Overview of Craniofacial Bone Implant Materials and Techniques

Author: Aegaeon Galefang.
Reviewed by Lathreas and Zennith.
2024-01

1 – Introduction

It is important for many to attain freedom over their physical form by shaping it to fit their own identity. For some, this identity involves the expression of non-human characteristics. If we are to make morphological freedom through personalised bodily modifications a reality, it is important that we are able to modify our facial appearances as requested, even if such requests demand drastic anatomical changes, for example the creation of a muzzle.

In the case of the construction of muzzles, the most apparent modification needed to achieve such a structure in humans is to overcome the large differences in craniofacial skeletal structure. Notably, the maxilla and the mandible need to be significantly lengthened, to result in their increased protrusion to form the muzzle’s distinctive appearance, alongside other smaller cranial shape changes to achieve an overall more aesthetically pleasing appearance (Freedom of Form Foundation Inc, 2018). The most straightforward method to achieve such changes would be to utilise implants. Using techniques currently available in plastic surgery, it would be theoretically possible to surgically remove relevant craniofacial skeletal structures and replace them with artificial implants, similarly to how large sections of the bone are replaced in maxillary and mandibular reconstruction procedures (Iyer & Thankappan, 2014; Kumar et al., 2016; Lee, Cripps & Rodriguez, 2022).

However, it is unclear whether currently available artificial implant materials are appropriate for such drastic bodily modifications, and the construction of an artificial muzzle may require more advanced implants that are constructed using genetic and tissue engineering. This review thus will (a) summarise the advantages and disadvantages of the various artificial implant materials currently available for use in patients, (b) suggest potential methods for the construction of improved, biologically engineered implants, and (c) discuss potential roadblocks in the development of said engineered implants.

2 – Current day measures – Artificial implants

In the process of designing a permanent artificial implant, one of the most important factors to consider is the material used. This is because the material’s properties dictate whether the implant will be able to achieve optimal osseointegration, defined as a lack of progressive relative movement between the implant and the bone it is attached to (Mavrogenis et al., 2009; Parithimarkalaignan & Padmanabhan, 2013). Additionally, the material determines whether the implant will be durable enough to resist the pressure exerted on it during regular function, and if it will have any adverse reactions with the internal environment of the human body (Mantripragada et al., 2013). The ideal material should satisfy all the aforementioned conditions. However, of the various implant materials used today, none are perfect, as they each possess their own unique flaws.

Stainless steel, an alloy consisting mainly of nickel, chromium, and molybdenum, is mostly used in constructing internal fixation devices like pins, screws, and bolts, utilised to keep fractured bones stabilised and in alignment during and after the healing process. Its frequent usage is mainly attributed to its relatively low cost and high tensile strength in comparison to some of the other materials used in the field (Barber et al., 2021). Additionally, the alloy can also be made to be biologically well tolerated by modifying the surface of the material through electropolishing, silanisation and the covalent attachment of polymers, reducing protein adsorption to prevent unintended immune system responses (Kang & Lee, 2007). Yet, a flaw of stainless steel is its potential to corrode within the biological environment over time, hence the tendency to restrict their usage to only temporary implants (Gotman, 1997). The presence of chloride ions and reduced sulphur compounds in bodily fluids have shown to potentially cause pitting and crevice corrosion in stainless steel implants, leading to the release of toxic elements like nickel and chromium into the bodily environment (Manivasagam, Dhinasekaran & Rajamanickam, 2010; Wang et al., 2019). Additionally, stainless steel is a lot stiffer compared to both human cortical bone and other implant materials, and thus may not be appropriate for larger implants, especially if the implant’s purpose is to replace and mimic the functions of skeletal structures, as a difference in flexibility may negatively affect energy absorption during impact loading (Tranquilli Leali et al., 2009). Therefore, even if the corrosion of stainless steel implants and its impacts are fully mitigated, the inherent mechanical properties make it likely inadequate for maxilla and mandible modifications.

Cobalt-chrome alloys, as the name implies, are primarily composed of cobalt and chromium, but may also include other metals like molybdenum to enhance the alloy’s mechanical properties. These alloys are ubiquitously used in knee and hip replacement implants, as they are highly resistant to degradation from both wear and corrosion, much more than that of stainless steel (Marti, 2000; Song, Park & Moriwaki, 2010). One concern of note however is the cobalt-chromium metal ions and wear debris, potentially released from metal-on-metal or metal-on-polymer interactions within joints and any other internal fixation devices (Mantripragada et al., 2013). The cobalt-chromium nanoparticles have been linked to overall implant loosening (Chen et al., 2021), carcinogenicity (Polyzois et al., 2012), and unfavourable inflammatory response such as tissue necrosis, ulceration, and allergic reactions (Eltit, Wang & Wang, 2019). As mitigation, joint implants using these alloys are applied alongside a ultrahigh molecular weight polyethylene bearing between moving parts to further reduce corrosion and wear, decreasing the production of the nanoparticles (Kurtz et al., 1999). Yet, cobalt-chrome alloys would likely be unsuitable as the material for maxilla and mandible modifications because they possess a much higher elastic modulus than even stainless steel, meaning that these alloys are even stiffer than stainless steel (Davim, 2019), suggesting that they may cause stress shielding to be more likely, in which surrounding bone density decreases as a result of the mismatch in the impact loading forces between the implant and surrounding bones. Alongside the potential toxicity of the wear debris, cobalt-chrome alloys are thus unlikely to serve as the top contender as the main material for most implant designs in our context.

Titanium and its alloys are used in a large variety of implants, from craniofacial implants and dental implants to bone fixation materials like nails, screws, and nuts. These alloys have been used in many medical applications due to their desirable mechanical properties: low density, high tensile strength to density ratio, and good resistance to corrosion due to the formation of titanium oxide layer (Leventhal, 1951). Additionally, titanium and its various alloys are physically more flexible than stainless steel, as they more closely match the elasticity modulus of bone, allowing them to perform better than many alternatives in many orthopaedic purposes (Barber et al., 2021; Mantripragada et al., 2013). Despite the advantages, the main concern with the material is their long-term performance. Though titanium is one of the most corrosion-resistant materials, cases of degradation caused by stress and corrosion have been recorded, as well as reports of resulting implant malfunction (Prestat & Thierry, 2021). Additionally, titanium alloys have a high coefficient of friction, causing them to undergo severe wear when it comes into contact with other metal surfaces (Lee, H. et al., 2018), which in turn leads to the eventual loosening of the implant through inflammatory responses. There are also concerns that the resulting debris from corrosion of wear would also release metals like aluminium and vanadium from certain alloys, which have been associated with health problems like neuropathy and bone softening (Gomes et al., 2011; Lin, Ju & Chern Lin, 2005). Fortunately, recent research on surface modifications have shown promise to decrease this friction coefficient and thus reduce the rate of wear of titanium alloy implants, allowing their life expectancy to become considerably extended (Lee, H. et al., 2018; Wypych et al., 2018). Given that the downsides are actively being mitigated at present, titanium and its alloys are thus very promising candidates as the main material for the maxilla and mandible implants, though their non-permanent lifespan due to regular corrosion and wear still leaves room for improvement.

Porous tantalum alloys have recently been identified as a potential alternative to titanium in prosthetic development. Not only does it possess similar favourable mechanical properties to that of titanium, it also has the added benefit of being a highly porous structure with an interconnected inner space which allows for bone-tissue ingrowth, permitting for a much higher degree of osseointegration than other implants (Huang, Pan & Qiu, 2021; Piglionico et al., 2020; Ren et al., 2015). Yet, as a relatively newly developed material, the full impact of tantalum implants is still uncertain. The material has seen some clinical applications, but there is a lack of long-term follow-ups on patients with these implants to verify their practicability (Huang, Pan & Qiu, 2021). Though given that there have not been any studies that reveal any early failures in tantalum’s clinical performance, it is likely that porous tantalum alloys are a promising candidate for future prosthetics.

Inert ceramics like aluminium oxide and zirconium oxide are an emerging alternative to conventional metal-based implant systems, but are already widely used in hip and knee replacements, maxillofacial reconstructions, middle ear bone replacement, and even artificial cornea replacements. Their broad applications are generally attributed to their high compressive strength, and their strong resistance to corrosion and wear (Mantripragada et al., 2013). The reason inert ceramics have not seen much popularity however is due to their association with catastrophic failures. Inert ceramics generally are more brittle than alternatives and have lower fracture strength and flexural strength. Additionally, their elastic modulus is similar to that of cobalt-chromium alloys, and thus are stiffer than cortical bone. They may be strong under forces of compression, but they have been shown to be likely to fracture easily in cases where the components suffer under forces that cause bending or flexing (Onur Kocadal et al., 2019; Rodriguez & Cooper, 2013). Recent work has shown that it is possible to create inert ceramics that are more resistant to cracking through reinforcement in combination with other materials, thus mitigating the largest issue with inert ceramics (de Freitas et al., 2020; Rodriguez & Cooper, 2013). However, this only further serves to make the material even more stiff, increasing their elastic modulus. Moreover, their long-term performance and success have not yet been sufficiently proven, as the field lacks clinical studies that study the long-term effects on patients with these implants. Ultimately, given their inflexibility, inert ceramics suffer from the same issue as stainless steel, where using it to replace and mimic the functions of the mandible and maxilla in implants as a large implant is inadequate, as it affects energy absorption during impact loading.

Non-oxide ceramics, which include materials such as silicon nitride and silicon carbide described as glass-ceramic composites, are another category of materials that are considered to be a rapidly developing contender in the field of artificial implants. The reason behind them showing great promise in prosthetics is due to their similarities in mechanical properties to that of the inert ceramics; they possess high compressive strength, and good wear and corrosion resistance. Furthermore, unlike inert ceramics, non-oxide ceramics have good fracture resistance, allowing them to potentially be more suitable in high-load medical applications (Du et al., 2022; Paione & Baino, 2023). Unfortunately, they also share a lot of vulnerabilities with inert ceramics that make them unsuitable for maxilla and mandible replacements. Non-oxide ceramics also have low flexibility, which may cause issues with impact loading as discussed previously. Moreover, the surface of non-oxide ceramics have been observed to chip off over time, resulting in a significant increase in wear and thus in vivo failure (Mazzocchi & Bellosi, 2008; Paione & Baino, 2023). Thus, it would be reasonable to conclude that non-oxide ceramics are unlikely to serve as the main component for the proposed implants.

Table 1. Mechanical properties of metallic and ceramic materials used for artificial implants. Modified from (Mantripragada et al., 2013) and (Thamaraiselvi & Rajeswari, 2004).

From looking at the extensive list of materials that are commercially accessible now and in the near future, it would be reasonable to conclude that if we are to carry out extensive modifications to the maxilla and mandible, titanium alloys or porous tantalum alloys would be the most suitable to use to construct the prosthetic. However, it cannot be denied that regardless of material, the main concern with artificial, prosthetic implants is that they may be more vulnerable to the effects of corrosion and wear, as they are not actively maintained, and in some cases even attacked by the patient’s bodily functions.

For these reasons, biological implants may be preferable for permanent modifications; successfully integrated bone grafts, made using the patient’s own cells, would function identically as native bone within patients, and thus have the intrinsic ability to repair, remodel, and regenerate.

3 – Cutting-edge possibilities – Engineered bone scaffolds and grafts

Stem cell and tissue engineering approaches have already made major strides towards growing biological implants that are able to mimic the form and functions of native bone. With respect to bone defects, these efforts have provided the possibility of using the patient’s own cells in combination with a scaffold for both bone repair and bone graft development. Similarly, it should also be possible to develop idealised maxillofacial implants using the same cutting-edge methods.

Scaffolds are a common starting point for projects in tissue engineering. In essence, they are mechanical constructs that act as carriers for cells and growth factors, to serve as a template for tissue regeneration, guiding the growth of new tissue. For reconstructive procedures, these scaffolds can be directly implanted into the site of injury, where the regeneration of tissues can be induced in vivo using the body’s own systems. Alternatively, the scaffolds can also be cultured in vitro to synthesise tissues to be implanted at a later date. Regardless of their utilisation, the scaffold needs to be able to support normal cell function and migration without eliciting severe immune responses, to be safely biodegradable within the internal environment as to allow cells to replace the scaffold, and finally to have adequate mechanical strength and porosity to allow for scaffold function and cell survivability (O Brien, 2011). All these criteria are why a large focus on scaffold research is going into the biomaterial from which the scaffold is fabricated.

3.1 – Scaffold Materials

The materials used in scaffolds must be biocompatible, biodegradable, as well as possess good mechanical properties to withstand the shear forces necessary for osteoconduction and thus maturation of the scaffold. Generally, the biomaterials in use at the time of writing that fit the criteria can be divided into natural and synthetic polymers, bioactive ceramics, and composites of the two (Chocholata, Kulda & Babuska, 2019).

Natural polymers such as polysaccharides and proteins are great at allowing cells to adhere to the scaffold structure and thus stimulating their growth, however they struggle with weak mechanical properties and unintended immunogenic responses. As an example, chitosan is widely used in scaffolds because it supports cell attachment and migration, whilst also being non-toxic and biocompatible (Nadia Kartikasari et al., 2016). Specifically in bone tissue engineering, it has been cited to promote the growth of osteoblasts and the mineralization of the scaffold as a whole (Asti & Gioglio, 2014). However, chitosan scaffolds have significantly lower compressive strength and modulus than natural bone tissue (Jana et al., 2012). Even as the scaffold matures and begins to integrate with the surrounding bone, sections of the scaffold that have not yet been replaced by native bone remain prone to fragile failure if under stress, and thus the material should not be used in vivo in load-bearing applications. Another polymer frequently used in bone scaffolds is collagen. As it acts as a major component in bone, it is extremely favourable as a candidate material. Yet, similar to chitosan, it also has low mechanical strength. Additionally, it also has the drawback of being extremely susceptible to biodegradation as it is a target for multiple enzymes naturally expressed by cells, to the point where the scaffold degrades before it has served its purpose (Rico-Llanos et al., 2021).

Synthetic polymers are based on polyesters, and their main advantage over their natural counterparts are that they can be produced with customisable structures, thus allowing the fine control over their degradation rate (O Brien, 2011). However, they are not as biocompatible as alternatives, and may suffer from adverse effects within the body environment. This is best illustrated with the example of the poly-α-hydroxy acid polymers: poly-lactic acid (PLA), poly-glycolic acid (PGA), and poly-lactide-co-glycolide (PLGA). They are normally used in resorbable surgical sutures, and are now also widely used for 3D scaffolds in bone tissue engineering, as these polymers are chemically versatile, allowing for ease of manipulation when designing scaffolds, as well as allowing for the modulation of their degradation rates. Yet, their issue is that they undergo a bulk erosion process, in which degradation occurs throughout the whole polymer simultaneously. This not only potentially causes scaffolds to fail prematurely, the acidic degradation products have also been recorded to cause strong inflammatory responses. There also have been issues with seeding and delivering cells into the scaffolds due to the high hydrophobicity of the surface of PLA and PLGA (Asti & Gioglio, 2014).

Bioactive ceramics consist of mechanically strong biomaterials like hydroxyapaitite (HAP), tricalcium phosphate (TCP), and certain compositions of bioactive glasses. They are also used frequently in bone tissue engineering, as well as orthopaedics and dentistry. As they are chemically similar to native mineralized bone, they show excellent corrosion resistance, high compression resistance, and allow for good differentiation and proliferation of introduced osteoblasts. Their drawbacks however, are that they are extremely brittle, having low fracture strength, and low mechanical reliability (Asti & Gioglio, 2014), which has caused issues with the mechanical loading and degradation rate control of these scaffolds (Berthiaume, Maguire & Yarmush, 2011).

(A table providing a brief overview of the advantages and disadvantages of specific materials used in tissue engineering is available at https://www.ncbi.nlm.nih.gov/pmc/articles/PMC6416573/)

As these materials each have their own flaws individually, the best scaffolds produced at the time of writing, which are most able to mimic the properties of real bones, are material composites; scaffolds made from a combination of ceramics and polymers, or of both synthetic and natural polymers, outperform those made from a single biomaterial, as they are able to capitalise on the advantages of each component. For example, adding PLGA to bioactive ceramic materials has shown to improve their fracture strength (Zhang et al., 2014), making this combination a much more preferable option than either or independently.

It is important to note that presently, many of the aforementioned types of scaffolds are further structured into hydrogels. Essentially, they are 3D flexible networks of polymers linked by a bridging agent that swell up in water or biological fluid, as a result of their interconnected porosity (Varaprasad et al., 2017). Scaffold materials are frequently structured as hydrogels in tissue engineering as this porosity allows for multiple benefits: the growth and proliferation of cells, the distribution of drugs and chemicals, and the diffusion of nutrients and waste products (Chocholata, Kulda & Babuska, 2019), all the while retaining good mechanical properties and biocompatibility.

3.2 – Scaffold Fabrication

Beyond the biomaterial the scaffold is constructed from, a good scaffold is also one that has a large network of interconnected pores. The optimal size of pores and degree of porosity is still in discussion, as the answer varies based on the purpose of the scaffold, and because it has to balance sufficient nutrient flow without sacrificing mechanical stability. Though, it is generally understood that pore size should be at least 100 microns to allow the movement of nutrients, and porosity should be above 75% to improve bone growth, nutrient transportation, and degradation rate of the scaffold (Abbasi et al., 2020; Murphy & O’Brien, 2010; Tang, Abdul Kadir & Ngadiman, 2020). Regardless, creating such pores without losing the structural integrity of the biomaterial is difficult, and thus requires special manufacturing techniques that not only preserve the scaffold’s load-bearing strength, but also allows for accurate control of the scaffold’s exterior shape (Kim, Shin & Lim, 2012). This is especially true when creating craniofacial scaffolds, as they have relatively more complex 3D structures compared to other regions of the body. Several methods of scaffold fabrication of note that have been able to achieve this are salt leaching, gas foaming, electrospinning, and 3D bioprinting.

Salt leaching is one of the more common methods utilised in industry at present, as the production process is simple and straightforward, avoiding the use of any computer-controlled procedures, whilst still allowing the pore size and porosity of the scaffold to be controllable. The technique simply involves dissolving the biomaterial polymer of use in an organic solvent, mixing it with a water-soluble porogen such as sodium chloride, then casting this mixture into a mould (Kim, Shin & Lim, 2012). Once the organic solvent evaporates, the scaffold is placed into water to leach the porogen away from the structure. In this method, we can control the amount of porosity by changing the amount of porogen in the mixture, and the sizes of the resulting pores depends on the size of the crystals that the porogen forms during solvent evaporation (Chocholata, Kulda & Babuska, 2019). However, the resulting scaffold is difficult to utilise in practice, as the organic solvents used are toxic, and any that remain on the scaffold could cause denaturation of the scaffold itself as well as a decrease in the activity of molecules through the structure (Rezwan et al., 2006). Additionally, the production process makes it impossible to add in additional pharmacological agents, thus we will be left unable to adjust the behaviour of tissues in the scaffold (Kim, Shin & Lim, 2012).

Gas foaming is used in some contexts as an improvement over salt leaching, as it has the advantage of avoiding usage of toxic organic solvents. As opposed to dissolving polymers within organic solvents, this method instead has polymers dissolved within an aqueous phase, then dissolved in dense carbon dioxide. Upon gas expansion, numerous pores would be successfully created within the scaffold structure (Ji et al., 2011). Yet, there are still many issues with this method. Notably, this technique can only be used with polymers that have a high solubility in carbon dioxide, which means hydrophilic and glassy polymers are incompatible with the method. This issue is significant especially if we intend to utilise polymers in the form of hydrogels as their function relies on hydrophilic functional groups attached to the polymeric backbone (Ahmed, 2015).

Electrospinning is one of many innovative techniques in recent development, in an attempt to create better ways to manufacture custom scaffolds. The method involves having a solution of a polymer injected with an electrical potential to create an overall charge imbalance, which then allows for a stable, steady deposition of electrospun fibres on a wide variety of substrates to result in a non-woven nanofibre scaffold (Venugopal et al., 2007; Xue et al., 2019). Compared to the two other methods more widely used in industry, electrospinning is ultimately a lot more versatile, as a larger variety of biopolymers are compatible with the method (Agarwal, Wendorff & Greiner, 2008). Additionally, it is also possible to control the thickness of individual fibres and their orientation, allowing for the creation of scaffolds with unique features at the nanoscale level (Xue et al., 2019), which may be useful in scaffold integration in vivo. Yet, the biggest obstacle with the wider use of electrospinning in tissue engineering is that while we are able to adjust the properties of the individual fibres, there is a lack of precision in the fibre deposition. Most successfully created scaffolds from electrospinning are either 2D structures or uniform 3D geometric shapes. Even then, there is still an inadequate control of the porosity of the created scaffolds (Flores-Rojas et al., 2023), which poses difficulties in terms of cell and nutrient flow into the structure.

3D bioprinting is arguably the most promising upcoming method in scaffold fabrication, aiming to create complex constructs through placing living cells and biomaterials layer by layer, to result in a 3D cell-laden scaffold (Zhang, J. et al., 2021) that ideally performs better than scaffolds constructed purely out of biomaterials. Generally, the process of creating these cell-laden scaffolds begins with obtaining both a cell culture and a printing model of the scaffold structure. Then, a bioink is created by mixing the prepared cell culture with hydrogels, along with other chemicals like growth factors if necessary (Yazdanpanah et al., 2022). This bioink is then used to create the scaffolds through 3D bioprinting. Specifically, there are 3 types of bioprinting techniques: inkjet, laser, and extrusion-based bioprinting, each with their own specific strengths, weaknesses, and limitations (Genova et al., 2020). Nevertheless, based on the provided 3D tissue model, a three-axis mechanical platform controls the movements of extruders in the required algorithm and shape, using the bioink mixture to create in the desired cell-laden tissue scaffold. Theoretically, these new scaffolds are overall much better than those created through traditional methods; the fact that a scaffold is printed according to a 3D model allows for extremely fine control over the porosity of the structure, and during testing these 3D-printed scaffolds have shown to be able to withstand physiologically relevant loads in rabbits (Yao et al., 2015). The only concern is that there has yet to be a conclusive test of the scaffolds themselves in humans (A. Maresca et al., 2023), hence the research has not shown thus far whether the same benefits can be claimed in clinical settings.

It is worth mentioning that under certain situations, there are physical limitations to the size of the scaffold that can be created using 3D bioprinting. Bones are metabolically active tissues, and thus require constant nutrient supply and waste removal to remain viable. Scaffolds that are too large in size cannot survive post-implantation in situ as they cannot be sustained by passive diffusion alone. Thus, researchers are actively trying to find solutions to induce vascularisation within scaffolds. One of the proposed solutions is to 3D print scaffolds that have vessels embedded within (Grigoryan et al., 2019). The issue with this method however, is that the process is extremely slow, requiring multiple days of printing, drastically increasing the likelihood of cells within the bioink dying before the scaffold can be completed (Zhang, J. et al., 2021). Fortunately, other methods such as adding in a maturation step using bioreactors are able to mitigate the concern with scaffold vascularisation.

Of the methods mentioned, 3D bioprinting still seems to be the most ideal, as the created scaffolds have shown to be capable and adequate when testing within models. Additionally, the accurate control it provides over the scaffold’s shape and structure makes it especially suitable for craniofacial implants as their skeletal structure is more complex than those elsewhere in the body. However, it is worth noting that this technique is still in its infancy, and would require a lot of testing to ensure its success and safety in vivo. Ultimately, the decision of which method to manufacture craniofacial implants need not be made immediately, but is without doubt a critical one that requires further investigation in the future.

3.3 – Scaffold Maturation

Once a scaffold has been manufactured, scaffold maturation can then be carried out if desired, transforming the artificially created construct into usable human tissues that are more suitable for implantation. For bone tissue generation, this will involve encouraging mesenchymal stem cells (MSCs) to differentiate into osteoblasts, the major cells responsible for laying down extracellular matrix and mineralizing the bone, by utilising growth factors and mechanical forces (Watson & Mikos, 2023). Typically, this is done in vitro within bioreactors, though it is also possible to instead leverage the body’s natural regenerative capacity to create bone tissues and mature these scaffolds in vivo, or even in situ.

In vitro maturation of scaffolds is currently the most established method of cultivating scaffolds, as the method allows for a tightly controlled and reproducible environment to generate tissues. This is achieved by incubating these scaffolds within bioreactors, which are devices that use mechanical means to influence biological processes (Plunkett & O’Brien, 2011). Specifically, they partially imitate the functions of the circulatory system, providing an environment where nutrient transfer and waste removal for the cells can occur, all the while allowing for the control of mechanical or shear forces applied to the cells, further promoting osteoblast differentiation in combination with several growth factors in the culture medium. There are many different types of bioreactor systems, optimised for the specific tissue engineered and their functional biomechanical environment (Zhang, J. et al., 2021), though the operating principle is the same. After several weeks in culture, the scaffold can then be harvested, and a surgery can be performed to implant the cell-laden mineralized bony scaffold (Watson & Mikos, 2023). It has already been shown that it is possible to reconstruct the entire facial bone in a skeletally mature large animal using this exact method, successfully reconstructing a large, anatomically precise load-bearing craniofacial defect using tissue engineering that holds up to implantation (Bhumiratana et al., 2016).

A major drawback however is that while these engineered scaffolds are porous by design as a result of the media flow within the in vitro bioreactors, they lack functional vasculature. This means that the implant constructs can potentially encounter premature failure as the cells within the scaffold are unable to efficiently carry out nutrient and waste exchange (Kaully et al., 2009). In many cases, the centre of the mineralised tissue within the implant is reliant on convective flow from the extravascular fluid once implanted until capillaries can grow into the scaffold, which causes the implant to only incorporate completely after many years (Roberts & Rosenbaum, 2012). Fortunately, current research has proposed many viable solutions to have the production pipeline to allow the in vitro maturation to also incorporate vasculature. For example, it has been discovered that co-transplantation of a combination of certain MSC sub-populations has resulted in the growth of de novo blood vessels in tissues (Zhao et al., 2023). Theoretically, it should therefore also be possible to utilise a coculture of these MSCs to induce the growth of blood vessels in engineered bone grafts, thus creating bone grafts that would be able to integrate more efficiently in patients.

Alternatively, instead of recreating the oxygen delivery, waste removal, and mechanical cues that occur within the body, some researchers have been more interested in using the body itself as an in vivobioreactor.” Utilising this tissue generation strategy, the scaffold is implanted at a distal site away from the maxilla or mandible. Over time, as cells migrate into the scaffold, and as capillaries form to provide nutrients and gas exchange, mineralised tissue is then generated within this scaffold, which can then be harvested with adjacent vasculature and transferred to the maxilla or mandible (Watson & Mikos, 2023). This has been performed successfully before in humans, with craniofacial bone tissue implants being created from scaffolds implanted in the rib periosteum (Cheng et al., 2006), the muscle (Warnke et al., 2006), and the gastrocolic omentum (Wiltfang et al., 2016). However, there are currently several concerns with this method of maturing scaffolds. Firstly, this strategy heavily relies on the capacity of the host for regeneration. With in vitro bioreactors, we can constantly supply the MSCs with potent growth factors. The same cannot be done inside of the patient’s body. Health conditions like old age, obesity, and diabetes can also severely affect the patient’s ability to regenerate (Alessio et al., 2020; Kim, H. et al., 2015; Yang, 2018), which would make this method inadequate for those with relevant underlying conditions. Additionally, given the nature of the method, the maturation process cannot be easily observed; it may be difficult to appreciate the mineralisation within the relevant region using x-rays, and repeat CT testing could lead to excessive radiation exposure for the patient (Watson & Mikos, 2023).

Similarly, it would also be possible to mature these scaffolds in situ, directly implanting them at the maxilla and mandible and having them develop over time. The overall process is similar to developing them elsewhere in the body, also sharing similar advantages and disadvantages (Cao et al., 2020). However, for our purposes, this strategy would additionally cause the patient to have experiences akin to that of a person with severe craniofacial bone defects. The scaffold before maturation is merely an imitation of craniofacial skeletal structures and would likely be unable to completely fulfil the physiological responsibilities of a fully developed implant, likely causing a lack of craniofacial usage for at least several weeks to several months.

Consolidating this information, it would be theoretically possible to generate individual-specific biological implants using techniques available in academia. Most likely, selecting the most appropriate material and the more feasible techniques, this implant would involve constructing an implant using a hydrogel consisting of a material composite, fabricated using 3D bioprinting, and cultivated within an in vitro bioreactor. To confirm this perspective and bring it beyond the realm of speculation, additional research efforts involving practical work are still necessary; for example, it is still unclear which material composite and growth factor combinations are best to stimulate osteogenesis, whether scaffolds produced through 3D bioprinting can be successfully matured within in vitro bioreactors, or if we are able to introduce vasculature into our matured implants. Additionally, even once it becomes possible to fabricate these biological implants, there are still concerns that need to be addressed before the adoption within clinical settings.

4 – Future concerns – potential roadblocks

We recognize that even when it is possible to create biological implants that completely mimic the properties of native bone, there are several concerns that would require additional research to alleviate and mitigate if necessary. 

For instance, the mechanical loading forces acting upon the implants warrants further investigation. It is well established that mechanical loading is a major regulator of bone mass and geometry, where increased mechanical forces promote bone formation (Bergmann et al., 2011; Hart et al., 2017), and decrease in loading is always associated with bone loss and skeletal fragility due to bone resorption (Uda et al., 2017)

For all bones, the amount of force acting upon it depends on several factors, including the specific shape and mass of bone, as well as the surrounding muscle structures. To ensure the long-term survivability of our modified craniofacial implant, it is necessary to ensure it undergoes the appropriate amount of loading to maintain its structure without further surgical intervention. Specifically, this would involve quantifying the difference in impact loading forces between human jawbones and animal jawbones in the routine mastication of food and other expected behaviours, investigating the level of such forces necessary for the lengthened mandible and maxilla to maintain their shape, then finally bridging the gap between human and animal jawbones, if any, by altering bone or muscle. 

At the time of writing, while there is literature detailing the changes when there are increased forces acting on human jawbones (Inoue et al., 2019), it is still unclear if the native craniofacial environment in humans can support a lengthened maxilla and mandible in the long run. Given that there are differences in the bite force between canines and humans (Kim, S. E. et al., 2018; Koc, Dogan & Bek, 2010; Takaki, Vieira & Bommarito, 2014), it is likely that the mechanical loading forces acting upon the implant will strongly influence its design.

Another concern is the integration of the dental environment and other relevant structures. A successful craniofacial implant should ensure quality of life of the patient, and a healthy dental environment is therefore crucial. The current strategy of replacing the maxilla and mandible would undoubtedly require additional modifications to the dentition of the patient, as it would both be impossible and undesirable to modify the length of craniofacial features without affecting the dental structures. At a cursory glance, there seems to be several viable strategies to either incorporate or replace the dental structures onto our prosthetic; the options range from simply replacing the entire environment with a completely artificial set of teeth to be combined with our artificial implant (Li, Wu & Lin, 2020), transplanting teeth onto their new positions within the mouth cavity during or after the surgery for the craniofacial implant (Clokie, Yau & Chano, 2001), to applying 3D bioprinting to dental tissue engineering and creating a new individualised set of teeth and all relevant structures (Ostrovidov et al., 2023). Of these methods, 3D bioprinting seems the most promising, though confirmation of this would necessitate additional review, as it is out of scope for this review. Regardless, craniofacial modifications, from the perspective of a patient, cannot be deemed successful without consideration of the dental environment.

Furthermore, the procedure to replace both the maxilla and mandible is inherently risky, as there are a wide variety of surgical complications associated with this type of surgery. Beyond the risks common to invasive procedures, there is an increased risk of haemorrhaging and neurological injuries occurring due to the large number of vessels and nerves in the craniofacial region. Additionally, there have also been many postoperative complications reported to be associated with maxillofacial surgeries that negatively impact a patient’s quality of life, including respiratory insufficiency and obstructive sleep apnea (Kim, Y., 2017; Liu & Costello, 2020). While the risk of these incidents can be decreased with proper technique exercised by trained and experienced surgeons, it is worth noting that it may become necessary to investigate alternative methods to surgery if it is found that the risks are unacceptable for certain patients, especially those with pre-existing conditions.

5 – Conclusion

With respect to the goal of achieving substantial craniofacial structure modification through implants, though it is possible to construct artificial prosthetics out of commercially accessible materials, it would more preferable to utilise biological implants generated from patient’s stem cells, as they are more suited for permanent fixtures; they are able to function in a manner identical to native bone, having the intrinsic ability to repair, remodel, and regenerate, all the while avoiding immune rejection. 

Using the tissue engineering techniques available within academia, such a biological implant is theoretically possible in humans. Through the fabrication and maturation of scaffolds, researchers have been able to create personalised bone tissue that successfully integrates with native tissue in animal models, and certain companies have even begun testing the implantation of 3D-bioprinted living tissues in clinical trials, showing that it is technically feasible to use the similar methods to generate ideal maxilla and mandible implants in humans. However, this is ultimately still conceptual, as while there are examples of clinical trials ongoing in other tissue systems, the technique is still in development with respect to bones. Investigating the impact these specific techniques have on a patient’s body and their immune system is still necessary. Beyond this, there are several concerns towards substantial craniofacial modifications in general that deserve to be investigated as well before these implants can be applied clinically. 

Overall, this review shows that while the path towards an artificial muzzle may encompass numerous challenges and problems, there is plenty of promise in the field of tissue engineering, implying a bright future for the freedom of form.

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