The in vivo degradation behavior of medical devices can be estimated to a certain extent by analyzing their in vitro degradation [1
]. The formation of a bone-like layer on the surface of a medical device is highly beneficial for encouraging bonding to bone [31
]. Thus, the degradation behavior and apatite formation of bioresorbable PLDLA composite pins reinforced with CGFs were investigated in this study to evaluate the feasibility for use in load-bearing orthopedic applications.
4.1. Degradation Behavior
Degradation behavior plays an important role in the success of resorbable implants for use in load-bearing applications. In addition to the effects on mechanical performance, the degradation mode can have a substantial effect on long-term clinical benefits [35
]; fragmentation degradation of pure polylactide has been shown to result in late complications [36
] and rapid delamination of phosphate glass fibers (PGFs) can lead to a mismatch with the rate of bone ingrowth [37
A significant finding of this work was the periodic degradation of the end face of the biocomposite pins which resulted in compelling nano- and micro-structures (Figure 1
). The periodic degradation occurred in four steps: (1) nano-cracks appeared around the fibers at the matrix interface; (2) a spiral delaminating type of degradation occurred on the core of the CGFs; (3) funnel-shaped and micro-tube structures appeared at the center of the CGFs; and (4) the ends of the CGFs eroded, shortening their length, and the polymer matrix became much more porous. The formation of nano-cracks was due to the dissolution of CGFs in the SBF in step 1. The capillary effect of these cracks led to more water being diffused into the matrix interphase [1
], which subsequently led to the formation of a hydrated layer on the surface or near the surface of the CGFs, resulting in a spiral delaminating type of degradation in step 2. As the CGFs underwent delamination the interphase boundary gradually weakened (Figure 1
and Figure 2
A–F), which presented as a pull-out effect during mechanical testing (Figure A2
). Also, the shear and flexural strength of the biocomposite pins gradually decreased over time as the materials degraded (Figure A1
). Previous studies have reported that delamination of PGFs [37
] occurs due to the formation of hydrated outer layers as a consequence of their dissolution mechanism. This is a similar mechanism to the delamination of silicate glass fibers whereby the outer layers undergo hydration reactions and breakage of the glass network [41
]. However, in this study, a dramatically different spiral delamination mode occurred which resulted in the formation of funnel-shaped and micro-tube structures, as identified in step 3 above. It has been speculated that the delamination relates to the processing technique [38
] and chemical composition [37
] of the fibers. The funnel-shaped and micro-tube structures formed due to bulk degradation of the CGFs, where the CGFs degraded internally from their cores and the hydrated layer did not leach into the surrounding medium, leaving it to act as a protective layer. Finally, as noted in step 4 above, the fibers shortened due to water molecules penetrating through the capillary-like channels that formed as the fibers degraded [37
]. Moreover, the capillary-like channels and micro-tubules also improved the migration rate of and degradation byproducts of the polymeric matrix, acting to inhibit core-accelerated bulk degradation [1
The CGFs located in the bulk of the biocomposite pins did not undergo delamination until week 78 (Figure 2
A–F). However, prior to the onset of delamination, water had diffused into the pins which caused more voids to form around the CGFs. The presence of CGFs accelerated the degradation rate of the pins (Figure 2
F,L) as a result of the interaction between the fibers and the polymer matrix whereby the hydroxyl ions attacked the ester groups in the matrix [21
This study has revealed that the addition of CGFs increased the degradation rate of pure PLDLA and increased the amount of apatite deposits on the surface during the degradation process (Figure 3
). The accelerated degradation was due to the interaction of alkali and acidic byproducts in the CGFs and polymeric matrix [21
]. The change in mass of the biocomposite pins was a result of polymer degradation, fiber dissolution, and apatite precipitation (Figure 3
A); apatite precipitation was the main contributor to the increased mass seen for both materials. The apatite precipitation was also a contributor towards the increased diameter of both pins (Figure 3
]. Due to the higher degradation rate of the biocomposite pins, their thermal properties, such as Tg, Tm, and crystallinity, were slightly lower than those of the pure PLDLA. The lower thermal properties corresponded to the difference in inherent viscosity between the pins. Thus, the content of CGFs could be altered in order to tailor the degradation rate of PLDLA for use in different load-bearing implants.
In comparison to PGFs [37
], the CGFs in this study underwent a drastically different form of spiral delamination and had a much lower degradation rate. This could improve the long-term strength retention of implants (Appendix Figure A1
and Figure A2
), making these materials useful for load-bearing applications.
4.2. Calcium Phosphate Formation
The ability of a material to form apatite in SBF is widely used to evaluate its bioactivity. An apatite layer deposited on the surface of a bioactive implant needs to demonstrate the following properties according to ISO 23317, which are similar to the properties of bone mineral: (1) being calcium deficient; (2) having a lower Ca/P atomic ratio than stoichiometric apatite; 3) containing impurities such as Mg2+, Na+, Cl−, and HCO3−; (4) having low crystallinity.
In this current study, an apatite-rich layer formed on the surface of both materials (Figure 4
). When the biocomposite pins were immersed in SBF there was clear precipitation of HA on the CGFs, which was due to the following reactions [1
]: (1) the acidic degradation of the polymeric matrix producing HA byproducts; (2) the dissolution of CGFs, whereby the alkali and alkali earth metals reacted with water to form basic hydroxides in the interphase resulting in a high local pH; (3) the neutralization of the products in (1) and (2); and (4) deposition of a Ca-P layer on the silicate-rich surface layer of the CGFs (Figure 5
The calcium phosphate on the surface of both pins produced nano-sheet and nano-needle apatite structures (Figure 1
, Figure 4
and Figure 5
), which is similar to the structure of bone mineral [42
]. Previous studies [5
] have demonstrated that the quantity and appearance of calcium phosphate crystals depends on the soaking medium, soaking time, and matrix. This aspect may be investigated in future studies.
The surface deposits were composed of carbonated HA doped with sodium and magnesium ions (Na+
), which was confirmed by XRD, FTIR, and EDS analysis (Figure 4
and Figure 5
). Due to the carbonate (CO32−
), the Na+
exchanged with the Ca2+
within the network of apatite (Figure 5
), leading to a calcium-deficient deposit of apatite. Magnesium is a minor but important component of bone, enamel, and dentine, and it plays an essential role in the bio-mineralization process [45
]. Compared with pure HA of the same crystallinity, Mg-doped HA implant materials offer better adhesion and proliferation of osteoblast cells [36
]. Therefore, a HA surface can layer presents superior osteoconductivity, can promote cellular function, and offers excellent biocompatibility [47
]. The apatite layer deposited on the biocomposite pins investigated in this study had a lower Ca/P ratio than the stoichiometric HA (1.67) (Figure 5
A,B), whereas the Ca/P ratio of the apatite layer deposited on the pure PLDLA was 1.82 (Figure 5
C). This indicated that the addition of CGFs enhanced the bioactivity of PLDLA in vitro. According to the broad peaks of 2θ = 32.0° in the XRD spectrum and 1039 cm−1
in the FTIR spectrum (Figure 6
), the low degree of splitting indicated the presence of apatite with low crystallinity [29
]. Thus, the apatite deposited on the biocomposite pins had similar characteristics to bone mineral.
The results of this study strongly support the use of these novel CGF-reinforced biocomposite pins in load-bearing orthopedic applications. Future studies may investigate the mechanism behind the formation of different crystal nanostructures on the surface of the biocomposite pins. This may relate to the selection of matrix type (altering the degradation rate and acidic byproducts), content of bioresorbable glass fibers (composition, processing method, and alkali byproducts) and soaking medium (prepared and controlled method). However, two of the key properties of the biocomposite pins demonstrated in this study that support their use for load-bearing orthopedic applications are their spiral delamination degradation mode and high bioactivity in vitro.
Related to the section above on how this research may be expanded in the future, there are some limitations with the methods used in this study that should be noted: (1) the study mainly focused on the degradation of CGFs in a PLDLA matrix, but did not analyze how different materials for the matrix may impact the degradation behavior. However, the degradation features of the PLDLA (L/D = 80/20) matrix in this study were shown to be comparable to matrices composed of different materials from the literature [1
]. This demonstrates the reliability of the results presented in this study and the practicality of the implants for load-bearing applications. (2) the implants analyzed in this study had a basic pin shape. More complex implants (screw thread, pores, or hollow core, etc.) may be considered in future studies. (3) A constraint of this study is that the effect of cell-related mechanisms, such as cytocompatibility in vitro and histological response in animal testing, was not considered [1
]. Studies have shown that the composition, structure, crystallinity, and degradation rate of bioglass fibers and mineralized HA are key factors that affect the proliferation and differentiation of osteoblasts [18
]. Therefore, future work may firstly consider the cytocompatibility of such reinforced bioresorbable composites reinforced with high strength CGFs.