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Nanostructured material formulated acrylic bone cements with enhanced drug release
Materials Science and Engineering C 58 (2016) 233–241
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Nanostructured material formulated acrylic bone cements with enhanced drug release Shou-Cang Shen a,⁎, Wai Kiong Ng a, Yuan-Cai Dong a, Junwei Ng a, Reginald Beng Hee Tan a,b,⁎⁎ a b
Institute of Chemical and Engineering Sciences, A*STAR (Agency for Science, Technology and Research), 1 Pesek Road, Jurong Island, Singapore 627833, Singapore Department of Chemical and Biomolecular Engineering, The National University of Singapore, 4 Engineering Drive 4, Singapore 117576, Singapore
a r t i c l e
i n f o
Article history: Received 6 February 2015 Received in revised form 19 June 2015 Accepted 11 August 2015 Available online 19 August 2015 Keywords: Bone cements Antibiotic Nanoparticles Drug delivery Diffusion network
a b s t r a c t To improve antibiotic properties, poly(methyl methacrylate) (PMMA)-based bone cements are formulated with antibiotic and nanostructured materials, such as hydroxyapatite (HAP) nanorods, carbon nanotubes (CNT) and mesoporous silica nanoparticles (MSN) as drug carriers. For nonporous HAP nanorods, the release of gentamicin (GTMC) is not obviously improved when the content of HAP is below 10%; while the high content of HAP shows detrimental to mechanical properties although the release of GTMC can be substantially increased. As a comparison, low content of hollow nanostructured CNT and MSN can enhance drug delivery efficiency. The presence of 5.3% of CNT in formulation can facilitate the release of more than 75% of GTMC in 80 days, however, its mechanical strength is seriously impaired. Among nanostructured drug carriers, antibiotic/MSN formulation can effectively improve drug delivery and exhibit well preserved mechanical properties. The hollow nanostructured materials are believed to build up nano-networks for antibiotic to diffuse from the bone cement matrix to surface and achieve sustained drug release. Based on MSN drug carrier in formulated bone cement, a binary delivery system is also investigated to release GTMC together with other antibiotics. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Due to the excellent tissue compatibility of poly(methylmethacrylate) (PMMA) and the rapid setting with achieved mechanical strength, PMMA bone cement has been widely used clinically to fix total joint prostheses (hip and knee replacements) to bone for decades. It is an integral part of almost all total knee replacements and also used in at least twothirds of total hip replacements [1]. With the rapid growth of the aged population, there is a dramatic increase in the burden of musculoskeletal disease, over 1,000,000 total hip and knee replacements are performed each year in the US alone. Post-operation infection is a clinical catastrophe, leading to the need for more complex surgery with much higher costs or the increase of morbidity [2]. As with all implant materials, bone cement carries an elevated risk of infection when implanted into the human body because of the possibility microorganism biofilms to form on an inert surface [3–6], usually requiring multiple surgeries for treatment [7,10]. To reduce this risk, an antibiotic (such as gentamicin) has been formulated with bone cement; however, it is known that as the antibiotic is mostly trapped within the substance of the cement,
⁎ Corresponding author. ⁎⁎ Correspondence to: R.B.H. Tan, Institute of Chemical and Engineering Sciences, A*STAR (Agency for Science, Technology and Research), 1 Pesek Road, Jurong Island, Singapore 627833, Singapore. E-mail addresses: [email protected] (S.-C. Shen), [email protected] (R.B.H. Tan).
http://dx.doi.org/10.1016/j.msec.2015.08.011 0928-4931/© 2015 Elsevier B.V. All rights reserved.
only a small proportion is available for diffusing out to the desired site for antibacterial activity [8]. The therapeutic antimicrobial activity drops rapidly below effective therapeutic levels after surgery within several days [9] and did not show significant advantages in reducing infection rate [10]. Therefore, it is desired to develop functional antibiotic bone cement that can enhance antimicrobial activity and/or prolong the therapeutic levels of antibiotic formulated in bone cement [11–13]. The sustained antimicrobial activity would be of significant impact in improving the results of joint replacement surgeries. Researchers have explored many approaches to improve the antimicrobial activity of bone cement. It has been reported that the presence of chitosan nanoparticles could effectively prevent viable bacteria from surviving on the surface of the bone cement [14] and the addition of silver nanoparticles to increase its antimicrobial activity [15]. However, those active ingredients cannot be released to protect surrounding tissues and often have adverse cytotoxic effects [16,17]. Many efforts have been made to improve the release of antibiotics from bone cement for applications in total joint replacement, including vacuum-mixing of commercially available antibiotic-impregnated bone cement, but the effects are limited [18–20]. M. Vallet-Regi et al. reported to enhance the release of GTMC from PMMA based bone cement by formulation with HAP [21,22]. Although the drug elution from the PMMA bone cement was enhanced, the high content of fillers could impair mechanical strength. MSN has been formulated with acrylic bone cement and the resulting mechanical behaviors were characterized [23]. Ormsby et al. [24] studied the incorporation of CNT into acrylic based bone cements
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and reinforced mechanical properties were found. However, the effect of these nanoparticles on delivery of antibiotics has rarely been studied. The sustained release of antibiotics from bone cement is essential to prevent the formation of biofilm on implanted bone cement and prosthesis parts. Thus, it is necessary to develop antibiotic bone cements with enhanced and prolonged antimicrobial activity to protect the surrounding tissues, as the operation sites are prone to infection for up to 6 to 8 weeks after surgery and usually need at least 6 weeks of parenteral antibiotic administration [25]. In this study, nanostructured materials, such as CNT, HAP nanorods and MSN, are formulated into PMMA based bone cement to enhance the release of antibiotics. The presence of hollow nanostructured particles in bone cement builds up diffusion networks and improves the loaded antibiotics to diffuse from bone cement to surface. Moreover, novel bone cement enabled effective delivery of gentamicin in combination with other antibiotics, such as vancomycin.
Table 1 Composition of commercial bone cements.
Powder Methylmethacrylate methacrylate copolymer Methyl methacrylate–stylene copolymer Zirconium dioxide Barium sulfate Benzoyl peroxide Gentamicin Liquid Methylmethacrylate N,N-dimethyl-p-toluidine Hydroquinone
Simplex P (w/w%)
Smartset HV (w/w%)
Smartset GHV (w/w%)
15.00
84.00
80.46
73.72
–
–
15.00
14.37
1.00 –
0.96 4.22
97.50 2.50
97.50 2.50
10.00 1.28 – 97.49 2.50 0.0075
% by weight (w/w) of powder component and liquid component.
2. Experimental 2.1. Materials Mesoporous silica nanoparticles (MSN) were prepared using fluorocarbon-surfactant-mediated synthesized as reported by Han et al. [26]. Typically, 0.5 g of Pluronic P123 and 1.4 g of FC-4 were dissolved in 80 ml of HCl solution (0.02 M), followed by the introduction of 2.0 g of TEOS under stirring. The solution was continuously stirred at 30 °C for 24 h and then transferred into a polypropylene bottle and kept at 100 °C for 24 h. The resultant solid was recovered by centrifuging and washed with deionized water twice, subsequently it was dried at 55 °C for 12 h. To remove the template molecules, the material was heated from room temperature to 550 °C at a heating rate of 2 °C/min and followed by calcination in air for 6 h. Nanorods of hydroxyapatite (HAP) were prepared by wet gel steaming method [27]. Typically, 15 ml of 0.5 M Ca(NO3)2·4H2O (Alfa Aesar) solution was added to 15 ml of 0.5 M (NH4)2HPO4 (Alfa Aesar) solution to achieve the desired Ca2+/PO3− ratio of 1.0 for precipitation 4 of di-calcium phosphate dihydrate (CaPO4·2H2O, Ca: P = 1:1) at 25 °C. The pH of the mixture was adjusted to 11.0 by using aqueous ammonia solution (NH4OH, 25 wt.%, Merck). The precipitate solid was recovered by filtration and then transferred to a 25 ml beaker, which was placed in an autoclave with a polytetrafluoroethylene (PTFE) liner. 5 ml of the ammonia solution was poured into the bottom of the PTFE cup and physically separated from the solid sample. The autoclave was then placed into an oven at 180 °C for 20 h. After the steam-assisted treatment, the resulting white solid was dispersed in water using ultrasound and recovered by centrifugation. The washing procedure was repeated twice. The obtained solid material was dried in an oven at 55 °C, and finally ground to a fine powder. The multiwall carbon nanotubes (CNT) were provided by Shenzhen Nanotech Port Co., Ltd. Three kinds of commercial bone cements were used: CMW Smartset-HV and CMW Smart GHV (DePuy International Ltd. UK) and Simplex P (Stryker Co, UK). The compositions of the bone cements, as indicated in the products' brochure, are shown in Table 1. 2.1.1. PBS buffer Buffered saline solution pH 7.4 was prepared as described in the British Pharmacopoeia (BP A79): anhydrous di-sodium hydrogen phosphate 2.38 g; anhydrous potassium hydrogen phosphate 0.19 g and sodium chloride 8 g to 1000 ml with deionized water (Milli-Q). 2.1.2. o-Phthaldialdehyde reagent o-Phthaldialdehyde reagent was prepared according to a procedure reported by Zhang et al. [28]. It was formulated by adding 2.5 g of o-phthaldialdehyde, 62.5 ml methanol and 3 ml of 2-mercaptoethanol to 560 ml of 0.04 M sodium borate in a deionized water solution. The
reagent was stored in the darkness and settled for at least 24 h prior to use. 2.2. Preparation of antibiotic-loaded bone cements GTMC was loaded onto the nanostructured materials (MSN, HAP or CNT) by wet impregnation. Typically, 0.20 g of GTMC was dissolved in 3 ml deionized water. 0.30 g of MSN powder was impregnated with GTMC solution under stirring and aged for 24 h. The mixture was dried under vacuum at room temperature under vacuum for 48 h. The dried GTMC loaded nanoparticles were ground to fine powder and a certain amount of GTMC loaded nanoparticles were mixed with commercial bone cement solid powder by manual grinding. When HAP and CNT are used as drug carriers, the same impregnation method was used and the drug-to-carrier ratio may be adjusted to avoid too high drug loading in the final formulation. Samples of A-1–A-9 prepared by using GTMC–MSN with different drug loadings are listed in Table 2. The samples of antibiotic loaded bone cement were prepared by mixing the powder with the liquid monomer in a ratio of 2 g/ml in a beaker in a laminar flow hood, in accordance with the manufacturer's instruction. Monomer liquid was added to the polymer–GTMC–MSN mixture in a bowl and was stirred using a spatula until the powder was fully wetted. The soft mixture was inserted into the mold with dimensions of 6 mm in diameter and 12 mm in height. The filled mold was pressed between two glass plates to harden overnight at room temperature. The hardened bone cement cylinders were pulled out of the mold and stored under sterile conditions at room temperature for the in-vitro drug release test and compression test. In addition, rectangular beams with dimension of 25 × 10 × 2 mm were prepared for a bending test and an antibacterial property test. 2.3. In vitro drug release study The drug release study was conducted by soaking two cylinder samples of each composition in 5 ml PBS buffer (pH 7.2). The sample Table 2 Composition of nanomaterial formulated antibiotic bone cements.
A-1 A-2 A-3 A-4 A-5 A-6 A-7 A8 A9
Simplex-P powder (g)
MMA (ml)
Filler %
Drug%
1.90 1.90 1.80 1.70 1.60 1.50 2.0 1.0 2.0
1.0 1.0 1.0 1.0 1.0 1.0 1.0 1.0 1.0
MSN 0 MSN 2.04 MSN 4.08 MSN 6.12 MSN 8.15 MSN 10.19 HAP 11.56 HAP 32.33 CNT 5.36
3.40 1.36 2.72 4.08 5.44 6.79 3.59 4.85 3.21
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Fig. 1. TEM images of (a) MSN, (b) nanorods of HAP and (c) CNT.
was put in an incubator shaker operated at 37 °C and 40 rpm. The release medium was withdrawn at predetermined time intervals, and replaced with fresh PBS buffer (5 ml) each time. The accumulative amount of gentamicin released was calculated based on the initial weight of the bone cement cylinder and the drug content. The gentamicin release was studied for 80 days. An indirect method was used for measurement of gentamicin concentration by a UV–vis spectrophotometer (Cary 50, Agilent Technologies) because gentamicin does not absorb ultraviolet nor visible light. The o-phthaldialdehyde was used as a derivatizing agent to react with the amino groups of gentamicin and yield chromophoric products. The reaction was carried out making 1 ml of our problem gentamicin in solution react with 1 ml of isopropanol (to avoid the precipitation of the products formed) and 1 ml of o-phthaldialdehyde reagent. After fully mixed, the concentration of gentamicin sulfate was determined by UV absorbance at 332 nm. 2.4. In vitro cytotoxicity assay 3T3 mouse fibroblast cells (3T3-Swiss albino, ATCC) were cultured in a complete growth culture medium in a 5% CO2 incubator. The
complete growth culture medium was prepared with Dulbecco's Modified Eagle's Medium (DMEM) supplemented with 10% fetal bovine serum, 1 mm L-glutamine and penicillin (100 U/ml). Cell viability testing was carried out via the reduction of the MTT reagent (3-[4,5-dimethyl-thiazol-2-yl]-2,5-diphenyltetrazolium bromide, Sigma). This assay provides a simple method for determining comparative cell viability using a standard microplate absorbance reader. The MTT assay was performed with the bone cement substrates placed at the bottom of a 96-well plate following the standard procedure with minor modifications. Control experiments were carried out using the complete growth culture medium only (serving as the non-toxic control) and 1% Triton X-100 (Sigma) (as the toxic control). 3T3 fibroblasts in the complete growth culture medium (100 μl) were seeded at a density of 104 cells/well in a 5% CO2 incubator for 24 h. The culture medium from each well was then removed and 100 μl of medium and 20 μl MTT solution (5 mg/ml in PBS) were then added to each well. After 4 h of incubation at 37 °C and 5% CO2, the media were removed and the formazan crystals were solubilized with 100 μl dimethyl sulfoxide (DMSO, Sigma) for 15 min. The optical absorbance was then measured at 570 nm on a microplate reader (Tecan GENios). Six samples were tested for each type of bone cement.
Fig. 2. Gentamicin release profiles of PMMA-based bone cement formulated with MSN (A-6), HAP (A-7, A-8) and CNT (A-9) as compared with antibiotic bone cement Smartset GHV.
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2.5. Testing of mechanical property The three point bending test was performed on the Instron Universal Materials Testing Machine (Model 5544). According to the standard test method of ASTM D790-3, the span length was 20 mm and the loading rate was 1 mm/min. The bending modulus (EB) was calculated according to the following equation: 3
EB ¼ L3 m=4bd ; where L is the support span (mm), b is the width of beam tested (mm), d is the depth of beam tested (mm), and m is the slope of the tangent to the initial straight-line portion of the load–deflection curve (N/mm). The compression tests were carried out on the bone cement cylinders with the same dimension as those for drug release investigation [29]. The compression force was applied along the axis using a crosshead speed of 5 mm/min. The compression strength was calculated from the obtained load–deformation curves. The compression strength (CS, MPa) was calculated using following equations: CSðMPaÞ ¼ FðNÞ=A m2 =1000; 000 where F is the applied load (N) at the highest point of the load–deflection curve and A is the cross-section area (m2) of the sample tested. 2.6. Characterization High resolution TEM images of nanostructured materials were taken by a TECNAI F20 (G2) (FEI, Philips Electron Optics, Holland) electron microscope at 200 kV. Nitrogen adsorption/desorption isotherms were measured by using an Autosorb-6B gas adsorption analyzer (Quantachrome) at a temperature of −196 °C. Before nitrogen adsorption–desorption measurements, each sample was heated at 40 °C under vacuum for 24 h. The specific surface areas of the samples were determined from the linear portion of the Brumauer–Emmett–Teller (BET) plots. The total pore volume, VT, was estimated from the amount adsorbed at a relative pressure id P/P0 of 0.95. 3. Results and discussion Fig. 1 shows the TEM images of nanostructured materials used as drug carriers and bone cement fillers in this study. It can be seen that the mesoporous silica nanoparticles (MSN) have orderly arranged pore channels (Fig. 1a) and carbon nanotubes (CNT) show a hollow structure with a single channel at the nanoscale (Fig. 1c). As a comparison, no porous structure can be observed in the HAP nanoparticles as the pore volume of HAP measured by N2 adsorption is 0.15 cm3/g only [27]. Fig. 2 displays the cumulative release profile of GTMC from different nanostructure drug carriers loaded with GTMC formulated in the Simplex-P bone cement and compared with the commercial antibiotic bone cement Smartset GHV. It is observed that the commercial antibiotic bone cement exhibits low GTMC release rate. Smartset GHV only shows about 5% of release on the first day and the release of antibiotic on following days is negligible. This result is consistent with reported investigation of commercially available GTMC-loaded acrylic bone cement and it was postulated that the elution mechanism of GTMC in this type of cement is a surface phenomenon [30]. Only surface adhered GTMC particles could be released. To enhance drug delivery efficiency, three types of nanostructured materials are used as drug carrier and formulated to bone cement. It is found that the presence of 5.36% of CNT (A-8) in functional bone cement can lead to more than 75% of GTMC release in 60 days and the formulation with 10.19% of MSN (A-6) also enhances the total drug release to 60%. As a comparison, the presence of 11.56% nonporous HAP nanoparticles only slightly enhances the total drug release to about 10%. When the content of the HAP nanoparticles
Fig. 3. Load–displacement curves of formulated bone cements.
in formulated bone cement is increased to 32.33%, 75% of GTMC could be released. This result is in agreement with literatures reported that the presence of about 30% of hydroxyapatite led to about 90% in 700 h [22]. However, the effect of high HAP content on mechanistic property was not reported. Fig. 3 shows the load–displacement curves of bone cement formulated with different nanostructured materials and compared with commercial Simplex-P bone cement. It is shown that the compression strength of MSN formulated bone cement (A-6) is well preserved as compared with the original Simplex-P, where the load–displacement curves are almost unchanged. As a comparison, the compression strength of HAP formulated bone cement (A-8) is reduced by about 50% although the high content of HAP could enhance the drug release. The large amount of HAP (32%) in bone cement formulation may cause the formation of a large void or a porosity structure [22], thus seriously reducing its compression strength. Moreover, the CNT formulated bone cement can only preserve about 10% of compression strength of the original Simplex-P bone cement. The detriment to biomechanical properties of bone cement limited the application of the nanostructured HAP and CNT in formulation of antibiotic bone cements for effective drug delivery. Thus, in this study, MSN was selected for further investigation. Fig. 4 displays the N2 adsorption–desorption isotherms of MSN and GTMC loaded MSN. The large pore volume of MSN (1.01 cm3/g) exhibits a high capacity of N2 adsorption as high as 511 cm3/g, STP at P/P0 of 0.98 and the sharp pore size distribution peak at 6 nm indicates its uniform pore diameter within the silica nanoparticles. When MSN is loaded with 40% of GTMC, the adsorption of N2 reduced significantly at whole
Fig. 4. N2 adsorption isotherms of (a) MSN (inset is the pore diameter distribution), (b) GTMC (40 wt.%) impregnated MSN (60 wt.%), and (c) solid powder of Simplex-p.
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Fig. 5. Cumulative GTMC release profile from Simplex-P bone cement formulated with MSN at difference loadings.
range of P/P0 and the total pore volume shrinks to 0.23 cm3/g. The results imply that the most pore volume of uniform pore channels is occupied by GTMC molecules. As a comparison, the PMMA based bone cement powder is a non-porous material as it showed negligible N2 adsorption. The nonporous structure of PMMA based polymers should be attributed to the low drug release efficiency as drug molecules are embedded in the condensed polymer based bone cement without porous diffusion pathways. When MSN loaded GTMC is formulated with the PMMA based polymer, the porous structure is used as drug carriers. Once the drug molecules are released to the medium, the void of uniform arranged mesoporous channels is expected to play a role as diffusion pathways for GTMC located at the inner matrix of bone cement to diffuse out. However, the content of MSN in the bone cement matrix can affect the efficiency of diffusion network. Fig. 5 shows the cumulative release profile of GTMC from MSN incorporated bone cement with different contents of MSN and GTMC. It is observed that, for the sample prepared by direct mixing of GTMC and bone cement powder, the release of GTMC is very limited. Only about 3% of GTMC released is observed in first day of immersion in PBS solution. No GTMC release is detectable in the following 80 days of investigation. When the sample is formulated with GTMC loaded MSN, the release from MSN-bone cement composite can be enhanced. However, the release rate is obviously affected by the content of MSN in the final composite. For the A-2 and A-3 samples with MSN contents of 2.04 wt.% and 4.08 wt.%, the release of GTMC only reaches 5% in the first day of immersion and the release stopped during the following investigation. The large portion of GTMC is still retained inside of the PMMA based bone cement matrix and no obvious enhancement on drug release character as compared with direct blending of GTMC and PMMA based bone cement. When MSN content is increased to 6.12 wt.% in the formulation, the release of GTMC from the sample of A-4 is significantly enhanced. After 7.3% of release in the first day, a sustained release of GTMC from MSN
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incorporated bone cement is observed in 80 days and reaches about 40% of release. A further increase of the content of MSN to 8.15 wt.% (A-5) results in a sustained release to 58% of GTMC in 80 days. When the MSN content is increased to 10.19 wt.%, the A-6 sample exhibits a faster release rate of GTMC in the middle stage between 10 and 30 days than that of A-5, and finally they reach a similar percentage of GTMC release after 80 days of immersion in PBS buffer. The MSN nanoparticles incorporated in bone cement is believed to form a nano-network pathway for GTMC diffusion from composite to surface and released to the medium. As shown in Fig. 6a, for GTMC formulated bone cement in the absence of MSN nanoparticles, GTMC particles embedded inside the PMMA matrix during polymerization cannot diffuse to the surface of bone cement for release. For the GTMC–MSN loaded bone cement composite, most of GTMC was entrapped inside the mesoporous channels of the MSN rod-like nanoparticles. The critical content of MSN inside bone cement is required to build up the network for GTMC to diffuse from the matrix to the medium. When the content of MSN is below 6 wt.%, most of the GTMC loaded MSN particles are isolated and also embedded in the bone cement matrix as shown in Fig. 6b. The release of GTMC from the bone cement composite cannot be obviously improved. Only those GTMC loaded MSN particles on the shadow surface of bone cement contribute to drug release. Once the MSN content increased to 6 wt.% or above, a nano-network can be built up through “particle–particle” contact (Fig. 6C) and allow the GTMC molecules to continuously diffuse from the PMMA based matrix to surface and be released to the medium, thus the release profile of GTMC is obviously improved as indicated in Fig. 5. When MSN content is increased to 8.15 wt.%, the diffusion network is more effective and the drug release is significantly enhanced. While the content of MSN is further increased to 10.19 wt.% (A-6), the maximum release of GTMC from the formulated bone cement is no higher than sample of A-5 with 8.15 wt.% of MSN. It should be noted that because the diffusion rate is also limited by the nano-sized pore channels in the network, the drug release is well controlled and achieves a sustained release of GTMC from MSN functionalized antibiotic bone cement in 80 days. This long term sustained release has not been achieved by using calciumphosphate bone substitutes [31,32] or a bioactive ceramic–polymer composite [33], which allowed GTMC to release much faster and complete most of release in the first several days or several hours. For PMMA polymer based bone cements formulated with MSN, the drug release kinetics is mostly controlled by diffusion, and thus the release could be expressed as [34]: ∞ ∞ X X Mt 4 8 ¼1 exp −Dα2n t exp −Dβ2m t 2 2 2 Mtot a αX n¼1 m¼0 l βm
ð1Þ
where Mt is the GTMC mass released at time t, Mtot is the total GTMC mass in the sample, a is the radius and l is the height of cylinder. The terms βm are defined as: βm = (2 m + 1)π / l and the terms αn are the positive roots of J0(aαn) = 0. Here J0 is the zero-order Bessel function of the first kind and aαn are the zeros of that function.
Fig. 6. Scheme of (a) GTMC mixed with bone cement (b) bone cement formulated with MSN at low loading and (c) effective diffusion network formed by MSN in PMMA based bone cement.
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Fig. 7. Plot of ln(1 − Mt / Mtot) as a function of time (Mt: the mass of GTMC released at time t; Mtot: the total GTMC mass in the sample).
For sustained and controlled release at long periods, only the first term of each sum will contribute to Eq. (1), thus: h i Mt 32 ¼ 1− 2 exp −D α21 þ β20 t 2 2 2 M∞ a l α β
ð2Þ
1 0
and Eq (2) can be expressed as Mt ln 1− ∝−D α21 þ β20 t: M∞
ð3Þ
When ln(1 − Mt / Mtot) is plotted as a function of time, it is observed that MSN formulated samples show good linear fit in this plot as shown in Fig. 7. The diffusion coefficients can be calculated by plotting ln(1 − Mt / Mtot) as a function of time: DA−1 ¼ 1:62 10−15 m2 =s DA−4 ¼ 1:55 10−13 m2 =s DA−5 ¼ 2:82 10−13 m2 =s DA−6 ¼ 4:96 10−13 m2 =s:
By the presence of 6 wt.% MSN in PMMA based bone cements, the diffusion coefficient is substantially increased by more than 95 folds. With the content of MSN increased to 10%, the nano-diffusion network built up by MSN becomes more efficient and the diffusion coefficient is further enhanced. In addition to the Simplex P bone cement, the Smartset HV bone cement is also functionalized with MSN fillers. Fig. 8 shows the release profiles of the GTMC–MSN functionalized Smartset bone cements with varied MSN contents and compared with available commercial gentamicin formulated bone cement Smartset GHV. After being functionalized with MSN filler, the drug release character of the Smartset bone cement is similar to that of the Simplex P bone cement. With 2 wt.% of MSN in formulation, the release profile is not obviously improved and comparable with the Smartset GHV antibiotic bone cement, because the effective diffusion network could not be formed while the MSN content is lower than the critical concentration. It is observed that, when the MSN content is increased to 6 wt.%, the release rate of GTMC is efficiently enhanced. 55% of GTMC could have sustained release in 80 days. Fig. 9 displays the bending modulus of MSN functionalized antibiotic bone cement with varied contents of MSN. It is observed that the bending modulus is slightly affected by the incorporation of MSN in the bone cement matrix. For the sample of A-6 with 10.19 wt.% of MSN, the
Fig. 8. Effect of MSN content on GTMC release from formulated Smartset HV bone cement.
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Fig. 9. Bending modulus (A) and compression strength (B) of Simplex-P and formulated with MSN at different loadings.
bending modulus is about 90% of the corresponding properties of the original bone cement while A-6 exhibits significantly improved drug release profiles. All of the MSN formulated samples preserve at least 80% of the original bending modulus as compared to the commercial Simplex-P bone cement. Furthermore, the compression strength of MSN functionalized antibiotic bone cements is also shown in Fig. 9. It is observed that the compression strength is negligibly affected by the incorporation of MSN into the antibiotic bone cement. The compressive strengths of MSN functionalized bone cements at all GTMC/MSN loadings were above the ASTM F541 and ISO 5833 minima of 70 MPa. Although some forms of antibiotic loaded bone cement were reported to exhibit excellent drug delivery profiles, they failed to be used in clinical practice because of their negative influence on the mechanical properties of the formulated bone cements [35,36]. It was reported that the mechanical properties of a PMMA–hydroxyapatite composite could be reinforced by a small amount of carbon nanotubes (0.1 wt.%), nevertheless, the performance on drug delivery was not conducted [37]. On other hand, it was found that the drug release rate could be significantly enhanced by incorporation of large amounts of hydroxyapatite (up to 35 wt.%) [38], however, we find that its compression strength decreased to about 50% of original PMMA based bone cement when 32 wt.% of hydroxyapatite is incorporated. It was reported that the presence of high antibiotic loading could facilitate the elution of drug from bone cements [38], the high loading of antibiotic in the PMMA based bone cement was found to detriment the mechanical properties. It was reported that when antibiotics were added by more than 4.8 wt.% by direct mixing, the cement strength was obviously affected and reduced by 27% [39], indicating that during artificial hip joint replacement, the dose of antibiotics should be concerned, in order to avoid affecting the strength of the bone cement and stability of the entire implant. In this work, it is noted that our formulation of antibiotics with MSN as drug carriers can preserve the mechanical properties and achieve significantly enhanced efficiency for drug release. The cytotoxicity of the nanostructured materials is essential to evaluate the modified antibiotic bone cement materials for clinical implantation application. The MTT cytotoxicity of CNT, HAP nanorods and MSN as well as the GTMC/MSN formulated bone cement B-6 sample before setting was assayed using 3T3 mouse fibroblasts and the measurement
is compared with the PMMA based bone cement without modification. The 3T3 cells are selected for this assay as their viable rates are substrate-dependent and they are non-specific cell lines. As displayed in Fig. 10, among these nanostructured materials, HAP nanorods showed the least cytotoxicity and 94% of 3T3 cell could be viable and CNT showed the highest toxicity with 85% cell viability. As HAP is one of major compositions of bone structure, its low cytotoxicity is reasonable. The induced cytotoxicity of CNT is among the most concerns for its application in biological systems and it has also attracted more attention in recent investigation [40–42]. As a comparison, MSN showed less cytotoxicity than CNT and cell viability with MSN is only slightly lower than that with HAP. 91 ± 5.2% of the 3T3 cells were viable in the presence of MSN particles in culture medium. When GTMC/MSN was mixed with the PMMA bone cement powder, the sample (A-6) exhibited least cytotoxicity. Although the sample was with 10.19% of MSN particles, 96 ± 3.2% of cell viability could be achieved. The results indicated that MSN formulation did not bring obvious cytotoxicity to commercial PMMA based bone cement. To cope with antibiotic-resisted microbial, binary antibiotic loaded MSN were formulated with PMMA-based bone cement and drug release
Fig. 10. Cell viability rate of cytotoxicity assay.
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Fig. 11. Drug release profiles from GTMC and vancomycin with MSN formulated Simplex-P bone cement.
profiles were investigated. Fig. 11 displays the release profiles of gentamicin and vancomycin (VCMC) respectively from a binary antibiotic loaded bone cement. The loading of antibiotic is 2.6 wt.% for gentamicin and VCMC equally and loading of MSN is the same as that of sample A-5. It was observed that the total release of GTMC can achieve 40% and meanwhile, VCMC also showed a conjugative release of up to 30%. The slightly lower release rate of VCMC than GTMC can be attributed to the larger molecular size of VCMC with a slower diffusion rate. The smaller molecular size of GTMC facilitated the diffusion through nanonetwork built up by MSN particles. The congestion of two types of molecules in nano-diffusion channels resulted in a slightly lower release rate of GTMC from combination formulation than that released from the GTMC/MSN formulated bone cements. The binary drug delivery system may enable a wide spectrum of antibiotic activity. 4. Conclusions PMMA based bone cements have been formulated with HAP nanorods, CNT and MSN as drug carrier to enhance the delivery efficiency of antibiotics. Although the presence of the CNT and the high loading of non-porous HAP as drug carriers can enhance the elution of antibiotics from formulated bone cement, the mechanical strength of PMMA based bone cement is seriously impaired and exhibited less potential for practical application. As a comparison, the MSN can effectively enhance the drug release profiles when the loading of MSN is at 6–12%, and the mechanical properties (bending and compression strength) of PMMA based bone cement are well preserved. The hollow nanostructured particles are believed to build up an effective network to enable antibiotics to diffuse from the bone cement matrix. Moreover, the MSN formulation also enables a binary delivery system to effectively deliver GTMC with other antibiotics (such as vancomycin). The excellent drug delivery efficiency and low cytotoxicity of MSN showed high potential to formulate effective antibiotic bone cements. Acknowledgment This work was generously supported by the Institute of Chemical Engineering and Sciences (ICES/14-422A02), and the Agency of Science Technology and Research (A*STAR), Singapore. References [1] N.P. Hailer, G. Garellick, J. Kärrholm, Uncemented and cemented primary total hip arthroplasty in the Swedish Hip Arthroplasty Register. Evaluation of 170,413 operations, Acta Orthop. 81 (2010) 34–41.
[2] W. Zimmerli, A. Trampuz, P.E. Ochsner, Prosthetic joint infections, N. Engl. J. Med. 315 (2004) 1645–1654. [3] R.P. Wenzel, T.M. Perl, The significance of nasal carriage of Staphylococcus aureus and incidence of post-operative wound infection, J. Hosp. Infect. 31 (1995) 13–24. [4] N. Goyal, A. Miller, M. Tripathi, J. Parvizi, Methicillin-resistant Staphylococcus aureus (MRSA) colonisation and pre-operative screening, Bone Joint J. 95 (2013) 4–9. [5] J.S. Everhart, E. Altneu, J.H. Calhoun, Medical comorbidities are independent preoperative risk factors for surgical infection after total joint arthroplasty, Clin. Orthop. Relat. Res. 2013 (471) (2013) 3112–3119. [6] M.R. Rasouli, M.G. Maltenfort, J.J. Purtill, W.J. Hozack, J. Parvizi, Has the rate of inhospital infections after total joint arthroplasty decreased? Clin. Orthop. Relat. Res. 471 (2013) 3102–3111. [7] A. Hernandez-Soria, Y. Xu, J.G. Matthew, R. Janine, F.R. Benjamin, B. Mathias, In vitro elution characteristics of antibiotic laden BoneSource™, hydroxyapatite bone cement, J. Biomater. Sci. Polym. Ed. 24 (2013) 797–806. [8] A. Konstantinos, K. Jens, Enhancement of antibiotic elution from acrylic bone cement, J. Biomed. Mater. Res. B Appl. Biomater. 90B (2009) 467–475. [9] J.G.E. Hendriks, J.R. van Horn, H.C. van der Mei, H.J. Busscher, Backgrounds of antibiotic-loaded bone cement and prosthesis-related infection, Biomaterials 25 (2004) 545–556. [10] R. Qadir, S. Sidhu, J.L. Ochsner, M.S. Meyer, G.F. Chimento, Risk stratified usage of antibiotic-loaded bone cement for primary total knee arthroplasty: short term infection outcomes with a standardized cement protocol, J. Arthroplasty 29 (2014) 1622–1624. [11] L. Boren, Z. Huan, W.L. Douglas, K.G. Vijay, K.A. Anand, B.B. Sarit, Sustained release of small molecules from carbon nanotube-reinforced monetite calcium phosphate cement, Mater. Sci. Eng. C 43 (2014) 92–96. [12] M. Marta, B. Alessandro, C.V. Maria, B.G.F. Carlotta, V. Enrica, Antibiotic-loaded acrylic bone cements: an in vitro study on the release mechanism and its efficacy, Mater. Sci. Eng. C 33 (2013) 3025–3032. [13] C.M. Ana, M.G. Lídia, R. Patrícia, A.V. Mário, J.A. António, F.B. Ana, A novel modified acrylic bone cement matrix. A step forward on antibiotic delivery against multiresistant bacteria responsible for prosthetic joint infections, Mater. Sci. Eng. 38 (2014) 218–226. [14] Z.L. Shi, K.G. Neoh, E.T. Kang, W. Wang, Antibacterial and mechanical properties of bone cement impregnated with chitosan nanoparticles, Biomaterials 27 (2006) 2440–2449. [15] H. Vik, K.J. Andersen, K. Julshamn, Neuropathy caused by silver absorption from arthroplasty cement, Lancet 13 (1985) (1985) 872. [16] P.V. AshaRani, G.L.K. Mun, M.P. Hande, S. Valiyaveettil, Cytotoxicity and genotoxicity of silver nanoparticles in human cells, ACS Nano 3 (2009) 279–290. [17] R. Lima, A.B. Seabra, N. Duránd, Silver nanoparticles: a brief review of cytotoxicity and genotoxicity of chemically and biogenically synthesized nanoparticles, J. Appl. Toxicol. 32 (2012) 867–879. [18] I. Landor, P. Vavrík, D. Jahoda, General principles of infection treatment in joint replacements, Acta Chir. Orthop. Traumatol. Cech. 72 (2005) 183–190. [19] W.A. Jiranek, A.D. Hanssen, A.S. Greenwald, Antibiotic-loaded bone cement for infection prophylaxis in total joint replacement, J. Bone Joint Surg. Am. 88 (2006) 2487–2500. [20] D. Neut, H. van de Belt, J.R. van Horn, H.C. van der Mei, H.J. Busscher, The effect of mixing on gentamicin release from polymethylmethacrylate bone cements, Acta Orthop. Scand. 74 (2003) 670–676. [21] R.P. del Real, S. Padilla, M. Vallet-Regi, Gentamicin release from hydroxyapatite/ poly(ethylmethacrylate)/poly(methyl methacrylate)composites, J. Biomed. Mater. Res. 52 (2000) 1–7. [22] S. Padilla1, R.P. del Real, M. Vallet-Regi, In vitro release of gentamicin from OHAp/ PEMA/PMMA Samples, J. Control. Release 83 (2002) 343–352. [23] J. Slane, J. Vivanco, D. Ebenstein, M. Squire, H.L. Ploeg, Multiscale characterization of acrylic bone cement modified with functionalized mesoporous silica nanoparticles, J. Mech. Behav. Biomed. Mater. 37 (2014) 141–152. [24] R. Ormsby, T. McNally, C. Mitchell, N.J. Dunne, Incorporation of multiwalled carbon nanotubes to acrylic based bone cements: effects on mechanical and thermal properties, J. Mech. Behav. Biomed. Mater. 3 (2010) 136–145. [25] A.D. Toms, D. Davidson, B.A. Masri, C.P. Duncan, The management of peri-prosthetic infection in total joint arthroplasty, J. Bone Joint Surg. 88 (2006) 149–155. [26] Y. Han, J.Y. Ying, Generalized fluorocarbon-surfactant-mediated synthesis of nanoparticles with various mesoporous structures, Angew. Chem. Int. Ed. 44 (2005) 288–292. [27] S.C. Shen, L. Chia, W.K. Ng, Y.C. Dong, R.B.H. Tan, Solid-phase steam-assisted synthesis of hydroxyapatite nanorods and nanoparticles, J. Mater. Sci. 22 (2010) 6059–6067. [28] X. Zhang, U.P. Wyss, D. Pichora, M.E.A. Goosen, Biodegradable controlled antibiotic release devices for osteomyelitis: optimization of release properties, J. Pharm. Pharmacol. 46 (1994) 718–724. [29] M.A. Puska, A.K. Kokkari, T.O. Närhi, P.K. Vallittu, Mechanical properties of oligomermodified acrylic bone cement, Biomaterials 24 (2003) 417–425. [30] G. Lewis, S. Janna, The in vitro elution of gentamicin sulfate from a commercially available gentamicin-loaded acrylic bone cement — VersaBond AB, J. Biomed. Mater. Res. Appl. Biomater. 71B (2004) 77–83. [31] H.P. Stallmann, C. Faber, A.L.J.J. Bronckers, A.V.N. Amerongen, P.I.L.M. Wuisman, In vitro gentamicin release from commercially available calcium-phosphate bone substitutes influence of carrier type on duration of the release profile, BMC Musculoskelet. Disord. 18 (2006) (8 pages). [32] U. Gbureck, E. Vorndran, F.A. Müller, J.E. Barralet, Low temperature direct 3D printed bioceramics and biocomposites as drug release matrices, J. Control. Release 122 (2007) 173–180.
S.-C. Shen et al. / Materials Science and Engineering C 58 (2016) 233–241 [33] H.Y. Li, J. Chang, Preparation, characterization and in vitro release of gentamicin from PHBV/wollastonite composite microspheres, J. Control. Release 107 (2005) 463–473. [34] P.L. Ritger, N.A. Peppas, A simple equation for description of solute release I. Fickian and non-Fickian release from non-swellable devices in the form of slabs, spheres, cylinders or discs, J. Control. Release 5 (1987) 23–36. [35] E. Ger, D. Dall, T. Miles, A. Forder, Bone cement and antibiotics, S. Afr. Med. J. 51 (1977) 276–279. [36] K.E. Marks, C.L. Nelson, E.P. Lautenschlager, Antibiotic-impregnated acrylic bone cement, J. Bone Joint Surg. Am. 58 (1976) 358–364. [37] M.K. Singh, T. Shokuhfar, J.J.D.A. Gracio, A.C.M.D. Sousa, J.M.D.F. Fereira, H. Garmestani, S. Ahzi, Hydroxyapatite modified with carbon-nanotube-reinforced poly(methyl methacrylate): a nanocomposite material for biomedical applications, Adv. Funct. Mater. 18 (2008) 694–700.
241
[38] R.P. del Real, S. Padilla, M. Vallet-Reg, Gentamicin release from hydroxyapatite/ poly(ethyl methacrylate)/poly(methyl methacrylate)composites, J. Biomed. Mater. Res. 52 (2000) 1–7. [39] C.S. Wu, F.T. Chiang, J.P. Hung, Effect of antibiotic additions on the mechanical properties of bone cement, Book Mater. Eng. Mech. Auto, Appl. Mech. Mater., 442 2014, pp. 147–151. [40] P. Anita, K. Brionna, T. Paul, Multi-walled carbon nanotubes induce cytotoxicity, genotoxicity and apoptosis in normal human dermal fibroblast cells, Ethn. Dis. 20 (2010) S1-65-72. [41] C.P. Firme, P.R. Bandaru, Toxicity issues in the application of carbon nanotubes to biological systems, Nanomedicine 6 (2010) 245–256. [42] Y. Seo, J. Hwang, J. Kim, Y. Yeong, M.P. Hwang, J. Choi, Antibacterial activity and cytotoxicity of multi-walled carbon nanotubes decorated with silver nanoparticles, Int. J. Nanomedicine 9 (2014) 4621–4629.
Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Nanostructured material formulated acrylic bone cements with enhanced drug release Shou-Cang Shen a,⁎, Wai Kiong Ng a, Yuan-Cai Dong a, Junwei Ng a, Reginald Beng Hee Tan a,b,⁎⁎ a b
Institute of Chemical and Engineering Sciences, A*STAR (Agency for Science, Technology and Research), 1 Pesek Road, Jurong Island, Singapore 627833, Singapore Department of Chemical and Biomolecular Engineering, The National University of Singapore, 4 Engineering Drive 4, Singapore 117576, Singapore
a r t i c l e
i n f o
Article history: Received 6 February 2015 Received in revised form 19 June 2015 Accepted 11 August 2015 Available online 19 August 2015 Keywords: Bone cements Antibiotic Nanoparticles Drug delivery Diffusion network
a b s t r a c t To improve antibiotic properties, poly(methyl methacrylate) (PMMA)-based bone cements are formulated with antibiotic and nanostructured materials, such as hydroxyapatite (HAP) nanorods, carbon nanotubes (CNT) and mesoporous silica nanoparticles (MSN) as drug carriers. For nonporous HAP nanorods, the release of gentamicin (GTMC) is not obviously improved when the content of HAP is below 10%; while the high content of HAP shows detrimental to mechanical properties although the release of GTMC can be substantially increased. As a comparison, low content of hollow nanostructured CNT and MSN can enhance drug delivery efficiency. The presence of 5.3% of CNT in formulation can facilitate the release of more than 75% of GTMC in 80 days, however, its mechanical strength is seriously impaired. Among nanostructured drug carriers, antibiotic/MSN formulation can effectively improve drug delivery and exhibit well preserved mechanical properties. The hollow nanostructured materials are believed to build up nano-networks for antibiotic to diffuse from the bone cement matrix to surface and achieve sustained drug release. Based on MSN drug carrier in formulated bone cement, a binary delivery system is also investigated to release GTMC together with other antibiotics. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Due to the excellent tissue compatibility of poly(methylmethacrylate) (PMMA) and the rapid setting with achieved mechanical strength, PMMA bone cement has been widely used clinically to fix total joint prostheses (hip and knee replacements) to bone for decades. It is an integral part of almost all total knee replacements and also used in at least twothirds of total hip replacements [1]. With the rapid growth of the aged population, there is a dramatic increase in the burden of musculoskeletal disease, over 1,000,000 total hip and knee replacements are performed each year in the US alone. Post-operation infection is a clinical catastrophe, leading to the need for more complex surgery with much higher costs or the increase of morbidity [2]. As with all implant materials, bone cement carries an elevated risk of infection when implanted into the human body because of the possibility microorganism biofilms to form on an inert surface [3–6], usually requiring multiple surgeries for treatment [7,10]. To reduce this risk, an antibiotic (such as gentamicin) has been formulated with bone cement; however, it is known that as the antibiotic is mostly trapped within the substance of the cement,
⁎ Corresponding author. ⁎⁎ Correspondence to: R.B.H. Tan, Institute of Chemical and Engineering Sciences, A*STAR (Agency for Science, Technology and Research), 1 Pesek Road, Jurong Island, Singapore 627833, Singapore. E-mail addresses: [email protected] (S.-C. Shen), [email protected] (R.B.H. Tan).
http://dx.doi.org/10.1016/j.msec.2015.08.011 0928-4931/© 2015 Elsevier B.V. All rights reserved.
only a small proportion is available for diffusing out to the desired site for antibacterial activity [8]. The therapeutic antimicrobial activity drops rapidly below effective therapeutic levels after surgery within several days [9] and did not show significant advantages in reducing infection rate [10]. Therefore, it is desired to develop functional antibiotic bone cement that can enhance antimicrobial activity and/or prolong the therapeutic levels of antibiotic formulated in bone cement [11–13]. The sustained antimicrobial activity would be of significant impact in improving the results of joint replacement surgeries. Researchers have explored many approaches to improve the antimicrobial activity of bone cement. It has been reported that the presence of chitosan nanoparticles could effectively prevent viable bacteria from surviving on the surface of the bone cement [14] and the addition of silver nanoparticles to increase its antimicrobial activity [15]. However, those active ingredients cannot be released to protect surrounding tissues and often have adverse cytotoxic effects [16,17]. Many efforts have been made to improve the release of antibiotics from bone cement for applications in total joint replacement, including vacuum-mixing of commercially available antibiotic-impregnated bone cement, but the effects are limited [18–20]. M. Vallet-Regi et al. reported to enhance the release of GTMC from PMMA based bone cement by formulation with HAP [21,22]. Although the drug elution from the PMMA bone cement was enhanced, the high content of fillers could impair mechanical strength. MSN has been formulated with acrylic bone cement and the resulting mechanical behaviors were characterized [23]. Ormsby et al. [24] studied the incorporation of CNT into acrylic based bone cements
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and reinforced mechanical properties were found. However, the effect of these nanoparticles on delivery of antibiotics has rarely been studied. The sustained release of antibiotics from bone cement is essential to prevent the formation of biofilm on implanted bone cement and prosthesis parts. Thus, it is necessary to develop antibiotic bone cements with enhanced and prolonged antimicrobial activity to protect the surrounding tissues, as the operation sites are prone to infection for up to 6 to 8 weeks after surgery and usually need at least 6 weeks of parenteral antibiotic administration [25]. In this study, nanostructured materials, such as CNT, HAP nanorods and MSN, are formulated into PMMA based bone cement to enhance the release of antibiotics. The presence of hollow nanostructured particles in bone cement builds up diffusion networks and improves the loaded antibiotics to diffuse from bone cement to surface. Moreover, novel bone cement enabled effective delivery of gentamicin in combination with other antibiotics, such as vancomycin.
Table 1 Composition of commercial bone cements.
Powder Methylmethacrylate methacrylate copolymer Methyl methacrylate–stylene copolymer Zirconium dioxide Barium sulfate Benzoyl peroxide Gentamicin Liquid Methylmethacrylate N,N-dimethyl-p-toluidine Hydroquinone
Simplex P (w/w%)
Smartset HV (w/w%)
Smartset GHV (w/w%)
15.00
84.00
80.46
73.72
–
–
15.00
14.37
1.00 –
0.96 4.22
97.50 2.50
97.50 2.50
10.00 1.28 – 97.49 2.50 0.0075
% by weight (w/w) of powder component and liquid component.
2. Experimental 2.1. Materials Mesoporous silica nanoparticles (MSN) were prepared using fluorocarbon-surfactant-mediated synthesized as reported by Han et al. [26]. Typically, 0.5 g of Pluronic P123 and 1.4 g of FC-4 were dissolved in 80 ml of HCl solution (0.02 M), followed by the introduction of 2.0 g of TEOS under stirring. The solution was continuously stirred at 30 °C for 24 h and then transferred into a polypropylene bottle and kept at 100 °C for 24 h. The resultant solid was recovered by centrifuging and washed with deionized water twice, subsequently it was dried at 55 °C for 12 h. To remove the template molecules, the material was heated from room temperature to 550 °C at a heating rate of 2 °C/min and followed by calcination in air for 6 h. Nanorods of hydroxyapatite (HAP) were prepared by wet gel steaming method [27]. Typically, 15 ml of 0.5 M Ca(NO3)2·4H2O (Alfa Aesar) solution was added to 15 ml of 0.5 M (NH4)2HPO4 (Alfa Aesar) solution to achieve the desired Ca2+/PO3− ratio of 1.0 for precipitation 4 of di-calcium phosphate dihydrate (CaPO4·2H2O, Ca: P = 1:1) at 25 °C. The pH of the mixture was adjusted to 11.0 by using aqueous ammonia solution (NH4OH, 25 wt.%, Merck). The precipitate solid was recovered by filtration and then transferred to a 25 ml beaker, which was placed in an autoclave with a polytetrafluoroethylene (PTFE) liner. 5 ml of the ammonia solution was poured into the bottom of the PTFE cup and physically separated from the solid sample. The autoclave was then placed into an oven at 180 °C for 20 h. After the steam-assisted treatment, the resulting white solid was dispersed in water using ultrasound and recovered by centrifugation. The washing procedure was repeated twice. The obtained solid material was dried in an oven at 55 °C, and finally ground to a fine powder. The multiwall carbon nanotubes (CNT) were provided by Shenzhen Nanotech Port Co., Ltd. Three kinds of commercial bone cements were used: CMW Smartset-HV and CMW Smart GHV (DePuy International Ltd. UK) and Simplex P (Stryker Co, UK). The compositions of the bone cements, as indicated in the products' brochure, are shown in Table 1. 2.1.1. PBS buffer Buffered saline solution pH 7.4 was prepared as described in the British Pharmacopoeia (BP A79): anhydrous di-sodium hydrogen phosphate 2.38 g; anhydrous potassium hydrogen phosphate 0.19 g and sodium chloride 8 g to 1000 ml with deionized water (Milli-Q). 2.1.2. o-Phthaldialdehyde reagent o-Phthaldialdehyde reagent was prepared according to a procedure reported by Zhang et al. [28]. It was formulated by adding 2.5 g of o-phthaldialdehyde, 62.5 ml methanol and 3 ml of 2-mercaptoethanol to 560 ml of 0.04 M sodium borate in a deionized water solution. The
reagent was stored in the darkness and settled for at least 24 h prior to use. 2.2. Preparation of antibiotic-loaded bone cements GTMC was loaded onto the nanostructured materials (MSN, HAP or CNT) by wet impregnation. Typically, 0.20 g of GTMC was dissolved in 3 ml deionized water. 0.30 g of MSN powder was impregnated with GTMC solution under stirring and aged for 24 h. The mixture was dried under vacuum at room temperature under vacuum for 48 h. The dried GTMC loaded nanoparticles were ground to fine powder and a certain amount of GTMC loaded nanoparticles were mixed with commercial bone cement solid powder by manual grinding. When HAP and CNT are used as drug carriers, the same impregnation method was used and the drug-to-carrier ratio may be adjusted to avoid too high drug loading in the final formulation. Samples of A-1–A-9 prepared by using GTMC–MSN with different drug loadings are listed in Table 2. The samples of antibiotic loaded bone cement were prepared by mixing the powder with the liquid monomer in a ratio of 2 g/ml in a beaker in a laminar flow hood, in accordance with the manufacturer's instruction. Monomer liquid was added to the polymer–GTMC–MSN mixture in a bowl and was stirred using a spatula until the powder was fully wetted. The soft mixture was inserted into the mold with dimensions of 6 mm in diameter and 12 mm in height. The filled mold was pressed between two glass plates to harden overnight at room temperature. The hardened bone cement cylinders were pulled out of the mold and stored under sterile conditions at room temperature for the in-vitro drug release test and compression test. In addition, rectangular beams with dimension of 25 × 10 × 2 mm were prepared for a bending test and an antibacterial property test. 2.3. In vitro drug release study The drug release study was conducted by soaking two cylinder samples of each composition in 5 ml PBS buffer (pH 7.2). The sample Table 2 Composition of nanomaterial formulated antibiotic bone cements.
A-1 A-2 A-3 A-4 A-5 A-6 A-7 A8 A9
Simplex-P powder (g)
MMA (ml)
Filler %
Drug%
1.90 1.90 1.80 1.70 1.60 1.50 2.0 1.0 2.0
1.0 1.0 1.0 1.0 1.0 1.0 1.0 1.0 1.0
MSN 0 MSN 2.04 MSN 4.08 MSN 6.12 MSN 8.15 MSN 10.19 HAP 11.56 HAP 32.33 CNT 5.36
3.40 1.36 2.72 4.08 5.44 6.79 3.59 4.85 3.21
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Fig. 1. TEM images of (a) MSN, (b) nanorods of HAP and (c) CNT.
was put in an incubator shaker operated at 37 °C and 40 rpm. The release medium was withdrawn at predetermined time intervals, and replaced with fresh PBS buffer (5 ml) each time. The accumulative amount of gentamicin released was calculated based on the initial weight of the bone cement cylinder and the drug content. The gentamicin release was studied for 80 days. An indirect method was used for measurement of gentamicin concentration by a UV–vis spectrophotometer (Cary 50, Agilent Technologies) because gentamicin does not absorb ultraviolet nor visible light. The o-phthaldialdehyde was used as a derivatizing agent to react with the amino groups of gentamicin and yield chromophoric products. The reaction was carried out making 1 ml of our problem gentamicin in solution react with 1 ml of isopropanol (to avoid the precipitation of the products formed) and 1 ml of o-phthaldialdehyde reagent. After fully mixed, the concentration of gentamicin sulfate was determined by UV absorbance at 332 nm. 2.4. In vitro cytotoxicity assay 3T3 mouse fibroblast cells (3T3-Swiss albino, ATCC) were cultured in a complete growth culture medium in a 5% CO2 incubator. The
complete growth culture medium was prepared with Dulbecco's Modified Eagle's Medium (DMEM) supplemented with 10% fetal bovine serum, 1 mm L-glutamine and penicillin (100 U/ml). Cell viability testing was carried out via the reduction of the MTT reagent (3-[4,5-dimethyl-thiazol-2-yl]-2,5-diphenyltetrazolium bromide, Sigma). This assay provides a simple method for determining comparative cell viability using a standard microplate absorbance reader. The MTT assay was performed with the bone cement substrates placed at the bottom of a 96-well plate following the standard procedure with minor modifications. Control experiments were carried out using the complete growth culture medium only (serving as the non-toxic control) and 1% Triton X-100 (Sigma) (as the toxic control). 3T3 fibroblasts in the complete growth culture medium (100 μl) were seeded at a density of 104 cells/well in a 5% CO2 incubator for 24 h. The culture medium from each well was then removed and 100 μl of medium and 20 μl MTT solution (5 mg/ml in PBS) were then added to each well. After 4 h of incubation at 37 °C and 5% CO2, the media were removed and the formazan crystals were solubilized with 100 μl dimethyl sulfoxide (DMSO, Sigma) for 15 min. The optical absorbance was then measured at 570 nm on a microplate reader (Tecan GENios). Six samples were tested for each type of bone cement.
Fig. 2. Gentamicin release profiles of PMMA-based bone cement formulated with MSN (A-6), HAP (A-7, A-8) and CNT (A-9) as compared with antibiotic bone cement Smartset GHV.
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2.5. Testing of mechanical property The three point bending test was performed on the Instron Universal Materials Testing Machine (Model 5544). According to the standard test method of ASTM D790-3, the span length was 20 mm and the loading rate was 1 mm/min. The bending modulus (EB) was calculated according to the following equation: 3
EB ¼ L3 m=4bd ; where L is the support span (mm), b is the width of beam tested (mm), d is the depth of beam tested (mm), and m is the slope of the tangent to the initial straight-line portion of the load–deflection curve (N/mm). The compression tests were carried out on the bone cement cylinders with the same dimension as those for drug release investigation [29]. The compression force was applied along the axis using a crosshead speed of 5 mm/min. The compression strength was calculated from the obtained load–deformation curves. The compression strength (CS, MPa) was calculated using following equations: CSðMPaÞ ¼ FðNÞ=A m2 =1000; 000 where F is the applied load (N) at the highest point of the load–deflection curve and A is the cross-section area (m2) of the sample tested. 2.6. Characterization High resolution TEM images of nanostructured materials were taken by a TECNAI F20 (G2) (FEI, Philips Electron Optics, Holland) electron microscope at 200 kV. Nitrogen adsorption/desorption isotherms were measured by using an Autosorb-6B gas adsorption analyzer (Quantachrome) at a temperature of −196 °C. Before nitrogen adsorption–desorption measurements, each sample was heated at 40 °C under vacuum for 24 h. The specific surface areas of the samples were determined from the linear portion of the Brumauer–Emmett–Teller (BET) plots. The total pore volume, VT, was estimated from the amount adsorbed at a relative pressure id P/P0 of 0.95. 3. Results and discussion Fig. 1 shows the TEM images of nanostructured materials used as drug carriers and bone cement fillers in this study. It can be seen that the mesoporous silica nanoparticles (MSN) have orderly arranged pore channels (Fig. 1a) and carbon nanotubes (CNT) show a hollow structure with a single channel at the nanoscale (Fig. 1c). As a comparison, no porous structure can be observed in the HAP nanoparticles as the pore volume of HAP measured by N2 adsorption is 0.15 cm3/g only [27]. Fig. 2 displays the cumulative release profile of GTMC from different nanostructure drug carriers loaded with GTMC formulated in the Simplex-P bone cement and compared with the commercial antibiotic bone cement Smartset GHV. It is observed that the commercial antibiotic bone cement exhibits low GTMC release rate. Smartset GHV only shows about 5% of release on the first day and the release of antibiotic on following days is negligible. This result is consistent with reported investigation of commercially available GTMC-loaded acrylic bone cement and it was postulated that the elution mechanism of GTMC in this type of cement is a surface phenomenon [30]. Only surface adhered GTMC particles could be released. To enhance drug delivery efficiency, three types of nanostructured materials are used as drug carrier and formulated to bone cement. It is found that the presence of 5.36% of CNT (A-8) in functional bone cement can lead to more than 75% of GTMC release in 60 days and the formulation with 10.19% of MSN (A-6) also enhances the total drug release to 60%. As a comparison, the presence of 11.56% nonporous HAP nanoparticles only slightly enhances the total drug release to about 10%. When the content of the HAP nanoparticles
Fig. 3. Load–displacement curves of formulated bone cements.
in formulated bone cement is increased to 32.33%, 75% of GTMC could be released. This result is in agreement with literatures reported that the presence of about 30% of hydroxyapatite led to about 90% in 700 h [22]. However, the effect of high HAP content on mechanistic property was not reported. Fig. 3 shows the load–displacement curves of bone cement formulated with different nanostructured materials and compared with commercial Simplex-P bone cement. It is shown that the compression strength of MSN formulated bone cement (A-6) is well preserved as compared with the original Simplex-P, where the load–displacement curves are almost unchanged. As a comparison, the compression strength of HAP formulated bone cement (A-8) is reduced by about 50% although the high content of HAP could enhance the drug release. The large amount of HAP (32%) in bone cement formulation may cause the formation of a large void or a porosity structure [22], thus seriously reducing its compression strength. Moreover, the CNT formulated bone cement can only preserve about 10% of compression strength of the original Simplex-P bone cement. The detriment to biomechanical properties of bone cement limited the application of the nanostructured HAP and CNT in formulation of antibiotic bone cements for effective drug delivery. Thus, in this study, MSN was selected for further investigation. Fig. 4 displays the N2 adsorption–desorption isotherms of MSN and GTMC loaded MSN. The large pore volume of MSN (1.01 cm3/g) exhibits a high capacity of N2 adsorption as high as 511 cm3/g, STP at P/P0 of 0.98 and the sharp pore size distribution peak at 6 nm indicates its uniform pore diameter within the silica nanoparticles. When MSN is loaded with 40% of GTMC, the adsorption of N2 reduced significantly at whole
Fig. 4. N2 adsorption isotherms of (a) MSN (inset is the pore diameter distribution), (b) GTMC (40 wt.%) impregnated MSN (60 wt.%), and (c) solid powder of Simplex-p.
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Fig. 5. Cumulative GTMC release profile from Simplex-P bone cement formulated with MSN at difference loadings.
range of P/P0 and the total pore volume shrinks to 0.23 cm3/g. The results imply that the most pore volume of uniform pore channels is occupied by GTMC molecules. As a comparison, the PMMA based bone cement powder is a non-porous material as it showed negligible N2 adsorption. The nonporous structure of PMMA based polymers should be attributed to the low drug release efficiency as drug molecules are embedded in the condensed polymer based bone cement without porous diffusion pathways. When MSN loaded GTMC is formulated with the PMMA based polymer, the porous structure is used as drug carriers. Once the drug molecules are released to the medium, the void of uniform arranged mesoporous channels is expected to play a role as diffusion pathways for GTMC located at the inner matrix of bone cement to diffuse out. However, the content of MSN in the bone cement matrix can affect the efficiency of diffusion network. Fig. 5 shows the cumulative release profile of GTMC from MSN incorporated bone cement with different contents of MSN and GTMC. It is observed that, for the sample prepared by direct mixing of GTMC and bone cement powder, the release of GTMC is very limited. Only about 3% of GTMC released is observed in first day of immersion in PBS solution. No GTMC release is detectable in the following 80 days of investigation. When the sample is formulated with GTMC loaded MSN, the release from MSN-bone cement composite can be enhanced. However, the release rate is obviously affected by the content of MSN in the final composite. For the A-2 and A-3 samples with MSN contents of 2.04 wt.% and 4.08 wt.%, the release of GTMC only reaches 5% in the first day of immersion and the release stopped during the following investigation. The large portion of GTMC is still retained inside of the PMMA based bone cement matrix and no obvious enhancement on drug release character as compared with direct blending of GTMC and PMMA based bone cement. When MSN content is increased to 6.12 wt.% in the formulation, the release of GTMC from the sample of A-4 is significantly enhanced. After 7.3% of release in the first day, a sustained release of GTMC from MSN
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incorporated bone cement is observed in 80 days and reaches about 40% of release. A further increase of the content of MSN to 8.15 wt.% (A-5) results in a sustained release to 58% of GTMC in 80 days. When the MSN content is increased to 10.19 wt.%, the A-6 sample exhibits a faster release rate of GTMC in the middle stage between 10 and 30 days than that of A-5, and finally they reach a similar percentage of GTMC release after 80 days of immersion in PBS buffer. The MSN nanoparticles incorporated in bone cement is believed to form a nano-network pathway for GTMC diffusion from composite to surface and released to the medium. As shown in Fig. 6a, for GTMC formulated bone cement in the absence of MSN nanoparticles, GTMC particles embedded inside the PMMA matrix during polymerization cannot diffuse to the surface of bone cement for release. For the GTMC–MSN loaded bone cement composite, most of GTMC was entrapped inside the mesoporous channels of the MSN rod-like nanoparticles. The critical content of MSN inside bone cement is required to build up the network for GTMC to diffuse from the matrix to the medium. When the content of MSN is below 6 wt.%, most of the GTMC loaded MSN particles are isolated and also embedded in the bone cement matrix as shown in Fig. 6b. The release of GTMC from the bone cement composite cannot be obviously improved. Only those GTMC loaded MSN particles on the shadow surface of bone cement contribute to drug release. Once the MSN content increased to 6 wt.% or above, a nano-network can be built up through “particle–particle” contact (Fig. 6C) and allow the GTMC molecules to continuously diffuse from the PMMA based matrix to surface and be released to the medium, thus the release profile of GTMC is obviously improved as indicated in Fig. 5. When MSN content is increased to 8.15 wt.%, the diffusion network is more effective and the drug release is significantly enhanced. While the content of MSN is further increased to 10.19 wt.% (A-6), the maximum release of GTMC from the formulated bone cement is no higher than sample of A-5 with 8.15 wt.% of MSN. It should be noted that because the diffusion rate is also limited by the nano-sized pore channels in the network, the drug release is well controlled and achieves a sustained release of GTMC from MSN functionalized antibiotic bone cement in 80 days. This long term sustained release has not been achieved by using calciumphosphate bone substitutes [31,32] or a bioactive ceramic–polymer composite [33], which allowed GTMC to release much faster and complete most of release in the first several days or several hours. For PMMA polymer based bone cements formulated with MSN, the drug release kinetics is mostly controlled by diffusion, and thus the release could be expressed as [34]: ∞ ∞ X X Mt 4 8 ¼1 exp −Dα2n t exp −Dβ2m t 2 2 2 Mtot a αX n¼1 m¼0 l βm
ð1Þ
where Mt is the GTMC mass released at time t, Mtot is the total GTMC mass in the sample, a is the radius and l is the height of cylinder. The terms βm are defined as: βm = (2 m + 1)π / l and the terms αn are the positive roots of J0(aαn) = 0. Here J0 is the zero-order Bessel function of the first kind and aαn are the zeros of that function.
Fig. 6. Scheme of (a) GTMC mixed with bone cement (b) bone cement formulated with MSN at low loading and (c) effective diffusion network formed by MSN in PMMA based bone cement.
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Fig. 7. Plot of ln(1 − Mt / Mtot) as a function of time (Mt: the mass of GTMC released at time t; Mtot: the total GTMC mass in the sample).
For sustained and controlled release at long periods, only the first term of each sum will contribute to Eq. (1), thus: h i Mt 32 ¼ 1− 2 exp −D α21 þ β20 t 2 2 2 M∞ a l α β
ð2Þ
1 0
and Eq (2) can be expressed as Mt ln 1− ∝−D α21 þ β20 t: M∞
ð3Þ
When ln(1 − Mt / Mtot) is plotted as a function of time, it is observed that MSN formulated samples show good linear fit in this plot as shown in Fig. 7. The diffusion coefficients can be calculated by plotting ln(1 − Mt / Mtot) as a function of time: DA−1 ¼ 1:62 10−15 m2 =s DA−4 ¼ 1:55 10−13 m2 =s DA−5 ¼ 2:82 10−13 m2 =s DA−6 ¼ 4:96 10−13 m2 =s:
By the presence of 6 wt.% MSN in PMMA based bone cements, the diffusion coefficient is substantially increased by more than 95 folds. With the content of MSN increased to 10%, the nano-diffusion network built up by MSN becomes more efficient and the diffusion coefficient is further enhanced. In addition to the Simplex P bone cement, the Smartset HV bone cement is also functionalized with MSN fillers. Fig. 8 shows the release profiles of the GTMC–MSN functionalized Smartset bone cements with varied MSN contents and compared with available commercial gentamicin formulated bone cement Smartset GHV. After being functionalized with MSN filler, the drug release character of the Smartset bone cement is similar to that of the Simplex P bone cement. With 2 wt.% of MSN in formulation, the release profile is not obviously improved and comparable with the Smartset GHV antibiotic bone cement, because the effective diffusion network could not be formed while the MSN content is lower than the critical concentration. It is observed that, when the MSN content is increased to 6 wt.%, the release rate of GTMC is efficiently enhanced. 55% of GTMC could have sustained release in 80 days. Fig. 9 displays the bending modulus of MSN functionalized antibiotic bone cement with varied contents of MSN. It is observed that the bending modulus is slightly affected by the incorporation of MSN in the bone cement matrix. For the sample of A-6 with 10.19 wt.% of MSN, the
Fig. 8. Effect of MSN content on GTMC release from formulated Smartset HV bone cement.
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Fig. 9. Bending modulus (A) and compression strength (B) of Simplex-P and formulated with MSN at different loadings.
bending modulus is about 90% of the corresponding properties of the original bone cement while A-6 exhibits significantly improved drug release profiles. All of the MSN formulated samples preserve at least 80% of the original bending modulus as compared to the commercial Simplex-P bone cement. Furthermore, the compression strength of MSN functionalized antibiotic bone cements is also shown in Fig. 9. It is observed that the compression strength is negligibly affected by the incorporation of MSN into the antibiotic bone cement. The compressive strengths of MSN functionalized bone cements at all GTMC/MSN loadings were above the ASTM F541 and ISO 5833 minima of 70 MPa. Although some forms of antibiotic loaded bone cement were reported to exhibit excellent drug delivery profiles, they failed to be used in clinical practice because of their negative influence on the mechanical properties of the formulated bone cements [35,36]. It was reported that the mechanical properties of a PMMA–hydroxyapatite composite could be reinforced by a small amount of carbon nanotubes (0.1 wt.%), nevertheless, the performance on drug delivery was not conducted [37]. On other hand, it was found that the drug release rate could be significantly enhanced by incorporation of large amounts of hydroxyapatite (up to 35 wt.%) [38], however, we find that its compression strength decreased to about 50% of original PMMA based bone cement when 32 wt.% of hydroxyapatite is incorporated. It was reported that the presence of high antibiotic loading could facilitate the elution of drug from bone cements [38], the high loading of antibiotic in the PMMA based bone cement was found to detriment the mechanical properties. It was reported that when antibiotics were added by more than 4.8 wt.% by direct mixing, the cement strength was obviously affected and reduced by 27% [39], indicating that during artificial hip joint replacement, the dose of antibiotics should be concerned, in order to avoid affecting the strength of the bone cement and stability of the entire implant. In this work, it is noted that our formulation of antibiotics with MSN as drug carriers can preserve the mechanical properties and achieve significantly enhanced efficiency for drug release. The cytotoxicity of the nanostructured materials is essential to evaluate the modified antibiotic bone cement materials for clinical implantation application. The MTT cytotoxicity of CNT, HAP nanorods and MSN as well as the GTMC/MSN formulated bone cement B-6 sample before setting was assayed using 3T3 mouse fibroblasts and the measurement
is compared with the PMMA based bone cement without modification. The 3T3 cells are selected for this assay as their viable rates are substrate-dependent and they are non-specific cell lines. As displayed in Fig. 10, among these nanostructured materials, HAP nanorods showed the least cytotoxicity and 94% of 3T3 cell could be viable and CNT showed the highest toxicity with 85% cell viability. As HAP is one of major compositions of bone structure, its low cytotoxicity is reasonable. The induced cytotoxicity of CNT is among the most concerns for its application in biological systems and it has also attracted more attention in recent investigation [40–42]. As a comparison, MSN showed less cytotoxicity than CNT and cell viability with MSN is only slightly lower than that with HAP. 91 ± 5.2% of the 3T3 cells were viable in the presence of MSN particles in culture medium. When GTMC/MSN was mixed with the PMMA bone cement powder, the sample (A-6) exhibited least cytotoxicity. Although the sample was with 10.19% of MSN particles, 96 ± 3.2% of cell viability could be achieved. The results indicated that MSN formulation did not bring obvious cytotoxicity to commercial PMMA based bone cement. To cope with antibiotic-resisted microbial, binary antibiotic loaded MSN were formulated with PMMA-based bone cement and drug release
Fig. 10. Cell viability rate of cytotoxicity assay.
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Fig. 11. Drug release profiles from GTMC and vancomycin with MSN formulated Simplex-P bone cement.
profiles were investigated. Fig. 11 displays the release profiles of gentamicin and vancomycin (VCMC) respectively from a binary antibiotic loaded bone cement. The loading of antibiotic is 2.6 wt.% for gentamicin and VCMC equally and loading of MSN is the same as that of sample A-5. It was observed that the total release of GTMC can achieve 40% and meanwhile, VCMC also showed a conjugative release of up to 30%. The slightly lower release rate of VCMC than GTMC can be attributed to the larger molecular size of VCMC with a slower diffusion rate. The smaller molecular size of GTMC facilitated the diffusion through nanonetwork built up by MSN particles. The congestion of two types of molecules in nano-diffusion channels resulted in a slightly lower release rate of GTMC from combination formulation than that released from the GTMC/MSN formulated bone cements. The binary drug delivery system may enable a wide spectrum of antibiotic activity. 4. Conclusions PMMA based bone cements have been formulated with HAP nanorods, CNT and MSN as drug carrier to enhance the delivery efficiency of antibiotics. Although the presence of the CNT and the high loading of non-porous HAP as drug carriers can enhance the elution of antibiotics from formulated bone cement, the mechanical strength of PMMA based bone cement is seriously impaired and exhibited less potential for practical application. As a comparison, the MSN can effectively enhance the drug release profiles when the loading of MSN is at 6–12%, and the mechanical properties (bending and compression strength) of PMMA based bone cement are well preserved. The hollow nanostructured particles are believed to build up an effective network to enable antibiotics to diffuse from the bone cement matrix. Moreover, the MSN formulation also enables a binary delivery system to effectively deliver GTMC with other antibiotics (such as vancomycin). The excellent drug delivery efficiency and low cytotoxicity of MSN showed high potential to formulate effective antibiotic bone cements. Acknowledgment This work was generously supported by the Institute of Chemical Engineering and Sciences (ICES/14-422A02), and the Agency of Science Technology and Research (A*STAR), Singapore. References [1] N.P. Hailer, G. Garellick, J. Kärrholm, Uncemented and cemented primary total hip arthroplasty in the Swedish Hip Arthroplasty Register. Evaluation of 170,413 operations, Acta Orthop. 81 (2010) 34–41.
[2] W. Zimmerli, A. Trampuz, P.E. Ochsner, Prosthetic joint infections, N. Engl. J. Med. 315 (2004) 1645–1654. [3] R.P. Wenzel, T.M. Perl, The significance of nasal carriage of Staphylococcus aureus and incidence of post-operative wound infection, J. Hosp. Infect. 31 (1995) 13–24. [4] N. Goyal, A. Miller, M. Tripathi, J. Parvizi, Methicillin-resistant Staphylococcus aureus (MRSA) colonisation and pre-operative screening, Bone Joint J. 95 (2013) 4–9. [5] J.S. Everhart, E. Altneu, J.H. Calhoun, Medical comorbidities are independent preoperative risk factors for surgical infection after total joint arthroplasty, Clin. Orthop. Relat. Res. 2013 (471) (2013) 3112–3119. [6] M.R. Rasouli, M.G. Maltenfort, J.J. Purtill, W.J. Hozack, J. Parvizi, Has the rate of inhospital infections after total joint arthroplasty decreased? Clin. Orthop. Relat. Res. 471 (2013) 3102–3111. [7] A. Hernandez-Soria, Y. Xu, J.G. Matthew, R. Janine, F.R. Benjamin, B. Mathias, In vitro elution characteristics of antibiotic laden BoneSource™, hydroxyapatite bone cement, J. Biomater. Sci. Polym. Ed. 24 (2013) 797–806. [8] A. Konstantinos, K. Jens, Enhancement of antibiotic elution from acrylic bone cement, J. Biomed. Mater. Res. B Appl. Biomater. 90B (2009) 467–475. [9] J.G.E. Hendriks, J.R. van Horn, H.C. van der Mei, H.J. Busscher, Backgrounds of antibiotic-loaded bone cement and prosthesis-related infection, Biomaterials 25 (2004) 545–556. [10] R. Qadir, S. Sidhu, J.L. Ochsner, M.S. Meyer, G.F. Chimento, Risk stratified usage of antibiotic-loaded bone cement for primary total knee arthroplasty: short term infection outcomes with a standardized cement protocol, J. Arthroplasty 29 (2014) 1622–1624. [11] L. Boren, Z. Huan, W.L. Douglas, K.G. Vijay, K.A. Anand, B.B. Sarit, Sustained release of small molecules from carbon nanotube-reinforced monetite calcium phosphate cement, Mater. Sci. Eng. C 43 (2014) 92–96. [12] M. Marta, B. Alessandro, C.V. Maria, B.G.F. Carlotta, V. Enrica, Antibiotic-loaded acrylic bone cements: an in vitro study on the release mechanism and its efficacy, Mater. Sci. Eng. C 33 (2013) 3025–3032. [13] C.M. Ana, M.G. Lídia, R. Patrícia, A.V. Mário, J.A. António, F.B. Ana, A novel modified acrylic bone cement matrix. A step forward on antibiotic delivery against multiresistant bacteria responsible for prosthetic joint infections, Mater. Sci. Eng. 38 (2014) 218–226. [14] Z.L. Shi, K.G. Neoh, E.T. Kang, W. Wang, Antibacterial and mechanical properties of bone cement impregnated with chitosan nanoparticles, Biomaterials 27 (2006) 2440–2449. [15] H. Vik, K.J. Andersen, K. Julshamn, Neuropathy caused by silver absorption from arthroplasty cement, Lancet 13 (1985) (1985) 872. [16] P.V. AshaRani, G.L.K. Mun, M.P. Hande, S. Valiyaveettil, Cytotoxicity and genotoxicity of silver nanoparticles in human cells, ACS Nano 3 (2009) 279–290. [17] R. Lima, A.B. Seabra, N. Duránd, Silver nanoparticles: a brief review of cytotoxicity and genotoxicity of chemically and biogenically synthesized nanoparticles, J. Appl. Toxicol. 32 (2012) 867–879. [18] I. Landor, P. Vavrík, D. Jahoda, General principles of infection treatment in joint replacements, Acta Chir. Orthop. Traumatol. Cech. 72 (2005) 183–190. [19] W.A. Jiranek, A.D. Hanssen, A.S. Greenwald, Antibiotic-loaded bone cement for infection prophylaxis in total joint replacement, J. Bone Joint Surg. Am. 88 (2006) 2487–2500. [20] D. Neut, H. van de Belt, J.R. van Horn, H.C. van der Mei, H.J. Busscher, The effect of mixing on gentamicin release from polymethylmethacrylate bone cements, Acta Orthop. Scand. 74 (2003) 670–676. [21] R.P. del Real, S. Padilla, M. Vallet-Regi, Gentamicin release from hydroxyapatite/ poly(ethylmethacrylate)/poly(methyl methacrylate)composites, J. Biomed. Mater. Res. 52 (2000) 1–7. [22] S. Padilla1, R.P. del Real, M. Vallet-Regi, In vitro release of gentamicin from OHAp/ PEMA/PMMA Samples, J. Control. Release 83 (2002) 343–352. [23] J. Slane, J. Vivanco, D. Ebenstein, M. Squire, H.L. Ploeg, Multiscale characterization of acrylic bone cement modified with functionalized mesoporous silica nanoparticles, J. Mech. Behav. Biomed. Mater. 37 (2014) 141–152. [24] R. Ormsby, T. McNally, C. Mitchell, N.J. Dunne, Incorporation of multiwalled carbon nanotubes to acrylic based bone cements: effects on mechanical and thermal properties, J. Mech. Behav. Biomed. Mater. 3 (2010) 136–145. [25] A.D. Toms, D. Davidson, B.A. Masri, C.P. Duncan, The management of peri-prosthetic infection in total joint arthroplasty, J. Bone Joint Surg. 88 (2006) 149–155. [26] Y. Han, J.Y. Ying, Generalized fluorocarbon-surfactant-mediated synthesis of nanoparticles with various mesoporous structures, Angew. Chem. Int. Ed. 44 (2005) 288–292. [27] S.C. Shen, L. Chia, W.K. Ng, Y.C. Dong, R.B.H. Tan, Solid-phase steam-assisted synthesis of hydroxyapatite nanorods and nanoparticles, J. Mater. Sci. 22 (2010) 6059–6067. [28] X. Zhang, U.P. Wyss, D. Pichora, M.E.A. Goosen, Biodegradable controlled antibiotic release devices for osteomyelitis: optimization of release properties, J. Pharm. Pharmacol. 46 (1994) 718–724. [29] M.A. Puska, A.K. Kokkari, T.O. Närhi, P.K. Vallittu, Mechanical properties of oligomermodified acrylic bone cement, Biomaterials 24 (2003) 417–425. [30] G. Lewis, S. Janna, The in vitro elution of gentamicin sulfate from a commercially available gentamicin-loaded acrylic bone cement — VersaBond AB, J. Biomed. Mater. Res. Appl. Biomater. 71B (2004) 77–83. [31] H.P. Stallmann, C. Faber, A.L.J.J. Bronckers, A.V.N. Amerongen, P.I.L.M. Wuisman, In vitro gentamicin release from commercially available calcium-phosphate bone substitutes influence of carrier type on duration of the release profile, BMC Musculoskelet. Disord. 18 (2006) (8 pages). [32] U. Gbureck, E. Vorndran, F.A. Müller, J.E. Barralet, Low temperature direct 3D printed bioceramics and biocomposites as drug release matrices, J. Control. Release 122 (2007) 173–180.
S.-C. Shen et al. / Materials Science and Engineering C 58 (2016) 233–241 [33] H.Y. Li, J. Chang, Preparation, characterization and in vitro release of gentamicin from PHBV/wollastonite composite microspheres, J. Control. Release 107 (2005) 463–473. [34] P.L. Ritger, N.A. Peppas, A simple equation for description of solute release I. Fickian and non-Fickian release from non-swellable devices in the form of slabs, spheres, cylinders or discs, J. Control. Release 5 (1987) 23–36. [35] E. Ger, D. Dall, T. Miles, A. Forder, Bone cement and antibiotics, S. Afr. Med. J. 51 (1977) 276–279. [36] K.E. Marks, C.L. Nelson, E.P. Lautenschlager, Antibiotic-impregnated acrylic bone cement, J. Bone Joint Surg. Am. 58 (1976) 358–364. [37] M.K. Singh, T. Shokuhfar, J.J.D.A. Gracio, A.C.M.D. Sousa, J.M.D.F. Fereira, H. Garmestani, S. Ahzi, Hydroxyapatite modified with carbon-nanotube-reinforced poly(methyl methacrylate): a nanocomposite material for biomedical applications, Adv. Funct. Mater. 18 (2008) 694–700.
241
[38] R.P. del Real, S. Padilla, M. Vallet-Reg, Gentamicin release from hydroxyapatite/ poly(ethyl methacrylate)/poly(methyl methacrylate)composites, J. Biomed. Mater. Res. 52 (2000) 1–7. [39] C.S. Wu, F.T. Chiang, J.P. Hung, Effect of antibiotic additions on the mechanical properties of bone cement, Book Mater. Eng. Mech. Auto, Appl. Mech. Mater., 442 2014, pp. 147–151. [40] P. Anita, K. Brionna, T. Paul, Multi-walled carbon nanotubes induce cytotoxicity, genotoxicity and apoptosis in normal human dermal fibroblast cells, Ethn. Dis. 20 (2010) S1-65-72. [41] C.P. Firme, P.R. Bandaru, Toxicity issues in the application of carbon nanotubes to biological systems, Nanomedicine 6 (2010) 245–256. [42] Y. Seo, J. Hwang, J. Kim, Y. Yeong, M.P. Hwang, J. Choi, Antibacterial activity and cytotoxicity of multi-walled carbon nanotubes decorated with silver nanoparticles, Int. J. Nanomedicine 9 (2014) 4621–4629.