journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

Available online at www.sciencedirect.com

www.elsevier.com/locate/jmbbm

Research Paper

Enhanced mechanical properties of a novel, injectable, fiber-reinforced brushite cement Stefan Maenza,e, Elke Kunischb, Mike Mu¨hlsta¨dta, Anne Bo¨hmc, Victoria Kopschb, Jo¨rg Bosserta, Raimund W. Kinneb,1, Klaus D. Jandta,d,e,n,1 a

Chair of Materials Science, Otto Schott Institute of Materials Research, Friedrich-Schiller-University Jena, Jena, Germany Experimental Rheumatology Unit, Department of Orthopedics, Jena University Hospital, Waldkrankenhaus “Rudolf Elle,” Eisenberg, Germany c Thuringian Institute of Textile and Plastics Research, Rudolstadt, Germany d Jena Center for Soft Matter (JCSM), Friedrich-Schiller-University Jena, Jena, Germany e Jena School for Microbial Communication (JSMC), Friedrich-Schiller-University Jena, Jena, Germany b

ar t ic l e in f o

abs tra ct

Article history:

Injectable, brushite-forming calcium phosphate cements (CPCs) have great potential as

Received 11 April 2014

bone replacement materials due to enhanced degradability and long-term inclusion in

Received in revised form

bone remodeling. However, the use of such brushite-forming CPCs in load-bearing areas is

22 July 2014

limited by their low mechanical strength. One approach to overcome this limitation is the

Accepted 28 July 2014

use of reinforcing fibers. Thus, an injectable, biodegradable, brushite-forming CPC based on

Available online 7 August 2014

beta-tricalcium phosphate/phosphoric acid with fiber reinforcement was developed for

Keywords:

minimally invasive surgery. The fibers (diameter 25 mm; length 0.25, 1 or 2 mm) were

Calcium phosphate cement

extruded from poly(l-lactide-co-glycolide) acid (PLGA) and added to the CPC (2.5, 5 or 7.5%

Fiber reinforcement Mechanical properties Injectability Biocompatibility

(w/w)). Independent of the fiber content, injectability of the CPC was retained up to a fiber length of 1 mm. The addition of all PLGA fiber types increased diametral tensile strength, biaxial flexural strength, and flexural strength by up to 25% (p r0.05 for the diametral tensile strength for the CPC with 5% (w/w) 1 mm fibers and the biaxial flexural strength of the CPC with 5% (w/w) 0.25 mm fibers). In contrast, the work of fracture strongly and significantly increased (po0.01) by up to 12.5-fold. At constant fiber content, the mechanical properties of the fiber-reinforced CPC were mostly augmented with increasing fiber length. Also, the addition of PLGA fibers to the brushite-forming CPC (up to 7.5% (w/w)) only transiently delayed cell growth and did not decrease cell viability. Fiber reinforcement of CPCs thus augments their mechanical strength while preserving the injectability and biocompatibility required for their application in modern surgery. & 2014 Elsevier Ltd. All rights reserved.

n Corresponding author at: Chair of Materials Science, Otto Schott Institute of Materials Research, Friedrich-Schiller-University Jena, Löbdergraben 32, 07743 Jena, Germany. Tel.: þ49 3641 947730; fax: þ49 3641 947732. E-mail address: [email protected] (K.D. Jandt). 1 Both authors have equally contributed to the study.

http://dx.doi.org/10.1016/j.jmbbm.2014.07.028 1751-6161/& 2014 Elsevier Ltd. All rights reserved.

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

1.

Introduction

Minimally invasive augmentation techniques with injectable cements, i.e. kyphoplasty and vertebroplasty, have become increasingly popular for the treatment of vertebral body fractures. Especially in load-bearing areas, injectable poly (methyl methacrylate) (PMMA) cements are most frequently used, which, however, lack bioactivity and biodegradability. In addition, their supra-physiological strength and Young's modulus may lead to critical loads and subsequent fractures in adjacent vertebral bodies (Hulme et al., 2006; Nouda et al., 2009; Trout et al., 2006; Uppin et al., 2003). Calcium phosphate cements (CPCs), as first described by LeGeros et al. and Brown/Chow in the 1980s (Brown and Chow, 1983; LeGeros et al., 1982), may represent a promising alternative in this context since they are biodegradable and have a Young's modulus comparable to that of cancellous bone (Burguera et al., 2006). However, commercially available, injectable CPCs show low mechanical strength and low fracture toughness, leading to premature damage of the cement at the implant site (Blattert et al., 2009; Wilke et al., 2006). The mechanical properties of CPCs are governed by their composition and microstructure. In general, apatite-forming CPCs have higher mechanical strength than brushite-forming CPCs (Dorozhkin, 2011). In addition, the mechanical properties of both types of CPCs increase with decreasing porosity (Engstrand et al., 2014; Tamimi et al., 2012) which in turn requires a high powder-to-liquid ratio in the preparation process. In order to obtain injectable CPCs, however, low powder-to-liquid ratios (o2.4) are necessary (Khairoun et al., 1998). Thus, for an injectable CPC a low powder-to-liquid ratio has to be maintained which on the other hand limits the mechanical strength of the resulting CPC. Another possibility to increase the mechanical strength and fracture toughness is the modification of the CPC with reinforcing fibers. Detailed reviews about fiber-reinforced CPCs (Canal and Ginebra, 2011; Krüger and Groll, 2012) describe the use of non-resorbable fibers consisting of aramide (Xu et al., 2000, 2001b), polyamide (dos Santos et al., 2000), polypropylene (Buchanan et al., 2007), glass (Xu et al., 2000) or carbon (Xu et al., 2001a, 2000), as well as the use of resorbable, biodegradable fibers, such as polycaprolactone, polylactide, polyglycolide or different copolymers thereof (Nair and Laurencin, 2007). If a completely resorbable CPC is to be developed, biodegradable fibers in conjunction with degradable calcium phosphate phases are desirable (see below). Biodegradable fibers consisting of poly(l-lactide-co-glycolide) acid (PLGA) are commercially available as clinical suture material. PLGA suture material with yarn diameters of 200– 350 mm has been used as biodegradable fiber reinforcement for CPCs in several studies (Burguera et al., 2005; Dagang et al., 2007; Gorst et al., 2006; Weir and Xu, 2010; Xu et al., 2007a, 2007b, 2000, 2006; Zhang and Xu, 2005; Zhao et al., 2010a, 2010b). In general, the addition of PLGA fibers results in increased strength and work of fracture (WOF) with the fiber volume fraction (Gorst et al., 2006; Xu et al., 2000, 2006) and the fiber length (Xu et al., 2000) as critical parameters. However, due to a complex interdepencency between diameter and/or length of the fibers, load transmission, and

329

injectability, the large diameter of commercial PLGA suture material limits their application in injectable CPCs. Unfortunately, there are currently only isolated reports using PLGA fibers with smaller diameters, e.g., the application of 16 mm diameter fibers to augment the mechanical strength of a CPC in a skull defect model (Losquadro et al., 2009). In addition to their mechanical strength, degradability and injectability of the CPCs are of central relevance for their use in minimally invasive vertebral surgery. However, all of the above-mentioned studies were focused on either non-injectable, fiber-reinforced, brushite-forming CPCs (Gorst et al., 2006) or on apatite-forming CPCs with only very restricted degradability in vivo (Apelt et al., 2004). The latter type of CPC may be of limited suitability as bone replacement material due to its quasi exclusion from long-term bone remodeling in vivo. Therefore, the aim of this study was to develop and investigate an injectable, brushite-forming CPC with biodegradable polymer fiber reinforcement, suitable for minimally invasive surgery. Therefore, PLGA fibers with different length were incorporated at different fiber content into a brushiteforming CPC and the resulting composite CPC was then tested for injectability, mechanical properties, and microstructure, as well as reciprocal interdependencies among these parameters. Initial assessment of the biocompatibility of the developed composites was also performed.

2.

Materials and methods

2.1.

Fabrication of PLGA fibers

PLGA fibers were prepared from the granulate material PURASORB PLG 1017 (Purac, Gorinchem, Netherlands). The extrusion of the fibers was performed using a mini extrusion system (RANDCASTLE EXTRUSION SYSTEMS INC, Cedar Grove, USA), equipped with a ceramic spinning nozzle at a temperature of 217 1C and an offtake velocity of 400 m/min (resulting diameter: 25 mm). Fibers were then cut to different lengths using a cutting mill PULVERISETTE 19 (FRITSCH GmbH, Idar-Oberstein, Germany) with two different sieve inserts (0.25 mm and 1 mm, hereafter called 0.25 mm fibers and 1 mm fibers). A third fiber length was obtained by manually cutting the fibers to a length of approximately 2 mm (hereafter called 2 mm fibers). The mechanical properties of the extruded fibers were measured according to DIN EN ISO 2062 (2009) applying a Zwick Z005 universal testing machine (Zwick GmbH & Co. KG, Ulm, Germany). The morphology of the fibers was characterized using a Zeiss Auriga 60 scanning electron microscope (Zeiss AG, Oberkochen, Germany). The fiber lengths after cutting were determined using the quantitative image analysis software Leica QWin V3 (Leica Microsystems GmbH, Wetzlar, Germany).

2.2.

Preparation of fiber-reinforced CPCs

The CPC powder consisted of 98.5% (w/w) β-tricalcium phosphate (β-TCP) and 1.5% (w/w) tetrasodium pyrophosphate, the liquid of an aqueous solution containing 3.0 M phosphoric

330

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

acid, 0.1 M sulfuric acid, and 0.1 M citric acid (all Sigma Aldrich, St. Louis, USA). CPC powders with different fiber content (2.5, 5, or 7.5% (w/w)) and different fiber length (0.25, 1, or 2 mm) were produced by mixing defined amounts of fibers und CPC powder in pure isopropanol with a high shear stirrer. Afterwards, the isopropanol was removed by evaporation. The respective fiber concentrations refer to the fiber weight content in the CPC starting powder. The powder-to-liquid ratio was 2.2 for all experiments. The CPCs were mixed with a spatula for 2 min. Composition and powder-to-liquid ratio of the CPCs were chosen on the basis of own previous work and the characteristics of an FDA-approved brushite-forming CPC (JectOS, Kasios, L’Union, France), suitable for vertebral injection during kyphoplasty in young patients with stable and recent traumatic vertebral body fractures.

2.3.

Injectability

The injectability of the pure and fiber-reinforced CPCs was measured according to a published method (Ginebra et al., 2007; Qi and Ye, 2009). Briefly, 5.5 g powder and 2.5 ml liquid were mixed in a bowl and transferred to 10 ml syringes with an inner diameter of 15 mm, carrying a nozzle with an inner diameter of 2 mm. Exactly 150 s after mixing, the CPC paste was extruded from the syringe with a crosshead speed of 30 mm//min using a Zwick Z020 universal testing machine (Zwick GmbH & Co. KG). The test was stopped automatically when the applied force reached 100 N. The injectability was calculated as the ratio between the CPC mass extruded from the syringe and the mass remaining inside the syringe.

2.4.

Mechanical testing

For each CPC composition, 3 types of mechanical tests were performed using a Zwick Z020 universal testing machine (Zwick GmbH & Co. KG): diametral tensile test, 3-point flexural test, and biaxial flexural test. The test specimens were prepared by injection of the CPC paste into silicon molds using 10 ml syringes as above without additional compression of the samples. After fabrication, the samples were first stored for 1 h at room temperature and then exposed to 37 1C and 100% humidity for 24 h prior to testing. For each CPC composition and mechanical test, n ¼10 samples were used. The diametral tensile test was performed with a crosshead speed of 1 mm/min on tablet-shaped samples with a diameter of 10 mm and a height of 6 mm. The diametral tensile strength (DTS) was calculated according to Eq. (1). DTS ¼ 2Fmax =ðd  h  πÞ

ð1Þ

where Fmax is the maximum force measured during the test, d is the diameter, and h the height of the sample. The biaxial flexural test was performed according to DIN EN ISO 6872 (2009). Disk-shaped specimens with a diameter of 16 mm and a height of 2 mm were placed on three balls (positions shifted by an angle of 1201) arranged in a circle with a diameter of 12 mm. The force was applied using a centered piston with a diameter of 1.45 mm and a crosshead speed of 1 mm/min. The biaxial flexural strength was

calculated according to Eq. (2): σ ¼ 0:2387Fmax ðX YÞ=b2

ð2Þ

with X ¼ ð1 þ υÞln ðd2 =d3 Þ2 þ ½ð1–υÞ=2ðd2 =d3 Þ2

ð3Þ

Y ¼ ð1 þ υÞ½1 þ ln ðd1 =d3 Þ2  þ ð1–υÞðd1 =d3 Þ2

ð4Þ

where Fmax is the maximum force measured during the test, b is the thickness of the sample at the origin of fracture, d1 is the diameter of the circle defined by the three balls (12 mm), d2 is the diameter of the loaded area (1.45 mm), d3 is the diameter of the sample and υ is the Poisson's ratio (assumed to be 0.25). A standard 3-point flexural test with a span of 16 mm and a crosshead speed of 1 mm/min was performed on beamshaped specimens (rectangular cross section) with a width of 4 mm, a height of 3 mm, and a length of 20 mm. Each test was stopped automatically after the force decreased to 20% of the maximum. From the 3-point-flexural test, flexural strength, flexural modulus of elasticity, and work of fracture were calculated. The WOF was calculated by integration of the stress–strain curve. The morphology of the fractured surfaces was characterized using a Zeiss Auriga 60 scanning electron microscope (Zeiss AG).

2.5.

X-ray diffraction

The phase compositions of 3 samples each of the starting powder, as well as the non-reinforced and 1 mm fiberreinforced, hardened CPC, were analyzed using an X-ray diffraction (XRD) system D5000 operated at 40 kV and 30 mA (Siemens AG, Berlin, Germany). The brushite and β-TCP content after hardening were calculated by Rietveld refinement using the software TOPAS 3 (Bruker AXS GmbH, Karlsruhe, Germany).

2.6.

Cytocompatibility testing

The chondrogenic cell line ATDC5 (Riken BRC cell bank, Tsukuba, Japan) was used as an indicator cell line for cytocompatibility tests. Cells were cultured at 37 1C in a humidified 5% CO2 atmosphere in Dulbecco's Modified Eagle Medium (DMEM; Invitrogen, Karlsruhe, Germany) containing 10% fetal bovine serum (FBS; Fischer Scientific GmbH, Dresden, Germany). Disk-shaped specimens with a diameter of 8 mm and a height of 1 mm were produced as described in the initial paragraph of Section 2.4. Prior to the cell proliferation test, the samples were autoclaved at 120 1C and afterwards prewashed for 24 h in sterilized phosphate-buffered saline (PBS) with repeated changes of the buffer solution. Cells (160 cells/mm2) were seeded onto either polystyrene, pure CPC or CPC modified with 2.5, 5, and 7.5% (w/w) 1 mm fibers in 48-well plates. After 1, 3, 7, and 10 days of cultivation, cell number (n¼ 6 independent assays) and cell viability (n ¼3 independent assays) were determined. The viability of the cells was assessed using fluorescein/ propidium iodide (both Sigma Aldrich) staining. In brief, cells were washed once with Dulbecco's phosphate-buffered saline (DPBS) and thereafter incubated for 2 min at 37 1C in 6 mM

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

fluorescein diacetate in DPBS, followed by staining with propidium iodide (0.2 mg/ml DPBS). Four representative images of each disc or well were taken using a fluorescence microscope Axiovert200M, an AxioCam MRm camera, and the AxioVision Rel4.8 program (all Zeiss AG) and the percentages of viable and dead cells were determined. The cell density was assessed using 40 ,6-diamidino-2phenylindole staining (DAPI; Sigma Aldrich). At the indicated time points, cells were fixed for 20 min in 4% PFA (Sigma Aldrich) and subsequently washed twice with PBS. Thereafter, the nuclei of the cells were stained with DAPI (0.1 mg/ml). Four randomly selected regions of each disc were photographed using an AxioCam MRm camera and the AxioVision Rel4.8 program (both Zeiss AG). The number of nuclei was counted in the four regions, averaged and normalized to the surface area.

2.7.

Statistics

All data were expressed as means7standard deviations. Mechanical properties (except for the WOF) and the fiber length distributions were analyzed for significant differences among groups by one-way ANOVA with Tukey post hoc comparisons for all 8 groups of the respective mechanical parameter. The WOF was analyzed for significant differences among all 8 groups using an ANOVA on ranks test with Tukey post hoc comparisons. The results from the cytocompatibility tests were analyzed for significant differences among groups using the Mann–Whitney U test. For all tests, the level of significance was set at pr0.05. All statistical tests were performed using the Sigmaplot software release 12.0 (Systat Software Inc., Chicago, USA).

3.

Results

3.1.

Fiber characterization

Young's modulus and tensile strength of the PLGA fibers (diameter: 25 mm) were 3.070.2 GPa and 258.4712.2 MPa, respectively (Table 1). The smooth fiber morphology after extrusion is shown in Fig. 1(a). Cutting of the fibers in a cutting mill resulted in pinches (Fig. 1(b)). In contrast, cutting of the fibers by scissors led to damage at the fiber endings (Fig. 1(c)). The fiber length distributions after cutting with the three different methods are shown in Fig. 1(d). The length of the so-called 0.25, 1, and 2 mm fibers was 352.87175.3, 597.67362.9, and 2302.57459.3 mm, respectively (Table 1). Statistical analyses showed significant differences (po0.001) between the fiber length obtained by the three methods. In addition, a change in the fiber morphology, i.e. an undulated deformation of the fiber surface, was noted after isopropanol treatment (see Fig. 1(e)).

3.2.

331

Table 1 – Properties of the PLGA fibers. Young's modulus Tensile strength Diameter Length [lm]

[GPa] [MPa] [lm] Cutting mill (0.25 mm) Cutting mill (1 mm) Scissors (2 mm)

3.070.2 258.4712.2 25.0 352.87175.3 597.67362.9 2302.07459.3

modified with 2 mm fibers was found to be injectable only up to a fiber content of 2.5% (w/w; Fig. 2(a)). For a higher fiber content, it was not possible to extrude a sufficient amount of CPC paste and thus not possible to produce the samples required for subsequent mechanical testing.

3.3.

Mechanical testing

The mean values of the DTS for all fiber-reinforced CPCs were numerically higher than for the pure CPC, but statistically significantly higher values were only observed for the CPC with 5% (w/w) 1 mm fibers; in addition, the latter CPC showed significantly higher values than the CPC with 7.5% (w/w) 1 mm fibers (Fig. 2(b); Table 2). The biaxial flexural strength of all fiber-reinforced CPCs compositions was consistently higher than that of the pure CPC, however, a statistically significant difference was only observed for the CPC with 5% (w/w) 0.25 mm fibers (Fig. 2(c); Table 2). Except for the CPC with 7.5% (w/w) 0.25 mm fibers, all fiber-reinforced CPCs showed a higher mean flexural strength than the pure CPC (Fig. 2(d); Table 2), however, without statistically significant differences among the groups. In comparison to the pure CPC, the flexural modulus was consistently lower for all fiber-reinforced CPCs, except for the CPC with 2.5% (w/w) 2 mm fibers, however without statistically significant differences (Fig. 2(e); Table 2). In contrast, the WOF increased with increasing fiber length and fiber content. Statistically significant differences compared to the pure CPC were detected for the CPC with 2.5% (w/w) 2 mm fibers, the CPC with 5% (w/w) 1 mm fibers, and CPC with 7.5% (w/ w) 1 mm fibers; in addition the CPC with 2.5% (w/w) 2 mm fibers showed significantly higher values than the CPC with 2.5% (w/w) 1 mm and 0.25 mm fibers (Fig. 2(f); Table 2). Compared to the pure CPC (Fig. 3(a) and (b)), the size and the shape of the brushite crystals in the fracture faces of the CPC matrix were not affected by the presence of the PLGA fibers (Fig. 3(c)–(h)). However, differences were observed among the CPCs with different fiber lengths. For the samples reinforced with 0.25 mm fibers or 1 mm fibers, either only pulled-out fiber endings or only very few fractured fibers were observed (Fig. 3(c)–(f)). In contrast, a substantial proportion of fractured fibers were visible in the CPCs reinforced with the 2 mm fibers (Fig. 3(g) and (h)).

Injectability 3.4.

Pure CPC and CPCs reinforced with fibers cut in the cutting mill (0.25 mm or 1 mm in length, 2.5, 5, and 7.5% (w/w)) showed an injectability of about 95% and can therefore be regarded as well injectable (Fig. 2(a)). In contrast, CPC

X-ray diffraction

In contrast to the starting powder, pure CPC, and CPCs reinforced with different percentages of 1 mm fibers showed clear peaks of brushite (Fig. 4). An increase of the fiber

332

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

Fig. 1 – (a) Extruded PLGA fiber with smooth fiber morphology; (b) extruded PLGA fiber chopped in a cutting mill with 0.25 mm sieve insert; a pinch is visible (white arrow); (c) extruded PLGA fiber after cutting with scissors, damage of the fiber ending is detectable (black arrow); (d) cumulative fiber length distribution after cutting with the 3 different approaches (np r0.001 vs. 0.25 mm, #pr0.001 vs. 1 mm); (e) undulated fiber morphology after isopropanol treatment.

content resulted in a higher degree of conversion from β-TCP to brushite, i.e. the proportion of brushite increased from 5672% (w/w) for the pure CPC to 6374% (w/w) for CPC reinforced with 7.5% (w/w) 1 mm fibers (Fig. 4).

3.5.

Cytocompatibility test

The mean cell viability was higher than 90% for all tested samples (except for the CPC modified with 7.5% (w/w) 1 mm fibers on days 3 and 10; 84% and 85%, respectively) without statistically significant differences among the different groups (Fig. 5(a)). Until day 7 after initiation of the culture, the density of ATDC5 cells on almost all CPCs was statistically significantly lower than that on polystyrene (Fig. 5(b)). Also, at any time

point (except for day 10) the mean cell density on the CPCs modified with PLGA fibers was lower compared to that on the pure CPC. Statistically significant differences compared to the pure CPC were observed on day 1 (for the CPC modified with 2.5% (w/w) 1 mm fibers), on day 3 (for all CPCs modified with fibers), and on day 7 (for the CPC modified with 5% (w/w) 1 mm fibers; Fig. 5(b)). In contrast, there were no significant differences between pure CPC and CPCs modified with PLGA fibers on day 10.

4.

Discussion

The treatment of vertebral body fractures by vertebroplasty or balloon kyphoplasty is most often performed with PMMA

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

333

Fig. 2 – Injectability and mechanical properties of fiber-reinforced CPCs with different fiber length and content: (a) Injectability (n ¼3); (b) diametral tensile strength (DTS) (n¼ 10); (c) biaxial flexural (n ¼ 10); (d) flexural strength (n¼ 10); (e) flexural modulus (n ¼10); (f) work of fracture (WOF) (n ¼ 10). npo0.05 vs. pure CPC; nnpo0.01 vs. pure CPC; §pr 0.05 vs. CPCþ5% (w/w) 1 mm; þþ p r0.01 vs. CPCþ2.5% (w/w) 2 mm.

cements. However, complications such as systemic reactions to non-polymerized toxic monomers (Aebli et al., 2002), thermally induced damage in surrounding tissues (Deramond et al., 1999) or the damage of adjacent vertebral bodies due to supra-physiological strength and Young's modulus of the PMMA cement (Hulme et al., 2006; Nouda et al., 2009; Trout et al., 2006; Uppin et al., 2003) have been reported. In addition, PMMA cements are neither bioactive

nor biodegradable. In this context, biodegradable, brushiteforming CPCs are an attractive alternative and may in some cases even show stronger bone induction than non-resorbable, HA-forming CPCs (Apelt et al., 2004). Clinical data (Blattert et al., 2009) indicate a high risk of fractures in the implanted CPCs. The failures were attributed to the low biomechanical resistance, especially against flexural and shear forces. Therefore, the application of CPCs (especially well-

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journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

Table 2 – Mechanical properties of CPCs with different fiber length and content. Mode of fiber cutting Pure CPC Cutting mill (0.25 mm) Cutting mill (1 mm) Scissors (2 mm)

Fiber content [% (w/w)]

2.5 5 7.5 2.5 5 7.5 2.5

DTS [MPa]

Biaxial flexural strength [MPa]

Flexural strength [MPa]

Flexural modulus [GPa]

WOF [J/m²]

1.2970.19 1.3370.16 1.4470.22 1.3770.13 1.5070.23 1.5870.24n 1.2970.211;§ 1.5470.19

6.9370.82 7.8071.39 8.6971.46n 7.3871.01 7.8471.14 7.9871.10 7.5571.11 8.0670.72

4.0870.85 4.3370.61 4.6170.62 3.7470.91 4.8571.09 5.0871.16 4.4870.32 4.8870.63

2.6370.76 2.2970.62 2.3170.44 1.8770.55 2.4170.65 2.2970.43 2.2070.56 2.6370.82

9.3272.48 17.49711.292;þþ 16.4772.93 19.4475.65nn 18.8576.762;þþ 51.83727.62 65.28746.65nn 117.74761.07nn

n

pr 0.05 vs. pure CPC. pr 0.01 vs. pure CPC. § pr 0.05 vs. CPCþ5% (w/w) 1 mm. þþ pr 0.01 vs. CPCþ2.5% (w/w) 2 mm). nn

degradable brushite-forming CPCs) in load-bearing areas is still limited. It is known from the literature that fiber-reinforcement can overcome this limitation by increasing the CPCs’ flexural strength and work of fracture (Gorst et al., 2006; Xu et al., 2001a). For the first time, this study describes a brushite-forming CPC with a degradable fiber-reinforcement which is injectable and, therefore, suitable for minimally invasive operation techniques.

4.1.

Injectability of the new CPC

The current study shows that the injectability of fiberreinforced CPCs depends on both fiber content and fiber length. Whereas the fiber content exclusively influenced the injectability of the 2 mm fibers (only injectable up to a fiber content of 2.5% (w/w)), the shorter fibers (0.25 mm and 1 mm length) remained injectable up to a fiber content of 7.5% (w/w). The reason for this difference may be that the ravels forming with fibers of all length, which can be unraveled by application of a high shear stirrer in the case of the shorter fibers, even resist extensive manual stirring in the case of the 2 mm fibers. In addition, the damaged and lacerated fiber endings obtained as a result of cutting the fibers with scissors (see Fig. 1c) may have contributed to the formation and/or stability of the ravels. However, it was necessary to cut the 2 mm by scissors because it was technically not possible to cut the fibers to a length longer than approximately 1 mm with the cutting mill. On the basis of the present results, an injectable CPC for minimally invasive surgery should thus contain PLGA fibers of a maximal length of 1 mm.

4.2.

Mechanical properties of the new CPC

In addition to their injectability, also the mechanical properties of the newly developed brushite-forming CPC appear to depend on both fiber length and fiber content (Fig. 2(b)–(f)). At constant fiber content, the mechanical properties of the brushite-forming CPCs were either numerically or significantly augmented by fibers of increasing length. Anisotropicity of the fiber distribution and/or orientation, a factor also potentially relevant for the mechanical properties of the modified CPC (Puska et al., 2004), did not appear to

play a role in the present study since no preferential distribution or orientation of the fibers was observed. Previous studies (Burguera et al., 2005; Canal and Ginebra, 2011; Xu et al., 2001a) have proposed three mechanisms, by which randomly distributed fibers (of different length) may reinforce a CPC: (i) bridging of cracks and stopping crack propagation; (ii) deflection of cracks, resulting in multiple cracking of the matrix; and (iii) frictional sliding against the matrix during fiber pull-out. It has to be noted, however, that these three mechanisms were postulated on the basis of data obtained using fibers with much higher thickness/mechanical strength than those applied in the present study. The above-mentioned three mechanisms are likely also valid for the interpretation of our results, although in the present study direct evidence was only obtained for the third mechanism. However, except for one other study with very little structural information (Losquadro et al., 2009), the present study for the first time uses PLGA fibers with a much smaller diameter (25 mm) and a resulting reduced loading capacity. When using such fibers, an additional mechanism has to be considered, i.e. loading of the fiber to its maximal capacity and subsequent energy absorption by fracturing. This fourth mechanism indeed played a role in the present study since substantial fiber fracturing was only observed in the case of 2 mm fibers. The mutual relationship between fiber length and fracturing has been discussed previously utilizing the concept of the critical fiber length. The critical fiber length lc is defined as the minimum length at which the fiber is loaded up to its ultimate (tensile) strength σf. Fibers with lengths longer than the critical fiber length will break while fibers shorter than the critical fiber length will be pulled out (Krüger and Groll, 2012). The critical fiber length (lc) can be estimated from the balance between shear and tensile stress, i.e. 2 π r l τf m ¼ r2 π σ f

ð5Þ

with lc ¼ 2 l ¼ σ f d=2 τf m

ð6Þ

where σf is the tensile strength of the fiber, d is the diameter of the fibers, and τfm is the interfacial strength between

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

335

Fig. 3 – Fracture faces of CPCs with different fibers tested in the 3-point-flexural test; ((a) and (b)) pure CPC; ((c) and (d)) CPC with 5% (w/w) 0.25 mm fibers w/o fractured fibers; ((e) and (f)) CPC with 5% (w/w) 1 mm fibers with scarce fractured fibers (nonfractured fiber shown in (f)); ((g) and (h)) CPC with 2.5% (w/w) 2 mm fibers, containing a substantial proportion of fractured fibers (fractured fiber ending shown in (h)).; please note the scale bars for 200 µm in the left column ((a), (c), (e) and (g)) and for 20 µm in the right column ((b), (d), (f) and (h)).

336

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

Fig. 4 – XRD patterns of CPC starting powder, pure CPC, and CPC reinforced with 1 mm fibers. All spectra were normalized to the highest peak. The symbols mark the positions of high intensity β-TCP peaks (þ, File number 70-2065, ICDD) and brushite peaks (*, File number 72-0713, ICDD).

Fig. 5 – Cell viability (a) and cell density (b) of the indicator cell line ATDC5 on pure CPC and on CPC reinforced with 2.5, 5 or 7.5% (w/w) 1 mm fibers. Polystyrene was used as a control. npr0.05 vs. pure CPC; nnpr 0.01 vs. pure CPC; # pr 0.05 vs. polystyrene; ##p r0.01 vs. polystyrene.

matrix and fiber. From Eq. (6) it is possible to calculate the critical fiber length (lc) for the fibers applied in this study using their tensile strength (258.4 MPa; see Table 1), their diameter (25 mm), and an interfacial strength of 1.1470.25 MPa, previously determined for a brushite-forming CPC and a PLGA suture material (Gorst et al., 2006). The result of this

calculation is a critical fiber length of 2.8 mm, somewhat exceeding the critical fiber length of approx. 2 mm observed in the present study. However, due to the usage of large diameter PLGA yarns, the experiments performed by Gorst may have underestimated the interfacial strength to a certain degree (Gorst et al., 2006). In addition, the fiber surface in the hardened CPC of the current study showed undulations following isopropanol treatment (see Fig. 3(c)), possibly resulting in an increased fiber surface and, if not properly considered, to an overestimation of the critical fiber length. Considering both factors, the “true” critical fiber length may be somewhat closer to the value of approx. 2 mm determined in this study. Until now, the mechanism of energy absorption by fiber fracture has not attracted major interest in the field of fiberreinforced CPCs. The reason may be that in previous studies PLGA fibers from suture material, showing a much larger diameter (ca. 300 mm) and consequently a much higher critical fiber length, have been used (approx. 28 mm; Burguera et al., 2005; Dagang et al., 2007; Gorst et al., 2006; Weir and Xu, 2010; Xu et al., 2007a, 2007b, 2000, 2006; Zhang and Xu, 2005; Zhao et al., 2010a, 2010b). Also, only two studies have directly addressed the relationship between the mechanical properties of the CPC composite and the fiber length in general (Xu et al., 2001a, 2000), and only one of them the effects of PLGA fibers with a tensile strength of about 1.1 GPa (Xu et al., 2001a), largely exceeding the values of the fibers used in the current study. Therefore, fibers have not been loaded to their maximal capacity and consequently frictional sliding, crack bridging and crack deflection have been the most important reinforcing mechanisms in previous studies. Since mechanical reinforcement of injectable CPCs for medical purposes requires small diameter, short to medium length, resorbable fibers, the proposed mechanism of energy absorption by fiber fracture is of high importance in this context. An optimized matrix-to-fiber load transfer, for example on the basis of augmented interfacial strength via increased surface roughness or chemical modification, may be a very attractive approach. Concerning the dependency of the mechanical properties of the CPCs on the PLGA fiber content, the mechanical parameters DTS, biaxial flexural strength, and flexural strength all showed peak values at a fiber content of 5% (w/w; Table 2; Fig. 2((b)–(d)) and not at the theoretically expected maximal fiber content of 7.5%. Based on published data, this can be explained by the presence of non-wetted fibers, unreacted calcium phosphate powder, and fiber ravels or agglomerates, known to occur more frequently in CPCs with increasing fiber content (Xu et al., 2001a). In contrast, the WOF increased with increasing fiber content. The considerable variation in WOF data can be explained by the stop criterion (at 20% of the maximum force). While the pure CPC and the CPCs reinforced with the 0.25 mm fibers showed the characteristics of a brittle material, the CPCs reinforced with the 1 mm and 2 mm fibers rather displayed features of a quasi-ductile material. Thus, the CPCs reinforced with 1 mm and 2 mm fibers were still able to bear load after reaching the maximum force. However, depending on the shape of the stress-strain curves for the individual CPC samples the strain at which the force decreased to 20% of the maximum differed strongly.

journal of the mechanical behavior of biomedical materials 39 (2014) 328 –338

Thus, all above-mentioned mechanical and morphological properties of the PLGA fibers (see Table 2; including the Young's modulus which may be particularly important for the bridging of cracks and the prevention of catastrophic failure of specimen upon loading) likely contribute to the resulting mechanical features of the modified CPC.

4.3.

Cytocompatibility of the new CPC

The newly developed CPCs showed high cytocompatibility, as documented by a cell viability Z90% for all tested samples up to a fiber content of 5% (and Z85% viability for samples with a fiber content of 7.5%). Although until day 7 of the in vitro culture the cell density of the indicator cell line ATDC5 was consistently lower on the pure CPC than on polystyrene, and in some cases even lower on the PLGA fiber-reinforced CPC (see also Ignatius and Claes, 1996), this difference was no longer present on day 10. This indicates that cell growth on the newly developed CPCs is somewhat delayed but not inhibited in the long run.

5.

Wissenschaft und Kultur (TMBWK), grant reference 62-4264 925/1/10/1/01 to K.D.J. We thank Dr. M. Müller from the OttoSchott-Institute for glass chemistry (Friedrich-Schiller-University Jena) for performing the XRD measurements.

r e f e r e nc e s

X-ray diffraction

The XRD measurements showed a higher degree of conversion from β-TCP to brushite with increasing fiber content. This can be explained by a decreased ratio between the net quantity of CPC powder (i.e., CPC powder without fibers) and the volume of the CPC liquid/phosphoric acid, conditions deliberately chosen in the present study to avoid the influence of an altered powder-to-liquid ratio on the biomechanical properties of the resulting CPC.

4.4.

337

Conclusion

The present study describes for the first time an injectable, PLGA fiber-reinforced, brushite-forming CPC for minimally invasive surgery. While injectability of the CPC was retained up to a fiber length of 1 mm, incorporation of PLGA fibers resulted in an increase of DTS, biaxial flexural strength, and flexural strength (all by up to 25%), as well as WOF (by up to 12.5-fold), mostly augmented with increasing fiber length. The newly developed CPCs showed high cytocompatibility and only transiently delayed cell growth. The fiber addition may allow a fine-tuning of the mechanical properties and the resorbability of the resulting CPC. The CPC may be in principle applicable as bone replacement in both loadbearing and non-load-bearing bone defects.

Acknowledgements We gratefully acknowledge the financial support by the Carl Zeiss Foundation (doctoral candidate scholarship to S.M.) and by the German Federal Ministry of Education and Research (BMBF FKZ 0316205C to J.B and K.D.J.; BMBF FKZ 035577D, 0316205B, and 13N12601 to R.W.K). We gratefully acknowledge the partial financial support of the Deutsche Forschungsgemeinschaft (DFG), grant reference INST 275/241-1 FUGG to K.D.J., and the Thüringer Ministerium für Bildung,

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