3.2.2 Study 2: Protein film development with controllable mechanical properties, degradation rates, and inflammation response ... 103
3.2.2.1 Crosslinking to control film physiochemical properties ... 104 3.2.2.2 Quantification of film crosslinking efficiency ... 106 3.2.2.3 Effect of crosslinking on film enzymatic degradation resistance ... 106 3.2.2.4 Effect of crosslinking on film mechanical properties ... 107 3.2.2.5 Effect of crosslinking on film inflammatory response ... 107 3.2.3 Study 3: Development of bilayered tubular scaffolds a porous outer layer and dense film layer ... 109
3.2.3.1 Effect of freezing direction and mandrel material on scaffold microarchitecture ... 110 3.2.3.2 Cell seeding, static culture, and compaction ... 111
3.3 Results ... 114
3.3.1 Study 1: Development of tubular scaffolds ... 114 3.3.1.1 Effect of freezing rate on tubular scaffold microarchitecture ... 114 3.3.2 Study 2: Protein film development with controllable mechanical properties, degradation rates, and inflammation response ... 116
3.3.2.1 Fabrication of dense, non-porous protein films ... 116 3.3.2.2 Quantification of film crosslinking density ... 117 3.3.2.3 Effect of crosslinking on film enzymatic degradation resistance ... 118 3.3.2.4 Effect of crosslinking on film mechanical properties ... 119 3.3.2.5 Effect of film crosslinking on film inflammatory response ... 122
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3.3.3 Study 3: Development of bilayered tubular scaffolds with a porous outer layer and dense film layer ... 123
3.3.3.1 Effect of freezing direction on resulting microstructure of bilayered scaffolds ... 124 3.3.3.2 Effect of mandrel material on resulting bilayered scaffold microarchitecture ... 125 3.3.3.3 Cell seeding, migration, and compaction within tubular scaffolds ... 127
3.4 Discussion ... 127 3.5 Conclusion ... 135
3.1 Introduction
Tissues are 3D self-assembled constructs developed to serve very precise needs based on their location. The organisation of the cells and ECM components in a tissue ultimately dictates the tissues potential function and is highly adapted to deal with any mechanical loads which it will encounter (Stella et al., 2010). This hierarchical structuring from the nanostructure to the macrostructure is a tightly regulated and complex process which has proven elusive to recreate fully in vitro (Fratzl and Weinkamer, 2007). Thus, bioengineers have resorted to developing biofabrication tools and techniques in order to aid in recreating tissues ex vivo. The ideal tissue engineered vascular graft (TEVG) would thus mimic native tissue microenvironment and structure, and be composed of a multi-layered lamellar structure (Fig. 3.1).
The biofabrication technique selected ultimately dictates the possible end architectural constraints. Multiple techniques now exist to generate structures with control over the nano/microstructure and with the capability to generate large, anatomically correct scaffolds with a suitable macrostructure. However, many of these techniques are limited to either synthetic polymers capable of withstanding the elevated temperatures required for 3D printing techniques (e.g fused deposition modelling) or require specialist photosensitive materials (e.g. stereolithography, two-photon polymerisation). Freezedrying, as a technique, is amenable to a broad range of materials and is highly suitable for thermosensitive natural polymers which may denature
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under elevated temperature conditions. However, freezedrying is often seen as limited in the scope of microstructural and macrostructural manipulation possible. Previous work has shown that although the microstructure can be manipulated, it is influenced by the composition (Ryan et al., 2014) as well as the specific freeze-drying conditions, such as freezing rate (O’Brien et al., 2004) and final freezing temperature (Haugh et al., 2009). Therefore, optimisation of microstructure may be achieved by altering these parameters. Hierarchical structured scaffolds can be fabricated through a combination of techniques to leverage the advantages of each while minimising the disadvantages. So far, this thesis has demonstrated that generating three dimensional porous scaffolds from collagen and elastin (Chapter 2) elicits a number of distinct advantages as a template for tissue regeneration. Utilising this data we next sought to further optimise this biomaterial by structuring it hierarchically. Specifically, we sought to create a physiologically relevant bilayered tubular architecture through generation of a porous outer layer (analogous to the tunica media) and a dense film inner layer (analogous to the tunica intima).
Traditional tissue engineered vascular grafts are created as a single homogenous layer in a tubular structure. Some of the earliest work utilising natural polymers for this application were in the form of easy to fabricate collagen gels. Indeed, a number of researchers have investigated SMC- seeded collagen type I gels which were compacted by the seeded cells over time to produce vascular grafts (L’Heureux et al., 1993; Seliktar et al., 2003a). However, due to the low polymer density and disorganised matrix formed, these grafts did not possess the required mechanical properties for the challenging haemodynamic environment (Barocas et al., 1998; Hirai et al., 1994). Consequently, the field progressed to incorporate synthetic polymer meshes within the natural polymers in order to improve the overall graft mechanical properties (Tillman et al., 2009; Tschoeke et al., 2008). However, despite the limitations associated with these early grafts, significant advances in natural polymer fabrication techniques over the last 20 years have enhanced the potential for development of a successful collagen-based tissue engineered vascular grafts. The hierarchical
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structuring of natural polymers, like collagen, can be achieved by altering the protein concentration between the layers to leverage the effect of density on mechanical properties (Caliari et al., 2011; Kumar et al., 2013). The integration of a dense luminal layer would further match the native architecture, would potentially increase mechanical properties, and provide a suitable surface for endothelialisation. While it is clear that emulation of natural vessel architecture and polymer composition has distinct advantages, the optimal combination of biomaterial and architecture for vascular tissue engineering applications has yet to be elucidated.
In addition to their inherently excellent biological properties, natural polymers offer the ability to easily modify their mechanical properties and degradation rates through crosslinking. Crosslinking, as a technique, offers the powerful ability of being able to tune a material’s properties independently of the density or microstructure by creating covalent bonds between nearby collagen molecules. Thus, a hierarchically structured scaffold can be further fine-tuned to match the desired mechanical properties. While the native collagen crosslinking pathway occurs through the enzyme lysyl oxidase in vivo, the harsh extraction procedures used to process collagen for medical device applications result in the degradation of these bonds. Thus, exogenous methods based upon physical, chemical, or alternative enzymatic methods are often utilised (Koh and Atala, 2004; Tierney et al., 2013). These exogenous crosslinks are capable of causing vastly improved mechanical properties but also result in cytotoxicity and/or unwanted cellular responses such as calcification. Crosslinking also can be used to tailor the degradation rate of the biomaterial where a slow degradation rate allows a gradual stress transfer to occur and is favourable from a mechanical viewpoint. However, biologically the most favourable outcome would be for the degradation rate to match the rate of new tissue growth which generally necessitates a faster degradation rate. The ideal solution involves striking a balance between the mechanical and biological suitability of the degradation rate for effective tissue repair. Interestingly, crosslinking has also being implicated as being involved in host-biomaterial interactions by modifying the immunogenic response (McDade et al., 2013).
100 3.1.1 Objectives
The overall objective of this chapter was to develop bilayered tubular scaffolds with controllable properties for use as a tissue engineered vascular graft. This chapter focused on optimising the architecture of the previously optimised biomaterial composition from Chapter 2. In line with these objectives, the specific aims of this study were:
Study 1: Develop the optimal biofabrication method to produce porous collagen-elastin tubular scaffolds using a custom designed mold and varying the freezing rate to control microstructure.
Study 2: Develop dense protein films with controllable mechanical properties and degradation rates through varied crosslinking procedures. Additionally, the effect of crosslinking on the inflammation response of macrophages was also examined in vitro.
Study 3: Develop bilayered tubular scaffolds by selecting the optimal methods from studies 1 and 2 above. The microstructure was further optimised through altering the freezing direction (horizontal v vertical) and mold materials during fabrication. Preliminary in vitro culture focused on examining tubular scaffold seeding and compaction on a custom constrainment rig during static culture.
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Figure 3.1 Graphical abstract of the chapter aims
The initial study focused on developing tubular porous scaffolds using the optimal composition from Chapter 2. Hierarchically structuring to mimic native architecture (left) was providing by a dense film layer (intima mimic) on the luminal aspect and a porous outer layer (media mimic) to form a bilayered tubular scaffold (right). Control over the mechanical properties, degradation rate, and inflammation response was examined via crosslinking. Validation of cells seeding and static culture effects were also performed.
3.2 Materials and Methods
3.2.1 Study 1: Development of tubular scaffolds
In order to be able to produce tubular scaffolds via freezedrying a custom designed mold was machined in Trinity College Dublin mechanical engineering workshop. Preliminary experiments focused on examining a number of alternative mandrel materials, altered outer mold diameters, and altered freezing cycles in order to achieve a reliable fabrication process. A collagen-elastin suspension (CE100) was created as per Section 2.2.1. The suspension was then pipetted into the mold (Fig. 3.2), which had an outer diameter of 11mm and an inner diameter of 5mm, and freeze-dried as per Section 2.2.1. Due to shrinkage of the scaffold during the freezedrying process the final wall thickness was approximately 2.5mm. This thickness was selected based upon the contraction data obtained in Section 2.3.4, where it was calculated that an initial wall thickness of 2.5mm would contract to <1mm after 14 days culture.
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Figure 3.2 Cross-section schematic of the mold used for producing tubular scaffolds.
The interstitial space between the mold outer casing and the mandrel is filled with the collagen-elastin suspension and freezedried. The mold casing and mandrel are both stainless steel to ensure they transfer heat efficiently to result in a homogenous freezing profile. Once freezedried the mold is easily disassembled and the tubular scaffold can be removed from the mandrel. To enable the production of multiple tubular scaffolds per freezedrying cycle the mold was designed with eight cavities. Additionally the external diameter of the mold cavities ranged from 9 mm to 12 mm to allow controllable variation in the wall thickness of the tubular scaffold walls (Fig. 3.3). The mandrel size could also be altered although for the scope of this project it was maintained as 5 mm.
Figure 3.3 Custom designed mold for tubular scaffold fabrication.
(A) The custom designed mold consists of two machined stainless steel plates which are held together via hex head screws (B) A 5mm diameter mandrel fits down the mold cavity to create an interstitial space where the protein suspension can be freezedried to form tubular scaffolds.
B A
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3.2.1.1 Effect of freezing rate on tubular scaffold microarchitecture
Two distinct freezing rates were examined to determine the optimal microarchitecture capable of supporting the growth of smooth muscle cells. The optimal pore size should be large enough to facilitate effective cell migration to within the scaffold structure. Samples were frozen via controlled freezing on the freeze-drier shelf at 1°C/min, or via liquid nitrogen immersion which would produce a freezing rate of >50°C/min. All samples were then transferred to the freeze-dryer and dried using a previously optimised drying profile (O’Brien et al., 2004). Microstructural analysis was performed via SEM as per Section 2.2.2.
3.2.2 Study 2: Protein film development with controllable mechanical