*3.1. Particulate Bioceramics*

Hydroxyapatite (HA) and β-tricalcium phosphate (TCP), have undergone intensive research for their use in bone applications [60]. HA (Ca10(PO4)6(OH)2) has a molar ratio of 1.67 Ca/P and is a bioactive ceramic that is typically used in coatings and can be incorporated into materials in particulate form. HA has been shown to improve the biocompatibility of biomaterials due to its similarity in structure and composition to bone and enamel [61]. It is osteoconductive and osteoinductive [60]. Tricalcium phosphate has three polymorphs; α-TCP, β-TCP, and α-TCP. In particular, β-TCP displays excellent biocompatibility, bioactivity, and bio-resorbability. Although HA has greater mechanical strength than β-TCP, the latter is more resorbable. This encourages faster growth of new bone around the implant.

Polyesters, such as PLLA, PCL, and PHBV can take up to 2 years to degrade, and ceramic incorporation has been found to decrease this. HA enhances the degradation rate of PCL, with mass loss in vitro increasing with HA content. In addition, the compressive yield strength and modulus of the material also increased almost linearly with HA content [62]. Up to 3 wt% nano-hydroxyapatite (n-HA) incorporated into a poly-L-lactic-co-glycolic acid (PLGA) matrix has been found to significantly increase the mechanical properties of the polymer [63,64]. It was concluded that a higher content of n-HA in the composite promoting enhanced crystallisation, but also caused greater agglomeration in the PLGA matrix, which resulted in a decrease in mechanical properties. The addition of 3 wt% n-HA enhanced the degradation performance of the material, making the material promising for use in clinical applications in comparison to pure PLGA in bone fracture internal fixation materials. Porous polymer-hydroxyapatite scaffolds for femur fracture treatment produced by 3D printing have presented promising results. Hydroxyapatite nanoparticles were used to enhance PLA polymer and subsequently 3D-printed into porous cylindrical structures. Comparing 5% to 25% HA content in PLA, the compressive modulus and elastic modulus increased by ~38% and ~92%, respectively, a 11% decrease in porosity of the scaffolds was also seen. These results illustrate the reinforcing behaviour of HA and the mechanical properties being a function of porosity [65]. The mechanical and thermal properties of injection-moulded poly(3-hydroxybutyrate-co-3hydroxyhexanoate) (P(3HBco-3HHx))/hydroxyapatite nanoparticles (n-HA) parts for use in bone tissue engineering and bone fixation were assessed [66]. The addition of 20% nHA improved the mechanical properties; specifically, the tensile moduli and flexural moduli by approximately 60%. A larger content of n-HA leads to agglomeration, with the ductility, toughness, and thermal

stability of the material declining. The addition of nHA up to 10% most closely resembled natural bone in terms of strength and ductility and demonstrated controlled agglomeration. Considering its biocompatibility, the copolymer composite can be appropriate as a bioresorbable implant for low-stress bone fixation sites. An alternative way to simple melt mixing of ceramics with polyesters is covalently linking the two. A novel study surface first grafted nano-HA to PLLA. This nano-HA/PLLA composite was then blended with PLLA. Surface-grafting HA to PLLA led to an improvement in mechanical properties, compared to similar non-grafted PLLA/HA composites due to the grafting providing strong linkages between the HA particles and the PLLA matrix [67]. Further works by the researchers saw surface modification of carbonate hydroxyapatite particles with PLGA, with processing as previously described [68]. The PLGA/g-CHAP nanocomposites displayed improved mechanical properties in comparison to unmodified PLGA. This was attributed to the strong interfacial bonding between PLGA and g-CHAP particles. At 2% of g-CHAP content, the fracture strain was increased from ~5% (for neat PLGA) to 20%. When g-CHAP content was between 2 and 15%, the composites showed enhanced tensile strength and fracture strain. The tensile strength decreased linearly with filler content beyond 20%.

In vitro, TCP incorporation into PLLA showed good biocompatibility and a rate of degradation consistent with bone healing [69]. In particular, material strength was maintained for 16 weeks, after which a decline was observed with no measurable strength of the material at 40 weeks. There was a decrease in tensile strength of the material over the investigated period. In vivo, these results were seen on a shorter time scale. The addition of 10% volume TCP filler to PLLA gave a biocomposite material with an extended strength and molecular weight retention period, both in vitro and in vivo [69]. The same experiments conducted on PHBV showed slower degradation rates, which was attributed to the hydrophobic nature of PHBV. Ternary blends of PCL/PGA/tricalcium phosphate (TCP) (80/10/10 and 70/10/20) displayed success for use in low load bearing applications, such as maxillofacial surgery [70]. Adhesion strength of the materials was tested using a previously established protocol, where investigated biomaterials were melted and applied to two bone sections before curing. The mechanical properties of the set fixation was then tested through compressive and tensile force [71]. The blends retain the adhesive strength of PCL whist having improved hydrophilicity. A higher incorporation of TCP also results in enhanced degradation and support for osteoblast growth. PLLA was strengthened with a poly(ε-caprolactone-co-L-lactide) copolyester [18]. A higher quantity of copolyester (PCL/LLA) in the blend provided a greater elongation at break with a concurrent decrease in the Young's modulus and strength. Use of a chain extender (Joncryl® ADR 4368, BASF, Thailand) enhanced phase compatibility of PLLA and the interspersed copolyester phase.

## *3.2. Glass Fibres*

Bioactive glasses were first developed in 1969 and represent a group of materials that have the capacity to bond with bone in physiological environments. They have been greatly studied due to their desirable characteristics, which include biocompatibility, degradation, and mechanical strength [72]. By changing the composition of glass, it is possible to obtain a range of mechanical properties and high controllability over the resorption and ion dissolution that resemble and complement the mineral content of bone. For instance, it is possible to decrease degradation rates in a phosphate glass composition by addition of hydration-resistant metal oxides, such as Al2O3, Fe2O3, and TiO2 [73,74]. The use of high aspect ratio fibrous bioactive glass structures, as reinforcing phases, has produced fully bioresorbable polyester composites with higher mechanical properties compared to composites with particulate reinforcement [74]; however, studies have shown variable results, with some showing a rapid loss of mechanical properties with degradation [75]. The use of fibres in bioresorbable composites has also been associated with an immediate osteoinductive effect owed to the presence of the filler in the outer surface of the polyester implant [76]. Glass fibre reinforced polyester composites have also been found to act

protectively to the polyester matrix against gamma radiation from deterioration of its mechanical properties [74].

The properties of the composite for a given polyester matrix and glass fibre reinforcement depend on the geometry (i.e., aspect ratio) of the filler, distribution within the matrix, volume fraction, and surface area of the composite and strength of the fibre–matrix interface. Therefore, fibre and composite manufacturing methods significantly influence the properties of the composites. Glass fibres are typically fabricated into continuous fibres via melt spinning or preform drawing manufacturing processes, with single filament production being more common [77]. The mechanical, thermal, and degradation properties of glass fibres vary from those of bulk glasses depending on processing temperatures, drawing speed and ratio, and viscosity of molten glass, which ultimately determine the fibre diameter. A smaller fibre diameter leads to increased dissolution times, a fact that is associated with the active surface area of the filler [74].

The fibres can be appropriately aligned within the polyester matrix to create long-fibre composites, or chopped and dispersed within the matrix to create short fibre composites. The use of long fibre reinforcement is associated with improved mechanical properties in the composites when measured along the axis of the fibre. During manufacturing, the fibrous preform has to be positioned within a mould cavity and infiltrated with polyester matrix, which might require complex tooling. Short fibre composites are easier to process with extrusion/injection or casting processes, with the filler is melt-mixed or dissolved with the polyester matrix prior to processing. Solvent casting methods have also been used to create short glass fibre polyester composites. The method of incorporation of glass fibres into the polyester matrix plays an important role in the obtained mechanical and degradation properties.

Long-fibre unidirectional (UD) woven mats and randomly orientated short fibre nonwoven mats (RM) of iron doped glass phosphate fibres as reinforcements in a PLA matrix have produced bioresorbable composites with enhanced mechanical properties [78]. UD matt reinforced composites with filler volume fraction of 20% revealed a faster depletion of mechanical properties during degradation compared to the randomly oriented short fibre matt composites of 30–40% volume fraction. A maximum modulus and strength for the RM and UD were 10 GPa/120 MPa and 11.5 GPa/130 MPa, respectively, falling short of the ideal properties for cortical bone. Increased concentration of fibres, despite allowing for enhancement of mechanical properties of the composite material in comparison to polymer alone, also lead to an increased degradation time. Further work demonstrated that ~30% fibre volume fraction of unidirectionally and randomly aligned fibres into PLA rods imparted the composite with a higher initial modulus, which succumbed to degradation faster than PLA alone [79]. The RM and UD-filled PLLA was manufactured into intramedullary nails and the mechanical properties were assessed [80]. The composite reinforced with unidirectionally aligned fibres provided enhanced mechanical properties compared to pure PLA rod. The method of processing via forging at 100 ◦C also improved the properties of the PLA matrices by influencing chain orientation. This method has a similar effect to the drawing process at low draw ratios. Fibre incorporation and material processing were jointly responsible for the property enhancement. The materials were also processed into bioresorbable screws with promising results [81,82].

Different treatment processes performed on phosphate glass fibres further impact their performance in composites. When short glass fibres were incorporated into PLA in the form of randomly oriented non-woven mats, there was a mass loss of 14% and 10%, respectively, for non-treated and heat-treated fibre composites over 6 weeks, in comparison to no mass loss seen for pure PLA. Incorporation of glass fibres enhanced the material modulus (2.5 GPa → ~ 5 GPa for both composites) and flexural strength significantly, with the latter matching that for cortical bone. Concerning the retention of mechanical properties over 6 weeks of in vitro degradation, the strength of PLA declined slightly, while the modulus was maintained. For the non-treated and treated samples, there was a significant decline in both modulus and flexural strength; 0.5–1 GPa and ~40 MPa, respectively. There was

no mass loss for PLA over the 6 weeks compared with ~12% and ~14% for heat-treated and non-treated, respectively. Heat treatment of the fibres led to a decreased dissolution rate. The PLA alone and heat-treated composites displayed higher cell viability due to their slower degradation [83]. Incorporation of short non-treated and heat-treated glass fibres in the form of randomly oriented non-woven mats was also investigated in polycaprolactone. The composite materials presented with a flexural strength and modulus of up to 30 MPa and 2.5 GPa, respectively, values that are comparable to those of the human trabecular bone. A higher mass loss was seen in the composites with a higher volume fraction (Vf 17/18%), 20% compared with 8% for (Vf 6.4%) over 5 weeks [84]. The rapid decrease in mechanical properties in glass fibre reinforced polyesters is attributed mainly to early hydration of the reinforcement due to weak interfacial interactions and polymer swelling during degradation that increases the internal stress of the system inducing early cracking and failure [85,86]. Coupling agents, including silanes, acids, and other agents that can create covalent bonds between filler and matrix material have been used to enhance the interfacial properties between phosphate glass and PLA matrices. The improved interfacial shear strength between the phosphate glasses and PLA matrix prevents early hydration of the filler and assists with the fibre/matrix load transfer, thereby improving the overall mechanical properties of the composite material [87].

Two main points can be seen from assessing composite material. Firstly, the method of incorporation of bioceramic fillers into the polymer matrices must be considered, as studies have shown certain processes can be disadvantageous to mechanical properties due to improper blending and homogeneity across the composite material [88]. Secondly, the volume fraction of the filler must be optimised, as excessive content can lead to agglomeration and depletion of mechanical properties.

## **4. Surface Enhancement**

## *4.1. Overview of Surface Enhancements*

Surface modifications for bone fixation has been an active area of research for decades and is commonly carried out in order to enhance physiological bone fixation, assist the healing process, and improve biocompatibility, functionality, and biological efficacy. The success or failure of the implant is dependent on the device and surrounding tissue at the implant interface [89]. Fabrication processes can change the surface composition in comparison to the material bulk. This may be due to oxidation or hydrolysis of surface groups and/or preferred molecular orientation of surface groups in order to minimise surface free energy. Such effects typically occur unevenly over a surface, which can impact the performance of the material [90]. It is thereby crucial to alter polymer surfaces to be able to regulate concurrent surface interactions and responses.

Controlling biocompatibility is an ongoing challenge with biomaterials, as synthetic and naturally occurring polymers quite often do not have the surface properties, which are required for specific applications [91]. Generally, surface enhancements are made with certain objectives. When considering polyester surface enhancements, they are challenging to modify due to their ease of degradation with chain scission, solvent sensitivity, and low heat stability [92,93]. Surface enhancements are typically conducted to increase or reduce [90]:


Surface engineering generally includes alteration of topographical (i.e., roughness) and chemical (i.e., coating) characteristics of a medical device. An increased roughness

is desirable as it leads to an increased surface area, which in turn gives a larger area for cell adhesion. It additionally enhances biomechanical interlocking between bone tissue and implant [89]. Calcium phosphates are typically the group of materials used for coating orthopaedic implants due to the excellent bioactivity of these ceramics. Together with the mechanical advantages of the substrate implant, there is an improvement of the implants overall performance [94].

#### *4.2. Surface Enhancements of Polyesters for Bone Fixation*

From reviewing literature and the commercial products, which are available on the market for bone fixation (Table 2), it is apparent that surface enhancements of polyesters used in bone fixation are not ye<sup>t</sup> employed. A vast array of in vivo and in vitro work has been conducted on bone tissue engineering applications, but this has not ye<sup>t</sup> been thoroughly investigated in bone fixation materials and devices. Figure 2 illustrates the array of techniques that can be used to alter surface properties. These techniques have been divided into three categories: (i) roughening, (ii) coatings, adhesions and depositions, and (iii) grafting. Each group of techniques allows for different topographical surface enhancements. It can be difficult to assess which technique is the best fit, for the purpose of enhancement, for a specific application. Work conducted on plasma, chemical, or laser methodologies investigated what treatment was most appropriate for the modification of PLA surfaces. The effects of each treatment was looked into mechanically (with surface roughness analysis), surface wettability, and chemically (via XPS). Chemical treatment caused the most drastic increase in surface roughness. Plasma treatment lead to an increase in roughness and was found to be dependent upon exposure time. Laser treatment appeared to decrease surface roughness when compared to the untreated PLA. Given these results, chemical modification may be an appropriate method to be used on PLA joint implant surfaces, as an increased surface roughness leads to an increased strength of the joints. Materials processed with plasma exhibit an oligomeric layer on the surface, which could be detrimental to the adhesive joint formation process. As laser-treated surfaces showed a decrease in surface roughness, this method was not deemed suitable to use with regards to material mechanical properties. Chemical modifications cause the least change in water contact angle, but plasma and laser treatments show a significant increase. Plasma treatment overall induces the higher surface energy of the three treatments. Oxygen content of the surfaces increases with all three modes, with plasma processing giving the highest. Plasma modifications are deemed to be the most beneficial of the three processes for polymer implants [95]. Plasma modifications are known to be an effective method to treat the surface of polymers for biomedical applications as these treatments can be selective, ye<sup>t</sup> not affect the bulk polymer characteristics [96,97]. As shown in Figure 2, plasma can be used to introduce surface roughness, graft surfaces, and deposit material onto a surface. Plasma surface modifications often offer a shorter processing time in comparison to other surface modification methods. Although both plasma treatment and plasma coating technologies are commercially available for modifying polymeric surface properties, their use in fracture fixation applications has not been significant to date.

**Figure 2.** Polymer surface modification methods (for biodegradable polymers) [89,91,98–101].

There is a large scope of data regarding the biocompatibility of polyester materials, specifically with regards to surface enhancement [101–105]. As the polymer biomaterials have interaction directly with an extracellular environment, it is vital that an adverse immune response is not provoked [91,100]. A few important surface characteristics can be altered to obtain the right biocompatibility control for the intended application. These include surface morphology, chemical structure and functional groups, interfacial free energy,wettability, cytotoxicity, and adhesion [91,105]. Plasma treated PDLLA was combined with collagen anchorage. The plasma surface property alterations included improved surface hydrophilicity and increased surface energy. As a result of plasma pre-treatment, there was more collagen fixated to the surface. Mouse fibroblast cells were used to assess cell affinity and showed good affinity with the plasma treated, collagen anchored samples in comparison to untreated surfaces [106]. Similar enhanced cell affinity behaviour on PDLLA has also been reported after treatment with anhydrous ammonia plasma treatment [107]. In contrast, PLA modified with medium pressure dielectric barrier discharge (DBD) plasma treatment illustrated an increase in hydrophilicity along with an increase in oxygen content, measured by contact angle measurements and XPS, respectively. Biologically, plasma modification of PLA led to increased initial cell attachment; however, after 7 days, there was no significant difference in activity on the untreated and treated samples.It was concluded that cell proliferation was not influenced by the application of plasma treatment to the surface [98]. The influence of plasma treatments on these polymer surfaces is dependent on the experiment parameters and materials, which will give a variation in cell attachment success.

A challenge with altering material surfaces is the side effect that may be imposed on the bulk material [91], however; there are surface enhancement techniques that can limit or even avoid this. A method investigated on polyester materials, which may limit the influence on the bulk, is extreme ultraviolet radiation (EUV). Degradation of the bulk material can be limited through using short wavelength radiation, which is in the extreme ultraviolet range; which, in turn, is only absorbed by a very thin layer of the polymer surface (<100 nm) [91]. Vapour phase (VP) grafting of N-vinylpyrrolidone was conducted on four types of biodegradable polymers; PLA, PCL, PLGA, and PTMC. Wettability of all materials was enhanced by the process and it was possible for the surface topographies not to be altered due to the thin graft layer applied. Additionally, film surfaces for the grafted materials with VP after 30 min was rougher than the original polymers and enhanced good cellular adhesion was noted with PLLA, PLGA, and PTMC [99]. Radio frequency (RF) plasma is an additional process for modification without bulk properties being affected [108]. Characteristics, including advantages and disadvantages of some commonly used surface enhancement treatments on polyester surfaces, are summarised in Table 2.


**Table 2.** Summary of surface enhancement treatments on polymers [89,102,105,109,110].

> The methods discussed highlight some of the breakthroughs, which have enabled triggering of specific responses, recruitment of the correct cells and stimulus for these cells to perform. These are additional functionalities to otherwise inert polyester biomaterials [89].

## **5. Current Market Products**

Table 3 summarises the current polyester-based biomedical devices used in bone fixation applications. It can be seen that the majority of devices are composed of PLA and its composites, which is to be expected, as PLA in clinical applications is the most commonly used biodegradable polymer [111]. This is predominantly due to its excellent biocompatibility, mechanical properties, and history of use in medical applications since the 1970s [112]. The majority of products that are detailed are used in small load bearing applications, which suggests that material enhancements to biodegradable polymers are still a requirement in order to be used in larger load applications. This may in part be due to the biodegradability and mechanical properties, prior to and during degradation, of these materials, when scaled into larger devices.

**Table 3.** Orthopaedic fracture fixation devices on the market.



**Table 3.** *Cont.*

## **6. Final Considerations and Perspective**

Table 4 summarises the enhancement strategies, material/s, and the resultant modified properties, as discussed in this review.

The main reasons for seeking to enhance the properties of materials for fixation are:



**Table 4.** Polyester enhancement strategies for bone fixation bioresorbable polyester materials.

> For the most part, enhancements have been applied using PLA or PLLA as a comparator material, aiming to retain strength, reduce brittleness, increase degradation to promote resorption, mitigate the acidic degradation products, and promote better integration between the polymer and the native tissue.

> The complexity arises from the inter-related nature of the properties, which are being manipulated: most of the enhancement methods previously described affect more than one of the properties of interest, and the device design can be used to mitigate some effects. For example, most devices are relatively thin, which enhances degradation and limits the volume of degradation products released at the implantation site.

> From Table 2 it is clear that two key material enhancement strategies have been adopted broadly adopted clinically, and so can be seen to offer benefits with clinical value:


The use of materials with anisotropic mechanical properties, through orientation of the polymer or through long fibre reinforcement, has not been adopted clinically, perhaps suggesting that the loading that the devices are put under in practice is too complex to be addressed by enhancement in one orientation.

In terms of future developments, there is ongoing interest in blending new polymers with polyhydroxyalkanoate-based materials, popular for their green production route, biocompatibility, and benign degradation products [113]. There is also ongoing interest in

refinements to particulate bioceramic additives, in terms of the additive itself, with Mg and Sr substituted HA systems explored, and in terms of the form of the reinforcement, with both nanoscale additives and short fibres being explored for enhanced biocompatibility and degradation profiles.

Surprisingly, surface modification techniques have not been significantly explored for fixation devices, despite the more general growth of interest in using surface functionalisation methods to enhance medical devices [114]. Reported work to date has predominantly focussed on promoting hydrophilicity on surfaces, but it is possible to apply specific molecules onto surfaces to elicit specific responses in vivo. The reasons for the application may be to generate a more biomimetic surface for attachment to the native tissue, or to release stimulatory or therapeutic molecules to the fracture site, and exploration of the potential for this in fixation applications could offer a way to further improve the performance of this important class of medical devices.

**Author Contributions:** All authors have read and agreed to the published version of the manuscript.

**Funding:** R.N. is funded by the European Union's Horizon 2020 research and innovation program under gran<sup>t</sup> agreemen<sup>t</sup> no. 814410. C.T. is funded by the UK Medical Research Council Confidence in Concept fund (MC\_PC\_18057).

**Conflicts of Interest:** The authors declare no conflict of interest.
