Upon insertion of a biomaterial into the body, all the features above described are involved. First, there is damage to the tissue during the insertion, and then an acute inflammatory response occurs. This is followed by the wound healing response that can proceed normally if the biomaterial is there for a short period of time, otherwise the response is distinct. The normal tissue repair cannot take place since there is a foreign body in the way. The first reaction of the body is to try to eliminate it. At this stage, the phagocytes come in and the acute inflammatory response may progress to chronic response. If it is not a degradable material or prone to phagocytosis, then the reaction will follow one of the two paths. The usual response consists in the formation of a fibrous capsule around the biomaterial by the fibroblasts in an attempt to eliminate it from the body. It is generally accepted that the thickness of the fibrous capsule is an indication of the material’s biocompatibility, i.e. a thinner capsule corresponds to a more biocompatible material [38]. There are two types of factors that can influence the wound healing process, the intrinsic factors and the extrinsic factors. The bulk nature of the biomaterial, its porosity, roughness and changes in the surface chemistry are intrinsic factors of the implant [39]. The extrinsic factors are, for example, the surgical procedure, the condition of the patient (diabetic, immunocompromised) and the anatomical location of the implant [4]. The less common response is for the inflammatory process to continue as a chronic inflammatory response and to progress to giant cells and granulomas. This occurs when the material is not biocompatible, and the host reaction is still trying to neutralize it.

At this stage some problems can occur associated with the host reaction to the biomaterial ultimately leading to infection. The formation of the fibrous capsule is an indication that the material is biocompatible and will occur as an early step in healing. Fibrous capsule formation needs to be considered when predicting device function [19]. For instance, if the device is to serve as a drug delivery system, the formation of the fibrous capsule may alter the permeability of the device and the diffusion of the drug so that the function of the device will not be anticipated from laboratory studies. Formation of a fibrous capsule generally does not occur with porous or textured materials [15]. If the interstices are large enough, local vascularized tissue will grow into the pores rather than form the capsule. Long-term percutaneous catheters have been coated with velour to facilitate ingrowth and anchorage. Vascular prostheses made of fabric demonstrate ingrowth of vascular-like tissue [37]. Joint replacement prostheses are sometimes coated with metal beads or wire mesh to facilitate ingrowth of bone and biological anchorage [37]. It is important to notice that acute inflammation and wound healing are necessary steps in the implantation of biomaterials and devices that enter the body, hence the effect of these responses on the function of the device must be considered.

Surface roughness (Fig. 4.4a) can essentially be divided into nano-roughness (<100 nm), micro-roughness (100 nm–100 μm) and macro-roughness (from 100 μm to several millimetres) [61]. The cells response to roughness will vary with the cell type and roughness scale (i.e. size, shape and distribution on the materials’ surface). Hence, surface roughness regulates the biological response of cells and tissues that are in contact with the implanted material. It provides cues to promote cell adhesion, orientation, migration and production of ECM. For instance, cells grown on micro-rough surfaces spread more and differentiate better than on smooth ones, as shown by their gene expression [62–64]. Also, on grooved materials, cells can orient themselves in the direction of grooves and proliferate faster than on a smooth surface [43, 65–70].

Nanotopography may provide superior biomimetic cell-modulating cues, as it may further resemble the natural ECM environment in which cells reside and interact. It has been shown that nanopatterned surfaces can improve protein adsorption and cellular response, since as on rougher surfaces, the focal adhesion points are located at cell edges, where the contact with the materials’ surface takes place [61, 71–77].

Porous structures (Fig. 4.4b) typically consist of irregularly shaped voids with interconnecting channels (voids). The pore size, porosity (the fraction of void volume) and surface area (also presented as a surface-to-volume ratio parameter when characterizing a biomaterial), pore shape, wall morphology (tortuosity) and pore interconnectivity are important requisites for a scaffold in tissue engineering applications. The tailoring of these properties allows for rapid cell attachment and proliferation, improve nutrient and cell transfer to the scaffolds’ center (cell and tissue ingrowth), as well as metabolite dispersal and new tissue formation [78–84]. The open pore structure and interconnected porosity have also been found to disrupt fibrous tissue deposition, minimize the formation of foreign-body capsules, improve tissue healing, and increase vascularization of the tissue surrounding the implant [85–90]. Pore size (microporosity: pore size <10 µm; macroporosity: pore size >50 µm) can be tailored to meet specific applications (Table 4.3); for instance the optimum size for neovascularization is around 5 µm; whereas for fibroblast ingrowth is in the range of 5–15 µm; for hepatocytes ingrowth close to 20 µm; for the regeneration of adult mammalian skin around 20–125 µm; for osteoid ingrowth around 40–100 µm and for bone regeneration around 100–350 µm. On the other hand, fibrovascular tissues require pores sizes greater than 500 µm for a rapid vascularization and for the survival of the transplanted cells [91, 92].

One of the most common techniques used to measure the surface energy of a material is the determination of contact angles, whereby the angle made by a drop of liquid deposited on a surface correlates with its hydrophobicity (Fig. 4.4c). The size and shape of the liquid droplet can be explained as a balance between the force with which the molecules of the liquid drop are being attracted to each other (designated by cohesive force) and, on the other hand, the attraction of the liquid molecules for the surface (adhesive force). The balance between these two opposing forces is expressed by the contact angle value (θ), which in turn translates the hydrophobicity of a given surface. Consequently, relatively more water wettable surfaces (i.e. exhibiting lower θ) are termed hydrophilic (high surface energy); conversely less wettable surfaces (with higher θ) are termed hydrophobic (low surface energy). To establish a borderline between these two terms, it is generally agreed that the criterion of a θ = 65° separates between the two regimes [100–105].

In the previous sections it was already explained that many proteins adsorb onto an implant surface immediately upon contact with it. This event modulates subsequent cell adhesion and/or foreign-host response (protein adsorption is the first step in the integration of an implanted device with its surrounding tissues). The dominant response will depend largely on the hydrophobicity (also referred to as surface energy or wettability) of the scaffold. The materials hydrophobicity thus affects protein adsorption, platelet adhesion/activation, blood coagulation and cell and bacterial adhesion. For instance, the adsorption of serum proteins (fibrinogen, fibronectin or vitronectin) later affect the adhesion of leukocytes, macrophages or platelets by preventing ECM proteins adsorption, ultimately leading to fibrous encapsulation; thus delaying or even preventing proper cell adhesion. Also, surface wettability (hydrophobicity) was found to regulate cytoskeletal organization and cell morphology. As a general rule, the more hydrophilic a surface is, the more cells adhere to it. Conversely, hydrophobic surfaces are more prone to protein-adsorption. This occurs due to the strong hydrophobic interactions between the material’s surface and the protein’s hydrophobic groups [6, 93, 106–113]. With respect to contact angle values, literature shows that polymer surfaces exhibiting moderate wettability (i.e. θ in the range 40–70°) enhance cell adhesion [106, 114, 115].

In the biomedical field, protein-surface interactions play key roles for the assembly of interfacial protein constructs, such as biosensors, activators and other functional components at the biological/electronic interface. In analytical sciences, non-specific protein adsorption on sensor surfaces, protein chips, or assay platforms represents a serious problem as adsorbed proteins are responsible for the degradation and thus decrease of the device performance. For example, in biosensors for in situ monitoring of cell culture or blood glucose levels in diabetic patients, specificity and sensitivity levels, and durability depend greatly on the protein adsorption onto the sensor surface. Likewise, in polymeric drug matrix delivery systems the accumulation of adsorbed proteins may impair the diffusion of the drug. In the design of biocompatible materials for surgical implants, adsorption of fibrinogen in particular should be suppressed, to avoid blood agglutination [116–118].

Surface chemistry is very important for the physiological interactions with biomaterials and it is closely related to surface energy (hydrophobicity); the surface free energy of a material originates from its surface functional groups and electrical charges. The material’s surface chemistry mediates cell response given that surface chemistry directly impacts on protein adsorption and interaction (i.e. surface chemistry affects the amount and conformation of the adsorbed proteins). All these factors play crucial roles in the later cell interactions and functions [119–122]. Table 4.4 lists some of the material’s most common surface groups and describes how they affect proteins, cells and tissue interactions.