Electrical resistivity of nanocrystalline Fe73.5Nb4.5Cr5Cu1B16 alloys

Journal of Controlled Release 170 (2013) 99–110

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New PLGA–P188–PLGA matrix enhances TGF-β3 release from pharmacologically active microcarriers and promotes chondrogenesis of mesenchymal stem cells Marie Morille a, b, e, Tran Van-Thanh a, e, Xavier Garric c, Jérôme Cayon d, Jean Coudane c, Danièle Noël b, Marie-Claire Venier-Julienne a, e, Claudia N. Montero-Menei a, e,⁎ a

LUNAM Université, Micro et Nanomédecines Biomimétiques (MINT), F-49933 Angers, France INSERM U844, Cellules Souches Mésenchymateuses, Environnement Articulaire et Immunothérapies de la Polyarthrite Rhumatoïde, F-34295 Montpellier, France c Institut des biomolécules Max Mousseron, CNRS UMR 5247, Montpellier, F-34093 France d Service Commun de Cytométrie et d'Analyse Nucléotidiques (SCCAN), F-49933 Angers, France e INSERM U1066, F-49933 Angers, France b

a r t i c l e

i n f o

Article history: Received 17 December 2012 Accepted 27 April 2013 Available online 3 May 2013 Keywords: Protein encapsulation Pharmacologically active microcarriers Mesenchymal stem cells Transforming growth factor Chondrogenic differentiation

a b s t r a c t The use of injectable scaffolding materials for in vivo tissue regeneration has raised great interest in various clinical applications because it allows cell implantation through minimally invasive surgical procedures. In case of cartilage repair, a tissue engineered construct should provide a support for the cell and allow sustained in situ delivery of bioactive factors capable of inducing cell differentiation into chondrocytes. Pharmacologically active microcarriers (PAMs), made of biodegradable poly(D,L-lactide–co-glycolide acid) (PLGA), are a unique system, which combines these properties in an adaptable and simple microdevice. However, a limitation of such scaffold is low and incomplete protein release that occurs using the hydrophobic PLGA based microspheres. To circumvent this problem, we developed a novel formulation of polymeric PAMs containing a P188 poloxamer, which protects the protein from denaturation and may positively affect chondrogenesis. This poloxamer was added as a free additive for protein complexation and as a component of the scaffold covalently linked to PLGA. This procedure allows getting a more hydrophilic scaffold but also retaining the protective polymer inside the microcarriers during their degradation. The novel PLGA–P188–PLGA PAMs presenting a fibronectin-covered surface allowed enhanced MSC survival and proliferation. When engineered with TGFβ3, they allowed the sustained release of 70% of the incorporated TGF-β3 over time. Importantly, they exerted superior chondrogenic differentiation potential compared to previous FN-PAM-PLGA-TGF-β3, as shown by an increased expression of specific cartilage markers such as cartilage type II, aggrecan and COMP. Therefore, this microdevice represents an efficient easy-to-handle and injectable tool for cartilage repair. © 2013 Elsevier B.V. All rights reserved.

1 . Introduction Cartilage is primarily composed of chondrocytes able to generate extracellular matrix proteins, mainly type II collagen and aggrecan, which provide the tissue its resistance to tensile and compressive strength. Degradation of extracellular matrix components is the result of cartilage dysfunction with age, injury, and diseases such as osteoarthritis (OA) [1,2]. Unlike other connective tissues, cartilage is not vascularized and suffers from poor healing capacity. Because current treatment options are not satisfying, there is a strong need for the development of new strategies to efficiently restore cartilage function and avoid surgery. A current method of treatment consists in

⁎ Corresponding author. Tel.: +33 244688536; fax: +33 244688546. E-mail address: [email protected] (C.N. Montero-Menei). 0168-3659/$ – see front matter © 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.jconrel.2013.04.017

drilling holes in the subchondral bone to allow influx of bone marrow cells, which fill and repair the lesion. Nevertheless, the resulting repair generally results in fibrocartilage on the long term, which doesn't fulfill the characteristics of hyaline cartilage [3,4]. One recent approach for cartilage repair is cell-based therapy using expanded chondrocytes harvested from the patient (autologous chondrocyte implantation (ACI)). However, limitation in the amount of tissue that can be harvested for cell isolation, morbidity at the harvest site and dedifferentiation during in vitro expansion have prompted researchers to focus on new strategies especially those based on the use of stem cells [5–7]. In this context, mesenchymal stem cells (MSCs) appear as an attractive cell source for cartilage engineering because of their accessibility from donors, the ease of isolation and in vitro expansion in high numbers [8]. Indeed, MSCs exhibit the capacity to differentiate into several lineages and in particular, to chondrocytes [9–11]. They are being evaluated in some phase I clinical trials but most of these

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studies rely on the implantation of cells alone. The drawback of such approach is the low survival of implanted cells and poor differentiation capacities, highlighting the need for tissue-engineered constructs [12,13]. An attractive approach in tissue engineering consists in constructing scaffolds which provide the transplanted cells the environmental cues, typically three-dimensional (3D) structure, extracellular matrix molecules and growth factors to increase cell survival and guide the in vivo cell fate [12]. To our knowledge, the majority of existing systems use combinations of devices to provide such micro-environmental cues rendering their translation to the clinic quite problematic [14–17]. Pharmacologically active microcarriers (PAMs) hold great promise as a smart, unique fibronectin-covered surface injectable microdevice allowing delivery of both cells and sustained release of growth factors. Indeed, PAMs are biocompatible and biodegradable microspheres (MS) engineered to continuously release an active protein and present a surface of extracellular matrix molecules supplying a three-dimensional (3D) scaffold for the transported cells [18]. These combined parameters are present in one single easy-to-use system that acts both on the transported cells and the surrounding tissue [19–22]. We previously demonstrated the potential of poly(lactic-co-glycolic acid) (PLGA)-based fibronectin (FN)-coated PAMs releasing transforming growth factor 3 (TGF-β3) for cartilage tissue engineering [23]. However, this formulation of PAMs allowed incomplete release of encapsulated protein. This was mainly due to protein-polymer interaction during the process and protein degradation related to acidic environment following degradation of PLGA [24,25]. Nevertheless, PLGA-based polymers still hold great promise and various strategies have been envisaged to improve protein release from such MS [26]. Indeed, controlled protein release over time still remains a technological challenge that needs to be overcome in order to propose an efficient and fully characterized pharmacologically active device. With the aim to enhance protection of the protein from degradation and more efficient release, our group developed a strategy based on a reversible nanoprecipitate of the protein in association with a triblock copolymer poloxamer P188 (poly(ethylene oxide) (PEO)– poly(propylene oxide) (PPO)–poly(ethylene oxide)PEO) leading to protein protection, especially during the formulation process [27–31]. After this, MS were prepared by a non-denaturing solid-in-oil-in-water (s/o/w) emulsion evaporation/extraction technique. Then, polyethylene glycol (PEG) segments were introduced into hydrophobic PLGA polyesters increasing protein release from thus formed PLGA–PEG– PLGA (ABA) triblock copolymer-based MS [29,32,33]. Nevertheless, the protein stabilization within the matrix during polymer degradation (pH drop) is not efficient [32]. The non-ionic surfactants P188 poloxamers should protect proteins more efficiently than PEG units alone. Indeed, the existence of both hydrophobic and hydrophilic moieties of various lengths, should allow fine-tuning of this balance to match the hydrophilic/hydrophobic profile of one or another protein depending of its amino acid composition [34,35]. To further improve protein release, a new advance of MS formulation could be to use both of the following forms of polymer P188: (i) free and as an additive nanoprecipitated with the protein and (ii) covalently linked to the initial matrix polymer, to retain this protective polymer inside the MS during degradation. The blending of poloxamer and PLGA was not considered here because it was previously shown that it increased the initial protein release [29]. We characterized this novel PAM formulation and further compared it to the previous PLGA–PAMs, in view of their physico-chemical properties: size, coating homogeneity, protein loading capacity, and protein release. The most commonly used medium to stimulate MSC chondrogenesis contains TGF-β, and although the three isoforms TGF-β1, TGF-β2 and TGF-β3 are known to induce chondrogenesis, investigators mostly use TGF-β1 or TGF-β3 [36]. Of these two isoforms, TGF-β3 is reported to have a higher chondrogenic potential than TGF-β1 and to lead to a more rapid differentiation [37]. We therefore have chosen to use TGF-β3 as a chondrogenic inductor

in this study. We next evaluated PAM's ability to influence MSC behavior in terms of cell proliferation and their impact on MSC chondrogenesis at the gene and protein levels when TGF-β3 was released from the PAMs. 2. Materials and methods 2.1 . Materials Polyvinyl alcohol (Mowiol® 4-88) was obtained from Kuraray Specialities Europe (Frankfurt, Germany). P188 poloxamer or Pluronic® F68 was kindly supplied by BASF (Levallois-Perret, France). Culture mediums, penicillin, streptomycin and trypsin were obtained from Lonza (Levallois, France). Uncapped (free carboxylic acid group at the terminal end) PLGA37.5/25 (Mn 14,000 Da) was provided by Phusis (Saint-Ismier, France). DL-lactide and glycolide were obtained from Purac (Gorinchem, The Netherlands) and poly(ethylene oxide) from Fluka. PLGA–P188–PLGA were synthesized by IBMM-CRBA CNRS UMR 5247 (Montpellier, France) as described in §2.2.1. Polytetrafluoroethylene (PTFE) filters Millex®-FH (pore size 0.45 μm) were obtained from Millipore (Millipore SA, Guyancourt, France). Chemical reagents in general were purchased from Sigma Aldrich (Saint-Quentin-Fallavier, France) unless otherwise stated. TGF-β3 was purchased from Peprotech (Paris, France). Human basic fibroblast growth factors (bFGF), as well as specific TGF-β3 DuoSet ELISA Development kits were purchased from R&D Systems (Lille, France). The SEAP Reporter Gene Assay, chemiluminescent kit was provided by Roche (Meylan, France). Biotinylated anti-mouse IgG antibodies were obtained from Vector laboratories (Burlingame, USA), streptavidin– fluoroprobe 547 from Interchim (Montluçon, France). 1.9 cm2 Costar ultra-low cluster plate was obtained from Corning (Avon, France) and 96-well flat-bottom culture plates from Nunc® (Dutscher, France). Cyquant cell proliferation assay® and SuperScript™ II Reverse Transcriptase were obtained from Invitrogen (France). Total RNA isolation Nucleospin® RNA II was purchased from Macherey Nagel (Hoerdt, France) and cDNAs were purified with Qiaquick PCR purification kit obtained from Qiagen (Courtaboeuf, France). iQ SYBR Green Supermix was acquired from Biorad. Anti aggrecan polyclonal rabbit antibody was obtained from Millipore (Molsheim, France) and anti-type II collagen monoclonal mouse antibody from Interchim (Montluçon, France). Ultravision detection system anti polyvalent HRP/DABkit and Mayer's hematoxylin were purchased from LabVision Microm (Francheville, France). 2.2 . Polymer synthesis and characterization 2.2.1. Synthesis The triblock copolymer PLGA–P188–PLGA (ABA copolymer) was prepared by ring-opening polymerization (ROP) of DL-lactide and glycolide using P188 as an initiator, and stannous octoate [Sn(Oct)2] as catalyst [35]. Briefly, precise amounts of P188, DL-lactide and glycolide were mixed and introduced into 100 mL round-bottom flasks with the catalyst. The mixture was heated to 140 °C and degassed by 15 vacuum-nitrogen purge cycles in order to remove the moisture and the oxygen, inhibitors of this polymerization. Flasks were then frozen at 0 °C and sealed under dynamic vacuum at 10 −3 mbar. Polymerization was allowed to proceed at 140 °C under constant agitation. After 5 days, the products were recovered by dissolution in dichloromethane and then precipitated by adding an equal volume of ethanol. Finally, the polymer was filtered, washed with cold ethanol and dried overnight at 45 °C under reduced pressure, up to constant weight. 2.2.2. Characterizations Number average molecular weight (Mn) and polydispersity index (PDI) of the copolymer were determined by size exclusion

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chromatography (SEC) using a Waters Inc. equipment fitted with a Plgel 5 μm mixed-C (60 cm) column as the stationary phase and a Waters 410 refractometric detector, eluted with dimethylformamide (DMF) at 1 mL·min − 1. Typically, samples were dissolved in DMF at 10 mg/mL and filtered on polytetrafluoroethylene (PTFE) filter Millex ®-FH (pore size 0.45 μm) prior to be injected. The number average molecular weight (Mn) and Weight Average Molecular Weight (Mw) were expressed according to calibration against poly(styrene) standards. 1 H NMR spectra were recorded at room temperature using an AMX300 Bruker spectrometer operating at 300 MHz. Deuterated DMSO was used as solvent, chemical shifts were expressed in ppm with respect to tetramethylsilane (TMS). 1H NMR: (300 MHz; d6-DMSO): δ (ppm) = 5.1 (m, 1H, CO–CH(CH3)–O), 4.8(m, 1H, CO– CH2–O), 3.6 (s, 4H, CH2–CH2–O), 1.5 (m, 3H, CO–CH(CH3)–O), 1.1 (d, 3H, CH(CH3)–CH2–O) (supplementary data). Differential scanning calorimetry (DSC) measurements were carried out under nitrogen on a Perkin Elmer Instrument DSC 6000 Thermal Analyzer. Samples were submitted to a first heating scan from 0 °C to 200 °C (10 °C/min) followed by a cooling (10 °C·min−1 from 200 °C to 100 °C, and 7 °C·min−1 from 100 °C to 0 °C) and a second heating scan to 200 °C (10 °C·min−1). Glass transition temperature (Tg), crystallization temperature (Tc), melting temperature (Tm) and melting enthalpy (ΔHm) were measured on the second heating ramp (supplementary data).

2.3. Nanoprecipitation and MS formulation PLGA MS were prepared using a solid/oil/water (s/o/w) emulsion solvent evaporation–extraction process previously described in [27,29,32]. Polymers were PLGA copolymer with a D,L-lactide: glycolide ratio of 75:25 and PLGA-P188-PLGA, which was synthesized by IBMM-CRBA CNRS UMR 5247 (Montpellier, France) as described in Section 2.2.1. The encapsulated protein was lysozyme from chicken egg white or TGF-β3. Protein loading was 1 μg of protein and 5 μg of human serum albumin (HSA)/mg of MS. The encapsulation of lysozyme was then performed as previously described [29]. TGF-β3 and HSA were nano-precipitated separately using a process previously described [27] but adapted to lyophilized TGF-β3. Briefly, 1.077 g of cold glycofurol (4 °C) was added to 10 μl of a TRIS–HCl 0.75 M, NaCl 2 M (pH = 7.4) solution containing 50 μg of TGF-β3 and various amounts of P188 depending on the protein-additive ratio (0.5 or 1 mg P188 for 1:10 or 1:20 ratio respectively). HSA nanoprecipitate with either a 1:10 or 1:20 protein:additive ratio was produced in a similar manner [19,23]. After 30 min at 4 °C, the nanoprecipitated proteins were recovered by centrifugation and dispersed in the organic phase (670 μL of 50 mg PLGA or PLGA–P188–PLGA dissolved in a 3:1 methylene chloride:acetone solution). The suspension was then emulsified in a poly(vinyl alcohol) aqueous solution (30 mL, 4% w/v at 1 °C) and mechanically stirred at 550 rpm for 1 min. After addition of 33 mL of deionized water and stirring for 10 min, the emulsion was added to 167 mL deionized water and stirred for 20 min to extract the organic solvent. Finally, the MS were filtered on a 0.45 μm High Volume Low Pressure (HVLP) type filter, washed and freeze-dried. MS without protein were prepared following the same process, and called blank-MS or blank-PAMs when covered with fibronectin. Encapsulation yield was determined after dissolution of MS in dimethyl sulfoxide (DMSO) during 1 h (5 mg, 3 batches). After this time, samples were centrifuged and residual DMSO was evaporated. The amounts of encapsulated bioactive lysozyme and TGF-β3 were evaluated with specific bioassays (described in Sections 2.9 and 2.11.2 respectively). Moreover, impact of DMSO on protein activity was tested using the bioassays described below, but no significant alteration of the proteins was induced by DMSO treatment (data not shown).

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2.4. Functionalization of MS surface: formulation of PAMs To obtain FN-PAMs, PLGA MS were coated with FN and poly-D-Lysine (PDL). Coating solutions were prepared in Dulbecco's Phosphate-Buffered Saline DPBS. The concentration of the coating molecules was 6 μg/mL of FN and 9 μg/mL of PDL (corresponding to a 60:40 ratio of FN:PDL). 5 mg of MS was suspended in DPBS and sonicated until full dispersion of the MS. The solution containing PDL and FN molecules was mixed to the MS suspension (final volume: 10 mL) and placed under rotation at 15 rpm at 37 °C during 1h30min. After coating, FN-PAMs were washed 3 times in sterile distilled water, lyophilized and kept at −20 °C. The microspheres were formulated in aseptic conditions and were thereafter incubated for 10 min with antibiotics after the coating step. Each tube was covered with sigmacote® to prevent product loss on the tube walls. To simplify FN-PAMs will thereafter be named simply PAMs but they always present the fibronectin covered surface. The fibronectin surface was characterized by confocal microscopy after FN immunostaining. Lyophilized PAMs (1 mg) were suspended in DPBS containing 4% bovine serum albumin (BSA), 0.2% Tween 20 (DPBSBT) and incubated for 30 min at room temperature (RT) under 15 rpm stirring. Samples were then washed three times with DPBS and centrifuged (9000 g, 5 min). Anti-FN mouse monoclonal antibody (100 μg/mL in DPBSBT) was incubated at 37 °C for 1.5 h under rotation. Samples were then washed 4 times before incubation with biotinylated anti-mouse IgG antibody (2.5 μg/mL in DPBS) for 1 h, at RT, under rotation. After three washes, samples were incubated with streptavidin–fluoroprobe 547 (1:500 in DPBS) at RT, for 40 min, under rotation. Samples were observed under confocal microscopy (Olympus FluoviewTM TU 300, Rungis, France). Three independent experiments were performed and every condition was observed in triplicate. 2.5. Physico-chemical characterization of MS and PAMs The average diameter and size distribution of the resulting MS were evaluated using a Multisizer® Coulter Counter (Beckman Coulter, Roissy, France). The zeta potential of MS and PAMs was determined using a Malvern Zetasizer® (Nano Series DTS 1060, Malvern Instruments S.A., Worcestershire, UK). The measure of zeta potential was achieved on MS suspension (0.3 mg/mL in NaCl 1 mM) after conversion of electrophoretic mobility values to ζ-potentials using Smoluchowski's equation. Results are presented as mean ± standard deviation, n = 3. Samples were finally washed three times with DPBS and observed under confocal microscopy (Olympus Fluoview® TU 300, Rungis, France). 2.6. hMSC culture hMSC cultures were established from bone marrow of patients undergoing hip replacement surgery, as previously described [38]. MSCs were isolated from patients after informed consent and approval by the Local Ethical Committee (registration number: DC-2009-1052). Briefly, cell suspensions were plated in a complete α-minimum essential medium (αMEM) supplemented with 10% fetal bovine serum (FBS), 2 mM glutamine, 100 U/mL penicillin, 100 mg/mL streptomycin and 1 ng/mL human basic fibroblast growth factor (bFGF). hMSCs were shown to be positive for CD44, CD73, CD90 and CD105 and negative for CD14, CD34 and CD45 and used between passages 3 and 4 [39,40]. 2.7. Adherence and viability of hMSCs after formation of MSC-PAM complexes hMSCs were washed with PBS, detached with 0.16% trypsin 0.02% EDTA solution (Lonza) and pelleted at 1400 rpm for 10 min. Cells

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were then suspended in culture medium supplemented with 3% FBS and plated in 1.9 cm 2 Costar ultra-low cluster plate with or without PAMs. When PAMs were added, 0.75 mg lyophilized PAMs were suspended in coated Eppendorf tubes containing α-MEM medium, 3% FBS for 15 min. PAM suspension was sonicated and briefly vortexed prior to add the cell suspension (0.25 × 10 6 cells/0.5 mg PAMs). Plates were incubated at 37 °C to allow cell attachment on PAM surface. For adhesion evaluation, PAMs/cell aggregates were recovered after 4 h, washed with α-MEM and pelleted at 200 g for 2 min. Cell adhesion was assessed using light microscopy and scanning electron microscopy. Samples were prepared for scanning electron microscopy analysis as previously described [18]. Briefly, PAMs were washed in PBS, fixed with glutaraldehyde 1% and osmium 1% and then dehydrated in a series of alcohol. Afterwards, they were soaked in hexamethyldisylasane, covered by a thick layer of carbon and finally observed. Cell viability was estimated at various time intervals (4 h, 24 h, 48 h and 7 days) and adherent live cells were quantified using the Cyquant cell proliferation assay® following the manufacturer's guidelines. 2.8. In vitro release of model protein lysozyme from MS In vitro release of lysozyme was determined by adding 250 μL of 0.05 M TRIS–HCl buffer, pH 7.4, containing 0.1% w/v BSA and 0.09% w/v NaCl to 5 mg of MS in capped tubes. Tubes were incubated in a shaking water bath (37 °C, 125 rpm). At determined time intervals, tubes were centrifuged for 5 min at 2800 g and supernatants (200 μL) were collected for analysis and replaced by fresh buffer. Percentage of biologically active released lysozyme was measured by enzymatic assay as described in Section 2.9. The ratio of cumulative release (in percent) was calculated based on the total amount of lysozyme obtained from the encapsulation yield. 2.9. Quantification of lysozyme The amount of active lysozyme was determined by measuring the turbidity change in a M. lysodeikticus bacterial cell suspension as previously reported [41]. Hundred microliters of a lysozyme solution was added to 2.9 mL of a 0.015% w/v M. lysodeikticus suspension in TRIS–HCl (0.01 M, pH 7.4) buffer solution. After incubation (37 °C, 4 h), the absorbance was measured at 450 nm. For the standard curve used for encapsulation ratio determination, 100 μL of lysozyme at 600 μg/mL (Tris 0.05 M, pH 7.4) were added to 10 mg blank microspheres dissolved in 0.9 mL DMSO. After 1 h, 3 mL of 0.01 M HCl was added. The solution was left to stand for one more hour. The solution was diluted and a volume corresponding to 5 to 50 ng lysozyme was withdrawn, completed to 100 μL with TRIS buffer and incubated with a M. lysodeikticus suspension for lysozyme activity determination. The standard curve used for lysozyme release was done with 100 μL of lysozyme solution containing 5 to 50 ng of lysozyme incubated in 2.9 mL 0.015% w/v Micrococcus lysodeikticus suspension in TRIS (0.05 M, pH 7.4) buffer solution (37 °C, 4 h). The absorbance of the suspension was then measured at 450 nm. During the release study, 100 μL of the diluted supernatant containing lysozyme was incubated with M. lysodeikticus. The amount of active protein was calculated using the standard curve. 2.10. In vitro release of TGF-β3 from MS In vitro release of TGF-β3 from MS was determined by adding 250 μL of phosphate buffer saline (PBS), pH 7.4, containing 1% w/v BSA to 5 mg of MS in capped tubes. Tubes were incubated in a shaking water bath (37 °C, 125 rpm). At determined time intervals, the tubes were centrifuged for 5 min at 2800 g and 200 μL of the

supernatant was collected for analysis and replaced by fresh buffer. Supernatants were stored at − 80 °C. 2.11. Quantification of TGF-β3 2.11.1. ELISA The quantity of TGF-β3 released was measured using a specific ELISA kit. The value of cumulative release was expressed as a percentage of the total amount of TGF-β3 initially encapsulated. 2.11.2. Bioassay To assess the biological activity of TGF-β3 released from MS, we performed a bioassay previously developed by Tesseur et al. [42]. The bioassay relies on the use of mouse fibroblasts isolated from TGF-β1 −/− mice (MFB-F11) stably transfected with a reporter plasmid consisting of TGF-β responsive Smad-binding elements coupled to the secreted alkaline phosphatase (SEAP) reporter gene. MFB-F11 fibroblasts were seeded at 3 × 10 4 cells/well in 96-well flat-bottom culture plates. After overnight incubation, cells were washed twice with PBS and incubated in 50 μL serum-free DMEM supplemented with 100 U/mL penicillin and 100 mg/mL streptomycin. After 2 h, 50 μL of TGF-β3 containing samples was added for another 24 h. Serial dilutions of TGF-β3 were added to other wells to determine the standard curve. SEAP activity was measured on 50 μL of culture supernatants using the SEAP Reporter Gene Assay, chemiluminescent kit according to the manufacturer's instructions. Chemiluminescence was detected using a micro plate luminometer and results were analyzed with the Ascent software for Fluoroscan (Ascent FL, Thermo Fisher Scientific, Cergy-Pontoise, France). This bioassay can reliably measure as little as 1 pg/mL TGF-β and its linear range extends well beyond 1000 pg/mL (for further details, see [42]). The ratio of cumulative release (in percent) was calculated based on the total amount of encapsulated TGF-β3. 2.12. Chondrogenic differentiation The chondrogenic differentiation was performed with the MS formulated with PLGA–P188–PLGA polymer matrix with a protein: additive ratio of 1:20 and compared to the previously used formulation [23] of PLGA matrix with a protein:additive 1:10 ratio. hMSCs were incubated with PAMs (PLGA PAMs/TGF-β3 or PLGA– P188–PLGA PAMs/TGF-β3 or blank PAMs) in conical tubes. Briefly, hMSCs (2.5 × 10 5 cells) and 0.5 mg PAMs were pelleted by centrifugation in 15 mL tubes, and cultured in chondrogenic medium. This medium consisted in DMEM supplemented with 0.1 mM dexamethasone, 0.17 mM ascorbic acid, 35 mM proline and 1% insulin-transferrin-selenic acid (ITS) supplement. Standard chondrogenesis of hMSCs was induced by culture in micropellets in chondrogenic medium supplemented with 10 ng/mL TGF-β3 as previously reported [43], and this condition was referred as the control for the RT-qPCR experiments (normalized to a value of 1). 2.13. Real-time RT-PCR After 21 days of culture in micropellets, hMSC/PAM complexes were washed in PBS and mechanically dissociated in lysis buffer. Total RNA from cell preparations was extracted, according to the recommendations of the manufacturer. Cells were lysed in 1% β-mercaptoethanol containing buffer and RNA extracted following treatment by DNAse to remove any traces of genomic DNA. First strand cDNA synthesis was performed with SuperScript™ II Reverse Transcriptase. Following first strand cDNA synthesis, cDNAs were purified, eluted in 50 μL RNAse free water. cDNAs (3.125 ng) were mixed with iQ SYBR Green Supermix and primer mix (0.2 mM) in a final volume of 10 μL. Amplification was carried on a Chromo4 thermocycler (Biorad) with a first denaturation step at 95 °C for

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Table 1 List of primers used for RT-qPCR. Gene

Full name

Accession number

Sequences

Col2a1 variant B

Type II procollagen variant B

NM_033150

ACAN

Aggrecan

HAPLN1

Link protein

NM_013227 NM_001135 NM_001884

Col10a1

Type X collagen

NM_000493

Alpl

Alkaline phosphatase

NM_000478

BGLAP

bone gamma-carboxyglutamate protein or osteocalcin

NM_199173

FabP4

Fatty acid binding protein 4

NM_001442

Lpl

Lipoprotein lipase

NM_000237

GAPDH

Glyceraldehyde-3-phosphate deshydrogenase

NM_002046

B2M

Beta-2-microglobulin

NM_004048

ACTB

Beta-actin

NM_001101

HSPCB

Heat shock protein S18

NM_007355

Forward 5′-gaggggatcgtggtgacaaagg-3′ Reverse 5′-ttgcattactcccaactgggcg-3′ Forward 5′-cgtgcatgccaaccagacg-3′ Reverse 5′-tggcagctccatgtcaggc-3′ Forward 5′-ttccacaagcacaaactttacacat-3′ Reverse 5′-gtgaaactgagttttgtataacctctcagt-3′ Forward 5′-tgctgccacaaatacccttt-3′ Reverse 5′-gtggaccaggagtaccttgc-3′ Forward 5′-ccacgtcttcacatttggtg-3′ Reverse 5′-gcagtgaagggcttcttgtc-3′ Forward 5′-ggcgctacctgtatcaatgg-3′ Reverse 5′-tcagccaactcgtcacagtc-3′ Forward 5′-atgggatggaaaatcaacca-3′ Reverse 5′-gtggaagtgacgcctttcat-3′ Forward 5′-gtccgtggctacctgtcatt-3′ Reverse 5′-tggatcgaggccagtaattc-3′ Forward 5′-agatccctccaaaatcaagt-3′ Reverse 5′-atgatcttgaggctgttgtc-3′ Forward 5′-cgcgctactctctctttct Reverse 5′-cagtaagtcaacttcaatgtcg Forward 5′-ccagatcatgtttgagacct Reverse 5′-ggcatacccctcgtagat Forward 5′-taccaaagtgatcctccatc Reverse 5′-ctcatctgaacccacatctt

3 min and 40 cycles of 95 °C for 10 s, 55 °C for 15 s and 72 °C for 15 s. After amplification, a melting curve of the products determined the specificity of the primers for the targeted genes. A mean cycle threshold value (Ct) was obtained from 2 measurements for each cDNA. Several housekeeping genes, glyceraldehyde-3-phosphate dehydrogenase (Gapdh, NM_002046), beta 2 Microglobulin precursor (B2M, NM_004048), beta actin (Actb, NM_001101), and heat shock 90 kD protein 1 beta (Hspcb, NM_007355) were tested for normalization (for primer sequences, see Table 1). Using GeNorm® freeware (http:// medgen.ugent.be/ejvdesomp/genorm/), Gapdh, B2M and Actb were determined to be the three most stable housekeeping genes. The relative transcript quantity (Q) was determined by the delta Ct method [Q = E (Ct min in all the samples tested-Ct of the sample)] where E is related to the primer efficiency (E = 2 if the primer efficiency = 100%). Relative quantities (Q) were normalized using the multiple normalization method described in Vandesompele et al. [44]. The reference technique for chondrogenesis of hMSC (cf. Section 2.12: micropellet in chondrogenic medium supplemented with 10 ng/mL TGF-β3) was defined as the control for the RT-qPCR experiments (normalized to a value of 1), and the other conditions were expressed in fold increase compared to this value. D0 indicates the basal expression in hMSC cultured in expansion condition (described in Section 2.6).

2.14. Histology and immunohistochemistry Samples were fixed in 4% paraformaldehyde for 24 h, washed in PBS and processed for routine histology. Paraffin-embedded samples sections (5 μm) were rehydrated through a gradient of ethanol and xylene and stained with hematoxylin–eosin. Immunohistochemistry was performed on sections using the Ultravision detection system anti-polyvalent HRP/DAB kit, according to the manufacturer's instructions. For type II collagen and Human Leukocyte Antigen immunostaining, samples were first incubated at 37 °C for 1 h with hyaluronidase 0.1% for epitope retrieval. Primary antibodies, anti-aggrecan polyclonal rabbit antibody (1:50;) or anti-type II collagen monoclonal mouse antibody (1:50;) were incubated for 1 h at room temperature. Samples were finally counterstained with Mayer's hematoxylin (LabVision) for 3 min and mounted

with Eukitt. Immunopositive extracellular matrix showed a brown staining. 2.15. Statistical analysis XLSTAT 2008 (Addinsoft Paris, France) was used for that purpose. Statistical significance for each experiment was determined by a Dunnett's test or a Student's t-test. The tests were considered as significant with p values of less than 0.05 3. Results 3.1. Polymer characterization With the aim of providing a new degradable copolymer with original drug release properties, we decided to copolymerize resorbable and biocompatible materials, precisely PLGA segments with poloxamer 188. The PLGA–P188–PLGA copolymer was characterized by proton nuclear magnetic resonance ( 1H NMR), size exclusion chromatography (SEC) and differential scanning calorimetry (DSC) (Supplementary data 1). The molecular weight of the PLGA block was determined by using the integration ratio of resonance of PEG units at 3.6 ppm, PGA at 4.8 ppm and PLA at 5.1 ppm in the 1H NMR spectra (Mn = 83,400). SEC analysis gave a different molecular weight value (Mn = 65,400). This slight difference is explained by the use of polystyrene standards for SEC calibration causing inaccurate measurements with aliphatic polyesters [45]. Moreover, the amphiphilic nature of the copolymer can modify its hydrodynamic volume and increase the retention time, giving underestimated results. Molecular weight calculated from 1H NMR spectra is therefore more accurate, because no PLGA copolymer (without P188) was formed during the Ring Opening Polymerisation (ROP) as confirmed by the monomodal chromatograms obtained in SEC analyses. As a consequence, values deduced from 1H NMR integrations can be considered as sound. DSC analyses of the copolymer showed a Tm of 165.4 °C due to the crystallinity of PLA sequences in PLGA blocks, a crystallization temperature at 105 °C, and no Tg was observed.

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3.2. MS/PAM characterization The morphological aspect of MS, determined by scanning electron microscopy (SEM), showed a spherical shape for both formulations but surface differences were evidenced, as PLGA–P188–PLGA MS appeared rougher than PLGA MS (Fig. 1A and B). The mean particle size observed for MS was similar for the two formulations: 58.4 ± 20.0 μm and 64.5 ± 18.8 μm for PLGA and PLGA–P188– PLGA, respectively (Fig. 1F). A combination of FN with the highly charged PDL molecules was added to MS to favor cell attachment, thus forming PAMs. As confirmed by confocal microscopy, the FN surface was homogeneous (Fig. 1C, D). Furthermore, both PLGA and PLGA–P188–PLGA PAMs showed positive zeta potential values with the FN surface that evolved from − 8.1 ± 2.3 mV for the two types of MS to + 7.9 ± 0.8 mV for the PLGA–P188–PLGA PAMs and + 15.4 ± 0.3 mV for the PLGA PAMs (Fig. 1E). These results were satisfactory for cell adhesion since a positively charged surface promotes adhesion of the cells.

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3.3. Characterization of hMSC/PAM complexes and cell proliferation Optical microscopy revealed an efficient attachment of hMSCs on PAMs, independently of the type of formulation (Fig. 2). Interestingly, 4 h after attachment, almost all the cells attached to the PAM's FN-covered surface and formed 3D complexes, with no free-floating cells. 24 h, 48 h, and even 7 days after cell attachment, hMSC/PAM complexes formed aggregates and no dead cells were observed throughout the follow-up (Fig. 3A). After 7 days of culture in the absence of PAM, hMSCs did not proliferate. On the contrary, cells cultured with either of the two types of PAMs proliferated, as shown by increased cell number (Fig. 3B). Importantly, the number of cells at 7 days was significantly higher when hMSCs were cultured with PLGA–P188–PLGA PAMs than with PLGA PAMs (Fig. 3B). 3.4. Encapsulation and in vitro release of lysozyme from MS and study of pH acidity We evaluated the impact of (i) the polymer matrix (PLGA or PLGA–P188–PLGA) and (ii) the amount of P188 additive associated in the nanoprecipitation step with lysozyme on the release profile of this protein from MS. The polymer matrix was first modified with a fixed protein-P188 ratio of 1:10 for the nanoprecipitation step (PLGA + P188 1:10 vs PLGA–P188–PLGA MS + P188 1:10, Fig. 4A). In these conditions, when PLGA–P188–PLGA was used as a matrix,

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the release of lysozyme from MS was significantly increased. At day 30, 51% of the encapsulated lysozyme was released compared to 17% for PLGA-based MS. At day 3, 12% of lysozyme was released from PLGA-MS and low subsequent release was observed after this initial burst (from 12% to 17% at day 30). By contrast, when using PLGA–P188–PLGA as a matrix, a more sustained release was observed from day 0 to day 15 (from 26% to 48%). The different ratios of poloxamer 188 as additive were then tested with PLGA–P188–PLGA polymer matrix. At 1:20 ratio, release of biologically active protein was enhanced. As observed in Fig. 4, lysozyme release was improved from 51% at day 30 with P188 1:10 (as previously mentioned) to 71% when the protein:P188 ratio was 1:20. In addition, as the acidic polymer degradation products are known to change the pH, the pH in the media recovered from PLGA MS or PLGA–P188–PLGA MS was measured. A pH decrease was initiated earlier for PLGA matrix than PLGA–P188–PLGA matrix to reach a pH value nearing 5 (limit of pH stability for lysozyme) at day 15, which slowly continued to decrease after this point (Fig. 4B).

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Fig. 2. Adherence of hMSCs on PAM formulated with (A) PLGA or (B, C) PLGA-P188-PLGA. hMSCs (0.25 × 106) were seeded with 0.5 mg PAM on ultra low cluster plates. A and B show images from optic microscopy, whereas C is an image from SEM of hMSC and PLGA–P188–PLGA PAMs co-culture.

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Fig. 1. Characterization of MS and PAMs. SEM imaging of (A) PLGA or (B) PLGA–P188– PLGA MS. (C) Differential interference contrast (DIC) microscopy image and (D) superposition of DIC and confocal microscopy images (anti-FN immunofluorescence) of PLGA–P188–PLGA based PAMs. Scale bars: 10 μm. (E), Surface charge of MS and PAMs as a function of the PLGA scaffold used (PLGA or PLGA–P188–PLGA). (F), Average diameter and size distribution of PLGA–P188–PLGA MS.

We evaluated the total amount of released TGF-β3 by ELISA and the bioactive amount of TGF-β3 using a specific bioassay (Fig. 5). The percent activity of TGF-β3 after nanoprecipitation was 101% ± 9.2 as detected by bioassay. To detect any loss in biological activity during the encapsulation process, the amount of active protein extracted from MS was determined after dissolution of MS in DMSO. The encapsulation yield of TGF-β3 in PLGA MS was 116 ± 22% and 89 ± 12% in PLGA-P188-PLGA MS. We first observed that the release profiles of bioactive TGF-β3 were similar for all formulations with a more pronounced initial release between day 0 and day 7 followed by a phase of slow release till day 30 (Fig. 5A, B, C). Modification of the polymer matrix with a fixed protein:P188 ratio of 1:10 for the nanoprecipitation step induced a low release of TGF-β3 from PLGA MS, corresponding to 25% of encapsulated TGF-β3 at day 30 (Fig. 5A). Use of PLGA–P188– PLGA MS at a protein:P188 ratio of 1:10 allowed an increased release

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of TGF-β3 by about 40% from day 0 to day 3 followed by a slow release from day 3 to day 30 (10%). The different ratios of protein:P188 additives were then tested with PLGA–P188–PLGA chosen as a polymer matrix (Fig. 5B). The bioactive TGF-β3 represented 56.2 ± 8.9% of the total amount of encapsulated TGF-β3 for PLGA MS (compared to 100% as measured by ELISA), whereas it was 96.7 ± 12.5% for PLGA–P188–PLGA + P188 1:20 MS, suggesting a more efficient protein protection during the encapsulation step with the last formulation. With the same matrix polymer (PLGA–P188–PLGA), and a protein: P188 ratio of 1:20, a total of 74% of bioactive TGF-β3 was released by day 30 and 60% being already released at day 6. Considering MS and PAMs, TGF-β3 release from PAMs was more sustained during the first 7 days compared to MS (32% release at day 3 for PAMs vs. 55% released from MS), suggesting that the FN-covered surface favored the retention of the protein inside the MS, at least during the initial phase of the study (Fig. 5C).

strongly suggested a higher rate of chondrogenic differentiation by continuous release of TGF-β3 using the new formulation of PLGA– P188–PLGA PAMs. Expression of collagen II and aggrecan proteins was then evaluated using a specific immunohistochemical staining (Fig. 7). PAM hollows can be seen in all preparations. Specific staining for type II collagen showed an intense and uniformly distributed brown staining when hMSCs were associated to PLGA–P188–PLGA PAMs (Fig. 7B). The staining intensity was weaker with PLGA PAMs and no specific type II collagen expression was seen for unloaded PLGA–P188–PLGA. By contrast, aggrecan staining can be observed when cells are associated with unloaded PLGA–P188–PLGA PAMs (in agreement with the expression of this protein at a basal level in hMSC) (Fig. 7F). Nevertheless, this staining was more intense when cells were cultured with PAMs releasing TGF-β3 (Fig. 7D, E). 4. Discussion

3.6. hMSC differentiation associated to PAMs +/− TGF-β3 hMSCs were allowed to adhere on PAMs (based on PLGA or PLGA– P188–PLGA matrix) releasing TGF-β3 or unloaded PAMs (without encapsulated growth factor) and cultured for 21 days. They were compared to hMSCs cultured in micropellets with TGF-β3 added every three days, the standard for in vitro chondrogenic differentiation. Chondrogenesis was evaluated by determining expression of specific markers as listed in Table 1. Higher expression of all tested chondrogenic markers (collagen type IIB variant, link protein, aggrecan, COMP) was observed when hMSCs were cultured with TGF-β3-releasing PAMs compared to micropellet cultures (Fig. 6A.). Moreover, significant up-regulation of collagen IIB and COMP was detected when hMSCs were cultured with PLGA–P188–PLGA PAMs (from 35 to 59 fold increased expression for collagen IIB for TGF-β3-releasing PLGA PAMs versus TGF-β3-releasing PLGA–P188– PLGA PAMs, respectively). Absence of differentiation was noticed when cells were cultured with unloaded PLGA-P188-PLGA or PLGA PAMs. Importantly, the osteogenic markers alkaline phosphatase and osteocalcin were expressed at very low levels whatever were the inductive conditions as compared to non-induced cells (D0) (Fig. 6B.). Interestingly, adipocytic markers were down-regulated, except for unloaded PLGA–P188–PLGA PAMs (Fig. 6C.). These results

In the absence of efficient and reproducible therapeutic options for cartilage repair, a number of scaffolds have been developed for tissue engineering strategies (Carticel®, Chondrotransplant® and Chondrosphere®, Bioseed®-C, ChondroCelect®). These products, which primarily rely on the implantation of autologous chondrocytes are being used or tested in clinics and hospitals. However, they have some limitations and have been criticized by healthcare professionals because of (i) the low quality of resulting cartilage (fibrous cartilage), (ii) its high cost (iii) the need for surgery and (iv) low reproducibility between patients. There is still a strong demand for an easy-to-handle and injectable system that could avoid surgery and fill the defects by forming a cartilaginous matrix with stable properties on the long term. The use of injectable PAMs, which combine in a unique and simple device the 3D FN-covered surface support for cell implantation and survival with the continuous supply of TGF-β3 for hMSC differentiation toward chondrocytes, may be useful to get a hyaline cartilage tissue with devoted functions. Moreover, this approach allows rapid adhesion of cells on their surface without the need of laborious cell culture. For therapeutic applications, association of PAM/cell complexes within an injectable hydrogel may represent an alternative to the current strategies. Within this line, preliminary results show a similar viability of hMSC/PAM complexes with or without a thermo-sensitive

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approved [20]. However, efficient and sustained release of proteins from PLGA MS remains a challenge mainly due to protein instability (adsorption, aggregation, and denaturation) during the formulation process or the release period [24,25]. To avoid these drawbacks, we

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Time (days) Fig. 4. Release study of the protein, illustrating the differences between PLGA versus PLGA-P188-PLGA polymers as a matrix. (A) Cumulated release of bioactive lysozyme. (B) pH measurement of the release buffer. PLGA + P188 1:10, PLGA–P188– PLGA + P188 1:10, PLGA–P188–PLGA + P188 1:20 were incubated in TRIS-HCL. The amount of lysozyme released was measured at different times in the medium (quantification limit of the bioassay 5 ng). In parallel the pH of this release medium was also tested. Each error bars represent the ±standard deviation of average percent cumulated values with n = 3 for each formulation.

Fig. 5. Cumulative release of total and bioactive TGF-β3 from microcarriers. (A) Influence of the matrix composition, (B) TGF-β3:P188 ratio, and (C) of the fibronectin surface surface on the protein release. PLGA + P188 1:10, PLGA–P188–PLGA + P188 1:10 and PLGA–P188–PLGA + P188 1:20 were incubated in PBS 1% BSA, and the TGF-β3 released at different time in the medium was tested by ELISA or bioassay. Each error bars represent the ± standard deviation of average percent cumulated values with n = 3 for each formulation.

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hydrogel (unpublished results). In the present study we therefore developed an efficient cell- and drug-delivery system that could allow in situ the chondrogenesis of hMSCs and the production of a high quality cartilaginous matrix. We have previously formulated PLGA-based PAMs releasing TGF-β3 which provided an appropriate scaffold for chondrogenic differentiation of hMSCs [23]. As PLGA is a biodegradable and biocompatible polymer, the use of PLGA based implantable devices is FDA

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Fig. 6. Chondrogenic differentiation of hMSCs cultured in vitro with PLGA PAMs or PLGA-P188-PLGA PAMs. Expression of (A) chondrocytic markers, (B) osteoblastic markers and (C) adipocytic markers normalized to housekeeping genes and then expressed in fold change compared to the standard condition (TGF-β3 in medium) which was normalized to 1. Dunnett's test, * indicates significant difference (p b 0.05), ** (p b 0.01), *** (p b 0.001), n = 3.

developed novel TGF-β3 PAMS using the polymer, P188 poloxamer under two forms: as an additive nanoprecipitated with the protein or covalently linked to PLGA macromolecules to form a triblock copolymer. The latter form should retain the P188 poloxamer within the matrix during its degradation, due to the use of a triblockin-triblock new polymer. To our knowledge, although poloxamers can be blended with protein or PLGA to enhance protein stability [29,46–48], the use of covalently linked PLGA-poloxamer for protein incorporation in PLGA-based scaffolds for tissue engineering has not been reported so far.

These novel PLGA–P188–PLGA PAMs, in the absence of any differentiation factor, provided a 3D FN covered surface with superior biocompatibility for hMSCs than previous PLGA PAMs, as shown by increased cell proliferation. This could be explained by the well documented cell survival/protecting effect of P188 [49]. These surfactants were shown to prevent the leakage of intracellular components [50–52], due to their amphiphilic nature, and especially to their hydrophobic PPO moiety, which allow their penetration into the membrane [53]. Additionally, some studies reported that stiffness, roughness, chemical nature and various other physical factors can

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Fig. 7. Expression of cartilage proteins by hMSC cultured on PAMs. hMSCs were cultured in micropellets with (A) unloaded PLGA–P188–PLGA PAMs, (B) TGF-β3 PLGA–P188–PLGA PAMs and (C) TGF–β3 PLGA PAMs. Formed complexes were embedded in paraffin after day 21 and sections were used for immunostaining against collagen II or aggrecan. Note the empty spaces signaling the PAMs hollows. Scale bars represent 50 μm.

promote better cell survival, proliferation and differentiation [12,54,55], suggesting that the difference in surface topography of PLGA–P188– PLGA PAMs may also explain higher cell survival/proliferation. Finally, there may also be more FN, which is involved in hMSC survival [56], covering the PAMs made with the triblock copolymer, but no major differences were observed by confocal microscopy. In order to study the mechanisms of protein release by the new polymer, we first evaluated the kinetic release of a model protein, lysozyme, previously used in our laboratory [29,32,41]. Although it cannot reflect the behavior of a specific therapeutic protein it may provide helpful insights on protein release at a lower cost. We used a low loading rate of active lysozyme (0.1%) as compared to previous studies (0.6%) to be close to physiological protein levels [29,32,41]. In parallel to the quantification of protein released in the release medium, we assessed the pH of the medium to correlate with MS degradation as done in Tran et al. to study PLGA–PEG–PLGA MS degradation [32]. Our results confirmed differences in the release profile according to the hydrophilic nature of the MS matrix, both for the model and therapeutic proteins. As an example, PLGA, which is hydrophobic, induced a higher TGF-β3 retention during the first day (4%) and low subsequent release to reach a total release of 18%. With the more hydrophilic PLGA–P188–PLGA matrix, the total protein release was 49% at the same P188 additive ratio (1:10), with a release of 28% at day 1. It is generally accepted that at least two phases can be distinguished in the release profile of a protein from PLGA devices. In the initial phase, drug release is governed by diffusion of the protein through an interconnecting network of aqueous pores, while at later stages erosion of the polymeric matrix plays a more prominent role [26]. Indeed, the increased and sustained protein release from the more hydrophilic PLGA–P188–PLGA MS may be explained by a better diffusion of water and proteins [33,57]. During the polymer degradation step, the pH drop in the release medium reflects the diffusion of degradation products. In PLGA–P188–PLGA matrix the exchange was facilitated therefore delaying protein degradation within the core of the MS. The remaining acidification could explain the loss of protein still observed (around 30%). When P188, which surrounds and protects the protein, was co-precipitated with TGF-β3 at the ratio of 1:20 of protein:additive polymer, a more complete and sustained release was achieved (over 70% in 30 days for PAMs). In this regard, it is important to note that the mean value of bioactive versus total

TGF-β3 ratio was around 56.2 ± 8.9% for the PLGA PAMs, whereas it was 96.7 ± 12.5% for PLGA–P188–PLGA + P188 1:20 PAMs. This new formulation therefore allows an almost complete bioactive TGF-β3 release from the PAMs probably due to a better diffusion of the protein and an enhanced protection within the core. The more efficient release of TGF-β3 by the novel PAMs should theoretically induce a higher rate of hMSC differentiation into chondrocytes. This was indeed the case. Compared to PLGA PAMs, PLGA–P188–PLGA PAMs induced significantly higher levels of expression of transcripts encoding for the secreted molecules specific for cartilage matrix such as collagen II, aggrecan, link and collagen X. The latter, a marker for hypertrophic chondrocytes, is known to be induced by TGF-β3; its expression suggests that terminal differentiation toward hypertrophy may occur with PAMs. Furthermore, aggregation of hMSCs at the PAM surface is likely to be important since it recapitulates the initial condensation phase observed during embryonic chondrogenesis. This is further suggested by the absence of expression of osteogenic or adipocytic markers confirming a chondrocyte specific differentiation profile. Moreover, contrary to the addition of TGF-β3 in the culture medium, which irrigates principally cells at the scaffold surface [58], the TGF-β3 release by PAMs allowed a better diffusion inside the aggregates that may explain the efficient chondrogenesis observed. Interestingly, the sustained release of TGF-β3 from PLGA-P188PLGA PAMs seemed to match the biological requirement for the program of chondrogenic differentiation [59,60]. Namely, an initial stimulation with high TGF-β3 amounts during the first week to induce the stem cell commitment was followed by a slow release necessary to maintain the differentiated phenotype. Continuous exposure of TGF-β3 for the first week is essential to induce chondrogenic differentiation of MSCs [59,60]. However, in vitro chondrogenic commitment of MSCs was shown to result in unstable phenotype with extensive mineralisation after in vivo implantation of MSCs [61]. Transient exposure to TGF-β3 may not be sufficient to promote differentiation and its absence could create a permissive environment to promote osteogenesis in vivo [62]. In this study, during the initial release phase, around 32 ng TGF-β3/mL/day was released from 0.5 mg PLGA–P188–PLGA PAMs and after day 7, PAMs released 2–3 ng/mL/day. In this regard, Bian et al. recently reported that a mean release of 3.5 ng TGF-β3/day from alginate MS associated to a

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hydrogel induced an efficient in vivo chondrogenic differentiation of MSCs after subcutaneous injection [59]. However, these systems are not injectable and implantation requires surgical methods. With the novel cell- and drug-delivery PAMs developed in this study, sustained release of bioactive TGF-β3 (representing a global mean of 11.6 ng/0.5 mg PAMs/day during 28 days) was associated with increased cartilage production in vitro. Even though release was less sustained after the first week, the continuous delivery of TGF-β3 by PAMs could help stabilize the chondrocyte phenotype on the long run and therefore represents a powerful approach not only to induce in situ the differentiation of MSCs but also to maintain the cartilage phenotype. Finally, besides its interest for enhanced protein protection, the intrinsic properties of the P188 poloxamer may positively affect chondrogenesis. P188 poloxamer was reported to exert a protective activity against necrosis of chondrocytes [63,64], reduce DNA fragmentation of bovine chondral explants exposed to injurious unconfined pressure [65], and stimulate matrix synthesis of OA-like chondrocytes [66]. The presence of P188 both as a free additive for protein complexation and as a component of the scaffold provides an additional value to the PAMs, since it is likely to contribute to stabilization or protection of newly formed chondrocytes, but also protect cartilage from degeneration or degradation in vivo after intra-articular injection. 5. Conclusion This study reports the formulation of novel polymeric PAMs composed of a hydrophilic matrix of PLGA–P188–PLGA associated with a 1:20 ratio of protein:P188, presenting a FN-covered surface and delivering TGFB3 over time. These novel PAMs allowed efficient and sustained protein release, improvement of hMSC survival and chondrogenic differentiation of hMSCs in vitro as well as formation of “hyaline-like cartilaginous tissue”. They hold great promises for cartilage tissue engineering and their efficiency for in vivo therapeutic strategies is currently being investigated. Supplementary data to this article can be found online at http:// dx.doi.org/10.1016/j.jconrel.2013.04.017. Acknowledgments This work was supported by INSERM Transfert. We want to gratefully acknowledge Emilie Gué for her participation on model protein encapsulation and release study, as well as Cedric Paniagua for polymer synthesis. Finally we are grateful to Dr Tony Wyss-Coray for the MFB-F11 cell line. We also thank the SCIAM ("Service Commun d'Imagerie et d'Analyse Microscopique") of Angers for confocal microscopy and scanning electron microscopy images. References [1] R.L. Mauck, S.B. Nicoll, S.L. Seyhan, G.A. Ateshian, C.T. Hung, Synergistic action of growth factors and dynamic loading for articular cartilage tissue engineering, Tissue Eng. 9 (2003) 597–611. [2] A.K. Williamson, A.C. Chen, R.L. Sah, Compressive properties and function–composition relationships of developing bovine articular cartilage, J. Orthop. Res. 19 (2001) 1113–1121. [3] J.R. Steadman, K.K. Briggs, J.J. Rodrigo, M.S. Kocher, T.J. Gill, W.G. Rodkey, Outcomes of microfracture for traumatic chondral defects of the knee: average 11-year follow-up, Arthroscopy 19 (2003). [4] T. Minas, A.H. Gomoll, R. Rosenberger, R.O. Royce, T. Bryant, Increased failure rate of autologous chondrocyte implantation after previous treatment with marrow stimulation techniques, Am. J. Sports Med. 37 (2009). [5] C.H. Chang, T.F. Kuo, C.C. Lin, C.H. Chou, K.H. Chen, F.H. Lin, H.C. Liu, Tissue engineering-based cartilage repair with allogenous chondrocytes and gelatin-chondroitin-hyaluronan tri-copolymer scaffold: a porcine model assessed at 18, 24, and 36 weeks, Biomaterials 27 (2006) 1876–1888. [6] C.K. Kuo, W.J. Li, R.L. Mauck, R.S. Tuan, Cartilage tissue engineering: its potential and uses, Curr. Opin. Rheumatol. 18 (2006) 64–73. [7] E.B. Hunziker, Articular cartilage repair: basic science and clinical progress. A review of the current status and prospects, Osteoarthr. Cartil. 10 (2002) 432–463.

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