1 Introduction
This chapter provides an introductory review of the techniques available for the production of scaffolds for Tissue Engineering (TE). Special attention will be addressed to the various techniques and to the biomaterials studied within this research area. Conventional and Additive Manufacturing (AM) techniques used for the production of scaffolds for TE will be described highlighting the advantages and limitations of each technique.
1.1 Tissue Engineering (TE)
With the increasing of aging of the worldwide population many healthcare problems arise. The evolution of medicine allows us to live longer than before although some cases of chronic diseases and traumatic events are still difficult to treat with traditional medical techniques. The use of transplanted organs and tissues is limited due to the lack of donors and to various risks involved in the transplantation process [Shin et al., 2003]. Autologous tissue transplantation (tissue from the patient) is the most reliable and safe way to treat patients. However, in most of the cases the site morbidity, limit availability of tissue and other complications related to the complexity of the surgical procedures limits this approach [Petite et al., 2000]. The use of tissues or organs from allogenic sources may carry severe drawbacks, such as rejection or transfer of disease [Damien and Parsons, 1991; Laurencin et al., 1999]. Moreover, the necessity for patients to take immunosuppressive drugs after the transplant is also a concern.
TE presents alternative approaches to address these problems. The most popular and accepted definition among the scientific community of TE is “an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function or a whole organ” [Langer and Vacanti, 1993]. Over the past two decades various approaches that aim to improve healing and restore functions of both tissues and organs have been developed. Three principal therapeutic strategies for treating diseased or injured tissues in patients are: (i) implantation of freshly isolated or cultured cells, (ii) implantation of tissues assembled in vitro from cells and scaffolds, and (iii) in situ tissue regeneration [Griffith and Naughton, 2002]. These approaches rely on three fundamental pillars:
cell, scaffolds and growth factors [Ikada, 2006; O'Brien, 2011]. The application of environmental factors (cytokines, growth factors, genetic manipulation, mechanical forces, physiochemical
factors, spatial and temporal signals, extracellular matrix (ECM) molecules, cell surface molecules) on an in vitro stage prior to the construct implantation can also enhance the engineered tissue [Khademhosseini et al., 2006].
Cells from different sources are investigated in the TE field. Adult or embryonic stem cells, capable of both self-renewal and differentiation into a variety of cell lineages, or a mixture of differentiated cells at different stages of maturation and from sources like autologous, allogenic (cells from a human donor) or xenogenic (cells from different species) are considered [Griffith and Naughton, 2002].
1.1.1 Scaffold-based TE approach
The scaffold-based approach is one of the most studied in TE field. This approach relies on the utilization of scaffolds to support cells proliferation and differentiation until a new tissue formation is achieved [Mano et al., 2007]. Generally, an in vitro pre-culture step is performed under controlled conditions. In some cases a bioreactor system is used to promote biophysical stimuli with the objective of applying mechanical, chemical or the combination of both solicitations to the constructs in order to enhance the tissue formation [Ellis et al., 2005]. After an initial period of time of in vitro culture, the living construct is implanted into the patient affected site where it is expected that the formation of new tissue is fully achieved and the scaffold fully degraded [David, 2004]. Other alternative approach uses the direct implantation of the scaffold on the affected site allowing cells from the surrounding tissue to migrate and produce the new tissue in situ [Ikada, 2006].
The success of the scaffold-based TE relies on some fundamental scaffold prerequisites [Moroni et al., 2008a; O'Brien, 2011]:
• Biocompatibility – Scaffold material(s) must be biocompatible and the degradation products should not induce allergic or inflammatory reactions. Cells should be able to adhere, proliferate, migrate and differentiate when in vitro cell culture step is performed.
Moreover, when acellular scaffold approach is used, cells from the surrounding tissue must be able to migrate towards or into the scaffold. When implanted, the tissue engineered constructs or scaffolds should not induce a severe host immune response, although a mild reaction is generally unavoidable.
• Biodegradability- [Vert et al., 1992; Woodruff and Hutmacher, 2010] - Scaffolds are temporary structures that are used to support cells while they build the natural ECM. The
degradation rate of the scaffold should be tuned such that sufficient structural integrity of the scaffold is retained until the newly grown tissue is able to support loads and stresses.
The degradation by-products should be non-toxic and should be able to be secreted by the diverse metabolic pathways from the body. The most frequent degradation mechanism observed in polymers is non-enzymatic hydrolysis, although in some cases enzymatic hydrolysis occurs.
• Architecture - Scaffold should possess a high and interconnected porosity as well as a highly porous surface and microstructure, allowing in vitro cell adhesion and in-growth and providing the necessary space for in vivo neo-vascularisation. In addition, pore interconnectivity affects the diffusion of nutrients and the removal of cells by-products [Puppi et al., 2010a]. Adequate scaffold porosity, pore size, pore distribution and macro, micro and nano size features should therefore be designed according to the tissue three- dimensional (3D) microenviroment and scaffold mechanical properties [Yeong et al., 2004].
• Mechanical properties – A TE scaffold should have sufficient mechanical strength and elasticity during in vitro culture to maintain the spaces required for cell in-growth and matrix formation, should retain structural stability and integrity allowing its proper handling by physicians during the implantation phase, and should match as closely as possible the mechanical properties of the host tissue to bear in vivo stresses and loading [Kohane and Langer, 2008; Puppi et al., 2010a].
• Surface properties – Scaffold surface morphology, hydrophilicity, energy and charge are factors that determine in vitro cell adhesion, migration, phenotype maintenance and intracellular signalling as well as in vivo cell recruitment and healing at the tissue-scaffold interface. Polymer surface engineering is a useful tool to improve scaffold multi- functionality and to design biomimetic materials able to interact with the surrounding environment by biomolecular recognition. Moreover, surface modification is critical to elicit specific cellular responses and direct new tissue formation [Puppi et al., 2010a].
• Anatomical shape - Scaffold should have an external geometry and size, matching those of tissue defect in order to achieve a good integration and distribution of mechanical loadings [Leong et al., 2008].
• Sterilization - An adequate sterilization technique should be selected according to the scaffold material in order to avoid physicochemical changes.
1.1.2 Biomaterials for scaffolds fabrication
The definition of biomaterial followed a long and complex path [Williams, 2009] until an acceptable version was published in the Dictionary of Biomaterials Science in 1999 [Williams, 1999] as by following: “A Biomaterial is a material intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ or function of the body”. Many biomaterials, such as synthetic ceramics (e.g hydroxyapatite (HA), tricalcium phosphate (TCP), silica-based calcium phosphate), natural minerals (e.g. coral ), metals (e.g. titanium and its alloys, stainless steel), synthetic polymers [e.g. poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly (ε- caprolactone) (PCL)], proteins (e.g. collagen, fibrin) and natural polysaccharide (e.g. alginate, agarose, chitosan and hyaluronic acid), are already clinically available in different products [Barrère et al., 2008]. Several are the natural and synthetic biomaterials investigated for the manufacturing of scaffolds for TE application [Mano et al., 2007; Nair and Laurencin, 2007; Place et al., 2009; Rezwan et al., 2006].
The most common polymers used to produce TE scaffolds are divided in two groups: the natural polymers and the synthetic polymers. Another group of materials generally used for TE applications are bioactive ceramics.
1.1.2.1 Natural polymers
Natural-based polymers are very attractive for TE applications due to their low toxicity, renewability, biological signalling and similarity with the natural ECM [Mano et al., 2007]. Despite these advantages, natural-based polymers can, in some cases and depending on the source, have impurities and endotoxins that can evoke undesirable immune responses following implantation [Mano et al., 2007]. Accurate purification procedures must be implemented in this case. Another limitation of natural polymers is the batch-to-batch variability that can limit the reproducibility of scaffold composition.
Polymers synthesized in living organisms are generally divided into: polysaccharides (starch, cellulose, chitosan, alginate, and hyaluronic acid), proteins (collagen, silk, fibrin, gelatin) and biopolyesters [Mano et al., 2007].
A wide range of polysaccharides, such as chitosan and alginate, have been proposed for the production of scaffolds for TE. Chitosan is the deacetylated form of chitin natural polymer with gel-forming capability, biodegradability and high adsorption capacity [Dash et al., 2011].
Moreover, it is biocompatible and non-toxic as well as endowed with antibacterial, antifungal and antitumor activity [Dash et al., 2011]. Various porous structures can be obtained with chitosan
making it suitable for cell ingrowth and osteoconduction [di Martino et al., 2005]. Applications in the TE field include bone, cartilage, intervertebral disc, liver and nerve [Dash et al., 2011; di Martino et al., 2005].
Collagen is the most widely used naturally occurring protein that can be found in the connective tissue of mammals such as skin, bone, cartilage, tendon, and ligament [Lee and Mooney, 2001]. A great amount of literature has investigated collagen for TE applications due to its cell binding properties. However, collagen presents some drawbacks such as, being expensive and fast degraded in vivo [Lee and Mooney, 2001]. Gelatin is a protein derived from denaturation of collagen and has tuneable degradation rates (depending on the degree of crosslinking) in vitro and it is commonly used for the controlled release of drug or growth factors and for both hard and soft TE scaffolds, either in the form of hydrogel or in combination with other polymers [Young et al., 2005].
Polyhydroxyalkanoates (PHA) are biodegradable and biocompatible aliphatic biopolyesters synthesized by many types of bacteria as carbon and energy reserve materials induced by unbalanced growth conditions [Chen, 2011; Puppi et al., 2010a]. PHA’s have been used for several biomedical applications including sutures, cardiovascular patches, orthopaedic pins, stents, and as scaffolds for cartilage, nerve, tendons and wound dressings [Chen and Wu, 2005]. The most extensively studied PHA is poly(3-hydroxybutyrate) (PHB) [Freier, 2006]. PHB was first isolated and characterized by Lemoigne in 1923 [Philip et al., 2007] and it is produced via biosynthesis of renewable carbon sources by bacterial fermentation. It has high crystallinity, it is extremely brittle and stiff, and relatively hydrophobic with slow hydrolysis in vitro and in vivo [Freier, 2006]. PHB has a very narrow processability window, and a low impact resistance that limits its use in load- bearing TE applications. Several copolymers and blends were tested to overcome these limitations [Chen et al., 2000]. Poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV) is an example of a copolymer of PHB containing segments of hydroxyvalerate (HV). PHBV is less crystalline than PHB and thus more flexible [Avella et al., 2000]. Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx) has shown better elasticity and processability compared with other PHAs [Chen, 2011].
1.1.2.2 Synthetic polymers
Synthetic polymers are produced under controlled conditions that allow for the obtaining of polymers with high purity and reproducibility, controlled chemical composition and mechanical properties and with tuneable degradation rates [Rezwan et al., 2006; Singh and Elisseeff, 2010].
Generally, the risk of immunogenicity, toxicity and infections is low when these types of materials are used [Rezwan et al., 2006]. Synthetic polyesters, including PCL, PLA, PGA and poly(lactic-co- glycolic acid) (PLGA) are extensively used for the production of scaffolds for TE applications.
PLA, PGA and their copolymers are degraded hydrolytically and degradation elements (lactic and glycolic acid) are normally identified in the metabolic pathways of the human body [Barrère et al., 2008]. The common use of these polymers may be attributed to the long history of products approved by the Food and Drug Administration (FDA) agency for diverse clinical applications, making these materials very attractive for TE applications despite the concerns about the acidic by- products resulting from degradation in vivo [Shoichet, 2010]. PLGA copolymers can be synthesized by direct polycondensation of lactic acid and glycolic acid or by ring-opening polymerization of lactide and glycolide [Felt et al., 2001; Konstantinos, 2008]. As biocompatible and biodegradable polymers with adjustable properties, such as degradation time and mechanical properties, they are generally used in biomedical applications such as sutures, orthopaedic fixation devices, drug delivery systems and scaffolds for TE applications [Konstantinos, 2008].
PCL is a hydrophobic, semi-crystalline polymer prepared by ring-opening polymerization of a cyclic monomer (ε-caprolactone) and its synthesis was firstly studied in 1934 [Natta et al., 1934].
Its good solubility, low melting point (59ºC – 64ºC), low glass transition temperature (-60ºC) and its better rheological and viscoelastic properties when compared to other polyesters, make PCL very attractive for processing with various techniques. However, the long periods of degradation (up to 3-4 years) impairs its application where a fast degradation is needed. Indeed, the initial applications of PCL included long-term drug delivery devices and slowly degradable sutures [Woodruff and Hutmacher, 2010]. PCL started to be investigated largely in the TE field due to the optimal processability with the majority of techniques used for scaffold production [Woodruff and Hutmacher, 2010]. Furthermore, since a significant number of PCL drug delivery devices has already FDA and CE mark registration, this might help when new products (e.g. a scaffold) are intended to be commercially available [Woodruff and Hutmacher, 2010]. Multiarm star-shaped poly(ε-caprolactone) (*PCL) polymers are composed of a number of linear homopolymer arms covalently connected to a central core and were recently proposed for TE applications [Puppi et al., 2010b]. Their small size, spherical structure and limited interaction between molecules confer different properties when compared to their linear equivalents with equal molecular weight [Puppi et al., 2010b], such as reduced viscosity, making these polymers promising for processing using techniques such as melt-electrospinning (melt-ES) and extrusion-based AM techniques.
1.1.2.3 Bioactive ceramics
Bioactive ceramics are solid inorganic materials that have a crystalline, partially crystalline or non- crystalline structure [Huang and Best, 2007]. Generally, these materials are clinically applied in
skeletal applications such as bone, teeth and joints, and they can be either resorbable or non- resorbable. The bioactive ceramics commonly studied for TE applications are calcium phosphates, bioactive glasses and glass-ceramics, due to their capability of promoting the formation of new bone on their surface. These materials are characterized by a high brittleness and hardness [Gerhardt and Boccaccini, 2010].
Calcium phosphates
Bone minerals are largely composed of calcium phosphates. Some of the bioactive ceramics studied in the TE field are calcium phosphates (HA, β-TCP and α -TCP). The most extensively, clinically used calcium phosphate is HA because of its chemical similarities to natural bone and to its biocompatibility, resorbability and osteoconductivity [Sopyan et al., 2007]. Moreover, HA scaffolds can be produced using a variety of techniques, such as conversion of natural bones, ceramic foaming technique, polymeric sponge method, gel casting of foams, microwave processing, slip casting, and electrophoretic deposition technique [Sopyan et al., 2007]. TCP has a chemical composition of Ca3(PO4)2 and has four polymorphs, and the most commonly studied types are α-TCP and β-TCP [Huang and Best, 2007].
Bioactive glasses
Bioactive glasses were first investigated for biomedical applications by Hench et al in the seventies [Hench et al., 1971]. They are silicates containing Na2O, CaO, SiO2 and P2O5 as main components [Hench, 1998]. These ceramics in contact with biological fluid or simulated body fluid (SBF) produce a bioactive hydroxycarbonated apatite layer that can bind to existing biological tissue [Kokubo and Takadama, 2007; Peitl et al., 2001].
1.1.2.4 Composite materials
The combination of natural and synthetic polymers and the combination of polymers and bioactive ceramics is commonly adopted for the fabrication of scaffolds for TE. The inclusion of various bioactive ceramics into polymer matrix allows the enhancement of mechanical properties, osteoconductivity and interaction of the scaffold with the surrounding tissue [Lam et al., 2008].
Moreover, bioactive ceramics can act as internal pH buffer inhibiting autocatalytic degradation of PCL, PLA, PLGA or their copolymers due to the acidic nature of their degradation by-products [Rezwan et al., 2006].
1.2 Techniques for Scaffolds Fabrication
Several techniques are used for the production of scaffolds for TE, and can be classified in two groups: conventional and AM techniques. All these techniques have some limitations and advantages over each other. It is impossible to select the optimal technique for the production of all types of scaffolds typology and therefore any selection should be performed according to the materials to be processed and to the scaffold-specific properties needed for the particular tissue application. Many research groups and companies are developing new technologies capable of overcoming some limitations by, for instance, combining different processing principles in the same apparatus.
Most of conventional techniques were developed in recent decades for the production of scaffolds with diverse morphology and composition. Techniques such as solvent-casting/particulate- leaching, gas foaming, fibre bonding, phase separation, melt moulding, freeze drying and many others have been used to produce scaffolds [Liu et al., 2007; Sachlos and Czernuszka, 2003;
Thomson et al., 2000]. However, these techniques generally present several drawbacks including the use of toxic organic solvents, low control over pore size and distribution, and low pore interconnectivity [Leong et al., 2003]. Moreover, these processes are generally complex and involve long periods for the fabrication of scaffold [Hutmacher, 2000]. In addition, the production of a 3D scaffolds is limited and the control of the 3D external shape and internal structure is poor.
AM techniques, also called rapid prototyping or solid freeform technologies, are gaining interest over conventional techniques thanks to their ability to customize the internal and external scaffold’s architecture. Some AM techniques, originally developed for the aerospace and automotive industries, have been explored for the production of TE scaffolds due to the feasibility of producing structures of high complexity and with high reproducibility [Hutmacher, 2000]. At present, research is focusing on the adaptation or development of new systems in order to respond to the demanding exigencies of scaffold production process. Conventional and AM techniques are described in detail in the following sub-chapters.
1.2.1 Solution electrospinning (ES)
The production of polymeric nanofibres can be achieved by diverse techniques, such as drawing, self-assembly, phase separation, template synthesis and electrospinning (ES) [Dahlin et al., 2011;
Detta et al., 2010c; Ondarçuhu and Joachim, 1998]. ES gained enormous attention due to the simplicity of the process. Unlike with other spinning techniques such as melt-, wet-, gel- and dry- spinning, this technique uses an electrical field to draw the fibre until its collection occurs [Frenot and Chronakis, 2003]. This process results in the production of continuous fibres with a diameter in the range from nano to tens of microns [Huang et al., 2003]. ES is a process that has been intensively investigated in the TE field because of the structural resemblance between ECM fibres present in several tissues and electrospun non-woven nanofibres meshes [Agarwal et al., 2008;
Liang et al., 2007; Murugan and Ramakrishna, 2006; Pham et al., 2006a]. Electrospun nanofibres have been used for many TE applications such as the engineering of skin, bone, cartilage, tendon, ligaments, cardiovascular and neural tissue [Dahlin et al., 2011; Fang et al., 2008]. Fibres of sub- micron and nano dimensions have a high surface area-to-volume ratio: a desirable feature for optimizing cellular adhesion [Dahlin et al., 2011]. A wide range of natural and synthetic biodegradable polymers and their blends have already been reported to be processed using ES techniques [Huang et al., 2003; Liang et al., 2007]. Moreover, it is also possible to include cells, growth factors and inorganic materials in the process [Chronakis, 2005; Huang et al., 2003; Sill and von Recum, 2008].
Solution ES is subject to some limitations, for example, the use of organic solvents, the low control over pore size and pore distribution and the high packing density of the electrospun meshes, limiting cell infiltration and migration through the electrospun mesh. Some approaches are currently being developed to overcome this limitation. The production of electrospun meshes containing a leachable polymer or salt, the use of ice crystals, or the combination with other techniques capable of producing fibres with larger fibre diameter [i.e. melt-electrospinning (melt- ES), wet-spinning and AM techniques] are some of the solutions proposed [Baker et al., 2008;
Chen et al., 2009; Kim et al., 2008; Martins et al., 2009; Park et al., 2008; Santos et al., 2008;
Tuzlakoglu et al., 2005; Yoon et al., 2009; Zhong et al., 2011].
The research community and industry focused on the solution-based ES rather than on melt-ES due to higher capital investment requirements and the difficulty in producing submicron fibres by melt- ES [33] (see sub-chapter 1.2.2).
The typical laboratory ES apparatus (Figure 1) requires a high voltage power supply, a syringe pump, a syringe, a stainless steel needle with a blunt tip and an electrical conductive collector [Teo and Ramakrishna, 2006].
Figure 1 - Typical electrospinning setup (basic laboratory configuration).
The complexity of the setup depends on the targeted mesh or fibre morphology. The power supplies that are used in an ES setup can use direct current (DC) or alternating current (AC), even if DC is more commonly used than AC. Generally, a single conductive spinneret is used, although multi-nozzle systems were also proposed in order to increase the production efficiency or to obtain multi-material meshes [Yamashita, 2008]. Nozzle-free systems allowing fibres collection onto a larger area and avoiding problems like needle clogging were also developed [Yarin and Zussman, 2004]. By acting on the material, size, geometry or movement of ES collector it is possible to vary the morphology of the produced fibres and the size and geometry of the resulting mesh [Detta et al., 2010b; Teo et al., 2011]. For instance, the use of rotating drums allows the mechanical winding of the collected nanofibre and so the achievement of fibre alignment and of tubular electrospun structures suitable for blood vessel or nerve TE [Carnell et al., 2008]. Ideally, the tangential velocity of the collecting drum should match that of the electrospun jet when it reaches the drum, in order to achieve fibre alignment. As shown by studies investigating the relationship between the rotation velocity of the collecting drum and the degree of alignment of the produced nanofibres [Edwards et al., 2010], the rotation speed of the drum should be very high (thousands of r.p.m.
depending on drum diameter).
Some other changes to the traditional ES setup have also been investigated, namely the application of auxiliary electrode systems allowing the manipulation of the electrical field [Teo and Ramakrishna, 2006].
The ES process is influenced by many parameters affecting fibre and mesh morphology and these are generally divided into three groups [Doshi and Reneker, 1995; Pham et al., 2006a]: a) solution
Power Supply
Syringe Pump
Syringe
Stainless steel needle Collector
parameters; b) processing parameters; and c) ambient parameters. Solution properties include parameters such as solution viscosity (associated with the polymer concentration and molecular weight), solution conductivity, solution surface tension and solution dielectric constant; the processing parameters that generally affect the ES process are solution feed rate, applied electrical voltage and distance between the nozzle and the collector. The ambient parameters mostly influence fibre morphology are temperature, humidity and air velocity [Doshi and Reneker, 1995].
The conventional ES set up allows for the collection of randomly oriented fibres. Several studies have proposed different solutions for the obtainment of a control over fibre orientation, such as the employment of rotating mandrel as collector or the employment of auxiliary electrodes to manipulate the electric field [Teo et al., 2011; Teo and Ramakrishna, 2006]. Studies performed within our group showed that it is possible to deposit in a controlled way a single ultrafine fibre by applying an uniform electric field, obtained by means of an electrical insulating spinneret [Chiellini et al., 2009; Errico et al., 2011]. Recently, Dalton et al [Dalton et al., 2008] developed a new ES technique called Direct Writing enabling the controlled deposition of a single electrospun fibre in order to develop scaffolds with controlled pore size and porosity in a layer-by-layer fashion. Until now this concept was only exploited using melt-ES technique, although, a high viscosity solution could also be used. Near field ES is another approach to collect fibres with a controlled orientation by employing a small distance between the needle tip and the collector. This allows to collect the electrospun fibre jet in the stable region and thus to avoid the bending instability region [Chang et al., 2008; Padmanabhan et al., 2011; Sun et al., 2006; Zheng et al., 2010].
1.2.1.1 ES for production of multi-scale scaffolds
Electrospun meshes are characterized by a fully interconnected network of pores that allow a good permeability to oxygen and nutrients necessary for cells colonization [Detta et al., 2010c].
However, the high packing density generally obtained in the electrospun meshes limits the proliferation and migration of cells in the inner part of the nanofibre membranes [Pham et al., 2006b; Sill and von Recum, 2008]. Several alternative solutions have been investigated in recent years. As previously discussed, Yang et al proposed electrospun meshes composed of layers of electrospun PCL/collagen fibres [Yang et al., 2009] intercalated layers of various types of cell for the production of skin substitutes. Other approaches include the combination of ES with diverse techniques such as AM (chapter 3), wet-spinning and others to produce multi-scale 3D scaffolds.
Kim et al [Kim et al., 2010] suggested the combination of melt-ES and solution ES to produce 3D micro/nano hybrid meshes of PLGA. The incorporation of the nanofibres into the microfibres web induced an increase in cell attachment and migration.
1.2.2 Melt-electrospinning (Melt-ES)
The first patent on melt-ES was granted in 1936 to Charles Norton [Norton, 1936], who proposed a new apparatus composed of a heating container for the melt of a viscous material, and a high voltage power supply for the generation of an electric field between an air-blasted jet of molten polymer and a collector (cylinder or flat surface). After that, no substantial work was published until 1991 when Larrondo and Manley presented some studies showing the feasibility of obtaining continuous fibres from polymer melts using an electric field [Larrondo and St. John Manley, 1981a, b, c]. In the last decade, some research groups become interested in this technology and started to further develop this technique. Different synthetic biodegradable polymers have already been processed using melt-ES to produce scaffolds for TE applications [Góra et al., 2011; Hutmacher and Dalton, 2011]. Conversely, natural polymers were never processed using this technique due to their low thermal stability. Fibres with a diameter of hundreds of microns down to 180nm [Zhmayev et al., 2010b] can be achieved.
Melt-ES is a technique that allows a high throughput production of fibres from polymer melt. This is related to the fact that the quantity of the polymer melted and submitted to the electrical field is the quantity of melt-electrospun mesh obtained (no solvent is involved in the process) [Detta et al., 2010a; Góra et al., 2011; Hutmacher and Dalton, 2011].
However, the production of meshes composed of ultra-fine fibres using melt-ES is in some cases limited due to the low whipping or bending instabilities of the melt jet because of the high viscosity of the polymer melt [Dalton et al., 2007; Karchin et al., 2011; Lyons et al., 2004] and the fast solidification of the fibre during stretching in the spinning region [Zhou et al., 2006]. Taking advantage of this drawback, Dalton et al [Dalton et al., 2008] proposed the collection of fibres with controlled patterns using translating collectors. The major goal of this approach is to collect fibres with high alignment and therefore to produce scaffolds with high control over pore size.
The morphology of fibres produced with melt-ES is influenced by many parameters such as: i) temperature at the nozzle; ii) temperature in the spinning area; iii) nozzle diameter; iv) nozzle-to- collector distance; and v) feed rate [Zhou et al., 2006]. In particular, fibre diameter is strongly affected by the processing temperatures (i.e. nozzle and spinning area temperatures). Some solutions are currently being investigated to reduce the fibre diameter obtained, thereby avoiding the rapid quenching or solidification of the polymer melt fibre. This phenomenon can be reduced by applying an air jet at a given temperature collinear with the direction of the formation of the melt-electrospun fibre [Zhmayev et al., 2010b]. Melt-ES principles are similar to those of solution ES, although some differences can be observed in the setup.
The melt-ES technique generally requires an auxiliary system that can control the polymer temperature during the entire process, making the setup more expensive when compared to solution ES. Temperature control systems allow the fusion of polymer by applying temperature above the melting temperature of the polymer. Some changes on the melt-ES technique, when compared to solution ES, can also be observed on the system applied to control the extrusion rate of the polymer. When complex extrusion and heating systems are used, the positive terminal of the high voltage power supply is generally connected to the collector and the extrusion nozzle is electrically grounded [Lyons et al., 2004]. This configuration is mainly used to protect the extruder from electrical discharges.
One of the biggest challenges in melt-ES is the correct and accurate control over the processing temperature. The temperature gradients along the melt-ES setup influence the morphology of the obtained fibre. In the literature, it is possible to find different temperature control systems such as water or oil circulating systems [Dalton et al., 2007; Dalton et al., 2008; Dalton et al., 2006b; Detta et al., 2010a], heating air gun [Dalton et al., 2007], electrical heating elements [Kadomae et al., 2009; Lyons et al., 2004], and lasers [Ogata et al., 2007a; Ogata et al., 2007b; Tian et al., 2009].
Moreover, the control of the spinning region temperature is also becoming considered in some studies [Zhmayev et al., 2010a].
The accurate control of extrusion rate of the polymer melt during melt-ES is a critical factor for producing fibres with good morphology and reproducibility. One of the developed strategies is to use a screw extruder system to control the flow rate of the polymer melt [Erisken et al., 2008a, b;
Lyons et al., 2004]. The development of a well-designed screw extruder system will allow a high throughput during melt-spun fibre production and more accurate control of the process. Other extrusion devices or solutions without extruder systems developed have limited control of the melt polymer feed into the electrical field [Deng et al., 2009; Ogata et al., 2007b; Rangkupan and Reneker, 2003]. Conversely, the employment of a syringe pump allows a better control of the extrusion rate [Dalton et al., 2006b; Detta et al., 2010a]. However, it requires an auxiliary water or oil bath to induce polymer melt and in this case the stabilization of the polymer melt temperature is relatively slow.
Since no organic solvents are used in the melt-ES technique, its application directly over fibroblast cells was proposed by Dalton et al [Dalton et al., 2006a]. The obtained results showed good cell viability despite the initial concerns regarding the high processing temperature. Melt-ES performed directly over cells requires proper manufacturing procedures to ensure the sterility along the process. Detta et al recently reported the development and assembly of a compact melt-ES system capable of being introduced inside a cell culture hood in order to perform all the necessary steps to produce living constructs with cells incorporated directly on the melt-electrospun meshes [Detta et
al., 2010a]. Different biodegradable polymers suitable to be processed in the melt state have been investigated for the production of melt-electrospun fibres. PCL and PCL-based copolymers [e.g.
poly(ethylene glycol)-b-PCL] are extensively studied for the production of melt-spun meshes for TE [Dalton et al., 2006a; Dalton et al., 2006b; Detta et al., 2010a]. PLA was also investigated although no application in TE was suggested [Ogata et al., 2007b]. PLGA was employed for the production of 3D scaffolds combining melt-ES and solution ES [Kim et al., 2010]. Moreover, some new polymers are being studied for the production of scaffolds using melt-ES technique. An example is a new biodegradable and thermostable polyurethane (PU) composed of PCL diol, 1,4- butane diisocyanate and 1,4-butanediol in a 1/4/3 molar ratio [Karchin et al., 2011]. With this polymer, it was possible to achieve nontoxic scaffolds composed of fibres with a diameter of 11.2 ± 2.3 µm and with mechanical properties suitable for soft tissue applications. A resume of the processed polymers, processing conditions and obtained fibre diameter using melt-ES for the production of scaffolds for TE is presented in Table 1.
Table 1 – Polymers, processing conditions and obtained fibre diameter using melt-ES techniques. Processing temperature [T]; Feed rate [FR]; Applied electrical voltage [Vapp]; Needle to collector [N-C]
Polymer Processing conditions Fibre diameter Ref.
PEG-b-PCL/PCL T=90ºC (water circulation) FR= 0.02 - 0.3 ml·h-1 Vapp= 20kV
N-C=10 cm
270±100 nm - 2.0±0.3 µm
[Dalton et al., 2007]
PEO-b-PCL T=85ºC (water circulation) FR= 0.03 - 0.3 ml·h-1 Vapp = 25 kV N-C=5 - 30 cm
1.5±0.3 - 1.7±0.3 µm [Dalton et al., 2006a]
PEG-b-PCL;
PEG-b- PCL/PCL;
T=90ºC (water circulation) FR= 0.005-0.05 ml·h-1 Vapp = 12kV
N-C=9 cm
2 µm [Detta et al., 2010a]
PCL T=90ºC (water circulation) FR= 0.005-0.02 ml·h-1 Vapp = 4-12kV N-C=2 – 6 cm
6 - 33 µm
Poly(L-lactic acid) (PLLA);
PLLA/ ethylene vinyl alcohol (EVOH)
T= n/a (CO2 laser) FR= n/a
Vapp = 25kV N-C= 5 cm
~ 3 µm (PLLA)
~ 1 µm (PLLA/EVOH)
[Tian et al., 2009]
PLA T= n/a (CO2 laser, 25W) FR= n/a
Vapp = 26 -41 kV N-C= 1 - 5 cm
~ 1 µm [Ogata et al., 2007b]
Table 1 – Cont.
Polymer Processing conditions Fibre diameter Ref.
PLGA T= 210ºC (oil circulation) FR= 1.4 to 5.4 ml·h-1 Vapp = 17.5 kV N-C= 8 cm
28 µm [Kim et al., 2010]
PU T= 220–240ºC (Electrical
heating elements) FR= n/a
Vapp = 30 kV N-C= 13 cm
11.2 ± 2.3 µm [Karchin et al., 2011]
1.2.3 Outlook over solution ES and melt-ES
The ES technique is widely studied for the production of scaffolds for a wide range of human tissues. However, some limitations of this technique, such as the difficulty of producing large scale 3D scaffolds without the use of any further processing technique may limit the application of electrospun meshes only to 2D applications, such as skin grafts, drug delivery patches, etc. Some modifications of the collectors are also studied for the production of 3D structures such as vascular grafts using rotating mandrel tubes or other special geometries used as collectors. Applications as load bearing scaffolds (e.g. bone scaffolds), are also purposed in literature although scaffolds application is limited due to the lack of adequate compressive strength.
The investigation of melt-ES for TE applications is still at the first stages, but the interest on this technique has increased over the last five years. In comparison with solution ES, it is possible to avoid the use of organic solvents and, scaffolds with larger pore size suitable for a better cell penetration can be obtained. In addition, the possibility to have a better control over the orientation of the collected fibres allows the production of 3D scaffolds with improved mechanical properties, control over the pore architecture and external geometry. Moreover, the high throughput that characterizes melt-ES allows a faster production of scaffolds with larger thickness as compared to solution ES. As shown by Eisken et al [Erisken et al., 2008a, b], who combined a twin-screw extruder with solution ES to produce multi-scale composite polymer/ceramic scaffolds, the combination of solution and melt-ES allows to exploit the advantages of the two techniques.
New technologies are expected to revolutionize the production of nanofibres. A technique was recently developed to produce nanofibres using centrifugal forces, instead of an electrical field, to generate the polymer jet and achieve the subsequent elongation of the fibre [Badrossamay et al., 2010; Sarkar et al., 2010; Wang et al., 2010c]. The use of centrifugal forces to induce polymer jets from melt or solution is proposed as an alternative to the conventional methods of melt and solution ES. This new technique allows a revolution in nanofibre production: the high quantities that can be
produced over time allow the scaling-up of the process. Some researchers are exploring also the combination of ES and centrifugal forces, termed electrocentrifugal spinning [Dabirian et al., 2011].
1.2.4 Wet-spinning
Wet spinning, as well as solution ES, involve the use of a suitable solvent to obtain a polymeric solution. In particular, in wet spinning, the polymeric solution is extruded through a spinneret immersed into a coagulation or precipitation bath and a continuous filament is produced [Harper and Petrie, 2003]. After precipitation, the filament undergoes further processing steps including stretching, washing, drying and winding until the final fibre is obtained [Carraher, 2010]. Variants of the previously mentioned processes can combine various approaches, for example, in dry-wet spinning the polymer solution is extruded trough the spinneret and an air-gap between the spinneret and the coagulation bath constitutes the first step in the solvent elimination.
The introduction of wet spinning into the TE field was proposed due to the limitations of some existing processes in producing fibres from natural polymers. Processes such as melt spinning that use high processing temperature might induce polymer degradation of natural materials such as gelatin or chitosan. Table 2 presents the polymers tested using the wet-spinning technique to produce fibres or scaffolds for TE applications.
Natural polymers have been extensively studied for the production of scaffolds and fibres for diverse application in TE and regenerative medicine using wet-spinning technique. The production of hybrid or composite fibres is also feasible by selecting suitable solvents. Hybrid fibres using alginate/chitosan are difficult to produce using the direct preparation of the spinning solution due to the formation of gels as a consequence of the ionic interactions of these two oppositely charged molecules [Watthanaphanit et al., 2008]. Watthanaphanit et al studied a new approach involving the preparation of a chitosan emulsion in order to avoid gelation, and the prepared alginate/chitosan fibres were proposed as drug delivery carriers.[Watthanaphanit et al., 2008]. Recently, a new method termed hydro-spinning was developed by Wang et al [Wang et al., 2010a] to produce alginate/chitosan hybrid fibres. The concept consisted of the precipitation of a chitosan solution into a sodium alginate coagulation bath under stirring, resulting in ribbon-like fibres. The same group presented also a technique termed spray-spinning [Wang et al., 2010b]. In this case, small drops of chitosan solution were air-sprayed into the previously described coagulation bath. In vivo experiments showed the biocompatibility of the developed alginate/chitosan fibres. Despite the possibility of rapidly producing fibres with these techniques, some limitations arise for the preparation of 3D scaffolds.
Table 2 - Natural and synthetic materials processed using wet-spinning technique for the production of scaffolds for TE applications. [R] – random fibres
Reference Natural [Funakoshi et al., 2005] [Yamane et al., 2005] [Heinemann et al., 2009] [Wang et al., 2010a; Wang et al., 2010b] [Watthanapha nit et al., 2008] Synthetic [Ellis and Chaudhuri, 2007] [Razal et al., 2009]
Non-Solvent CaCl2 dissolved in water/methanol solution CaCl2 dissolved in water/methanol solution n/a Aqueous sodium alginate solution CaCl2 dissolved in water/methanol solution + MeOH. Water Isopropanol
Solvent Aqueous acetic acid solution Aqueous acetic acid solution n/a Acetic acid/Sodium acetate Aqueous sodium alginate solution (with or without chitosan-citrate emulsion) 1-methyl-2- pyrrolidinone or dioxane Chloroform
Target tissue Ligament Cartilage Bone Drug release fibres Bone Muscle
Properties Fibre diameter = 30 µm Fibre diameter = 30 µm Fibre diameter = 20 µm Fibre diameter = 4.19±2.43 -9. ±3.83 µm Fibre diameter = 50 – 200 µm Fibre outer diameter = 700 µm and inner diameter = 250 µm Fibre diameter = 25 - 35 µm
Fibre orientation 3D Scaffolds; 3D Scaffolds; 3D Scaffolds; [R] Ribbon-like fibers Fibres Hollow fibre Fibres
Polymers Chitosan/ hyaluronan Chitosan/ hyaluronan Chitosan with collagen coating alginate/chitosan alginate/chitosan PLGA PLA:PLGA
Table 2 – Cont.
Reference Synthetic (Cont.) [Razal et al., 2009] [Morgan et al., 2007] [Rissanen et al., 2010] [Gao et al., 2007] [Yilgor et al., 2010] [Williamson et al., 2006] [Williamson and Coombes, 2004] [Puppi et al., 2011]
Non-Solvent Isopropanol Water Ethanol Isopropyl alcohol or a mixture of isopropyl alcohol with methanol Methanol Methanol Methanol Methanol, ethanol and water
Solvent Chloroform 1-methyl-2- pyrrolidinone Dichloromethane Chloroform Chloroform Acetone Acetone Acetone or chloroform
Target tissue Muscle Bone n/a n/a Bone Vascular Soft TE applications Bone
Properties Fibre diameter = 25 - 35 µm Fibre outer diameter= 770±60 µm and inner diameter=500±30 µm Fibre diameter = 46 - 70 µm Fibre diameter = 50 - 180 µm n/a n/a Fibre diameter = 150 - 190 µm Fibre diameter = 100 – 250 µm
Fibre orientation Fibres Hollow fibre Drug delivery fibres Drug delivery fibres 3D Scaffolds; [R] Tube produced with luminal fibres Single fibre Gravity spinning 3D scaffolds [R]; Loaded with clodronic acid disodiumand
Polymers PLA:PLGA PDLLGA PLDLLA PLDLA PLLA PCL PCL PCL *PCL; *PCL/nHA
Table 2 – Cont.
Reference Hybrid [Pashkuleva et al., 2010] [Malheiro et al., 2010] [Tuzlakoglu et al., 2010] [Leonor et al., 2011] [Yilgor et al., 2009]
Non-Solvent Water Methanol Methanol Methanol or calcium silicate solution (Si(OC2H5)4:H2O:C 2H5OH:HCl:CaCl2) Aqueous solution of Na2SO4 and NaOH
Solvent Dimethyl sulfoxide Formic acid/acetone Chloroform Chloroform Aqueous acetic acid solution
Target tissue Nonload bearing applications TE applications Bone Bone Bone
Properties Fibre = 100 µm Porosity=80.5% Fibre = 113 ± 29 to 140 ± 30 µm (dry conditions) and 136 ± 28 to 374 ± 66 µm (wet conditions) Fibre = 100 µm pore size = 250 µm Fibre = 100 µm Porosity= 57%
Fibre orientation 3DScaffolds; [R]. 3DScaffolds; [R]. 3DScaffolds; [R]. 3DScaffolds; [R]. R non-woven
Polymers Starch/ EVOH PCL/chitosan Starch/PCL Starch/PCL Chitosan; Chitosan/PEO
Yilgor et al proposed the development of wet-spun 3D scaffolds composed of chitosan and chitosan/poly(ethylene oxide) (PEO) fibres, loaded with PLGA nanoparticles containing bone morphogenetic protein-2 (BMP-2) and PHBV nanocapsules containing bone morphogenetic protein-7 (BMP-7) that were either adsorbed on the surface or incorporated into the fibres [Yilgor et al., 2009]. The use of nanoparticles made of different materials allowed a sequential delivery of the BMPs, inducing an improved differentiation of rat bone marrow mesenchymal stem cells (MSCs). Moreover, the presence of the nanoparticles on the surface of the wet-spun fibres resulted in improved alkaline phosphatase (ALP) activity but in lower cell vitality. The opposite effect was observed in the fibres containing the nanoparticles in their interior. A further study performed by the same authors [Yilgor et al., 2010] on PCL scaffolds compared the architectures produced using either an AM technique (3D plotting) or wet-spinning. The scaffolds were coated with the previously mentioned nanoparticles, and results showed that wet-spun scaffolds composed of randomly oriented fibres showed a better cell proliferation and differentiation when compared to AM scaffolds composed of aligned fibres.
Since no heating is involved, wet-spinning allows the production of fibres and scaffolds loaded with bioactive agents (protein, peptides, drugs, etc.). Single fibres of poly(L-lactic acid) (PLLA) were purposed for the delivery of an anti-cancer drug (5-fluorouracil) [Gao et al., 2007]. The mild processing conditions and the absence of heating allowed the maintenance of drug efficacy.
Another group of polymer studied for the production of wet-spun fibres are polylactide stereocopolymers [poly(L,D-lactide) (PLDLA) and poly(L,DL-lactide) (PLDLLA)] [Rissanen et al., 2009]. The production of bovine serum albumin loaded fibres was studied by combining this protein with two different polylactide polymers, namely (PLDLA) and (PLDLLA) [Rissanen et al., 2010]. However, the produced filaments showed unsatisfactory mechanical properties and a slow release of the loaded protein. Recently, our group developed *PCL wet spun non-woven structures loaded with a bisphosphonate and HA for application as TE scaffolds able to inhibit bone resorption associated with various bone diseases, such as tumour-associated altered bone metabolism or osteoporosis [Puppi et al., 2011].
Another advantage of the wet-spinning technique is the feasibility of producing composite fibres made of a combination of natural and synthetic polymers dissolved in the same organic solvent or in a co-solvents system. As an example, Starch was combined with PCL in composite fibres constituting 3D scaffolds (30/70 wt %) [Santos et al., 2010]. The developed scaffolds were treated with gas plasma showing higher cell viability and ALP activity when compared to the non-treated samples. A thermosensitive starch-based composite scaffold [starch/poly(ethylene-co-vinyl alcohol)] was developed using a combination of wet-spinning and fibre-bonding techniques [Pashkuleva et al., 2010]. After the production by wet-spinning, scaffolds were dehydrated, left to
dry and a subsequent compression process (fibre bonding) was applied in order to improve the cohesion among fibres. Further plasma treatment was performed on the prepared scaffolds and improvements in cell vitality were observed when compared to untreated ones. PCL/chitosan composite fibres with various weight ratios were developed by Malheiro et al [Malheiro et al., 2010] by employing a co-solvent system of formic acid and acetone. Results showed various fibre roughnesses, depending on the PCL/chitosan blend percentage used.
Another processing way to fabricate functionalized wet spun fibres involves the use of a coagulation bath containing molecules capable to react with the polymer involved. Leonor et al [Leonor et al., 2011] studied the preparation of bioactive scaffolds capable of inducing osteogenesis by preparing starch/PCL fibres precipitated into a coagulation bath containing calcium silicate. The functionalization of the fibres during phase inversion induced the presence of silanol (Si–OH) groups on the fibre surface. In vitro tests performed with goat MSCs showed that the presence of the functional groups induced an improved differentiation towards osteoblastic phenotype when compared to control fibres. Other types of functionalization can be performed after the production of the wet-spun fibres. Scaffolds produced from chitosan wet-spun fibres were coated with collagen type I and studied for bone TE [Heinemann et al., 2009]. Collagen coating allowed an improvement of the biocompatibility of these scaffolds in vitro.
The possibility of producing hollow fibres is also one of the potential benefits of the wet-spinning process. Hollow fibres combined with MSCs have been proposed as scaffolds for bone regeneration [Morgan et al., 2007]. Results of implanted poly(DL-lactide-co-glycolide) (PDLLGA) hollow fibres pre-seeded with MSCs into mice showed collagen type I deposition, osteoid formation and bone mineralisation [Morgan et al., 2007]. Another study suggested the application of PLGA hollow fibres to provide better mass transfer for cell culture and pseudo-vascularisation of the scaffolds in vitro and possible angiogenesis for in vivo application [Ellis and Chaudhuri, 2007].
Most published studies propose wet-spun scaffolds for bone TE [Heinemann et al., 2009; Leonor et al., 2011; Puppi et al., 2011; Yilgor et al., 2010; Yilgor et al., 2009], although some other applications, such as muscle [Razal et al., 2009], ligaments [Funakoshi et al., 2005], cartilage [Yamane et al., 2005] and vascular TE applications [Williamson et al., 2006], are also considered.
Razal et al [Razal et al., 2009] showed that the use of wet-spun biodegradable fibres seeded with myoblasts could be an ideal approach for muscle repair.
Vascular grafts production was investigated by combining wet-spinning and ES techniques [Williamson et al., 2006]. A first layer of wet spun PCL fibres were wrapped around a mandrel and a second layer of PU was electrospun over the PCL fibres. The developed scaffolds were seeded
with human umbilical vein endothelial cells showing good cell attachment and proliferation with the formation of a cells monolayer.
Wet-spun fibres were also investigated as scaffolds for cartilage TE [Yamane et al., 2005].
Chitosan fibres were produced using a calcium coagulation bath, then stretched and wound around a roller [Yamane et al., 2005]. Chitosan fibres coated with hyaluronic acid were also produced using a secondary coagulation bath prior to the stretching and winding phases. Fibres were then cutted and stacked in a perpendicular arrangement. Chondrocytes seeded scaffolds showed better adhesion, proliferation and production of new ECM on the hybrid chitosan/ hyaluronic acid scaffolds when compared to chitosan scaffolds.
In conclusion, the feasibility of producing fibres from a wide range of natural and synthetic polymers and composite fibres capable of incorporating bioactive agents makes wet spinning a promising technique for developing TE scaffolds. Wet-spun fibres functioning as controlled drug releasing systems can aid to a faster healing of a wound site [Polacco et al., 2002]. Moreover, the feasibility of producing fibres of various formats, (e.g. hollow fibres, porous fibres and solid fibres), and with various diameters allows to a have a good control over drug release kinetics.
However, the poor control over internal and external structure limits its application in the fabrication of complex 3D structures. Moreover, the low mechanical properties of the typical non- woven structures produced by wet spinning do not allow their application on load-bearing sites.
The advantages and limitations of this technique are summarized in Table 3. Further improvements of this technique are being implemented and tested for the production of 3D scaffolds with controlled pore size and external geometry (see chapters 5, 6 and 7).
Table 2 – Advantages and limitations of wet-spinning technique for the production of scaffolds for TE
Advantages Disadvantages
• A wide range of natural and synthetic polymers can be processed;
• Possibility of loading bioactive agents;
• Possibility of developing drug release scaffolds;
• Thermal degradation is avoided (no heating);
• Various fibre forms (hollow, solid) and various fibre diameters can be produced.
• Non suitable for fibre alignment;
• Limited control over the scaffold’s external geometry and internal architecture of pores;
• 3D scaffolds with poor mechanical properties;
• Step of residual solvent and non- solvent removal necessary in some cases.
1.2.5 Additive manufacturing (AM) techniques
AM techniques, also known as solid freeform fabrication (SFF) techniques, are based on the construction of 3D objects built layer-by-layer. Using AM techniques, it is possible to produce complex 3D structures with good mechanical properties and a predefined internal architecture and external geometry. These techniques allow for the production of a fully interconnected network of pores with customizable size, shape and distribution. 3D models used in AM are generally treated or designed using Computer Aided Design (CAD) and computer-aided manufacturing (CAM) software. The 3D data for the production of scaffolds with complex external geometry and internal pore architectures can derive from medical imaging techniques such as computer tomography (CT) and magnetic resonance imaging (MRI) that nowadays are extensively used for diagnostic purposes [Hieu et al., 2005; Sun and Lal, 2002]. Other approaches for the design of a complex network of pores involve mathematical equations [Gabbrielli et al., 2008] and models developed for scaffold topological (shape) optimization [Almeida and da Silva Bártolo, 2010].
After data treatment, the information is generally converted to a stereolitography-file (STL-file) (containing the information of the surface geometry of a 3D object) from the major part of the computer-aided design CAD software. This neutral file is composed of different information regarding the 3D model. This model is tasselled into layers originating and slice file (SLI). This file contains the information related to each layer that is built by the AM systems. Various patterns can be programmed in order to obtain the desired internal structure. Generally, the calculation of the SLI file and subsequent generation of the patterns is performed by customized softwares.
AM techniques may be divided into four categories: i) Stereolithography (SLA); ii) Selective Laser Sintering (SLS); iii) Three-Dimensional Printing (3DP); and iv) Fused Deposition Modeling (FDM) [Hutmacher et al., 2004]. A plethora of new AM techniques has been developed on the basis of the initial techniques developed for other industrial purposes. They were adapted to enhance material processing and to operate on a wider range of materials. The base techniques used on the TE field are described below and a general overview of the new approaches is also reported.
1.2.5.1 Stereolithography (SLA)
SLA is an AM technique that uses an ultraviolet light or laser to polymerize a photosensitive polymer by selectively irradiating a layer of polymer. The first developments on the production of 3D models obtained by the photo polymerization of a liquid resin using ultraviolet (UV) light were proposed by Kodama [Kodama, 1981], who developed two approaches to polymerize subsequent layers until a 3D solid model was obtained: the first used masks for the definition of each layer and the second an optical fibre to conduct light to selectively polymerize the resin inside a container. In
the latter, a predefined pattern was achieved by controlling the movement of the fibre in the X and Y axis. In 1984, Hull attributed the denomination of SLA to the process of producing 3D solid objects by creating cross-sectional patterns [Hull, 1986; Hull, 1990]. Generally, the SLA process involves the construction of 3D structures applying a bottom-up principle. A vat containing the photo-polymerizable materials includes a construction platform, whereas the 3D model is built layer upon layer. After the polymerization of the first layer, the platform lowers and a recoat bar places a new uniform layer of resin on top of the previously built one. The distance by which the platform descends depends on the polymerized thickness. The polymerization of the second layers is performed by overlapping a percentage of the previously built one in order to prevent delamination of the layers. The process is repeated until the 3D object is built. Further processing steps include the removal of the non-polymerized resin and the post-curing of the green part to finalise curing of the 3D object and to improve polymerization between layers and to reduce of surface irregularities [Hutmacher et al., 2004; Melchels et al., 2010b].
A recently developed technique involves a digital light projector source to direct light using a Digital Micromirror Device™ (DMD) constituted by an array of mirrors that selectively divert the light to the vat containing the photo polymerizable polymer [Melchels et al., 2010b]. This technique uses a smaller quantity of photo polymer when compared to conventional SLA equipments, although some limitations of the scaffold construction might happen due to the mechanical forces produced when the layer that is being built is detached from the glass on which the polymer is being irradiated.
The possibility of producing highly complex structures capable of improving the performance of scaffolds both in vitro and/or in vivo is one of the advantages of SLA techniques. However, to achieve the optimal viscosity in the photocrosslinkable liquid resin, toxic diluents are generally used [Elomaa et al., 2011]. Moreover, the commercially available resins (epoxy-based or acrylate- based) generally processed with SLA present limited biocompatibility and biodegradability [Jansen et al., 2008; Melchels et al., 2010b]. The first attempts to use SLA technology for the production of 3D scaffolds was performed using an indirect approach due to the limited biocompatibility of the used resins. Levy et al employed a slurry consisting of a photocurable acrylic resin with a suspension of HA powder [Levy et al., 1999]. After the production of the scaffold by the polymerization of the acrylic resin, a further step of sintering was performed to extract the resin and the final obtained structures were made of HA. The production of a “lost-mould” using SLA techniques was another approach developed by Chu et al [Chu et al., 2001]. The steps involved in the scaffold production comprised the production of the epoxy mould and a subsequent infiltration of the mould with a 40 vol% HA suspension in propoxylated neopentyl glycol diacrylate and iso- bornyl acrylate. After infiltration, the moulds and the binders used were removed by pyrolysis and
a further sinterization of the scaffold was performed. The direct use of SLA technology to produce scaffolds was only possible after the development of new biocompatible and/or biodegradable materials. However, the amount of photopolymerizable biomaterials processed using SLA with suitable biocompatibility and biodegradability properties is still limited [Hutmacher et al., 2004].
Only recently, some new biodegradable and biocompatible polymers have been developed for scaffolds production. Fumaric acid monoethyl ester functionalized with three-armed poly(D,L- lactide) (PDLLA) oligomers and N -vinyl-2-pyrrolidone was proposed for the production of scaffolds for TE [Jansen et al., 2008], that were prepared using a gyroid architecture of pores calculated mathematically in order to achieve a larger specific surface area and high interconnectivity. Mouse fibroblasts were used to evaluate the biocompatibility of the synthesized polymeric material, but no tests were performed on the SLA produced scaffolds. Star-shaped (PDLLA) oligomers were functionalized using methacryloyl chloride and this resin was diluted with ethyl lactate and combined with a photo-crosslinker, and used to produce scaffolds and films using SLA equipment [Melchels et al., 2009]. The films prepared with SLA were seeded with murine preosteoblast, and cells adhered and proliferated well, although cell number was significantly lower than the control grown on tissue culture polystyrene (TCP). The previously mentioned PDLLA based resin was used for the production of gyroid scaffolds, and compared to salt leaching scaffolds with equivalent porosity [Melchels et al., 2010a]. Results obtained from static and dynamic cell culture indicated the advantage of producing a highly interconnected network of pores obtained by the SLA process when compared to salt leaching scaffolds. Moreover, the uniform dispersion of cells on the SLA scaffolds was correlated to the pore architecture (gyroid channels).
Natural or synthetic hydrogels are highly hydrated polymer networks that when combined with cells and with biomimetic and ECM components are capable of triggering cell responses and inducing the formation of new tissue [Fedorovich et al., 2007]. Hydrogels that encapsulate cells are another type of living construct that are being developed for TE applications by crosslinking hydrogels with UV light [Tsang and Bhatia, 2004]. The success of the production of scaffolds that encapsulate cells depends on the UV light intensity, the exposure time for scaffold polymerization, and free radicals that can be generated from the photo-initiator [Lu et al., 2006].
Photopolymerizable hydrogels, such as PEO with poly(ethylene glycol) (PEG) dimethacrylate, were used for the production of constructs encapsulating Chinese hamster ovary cells [Dhariwala et al., 2004] by means of SLA. Despite some cytotoxicity being induced by the photoinitiatior, the lowest initiator amount tested allowed the achievement of good cell vitality. The mechanical properties of the produced structures were ideal for the replacement of soft tissue. Further improvements studied by Arcute et al [Arcaute et al., 2006] using PEG dimethacrylate allowed an increase in encapsulated human dermal fibroblasts survival observed after 24 hours.
