Controlled release of gentamicin from calcium phosphate—poly(lactic acid-co-glycolic acid) composite bone cement

ARTICLE IN PRESS

Biomaterials 27 (2006) 4239–4249 www.elsevier.com/locate/biomaterials

Controlled release of gentamicin from calcium phosphate— poly(lactic acid-co-glycolic acid) composite bone cement Julia Schniedersa, Uwe Gbureckb,, Roger Thullb, Thomas Kissela a

Department of Pharmaceutics and Biopharmacy, Philipps-University Marburg, Ketzerbach 63, 35032 Marburg, Germany Department of Functional Materials in Medicine and Dentistry, University of Wu¨rzburg, Pleicherwall 2, 97070 Wu¨rzburg, Germany

b

Received 12 January 2006; accepted 21 March 2006

Abstract Modification of a self setting bone cement with biodegradable microspheres to achieve controlled local release of antibiotics without compromising mechanical properties was investigated. Different biodegradable microsphere batches were prepared from poly(lactic-coglycolic acid) (PLGA) using a spray-drying technique to encapsulate gentamicin crobefate varying PLGA composition and drug loading. Microsphere properties such as surface morphology, particle size and antibiotic drug release profiles were characterized. Microspheres were mixed with an apatitic calcium phosphate bone cement to generate an antibiotic drug delivery system for treatment of bone defects. All batches of cement/microsphere composites showed an unchanged compressive strength of 60 MPa and no increase in setting time. Antibiotic release increased with increasing drug loading of the microspheres up to 30% (w/w). Drug burst of gentamicin crobefate in the microspheres was abolished in cement/microsphere composites yielding nearly zero order release profiles. Modification of calcium phosphate cements using biodegradable microspheres proved to be an efficient drug delivery system allowing a broad range of 10–30% drug loading with uncompromised mechanical properties. r 2006 Elsevier Ltd. All rights reserved. Keywords: Bone cement; Hydroxyapatite; Biodegradation; Microspheres; Antimicrobial; Drug delivery

1. Introduction Conventional treatment of acute and chronic bacterial osteomyelitis includes surgical removal of necrotic bone tissue and repeated irrigations combined with high systemic doses of antibiotic drug substances for prolonged periods of time [1]. This treatment frequently causes unwanted side effects and often fails to cure bacterial bone infections. Therefore, drug delivery systems providing controlled release of antibiotics were studied for local therapy as an alternative [2]. Currently used local antibiotic delivery systems are based on poly(methyl-methacrylate) (PMMA) beads, chains or self setting PMMA bone cement [3]. A disadvantage of these materials is the need for surgical removal of the spent devices and the risk of induction of antibiotic resistance when insufficient doses are adminisCorresponding author. Tel.: +49 931 201 73550; fax: +49 931 201 73500. E-mail address: [email protected] (U. Gbureck).

0142-9612/$ - see front matter r 2006 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2006.03.032

tered [4]. In addition, PMMA cements are known to cause cytotoxic effects due to residual monomers, and tissue necrosis due to exothermic polymerization reaction temperatures exceeding 70 1C [5]. More recently drug delivery systems were developed from biodegradable and osteoconductive materials such as degradable polymers or calcium phosphate compounds [6]. Research efforts were focused on ceramic materials, e.g. sintered hydroxyapatite or self-setting calcium phosphate cements (CPC), as antibiotic carriers for the treatment of bone infections due to their similarity with the mineral phase of bone [7,8]. A major advantage of ceramic bone cements is their mechanical strength after setting to reinforce damaged bone structures and their application as injectable cement paste [9]. Calcium phosphate cement modification with various antibiotics (e.g. cephalosporin, gentamicin, vancomycin or tetracycline) was performed in the past by adding the water-soluble salts either to the liquid or solid cement phase during cement mixing [2,10–12]. This can cause problems due to an interaction

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of the water-soluble drug and the setting reaction of the cement after adsorption of the drug molecules at the particle surface [10,12]. This behavior was correlated with a higher cement porosity and lower degree of conversion or with the presence of chloride ions from the antibiotic salt (e.g. tetracycline–HCl), which formed a halogenide substituted hydroxyapatite as reaction product [13]. We hypothesized that interactions of antibiotics with cement setting could be prevented by microencapsulating drugs in biodegradable microspheres yielding cement/ polymer composites with satisfactory mechanical properties. Moreover, it was anticipated that diffusional release of antibiotics could be controlled by the microencapsulation. Polymers suitable for microsphere production are chitosan [14], coralline hydroxyapatite-gelatin [15] composites, and poly(lactic acid-co-glycolic acid) polymers (PLGA) [16]. PLGA was successfully used for microencapsulation of tetracycline [17], rifampicin [18] or vancomycin [19]. PLGA was used in this study due to its biocompatibility and biodegradability to physiological metabolites [20] and the ability to control degradation time by using different molecular weights or compositions. Degradation times of one month are useful to prevent antibiotic resistance due to insufficient polymer degradation or low drug concentrations, like it was described for PMMA beads [21]. Spray drying was used to generate biodegradable microspheres containing gentamicin crobefate as a poorly soluble drug substance to prolong the release rate and to reduce burst release [22]. The microspheres were then incorporated in various ratios into an apatitic cement matrix made from tetracalcium phosphate (TTCP) and dicalcium phosphate anhydrate (DCPA). The setting behavior, mechanical performance and the release kinetics of cement/microsphere composites were investigated and compared to the non-modified cement. Biocompatibility of the materials was tested with L929 fibroblasts using a MTT assay under in vitro conditions [23].

2. Materials and methods 2.1. Materials Biodegradable poly(lactic-co-glycolic acid), PLGA 50:50 with different molecular weights and hydrophilicity were purchased from Boehringer Ingelheim (Ingelheim, Germany) (see Table 1). DCPA and calcium carbonate were acquired from Malinckrodt-Baker (Griesheim, Germany). Gentamicin crobefate (EMD 46217) was a gift from Biomet Merck (Darmstadt, Germany). All other chemicals used in this study were of analytical grade. TTCP was prepared by sintering an equimolar mixture of DCPA and calcium carbonate at 1500 1C for 18 h as described previously [12] followed by quenching in air and 10 min dry grinding with a planetary ball mill (PM400, Retsch, Haan, Germany). DCPA was ground in ethanol for 24 h followed by drying in vacuum. Both cement components were mixed in the ball mill in an equimolar ratio with the addition of approx. 1 wt% sodium phosphate powder as setting accelerator.

2.2. Preparation of microspheres Drug loaded microspheres were prepared using a spray drying method [24]. A two-fluid nozzle with a diameter of 0.5 mm was used to manufacture placebo microspheres (MS) as well as gentamicin crobefate loaded microspheres (GC–MS) in a Bu¨chi 190 Mini Spray Dryer (Flawil, Switzerland). The settings were as follows: inlet temperature 48–50 1C, outlet temperature 32–39 1C, air flow 800 Nl/h, aspirator setting
30 mbar and feed rate 8–10 ml/min. MS batches were collected and subsequently dried at 25 1C in vacuum to remove residual organic solvents. Placebo MS were prepared from 5% (w/V) polymer solutions in 100 ml dichloromethane (DCM). For the GC–MS the drug was dissolved in three different concentrations (10, 20 and 30% (w/w) in a solution of methanol (MeOH) and DCM (1:4) and mixed with 100 ml of a 5% (w/V) polymer solution.

2.3. Characterization of microspheres Particle size and size distributions of the MS were measured in isopropanol using a laser particle sizer Horiba LA-300 (Kyoto, Japan) based on a laser light scattering technique. Approximately 100 mg were suspended in 200 ml isopropanol and sonicated for 15 min in an ultrasound bath (Sonorex RK100 H, Bandelin, Berlin, Germany). Each sample was measured in triplicate. The weighted average of volume distribution (D[4.3]) was used to describe the particle size.

Table 1 Summary of characteristics from placebo and gentamicin crobefate (GC) loaded microspheres (n ¼ 3) Microparticle batch

Polymer

Mw (Da)

Average particles size (mm)a

Yield (%)b

Drug loading (%)c

Tg (1C)

(I) Variation of polymers MS-1 MS-2 MS-3 MS-4 MS-5 MS-6

RG RG RG RG RG RG

502 H 502 503 H 503 504 H 504

14 000 14 500 28 000 35 000 53 500 55 000

7.170.4 9.070.0 13.070.6 14.470.1

6.7 10.7 30.3 39.7 27.0 14.0

— — — — — —

38.3 37.0 41.9 40.6 42.7 n.d.

(II) Variation of theoretical loading GC-1 RG 503 H GC-2 RG 503 H GC-3 RG 503 H GC-4 RG 503

28 000 28 000 28 000 35 000

9.270.3 7.770.5 5.470.2 10.171.0

36.0 37.4 41.4 34.9

10.0 20.0 30.0 10.0

41.4 41.6 42.8 42.0

a

Determined by laser diffractometry. 100 Wmp/(Wdrug+Wpolymer), determined gravimetrically. c Theoretical loading. d No result due to no particle formation. b

d d

ARTICLE IN PRESS J. Schnieders et al. / Biomaterials 27 (2006) 4239–4249 Encapsulation efficiency and drug loading were determined spectrophotometrically after extraction from the microspheres. Briefly, 10 mg of gentamicin crobefate loaded microspheres were dissolved in 1 ml DCM and subsequently 5 ml of ultra pure water was added. This mixture was placed in a rotating shaker (Rotatherm, Liebisch, Bielefeld, Germany) at 30 rpm and 37 1C for 2 h. After centrifugation the aqueous supernatant was analyzed spectrophotometrically using calibration curves on a Shimadzu UV 160 spectrophotometer (Shimadzu Europe Ltd/Deutschland GmbH, Duisburg, Germany) at 344 nm, with a sensitivity of 2 mg/ml and 20 mg/ml as limit of detection. Each sample was measured in triplicate. Encapsulation efficiency was expressed as follows:

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means of the Joint Committee for Powder Diffraction Studies (JCPDS) reference patterns for TTCP (PDF Ref. 25-1137), HA (PDF Ref. 09-0432) and DCPA (PDF Ref. 09-0080). Quantitative phase compositions of the materials were calculated by means of total Rietveld refinement analysis with the TOPAS software (Bruker AXS, Karlsruhe, Germany). As references database structures of TTCP, HA and DCPA were used together with a Chebychev forth-order background model and a Cu Ka emission profile.

2.6. Gentamicin crobefate release from microspheres and composites under in vitro conditions

Encapsulation efficiency ¼ ðactual drug loading=theoretical drug loadingÞ  100. The morphology of microspheres and cement devices was analyzed by scanning electron microscopy (SEM) using a Hitachi S 510 scanning electron microscope (Hitachi Denshi GmbH, Rodgau, Germany). Dried cement devices were cut using a razor-blade and then mounted on aluminum stubs using double-sided adhesive tape. The samples were sputter coated three times with a gold layer at 25 mA in an argon atmosphere at 0.3 mPa for 2 min (Sputter Coater S 150, Edwards/Kiese, BOC Edwards GmbH, Kirchheim, Germany) and analyzed with regard to surface morphology with accelerating voltage of 25 kV. Photographs were taken using a Pentax MX with a Pentax M 40 mm (1:2.8) objective (Tokyo, Japan). Glass transition temperatures (Tg) were measured using a DSC 7 differential scanning calorimeter (Perkin-Elmer, Wiesbaden Germany). Placebo and gentamicin crobefate loaded microsphere samples were sealed in aluminum pans and heated twice under nitrogen atmosphere. The resulting thermograms covering a range of
10–80 1C were recorded at heating rates of 10 1C/min. The second run was used for Tg calculation referring to the midpoint temperature. Calibration of the system was performed using an indium standard.

2.4. Preparation of cement cylinders Cement cylinders were prepared as described previously [12]. Briefly placebo or gentamicin crobefate loaded MS were mixed with the cement for 5 min using mortar and pistil in a 1:9 ratio(10% (w/w). Cement cylinders with 6 mm in diameter and 12 mm in height were obtained at a powder-liquid-ratio (P/L) of 3.3 using ultra pure water. Samples were prepared by mixing 800 mg of cement for the placebo devices and 80 mg of GC–MS added to 720 mg of cement for loaded devices with the required liquid volume in a nitrile rubber mixing container on a vibratory shaker Thermolyne Maxi Type 37600 (Barnstead International, Dubuque, Iowa, USA) for 15 s. The cement paste was transferred into stainless steel molds (6 mm diameter), and biaxially compressed using a self-made cantilever device at a pressure of 2.7 MPa for 5 s, followed by a load of 700 kPa for 2 h at 37 1C and 100% humidity to form cylinders with a height of 12 mm [25]. Specimen were removed from the moulds and stored in an incubator (GFL Type 1008, Burgwedel, Germany) at 100% humidity and 37 1C for additional 22 h prior to testing.

2.5. Characterization of cement cylinders Mechanical strengths of wet samples were measured at a crosshead speed of 1 mm/min after 24 h hardening at 37 1C, using a static mechanical testing device Zwick 1440 (Zwick, Ulm, Germany) with a 5 kN load cell. The initial setting time of the cements was measured in a humidity chamber at 37 1C and 490% humidity and in normal laboratory atmosphere (20–23 1C and 50–60% humidity), respectively using the Gilmore needle test with a needle of 113.98 g and 2.117 mm diameter according to ASTM standard [26]. X-ray diffraction patterns of the set cements were recorded on a D 5005 diffractometer (Siemens, Karlsruhe Germany). Data were collected from 2y ¼ 202401 with a step size of 0.021 and a normalized count time of 1 s/ step using Cu Ka radiation. The phase composition was analyzed by

Drug release was determined by suspending 10 mg of drug loaded microspheres or one cement device in 5 ml of PBS (pH 7.4). The particles were wetted prior to suspension using PBS containing 0.05% Myrj 52. The 15 ml Pyrex-vials were placed in an orbital shaker maintained at 37 1C and rotated at 30 rpm (Liebisch, Bielefeld, Germany). Samples were withdrawn at regular time intervals after centrifugation with 5000 rpm for 10 min using a laboratory centrifuge Sigma 203 (Sigma, Osterode am Harz, Germany) and replaced by fresh medium. The samples were analyzed spectrophotometrically at 344 nm with a Shimadzu UV 160 instrument (Shimadzu Europe Ltd/Deutschland GmbH, Duisburg, Germany).

2.7. In vitro cytotoxicity testing in L929 fibroblasts In vitro cytotoxicity of the polymers was evaluated using an MTT assay as described previously [23,27]. Briefly, L 929 cells were seeded into 96-well microtiter plates (NunclonTM, Nunc, Wiesbaden, Germany) at a density of 8000 cells/well. After 24 h the culture medium was replaced with 100 ml/ well of serial dilutions of polymer stock solutions in antibiotic-free DMEM (n ¼ 8). Polymer extracts were prepared by using up to 10 mg MS per milliliter medium. The extracts were incubated at 37 1C for 24 h in a Rotatherm apparatus. If necessary the extracts were neutralized and then sterilized by filtration. After an incubation period of 24 h MTT (3-(4,5dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide) was dissolved in phosphate buffered saline at 5 mg/ml and 20 ml was added to each well reaching a final concentration of 0.5 mg MTT/ml. After an incubation time of 4 h un-reacted dye was removed by aspiration and the purple formazan product was dissolved in 200 ml/well dimethylsulfoxide and quantified by a plate reader (Titertek Plus MS 212, ICN Biomedicals, Eschwege, Germany) at wavelengths of 570 and 690 nm. The relative cell viability related to control wells containing cell culture medium without polymer was calculated by [A] test/[A] control  100, with [A] as the concentration of viable cells. Poly(ethylene imine) 25 kDa was used as a positive control.

2.8. Calculations and analysis of data Data were collected in a Microsofts Excel 2000 database and results were presented as means and standard deviations of at least 3 experiments using the Origins 7.0 software. Significance between the mean values was calculated using ANOVA one-way analysis (GraphPad InStat 3.06, GraphPad Software, USA). Probability values pp0:05 were considered as significant.

3. Results and discussion 3.1. Microspheres preparation and characterization The microencapsulation of gentamicin crobefate using biodegradable PLGA was performed with a spray drying method as described previously [24]. Spray drying generates spherical particles with high encapsulation efficiencies compared to other encapsulation methods, such as

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20

MS-1 MS-2 MS-3 MS-4

18 16 14

Volume (%)

solvent evaporation [28]. High encapsulation efficiency and drug loading are requirements for composite formation with apatitic calcium phosphate cement as described later. Also particle sizes o15 mm were easily attained with narrow size distributions. The admixed MS will reside in the porous structure of the cement matrix with pore sizes in the 1–20 mm range, and to avoid destruction they need to be in the size range of 5–15 mm. As shown in Table 1 placebo and gentamicin crobefate loaded microspheres with a weight average of volume distribution (D[4.3]) in the range of 5–15 mm were obtained by spray drying. Monomodal particle size distributions were observed in all cases (Fig. 1). As expected, PLGA molecular weight affected both yields and particle sizes drastically. Microsphere yield increased from 6.7% up to 39.7% by increasing the Mw and also drug loading affected the yield in a positive manner. The yields of placebo and gentamicin crobefate loaded MS were in the expected range for small laboratory spray dryers [29,30]. With PLGA in the Mw range of 14–35 kDa both placebo and drug loaded particles yielded sizes in the range of 7.170.4 and 14.470.1 mm for the unloaded and 9.270.3 and 5.470.2 mm for the gentamicin-loaded MS, respectively. The particle size was affected by the viscosity of the feed suspension and both GC–MS and placebo MS showed comparable weighted averages of volume distribution (D[4.3]) and monomodal size distributions. Above M w 435; 000 Da the feed became too viscous and hence insufficient microsphere formation was noted. The encapsulation efficiency of GC–MS prepared by spray drying was found to be approximately 100% (w/w), e.g. the effective drug loading from batch GC-4 was 98.672.2% and similar to results obtained for ampicillin. By contrast, microencapsulation of gentamicin sulfate yielded much lower encapsulation efficiencies of 23.671.0% using a spray drying method at 10% loading level [31]. In this study gentamicin sulfate was incorporated into the feed not as solid but as an emulsion. Our results demonstrate that spray drying is a suitable technique for the preparation of GC–MS with appropriate size and high encapsulation efficiency. To verify the spray-drying results regarding microsphere size and distribution and to characterize particle morphology, SEM was employed and micrographs are presented in Figs. 2 and 3. Placebo particles produced with polymers with a lower Mw showed all a smooth surface and spherical particles morphology (Fig. 2A–D). MS prepared from PLGA with M w 435; 000 showed strong agglomeration and discrete particle formation could not be observed (Fig. 2E and F). This observation is not unexpected since homogenous spray formation is strongly influenced by the viscosity of the feed suspension. Based on results with different PLGA regarding MS sizes and morphology, GC–MS were manufactured with RG 503 H and RG 503, because this particle size range was considered as useful for the preparation of composites [32].

12 10 8 6 4 2 0 1

(A)

10

100

Particle size [µm] 20

GC-1 GC-2 GC-3 GC-4

18 16 14

Volume (%)

4242

12 10 8 6 4 2 0 1

(B)

10

100

Particle size [µm]

Fig. 1. Particle size distribution according to laser light scattering measurements of spray dried microspheres: (A) Placebo (MS) particles using matrix polymers with different molecular weights and (B) gentamicin crobefate (GS) loaded microspheres with different drug loadings.

With both types of PLGA the preparation of drug loaded MS by spray drying was easily achieved. The SEM micrographs shown in Fig. 3 allow an overview of the morphology of GC-loaded particles as a function of different drug loadings (10–30% (w/w)). All loaded batches showed spherical morphology with smooth particle surface, demonstrating that GC was completely entrapped with the polymer matrix and unencapsulated gentamicin was not attached to the surface. In comparison, SEM images from gentamicin sulfate-loaded MS published recently look similar, but do also show some indentations on the surface, possibly caused by the gentamicin sulfate [31]. The average size of particles was comparable to D[3,4]. Based on these results GC–MS were regarded suitable for the preparation of CPC. To minimize interactions between the hydrophilic gentamicin sulfate and PLGA, we selected a more

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Fig. 2. SEM micrographs of spray dried placebo microspheres (MS) batches made of six different 5% (w/v) polymer solutions: (A) MS-1: RG 502 H, (B) MS-2: RG 502, (C) MS-3: RG 503 H, (D) MS-4 RG 503, (E) MS-5: RG 504 H, (F) MS-6: RG 504.

lipophilic salt combination, namely gentamicin crobefate (Fig. 4). Using micro-encapsulation it was anticipated that GC would not interfere with the setting process of the bone cement. On the other hand gentamicin sulfate was shown to interact with PLGA carboxylic end groups [31]. The distribution of gentamicin crobefate in PLGA microspheres did not affect the glass transition temperatures (Tg) as shown in Table 1 suggesting that GC–MS consist of a solid dispersion morphology. The Tg values for placebo MS ranged from 41.5 to 42.0 1C and drug-loaded particles showed similar values.

3.2. In vitro release properties of gentamicin crobefate loaded microspheres Spray dried MS produced from RG 503 H and loaded with different amounts gentamicin crobefate, were used to investigate the drug release profiles as function of drug loading. For local treatment of bone infections, gentamicin implant and microspheres were studied either using the sulfate salt form [33] or the free base. While MS containing gentamicin base showed good encapsulation efficiencies, interactions between drug and PLGA caused stability

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Fig. 3. SEM micrographs of spray dried microspheres (MS) batches loaded with different amounts of gentamicin crobefate (GC) and made of RG 503 H and RG 503: (A) GC-1: 10% (w/w) GC in RG 503 H, (B) GC-2: 20% (w/w) GC in RG 503 H, (C) GC-3: 30% (w/w) GC in RG 503 H and (D) GC-4: 10% (w/w) GC in RG 503.

OMe

O

OMe E

H2 O3 PO O Fig. 4. Chemical structure of crobefate.

problems. With gentamicin sulfate-loaded particles, drug stability was not an issue, however, initial drug release (burst release) resulting from insufficient encapsulation and high water solubility of gentamicin sulfate were more problematic. GC, a poorly water-soluble salt of gentamicin, is commercially available as combination with gentamicin sulfate in collagen mats for the prevention of bone and tissue infections [22]. We hypothesized that this gentamicin salt might posses more controlled release characteristics than the sulfate, as it was shown in the literature [24]. The gentamicin crobefate release from PLGA microspheres was determined before and after the addition to the

calcium phosphate bone cement. As shown in Fig. 5, the drug release from PLGA microspheres under in vitro conditions was characterized by a triphasic drug release kinetic (Fig. 5A) as often described in the literature [34]. The initial burst ranged from 6.574.0% up to 26.270.4% in the first 24 h depending on drug loading. As expected higher drug loading caused higher drug bursts, but the values are much lower than those in gentamicin sulfate microparticles where bursts up to 60% were obtained at lower loadings [31]. The second phase (plateau-phase) lasted approximately 7–12 days with only low doses released during this period of time followed by the erosion phase, in which polymer degradation occurred and drug release could be observed. The plateau-phase decreased with increasing drug loading, in accordance with the generally postulated pore diffusion mechanism. By contrast, GC release from MS/cement composites showed nearly a zero-order kinetic, which was characterized by a slower, but linear release over 100 days without initial drug burst (Fig. 5B). Burst release was strongly reduced for all composites to values below 2% in the first 24 h as shown in Fig. 6. This can be explained by the lower water solubility of GC and the slower drug diffusion out of the cement matrix. The linear release profile can be attributed to the embedding of the MS into the cement

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30

GC-1: 10 % GC in RG 503 H GC-2: 20 % GC in RG 503 H GC-3: 30 % GC in RG 503 H

25

Gentamicin release from MS Gentamicin release from MS after embedding into cement

20

Burst release (%)

Cumulative release (mg)

4

3

2

15

10

5

1

0

0

GC-1 10 %

0

20

40

60

80

5

GC-1: CPC + 10% (w/w) GC GC-2: CPC + 10% (w/w) GC GC-3: CPC + 10% (w/w) GC

4

Cumulative release (mg)

GC-3 30 %

Fig. 6. Gentamicin crobefate burst release from biodegradable microspheres with 10%, 20% and 30% drug before and after embedding in calcium phosphate cement. Values obtained after 24 h of incubation (n ¼ 3).

Time (d)

(A)

GC-2 20 %

100

Table 2 Gentamicin crobefate (GC) concentrations released from microspheres after admixing to calcium phosphate bone cement (n ¼ 3)

3

Batch

GC released after 8 h

x-fold higher then MIC (1–2 mg/ml)

GC released after 50 d

x-fold higher then MIC (1–2 mg/ml)

GC-1 (10%) GC-2 (20%) GC-3 (30%)

0.470.2 3.070.2 7.371.2

1/5–1/2.5 3–1.5 7–3.5

79.3710.4 128.073.0 124.0711.6

80–40 130–65 120–60

2

1

0 0

(B)

20

40

60

80

100

Time (d)

Fig. 5. In vitro release profiles from poly(lactide-co-glycolide) microspheres before (A) and after (B) embedding in cement devices. Three different drug loadings of gentamicin crobefate (GC) are shown (-’-)GC 10% (w/w), (-K-)GC 20% (w/w), (-m-)GC 30% (w/w) (n ¼ 3).

matrix. Another reason might be the adsorption of the released drug on the cement matrix, which might also prolong the release kinetics [35]. All these factors contributed to the linear kinetic profile compared to the triphasic release profile from the plain drug loaded particles. Extremely high release rates of antibiotics were observed in studies published in the literature, where 55–95% of drug release occurred between 3 and 7 d, independent of the type of the cement, namely apatite or brushite [33]. To our knowledge, this is the first example of an antibiotic cement composite demonstration both low reduced drug burst and linear release kinetics. The minimal inhibition concentration (MIC) of gentamicin in vitro could be reached at all time points during the in vitro release study. All composites released much higher concentrations of GC than the required concentration to eradicate bacteria above the MIC of o1–2 mg/ml (Table 2).

These results demonstrate the feasibility a drug delivery system based on cement/MS composites. 3.3. Mechanical properties of hardened microsphere/cement composites Ceramic materials, such as hydroxyapatite have received increasing interest recently [8]. The addition of antibiotics to calcium phosphate bone cements has been reported earlier [13]. The main drawback of this approach seems to be the negative effects of antibiotics on mechanical properties of cement composites, causing a decrease in compressive strength. We speculated that isolation of antibiotic by microencapsulation into a biodegradable polymer should prevent interactions between the cement and antibiotic during the setting phase and hence lead to composites with sufficient mechanical strength. Properties of the composites such as setting time and compressive strength with and without the addition of antibiotic-loaded MS are displayed in Fig. 7A. The MS free cement showed an initial setting time of 5 min. Addition of unencapsulated gentamicin crobefate (1% (w/w)) abolished cement setting, since GC is a very effective inhibitor of HA crystal growth. This is in agreement with other gentamicin

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MS-3 MS-4 GC-1 GC-2 GC-3 CPC GC GS 1% GS 2% GS 3% 0

5

10

(A)

15

20

25

Setting time (min)

90 80

**

Compressive strength (MPa)

70 60 50 40 30 20 10 0

(B)

MS-1

MS-2

MS-3

MS-4

MS-5

90 80

**

** *

Compressive strength (Mpa)

70 60 50 40 30 20 10 0

(C)

GC-1

GC-2

GC-3

CPC

GC

GS

Fig. 7. Basic characteristics of cement devices after mixing with placebo and antibiotic loaded microspheres: (A) setting time, (B) compressive strength with different polymers used for microsphere batches, (C) compressive strength regarding different drug loadings. MS ¼ placebo particles, GC ¼ Gentamicin crobefate-loaded particles, CPC ¼ calcium phosphate cement, GS ¼ gentamicin sulfate, added only as solution. (*p0.05; **p0.005).

salts [12]. By admixing gentamicin sulfate in increasing amounts to an apatitic cement, setting of the cement was achieved but at increased setting times up to 26 min. In contrast, no change of setting times was noted for the addition of placebo MS batches MS-3, MS-4 or drug-

containing batches GC-1 and GC-2. A slightly, but still acceptable increase of the initial setting time for GC-3 to 8 min was observed. In the literature methods were described, which use prolonged setting times as a means to improve fitting of the cement paste to the defect site. The setting time, however, should not exceed 12–15 min for reasons of clinical handling [36]. At the same time, the compressive strength of cements mixed with MS of up to 30% (w/w) GC was not affected compared to the control group with 52.0875.1 MPa (Fig. 7C). The compressive strength increased significantly (pp0:05 and pp0:005) up to 70 MPa by the addition of drug-loaded MS to the cement, independent of the batch. No significant differences between the three GC-loaded groups could be seen. The compressive strength was affected by the polymers of the admixed particles with regard to the Mw. A significant (pp0:05) influence on compressive strength was seen by using different polymers. With increasing the polymer molecular weight an increase in compressive strength of the tested cement samples was observed (Fig. 7B). It is known from the literature, that PLGA with high molecular weights are used for tissue engineering due to their better stability, also in combination with CPC [37]. The limit of MS addition with respect to mechanical properties is reached at about 20% (w/w) drug-loaded particles to the cement which then resulted in a decrease of compressive strength to 28.13711.37 MPa (data not shown) after addition of 30% (w/w) drug-loaded particles. After exhaustive setting, the cement/MS composites were analyzed using an X-ray diffraction technique to study the conversion to HA. Results of the phase analysis according to Rietveld refinement analysis of the cement after setting are given in Table 3. All materials converted to nanocrystalline hydroxyapatite (76–86% (w/w)) within 24 h with crystal sizes of approximately 13–17 nm. The latter results are based on peak broadening in X-ray diffraction analysis. A minor phase of non-reacted TTCP (10–22%) could be detected while the DCPA reactant was nearly completely consumed such that the precipitated HA is thought to be of calcium deficient composition. Finally, the cement/MS composites were analyzed using SEM to investigate the structure and integrity of MS embedded in the cement composite. Additionally, the degradation of the biodegradable MS was demonstrated after incubation under in vitro conditions. Therefore, cross sections of the composites before and after the in vitro study were taken. While initially the MS are not damaged and remain intact (Fig. 8A), they were completely degraded after 100 d of incubation and are not visible at the cross section site (Fig. 8B). This confirms that GC–MS released their drug load completely. 3.4. Biocompatibility study under in vitro conditions Different batches of placebo MS (3,4,5,6) and GC containing MS (GC-4) were subjected to biocompatibility

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Table 3 Phase composition of cements according to Rietveld refinement analysis Sample

Microparticle batch

Polymer

Drug loading (%)

HA % (w/w)

TTCP % (w/w)

DCPA % (w/w)

Size HA (nm)

Ref. (no GC)

—

—

—

84.8

13.9

1.3

17.270.3

Variation of polymers CPC+MS-3 MS-3 CPC+MS-4 MS-4

RG 503 H RG 503

— —

76.1 79.0

22.8 19.5

1.1 1.5

16.170.5 15.370.3

Variation of theoretical drug loading CPC+GC-1 GC-1 CPC+GC-2 GC-2 CPC+GC-3 GC-3 CPC+GC-4 GC-4

RG RG RG RG

10 20 30 10

85.0 85.9 86.0 79.0

14.7 9.7 9.8 18.4

0.3 4.5 4.2 2.6

14.870.3 13.670.3 14.470.4 15.470.4

503 H 503 H 503 H 503

By the addition of free gentamicin crobefate to the CPC, no setting occurred, therefore no results in phase composition are shown here.

120

Cell Viability (%)

100

MP-3 MP-4 MP-5 MP-6 GC-4

80

60

40

20

0 0

2

4

6

8

10

Concentration of Microparticles (mg/ml)

Fig. 9. Cytotoxicity of placebo and gentamicin crobefate containing microspheres to L 929 fibroblasts as measured by MTT assay. Cells were incubated with increasing concentrations of particle-extracts for 24 h. No significantly differences were detectable (n ¼ 8).

microspheres/ml incubation medium both placebo and drug-loaded MS did not show signs of impaired cell viability. A slightly increase of cell viability was observed for GC-4 batch due to a growth stimulatory effect of the added gentamicin. These results are not unexpected since both PLGA MS and calcium phosphate cement are FDA approved and GC is used clinically [38]. Fig. 8. SEM micrographs of calcium phosphate cement devices embedded with spray dried microsphere batch GC-1made of RG 503 H (PLGA) polymer with 10% (w/w) gentamicin crobefate: (A) before and (B) after 100 d of incubation in PBS, 37 1C, 30 rpm.

testing under in vitro conditions. The materials were incubated with cell culture medium according to USP XXV for 24 h. Serial dilutions of these extracts were studied in a standard cell line of mouse fibroblasts (L 929) with positive and negative controls. Cell viability was tested using a MTT assay after 24 h of incubation as shown in Fig. 9. Even at extract concentrations as high as 10 mg

4. Conclusions The feasibility of a self-setting bone cement composite containing biodegradable gentamicin crobefate-loaded microspheres was demonstrated in this study. Drug-loaded microspheres were produced by spray drying with a monomodal distribution and a high encapsulation efficiency of nearly 100%, which reduced initial drug burst. Gentamicin crobefate release was controlled by a combination of pore diffusion and polymer erosion, which resulted in zero-order release kinetics for cement/microsphere composites. Moreover, no change of the setting properties

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and mechanical performance of the cement was observed due to an effective decoupling of cement setting and drug release. Composites made from calcium phosphate cement and drug-loaded microspheres could be of interest in prevention and treatment of bone infections. A significant advantage of this drug delivery system providing local antibiotic release would be its biodegradability avoiding secondary surgery.

[15]

[16]

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Acknowledgments The authors wish to thank Biomet Deutschland GmbH, Berlin, for providing the gentamicin crobefate. The authors would also like to thank the European Society of Biomaterials for the Student poster award 2003 and the related grant cordially.

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