied to this surface can be divided into three components along the x, y and z axes: there are respectively $\sigma_{xy}$, $\sigma_{yy}$ and $\sigma_{zy}$. More generally, the notation $\sigma_{ij}$ is used, where $i$ is the stress direction and $j$ the direction of the normal $\mathbf{n}$ to the surface $dS$. The tangential stresses are $\sigma_{xy}$ and $\sigma_{zy}$ and the normal stress is $\sigma_{yy}$. For a force $df$ of any direction applied to the surface $dS$ of normal vector $\mathbf{n}$, the stress $\sigma_n$ applied to $dS$ is written as:

$$ \sigma_{\mathbf{n}} = \sigma \cdot \mathbf{n} = \frac{df}{dS} $$

$$ \text{In matrix notation: } \begin{pmatrix} \sigma_{xx} & \sigma_{xy} & \sigma_{xz} \\ \sigma_{yx} & \sigma_{yy} & \sigma_{yz} \\ \sigma_{zx} & \sigma_{zy} & \sigma_{zz} \end{pmatrix} \cdot \begin{pmatrix} n_x \\ n_y \\ z \end{pmatrix} = \frac{1}{dS} \times \begin{pmatrix} df_x \\ df_y \\ df_z \end{pmatrix} $$
σ is the total stress tensor and can be written as:

$$ \sigma = \tau - p \delta \qquad (2.8) $$

where p is the hydrostatic pressure, δ the Kronecker delta (δ = δᵢⱼ, with δᵢⱼ = 1 if i = j and 0 if i ≠ j) and τ is the stress tensor of viscosity which appears when the fluid is moving. Under hydrostatic condition, normal stress σₙ = −p only remains. The pressure p has a positive sign in the case of a traction. For a fluid at rest, the only non-zero terms of the tensor are the diagonal terms:

$$ \sigma_{xx} = \sigma_{yy} = \sigma_{zz} = -p $$

### 2.1.2.1.2 Oscillatory shear flow

To simply explore the viscoelastic properties of a material in the linear response regime, it is sufficient to shear it in an oscillating manner. If the fluid sample is subjected to a sinusoidal strain of amplitude γ₀ and pulsation ω:

$$ \gamma(t) = \gamma_0 e^{(i\omega t)} $$

The response stress σ(t) is written as:

$$ \sigma(t) = \sigma_0 e^{(i\omega t + \alpha)} $$

where α is the phase angle which allows the characterization of the material on a "viscoelasticity" scale. By analogy with the Hooke law, a complex modulus Ḡ(ω) can be defined by:

$$ \bar{G}(\omega) = \frac{\sigma(t)}{\gamma(t)} = \frac{\sigma_0}{\gamma_0} e^{(i\alpha)} $$

$$ \bar{G}(\omega) = G'(\omega) + iG''(\omega) $$

**NB:** linearity implies that G does not depend of the magnitude of γ₀. In practice, it can be achieved for small enough values of the shear rate.

This equation applies for small strains while the material structure is not modified. The total stress σ(t) can be viewed as the combination of a viscous (imaginary) and an elastic (real) parts. Moreover, the real part of the complex modulus Ḡ(ω), G', refers to the in-phase component of the stress-strain relation and is associated to the elastic energy stored per unit of volume of the material during a deformation cycle. It is called the storage modulus or the elastic modulus. The imaginary part of the complex modulus, G'', refering to the out-of-phase component is the loss modulus, or viscous modulus, associated to the viscous dissipation of energy. Those two parts are related by the phase shift α expressed as:

$$ \tan(\alpha) = \frac{G''}{G'} $$
If the fluid is purely viscous the phase angle tends to 90°:

$$G'(\omega) = 0$$

$$G''(\omega) = \eta \times \omega$$

If the material is purely elastic, the phase angle will tend to zero.

$$G'(\omega) = G$$

$$G''(\omega) = 0$$

## Maxwell model

The basic model for viscoelastic fluids is the Maxwell model. It consists in combining a linear elasticity and a linear viscosity models. For the linear elasticity model, it is postulated that the stress is proportional to the deformation (Hooke law), the coefficient of proportionality being the Young's modulus which is modeled by a spring of stiffness E. For Newtonian fluids, the stress is proportional to the deformation rate where the coefficient of proportionality is the viscosity η. This is modeled by a dashpot. By putting those two elements in series, we obtain the one-dimensional principle of the Maxwell model (Fig 2.17).

<figure><img src="image_35.png" /><figcaption>Figure 2.17: Scheme of a Maxwell liquid model: a viscoelastic fluid is modeled by a dashpot with a constant η which represents the viscous part put in series with a spring of stiffness E for the elastic part.</figcaption></figure>

The spring and the dashpot obey respectively the Hooke law and the Newton law given by:

$$\sigma_s = E \times \gamma_s$$

$$\sigma_d = \eta \times \frac{d\gamma_d}{dt}$$

where $\sigma_i$ and $\gamma_i$ are the respective stress and strain of spring (s) and dashpot (d). Equilibrium requires that the stress is the same in both elements which means: $\sigma = \sigma_s = \sigma_d$. Strains and shear rates are added:

$$\gamma = \gamma_s + \gamma_d$$

$$\dot{\gamma} = \frac{d\gamma}{dt} = \frac{d\gamma_s}{dt} + \frac{d\gamma_d}{dt}$$

In the one-dimensional case, we thus obtain the following differential equation:

$$\sigma + \frac{\eta}{E} \frac{d\sigma}{dt} = \eta_M \dot{\gamma}$$
At the start of the stress-relaxation experiment, $t = 0$, and $\sigma(t=0) = \sigma_0$ is the initial stress. Solving the last differential equation, we obtain:

$$\sigma(t) = \sigma_0 e^{-\frac{E}{\eta}t}$$

The stress decays exponentially with a characteristic relaxation time $\lambda$ defined as:

$$\lambda = \frac{\eta}{E}$$

The storage $G'$ and loss $G''$ modulii are thus given by:

$$G' = E \times \frac{\omega^2 \lambda^2}{1 + \omega^2 \lambda^2}$$

$$G'' = E \times \frac{\omega \lambda}{1 + \omega^2 \lambda^2}$$

<figure><img src="image_36.png" /><figcaption>Figure 2.18: Maxwell fluid behavior for an oscillatory shear: the storage modulus $G'$ (plain line) and the loss modulus $G''$ (dotted line) intersect at the pulsation $\omega = 1/\lambda$</figcaption></figure>

At low frequencies, the loss modulus $G''$ is higher than the storage modulus $G'$ (Fig 2.18). In this case, the viscous behavior is predominant and: $G' \propto \omega^2$ and $G'' \propto \omega$. When the frequency increases, $G'$ tends to a plateau and $G''$ decreases rapidly; the elastic behavior is then predominant. At the intersection point, $G' = G''$ and $\omega\lambda = 1$.

## Constitutive equations

To generalize linear models (Maxwell model) to non-linear behaviour models as for real fluids, the constitutive equation must be conformed to several principles [4]:

- **Principle of coordinates invariance:** the different constitutive variables (stress, strain, rates etc) must be expressed as tensors to ensure the invariance of the equation relationship from a coordinate frame to another (ex: Cartesian to Spherical coordinates)
- **Principle of invariance:** as a consequence from the previous principle, when passing the constitutive equation from a coordinate system to another, their orientation and scale must be invariant.

- **Principle of objectivity:** the deformation of a material is not affected by any frame rotation but only by the deformation itself.

This last principle is not applied in linear models. It is thus necessary to modify the linear Maxwell model by replacing the time derivative with an invariant one called the convected derivative:

$$ \frac{D\sigma}{Dt} = \frac{\partial\sigma}{\partial t} + \nabla\sigma $$

where

$$ \nabla\sigma = (\nu \cdot \nabla)\sigma - (\nabla\nu)\sigma - \sigma(\nabla\nu)^T $$

Frame invariant models are thus derived from linear one by writing the equation in a frame of reference that moves with a material element. The three most common frame invariant time derivatives are the upper- and lower-convected and the corotational time derivative. In the following, we will present the Upper Convected Maxwell, Oldroyd-B and White-Mezner models.

## Upper-Convected Maxwell Model

The linear Maxwell model can be generalized for arbitrary flows by writing its frame-invariant analogue which is called the Upper-Convected Maxwell (UCM) model that takes into account the relaxation time $\lambda$. We have from the previously equation 2.8:

$$ \sigma = \tau - p\delta $$

now, the stress tensor of viscosity $\tau$ obeys the UCM:

$$ \tau + \lambda \frac{\nabla \tau}{\nabla \cdot \tau} = \eta_0 \dot{\gamma} $$

where $\eta_0$ is the zero-shear polymer viscosity and $\nabla$ the upper-convected derivative. By combining a basic kinetic theory and a constitutive equation such as UCM equation, it is possible to link the zero-shear polymeric viscosity $\eta_0$ and the relaxation time $\lambda$ to the microscopic properties of polymers. Briefly, polymer molecules are modeled by hookean dumbbells consisting of two spherical beads connected by an elastic spring. In the case of very dilute polymer solutions, the dumbbells only interact with each other through hydrodynamic interactions or intermolecular forces. After several steps of equations (motion, diffusion, forces etc), we can obtain the relationship: $\eta_0 = \lambda nk_B T$ with $n$ the number of dumbbells, $k_B$ the Boltzmann constant and $T$ the temperature.

This model takes into account the memory effects of materials, but it does not consider the solvent viscosity or the shear rate dependence of the viscosity as the following models.
Oldroyd-B Model

The Oldroy-B model is the extension of the UCM model in which the solvent viscosity η_s is taken into account. As for the UCM model, the viscoelastic fluid is modeled by a suspension of dumbbells connected by infinitely extensible springs in a newtonian solvent. If the tensor of viscosity τ obeys to the UCM, we obtain the set of equations called the Oldroyd-B model which is formulated via the total deviatoric stress σ = η_s ˙γ + τ which satisfies:

$$
\sigma + \lambda \nabla \sigma = \eta(\dot{\gamma} + \lambda_r \dot{\gamma})
$$

where λ_r is an independent time scale called the retardation time given by: λ_r = λ(η_s/η);
and η the total viscosity which is simply: η = η_s + η_0. This model can be simplified if:

- λ_r = 0: it reduces to the UCM model previously described

- λ = λ_r: it assimilates to the case of a newtonian fluid with the viscosity η₀

- λ = 0: it simplifies to a second-order fluid with a vanishing second normal stress coefficient.

When viscoelastic fluids are subjected to shear flow, normal stresses arise and are respon-
sible for spectacular physical phenomena. The mechanism of those stresses can be ex-
plained by considering folded polymers chains at rest. While subjected to the action of
a local shear, the chains will tend to align to the streamlines. However, at the same time,
entropic forces are acting to make the elongated polymer chain return to its folded con-
formation. This leads to an extra tension in the direction of the flow. We have already
defined the normal stresses differences N₁ and N₂ which can also be related to the shear
rate for a linear shear flow v(x, t) = (γy, 0, 0) by:

$$
N_1 = \tau_{11} - \tau_{22} = \psi_1 \dot{\gamma}^2
$$

$$
N_2 = \sigma_{22} - \sigma_{33} = \psi_2 \dot{\gamma}^2
$$

where ψ₁ and ψ₂ are respectively the first and second normal stress difference coeffi-
cients. From the Oldroyd-B model equation, this first coefficient ψ₁ is linear in the poly-
mer relaxation time λ and is given by: ψ₁ = 2λη₀ (≥ 0). In this model, there is no sec-
ond normal stress difference coefficient (ψ₂=0) as there are no transverse normal stresses.
Moreover, the viscosity is constant and always equal to the zero-shear rate viscosity which
is not representative of the real polymer solutions viscosity that often exhibits shear-
thinning behavior. This is why, further models that considers those characteristic must
be used.

In the case of newtonian fluids (ex: water), N₁ and N₂ are equal to zero. For non-newtonian
fluids, the existence of those normal stress differences is responsible for spectacular phe-
nomena as the Weissenberg effect or the die swell that we will observe and describe in
more detail in the Chapter 4 4.3.3.

White Metzner Model

Shear thinning fluids have a viscosity which is a decreasing function of the shear rate
˙γ. Many polymer solutions exhibit this type of behavior such as blood. This can be
easy interpreted by the presence of polymeric structures that tend to align in the flow direction while increasing the shear rate, which leads to a viscosity decrease. The White-Metzner model is suitable for this kind of fluid as it is a modified version of the UCM equation, which assumes a shear-rate dependent shear viscosity and relaxation time, here described using a Cross law:

$$ \eta(\dot{\gamma}) = \frac{\eta_0}{1 + \left(\frac{\dot{\gamma}}{\dot{\gamma}_0}\right)^{1-n}} $$

### 2.1.2.2 Rheometer geometries

Several geometries of rheometers are available to measure those characteristic quantities. The appropriate geometry is chosen in function of solution concentrations, viscosity, pH or volume. For each geometry, geometric shape constants K relating to the viscosity can be defined by:

$$ \eta = \frac{\sigma}{\dot{\gamma}} $$

where: $\sigma = M \times K_\sigma$

and: $\dot{\gamma} = \Omega \times K_\gamma$

with M the torque in N.m, $\Omega$ the angular velocity in rad/s, $K_\gamma$ the strain constant and $K_\sigma$ the stress constant.

<figure><img src="image_37.png" /><figcaption>Figure 2.19: Rheometer geometries: left: Double gap Couette cell, middle: Parallel plates and right: Cone and plate</figcaption></figure>
Double gap Couette cell geometry

Double gap Couette cell geometry is a concentric cylinder system suitable for solutions of low viscosity. A fixed cylinder is mounted in the center of the cup and a mobile cylinder is introduced inside the cell. This double system increases the shear area and lower torques can be determined. It was used in the case of alginate solutions of concentration lower than 1%. For this geometry, the geometric constants are:

$$
\begin{align*}
K_{\gamma} &= \frac{R_{1}^{2} + R_{2}^{2}}{R_{2}^{2} - R_{1}^{2}} \\
K_{\sigma} &= \frac{(R_{1}^{2} + R_{2}^{2})}{4\pi h R_{2}^{2}(R_{1}^{2} + R_{3}^{2})}
\end{align*}
$$

where Rᵢ are the different radii and h the height of fluid in meter. For fluids with higher viscosity, other geometries must be used.

Cone-plate geometry

In a cone-plate geometry, the fluid is sheared between a plate and a cone of a radius R which makes an angle α (less than 4°) with the plate. The cone is truncated and has no contact with the lower plate. The resulting air gap corresponds to the truncation, and is generally very small (few microns). The main advantage of this geometry is the homogeneous shear rate induced in the sample. The two constants K are given by:

$$
\begin{align*}
K_{\gamma} &= \frac{1}{\alpha} \\
K_{\sigma} &= \frac{3}{2\pi R^3}
\end{align*}
$$

During the experiments, the sample is kept in a humid atmosphere at 20°C thanks to a hermetically sealed cell. In our case, the rheometric tests were carried out mainly in oscillatory regime to determine the elastic modulus G' and the viscous modulus G''. These two moduli depend on the frequency of stress. The pulsation varied between 0.01 and 100 rad/s.

This rheometer geometry was also used to measure the solution viscosity and normal stresses at different shear rates ranging from 0.01 to 100s⁻¹.

Glass parallel plate geometry

The parallel plate geometry can be considered as a simplified version of the cone and
plate, having an angle of 0°. The fluid is constrained in the narrow gap between the two
surfaces. For collagen solutions, the acid acetic may corrode the elements of the previous
geometry (cone and plate) made of steel. We therefore used a glass parallel plate geometry
that was available in the laboratory to study our collagen solution at 5 mg/mL (Chapter 5,
5.2 p 156). A small amount of solution is needed in this case which is a great advantage.
However, the shear rate is not uniform and the apparent viscosity must be calculated by
using a model. In fact, the shear rate given by the rheometer is calculated at the periphery of the plate and the two constants K are then given by:

$$K_Y = \frac{R}{e}$$

$$K_σ = \frac{2}{\pi R^3}$$

### 2.1.3 Mechanical testing of extruded threads

Mechanical properties of extruded threads were investigated in order to assess the effect of fibrillogenesis buffer, culture medium, cells seeding or collagen concentration.

Tensile testing of extruded threads without cells were performed using a BOSE® Electro-Force® 3200 Series II Test Instrument at the Institut des Nanosciences de Paris (INSP). Specimens were attached horizontally to the stainless steel crossheads by setting each extremity on a cyanoacrylate glue drop and filter papers. Sample and crossheads were immersed in a glass Petri dish filled with either fibrillogenesis buffer or culture medium at room temperature. Initial threads diameter, d0, was determined by optical microscopy assuming full cylindrical symmetry. Mechanical loading of the thread was achieved by driving one crosshead at a constant velocity of 0.1 mm per second with a stepping motor (0.1-micron steps). The loading force F was measured with a 220g load cell. The cross-section S of the thread was calculated from the diameter measurements by bright field microscope and by assuming that threads are cylindrical. We defined the nominal tensile stress as:

$$\sigma(Pa) = \frac{F(N)}{S(m^2)}$$

The nominal strain is defined as

$$\epsilon = \frac{L - L_0}{L_0}$$

where L is the distance between the glue points and L0 the initial length. All measurements were performed up to the specimens rupture. The resulting stress-strain curves were thus plotted and analyzed. The Young Modulus E (Pa) of a sample is the stress-strain curve slope for small deformations and is obtained from the Hooke law:

$$\sigma(Pa) = E \times \epsilon$$

Tensile testing of threads put in cell culture were performed at the laboratory Sciences et Ingénierie de la Matière Molle-Physico-chimie des Polymères et des Milieux Dispersés (SIMM) with the help of Jingwen Zhao. In fact, with the previous instrument, threads cultured with cells splitted off from the crossheads as the cyanoacrylate was not sufficient to fix them. Instron 3343 machine is equipped with clamping jaws where sides are covered by Velcro straps. The threads with cells were put between the Velcro covered jaws vertically. The threads were immersed thanks to a tubing filled with PBS 1X solution which integrates the whole system. All tensile experiments were carried out at a velocity of 0.1mm/s with at least three measurements per condition to check the reproducibility. The set-up configuration requires a further data processing. Indeed, as the thread is immersed in the PBS solution vertically, the buoyant force exerted on the thread and equal
<figure><img src="image_38.png" /><figcaption>Figure 2.20: Mechanical testing configurations: A) horizontal for threads without cells and fixed with cyanoacrylate and filter paper; B) vertical for threads with cells fixed with velcro coated clamping jaws</figcaption></figure>

to the weight of the fluid that the body displaces, needs to be taken into account. Briefly, after the thread rupture, data were still acquired to collect the buoyancy contribution. This was then subtracted to the force signal in order to have only the thread mechanical properties. The following data processing is then the same as the horizontal set-up.

For both configurations, Young's modulus, tensile strength, and elongation at break of the threads were derived from the stress–strain curves as described in the sections 1.1.1.2.2 and 2.1.3.
## 2.2 Materials

In this section, the collagen and alginate chemical properties will be described. The different protocols of purification, preparation, concentration etc. will be detailed for pure and mixed systems.

### 2.2.1 Collagen solutions

#### 2.2.1.1 Collagen type I and its fibrillogenesis

Collagen is a molecule of the extracellular matrix and has large domains organized in triple helix. There are 28 types of collagen that differ in the composition of the polypeptide chains, their three-dimensional structure and their distribution in connective tissues [5]. In most collagens, the three polypeptide chains which form a α helix are identical. If the chains are different, the collagens are said to be heterotrimeric. The 28 types of collagens are divided into seven families that group together collagens with similar structural behaviors (Fig 2.21).

<table><thead><tr><th>Family</th><th>Collagen types</th><th>Scheme</th></tr></thead><tbody><tr><td>Fibril forming</td><td>I, II, III, V, XI, XXIV, XXVII</td><td>Collagen I<br>Collagen I fibril</td></tr><tr><td>Networking forming</td><td>IV, VIII, X</td><td>Collagen X<br>Collagen X network</td></tr><tr><td>FACIT</td><td>IX, XII, XIV, XVI, XIX, XX, XXI, XXII</td><td>Collagen IX<br>Collagen IX on collagen II fibril</td></tr><tr><td>Anchoring</td><td>VII</td><td>Collagen VII</td></tr><tr><td>Transmembrane</td><td>XIII, XVII, XXIII, XXV</td><td>Collagen XIII<br>Collagen XVII</td></tr><tr><td>Beaded filaments</td><td>VI, XVI, XXVIII</td><td>Collagen VI<br>Collagen VI filament</td></tr><tr><td>Others with unique function</td><td>XV, XVIII, XXVIII</td><td></td></tr></tbody></table>

Figure 2.21: Table and schemes of the seven collagen families. Blue dots are non-collagenous domains, the black line represents triple helix domain (Glycine-X-Y). Schemes adapted from [6]

Some collagens are specific to a tissue type and others are more ubiquitous. The most abundant collagen is fibrillar collagen type I and it is the one of interest in this project.
In vivo, collagen is synthesized by the cells. The synthesis begins by the production of the alpha chains in the ribosomes of the rough endoplasmic reticulum in which post-translational modifications take place (Fig 2.22 1).

<figure><img src="image_39.png" /><figcaption>Figure 2.22: Collagen biosynthesis scheme: 1) Alpha chains synthesis, 2) Winding of the three alpha chains (procollagen molecule), 3) Removing terminal polypeptides, 4) Collagen molecule assembly. From https://mmegias.webs.uvigo.es/02-english/5-celulas/2-componentes_proteinas.php</figcaption></figure>

It is followed by several steps: the hydroxylation of proline and lysine residues, the glycolysation of certain hydroxylysines and the formation of disulfide bridges between the C-terminal propeptides (which causes the three chains coiling in the N-terminal direction). The resulting molecule is called procollagen (Fig 2.22 2). Procollagen is soluble in physiological condition thanks to the presence of non-helical and charged propeptides. At this level, the procollagen molecules are found in vesicles in the Golgi apparatus before being secreted outside the cell. At the end of the process, enzymatic cleavage of the terminal N and C propeptides (Fig 2.22 3) globular domains takes place, resulting in the formation of fibrils (Fig 2.22 4)). The loss of the globular domains leads to a decrease of the solubility under physiological conditions. The cleavage of the ends exposes the telopeptide sequences, which are small, non-helical domains. These telopeptides contain the binding sites for fibrillogenesis [7]. The collagen fibrils are spontaneously formed by the triple collagen helices, soluble at acidic pH but not soluble in a physiological medium during fibrillogenesis.

Collagen type I triple helix is composed of two strands α1 which have an identical sequence of 1056 amino acids and one strand α2 with a sequence of 1023 amino acids. These three strands are formed in their central part by the repetition of a sequence of three amino acids Glycine-X-Y. The glycine points towards the center of the triple helix while the larger residues of X and Y amino acids point to the outside of the triple helix, as shown in Figure 2.23 (A), and are therefore available for interactions with other molecules or with the solvent. Proline and hydroxyproline are the most encountered (30% of cases) in the X and Y positions respectively.
<figure><img src="image_40.png" /><figcaption>Figure 2.23: Collagen type I triple helix scheme. Left: Viewed along the molecular axis (-carbons only), showing the paths of the individual polypeptide chains and the locations of residues in the GlyX-Y triplets (G = Gly; from [8]).</figcaption></figure>

The amino acid sequence corresponds to the primary structure (Fig 2.23 bottom right) and it is the folding of the left helix chain of 8.68 nm pitch, which forms the secondary structure. The folding is driven by steric repulsions between the proline residues in the X-position and the hydroxyproline residues in the Y-position. Between two residues, the axial distance is 0.286 nm (Fig 2.23 middle right). Then, the tertiary structure is formed by the folding of the secondary structures in space and the quaternary structure by the association of monomers (the chains α) between them. Indeed, the presence of numerous glycines along the chains α allows them to associate to form a right triple helix. The triple helix then has a length of 300 nm and a diameter of approximately 1.5 nm (Fig 2.23 top right). Hydrogen bonds between the glycine of one chain and the carboxyl groups of another chain ensure the stability of the triple helix. The ends of the collagen type I molecule, called telopeptides, are non-helical regions.

### 2.2.1.2 Chemical properties

#### Solubility

Collagen molecules of young tissues are soluble at acidic pH at room temperature. There are two different ways to induce the fibrillogenesis process: in physiological condition at 4 °C by increasing the temperature to 37°C; or by increasing the pH if the collagen solution is in acidic solutions. In the case of fibrillogenesis induced by an increase of pH, the collagen solutions are subjected to other solutions at neutral pH. This induces the precipitation of the collagen molecules which assemble into fibrils. Depending on the physicochemical conditions (pH, ionic strength), the size of the fibrils will vary, which may have consequences on the mechanical properties as we will see in Chapter 5 (5.3 p 157).

As mentioned above, the precipitation of the collagen molecules can also be induced by a thermal increasing. Gross and Kirk [9] and Jackson and Fessler [10] were the first to study this phenomenon and showed that collagen in neutral saline at 4 °C formed a fibrillar gel when the temperature rose to 37 °C for relatively low starting concentrations between 0.2 and 2 mg/mL. However, they found that the collagen could be denatured before reaching 37 °C and that the precipitation showed inhomogeneities. This second way to induce fibrillogenesis will be not used in our study.
## Collagen net charge

From the physico-chemical point of view, the collagen molecule is a polyelectrolyte with an amphoteric character. Indeed, it can behave both as an acid (proton donor) or as a base (proton acceptor). Knowing the amino acid sequence of the three α-chains of the collagen molecule, it is possible to assess the distribution of the charges along the molecule which appears relatively homogeneous. Furthermore, Frederic Gobeaux estimated the net charge of the proteins as a function of the pH by using a spreadsheet provided by the University of Maine [11].

<figure><img src="image_41.png" /><figcaption>Figure 2.24: Collagen net charge plot calculated from its amino acid sequence as a function of pH. The calculation was performed with Charge Calculator III [11]</figcaption></figure>

The results of those calculations are presented in Figure 2.24. In acetic acid at 500 mM, at pH 2.5, the net charge of collagen is at its maximum of 259 positive charge. The electrostatic repulsions are therefore maximal and this explains why in those conditions the collagen is soluble. On the other hand, at physiological pH (pH 7.4) and with only 38 positive charges, this condition is very close to the isoelectric point (pH ≈ 9.8) and we understand why the collagen precipitates in the form of fibrils when acid solutions are neutralized. This calculation remains very approximate, because the isoelectric point depends strongly on the ions present in solution and their ionic strength, which is not taken into account here.

## Flexibility and persistence length

The triple helix of collagen can be considered as a rod which is more or less flexible. This criterion may be important to interpret the collagen behavior in solution. Indeed, the flexibility impacts the conditions of the isotropic/nematic transition and the rheological properties of macromolecules in solution.

A critical parameter to characterize the flexibility of a polymer is its persistence length P, which describes the length at which two tangent vectors are correlated. The flexibility α is described as the ratio between the total contour length of the molecule L and its persistence length P:

$$ \alpha = \frac{L(nm)}{P(nm)} $$
When $\alpha \ll 1$, the polymer is very rigid and can be modeled by a rod, while if $\alpha \gg 1$, it is very flexible. In this case, the polymer is modeled by a chain of articulated rigid segments (Kühn segments) or by the wormlike chain model. If the two quantities L and P are of the same order of magnitude, the term semi-flexible "molecule" is used. However, the persistence length of a polyelectrolyte can vary depending on its concentration, the solvent (ionic strength), or the temperature. Also, there are various values for collagen type I in the literature, depending on the experimental conditions, but also according to the method used. Some of them are reported in Table 2.2.

<table><caption>Table 2.2: Persistence length for different buffers found in the literature. (FFB: Fibril Forming Buffer and PBS: Phosphate Buffer Saline)</caption><thead><tr><th>C (mg/mL)</th><th>Buffer</th><th>pH</th><th>P length (nm)</th><th>Ref</th></tr></thead><tbody><tr><td>0.08 - 0.6</td><td>0.3 M Acetate + 6 mM NaCl</td><td>4</td><td>170</td><td>[12]</td></tr><tr><td>0.4 - 1.5</td><td>0.06 M potassium acetate buffer</td><td>4.8</td><td>130</td><td>[13]</td></tr><tr><td>0.02 - 0.03</td><td>glycerol</td><td>NA</td><td>57</td><td>[14]</td></tr><tr><td>0.0051</td><td>Nanopure Water</td><td>7</td><td>12.5</td><td>[15]</td></tr><tr><td>6.4 10<sup>-4</sup></td><td>FFB</td><td>7</td><td>135</td><td>[15]</td></tr><tr><td>0.00051</td><td>PBS</td><td>7</td><td>165</td><td>[15]</td></tr></tbody></table>

The most plausible values in our case are those of the order of 165 nm, measured in PBS [15]. In this condition, the contour length is about 260nm which leads to a flexibility $\alpha$ of 1.6. This means that collagen triple helix molecules can be considered as "semi-flexible".

### 2.2.1.3 Collagen solutions preparation

Type I collagen was extracted from young wild-type non-treated Wistar or Sprague-Dawley rat tail tendons. The rats cadavers come from biological laboratories. They were initially sacrificed as controls of various experiments and not only for collagen purification. As they are a few months old, the tendon collagen is very weakly crosslinked. In the following part, the whole protocol of collagen preparation is described.

### 2.2.1.4 Collagen extraction and purification

The rat tail is travelled by tendons running longitudinally along its length (2.25, left scheme). To collect those tendons, the first step consists in cutting partially the skin with sterile clamps (about 1 cm from the rail extremity), twisting it and moving apart carefully the two pieces until bundles of white tendons are exposed and separated (2.25, right image). Those are the rat tail tendons.

This step is repeated every centimeter until 2-3 cm from the rat tail basis is reached. The excised tendons are then washed several times with phosphate-buffered saline (PBS) to remove cells and traces of blood coming from the veins and artery. The remaining cellular wastes are lysed by the addition of NaCl 4 M solution. Tendons are rinsed again with PBS and solubilized overnight in 500 mM acetic acid (pH=2.5). A viscous mixture containing proteoglycans, collagen, small proteins and some insoluble fibers is thus obtained. The collagen solution is purified by successive and selective precipitations. By centrifugation at high speed, the insoluble fibers are first removed. Then, the proteoglycans and polyelectrolytes of high molecular weight are precipitated by a slow step by step addition of a 4 M NaCl solution up to a final concentration of 300 mM. They can then be separated by a new centrifugation. Finally, the collagen is precipitated by a second addition of 4M NaCl solution up to a final concentration of 600 mM. After centrifuging at
<figure><img src="image_42.png" /><figcaption>Figure 2.25: Rat tail schemes of sagittal (top left) and transverse (bottom left) sections. The rat tail is composed of several caudal vertebrae separated by intervertebral discs. All along its length it is travelled by tendons, veins and the ventral artery. The extraction of rat tail tendons is performed by cutting every centimeter the rat tail and moving apart the two pieces until tendons fascicles are exposed (right).</figcaption></figure>

low speed, pellets containing the collagen I is collected. Those pellets are solubilized and dialyzed several times in 500 mM acetic acid in order to remove NaCl. The collagen solutions are kept at 4 °C and centrifuged at 41,000xg for 4 h before use. After each collagen extraction, the protein purity is checked and the collagen concentration is quantified.

### 2.2.1.5 Collagen solutions dosage

Collagen concentration is determined from the acidic solution by assessing the amount of hydroxyproline. In fact, the hydroxyproline (hp) is assumed to be a quasi-specific amino acid of the collagen extracted from the rat tail tendons and representing approximately 13% of its molecular weight. To dose this amount, a volume of 50 µL of collagen solution is mixed with a the same volume of hydrochloric acid at 37% in order to release the hydroxyprolines. The mix is placed in an oven at 106 °C overnight to hydrolyse the collagen. The solvent is then evaporated and the residue obtained is re-suspended in 1 mL water. This solution is further diluted with water to obtain three samples of 400 µL of different concentration ratios(1/40, 1/20, 1/10). In each of those samples, a volume of 200 µL of Chloramine-T, perchloric acid and DMBA (4 dimethylamino-benzaldehyde) is added to provoke a colored reaction between yellow and red. The color is related to the amount of hydroxyproline present in the solution. Finally, the absorption at 557 nm is measured for each sample. A reference curve is obtained for a hydroxyproline solution which concentration is known and equal to 20 µg/mL. This curve correlates the hydroxyproline concentration with absorption. Taking into account the dilution of each sample and knowing that the type I collagen molecule contains 13% of hydroxyproline, the collagen concentration $C_{collagen}$ is calculated as follows:

$$C_{collagen} = 7.6 \times f_{D} \times C_{hp}$$