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Pulsed laser deposition of magnesium-doped calcium phosphate coatings on porous polycaprolactone scaffolds produced by rapid prototyping
Materials Letters 148 (2015) 178–183
Contents lists available at ScienceDirect
Materials Letters journal homepage: www.elsevier.com/locate/matlet
Pulsed laser deposition of magnesium-doped calcium phosphate coatings on porous polycaprolactone scaffolds produced by rapid prototyping Marta Vandrovcova a, Timothy E.L. Douglas b,1, Waldemar Mróz c, Olga Musial b, David Schaubroeck d, Boguslaw Budner c, Renata Syroka c, Peter Dubruel b, Lucie Bacakova a,n a
Dept. of Biomaterials and Tissue Engineering, Institute of Physiology of the Czech Academy of Sciences, Videnska 1083, 14220 Prague 4, Czech Republic Polymer Chemistry and Biomaterials (PBM) Group, Department of Organic Chemistry, University of Ghent, Krijgslaan 281 S4, 9000 Gent, Belgium Institute of Optoelectronics, Military University of Technology, Gen. S. Kaliskiego 2, 00-908 Warsaw, Poland d Center for Microsystems Technology (CMST), ELIS, imec, Technologiepark 914a, 9052 Ghent, Belgium b c
art ic l e i nf o
a b s t r a c t
Article history: Received 14 November 2014 Accepted 14 February 2015 Available online 23 February 2015
Polycaprolactone (PCL) is a biodegradable and biocompatible polyester whose low melting point facilitates production of 3D porous scaffolds with precisely defined dimensions and internal architecture by rapid prototyping techniques. To improve the suitability of such PCL scaffolds for bone regeneration applications, they were coated with inorganic layers of calcium phosphate (CaP) and CaP doped with 0.6% w/v magnesium (CaP þ Mg) using pulsed laser deposition (PLD) and characterized in vitro using osteoblast-like Saos-2 cells. Saos-2 cells were able to adhere to all scaffolds. CaP þMg coatings significantly increased activity of alkaline phosphatase (ALP), an early differentiation marker, after 7 days. However, gene expression of ALP after 7 days was markedly lower on the same scaffolds. These data show the feasibility of coating PCL with CaP layers by PLD and the possibility of influencing osteoblastic differentiation by magnesium doping of the CaP coating. & 2015 Published by Elsevier B.V.
Keywords: Biomaterials Composite materials Mineralization Osteoblasts Bone tissue engineering
1. Introduction Polycaprolactone (PCL) is a popular polymer in tissue engineering (TE) due to its degradability and low melting point, allowing easy processing into porous 3D structures by methods such as BioPlottings, a rapid prototyping technique. Thus allows layer-bylayer production of scaffolds with highly defined dimensions and internal architecture (e.g. porosity, pore size). However, poor cell attachment and proliferation on PCL necessitates surface modification, for example by plasma treatment [1], sometimes combined with covalent coupling of gelatin and fibronectin [2–4]. An alternative surface treatment strategy is deposition of calcium phosphate (CaP) coatings (for reviews, see [5,6]). Certain types of CaP such as hydroxyapatite (HA) (Ca10(PO4)6(OH)2), have been coated onto porous titanium-based implants to improve osseointegration [7]. One method to deposit CaP on non-smooth and porous n
Corresponding author. Tel.: þ42 296443743; fax: þ42 296442488. E-mail address: [email protected] (L. Bacakova). Current address: Nano & Biophotonics Group, Department of Molecular Biotechology, Coupure Links 653, 9000 Ghent, Belgium. 1
http://dx.doi.org/10.1016/j.matlet.2015.02.074 0167-577X/& 2015 Published by Elsevier B.V.
substrates is pulsed laser deposition (PLD). PLD has been used to coat metallic biomaterials [8–11], but PLD coating of polymeric biomaterials remains relatively unexplored. PLD permits doping with elements such as magnesium (Mg) without adversely affecting stability and biocompatibility of HA coatings [12,13]. Mg in solution has improved proliferation of human bone marrow stromal cells and extracellular matrix production and mineralization [14]. Positive effects of Mg enrichment of inorganic biomaterials in vitro have been reported. Mg-enriched HA has stimulated osteoblast adhesion [15] and proliferation [16–18]. Osteogenic differentiation of stem cells from both human bone marrow and human adipose tissue was superior on the Mg-containing bioceramic akermanite compared to the commonly used CaP bone replacement material beta-tricalcium phosphate (β-TCP) [19–21]. In another study, Mg enrichment of βTCP enhanced proliferation and osteogenic differentiation of MC3T3-E1 osteoblast-like cells [22]. Positive effects of Mg in vivo have also been reported. HA coatings enriched with magnesium on titanium implants promoted early osseointegration after 2 weeks compared to pure HA coatings [23]. Mg enrichment of TiO2 implant coatings caused increased gene expression of osteogenic differentiation markers in surrounding bone [24]. However, the effect of Mg
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Fig. 1. First row from top: porous PCL scaffold of dimensions 1 cm 1 cm produced by rapid prototyping (viewed from above). Second and third rows from top: SEM micrographs of uncoated PCL scaffolds (left column) and scaffolds coated with calcium phosphate (CaP) (middle column) and CaP doped with magnesium (CaP þ Mg) (right column). Fourth row from top: SEM micrographs showing side view of CaP layer (left) and CaP þMg layer (right). Scale bars: second row: 1 mm; third and fourth row: 1 μm.
Table 1 Parameters for the deposition of layers of calcium phosphate (CaP) and CaP doped with 0.6% Mg (CaPþ Mg). Distance from sample to target Target material Deposition temperature of implant Deposition time Number of laser shots Laser pulse duration Average pressure in experimental chamber Atmosphere Laser pulse energy Laser pulse energy on target Focus Equivalent deposition thickness measured on flat silica
657 2 mm HA or HA þ0.6% Mg 20 1C (room temperature) 144 min 15000 15–20 ns 3.2 10 2 mbar Water vapor 3007 10 mJ 203 7 10 mJ 1.3 7.9 mm2 830 7 30 nm
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on osteogenic differentiation as a component of CaP coatings, especially coatings on polymer scaffolds, remains relatively unexplored. In this study, bioplotted PCL scaffolds were coated with CaP and CaP doped with Mg (CaPþMg). Mg content was set at 0.6 w/w %, which is similar to the Mg contents in bone, dentine and enamel (0.47%, 1.11% and 0.3%, respectively [25]) and sufficient to enhance osteoblast differentiation on octacalcium phosphate (OCP) coatings [10]. In vivo, HA enriched with lower weight percentages of Mg has promoted osseointegration and bone ingrowth [22,26]. Furthermore, our previous work on PLD of Mg-containing HA coatings suggests effective incorporation of 0.6 w/w % Mg into the HA structure [13]. In this work, Mg incorporation in the deposited CaPþ Mg coatings on PCL scaffolds was conclusively proved. To our best knowledge, there are very few, if any studies on PLD coating of polymer scaffolds with CaP. Coated scaffolds were characterized biologically in vitro using osteoblast-like Saos-2 cells. Effects of surface modification on gene expression of the osteogenic markers collagen type I, alkaline phosphatase (ALP), and osteocalcin were determined using quantitative real-time polymerase chain reaction (q-PCR). Activity of ALP was also measured.
5 min. After halting the reaction by adding 0.8 ml 1 M NaOH the absorbance was measured at 405 nm. The number of viable cells measured by CellTiter 96s AQueous One Solution Cell Proliferation Assay (MTS; Promega, Cat. No. G3581) was used to relativize the value of ALP activity according to the manufacturer’s protocol. PCR data and ALP activity values were presented as mean 7S.D. (Standard Deviation) from 4 and 9 measurements, respectively.
2. Materials and methods Production of PCL porous scaffolds by rapid prototyping, coating by PLD, scanning electron microscopy (SEM), energy-dispersive spectroscopy (EDS) and Attenuated Total Reflectance Fourier Transform Infrared Spectroscopy (ATR-FTIR) analysis: Porous scaffolds of PCL (MW ¼80,000 g/mol) of dimensions 1 cm 1 cm (Fig. 1), pore side length 0.5 mm were prepared using a Bioscaffolder device (Sys-Eng) as described previously [1–3]. Production of HA and HA þMg (0.6 w/w %) targets and their deposition by the PLD method was performed as described previously [12,27]. Vapor deposition parameters are shown in Table 1. Three experimental groups were prepared: uncoated PCL, PCL coated with CaP and PCL coated with CaPþ Mg. SEM and EDS were performed on a JSM-5600 instrument (JEOL) in secondary electron mode (SEI) equipped with an electron microprobe JED 2300 and an EDS detector. Samples were coated with a thin gold layer as described previously [28]. EDS analysis was performed in the mapping mode. Three measurements were performed on each sample on random spots (65 mm by 50 mm). ATR-FTIR measurements were performed using a Nicolet IS50 model (ThermoSCIENTIFIC), wavenumber range 4000–400 cm 1, resolution 4 cm 1. Cell culture studies: Prior to cell seeding, samples were sterilized with ethylene oxide as described previously [29]. Samples in 24-well polystyrene cell culture plates were seeded with Saos-2 cells (European Collection of Cell Cultures, Cat. No. 89050205), suspended in McCoy medium (Sigma, Cat. No. M4892) supplemented with 15% fetal bovine serum (Sebak GmbH), and gentamicin (40 mg/ml, LEK) at a density of 100,000 cells per scaffold and 2 ml of the medium. The cells were allowed to adhere for 2 h before medium (2 mL) was added. After 24 h, well plates were transferred to dynamic conditions using a MiniOrbital Shaker SSM1 (planar rotation, speed 70 rpm) and cultured for 7, and 14 days at 37 1C in a humidified air atmosphere containing 5% CO2. For q-PCR, RNA extraction and reverse transcription were performed after 7 and 14 days as described previously [30]. Endogenous GAPDH served as a so-called housekeeping gene to normalize expression levels. Primers were designed according to the literature [31,32]. ALP activity was evaluated on day 7 after seeding. Samples were rinsed twice with phosphate-buffered saline (PBS), and transferred to new 24-well plates. The cells were incubated with 1 ml of 270 mM pnitrophenyl-phosphate in 50 mM glycine, 1 mM MgCl2, pH 10.5 for
Fig. 2. Attenuated Total Reflectance Fourier Transform Infrared spectroscopy (ATRFTIR) examination; from top: uncoated PCL scaffolds, HA target, scaffolds coated with calcium phosphate (CaP), and scaffolds coated with CaP doped with magnesium (CaP þ Mg). Spectra of HA and HA þMg targets were identical, hence, only the ATR-FTIR spectrum of the HA target is shown.
M. Vandrovcova et al. / Materials Letters 148 (2015) 178–183
Multiple comparison procedures were performed with ANOVA, using the Student-Newman–Keuls method. Pr 0.05 was considered significant.
3. Results and discussion SEM analysis confirmed that both CaP and CaPþ Mg coatings appeared similar and adherent to the substrate. Hence, Mg does not affect coating morphology. At the top of deposited layers some grains were visible. Most were roughly spherical with a diameter of approximately 1 μm. Some elongated cuboid-like deposits approximately 1–2 μm in length were also observed (Fig. 1). The thicknesses of the CaP coatings were variable. The maximum thicknesses (approximately 1–1.5 mm) of the CaP þMg coating, were observed on surfaces perpendicular to the expanding plasma (Fig. 1, fourth row, right). Thicknesses of CaP coatings on non-perpendicular surfaces were lower, which can be explained by greater plasma expansion and material transfer in the perpendicular direction. In the chosen cross section of the CaP coating (Fig. 1, fourth row, left) the thickness of deposited coating was approximately 0.5 mm. The purpose of the CaP coatings is to promote cell attachment and proliferation, for which a 0.5 mm thick coating is sufficient. After implantation in vivo, resorption of PCL scaffolds and CaP coating is expected.
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EDS analysis demonstrated the absence of Mg on uncoated scaffolds. CaPþ Mg-coated scaffolds contained 0.59 70.13% mass percentage of Mg. ATR-FTIR analysis also proved CaP coating deposition (Fig. 2). Strong bands typical for PCL at 1750 cm 1 and 3000–2850 cm 1 (C ¼O carbonyl stretching and C–H asymmetrical and symmetrical stretching, respectively [33]), were weaker in coated samples. Spectra of PCL scaffolds coated with CaP and CaP þMg also showed no obvious differences, demonstrating that magnesium did not significantly influence the type of CaP phase formed. Bands observed in the regions 600–500 cm 1 and 1200– 800 cm 1 are characteristic for phosphate υ4 and υ3 stretching, respectively [34]. For HA, these bands have a fine structure [12]. In deposited films these bands were registered as one broad band which indicates that an amorphous CaP is the dominant phase. The low crystallinity of the CaP coatings is connected with their deposition at room temperature, which was necessary due to the low degradation temperature of the PCL scaffolds. To avoid reduction of –OH groups, the process of deposition of CaP coatings by the PLD method was carried out in a water vapor atmosphere. The broad band with a main peak at approximately 3500 cm 1 is characteristic for stretching modes of –OH coming from absorbed H2O as well from hydroxyl groups present in the CaP structure. This peak is slightly less intense for CaP coatings doped with Mg. Our earlier measurements of HA coatings doped with Mg also
Fig. 3. Leica TCS SP2 confocal microscope, objective 10 , bar ¼250 μm, cells on 3D scaffolds on day 3 after seeding. Legend: Turquoise blue¼ surface of the scaffold, red color ¼ maximal depth (in mm), U¼ upper layer, L ¼lower layer, Asterisk¼pore. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
Fig. 4. (A) Activity of alkaline phosphatase (ALP) expressed as a ratio of absorbance of ALP product and absorbance of MTS test (number of cells). ALP activity was determined in Saos-2 cells cultured on untreated PCL samples and PCL coated with CaP and CaP þ Mg and compared to ALP activity of cells cultured on PS on day 7 after seeding. Mean 7 standard deviation from 9 measurements (3 samples). *,p r 0.05 compared to CaP þMg and PS, respectively. (B and C) Gene expression of collagen type I (col 1), early osteogenic differentiation marker alkaline phosphatase (ALP) and late osteogenic differentiation marker osteocalcin (OC) after culture for 7 days (B) and 14 days (C) on samples of uncoated PCL, PCL coated with CaP and PCL coated with CaP þMg. Error bars show standard deviation.
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suggested that doping of CaP coatings with Mg increases their hydrophobicity [12]. Confocal microscopy on day 3 after seeding revealed well spread cells homogenously distributed on the sample struts (Fig. 3). The overall relative activity of ALP was significantly higher on samples coated with CaPþ Mg than on all other samples (Fig. 4A). In a study on MC3T3-E1 osteoblast-like cells cultured on Mg-substituted βTCP, ALP activity was increased significantly after 7 and 14 days at a substitution degree of Z14 mol%, but not at lower substitution degrees [22]. In another study on MG-63 cells cultured on apatitic coatings substituted with approximately 2 at% Mg, collagen I production but not ALP activity increased after 7 days [17]. In another study, it was reported that MC3T3-E1 osteoblast-like cells displayed higher ALP activity after 7 and 14 days on HA coatings on titanium substituted with approximately 5% Mg [23,35]. Positive effects of Mg on osteogenic cell differentiation have also been reported for titanium implanted with Mg ions [36], and gelatin sponges with various amounts of incorporated magnesium calcium phosphate [37]. In the present study, a lower Mg content than in the aforementioned studies (0.6%) was sufficient to increase ALP activity of Saos-2 after 7 days. Comparing the results of the present study to those obtained by other groups, it can be speculated that the effect of Mg on ALP activity is dependent on cell type. It should also be mentioned that relative activity of ALP was higher on all the tested samples than on control polystyrene. Q-PCR results were similar after both 7 and 14 days (Fig. 4B and C). ALP gene expression displayed an opposite tendency to ALP activity, possibly due to the timing of individual steps during translation. ALP is an early marker of osteogenic differentiation. First, mRNA is transcripted from DNA and amplified, then translated into protein. When the amount of protein is sufficient, mRNA undergoes turnover and disappears, however the protein remains. Collagen type I expression was significantly upregulated on all PCL samples in comparison with PS, but was the highest on uncoated PCL. There were no significant differences in osteocalcin expression between the tested samples, with the exception of PCL samples coated with CaP, on which expression was lower. These results suggest that Mg-enhanced CaP coatings on PCL enhance osteogenic differentiation, while Mg-free coatings do not when compared to pure PCL. Further work is needed to determine the optimum Mg content to promote differentiation and assess the ability of coated scaffolds to support bone regeneration in vivo. The results prove the principle of coating polymer scaffolds with CaP using PLD.
4. Conclusions CaP and CaP þMg coatings were successfully deposited by PLD at room temperature on porous PCL scaffolds fabricated by rapid prototyping. All PCL scaffolds supported Saos-2 cell attachment and growth. Activity of the early osteogenic differentiation marker ALP was highest on samples coated with CaPþ Mg. The results show that CaP coatings can improve behaviour of bone cells seeded on PCL with a view to bone TE.
Acknowledgement T.E.L.D. acknowledges FWO, Belgium for a postdoctoral fellowship. M.V. and L.B. acknowledge Grant Agency of the Czech Republic for financial support (“Center of Excellence”, grant no. P108/12/G108). W.M. acknowledges financial support from the Polish Part of the Eureka project no. 57/N-Eureka/2007.
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Contents lists available at ScienceDirect
Materials Letters journal homepage: www.elsevier.com/locate/matlet
Pulsed laser deposition of magnesium-doped calcium phosphate coatings on porous polycaprolactone scaffolds produced by rapid prototyping Marta Vandrovcova a, Timothy E.L. Douglas b,1, Waldemar Mróz c, Olga Musial b, David Schaubroeck d, Boguslaw Budner c, Renata Syroka c, Peter Dubruel b, Lucie Bacakova a,n a
Dept. of Biomaterials and Tissue Engineering, Institute of Physiology of the Czech Academy of Sciences, Videnska 1083, 14220 Prague 4, Czech Republic Polymer Chemistry and Biomaterials (PBM) Group, Department of Organic Chemistry, University of Ghent, Krijgslaan 281 S4, 9000 Gent, Belgium Institute of Optoelectronics, Military University of Technology, Gen. S. Kaliskiego 2, 00-908 Warsaw, Poland d Center for Microsystems Technology (CMST), ELIS, imec, Technologiepark 914a, 9052 Ghent, Belgium b c
art ic l e i nf o
a b s t r a c t
Article history: Received 14 November 2014 Accepted 14 February 2015 Available online 23 February 2015
Polycaprolactone (PCL) is a biodegradable and biocompatible polyester whose low melting point facilitates production of 3D porous scaffolds with precisely defined dimensions and internal architecture by rapid prototyping techniques. To improve the suitability of such PCL scaffolds for bone regeneration applications, they were coated with inorganic layers of calcium phosphate (CaP) and CaP doped with 0.6% w/v magnesium (CaP þ Mg) using pulsed laser deposition (PLD) and characterized in vitro using osteoblast-like Saos-2 cells. Saos-2 cells were able to adhere to all scaffolds. CaP þMg coatings significantly increased activity of alkaline phosphatase (ALP), an early differentiation marker, after 7 days. However, gene expression of ALP after 7 days was markedly lower on the same scaffolds. These data show the feasibility of coating PCL with CaP layers by PLD and the possibility of influencing osteoblastic differentiation by magnesium doping of the CaP coating. & 2015 Published by Elsevier B.V.
Keywords: Biomaterials Composite materials Mineralization Osteoblasts Bone tissue engineering
1. Introduction Polycaprolactone (PCL) is a popular polymer in tissue engineering (TE) due to its degradability and low melting point, allowing easy processing into porous 3D structures by methods such as BioPlottings, a rapid prototyping technique. Thus allows layer-bylayer production of scaffolds with highly defined dimensions and internal architecture (e.g. porosity, pore size). However, poor cell attachment and proliferation on PCL necessitates surface modification, for example by plasma treatment [1], sometimes combined with covalent coupling of gelatin and fibronectin [2–4]. An alternative surface treatment strategy is deposition of calcium phosphate (CaP) coatings (for reviews, see [5,6]). Certain types of CaP such as hydroxyapatite (HA) (Ca10(PO4)6(OH)2), have been coated onto porous titanium-based implants to improve osseointegration [7]. One method to deposit CaP on non-smooth and porous n
Corresponding author. Tel.: þ42 296443743; fax: þ42 296442488. E-mail address: [email protected] (L. Bacakova). Current address: Nano & Biophotonics Group, Department of Molecular Biotechology, Coupure Links 653, 9000 Ghent, Belgium. 1
http://dx.doi.org/10.1016/j.matlet.2015.02.074 0167-577X/& 2015 Published by Elsevier B.V.
substrates is pulsed laser deposition (PLD). PLD has been used to coat metallic biomaterials [8–11], but PLD coating of polymeric biomaterials remains relatively unexplored. PLD permits doping with elements such as magnesium (Mg) without adversely affecting stability and biocompatibility of HA coatings [12,13]. Mg in solution has improved proliferation of human bone marrow stromal cells and extracellular matrix production and mineralization [14]. Positive effects of Mg enrichment of inorganic biomaterials in vitro have been reported. Mg-enriched HA has stimulated osteoblast adhesion [15] and proliferation [16–18]. Osteogenic differentiation of stem cells from both human bone marrow and human adipose tissue was superior on the Mg-containing bioceramic akermanite compared to the commonly used CaP bone replacement material beta-tricalcium phosphate (β-TCP) [19–21]. In another study, Mg enrichment of βTCP enhanced proliferation and osteogenic differentiation of MC3T3-E1 osteoblast-like cells [22]. Positive effects of Mg in vivo have also been reported. HA coatings enriched with magnesium on titanium implants promoted early osseointegration after 2 weeks compared to pure HA coatings [23]. Mg enrichment of TiO2 implant coatings caused increased gene expression of osteogenic differentiation markers in surrounding bone [24]. However, the effect of Mg
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Fig. 1. First row from top: porous PCL scaffold of dimensions 1 cm 1 cm produced by rapid prototyping (viewed from above). Second and third rows from top: SEM micrographs of uncoated PCL scaffolds (left column) and scaffolds coated with calcium phosphate (CaP) (middle column) and CaP doped with magnesium (CaP þ Mg) (right column). Fourth row from top: SEM micrographs showing side view of CaP layer (left) and CaP þMg layer (right). Scale bars: second row: 1 mm; third and fourth row: 1 μm.
Table 1 Parameters for the deposition of layers of calcium phosphate (CaP) and CaP doped with 0.6% Mg (CaPþ Mg). Distance from sample to target Target material Deposition temperature of implant Deposition time Number of laser shots Laser pulse duration Average pressure in experimental chamber Atmosphere Laser pulse energy Laser pulse energy on target Focus Equivalent deposition thickness measured on flat silica
657 2 mm HA or HA þ0.6% Mg 20 1C (room temperature) 144 min 15000 15–20 ns 3.2 10 2 mbar Water vapor 3007 10 mJ 203 7 10 mJ 1.3 7.9 mm2 830 7 30 nm
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on osteogenic differentiation as a component of CaP coatings, especially coatings on polymer scaffolds, remains relatively unexplored. In this study, bioplotted PCL scaffolds were coated with CaP and CaP doped with Mg (CaPþMg). Mg content was set at 0.6 w/w %, which is similar to the Mg contents in bone, dentine and enamel (0.47%, 1.11% and 0.3%, respectively [25]) and sufficient to enhance osteoblast differentiation on octacalcium phosphate (OCP) coatings [10]. In vivo, HA enriched with lower weight percentages of Mg has promoted osseointegration and bone ingrowth [22,26]. Furthermore, our previous work on PLD of Mg-containing HA coatings suggests effective incorporation of 0.6 w/w % Mg into the HA structure [13]. In this work, Mg incorporation in the deposited CaPþ Mg coatings on PCL scaffolds was conclusively proved. To our best knowledge, there are very few, if any studies on PLD coating of polymer scaffolds with CaP. Coated scaffolds were characterized biologically in vitro using osteoblast-like Saos-2 cells. Effects of surface modification on gene expression of the osteogenic markers collagen type I, alkaline phosphatase (ALP), and osteocalcin were determined using quantitative real-time polymerase chain reaction (q-PCR). Activity of ALP was also measured.
5 min. After halting the reaction by adding 0.8 ml 1 M NaOH the absorbance was measured at 405 nm. The number of viable cells measured by CellTiter 96s AQueous One Solution Cell Proliferation Assay (MTS; Promega, Cat. No. G3581) was used to relativize the value of ALP activity according to the manufacturer’s protocol. PCR data and ALP activity values were presented as mean 7S.D. (Standard Deviation) from 4 and 9 measurements, respectively.
2. Materials and methods Production of PCL porous scaffolds by rapid prototyping, coating by PLD, scanning electron microscopy (SEM), energy-dispersive spectroscopy (EDS) and Attenuated Total Reflectance Fourier Transform Infrared Spectroscopy (ATR-FTIR) analysis: Porous scaffolds of PCL (MW ¼80,000 g/mol) of dimensions 1 cm 1 cm (Fig. 1), pore side length 0.5 mm were prepared using a Bioscaffolder device (Sys-Eng) as described previously [1–3]. Production of HA and HA þMg (0.6 w/w %) targets and their deposition by the PLD method was performed as described previously [12,27]. Vapor deposition parameters are shown in Table 1. Three experimental groups were prepared: uncoated PCL, PCL coated with CaP and PCL coated with CaPþ Mg. SEM and EDS were performed on a JSM-5600 instrument (JEOL) in secondary electron mode (SEI) equipped with an electron microprobe JED 2300 and an EDS detector. Samples were coated with a thin gold layer as described previously [28]. EDS analysis was performed in the mapping mode. Three measurements were performed on each sample on random spots (65 mm by 50 mm). ATR-FTIR measurements were performed using a Nicolet IS50 model (ThermoSCIENTIFIC), wavenumber range 4000–400 cm 1, resolution 4 cm 1. Cell culture studies: Prior to cell seeding, samples were sterilized with ethylene oxide as described previously [29]. Samples in 24-well polystyrene cell culture plates were seeded with Saos-2 cells (European Collection of Cell Cultures, Cat. No. 89050205), suspended in McCoy medium (Sigma, Cat. No. M4892) supplemented with 15% fetal bovine serum (Sebak GmbH), and gentamicin (40 mg/ml, LEK) at a density of 100,000 cells per scaffold and 2 ml of the medium. The cells were allowed to adhere for 2 h before medium (2 mL) was added. After 24 h, well plates were transferred to dynamic conditions using a MiniOrbital Shaker SSM1 (planar rotation, speed 70 rpm) and cultured for 7, and 14 days at 37 1C in a humidified air atmosphere containing 5% CO2. For q-PCR, RNA extraction and reverse transcription were performed after 7 and 14 days as described previously [30]. Endogenous GAPDH served as a so-called housekeeping gene to normalize expression levels. Primers were designed according to the literature [31,32]. ALP activity was evaluated on day 7 after seeding. Samples were rinsed twice with phosphate-buffered saline (PBS), and transferred to new 24-well plates. The cells were incubated with 1 ml of 270 mM pnitrophenyl-phosphate in 50 mM glycine, 1 mM MgCl2, pH 10.5 for
Fig. 2. Attenuated Total Reflectance Fourier Transform Infrared spectroscopy (ATRFTIR) examination; from top: uncoated PCL scaffolds, HA target, scaffolds coated with calcium phosphate (CaP), and scaffolds coated with CaP doped with magnesium (CaP þ Mg). Spectra of HA and HA þMg targets were identical, hence, only the ATR-FTIR spectrum of the HA target is shown.
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Multiple comparison procedures were performed with ANOVA, using the Student-Newman–Keuls method. Pr 0.05 was considered significant.
3. Results and discussion SEM analysis confirmed that both CaP and CaPþ Mg coatings appeared similar and adherent to the substrate. Hence, Mg does not affect coating morphology. At the top of deposited layers some grains were visible. Most were roughly spherical with a diameter of approximately 1 μm. Some elongated cuboid-like deposits approximately 1–2 μm in length were also observed (Fig. 1). The thicknesses of the CaP coatings were variable. The maximum thicknesses (approximately 1–1.5 mm) of the CaP þMg coating, were observed on surfaces perpendicular to the expanding plasma (Fig. 1, fourth row, right). Thicknesses of CaP coatings on non-perpendicular surfaces were lower, which can be explained by greater plasma expansion and material transfer in the perpendicular direction. In the chosen cross section of the CaP coating (Fig. 1, fourth row, left) the thickness of deposited coating was approximately 0.5 mm. The purpose of the CaP coatings is to promote cell attachment and proliferation, for which a 0.5 mm thick coating is sufficient. After implantation in vivo, resorption of PCL scaffolds and CaP coating is expected.
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EDS analysis demonstrated the absence of Mg on uncoated scaffolds. CaPþ Mg-coated scaffolds contained 0.59 70.13% mass percentage of Mg. ATR-FTIR analysis also proved CaP coating deposition (Fig. 2). Strong bands typical for PCL at 1750 cm 1 and 3000–2850 cm 1 (C ¼O carbonyl stretching and C–H asymmetrical and symmetrical stretching, respectively [33]), were weaker in coated samples. Spectra of PCL scaffolds coated with CaP and CaP þMg also showed no obvious differences, demonstrating that magnesium did not significantly influence the type of CaP phase formed. Bands observed in the regions 600–500 cm 1 and 1200– 800 cm 1 are characteristic for phosphate υ4 and υ3 stretching, respectively [34]. For HA, these bands have a fine structure [12]. In deposited films these bands were registered as one broad band which indicates that an amorphous CaP is the dominant phase. The low crystallinity of the CaP coatings is connected with their deposition at room temperature, which was necessary due to the low degradation temperature of the PCL scaffolds. To avoid reduction of –OH groups, the process of deposition of CaP coatings by the PLD method was carried out in a water vapor atmosphere. The broad band with a main peak at approximately 3500 cm 1 is characteristic for stretching modes of –OH coming from absorbed H2O as well from hydroxyl groups present in the CaP structure. This peak is slightly less intense for CaP coatings doped with Mg. Our earlier measurements of HA coatings doped with Mg also
Fig. 3. Leica TCS SP2 confocal microscope, objective 10 , bar ¼250 μm, cells on 3D scaffolds on day 3 after seeding. Legend: Turquoise blue¼ surface of the scaffold, red color ¼ maximal depth (in mm), U¼ upper layer, L ¼lower layer, Asterisk¼pore. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
Fig. 4. (A) Activity of alkaline phosphatase (ALP) expressed as a ratio of absorbance of ALP product and absorbance of MTS test (number of cells). ALP activity was determined in Saos-2 cells cultured on untreated PCL samples and PCL coated with CaP and CaP þ Mg and compared to ALP activity of cells cultured on PS on day 7 after seeding. Mean 7 standard deviation from 9 measurements (3 samples). *,p r 0.05 compared to CaP þMg and PS, respectively. (B and C) Gene expression of collagen type I (col 1), early osteogenic differentiation marker alkaline phosphatase (ALP) and late osteogenic differentiation marker osteocalcin (OC) after culture for 7 days (B) and 14 days (C) on samples of uncoated PCL, PCL coated with CaP and PCL coated with CaP þMg. Error bars show standard deviation.
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suggested that doping of CaP coatings with Mg increases their hydrophobicity [12]. Confocal microscopy on day 3 after seeding revealed well spread cells homogenously distributed on the sample struts (Fig. 3). The overall relative activity of ALP was significantly higher on samples coated with CaPþ Mg than on all other samples (Fig. 4A). In a study on MC3T3-E1 osteoblast-like cells cultured on Mg-substituted βTCP, ALP activity was increased significantly after 7 and 14 days at a substitution degree of Z14 mol%, but not at lower substitution degrees [22]. In another study on MG-63 cells cultured on apatitic coatings substituted with approximately 2 at% Mg, collagen I production but not ALP activity increased after 7 days [17]. In another study, it was reported that MC3T3-E1 osteoblast-like cells displayed higher ALP activity after 7 and 14 days on HA coatings on titanium substituted with approximately 5% Mg [23,35]. Positive effects of Mg on osteogenic cell differentiation have also been reported for titanium implanted with Mg ions [36], and gelatin sponges with various amounts of incorporated magnesium calcium phosphate [37]. In the present study, a lower Mg content than in the aforementioned studies (0.6%) was sufficient to increase ALP activity of Saos-2 after 7 days. Comparing the results of the present study to those obtained by other groups, it can be speculated that the effect of Mg on ALP activity is dependent on cell type. It should also be mentioned that relative activity of ALP was higher on all the tested samples than on control polystyrene. Q-PCR results were similar after both 7 and 14 days (Fig. 4B and C). ALP gene expression displayed an opposite tendency to ALP activity, possibly due to the timing of individual steps during translation. ALP is an early marker of osteogenic differentiation. First, mRNA is transcripted from DNA and amplified, then translated into protein. When the amount of protein is sufficient, mRNA undergoes turnover and disappears, however the protein remains. Collagen type I expression was significantly upregulated on all PCL samples in comparison with PS, but was the highest on uncoated PCL. There were no significant differences in osteocalcin expression between the tested samples, with the exception of PCL samples coated with CaP, on which expression was lower. These results suggest that Mg-enhanced CaP coatings on PCL enhance osteogenic differentiation, while Mg-free coatings do not when compared to pure PCL. Further work is needed to determine the optimum Mg content to promote differentiation and assess the ability of coated scaffolds to support bone regeneration in vivo. The results prove the principle of coating polymer scaffolds with CaP using PLD.
4. Conclusions CaP and CaP þMg coatings were successfully deposited by PLD at room temperature on porous PCL scaffolds fabricated by rapid prototyping. All PCL scaffolds supported Saos-2 cell attachment and growth. Activity of the early osteogenic differentiation marker ALP was highest on samples coated with CaPþ Mg. The results show that CaP coatings can improve behaviour of bone cells seeded on PCL with a view to bone TE.
Acknowledgement T.E.L.D. acknowledges FWO, Belgium for a postdoctoral fellowship. M.V. and L.B. acknowledge Grant Agency of the Czech Republic for financial support (“Center of Excellence”, grant no. P108/12/G108). W.M. acknowledges financial support from the Polish Part of the Eureka project no. 57/N-Eureka/2007.
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