The term 'cranio-facial implants' has been used to describe endosseous implants inserted in the mastoid, orbital and nasal regions; and although some authors exclude the upper and lower jaws, both are parts of the craniofacial skeleton1. Cranio-facial implants have been classified in two groups: intra-oral dental implants, which are well-developed and extensively studied; and extra-oral implants, which are not as developed or studied as the intra-oral ones. The development of extra-oral implants has been slower in terms of design and applications because this type of implants has limited demand compared to the intra-oral implants. Furthermore, the placement of extra-oral implants can only be done in an operating room, and requires trained oral and maxillofacial surgeons. Additionally, the fabrication of these prostheses has to be performed by trained prosthodontists resulting in extra-oral implants being more complex and costly than the intra-oral counterparts2.

The loss of craniofacial tissues can result from neoplasm, trauma, tumor or cyst resection, infectious diseases, nonunion fractures, and congenital or developmental conditions (i.e., cleft palate defects); which results in serious functional, aesthetic and psychological sequelae. In these situations, the absence of hard and soft tissues can be disfiguring, and in many cases, it compromises basic functions such as mastication, speech, swallowing, leading to limited thermal and physical protection of important anatomical structures (i.e., brain, nerves, arteries, veins)3-6. In the United States, there is a clinical need for craniofacial bone regeneration, and more than 30,000 surgical procedures are performed each year to repair craniofacial bone defects7. Data also revealed that over 1 million skeletal-related craniofacial procedures were performed in 2002, including 16,338 craniotomies and 32,043 post-traumatic facial reconstructions3. The treatment of craniofacial bone defects remains challenging in terms of providing protection to the brain, preventing infection and maintaining adequate appearance. Consequently, the outcome of craniofacial bone reconstruction is thought to be dependent on surgical skills, the quality of adjacent soft tissues, the size and location of the defect and the choice of repair method4,7,8.

The long-term success of dental, facial, orbital, or auricular prostheses beyond primary reconstruction is dependent on the maintenance of its anchorage function. Providing adequate retention and support of the implant has been a constant challenge; as the inherent mechanical retention within the defect or the use of adhesive systems has proven to be either problematic or unacceptable3,4,7-10. Additional hurdles in the treatment of craniofacial bone defects include the presence of bacteria from the oral and sinus cavities, the ability of the implant to withstand mechanical stresses from the masticatory function, and the challenge on finding a cost-effective solution2-4,7,10.

In order to understand the causes of cranio-facial and dental implant failure it is first important to understand the process of implantation and the healing of the bone around the implant when the process is successful. Bone healing around an implant follows a similar process to fracture healing but is highly dependent on the surface characteristics of the implant. Blood at the implant surface supports the deposition of proteins, which is followed by coagulation, inflammation and tissue formation, all of which are regulated by the surface chemistry and topography of the implant11. Within seconds of blood contact with the implant, proteins adsorb to the implant surface which then allows for platelets to become activated and bind to the adsorbed protein. A clot then forms, which contains many signaling molecules which influence the migration of monocytes, neutrophils (both involved in inflammation), and mesenchymal cells (cells that can differentiate into osteoblasts) towards the implant surface12. When neutrophils and macrophages are activated they migrate to the implant site from nearby capillary beds release inflammatory mediators which are necessary for the initiation of bone formation. Members of the tissue growth factor β (TGF-β) superfamily are also expressed within 24 hours of implantation, including bone morphogenetic protiens (BMPs) and growth and differentiation factors (GDFs). These signaling molecules allow for the recruitment, migration, and differentiation of mesenchymal cells, which take part in the formation of woven bone13. Woven bone is then remodeled resulting in the formation of mature bone, the desired end result13.

When proper bone healing around the implant does not occur, implant failure results. There are many causes of implant failure, but most can be broken up into two main categories: aseptic loosening and infection14. Aseptic loosening, not associated with infection, can result from the inability of the bone to integrate with the adjacent implant surface15,16. Inadequate bone integration with the implant can lead to implant migration which stimulates the foreign body reaction and can lead to infection and tissue necrosis. Modifying the surface roughness and chemistry of the implant can affect the ability of the implant to induce strong osseointegration15. Roughening treatments have been employed in attempts to increase cell attachment and cell proliferation17. Moreover, bioactive coatings have been added to implants to further improve cell attachment and differentiation while reducing the likelihood of loosening15.

While prosthetic implant infection (PII) can arise perioperatively or postoperatively, the majority are due to the introduction of bacteria directly into the patient during or soon after the surgery (perioperatively) and occur within 3 months of implantation18. The bacteria can originate from skin flora present around the site of surgery, from the bacteria present in the mouth, or from the surgeon. Furthermore, the conditions of the surgical wound, such as clotted blood and compromised soft tissue, make the site ideal for bacterial colonization. In the cases of acute infection, local cellulitis is produced which leads to the death of leukocytes, an increase in bone pressure, a decrease in pH, and a decrease in oxygen tension. Blood circulation is thus compromised which ultimately leads to the necrosis of large segments of bone19. Chronic PII extending from the bone-implant interface into the surrounding native bone and marrow is termed osteomyelitis and can result in significant bone loss and implant damage in severe cases20. Postoperative PII is usually caused by a single bacterial species (monomicrobiotic), the most common of which is Staphylococcus aureus. Infections caused by these microorganisms are becoming more worrisome due to the increase in multiple-antibiotic-resistant strains, such as methicillin-resistant S. aureus, which is now the most commonly isolated nosoco-mial bacterial pathogen in most of the world. While the immature or compromised immune system of the host is the primary cause of initial infection, the development of the infection into a persistent and chronic one is generally caused by other species of bacteria, such as Enterococcus spp., Strepto-coccus spp., Pseudomonas aeruginosa, Enterobacter spp., Mycobacterium spp., and Candida spp19. Bacteria present at the implant can also lead to biofilm formation which is of much concern due to the difficulty in eradicating them. Several strains of bacteria, especially those found in the oral cavity, are capable of forming biofilms. Biofilms form when they attach to the material surface and then begin to grow in multilayered cell cluster which are then surrounded by a slimy matrix produced by the bacterial cells21. Biofilms can be a thin single layer of cells, or they can be thick with complex architecture in which the microcolonies form distinct pillars or mushroom-shaped structures. Between these pillars runs an intricate channel network through which nutrients can be transported, even to the deepest areas of the biofilm22. One benefit for bacteria that exist in biofilms is that the extracellular matrix is able to seize and concentrate several environmental nutrients. Furthermore, the bacteria that grow in biofilms are more resistant to several removal tactics, such as elimination by antimicrobial or antifouling agents (mediated by low metabolic levels and downregulated rates of cell division), shear stress, host phagocytic clearance, and host oxygen radical and protease defense23. Biofilms are also able to slow the infiltration of some antimicrobial agents, and in many cases, inactivate them in the process. The biofilm can also prevent the host inflammatory molecules from entering the biofilm, thus leading to a resistance to the host response. The host response itself can lead to host cell lysis and subsequent damage to the host tissue, resulting in the release of host cell components which act as nutrients for the bacteria. Finally, the biofilm has the potential for dispersion by way of detachment. This means that the microcolonies that exist in the pillars can detach under the direction of mechanical fluid shear or through a genetically programmed response that mediates the detachment process. The detached microcolonies can then travel under the direction of the fluid flow and attach and promote biofilm formation to other areas in the host that were previously uninfected19.

Titanium is the most commonly used material for bone-contacting dental implants due to its high biocompatibility and good mechanical properties. Titanium and its alloys spontaneously form a titanium oxide (TiO₂) layer on their surfaces which contribute to many of their excellent properties such as corrosion resistance. Furthermore, compared to other metals, titanium has a relatively low modulus, reducing the potential for stress shielding, as well as good fatigue strength15,24,25. It has been shown to produce very little fibrous tissue which allows for bone to easily grow on its surface26. Moreover, it has been shown to have a high capacity to join with bone. However, simple machined surfaces require several months of healing before bone integration occurs which means that there is a latency period of several months before the implant can undergo mechanical loading. To this end, several surface modification techniques have been employed to shorten the time between implantation and use of the site, summarized in Table 127.

Grit-blasting and acid etching among the most commonly employed surface modification techniques used in commercially available implants. Sand blasting increases the surface area of the implant compared to machined surfaces. This increase in surface area has been shown to improve cell attachment and proliferation which results in increased implant stability15,24,28. SLActive (Institut Staumann AG, Basel, Switzerland), for example, has been shown to increase wound-healing rate when compared to SLA (Institut Staumann AG) which may be attributed to its greater hydrophilic surface which results from its thicker oxide layer29-31. In these animal studies, a healing chamber model was used as opposed to appositional bone formation which is what is typically seen for screw root form implants. Electrochemical anodization is another chemical surface modification method that has been employed. It increases the surface microtexture and changes the surface chemistry of the implant resulting in a TiO₂ layer that is several orders of magnitude thicker than a passivated surfaces32. This surface modification, seen in the TiUnite implant, has been shown to increase the host/implant response at early implantation times relative to other surfaces33-36.

The addition of a ceramic coating to the roughened surface has gained much popularity due to their increased osseoconductivity. Integra-CP (Bicon Dental Implants, Boston, MA, USA), for example, employs a plasma sprayed hydroxyapatite (HA) coating which results in an irregular surface. The process involves blasting the surface with HA particles at the implant surface at high temperatures, resulting in a cracked coating as the coating undergoes rapid cooling. While this coating has shown enhanced bone-to-implant contact magnitudes at early implant times in vivo, the technique has compromised bone-coating interface mechanical properties in addition to non-uniformity in degradation after long periods in function37-43. For these reasons, these types of coatings have fallen out of favor in dentistry. Alternatives have been recently employed in commercially available implants. This includes the discrete crystalline deposition of nano-sized calcium phosphate as seen in the NanoTite implant. In this surface modification, the titanium surface undergoes a dual acid-etched treatment followed by nano-HA deposition. A clincial pilot study showed higher bone implant contact (BIC) after two months44. The addition of a fluoride treatment to roughened titanium surfaces has also been employed as seen in Osseospeed. Fluoride treatment has been shown to enhance gene expression in cell arrays and enhance host-to-implant response at early implant times45. An in vitro study showed that after 14 days of culture with osteoblast cells, the cells grown on Osseospeed expressed increased levels of alkaline phosphatase activity and collagen I production compared to TiOblast and tissue culture plastic. Further, a higher number of calcium crystals were evident on the Osseospeed substrates46. An in vivo animal study conducted in canine mandibles, the Osseospeed implant was shown to have a higher BIC compared to the TiOblast implant47. While several advances in surface modification have been made in order to improve implant osseointegration, no treatments address the issue of infection to the best of our knowledge. While some claim to be bacteria-proof due to their tight interlocking, the implant itself does not prevent bacterial attachment which can lead to biofilm formation and finally implant failure.

Several synthetic polymers and chitosan, have exhibited antibacterial properties and can be used as coatings on dental implants to prevent infection at the implant site. Martin et. al.57, developed a poly (dimethylaminomethyl styrene) coating, deposited by chemical vapor deposition, that was effective in killing 99.9999% of Escherichia coli and Bacillus subtilis bacteria. While the coating was deposited on a nylon substrate, the authors suggest that the coating could be used for other biomedical implants, including dental and other craniofacial implants. Another polymer coating, a poly(L-lysine)-grafted-poly(ethylene glycol) copolymer was shown by Harris et al.76, to reduce S. aureus adhesion as well as osteoblast adhesion. However, the addition of an arginine-glycine-aspartate peptide showed improved osteoblast attachment without inhibiting its antibacterial activity. An antibacterial polymer, Poly(N,N-dimethyl-N-(ethoxycarbonylmethyl)-N-(2'-(meth acry loyloxy)ethyl)-ammonium bromide), or (pCBMA-1 C2), was attached as a coating by a grafting technique known as surface-initiated atom transfer radical polymerization. The coating was shown to kill 99.9% E. coli bacteria followed by a release of the bacterial cells upon hydrolysis77. The coating shows promise, but bacteria more relevant to dental and craniofacial implants would need to be tested.

Polymers have also been combined with antibiotics to develop antimicrobial polymeric coatings. Al-Deyab, for example, soaked electrospun nylon-6/chitosan (nylon-6/Ch) nanofibers in an aqueous solution of glycidyltrimethylammonium chloride, an antibacterial agent, to make make nylon-6/N-[(2-hydroxy-3-trimethylammonium) propyl] chitosan chloride. The antibacterial efficacy of the fibers were tested against E. coli, P. aeruginosa and S. aureus and showed to negatively affect bacterial replication and induce cell damage. Significant zones of inhibition were also observed78. Grafted Allylamine, N-allylmethmylamine (AMA) and N,N-dimethylamine (DMAA) monomers with the addition of the antibiotic triclosan (TC) were tested for their antibacterial activity against S. aureus and E. coli. Bílek et al.79, found that the grafted AMA and DMAA with TC showed antibacterial activity against both bacterial strains and that they actually exhibited greater antibacterial activity against the gram-positive S. aureus. Zhao et al.80, developed a Poly(N-hydroxyethylacrylamide)/Salicylate hydrogel (polyHEAA/SA) coating that released antibacterial SA compounds resulting in a polymeric coating that resisted the attachment of S. epidermis and E. coli after 24 hours. The coating also inhibited bacterial growth which was determined by measuring the optical density of the bacteria.

Alternative antibacterial agents, such as zinc oxide (ZnO) and silver (Ag), have been combined with polymers as well. Liu and Kim81, added ZnO and Ag to genipin-crosslinked chitosan/poly(ethylene glycol) (GC/PEG) hydrogel matrix. The nanocomposites showed enhanced antibacterial activity against gram-negative E. coli and P. aeruginosa as well as gram-positive S. aureus and B. subtilis over GC/PEG alone. Adding ZnO increased the zone of inhibition of the copolymer, which was further enhanced by the addition of Ag81. Antibacterial metals have also been deposited directly onto implant surfaces. In one study, zirconium oxid (ZrO₂), ZrO₂-copper (ZrO₂-Cu) and ZrO₂-Ag coating were deposited onto pure-Ti substrates using pulsed unbalanced magnetron sputtering using high-purity Zr, Ag and Cu targets. When tested against S. aures and Actinobacillus actinomycetemcomitans, ZrO₂ surfaces showed less bacterial attachment and proliferation than the ZnO₂-Cu surfaces, which in turn, showed less bacterial proliferation than ZrO₂ coatings82. Titanium oxide (TiO₂) and zinc-doped TiO₂ (Ti(Zn)O₂) and ZnO coatings deposited by cathodic arc deposition have also been developed. The ZnO coatings had the least bacterial attach ment and the Ti(Zn)O₂ coating had less than the TiO₂ coating. However, when tested for osteoblast attachment, ZnO showed the least number of osteoblasts while TiO₂ showed the most.

The presence of infection is an important parameter that must be considered for nearly any reconstructive technique. Infections following traumatic craniofacial injuries are common, and although success rate of dental implants are high, biomaterials for the restoration of oral function are prone to biofilm formation, and failure is commonly associated with bacterial infections. The consequences of implant associated infection are significant and usually require revision surgery, with removal of the implant, prolonged antibiotic treatment, impaired oral function, and in extreme cases even death. Furthermore, bacterial colonization of dental implants can lead to inflammatory reactions which prevent or result in loss of osseointegration. For those reasons, various approaches to reduce the rate of infection have been investigated. Therefore, antibiotic delivery and antibacterial surfaces may thus be important aspects of tissue engineering strategies in the craniofacial complex18,50,83-86.

In this sense, the purpose of these bioactive surfaces would be to disrupt the metabolic machinery of the microbes or to prevent bacterial adhesion to the implant and, consequently, the development of biofilms. Different surface modification approaches and techniques have been developed for this purpose. Materials which promote the colonization of host tissue and suppress the colonization of bacterial species are often studied. These materials have many different mechanisms of action; some can interfere with bacterial adhesion by modifying surface energy, have surface immobilized molecules that are bactericidal, are photocatalytic, or more commonly, release metal ions or antibiotics18,83.

Many drugs and coatings have been developed to create antibacterial surfaces that either kill bacteria or prevent their attachment to implant surfaces. A handful of polymers are known to kill bacteria or prevent them from attaching and can therefore be used as coatings for antibacterial purposes87. There are also several low molecular weight molecules and inorganic ions that are known to also be antibacterial in solution and can either be released in a controlled manner or be grafted by covalent immobilization onto the implant surface87. The disadvantage to the controlled release method is that the duration of the antibacterial action is limited by loading and release kinetics88. Polymers can be coated onto medical implants by many methods including dip coating, spin coating, and layer-by-layer plasma polymerization which allows for a great variety of polymers to be applied onto material surfaces for the purposes of antibacterial action87. The majority of synthetic polymers, and the natural polymer chitosan, that have been reported to be antibacterial are cationic. The release of antibiotics from polymer coatings has also been extensively investigated. By releasing the drugs locally, as opposed to systemically, higher local doses can be administered without the risk of exceeding the systemic toxicity levels of the drug which could result in renal and liver complications89. When considering the release mechanism for these systems, it is important to consider the release kinetics of the drug; a fast release allows for a high dose but only for a short period of action whereas a slow release may not reach the required therapeutic levels and could also result in bacterial resistance87. According to Vasilev et al.87 the ideal release coating should provide a fast initial release within the first 6 hours, that will protect the site while the immune system is weakened, followed by a slow release. The use of a polymer matrix that degrades in the body allows for a combination delivery by diffusion and polymer matrix erosion90. Some common antibiotics used in these polymeric coatings are gentamicin, norfloxacin, cefazolin, amikacin, and vancomycin87. Nitric oxide has also been used in release systems as an antibacterial agent and has been shown to reduce the adhesion of P. aeruginosa, S. aureus, and S. epidermis87.

Although current implant materials for the reconstruction of craniofacial bone defects have shown favorable results in most craniofacial and dental applications, the presence of complications related with infection and poor osseointegration still represent a challenge in the biomedical field. Different clinical circumstances may present different challenges; however, the multitude of dissimilar solutions for the repair of bone defects that have been proposed during the last few decades highlights the fact that an ideal solution has yet to be defined. There remains a need to develop strategies that will further reduce implant failure while simultaneously addressing different problems and causes of complications in a cost-effective manner.

Biomaterials for craniofacial bone repair and dental applications such as titanium, HA, bioactive glass, biocompatible and biodegradable polymers, and composites have been widely studied and used in clinical applications. Ongoing developments indicate that the tissue engineering field is moving towards the development of biomaterials with improved surfaces that will stimulate bone formation and avoid infections though the incorporation of surface modification techniques and antibacterial coatings and agents, as well as the incorporation of growth factors, stem cells and other pharmacological drugs. Scientists are paying far closer attention to the biological interface in terms of both the specific cellular and vascular responses necessary for stable osseointegration as well as the unique microbial strains in the oral space and the necessary steps to prevent biofilm formation to avoid infection related complications. Through the application of principles of engineering and biology, optimization and further investigation of surface manipulation and coating techniques is necessary in order to develop craniofacial bone substitutes that will restore, maintain and improve tissue function while combating infection.