The effect of bisphosphonate-modified poly(L-lactic acid) micro-fibers on the stiffness, bending strength and toughness of calcium phosphate cements.
The main goal of this research is to research if bisphosphonate modification of poly(L-lactic acid) micro-fibers has a significant effect on the stiffness, bending strength and toughness of calcium phosphate cements.
The current generation of synthetic bioceramic bone substitutes such as calcium phosphate cements (CPCs) has poor mechanical properties, which limits the use of these materials to non-load-bearing skeletal sites. These poor mechanical properties can be overcome by adding poly(L)-lactic acid (PLLA) fibers, as these are biodegradable polymers with relatively good strength and biocompatibility. The adhesion between the PLLA fibers and the CPCs is, however, not optimal due to the difference in polarity between these two types of materials. PLLA is a hydrophobic polymer, whereas CPC is a hydrophilic ceramic. Therefore, this study researched the effect of surface modification of PLLA fibers on the mechanical properties of fiber reinforced CPCs. More specifically, the goal of this study was to research if surface modification of poly(L-lactic acid) micro-fibers with calcium-binding bisphosphonate (BP) groups using a polydopamine (PDA)-based immobilisation method had an effect on the stiffness, bending strength and toughness of calcium phosphate cements. To this end, CPCs with or without PLLA fibers were prepared. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP). The fibers were cut mechanically as well as manually. All samples were tested mechanically using a three-point bending test. We hypothesized that the stiffness, bending strength and toughness of calcium phosphate cements reinforced with surface-modified PLLA fibers would be significantly higher than the cements reinforced with unmodified poly(L-lactic acid) microfibers. Our results showed that the fiber modification had no statistically significant effect on either the stiffness, bending strength or toughness of the CPCs reinforced with mechanically or manually cut PLLA fibers. Summarizing, the added value of surface modification of the PLLA fibers using BP coatings was not confirmed in the current study.
In the oromaxillofacial area bony defects may exist due to infections, tumors, cysts, fractures, periodontal disease or congenital malformation. These defects may occur at different locations such as the marginal bone, intra-alveolar bone and/or at the furcations of teeth. Applying bone substitution materials preserves the alveolar ridge, prevents resorption of bone and provides sufficient bone for immediate or subsequent implant placement.
Different treatments have been developed to increase bone volume; for example bone grafts (autogenous and allogenic block grafts) and barrier membranes for guided bone regeneration. The use of autogenous bone grafts is still the gold standard for the reconstruction of large bone defects. Autologous bone possesses optimal osteoconductive and osteogenic properties. It can be retrieved from several donor sites: crista iliaca, the ribs, the skull, the tibia and several sites in the oral cavity[1].
However, several other materials are available for the use of bone substitution such as animal-derived bone substitutes and synthetic bone substitutes. One widely used xenogenic graft is an animal-derived bone substitute such as Bio-Oss, which is a deproteinized sterilized bovine bone. No local T-cell or B-cell inflammatory reactions have been reported with the use of this inorganic bone. The material has shown to be biocompatible and osteoconductive and is widely used in the medical sector. The disadvantage of the use of these grafts is the extensive procedures in order to obtain and place the materials [2]. This is why synthetical bone substitutes have been on the rise for many years. These substitutes can simplify the process of retrieving the donor bone and make the procedure less invasive.
Ideally a synthetic bone substitute is biocompatible, supports new bone formation, shows minimal fibrotic reaction and can be remodelled. The first main category of bio-ceramics is the bioactive glass. This is a hard, non-porous, material and has osteointegrative and osteoinductive properties. This glass is based on a glass-forming SiO2 network. The key compositional features which are responsible for the bioactivity are the low SiO2 content, the high Na2O and CaO content and the high CaO/P2O5 ratio[3]. A mechanically strong bond is formed between the glass and bone through hydroxyapatite crystals. The main disadvantage of these materials is their poor mechanical strength.
The second main category of bio-ceramics is the calcium phosphate ceramics (CaP ceramics). CaP ceramics are synthetic materials and have been used since the 1980s [4]. These materials do not have osteoinductive or osteogenic properties, but are osteoconductive. The two most widely used calcium phosphate based bio-ceramics are hydroxyapatites (HAP) and β-tricalcium phosphates (β-TCP). Hydroxyapatite (HA) is a very biocompatible ceramic, which is known to be brittle and undergoes a slow resorption process. β-tricalcium phosphate is known to have high osteoconductivity and a fast resorption process [5]. A particularly interesting class of CaP ceramics are the CaP cements; they are injectable pastes, which harden after application into defects.
Calcium phosphate cements have been researched since the 1980s and are often used as bone substitutes. These materials harden in vivo and have strong similarities with the mineral phase of bone and teeth. They are osteoconductive, which induces growth of bony tissue in order to repair damaged bone. One major advantage is that upon mixing the calcium phosphate powder with a liquid phase, they become mouldable into every shape and size and can easily be injected in the site of bone damage. Once placed it is desirable that the material is resorbed and replaced by regenerating bone.
Currently, several formulations have been established which can transform into either hydroxyapatite (HA:Ca10(PO4)6(OH)2) or brushite (DCPD: CaHPO4.2H2O) cements. In comparison to the brushite cements, the apatitic cements have a higher mechanical strength. It also takes longer before the apatitic cements are dissolved [6]. Another advantage of the CPC is the ability to be injected, which is characterized by the capacity of a paste to be forced through a needle without causing demixing[7]. Surgery is minimally invasive and complex shaped defects will be filled adequately [8]. Unfortunately, the mechanical properties of CPCs are poor in terms of strength, ductility, toughness and fatigue resistance. This limits the use of CPCs to non-stress-bearing sites [9]. This is why one strategy to reinforce these materials is by addition of fibers in the cement matrix.
Due to their brittle nature, CPCs typically exhibit sudden fractures without any previous plastic deformation. Therefore the load required to fracture the cement is very low. One way to increase the mechanical strength and fracture toughness is by reinforcing the CPCs with fibers. This idea is taken from civil engineering where they use fibers to reinforce concrete.
In CPCs failure will occur due to the fast propagation of a single crack while in fiber reinforced CPCs the composite will bear increasing loads after the matrix begins to crack. This occurs if the resistance of the pulled-out fibers at first crack is higher than the load at first crack. Since the matrix does not resist any tension at fracture. The fibers carry the entire load taken by the composite ant they will transfer stresses to the matrix through bond stresses when the load on the composite is increased.
A network of multiple cracks will occur by fiber pull-out and fiber bridging, therefore hampering the rapid propagation of the initial crack. This leads to the bending of the material instead of its’ fracture. This process of multiple cracking will progress until either the fibers fail or the accumulated debonding will occur in fiber pull-out.
As shown in figure 1, the increasing strength results in a fast crack and a low amount of plastic deformation with monolithic CPC. However, the fiber reinforced CPC increases in strength with increasing stress, which makes it a strain hardening composite[6].
Interface is an important aspect in fiber reinforced composites. The interface bonding between the fiber and the matrix needs to be of an adequate strength for the load to take place between the two elements of the composite. A bond that is too weak will lead to fiber pull-out, while a strong bond will determine fiber fracture.
Fig. 1 Stress-strain curves for a brittle behaving CPC and fiber reinforced composites displaying quasibrittle tension softening or strain hardening behavior [6]
Also, another important parameter is the length and diameter of the fibers, which influences the transfer of the stress from the matrix to the fiber. Fiber length is relevant because Fibers that are too short do not fully utilize their reinforcing capability. This leads to pull-out before the fiber fractures. A critical fiber length and volume is necessary for efficient ductilization and to induce multiple cracking.
Biocompatibility relates to the ability of a biomaterial to exhibit its desired function towards a medical therapy, without causing any undesirable systemic or local responses in the recipient of the therapy [10]. Since the fibers for calcium phosphate reinforcement are used in the medical sector their biocompatibility is a crucial aspect that needs to be considered. Thus, fibers can be degradable or non-degradable. Reinforcing CPCs with degradable fibers follows a different strategy compared to using non-degradable fibers. This first category of CPCs is focused on providing temporary reinforcement. Bone is able to grow into the macropores that are created after the fibers are dissolved. The strength lost by fiber degradation is ideally compensated by the formation of new bone. Poly(L-lactic acid) (PLLA) is a commonly used biodegradable polymer for fiber reinforcement due to its relatively good strength and biocompatibility. Research has shown that porous PLLA has a strong influence on the formation of new bone by degradation and absorption of PLLA [11]. The second category of CPCs reinforced with non-degradable fibers is mainly focused on mechanical reinforcement. A non-resorbable fiber such as polyamide 6.6 (PA6.6) is widely used. These fibers have a high failure strain and a low elastic modulus. The combination of these parameters results in a high work of fracture as a measure for the toughness [9].
Fiber length, fiber volume fraction, and fiber strength are important parameters that control the mechanical properties of fiber reinforced CPCs [12]. Long fibers are needed to bridge large cracks at non-stress-bearing sites; however the volume fraction of longer fibers can be much smaller in comparison to the volume fraction of shorter fibers. The long fibers decrease the injectability of the mix significantly. In order for fibers to be used effectively for fiber reinforcement, a critical minimum length (lc) can be calculated as shown in the equation below.
(1)
where lc represents the critical minimum length of the fiber, df is the fiber diameter, f,B is the fiber strength and i is the interfacial shear strength between the fiber and matrix.
The equation shows that the critical length decreases with increasing adhesion of the fiber to the matrix. If the interfacial shear strength increases, the critical length (lc) will decrease. In summary, when the adhesion of the fiber to the matrix is improved, the fibers can be shorter. A shorter fiber also improves the workability of the cement [6].
The interfacial shear strength (ISS) is essential for the biomechanical properties of fiber reinforced CPCs. The fiber and matrix work together and therefore the synergy of this two-component system is crucial. By shear deformation the load is transferred at the matrix-fiber interface. The adhesive forces between the fiber surface and the matrix have been known to have an efficient influence on stress transfer between the two phases [9]. When a weak bond exists between the fibers and the matrix, it is possible that the fibers slip out at relatively low loads and lose their fiber-bridging efficacy. In this situation, toughening the system by adding fibers loses its purpose. However, if the bond is too strong, fibers will break before they can dissipate energy through a sliding out mechanism. Polyesters are commonly used as fiber material for reinforcement of calcium phosphate cements. Nevertheless, PLLA is a hydrophobic material, which means that PLLA fibers do not easily adhere to hydrophilic CPCs.
Several treatments have been applied in order to improve the matrix-fiber interface. One of these treatments is the mechanical anchoring of the fibers. Anchoring in the matrix can be achieved by fibers with specifically designed morphology. Also roughness of the surface of the fiber can lead to friction and therefore mechanical anchoring.
Another approach of improving the interface is modifying the fiber surface with different functional groups. Surface functionalization by plasma cleans the surface and subsequently generates new polar groups. An etch effect leads to a roughening of the fiber surface and wetting by aqueous media is improved[6].
A third approach will be investigated in the current study, which involves specific chemical binding of the fibers to the matrix. Currently, a novel approach for improving the adhesion of fibers to the cement matrix is researched at the Department of Biomaterials at Radboudumc. This method involves polydopamine-assisted immobilization of bisphosphonate groups onto PLLA polyester fibers to improve the adhesion of PLLA fibers to calcium phosphate cement. Bisphosphonates (BP) are the most commonly prescribed drugs for osteoporosis. They display a specific and very strong adhesion for calcium ions as present in the mineral phase of bone[13]. This might suggest that fibers modified with BPs will have an improved affinity to the CPCs and therefore increase the interfacial shear strength. To immobilize these calcium-binding BP groups onto PLLA fibers, polydopamine immobilization was used in the current study. Polydopamine is advocated as a universal glue due to its adhesive properties [14]. In nature, the proteins that are produced and secreted by mussels have a good adhesive property and dopamine is a compound that retains certain functionalities similar to this class of mussel inspired adhesives, which makes it a great candidate for coating different types of substrates.
The current generation of synthetic bio-ceramic bone substitutes such as calcium phosphate cements (CPCs) has poor mechanical properties, which limits the use of these materials to non-load-bearing sites. Fiber reinforcement using degradable polyester fibers is proposed to overcome this problem, but conventional polyester fibers such as poly(L-lactic acid) are hydrophobic and do not adhere to calcium phosphate ceramics. Consequently, fiber reinforcement of CPCs using polyester fibers is not yet optimal.
Can the stiffness, bending strength and toughness of fiber reinforced calcium phosphate cements CPCs be improved by coating poly(L-lactic acid) micro-fibers with bisphosphonates using a polydopamine based mobilisation method?
To study the effect of bisphosphonate modification of poly(L-lactic acid) micro-fibers on the stiffness, bending strength and toughness of calcium phosphate cements.
The starting material for this study are poly(L-lactic acid) microfibers with a diameter of 11 um and a length of 6. These fibers were cut both manually and mechanically into fiber lengths of 2.0 mm. The mechanically cut fibers were trimmed with a Pulverisette 19 cutting mill while the manually cut fibers were shortened with a carbon steel surgical blade. The next step in the study refers to the chemical modification of the fiber surface. To this end, the fibers are initially coated with polydopamine, followed by conjugation of bisphosphonate (BP) groups (more specifically alendronate groups) onto the fiber surface. Following this, an apatite-forming calcium phosphate cement was mixed with a) unmodified, b) polydopamine-modified poly(L-lactic acid) or c) bisphosphonate-modified poly(L-lactic acid) microfibers. Rectangular CPC samples (4 × 4 × 20 mm) were made consisting of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with either PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP). A sample size of n=10 was applied to all the groups. The stiffness, bending strength and toughness of the samples were measured in a three point bending test in which force versus the displacement was measured. The toughness, flexural modulus and flexural strength of the tested specimens were calculated from these force-displacement curves and analysed statistically using one-way analysis of variance (one-way ANOVA) followed by a Tukey post hoc test.
Poly(L-lactic acid) micro-fibers were obtained from Trevira (Bobingen, Germany) and α-tricalcium phosphate (α-TCP) micro-particles were obtained from CAM Bioceramics (Leiden, The Netherlands). Amino-bisphosphoante (BP) alendronic acid was purchased from AK Scientific. All other chemicals such as ethylenediamine (EDA) were obtained from Sigma Aldrich, unless stated otherwise.
The mechanically cut fibers were cut into 2.0 mm using a cutting mill (Fritsch, Pulverisette 19). The manually cut fibers were cut into 2.0 mm using a scalpel. For surface modification, each type of fiber was immersed in an aqueous coating solution at room temperature. The reaction solution was obtained by dissolving dopamine hydrochloride (2 mg/mL) in 10 mM Tris buffer (pH 8.5). The solution was added to the fibers and the reaction took place for 24 hours under mild shaking (150 rpm). After 24 h coating time, the loosely bound polydopamine was washed two times with Tris buffer. The solution was added once more to the fibers and the same reaction took place. After 24 h coating time, the loosely bound polydopamine was washed two times with a Tris buffer. All the samples were freeze dried.
The mechanically cut fibers were cut into 2.0 mm using a cutting mill (Fritsch, Pulverisette 19). The manually cut were cut into 2.0 mm using a scalpel. For surface modification, each type of fiber was immersed in an aqueous coating solution at room temperature. The reaction solution was obtained by dissolving dopamine hydrochloride (2 mg/mL) in 10 mM Tris buffer (pH 8.5). The solution was added over the fibers and the reaction takes place for 24 hours, under mild shaking (150 rpm). After 24 h deposition time, the loosely bound polydopamine was washed two times with Tris buffer. A solution of PDA/EDA/BP was prepared by dissolving dopamine hydrochloride (2 mg/mL) and ethylenediamine (111 µL) in 10 mM Tris buffer (pH 8.5). The solution was added to the fibers. By using this technique of PDA/EDA coating, a high density of amine groups was obtained on the substrate surface. Subsequently a 5% glutaraldehyde v/v concentration was added to the fibers. Glutaraldehyde acts as crosslinker for the reaction between the amine groups of alendronic acid and the amine groups present on the surface of fibers. The reaction took place for 24 hours. After 24 h coating time the fibers were washed again with Tris buffer, the supernatant was removed and the fibers were freeze-dried.
All fibers were weighed in a plastic container. α-tricalcium phosphate (α-TCP) and 4% (w/v) Na2HPO4 solution were added to the fibers and manually mixed to disperse the fibers in the cement matrix. After mixing, the blend in paste form was immediately poured into a silicon mold of 4x4x20 mm. The weight fraction of the fibers was 2.5%. The composite mixture in the mold was covered with a glass slide and stored at room temperature for 24 hours. Subsequently, the hardened composite samples were removed from the mold and incubated in phosphate buffered saline (PBS) for 3 days prior to testing the mechanical properties. The fiber-free CPC samples were prepared in the same way without adding the fibers to the mixture.
Rectangular samples were produced (4 × 4 × 20 mm) followed by a three-point bending test [8]. The samples were soaked first in PBS for three days prior to the three-point bending test. The tests were performed on a mechanical test bench using a load cell of 100 N and at a crosshead speed of 1 mm/min for the samples containing fibers and 0.2 mm/min for the fiber-free CPC specimens. The tests were stopped at an extension of 2 mm.
Fig. 2 Three-point bending test 
Before testing, each sample was measured in order to get the correct height and width. As seen in figure 2, the sample was placed on two supporting pins a set distance and a third loading pin was lowered from above at a constant rate until the sample fractures. After fracture, the mechanical properties were evaluated from the resulting load–displacement curve:
The flexural strength of a material is the maximum stress before failure occurs.
Flexural strength: Sc = 3PmaxL/(2b h2) (2)
in which Pmax is the maximum load on the load–displacement curve, L is the flexure span, b is the sample width and h is the sample thickness.
The flexural modulus defines the relation between stress (force per unit) and strain (proportional deformation) in the material. The modulus of living bone is quite low and ranges between 15GPa and 20 GPa.
Flexural modulus: E = mL3/[4b h3] (3)
in which m is the slope of the tangent to the initial straight-line portion of the load–displacement curve (N/m), b is the sample width and h is the sample thickness.
The Work of Fracture (WOF) is used to characterize the toughness of a material. The correlated energy dissipation results in a larger Work of Fracture. [6]
in which A is the curve under the load–displacement (the work done by the applied load to deflect and fracture the sample), b is the sample width and h is the sample thickness.
Work-of-fracture: WOF = A/(bh) (4)
Table 1 An overview of the experimental groups of CPCs with or without mechanically/manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
| Abbreviation | PLLA w/t% | PDA | BP | ||
| CPC-PLLA | 2,5% (mechanically cut) | no | no | ||
| CPC-PLLA-PDA | 2,5% (mechanically cut) | yes | no | ||
| CPC-PLLA-BP | 2,5% (mechanically cut) | yes | yes | ||
| CPC-PLLA | 2,5% (manually cut) | no | no | ||
| CPC-PLLA-PDA | 2,5% (manually cut) | yes | no | ||
| CPC-PLLA-BP | 2,5% (manually cut) | yes | yes | ||
| CPC | 0 % | no | no | ||
SEM
The fractured surfaces of the samples were collected to study the morphology after the bending tests. We used scanning electron microscopy (Zeiss), which was operated at an accelerating voltage of 10 kV and a working distance of 15 mm.
Data are presented as mean ± standard deviation. A one-way analysis of variance (ANOVA) followed by a Tukey post hoc test was performed to identify significant differences. Results were considered significant if P < 0.05. Calculations were performed using GraphPad Instat®.
9. Results
In figure 3 representative load-displacement curves of the three-point bending tests are shown of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA(CPC-PLLA-PDA) or BP( CPC-PLLA-BP).
The CPC without mechanically cut PLLA fibers displayed a very low amount of plastic deformation. An extension of 0,1 mm was reached, after which the load dropped instantaneously. For all CPC groups with mechanically cut PLLA fibers the extension reached up to 1 mm and for some samples even up to 2 mm. After the initial linear elastic phase a phase of plastic deformation is seen in which the curve moves up and down repeatedly. With increasing displacement the load decreases until it almost reaches 0 N. The maximum load was at an average of 15 N for each of the CPCs with mechanically cut PLLA fibers.
Fig. 3 Load-displacement curves of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
In figure 4 representative load-displacement curves of the three-point bending tests are shown of CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA(CPC-PLLA-PDA) or BP( CPC-PLLA-BP).
The CPC without mechanically cut PLLA fibers displayed a very low amount of plastic deformation. An extension of 0,1 mm was reached, after which the load dropped to 0 N. For all CPC groups with manually cut PLLA fibers the extension reached up to 1 mm and for some samples even up to 3 mm. After the initial linear elastic phase, a phase of plastic deformation is seen in which the curve moves up and down repeatedly. With increasing displacement the load decreases until it almost reaches 0 N. The phase of plastic deformation starts at a different load for each of the CPCs with manually cut PLLA fibers. CPC-PLLA enters this phase at an average load of 13 N, CPC-PLLA-PDA enters this phase at an average load of 11 N and CPC-PLLA-BP enters this phase at an average load of 8 N.
Fig. 4 Load-displacement curves of CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
In figure 5 the flexural strength as calculated using formula 2 is shown for CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 5 Flexural strength of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
The average flexural strength of all the experimental groups was about 6 MPa. Differences between all experimental groups were statistically insignificant (p>0,05).
In figure 6 the flexural strength as calculated using formula 2 is shown for CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 6 Flexural strength of CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
The average flexural strength of all the groups is around 6 MPa. Differences between all experimental groups were statistically insignificant (p>0,05).
In figure 7 the flexural modulus as calculated using formula 3 is shown for CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 7 Flexural modulus of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
All the groups have an average flexural modulus of around 1300 MPa. Differences between all experimental groups were statistically insignificant (p>0,05).
In figure 8 the flexural modulus as calculated using formula 3 is shown for CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 8 Flexural modulus of CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP). “*” indicates significant difference relative to the two groups connected by the line (p< 0,05)
The flexural modulus for all the groups was statistically insignificant for all groups except CPC-PLLA-PDA for which a significantly lower stiffness was measured as compared to unreinforced CPC and CPC-PLLA-PDA samples (p<0,05).
In figure 9 the Work of Fracture (as a measure for toughness) as calculated using formula 4 is shown for CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 9 Work of Fracture (J/m2) of CPCs with or without mechanically cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP). “*” indicates significant difference relative to the two groups connected by the line (p< 0,05)
The Work of Fracture of the CPC specimens without mechanically cut PLLA fibers was significantly lower than the work of fracture for all the groups containing mechanically cut PLLA fibers (p<0,05). Between these three groups the Work of Fracture of the CPC-PLLA specimens was statistically lower than the Work of Fracture of the CPC-PLLA-PDA samples (p<0,05).
In figure 10 the Work of Fracture (as a measure for toughness) as calculated from formula 4 is shown for CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP).
Fig. 10 Work of Fracture (J/m2) of CPCs with or without manually cut PLLA fibers. The fibers were left unmodified (CPC-PLLA) or coated with PDA (CPC-PLLA-PDA) or BP (CPC-PLLA-BP). “*” indicates significant difference relative to the two groups connected by the line (p< 0,05)
The Work of Fracture of the CPCs without manually cut PLLA fibers was significantly lower than the Work of Fracture of all the groups containing manually cut PLLA fibers (p<0,05). Differences between the CPC specimens with manually cut PLLA fibers groups were statistically insignificant (p>0,05). The Work of Fracture of all these groups are around 700 J/m2 in comparison to the CPC group with a much lower Work of Fracture of ~15 J/m2.
In figure 11 a scanning micrograph is shown of a sample of CPCs without any PLLA fibers.
Fig. 11 Scanning electron micrograph of CPCs without PLLA fibers.
Nanopores are visible throughout the entire cement. A structure of entangled platelet-shaped hydroxyapatite crystallites was clearly visible, which is characteristic for the microstructure of self-hardening CPCs.
| 2,5 % Manually cut | 2,5 % Mechanically cut | |
|
PLLA |
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