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Scuola di Ingegneria
Corso di Laurea Magistrale in Ingegneria Biomedica
Tesi di Laurea Magistrale
Fabrication and characterization of microporous,
photo-crosslinked poly(trimethylene) carbonate and
nano-hydroxyapatite composite.
3D Printing and micro-porosity and roughness
effects on the osteogenic differentiation of hBMSCs
Relatori
Prof. Dr. Giovanni Vozzi
Prof. Dr. Dirk W. Grijpma
Ing. Mike A. Geven
Candidato
Anna Lapomarda
Anno accademico 2016-2017
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Index
Introduction ... 6
1. Bone tissue engineering ... 9
1.1 Bone Tissue ... 11
1.1.1 Bone components and physiology ... 11
1.1.2 Cellular classification and bone remodelling process ... 13
1.1.3 Mechanical properties of bone ... 14
1.2 The scaffold: a key role in BTE ... 16
1.3 Materials ... 19
1.3.1 Biodegradable polymers ... 19
1.3.2 Bioactive ceramics ... 24
1.3.3 Composite of biodegradable polymers with nHA ... 26
1.3.4 Scaffolds as drug delivery systems for BTE ... 27
1.4 Fabrication techniques for BTE scaffolds ... 29
1.4.1 Conventional fabrication techniques ... 30
1.4.2 Rapid prototyping techniques ... 34
1.4.3 Low-temperature deposition manufacturing. ... 42
1.5 Aim of the thesis ... 45
2. MATERIALS AND METHODS: Fabrication and characterization of micro-porous films and 3D scaffolds ... 47
2.1 Materials ... 47
2.2 Methods ... 47
2.3 Synthesis of three-armed PTMC-MA ... 57
2.4 Preparation of photo-crosslinked PTMC-MA and nHA composite networks ... 59
2.5 Icariin-loaded microporous composite films ... 60
2.5.1 Icariin dispersion ... 60
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2.6 Low-temperature Deposition Modelling ... 62
2.6.1 Optimization of the printing parameters ... 63
2.7 Ink and film characterization ... 66
2.7.1 Thermogravimetric analysis ... 66
2.7.2 Porosity and swelling degree measurement ... 67
2.7.3 Mechanical Properties ... 67
2.7.4 Rheological properties ... 68
2.7.5 Micro porosity and surface roughness ... 69
2.8 Scaffold characterization ... 69
2.9 Characterization of icariin-loaded films ... 69
2.9.1 Icariin extraction in methanol ... 69
2.9.2 Drug release experiments ... 71
2.9.3 Extra experiments ... 72
3 MATERIALS AND METHODS: Effect of micro-porosity and surface roughness on hBMSCs osteogenic differentiantion ... 73
3.1 Materials ... 73
3.2 Additional film characterization ... 74
3.2.1 Calcium release ... 74
3.2.2 Protein adsorption ... 74
3.3 In vitro cell investigation ... 75
3.3.1 Cell adhesion ... 75
3.3.2 Metabolic activity ... 76
3.3.3 Cell viability ... 76
3.3.4 Cell adhesion, proliferation and spreading imaged by SEM ... 76
3.3.5 ALP activity and staining ... 77
3.3.6 Degree of mineralization ... 78 4- RESULTS AND DISCUSSION: Characterization of ink, films and scaffolds
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4.1 Synthesis of three-armed PTMC-MA ... 79
4.2 Rheological properties of the ink ... 81
4.3 Physical characterization of PTMC-MA/nHA composite material and microporous films ... 82
4.4 Microstructural and surface characterization ... 83
4.5 Mechanical Properties ... 86
4.6 Printing optimization ... 89
4.7 Analysis of printed scaffolds ... 91
5. RESULTS AND DISCUSSION: Effect of micro-porosity and surface roughness on hBMSCs osteogenic differentiation ... 97
5.1 Film characterization: calcium release and protein adsorption ... 97
5.2 In vitro differentiation investigation ... 98
5.2.1 Cell viability and proliferation ... 98
... Errore. Il segnalibro non è definito. 5.2.2 Osteogenic differentiation ... 102
6. RESULTS AND DISCUSSION: Characterization of icariin-loaded films loaded 107 6.1 Characterization of photo-crosslinked PTMC-MA/nHA-IC films ... 107
6.1.1 Physical Characterization ... 107
6.2 Mechanical characterization ... 108
6.2.1 Icariin loading and fate in PTMC-MA/nHA-IC films ... 110
6.3 Characterization of icariin swell-coated films ... 113
6.3.1 Physical and mechanical properties ... 113
6.3.2 Icariin extraction in MeoH ... 113
6.3.3 Drug release ... 113
Conclusions ... 116
APPENDIX A ... 118
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Introduction
Bone tissue is an essential organ playing key roles in several critical functions in human physiology including movement and support, protection of other critical organs (such as the brain), blood production and pH regulation, mineral storage and others. Bone diseases such as osteogenesis imperfecta, osteoarthritis, fractures and primary tumor resection lead to or induce complex bone defects or voids. These cannot be naturally restored by the natural bone and represent a challenge to repair. The conventional orthopaedic treatments, involves transplantation of autograft and allografts. Even though autografts are considered as the ‘gold standard’ material for bone tissue repair, they are limited in size and availability. Moreover they often result in donor site morbidity. The major risk associate with the use of allografts is the risk of disease and infection transmission and of immunogenic response. Moreover, and in spite of the fact that material science technology has resulted in clear improvements in the field of bone substitution medicine, no adequate bone substitute has been developed and hence large bone defects/injuries still represent a major challenge for orthopaedic and reconstructive surgeons. In these context bone tissue engineering is emerging as a novel and valid approach to the current therapies for bone regeneration and substitution. The scaffolds plays a critical role in the successful engineering for bone tissue, providing a suitable environment for cell migration, proliferation and differentiation towards the desired tissue phenotype. An ideal scaffold for bone tissue applications should mimic the morphology, structure, physicochemical properties and functions of native bone. Moreover it should be made of a biodegradable material so as to create a template to guide and promote the formation of new bone and de novo vascularized bone tissue in vivo. Thus the choice of the materials and the technique of fabrications are essential to the success of the final scaffold. The combination of additive manufacturing techniques with bioresorbable composite materials open up possibilities for the fabrication of micro-porous scaffold with a designed macro-porosity for the reconstruction of bone defects. Apart from a designed macro-porosity for tissue ingrowth, a micro-porosity in the implant struts will be beneficial for nutrient diffusion, protein adsorption and drug loading and release.
In this work three bioresorbable, micro-porous and photo-crosslinked composite materials of poly(trimethylene) carbonate and nano-sized hydroxyapatite, with different porosities (of 44, 52 and 71 %) and surface roughness were produced. The ethylene
7 carbonate was used as solvent and porogen. To allow a straightforward comparison between the different surface roughness and micro-porosities, the three composites were be cast in the shape of films and subsequently characterized mechanically, thermally, microscopically and gravimetrically.
In vitro experiments were subsequently carried out in order to investigate the effects of different porosities and surface roughness on the osteogenic differentiation of human bone marrow mesenchymal stem cells. Cell adhesion, cell viability, degree of mineralization and osteogenic differentiation were investigated.
In this work we also report on a novel low deposition modelling setup for the extrusion of the different composite materials. A new cooling system and an anti-condensation system were designed in order to promote the crystallization of the ethylene carbonate and the removal of water on the working space. After the crystallization of the ethylene carbonate and the photo-crosslinking of the poly(trimethylene) carbonate matrix, the ethylene carbonate was extracted in water. With this innovative technique homogenous 3D micro- and macro-porous scaffolds can obtained.
Finally in order to increase the osteinductivity of the composite material a natural drug, icariin, was loaded by two different methods. This new material was then characterized mechanically, thermally, microscopically and gravimetrically.
The thesis is organized as follows.
Chapter 1: introduces the background of bone anatomy and physiology. It provides an overview of the current materials and biofabrication techniques used in bone tissue engineering with their advantages and limitations.
Chapter 2: presents the materials and methods used to produce the composite material with and without icariin. Furthermore a description of the novel low deposition modelling method used to fabricate 3D scaffolds is given.
Chapter 3: presents the materials and methods used for in vitro experiments using human bone marrow mesenchymal stem cells seeded on the three films with different porosities and roughness .
Chapter 4: presents and discusses the results obtained for the mechanical, thermal, microscopical and gravimetrical characterization of the composite materials, the micro-porous films and micro- and macro-porous 3D structures.
8 Chapter 5: provides the results obtained for in vitro experiments described in chapter 3.
Chapter 6: presents and discusses the results obtained for the mechanical, thermal, microscopical and gravimetrical characterization of the composite materials, the micro-porous films loaded with icariin.
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1. Bone tissue engineering
Tissue engineering is an innovative scientific field whose first definition dates back to 1987[1]. It is an interdisciplinary field which combines the knowledge from different fields like medicine, material science and engineering in order to develop biological substitutes or scaffolds that restore, maintain or improve tissue function [2-4]. Its application on bone tissue, hence the name Bone Tissue Engineering (BTE), represents an alternative and innovative strategy for applications in regenerative medicine. BTE principles can be applied to treat critical sized bone defects caused by trauma, tumours, or congenital skeletal disorders that cannot self-regenerate, promoting the natural regeneration process of the bone. Its goal is to eliminate all the disadvantages of the conventional orthopaedic treatments, which involves transplantation of autografts and allograft. Autografts are considered as the ‘gold standard’ material for bone tissue repair. However the amount of autografts from patient bone, with adequate properties, is not always sufficient. Moreover extra surgery, to harvest the bone from a donor site, is needed with possible complications associated with it. The major risk associate with the use of allografts is a disease and infection transmission from donor to patient (e.g. hepatitis C and HIV). Therefore their use often involve the need for immunosuppressant drug administration[3-5]. A tissue engineering scaffold plays an essential role in regenerating bone tissue. As the natural extracellular matrix (ECM) surrounding the cells, the scaffold represents the framework for dissociated cells to reform an appropriate tissue structure.
In order to control and to incite the bone regrowth, an optimal scaffold should have similar mechanical properties and topographic characteristics to the natural bone. Moreover it is optimally made of a biodegradable material so as to create a template to guide and promote the formation of new bone and de novo vascularized bone tissue [6].
To induce the formation of new tissue multiple strategies have been adopted such as cell transplantation, acellular scaffolds, gene therapy, stem cells therapy and growth factor delivery[7]. However, two primary tissue engineering strategies have been emerged as the most promising approaches:
i) Implantation of cellularized scaffold (Figure 1.1). Before the implantation, mesenchymal stem cells (MSCs), isolated from the patient, are firstly expanded in vitro on two-dimensional surfaces and then seeded onto a 3D synthetic scaffold. The cell constructs are further cultivated in bioreactors to provide
10 optimal and controlled conditions for organization into a functioning tissue. Once the a functioning tissue has been successfully engineered, the construct is transplanted into the osseous defect to restore functions[2,6-7].
ii) Implantation of acellular scaffold, in the damaged site immediately after the injury. The advantage of this approach is the reduced number of surgeries needed, resulting in a shorter recovery time for the patient and reduced risks of surgical problems[2,6-8].
In both techniques the scaffold could be loaded by growth factors, small molecules and micro- and/or nano-particles in order to increase bone recovering capabilities.
Figure 1. 1 Scheme of the tissue engineering approach that involves seeding cells within a 3D porous biodegradable scaffold and the following implantation in the human body [9].
After the scaffold is implanted into the host it can be vascularized naturally, allowing the cells to continue their growth. Since the scaffold material is biodegradable,
11 the scaffold is resorbed over time, leaving three-dimensional structures of functional cells from the donor, blood vessels and supporting tissue [6].
In this work, the second strategy of acellular scaffolds will be applied. A more detailed description of this process will be given in the following chapters (chapter 2). To do so, an overview is given of the structure of bone, the role of a scaffold in bone regeneration and materials and processing methods for scaffolds.
1.1 Bone Tissue
1.1.1 Bone components and physiology
Bone is a complex tissue which carries out several important functions in the body: (i) it provides a structural and mechanical support to soft tissues serving as a lever for muscle action, (ii) it protects vital organs like the brain and the spinal cord, (iii) it supports haematopoiesis and intervenes in the blood pH regulation and (iv) it stores healing cells and mineral ions [7, 10,11].
Bone tissue is made of water, organic and inorganic components. They represent respectively 20 %, 45 % and 35 % of the wet weight of mature cortical bone. In detail, bone tissue is made of a fibrous, organic matrix, which is permeated by large stores of inorganic salts. The organic component of bone is made of type I collagen (90-95%), glycosaminoglicans (l%) and proteins (5%). The inorganic part of bone consists mainly of a mixture of hydroxyapatite (HA) nano-scale crystals (Ca10(PO4)6(OH)2) [12-13] and calcium carbonate with lesser quantities of sodium, magnesium and fluoride.
Macroscopically bones are generally classified by their shape: long, short flat or irregular. Considering the porosity it is possible to classify bone as cortical and cancellous. Cortical bone is dense and compact with a porosity of approximately of 5-30%. It is composed of closely packed osteons (haversian system) so tightly assembled together that it appears as a solid mass. The osteon is made of a central canal, called osteonic or haversian canal, which is surrounded by concentric rings (lamellae) of matrix. Between these rings of matrix, bone cells (osteocytes) are located in spaces called lacunae. Small channels (canaliculi) radiate from the lacunae to the osteonic canal to provide passageways through the hard matrix. The osteonic canals contain blood vessel, parallel to the long axis of the bone, that are interconnected with vessels on the surface of the bone (Figure 1.2). The cortical bone, which comprises 80% of the skeleton, mostly provides mechanical strength, having high resistance to bending and
12 torsion, and protection. Bone represents a reservoir of several minerals (such as calcium phosphate, magnesium fluoride) in the human body. In case of a severe or prolonged mineral deficit it can also participate in metabolic responses [7,14].
Cancellous bone is less dense and more elastic than compact bone with a porosity ranging from approximately 5 to more than 90%. Mostly present in the epiphyses of bones, it consists of a 3-dimensional network of bony plates (trabeculae) and struts. Trabeculae divide the interior volume of bone into intercommunicating pores of different dimensions producing a structure of variable porosity and apparent density[12]. It may appear that trabeculae are arranged in a random manner instead they are organized to provide maximum strength. Trabeculae represent the structural support of the cancellous bone. They follow indeed the lines of stress and can realign if the direction of stress changes over prolonged time [7]. Cancellous bone represents 20% of the skeletal mass and it mainly contributes to mechanical support, particularly in bones such as the vertebrae. Moreover it provides an initial supply of minerals in acute deficiency states [14].
Figure 1. 2 Internal organization of the bone in a representative long bone. The basic anatomy and differences in cortical and cancellous bone (left) and micro-structural features in bone (middle and right) are illustrated [15].
Although the cortical and the trabecular bone are macroscopically and microscopically different, the two forms of bone are identical in their chemical composition [7,14].
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1.1.2 Cellular classification and bone remodelling process
Considering their functions and morphology, bone cells are classified in: osteoblasts, osteoclasts and osteocytes.
Osteoblasts are differentiated from multipotent mesenchymal stem cells. Their
primary function is to produce and secrete organic and inorganic bone ECM, known as the osteoid, and to regulate osteoclasts activity. Toward the end of the matrix-secreting period, 15% of mature osteoblasts are entrapped in their own calcified matrix and they change their phenotype matrix differentiating into osteocytes. The remaining osteoblasts remain on the bone surface, becoming flat lining cells [7,10,14,16].
Osteocytes are the most abundant cells in bone, derived from osteoblast. They possess a high number of cytoplasmatic extensions or filopodia. These serve to interconnect the osteocytes and to connect them with the bone-lining cells, forming a network of thin canaliculi permeating the entire bone matrix. Their exact function is not completely clear, eventually they stop generating osteoid and play a key role in mechanotransduction [7,14,16].
Finally, osteoclasts are giant multinucleated cells derived from fusion of mononuclear hemopoietic stem cells. Their primary function is to secrete acids and proteolytic enzymes which erode bone ECM under the influence of chemical cues. Osteoclasts form deep folds in their plasma membrane in the area facing the bone matrix and the surrounding zone of attachment (called sealing zone).An activated osteoclast is able to resorb 200,000 μm3/day, an amount of bone formed by seven to ten generations of osteoblasts [7,14,16].
The process of production and resorption of bone ECM, regulated by the coupled action of osteoblasts and osteoclasts, is called bone remodelling. It is responsible for maintaining skeletal integrity repairing defects like micro-fractures, healing, blood calcium regulation, and accommodation of changes in bone stress profiles. Moreover it regulates the adjustment of the architectural and hence mechanical properties of bone as a function of mechanical and chemical signalling.
Bone remodelling is a complex process characterized by three distinct phases: resorption of ECM by osteoclast, reversal, and formation of ECM by osteoblasts. Resorption is initiated with the migration of partially differentiated mononuclear preosteoclasts to the bone surface where they form multinucleated osteoclasts. Osteoclasts resorb bone by acidification and proteolysis of the bone matrix and of the
14 hydroxyapatite crystals encapsulated within the sealing zone of the osteoclasts. Hereafter, during the reversal phase, mononuclear cells appear on the bone surface. These cells prepare the surface for new osteoblasts to begin bone formation and provide signals for osteoblast differentiation and migration. Then osteoblasts firstly produce osteoid by rapidly depositing collagen. This is followed by an increase in the mineralization rate to equal that of collagen synthesis. In the final stage the rate of collagen synthesis decreases and mineralization continues until the osteoid becomes fully mineralized. Osteoblasts lay down bone until the resorbed bone is completely replaced by new bone.
In a homeostatic equilibrium the phases of bone resorption and formation are balanced, occurring at the same rate, so that old bone is continuously replaced by new tissue in order to adapt the bone properties to mechanical load and strain. The stages of the remodelling cycle have different durations. Resorption probably continues for about 2 weeks, the reversal phase may last up to 4 or 5 weeks, while formation can continue for 4 months until the new bone structural unit is completely created.[14,16-17]
1.1.3 Mechanical properties of bone
The mechanical properties of bone strongly depend on its sophisticated macro-, micro-and nano-scale hierarchical structure (Figure 1.3). For instance the macro-scale organization of osteons, osteoid and haversian canals provide long bones with their characteristic mechanical anisotropy [7]. Another example are the HA crystals in bone, which have a length of 30–200 nm, thickness of 2–10 nm [12-13] and bind type I collagen (300 µm long). Both the size and orientation of the HA crystals are directed by the collagen template, and the precise structural relationship between the collagen and HA is critical to the resilience and strength of bone [7].
15 Figure 1. 3 Hierarchical structural organization of the bone [18]
Thus, bone is globally an anisotropic material, since its elastic properties and strength are dependent upon the orientation of the bone microstructure with respect to the direction of loading. In addition, bone is a viscoelastic material, since its elastic properties and strength are dependent upon both the rate and the duration of the applied loading. For instance bone elastic modulus increases with the loading rate [12, 19].
Table 1.1 shows the mechanical properties of bone. Cortical bone is more stiff and anisotropic than trabecular bone. It shows indeed, two different elastic moduli (E) depending on the direction of the stress (i.e. along the longitudinal and transverse axis). The mechanical properties of trabecular bone vary widely and are a function of the density and porosity of the trabeculae [7].
Table 1.1 Mechanical properties of cortical and trabecular bone [12, 20] .The elastic modulus of the bone has been indicated with letter E.
Property Cortical bone Trabecular bone
Longitudinal E (GPa) 17 0.5 – 0.05
Transversal E (GPa) 11.5 0.5 – 0.05
Compressive strength (MPa) 100 - 230 2 -12 Tensile Strength (MPa) 50 – 150 10 - 20 Fracture toughness (MPa m ½) 2 – 12 -
Strain to failure (%) 1 - 3 5 - 7
Moreover the mechanical properties are a function of the age and health state of the bone. Osteoporosis and osteoarthritis for instance induce a change in the bone in
16 terms of composition, porosity and degree of mineralization [21]. All these factors significantly affect the elastic modulus of bone. Furthermore mechanical properties of bone increase markedly from birth up to 40 years of age because of the increasing degree of mineralization. On the contrary in elderly people (average age 71) bone becomes more porous and less dense, losing its mechanical properties in terms of elastic modulus, strength and deformation to failure [12].
1.2 The scaffold: a key role in BTE
Scaffolds play the central role in BTE and its properties strongly influence the outcome of de novo bone re-growth. It temporary replaces the ECM of the bone tissue. ECM provides the cells a specific environment and a 3D architecture. Furthermore, it serves as a reservoir of water, nutrients and growth factors. The scaffold, acts as a temporary matrix for the cell proliferation and natural ECM deposition [4].
An ideal BTE scaffold should have the following properties:
(i) BIOCOMPATIBILITY and BIODEGRADABILITY: Scaffolds should be well integrated in the host’s tissue and should degrade without eliciting an immune response and adverse effects. In order to prevent biological response like sepsis or cancer recurrence it is possible to load a scaffold with active molecules like antibiotics [7]. The degradation should be controllable in order to match the natural tissue growth rate of the host. The degradation and the resorption kinetics have to be controlled is such a way that the bioresorbable scaffold retains its physical properties for at least 3-6 months. Thereafter the scaffold matrix can start losing its mechanical properties and should be metabolized by the body without a reaction after 12-18 months. During this time, mechanical properties should still be sufficient to support the newly formed tissue. It should possess sufficient strength and stiffness to function for a period until in vivo tissue ingrowth has replaced the slowly vanishing scaffold matrix[4, 20].
(ii) OSTEOGENEITY, OSTEOCONDUCTIVITY, OSTEOINDUCTIVITY and OSTEOPROMOTION: When the portion of bone to regenerate is large, natural osteoinduction, -conduction and -geneity of bone combined with a biodegradable scaffold may not be enough. Because of this the scaffold should present the aforementioned properties [4]. Osteogeneity is the capability of the scaffold to form bone tissue de novo [22]. Osteoconductivity is the capacity of a
17 material to guide bone forming tissue into a defect [23]. Osteoinduction is the process by which stem and osteoprogenitor cells are recruited to a bone healing site, and stimulated to undergo the osteogenic differentiation pathway.[4,24]. In order to regenerate a bony defect a scaffold should be at least osteoconductive.
In this regard the scaffold can release in a controlled way bioactive molecules like growth factors (e.g. bone morphogenetic proteins (BMPs) [4,7] or natural drugs like icariin[25,26]. These last molecules can have a role in osteoinduction, osteoconduction and osteopromotion (ability of the material to promote the de novo bone formation), they also accelerate ECM production and influence the subsequent tissue integration [4].
(iii) POROSITY and PORE SIZE: Porosity and pore size both at the macroscopic and the microscopic level, are important morphological properties of a scaffold for bone regeneration [27]. Indeed the presence of pores in the scaffolds increases the surface area which promotes mass transfer, cell mobility and in-growth, vascularization and consequently tissue development. Pores should be interconnected for an effective diffusion of nutrients and for the removal of metabolic waste resulting from the activity of cells that had meanwhile grown into the scaffold. This is of particular importance regarding bone tissue engineering because, due to bone metabolic characteristics, high rates of mass transfer are expected to occur, even under in vitro culture conditions. However, the degree of porosity influences other properties of the scaffolds such as the rate of biodegradability and mechanical stability. [4,6-7,20] Scaffolds fabricated from biomaterials with a high degradation rate should not have high porosities (90%), since rapid depletion of the biomaterial will occur compromising the mechanical and structural integrity before substitution by new bone. In contrast, scaffolds fabricated from biomaterials with low degradation rates and robust mechanical properties can be highly porous, because the higher pore surface area interacting with the host tissue can accelerate degradation due to macrophages via oxidation and/or hydrolysis. It has been shown that in vitro lower porosity enhances osteogenesis due to cell aggregation and suppressed proliferation. In vivo high porosity and large pores enhance bone ingrowth and osteointegration of the implant after surgery [27].The minimum recommended pore size for a scaffold is 100 μm due to cell size, migration requirements and transport. However other studies demonstrated that implants with pores > 300
18 μm allow better vascularization and higher oxygenation and consequently better osteogenesis.
There is an upper limit in porosity and pore size set by constraints associated with mechanical properties. An increase in the void volume results in a reduction in mechanical strength of the scaffold, for this reason the porosity should always be balanced with the mechanical requirements of the bone [4,20,27]. (iv) SURFACE PROPERTIES: Surface properties, both chemical and topographical, can have consequences on the cellular adhesion and proliferation. Chemical interactions are governing the adherence of cells and proteins to the scaffolds material.Topographical properties on nano-/micro-scale formulate binding sites to actively regulate and control cell and tissue behaviour while interacting with host cells. Scaffolds should also possess a similar macro structure to what is found in natural bone. This feature may provide space for the growth of cells and new tissues, as well as for the carriers of growth factors [28].
An important topographical property is the surface roughness of the scaffold. A rougher surface enhances attachment, proliferation and differentiation of anchorage dependent bone forming cells. As surface roughness increases the surface area, it contributes to more bone inducing protein adsorption [27-29].
(v) MECHANICAL PROPERTIES: The mechanical features of the scaffold can significantly influence its osteointegration with surrounding tissues, as well as cell behaviour on the scaffold surface [28]. In vitro, scaffolds should have sufficient mechanical strength to withstand the hydrostatic pressures and to maintain the spaces required for cell in-growth and matrix production. The mechanical properties of the implanted construct should ideally match those of living bone in vivo. In this way the scaffold can provide temporary mechanical support in the injured site without showing symptoms of fatigue or failure. Moreover they should be retained in the space they were designated for and to provide the tissue with adequate space for growth. A skeletal scaffold should be designed in such a way that stresses are evenly distributed to produce the equivalent of native bone [4,7,20].
(i) STERILITY and REPRODUCIBILITY: the scaffold should be sterilizable without loss of bioactivity. For clinical applications, it must be possible to
19 manufacture the scaffold with a reproducible and controlled technique at an economic cost and speed [7,20].
To meet these requirements it is important to not only apply an adequate material but also the processing technique used to fabricate the scaffold. However it is apparent from the preceding requirements, that some are contradicting, for example the desired high porosity and pore size or the degradability and the mechanical properties. This leads to the conclusion that designing an ideal BTE scaffold always involves a trade off between the different requirements.
1.3 Materials
The choice of the most appropriate material to produce a scaffold for tissue engineering applications is one of the most important aspects to be considered, since its properties will determine, to a great extent, the properties of the final scaffold. Up to now several materials such as metals, ceramics and polymers from both natural or synthetic origins have been proposed. However, metals and most of the ceramics are not biodegradable, thus the selection is reduced to a small number of ceramics and to biodegradable polymers and composite materials [4].
1.3.1 Biodegradable polymers
There are two types of biodegradable polymers: natural and synthetic. Natural polymers include polysaccharides (starch, alginate, chitin/chitosan, hyaluronic acid derivates) or proteins (soy, collagen, fibrin gels, silk). Synthetic biodegradable polymers include saturated aliphatic polyesters, polyfumarates, polyhydroxyalkanoates and polycarbonates. Synthetic polymers show several advantages: they can be produced under controlled conditions and exhibit predictable and reproducible mechanical and physical properties such as tensile strength, elastic modulus and degradation rate. Moreover it is possible to control the amount of impurities in the material, thus reducing risks such as toxicity, immunogenicity and infections. [8]
This paragraph provides an overview on the synthetic biodegradable polymers used for 3D scaffolds in BTE. Particularly it will focused on the properties of poly(trimethylene carbonate) (PTMC) and its applications, since this material will be used in the project described in this thesis.
20 - SATURATED ALIPHATIC POLYESTERS
Saturated aliphatic polyesters include poly(lactic acid) (PLA), poly(glycolic acid) (PGA), as well as poly(lactic-co-glycolide) (PLGA) copolymers and poly(ε-caprolactone) (PCL).
PGA, PLA and PCL are the most used biodegradable materials for 3D scaffolds in tissue engineering due to their biocompatibility, tuneable degradation and mechanical properties. A feature of particular interest of polyester-based materials is the removal of monomeric components in vivo by natural pathways. The body contains indeed highly regulated mechanisms for completely removing the formed monomeric components after their degradation. The main drawback of this group of polymers is the release of acidic degradation products. The degradation of polyester-based materials involves hydrolysis of the ester bonds leading to chain scission which forms acidic by-products. This can cause a strong inflammatory response in vivo. In addition all these polymers undergo a bulk erosion process such that it can cause scaffolds to fail prematurely. By to copolymerization of lactide and glycolede, there is better control over degradation rates and the mechanical properties of PLGA. It has therefore been used in biodegradable sutures which have been approved by US Food and Drug Administration. On the other hand PCL with high molecular weight (e.g. 50000 g/mol) degrades slowly requiring 3 years for complete removal from the host body. For this reason PCL is used for long term applications [8].
- POLYFUMARATES
Polyfumarates are unsaturated linear polyesters. In BTE applications, mostly poly(propylene fumarate) (PPF) is applied. Like the previous group of polymers, the degradation products of PPF (i.e. propylene glycol and fumaric acid) are biocompatible and readily removed from the body.
The double bond along the backbone of the polymer permits cross-linking in situ, which causes mouldable composites of PPF to harden within 10–15 min. For this reason PPF has been suggested for use as an injectable bone replacement. It also has been used as a substrate for osteoblast cultures. The critical issue of PPF is the preservation of the double bonds and the control of molecular weight during synthesis.[8]
21 Polyhydroxyalkanoates (PHA) are a group of aliphatic polyesters as well, but produced by microorganisms under unbalanced growth conditions. They are generally biodegraded via hydrolysis and are thermo processable, making them attractive as biomaterials for applications in medical devices and bone tissue engineering.
The mechanical properties and the biocompatibility of PHA polymers can be modified combining them with other polymers, enzymes or inorganic materials. However, a drawback of some PHA polymers is their limited availability and the time-consuming extraction procedure from bacterial cultures that is required for obtaining sufficient processing amounts [8].
- POLYCARBONATES
Polycarbonates that are largely applied in bony defects are rigid materials such as poly(bisphenol A carbonate) (poly(BPAC)). This polymer is known for its good processability and excellent mechanical properties. However, poly(BPAC) is essentially non-degradable [70]. An interesting, biodegradable and novel alternative polycarbonate for BTE is PTMC.
Figure 1. 4 Chemical structure of linear PTMC (a) and three-armed PTMC functionalized by methacrylic anhydride (b)
PTMC is a linear polymer (Figure 1.4a) obtained by ring opening polymerization of the cyclic trimethylene carbonate (TMC). It is a hydrophobic and amorphous polymer with a glass transition temperature (Tg) of around -20°C [30].
PTMC is a biocompatible polymer characterized by tuneable chemical and physical properties that make it a good candidate for several tissue engineering applications. By changing the molecular weight of PTMC ( Mn ), it is possible to modulate its mechanical properties and the degradation rates. Pêgo et al. demonstrated that the mechanical properties of PTMC (e.g. elastic modulus, strain at yield, strain at break) increase with increasing molecular weight (Table 2). Low molecular weight
22 PTMC, (with Mn below 100,000), possesses inferior mechanical properties to very high molecular weight PTMC (with Mn above 200,000) (Table 1.2). Low molecular weight PTMC possesses indeed a high number of chain ends per unit volume which reduce the packing efficiency of the polymer chains preventing crystallization. With an increase in molecular weight, a higher percentage of the structure can be permanently oriented by stretching, resulting in a higher ultimate tensile strength. Thus high molecular weight PTMC is flexible and tough, with rubber-like properties, showing good recovery after mechanical deformation [30].
,
Because of its mechanical features, linear PTMC matches the mechanical behaviour of human soft tissues. In this regard linear PTMC has been used to fabricate vascular tissue scaffolds[31].
One of the most important properties of PTMC is that, differently from saturated aliphatic polyesters, it degrades by enzymatic surface erosion in vitro and in vivo. This confers higher and mainly more durable mechanical stability to PTMC-based scaffolds. Moreover no formation of acidic degradation products occurs. This reduces the risk of an inflammatory response in the host’s body [32]. The degradation rate of PTMC is influenced both in vitro and in vivo by the molecular weight of PTMC. In both cases the rate mass loss of high molecular weight PTMC specimens is higher than that of lower molecular weight [32-33].
In order to improve the performances of linear PTMC, in terms of mechanical properties and rate of degradation, different approaches have been explored. These include copolymerization, addition of calcium phosphates[71] to obtain composite materials, and end-group functionalization of PTMC oligomers with methacrylate groups (PTMC-MA) and subsequent photo-crosslinking.
Table 1.2 Molecular weights ( Mn) and stress–strain behavior of compression molded PTMC films [30]. Particularly
σyield, σbreak and σmax are respectively the stress at yield, at break and the maximum stress. εyield, and εbreak are
23 By photo-crosslinking PTMC-MA with UV light [30] form-stable networks with high resistance to creep properties were obtained [33]. Varying the macromer molecular weight, PTMC networks with a wide range of properties can be prepared [36]. All PTMC networks are transparent and amorphous and their Tg increases with macromer molecular weight, ranging from 8 °C for networks prepared from 1000 g/mol macromers to -15 °C for networks prepared from 41,000 g/mol macromers. The decrease in glass transition temperature is due to the increasing mobility of longer network chains.
With increasing macromer molecular weight the elastic modulus of the networks decreases from 314 MPa (networks prepared from macromers with molecular weights lower than 1800 g/mol) to 5 MPa (networks prepared from macromers with molecular weights higher than 10000 g/mol), while their elongation at break increases reaching a very high value of 1200 % (Figure 1.5). Thus low molecular weight PTMC macromers yield relatively stiff and brittle networks with low elongation at break. On the contrary networks obtained from high molecular weight macromers result in flexible, tear resistant and rubber-like elastomers Therefore networks obtained from low molecular weight PTMC macromers are promising candidates for use as biodegradable bone tissue implants. Whereas the tough and elastic networks prepared with high molecular weight macromers are very well suited for use as biodegradable soft tissue implants like blood vessel scaffolds [34].
Figure 1. 5 Tensile stress–strain curves for PTMC networks obtained by photo-crosslinking three-armed macromers of different molecular weights (The number above each stress-strain curve represents the molecular weights of the macromers in g/mol) [34]
By copolymerization of TMC with cyclic esters like ε-caprolactone [35] or D,L-lactide [36] it is possible to increase or reduce the rate mass loss and surface erosion
24 typical of PTMC. This makes PTMC attractive for its versatility to prepare short- and long-term degradable devices for tissue engineering.
The molecular weight of both linear PTMC and of macromers for photo-crosslinked PTMC-MA networks affects the degradation rate of the final PTMC product. Rongen et al. [37] showed that PTMC-MA networks with higher molecular weights (27 kg/mol) are characterized by an increased degradation rate compared to networks prepared from macromers of the lowest molecular weight (13 kg/mol). Crosslinking macromers of lower molecular weight results, indeed, in shorter PTMC chains between poly(methyl methacrylate) crosslinking chains and an increased crosslinking density. This could limit the accessibility for enzymatic attack on the PTMC chain.
PTMC is well suited to make composite materials for BTE applications. Geven et al. prepared a composite material made of low molecular weight, three-armed PTMC-MA of 10000 g/mol and nano-sized hydroxyapatite crystals (nHA) [38]. This material has been processed by low-temperature deposition modelling [39] and stereolithography [40]; obtaining in the last case patient-specific composite orbital floor implants. It was shown that…
1.3.2 Bioactive ceramics
Bioactive ceramics are a group of materials that include calcium phosphate-based materials and bioglasses. Bioactivity is the feature shared by all the materials belonging to this category. Bioactivity is the ability of the material to directly bond with the newly formed bone [41]. Upon the implantation of the bioactive implant, the surface forms a biologically active hydroxyl carbonate apatite layer which provides the bonding interface with tissues. This layer is chemically and structurally equivalent to the mineral phase in bone, providing interfacial bonding [8].
- BIOACTIVE GLASS
The basic constituents of most types of bioactive glass are silicon dioxide (SiO2), sodium oxide (Na2O), calcium oxide (CaO), and phosphorus pentoxide (P2O5) [8].
One key reason that makes bioactive glass a relevant scaffold material is the possibility of controlling the rate of bioresorption by varying either the composition or
25 the thermal or environmental processing history, [8]. Bioactive glass shows similar properties to natural bone in terms of composition, osteoconductivity and cell delivery capabilities [28]. Furthermore bioactive glass has been found to support enzyme activity, vascularization, osteoblast adhesion, growth and differentiation. The above-mentioned advantages are the reasons why bioactive glass is successfully used in clinical treatments of periodontal disease and as a bone filler material. Bioactive glass implants have also been used to replace damaged middle ear bones, restoring hearing to patients [8].
However, a drawback of bioactive glass is the low fracture toughness and mechanical strength, especially when the glass is in a porous form. Hence, bioactive glass alone has limited application in load-bearing situations and it is more often applied in a composite [8,41].
- CALCIUM PHOSPATE-BASED BIOMATERIALS
Calcium phosphate-based biomaterials (CP) is a group of materials which includes hydroxyapatite (HA), β-tricalcium phosphate (Ca3(PO4)2) (β-TCP) and biphasic calcium phosphates which are obtained by a combination of HA and β-TCP.
This class of material is widely used in BTE because of their similarity in composition to the bone mineral (nano-crystals of apatite). Furthermore these materials are similar to the natural bone in terms of biodegradability, osteoconductivity and bioactivity. Due to this last property, CP promote formation of a carbonate HA layer on the surface of the implanted scaffold. This layer attracts protein to which cells adhere, proliferate, and differentiate, leading to matrix production and bio-mineralization and thus formation of new bone [41].
Although CP show several advantages, they have some major drawbacks. They present a low mechanical stability showing a brittle and fragile behaviour and they are difficult to process. For these reasons they are not suitable to be used in the regeneration of large bone defects. Furthermore, due to in vivo biological activity, such as the osteoclastic activity which may vary by patient and defect, their degradation/dissolution rates are difficult to predict. This could compromise the mechanical stability of the construct, which is already low by itself, if it degrades faster than expected. At the same time, this would dramatically increase the extracellular concentrations of calcium and phosphorus which can cause cellular death [4].
26 Additionally CP, by themselves, are not osteoinductive. However, osteoinductive properties can be introduced to CP by two methods:
- designing the CP with appropriate geometry, topography and with proper features such as macroporosity and/or microporosity which allow for the entrapment and concentration of circulating growth factors or osteoprogenitor cells responsible for bone formation[41]. Combining CP with growth factors (e.g. bone morphogenetic proteins or mesenchymal cells) or bioactive proteins (such as collagen) or drugs like icariin. [41].
1.3.3 Composite of biodegradable polymers with nHA
Composite materials, obtained combining biodegradable polymers and bioactive CP, allow to mash the advantages of the two classes of materials in order to obtain optimized biodegradable andosteoconductive scaffolds for BTE. In the last years several combinations of polymers and CP have been used. Among the calcium phosphate materials the most used for BTE scaffolds are HA and β-TCP.
In our application we decided to use nHA. As previously described in the paragraph 1.1.3 the nanometer size of the HA in natural bone affects the mechanical properties of the bone. It has been demonstrated that better osteoconductivity could be achieved if CP crystals would resemble bone minerals not only in composition but also in size and morphology. In this regard nano-sized HA can have beneficial properties due to its small size and high surface area. Moreover protein adsorption and osteoblast adhesion increases on materials based on nano-sized ceramic crystals compared to those based on traditional micron-sized ceramic crystals [42-43]. As evidenced by Zhang et al. [44-45] CP drive the osteogenesis process in host bone through its surface architecture. The presence of CP in a composite material will influence the topographic features of the surface of the scaffold, such as the surface roughness. This affects bone formation by virtue of the enlarged surface area of the scaffold, since it results in higher protein adsorption, higher calcium and phosphate ion release and induced surface mineralization [44]. Particularly, with decreasing dimension of the CP crystals in a composite scaffold the surface area that affects the osteogenic differentiation of mesenchymal stem cells will increase and subsequently higher osteoinductive potential of the final scaffold in vivo can be expected. [45] Wei et al. have shown…
27 Figure 1. 6 TEM image of needle-like apatite crystal used in this thesis [46].
For these reasons we decided to use needle-like HA crystals of 200 to 400 nm long and 20 to 50 nm wide. These dimensions are in the same size range as the HA in bone. (Figure 1.6)
The presence of HA in the composite affects the surface architecture of the scaffold, playing a role in the osteoinduction. Precisely, the incorporation of nano-sized HA into PLLA micro-porous scaffold improved the mechanical properties and the protein adsorption [47]. Barbieri et al. showed that the presence of nHA (40 wt % ) into poly(D,L-lactide) provides a structural composite with a microstructured surface enhancing calcium release and surface mineralization which ultimately lead to an osteoinductive material [48]..
Also other authors have shown that the presence of nHA in a composite enhances the mechanical properties of the scaffold.[38]. Particularly, with an increasing content of nHA in the composite the elastic modulus, yield strength and tensile strength increases. Therefore the higher is concentration of nHA in the composite the higher is the osteopromotive and osteogenic potential of the composite material [38,49–50].
Thus by varying the content and the dimension of crystals of nHA it is possible to obtain scaffolds for BTE with tuneable mechanical and osteoinductive/osteogenic properties.
1.3.4 Scaffolds as drug delivery systems for BTE
As previously discussed, the combination of CP with growth factors (GF) or drugs can imbue an osteoinductive potential to scaffolds for bone repair applications. Not only can it be applied to CP, but several strategies have been developed for delivery of osteoinductive or –promotive compounds and proteins into a bony defect site.
28 Commonly GF, especially bone morphogenetic proteins (BMPs), are the most used osteoinductive substances in BTE scaffolds. GF play an important role in the bone repair process. The rationale for therapeutic use of GF in the stimulation of bone healing is based on the hypothesis that through appropriate signalling they may induce or accelerate the entire healing process. GF may be administered either via a protein therapy, resulting in direct GF delivery to the regeneration site, or via gene therapy, delivering GF into cells through their encoding genes [51].
Although widely used, these molecules show some limitations in their clinical use. BMPs are characterized by a rapid degradation losing their bioactivity in time. They are very sensitive to temperature so the sterilization process of scaffolds increases the rate of degradation [52]. Moreover their retention time at the defect site is minimal and this arouses a rapid diffusion of BMPs from the target tissue, resulting in insufficient bone regeneration [26,53]. To provide a longer-term effect at the target site, a stronger anchoring of the growth factor to an implantable device surface is a strategy that is commonly explored [53]. Additionally it has been shown that GF in general could have a role in tumour promotion when applied at a site of tumours bone removal. Furthermore, it is still unknown which possible genetic alterations in humans can occur following the application of GF. For these reasons their application during pregnancy and childhood must be carefully considered [51]. Finally, considering the high cost and the huge amount necessary, a substitute material for GF should be found.
In this regard, icariin could be a valid alternative to GF. Although it does not directly induce bone growth, it is an osteopromotive compound capable of enhancing osteoinductivity. Icariin (Figure 1.7) is the principal component of the seven flavonoid glycosides in Herba Epimedii [63], an important medicinal plant whose beneficial properties are well known in the Chinese traditional medicine. Icariin is pharmacologically active and shows therapeutic capacities such as osteoprotection, neuroprotection, cardiovascular protection, anti-cancer function, anti-inflammatory function, immunoprotective effect and enhancement of the reproductive function [25].
29 Figure 1. 7 Chemical structure of icariin
Already used in China as anti-osteoporosis drug, icariin can play an active role in the bone remodelling process when applied. It reduces bone resorption inhibiting the proliferation and differentiation of osteoclasts and it promotes the bone formation accordingly [25]. Unlike GF, icariin is a stable compound. It maintains its bioactivity perfectly after sterilization and its delivery is constant over 7 days [52,54]. This properties make icariin a promising drug candidate in BTE
Previous studies showed that icariin alone enhances in vitro osteogenic differentiation inducing bone mineralization. While in vivo, icariin, contained in calcium phosphate cement tablets, accelerates bone regeneration and angiogenesis [26]. Icariin, loaded in the biodegradable poly(3-hydroxybutyrate-co-3-hydroxyvalerate), stimulates in vitro the proliferation of human osteoblasts and preosteoblasts [52]. In other work it has been shown that the introduction of icariin in chitosan/nHA scaffolds only negligibly affect the properties of the scaffold. Particularly the presence of icariin causes only a negligible decrease of the compressive modulus of the scaffolds. Furthermore icariin does not affect the morphology and the porosity of the scaffold [54-55].
Thus, considering its osteopromotive capability, chemical stability, the lack of carcinogenic effects and the low cost, icariin can be considered a valid alternative to GF.
1.4 Fabrication techniques for BTE scaffolds
For the fabrication of scaffolds for BTE, several techniques are available. In this respect, the fabrication technique must not adversely affect the materials properties, specifically their biocompatibility or chemical properties. The technique should be
30 accurate and consistent, regarding the porosity, pore size, pore distribution and interconnectivity in the fabricated scaffolds. Different scaffold batches should exhibit minimal variations in their properties when processed from the same set of processing parameters and conditions [4]. Numerous techniques have been developed and employed to fabricate 3D scaffolds for BTE applications. These can be divided into two principal categories: conventional fabrication techniques and solid freeform fabrication techniques (also called ‘rapid prototyping’). Each of these techniques produces different features of internal architecture of the final scaffold, such as pore size, pore structure and interconnectivity, as well as mechanical properties.
This paragraph provides an overview on the conventional approaches and the rapid prototyping techniques that are used for producing BTE scaffolds. For each technique it will focus on the operating principles and relevant advantages and limitations.
Particularly it will be focused on the low temperature deposition modelling process since it will be the technique used to produce scaffolds for this project.
1.4.1 Conventional fabrication techniques
- SOLVENT CASTING
Solvent casting of composites involves dissolution of the polymer-ceramic particle mixture in an organic solvent, and casting the resulting mixture into a predefined 3D mould. The solvent subsequently evaporates, leaving a scaffold behind. The advantage of this method is that the preparation process is easy and does not require expensive equipment. However, there are two major disadvantages. First, with this technique it is only possible to form scaffolds with simple shapes (e.g. flat sheets, tubes, or cubes). Second, the residual solvents left in the scaffold material could denature proteins, and thus be harmful to cells and biological tissues [5].
- SOLVENT CASTING/PARTICULATE LEACHING
This approach involves casting a mixture of polymer solution and porogen particles such as sieved salt or sugar particles, and inorganic granules to fabricate porous membranes or 3D networks (Figure 1.8). The size of porogen particles and the ratio of polymer to porogen directly control the internal pore size and porosity of the final scaffold respectively. After solvent evaporation, the dried scaffolds are
31 fractionated in water or a suitable solvent to remove the particulates. Once the porogen particles have been completely leached out of the mixture, a porous structure is obtained. This method has both advantages and disadvantages similar to the solvent casting technique [5]. Additionally, the pore size distribution can be broad and pores are generally not completely interconnected.
Figure 1. 8 Schematic presentation of solvent casting/particulate leaching technique [28]
- FREEZE DRYING
As the previous method, the freeze drying technique also requires the use of organic solvents or water to produce a porous scaffold but does not require the use of porogen particles. First, a synthetic polymer is dissolved into the solvent. Subsequently, the solution is poured into moulds of specified dimensions and is frozen (e.g. with liquid nitrogen). The frozen polymer solution is lyophilised to produce porous scaffolds of highly interconnected pores with porosities being up to 90 %. One of the great benefits of this technique is the ability to fabricate a scaffold without the use of a high temperature. Further, the pore size and the morphology of the scaffolds depend on specific processing parameters, including the freezing rate, temperature and polymer concentration. However, the spongy scaffolds produced by this technique exhibit a porous structure of irregular and small pore size, typically ranging from 15 to 35 μm [5].
32 This approach involves the use of a volatile organic solvent of a low melting point to dissolve the polymer mixed with or without ceramic particles. To induce phase separation, the polymer solution is first cooled rapidly. This leads to the solidification of solvent, which forces the polymer solute into the interstitial spaces. Subsequently, a porous scaffold is obtained after the evaporation of solvent via sublimation. The principal advantage of this technique is that it is possible to obtain a variety of scaffold architectures by changing the types of polymer and solvent, polymer concentration and phase separation temperature. With this approach it is possible to produce scaffolds of 95% porosity but with typical pore size of lower than 200 µm which limits its utility in BTE [5].
- GAS FOAMING PROCESS
This technique employs a gas as a porogen to create interconnected pores in order to eliminate the use of organic solvents, the residual of which might result in an inflammatory response after implantation. First, a polymer is placed in a chamber and then saturated with high-pressure CO2, a non-toxic and non-flammable gas. As the pressure is rapidly dropped, the nucleation and formation of pores occur as a result of the thermodynamic instability in the gas/polymer system. The fabrication parameters such as temperature, pressure, degree of saturate and depressurisation time have a great influence on the pore morphology and pore size of the scaffolds. The negative aspects of this process include the use of the excessive heat during compression moulding; closed, non-interconnected pore structures, and a nonporous skin layer at the surface of the final product [5].
- ELECTROSPINNING
This technique allows the fabrication of non-woven scaffolds with a high porosity from a polymer solution. In the electrospinning process, a high voltage electric field is generated between a polymer solution held by its surface tension at the end of a syringe (or a capillary tube) and a collection target (Figure 1.9). Charge is induced on the solution surface by the electric field. As the intensity of the electric field increases, the hemispherical surface of the solution at the tip of the capillary tube elongates to form a conical shape known as Taylor cone. When the electric field reaches a critical value at which the repulsive surface charges overcome the surface tension force, a charged jet of the solution is ejected from the tip of the cone. As the jet is propelled
33 through the air, its diameter decreases as a result of stretching and solvent evaporation. Decreasing the jet diameter, the surface charge density increases and the resulting high repulsive forces split the jet into several smaller jets. This process may be carried out several times,leading to many small jets which are finally accumulated on the surface of the collector, resulting in nonwoven random oriented fibers with diameters on the nanometer scale. By changing the values of the variables, including solution viscosity, polymer charge density, polymer molecular weight and electric field strength, it’s possible to control the fibre diameter and morphology. With this technique highly porous scaffolds with interconnected pores can be produced. The disadvantage of this technique is that several polymers can only be processed by the use of organic solvents, which could be toxic to cells if not completely removed [5,56].
Figure 1. 9 Schematic overview of an electrospinning setup [56]
- POWDER-FORMING PROCESSES
The powder-forming process allows for fabrication of porous ceramic and glass scaffolds. In this process, a suspension of ceramic particles in a liquid (such as water or ethanol) called slurry is used to prepare green bodies. Different materials can be used as porogen such as sucrose, gelatine, poly(methyl methacrylate) microbeads and a wetting agent (like surfactant). They leave behind a porosity when they are evaporated or burned out during sintering. In addition, the presence of binders such as polysaccharides poly(vinyl alcohol) (PVA) and poly(vinyl butyl) in slurries plays an important role in improving the strength of the green body before the product is sintered. Using this technique scaffolds with a porosity of 90 % and pore size ranging from 510 and 720 μm have been produced [5].
34 - SOL–GEL TECHNIQUES
This is a versatile process, involving formation of a sol by the addition of a surfactant, followed by condensation and gelation reactions. Using the sol–gel process, it is possible to fabricate ceramic or glass materials in a variety of forms, including ultra-fine or spherical-shaped powders, thin-film coatings, ceramic fibres, microporous inorganic membranes, monolithic ceramics and glasses, and highly porous aerogel materials. Despite its advantages, the sol–gel technique does not produce porous ceramics of high mechanical strength. Even including particles inside like bioglass, it’s not possible to achieve the same mechanical properties of bone [5].
Although by conventional fabrication techniques it is possible to produce various types of scaffolds, most of them are incapable of producing structures with a fully continuous interconnectivity and uniform pore morphology within a scaffold. Additionally, the pore size, pore geometry and spatial distribution cannot be precisely controlled in these conventional processes. Some conventional techniques indeed are performed manually, with poor reproducibility. Another limitation of most conventional fabrication methods is the need of an organic solvent to dissolve polymers and other chemicals, as well as the use of porogens to create pore structures. Most solvents and porogens are toxic, and their residues in the scaffold may cause severe inflammatory responses [5].
1.4.2 Rapid prototyping techniques
All the techniques that will be described in the following paragraphs are based on a computer-aided design (CAD) to fabricate custom-made devices, like scaffolds, directly from computer data. The designed geometry of the final structure that is to be built can be either devised using 3D drawing computer software, described through mathematical equations, or derived from scanning data of clinical imaging technologies such as magnetic resonance imaging (MRI) or tomography techniques. The possibility of using clinical imaging data from a patient, make these manufacturing technologies particularly useful for many applications in biomedical engineering, since it’s possible to fabricate patient-specific implants[40]. This increases the performance and the efficacy of the structures in vivo (reference).
35 (i) Development of a CAD-file: The CAD-file describes the geometry and the size of the structure to be built. The information needed for the device are contained in standard tessellation language (STL) files. They contain a list of coordinates of points that together describe the surface of the designed 3D structure.
(ii) Slicing of the structure: The designed structure is virtually sliced into layers of a thickness equal to the printing resolution (printed layer thickness) which depends on the resolution of the used technique.
(iii) Fabrication of the structure: These data are finally sent to the apparatus and the structure is then fabricated layer by layer.
With these approaches it is possible to obtain complex scaffolds whose architecture is manufactured layer by-layer building via the processing of solid sheet, liquid or powder materials stocks according to its computerised cross-sectional 3D image. Solid freeform fabrication techniques have significant advantages overconventional techniques in terms of consistency, reproducibility of designed scaffolds and the precise control over the architecture of 3D scaffolds such as internal structure, geometry, pore sizes and spatial distribution. In this way by using these scaffolds both biological and mechanical performances of tissue-engineered constructs can be improved [5]
- STEREOLITHOGRAPHY
Stereolithography (SLA) is based on the spatially controlled solidification of a liquid resin by polymerisation. An SLA apparatus consists of a tank of photo-sensitive liquid resin, a moveable build platform (in the directions x, y, z), a computer controlled ultraviolet (UV) laser or a digital light projector to irradiate the resin and a dynamic mirror system (Figure 1.10).
36 Figure 1. 10 Schematic representation of a stereolithography apparatus [5]
The photo-sensitive liquid resin in the stereolithograph is exposed to the UV laser which photo-polymerizes or -crosslinks the resin layer-by-layer into a 3D scaffold. Once one layer is completely solidified onto the build platform, it is vertically lowered with a small distance into the resin-filled vat. Subsequently, liquid resin covering the previous layer, will form the next layer after UV laser exposure. These steps are repeated until a complete 3D part is formed. Finally, uncured resin is washed off and the scaffold can be post-cured under UV light, yielding a fully cured part.
According to the build orientation and the method of illumination of the liquid resin, it’s possible to distinguish two SLA setups called respectively bottom-up and top-down setup as shown in Figure 1.11. The system depicted in Figure 1.10 is an example of a bottom-up setup. In both systems objects are built in a layer-by-layer manner by spatially controlled photo-polymerization of the photo-sensitive resin.[58]
37 Figure 1. 11 Schemes of two types of SLA setups. Left: a bottom-up system with scanning laser. Right: a top-down setup with digital light projection [58].
Even though the structures obtained with top-down setup are subjected to larger mechanical forces, as they have to be separated from the bottom plate of the resin bath after illumination of each layer, this approach has several advantages over the bottom-up systems. For instance the surface being illuminated is always smooth, only small amounts of resin are required, and the illuminated layer is not exposed to the atmosphere, so oxygen inhibition of the photo-polymerization is limited [58].
SLA is widely used for BTE scaffolds. It has been applied to biodegradable polymers such as poly(propylene fumarate), photo-crosslinkable poly(ε-captolactone) (PCL) and poly-DL-lactide (PDLLA). The processing of bioceramics has also been reported for SLA (as suspension of hydroxyapatite and a low viscosity acrylate resin) and polymer/ceramic composites such as poly(trimethylene carbonate) and nano-hydroxyapatite [5,40].
SLA is a versatile process that allows for the design of complex structures with different geometries, different sizes (from submicron to decimetre) and a high surface finish. Compared with other solid freeform fabrication techniques, SLA is a fast technique which shows excellent reproducibility, producing nearly identical built architectures with high resolution (of about 50 μm) and accuracy. Despite the many advantages, the use of photo-sensitive resins is primarily considered a limitation of this process. It can cause problems of toxicity as skin irritation and cytotoxicity. Another disadvantage of this process is associated with the shrinkage of the formed polymer networks during polymerisation[5].
