Encapsulation of Osteoblast Seeded Microcarriers ... - ACS Publications

Biomacromolecules 2005, 6, 1608-1614

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Encapsulation of Osteoblast Seeded Microcarriers into Injectable, Photopolymerizable Three-Dimensional Scaffolds Based on D,L-Lactide and E-Caprolactone Heidi A. Declercq,† Tomasz L. Gorski,‡ Se´ verine P. Tielens,† Etienne H. Schacht,‡ and Maria J. Cornelissen*,† Department of Anatomy, Embryology, Histology and Medical Physics, Louis Pasteurlaan 2, Ghent University, B-9000 Ghent, Belgium, and Polymer Material Research Group, Krijgslaan 281 (S4), Ghent University, B-9000 Ghent, Belgium Received January 17, 2005; Revised Manuscript Received March 2, 2005

UMR-106 seeded microcarriers were encapsulated into in situ, photopolymerizable three-dimensional scaffolds based on D,L-lactide and -caprolactone. UMR-106 and rat bone marrow cells proliferated and differentiated well on the microcarriers. The microcarriers were completely colonized after 14 days in culture. The viscous polymer paste allowed to mix the UMR-106 seeded microcarriers and gelatin (porosigen) properly. After the photopolymerization process, microcarriers and gelatin were evenly distributed throughout the scaffold. Gelatin was leached out within 7 h, and a porous scaffold was obtained. The microcarriers remained in the scaffold even after 7 days which demonstrates that they were well entrapped in the polymer. Increasing the amount of entrapped microcarriers (20-50%) leads to scaffolds with a reduced cross-linking. Hence, the microcarriers leached out. The encapsulated UMR-106 cells did not show pyknotic nuclei which demonstrates that the photopolymerization and handling the viscous polymer/gelatin/microcarrier paste is not detrimental for the cells. 1. Introduction An estimated one million people per year require the treatment of critical size bone defects, defects that do not heal spontaneously. Up till now, current treatments include auto- and allografts, but these procedures have many limitations. Tissue engineering approaches offer new possibilities for functional and structural restoration of the damaged or lost tissue. Tissue engineering of bone requires a suitable osteoconductive and -inductive matrix1-3 and dependent on the severity of the defect, additional sources of osteogenic cells. Traditionally, the osteoconductive matrix is a prefabricated three-dimensional scaffold composed of degradable polymers and/or ceramics, but an injectable, in situ forming scaffold may provide many advantages.4 The moldable paste can fit exactly in irregular bone defects. Biocompatible in situ photopolymerizable polymers for tissue engineering applications have the potential to reduce the invasiveness and cost of many surgical procedures. For instance, instead of making large incisions to implant cell and polymer constructs, physicians could potentially reduce costs and surgical trauma by injecting efficiently crosslinking polymers mixed with osteoinductive agents (calcium phosphate cements, bone morphogenetic proteins, and precusor cells) in situ in tissue defects by minimally invasive * Corresponding Author. Tel.: 32/9/264 9241. Fax: 32/9/264 9498. E-mail: [email protected]. † Department of Anatomy, Embryology, Histology and Medical Physics. ‡ Polymer Material Research Group.

procedures.5,6 To permit cell infiltration, porosigens such as gelatin can be added to achieve an interconnected porous scaffold. Although a variety of cell sources may be used, an autologous patient biopsy (bone marrow) can be preferable as this decreases the risk of immunogenic response and disease transfer. To reduce the size of the required biopsy, prior to encapsulating in 3-D scaffolds, the cells must be expanded to obtain a sufficient cell number. To mix cells with the viscous polymer paste and inject into the site of injury, cells should be protected from handling the paste. Cells are often encapsulated in hydrogels, e.g., agarose, alginate, fibrin, gelatin, or photopolymerizable hydrogels, e.g., PEO-dimethacrylate.7-10 Highly swollen hydrogels are capable of suspending cells three-dimensionally and supporting nutrient diffusion to encapsulated cells but may not provide an ideal environment for anchoragedependent osteoblasts (osteoprogenitor cells).4 Second, the mechanical stability of the hydrogels could be too weak to protect the encapsulated cells from handling the viscous polymer/hydrogel mix. For epithelial-like cells or cells showing a polarity (e.g., osteoblast-like cells), macroporous microcarriers could be promising. The use of macroporous gelatin-based microcarriers in tissue engineering strategies offers several advantages: (1) cells can be expanded in vitro, (2) biodegradable, (3) cells are in a “natural” 3-D environment, and (4) cells within the interior of the microcarrier are in a protected environment (protected from handling).

10.1021/bm050031s CCC: $30.25 © 2005 American Chemical Society Published on Web 04/08/2005

Encapsulation of Osteoblast Seeded Microcarriers

In the present paper, UMR-106 and rat bone marrow cells were seeded on macroporous Cultispher S microcarriers. The colonization, viability, and differentiation was evaluated. Cultispher S microcarriers seeded with UMR-106 cells were encapsulated into in situ photopolymerizable polyester scaffolds based on D,L-lactide and -caprolactone. The following fundamental questions related to the encapsulation of UMR106 seeded microcarriers into injectable, in situ photopolymerizable polyester scaffolds will be answered: (1) Will the microcarriers be evenly distributed in the 3-D scaffold? (2) Do the microcarriers remain in the scaffold while the gelatin (porosigen) is leached out? (3) Do cells remain viable after the photoencapsulation process? 2. Materials and Methods 2.1. Materials. DMEM GlutaMAX-I (high glucose) (Cat No. 61965-026), MEM Alpha Medium (+ GlutaMAX-I) (Cat No. 32561029), fetal bovine serum (FBS heat inactivated, E. C. approved), penicillin-streptomycin (10 000 U/mL-10 000 µg/mL), Fungizone, and trypsine-EDTA were purchased from Gibco BRL (Life Technologies, Merelbeke, Belgium). L-Ascorbic acid 2-phosphate, dexamethasone (Cat No. D-2915), 1,25-dihydroxyvitamin D3 (Cat No. D-1530-10UG), and MTT (thiazolyl blue tetrazolium (M-5655)) were supplemented from Sigma (Sigma-Aldrich NV/SA, Bornem, Belgium). Tissue culture dishes were purchased from Greiner bio-one (Wemmel, Belgium). Cultispher S was purchased from Percell Biolytica. Gelatin was from Gelatines Rousselot. 20 mL vials and 50 mL spinner flasks were respectively from Packard Bioscience and Schott Duran. The visible light curing unit, 3M Unitek (500 mW/cm2), Ortholux XT was from 3M Dental Products. UltraClear and MountingClear were obtained from J. T. Baker (Klinipath, Geel, Belgium). The rat osteocalcin EIA kit BT 490 was purchased from Biomedical Technologies, Inc. (Sanbio, Uden, The Netherlands). 2.2. Methods. 2.2.1. Cell Cultures. 2.2.1.1. Culture of UMR-106 Osteosarcoma Cells. UMR-106 rat osteosarcoma cells (ETCC) were routinely grown in monolayer culture in DMEM GlutaMAX-I containing 10 vol % fetal bovine serum, 0.5 vol % penicillinstreptomycin and 1 vol % Fungizone at 5% CO2/95% air, 37 °C. (DMEM medium).11 2.2.1.2. Isolation and Culture of Bone Marrow Derived Cells. Bone marrow cells were obtained from the tibiae and femora of young (6 weeks old) adult male Wistar rats. Skin, soft connective tissue, and periosteum were removed. Tibiae and femora were washed and the diaphyses were cut free of epiphyseal cartilage. The bone marrow was flushed out repeatedly with MEM Alpha medium (containing 10 vol % fetal bovine serum, 0.5 vol % penicillin-streptomycin, 1 vol % Fungizone) and centrifuged (10 min, 1000 rpm). The cell pellet was resuspended in MEM Alpha medium supplemented with 100 µM L-ascorbic acid 2-phosphate and 10 nM dexamethasone and seeded in T75 tissue culture dishes. After 24 h, the medium was changed to remove nonadherent

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cells.11 The cells were cultured until confluence (5% CO2/ 95% air, 37 °C). 2.2.2. Toxicity of Dental Lamp. UMR-106 were seeded in 96-well tissue culture plates at a density of 25 000 cells/ well and 10 0000 cells/well in 200 µL of complete DMEM medium. After 24 h, the regular medium was withdrawn and the cells were irradiated for 0, 10, 20, 40, 80, 160, and 240 s (in triplicate) with the dental lamp curing unit (500 mW/ cm2 blue light). A total of 200 µL of culture medium was added, and the cells were incubated at 37 °C. The number of viable cells was determined by MTT-assay immediately or after 2 days incubation. MTT containing medium (0.5 mg/ mL) was added to the cells and incubated. After 4 h, the MTT containing solution was withdrawn and 200 µL of 1% Triton X-100 in 2-propanol/0.4 N HCl was added. The formazan was dissolved by shaking at 37 °C for 30 min. The absorbance was measured at 580 nm (Universal Microplate Reader EL 800, BIO-TEK instruments Inc.) and the viability was calculated as percentage of the control. 2.2.3. Microcarrier Seeding and Expansion. Macroporous Cultispher S microcarrier beads (gelatin-based), diameter 130-380 µm, were prepared according to the manufacturer’s instructions. Briefly, 0.09 g of microcarriers/vial were hydrated prior to use in PBS and steam-sterilized (15 min, 121 °C). They were rinsed in PBS and twice with MEM Alpha medium (37 °C). The microcarriers were transferred and divided over three wells of a 12 well tissue culture dish (for suspension culture). 2 million UMR-106 and rat bone marrow cells were seeded/0.09 g of CultiSpher in MEM Alpha medium respectively MEM Alpha medium with L-ascorbic acid 2-phosphate and dexamethasone. After 24 h, the microcarriers were transferred to a 50 mL spinner flask (working volume 5 mL, end concentration 18 mg Cultispher S/mL), and the stirring speed was set at 55 rpm. The cultures were maintained at 37 °C, 5% CO2 for 28 days. 2.2.3.1. Microcarrier Colonization. The colonization of the cells on the microcarriers was assessed on days 1, 5, 7, 14, 21, and 28 using propidium iodide staining and histology. 2.2.3.1.1. Propidium Iodide Staining. Samples were taken from the microcarrier spinner cultures and transferred to Eppendorf tubes. After rinsing with Ringer solution, they were fixed with 4% phosphate (10 mM) buffered formaldehyde at 4 °C (10 min). The microcarriers were rinsed with Ringer solution and 0.5 mL Ringer containing propidium iodide (10 µL PI/10 mL Ringer) was added for 5 min. After rinsing with Ringer solution, the microcarriers were transferred to slides and mounted with DABCO containing glycerin and visualized using a fluorescence microscope (excitation 536 nm, emission 617 nm). 2.2.3.1.2. Histology. Samples were taken from the microcarrier spinner cultures. After rinsing with Ringer solution, they were fixed with 4% phosphate (10 mM) buffered formaldehyde (pH 6.9) (4 °C, 24 h). They were dehydrated in a graded alcohol series and embedded in paraffin. 5-7 µm sections were made, stained (hematoxylin & eosin and Masson’s Trichrome), and mounted (with Canada balsam). 2.2.3.2. Cell Viability. MTT analysis was used to monitor cell viability. On the specific time points, 4 mg samples from spinner microcarrier cultures were taken and transferred to

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Eppendorf tubes, and the medium was replaced by MTT containing medium (0.5 mg/mL). After 4 h incubation, the medium was withdrawn, the microcarriers were rinsed twice with Ringer solution, and 0.5 mL of 1% Triton X-100 in 2-propanol/0.4 N HCl was added. The formazan was dissolved by shaking at 37 °C for 30 min. A total of 300 µL was pipetted in a 96 well tissue culture dish, and the absorbance was measured at 580 nm (Universal Microplate Reader EL 800, BIO-TEK instruments Inc.). 2.2.3.3. Cell Differentiation (Osteocalcin Assay). 48 h prior to assaying (21 days), microcarriers seeded with rat bone marrow cells were placed into 5 mL of serum-free osteogenic medium containing 10-9 M 1,25-dihydroxyvitamin D3. After 48 h, the medium was harvested, and osteocalcin was measured at 450 nm by an enzyme immuno assay according to the company’s instructions. 2.2.4. Synthesis of Methacrylate-Endcapped Poly(D,Llactide-co--caprolactone) Oligomer. Linear, telechelic poly(D,L-lactide-co--caprolactone) with hydroxyl groups at the termini was synthesized by the ring-opening polymerization of D,L-lactide (LA) and -caprolactone (CL) in the presence of 1,6-hexanediol as an initiator and zinc acetate as a catalyst. Polymerization was performed in bulk at 140 °C for 48 h and on a 15 g scale. A glass tube was silanized with dichlorodimethylsilane, dried in an oven, and cooled in a desiccator. A typical experimental procedure was as follows: the glass tube was charged with predetermined amounts of comonomers (LA/CL molar ratio 50/50), the initiator, and the catalyst. The target number average molecular weight was 2700 g mol-1. The tube was degassed three times, sealed, and placed in a constant-temperature oven. Polymerization was terminated by cooling the tube in a refrigerator. In the second step, the poly(D,L-lactide-co--caprolactone) diol was endcapped using an excess of methacryloyl chloride (2-folds excess to the total number of hydroxyl groups derived from the oligomer) in the presence of triethylamine (1.5 molar excess to the amount of methacryloyl chloride) in methylene chloride (2:1 volume ratio) to give the corresponding methacrylate-endcapped oligomer. The esterification proceeded at room temperature for 24 h. The triethylamine hydrochloride produced from the reaction was removed by filtration and the remaining solution was concentrated and stored in a freezer for 1 h, and again a white solid was filtrated. Finally, the remaining solution was dialyzed (Spectra/Por Membrane, MWCO: 1000) in dry acetone for 2 days.12,13 The yield was 62%. 2.2.5. Encapsulation of Osteoblast Seeded Microcarriers into the Polymer Scaffold. The methacrylate-endcapped copolymers can be converted into a solid, 3-D polymer network by visible-light irradiation in the presence of D,L-camphorquinone/ethyl 4-dimethylaminobenzoate initiator system. By adding leachable particulates (porosigens), a porous 3-D network can be obtained.12-18 The porosigen gelatin was sieved with ASTM E11-70 standard testing sieves with openings of 250 µm and 355 µm, placed on the Retsch AS 200 sieved shaker until particulates with a defined particle size (250-355 µm) were obtained.

Declercq et al.

Methacrylate-endcapped oligomers (LA50-CL50-HXD20/1bismethacrylate), 15 wt % 2-hydroxyethyl methacrylate (HEMA), catalyst (0.6 wt % D,L-camphorquinone and 0.7 wt % ethyl 4-dimethylaminobenzoate), and gelatin (particle size 250-355 µm) were sterilized by ethylene oxide (12 h, 60 °C, 48 h air).12,13 HEMA was added to the catalyst until dissolved. The HEMA/catalyst solution followed by gelatin is added to the viscous prepolymer and mixed thoroughly. Before encapsulation, the medium of the UMR-106 seeded microcarriers is removed. The microcarriers are further dried by pipetting them on filter paper. The “dry” microcarriers are added to the catalyst/HEMA/polymer/gelatin paste and carefully mixed. The viscous paste is then brought into cylindrical Teflon molds (diameter 5 mm, height 3 mm), and the prepolymer is photopolymerized for 20 s at one side with the dental lamp curing unit. After polymerization, the scaffolds are removed from the mold and sectioned to 1.5 mm height by means of a scalpel. The scaffolds are placed in MEM Alpha medium and incubated at 37 °C for a culture period of 7 days. The medium is replaced after 4-7 h and the following days to remove excess gelatin. A typical experimental procedure was as follows. Scaffolds with apparent porosity of 68 were prepared with respectively 20% and 50% microcarriers versus polymer. Scaffolds with 20% microcarriers initially contained 0.473 g of polymer, 0.8681 g of gelatin, and 0.09 g of microcarriers. Scaffolds with 50% carriers initially contained 0.355 g of polymer, 0.64833 g of gelatin, and 0.18 g of microcarriers. 2.2.5.1. Scaffold Characterization (Porosity). To determine the density of the solid polymeric material, the pycnometric measurements were performed in triplicate. The density of polymeric material is 1.13 g/cm3. The density of gelatin particles is 0.97 g/cm3.17 The porosity of the porous scaffolds was calculated from the theoretical apparent densities of the scaffolds using the following equation described by Mikos et al.: ) [Wg/dg]/[(1 - Wg)/dp + wp/ dp] where is the theoretical prediction of an apparent porosity, wg is the porosigen weight fraction, and dg and dp are the density of respectively the porosigen and the polymer.16 2.2.5.2. Histology. A 7 day culture period was chosen to detect if the encapsulated microcarriers were still present in the scaffold. For the viability, we studied the encapsulated UMR-106 cells immediately after cross-linking. Immediately after cross-linking and 7 h, 24 h, and 7 days after crosslinking, scaffolds were fixed with 4% formaldehyde buffered with 10 mM phosphate (pH 6.9; 4 °C, 24 h). They were dehydrated in a graded alcohol series (UltraClear is used instead of toluene because toluene dissolves the polymer) and embedded in paraffin. 5-7 µm sections were made, stained with hematoxylin and eosin, and mounted with MountingClear. 3. Results 3.1. Toxicity of Dental Lamp. UMR-106 cells showed a 100% viability even after 240 s of irradiation with visible light for both high and low seeding density as evaluated by MTT-assay (data not shown).



Figure 1. Cell viability (proliferation curves) of UMR-106 and rat bone marrow cells grown on Cultispher S microcarriers. On specific time points, the amount of viable cells on 4 mg of cell seeded microcarriers was quantified by MTT assay. Mean and SD of two independent experiments in duplicate.

3.2. Microcarrier Expansion and Colonization. The 24 h static seeding regime allowed the cells to attach to the microcarriers and yielded a relatively homogeneous cell distribution between the microcarriers (seeded with UMR106 and rat bone marrow cells) within the culture as assessed after 1 day by MTT staining, PI staining, and histology. Static seeding was shown to improve the seeding efficiency in comparison to dynamic seeding (data not shown). UMR106 cells attached slightly better on the microcarriers than the rat bone marrow cells, although not significantly. For both UMR-106 and rat bone marrow cells, cells remained viable and the cell number increased over time, as demonstrated by MTT assay (Figure 1). The amount of cells on the microcarriers reached a plateau after 14 days (Figure 1). UMR-106 had a higher proliferation rate than the rat bone marrow cells. On day 21, the amount of cells on the microcarriers was equal or declined as shown by MTT. After 14 days, the cells reached a confluent state on the microcarrier. The cells are not only present around the microcarrier (Figure 2a-d) but the microcarrier is also fully colonized in the interior (Figure 2c,d). Although most microcarriers are completely colonized by cells, some microcarriers only have a limited amount of cells (Figure 2e,f). Microcarriers seeded with bone marrow as well as UMR106 cells formed bridges between the microcarriers and large clusters were made as shown in Figure 2e,f,h. The microcarrier aggregates ranged from two cell seeded microcarriers to more than a dozen microcarriers and started only after 14 days in culture. Bone marrow derived cells not only proliferated on the microcarriers but were also able to differentiate. An abundant extracellular matrix is formed (Figure 2g,h). The extracellular matrix spanned multiple microcarrier beads and formed bridges resulting in large aggregates (Figure 2h). Osteocytes are clearly seen embedded in the extracellular matrix (Figure 2g). The cells were highly differentiated as osteocalcin was secreted in the culture medium (4.2 ( 0.3 ng/mL). 3.3. Encapsulation of Microcarriers into the ThreeDimensional Scaffold. From the above shown cell proliferation curves and previous results concerning the osteoconductivity of the scaffolds, we decided to incorporate UMR106 seeded microcarriers cultured for 14 days in a dynamic

Figure 2. Propidium iodide (a, b), H&E (c, e, g, h), and Trichrome Masson (d, f) staining of UMR-106 (a, c, e) and rat bone marrow cells (b, d, f, g, h) grown on Cultispher S microcarriers after 14 days (a, b, c, e) and 28 days (d, f, g, h).

system into scaffolds with an apparent porosity of 68 obtained by gelatin leaching.19 The viscous polymer allowed us to mix the microcarriers properly with the polymer, the catalyst/initiator, and the porosigen gelatin. Immediately after cross-linking, the microcarriers and gelatin are evenly distributed throughout the polymer scaffold (Figure 3a,d). After 4-7 h incubation in culture medium, the gelatin is leached out completely and a porous osteoconductive 3-D scaffold is formed. Nevertheless, the microcarriers are encapsulated and remain in the 3-D scaffold (Figure 3b,e). Even after a culture period of 7 days, the microcarriers are still present in the scaffold (Figure 3c,f). This demonstrates that the microcarriers are well entrapped in the cross-linked polymer scaffold. Mechanical data of cross-linked 3-D scaffolds are described by Gorski et al.18 The photopolymerization evironment and the molding/ mixing of the microcarriers into the polymer/gelatin paste is not detrimental to the mechanical stability of the microcarriers. When the amount of microcarriers to polymer is 20%, the scaffold is well cross-linked, as described above. Increasing the microcarrier/polymer ratio to 50% leads to inferior crosslinked scaffolds (Figure 3d-f). Even here are the microcar-

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Figure 3. UMR-106 seeded microcarriers encapsulated in threedimensional LA/CL scaffolds immediately after cross-linking (a, d), after 7 h (b, e), and 7 days (c, f) incubation in culture medium. 20% (a, b, c) versus 50% (d, e, f) microcarrier/polymer ratio. H&E stain.

Figure 4. UMR-106 seeded microcarriers encapsulated in threedimensional LA/CL scaffolds. H&E stain.

riers evenly distributed in the polymer matrix, but they are more packed together in the scaffold. Nevertheless, some microcarriers leach out as a result of the reduced cross-linking of the scaffold. 3.3.1. Cell Viability. UMR-106 cells seeded onto microcarriers and encapsulated into the three-dimensional scaffold were analyzed by light microscopy immediately after crosslinking. Inside the microcarriers, cells are clearly present as seen in Figure 4. The cells are still attached to the interior of the microcarrier and no pyknotic nuclei, indicative for dead cells, were observed. The photopolymerization environment and the handling (molding) of the microcarrier/polymer/gelatin paste is not detrimental to the initial viability of the encapsulated UMR-106 cells. 4. Discussion The increasing popularity of arthroscopic procedures in orthopaedics and the requirement to bridge irregular bone

Declercq et al.

defects resulted in great interest in fixation materials (both ceramic and polymeric-based) that are injectable, in situ forming, and biodegradable. However, some critical size bone defects require an additional source of osteogenic cells. Hence, the combination of injectable, in situ forming scaffolds and osteogenic cells is a promising approach. Some examples of injectable materials that harden in situ are the following. The most commonly used injectable bone material poly(methyl methacrylate) (PMMA) is not biodegradable and suffers from high curing temperatures. If the setting reaction involves a temperature change, the increase or decrease should be as small as possible to reduce damage to the surrounding tissue. Calcium phosphate pastes (e.g., octa calcium phosphate, tricalcium phosphate, etc.) are undergoing after several minutes a nonexothermic setting to form materials with high compressive strength. Injectable polymers must be cross-linked in situ. This curing is usually initiated either chemically or via the use of light.5 In our group, a variety of in situ, photopolymerizable 3-D scaffolds were synthesized with variables of polymer type (lactide, -caprolactone, trimethylene carbonate, etc.), copolymer ratio, and intitiator systems.12,13 These scaffolds were compared regarding to degradation time, mechanical strength, biocompatibility, osteoconductivity and osteoblast behavior. Scaffolds with different apparent porosities (50, 60, and 70) and different porosigens (gelatin, sodium chloride, and sugar) were made, and their osteoconductivity was compared by SEM and histological analysis. The in situ, photopolymerizable scaffold based on D,L-lactide and -caprolactone that is also used in this study showed good results concerning the toxicity according to the ISO norms. Scaffolds with a porosity of 70 obtained by gelatin leaching showed a good interconnected pore morphology. Cellular infiltration was inferior in scaffolds obtained by sodium chloride and sugar leaching. Rat bone marrow derived cells attached and differentiated on these porous 3-D scaffolds based on LA/ CL and gelatin leaching. After three weeks, the cells expressed high alkaline phosphatase activity, formed an abundant extracellular matrix, secreted osteocalcin and were able to mineralize the extracellular matrix.19 However, the bone formation was only limitited to the edges of the scaffold. Hence, encapsulation of cells in the 3-D scaffold would be promising. However, cells should be protected from the photopolymerization process and also from handling the viscous polymer paste. Then, in vivo, bone formation could start already in the center of the implant while the polymeric scaffold is degrading. Microcarriers loaded with cells offer the advantage of being able to readily place large number of cells directly and controllably in an injectable and in situ forming scaffold. The encapsulation of cells (chondrocytes and osteoblasts) is mostly limited to hydrogels (agarose, gelatin, alginate, fibrin, etc.) or photopolymerizable hydrogels.7-10,20 These gel encapsulated cells showed a good viability. Although some diffusion problems exist. An increased macromer concentration resulted in a decreased viability.4,7,8 Most hydrogels degrade very fast and are often used as a cell delivery system. The cells are released within a few hours to days into the site of implantation. In this study, we have


chosen a system where the cells can be cultured in vitro for a prolonged time (several weeks) before encapsulation into the implant site. Also the mechanical stability of a hydrogel might be not high enough to protect the encapsulated cells from handling the viscous polymer/hydrogel mix. Hence, for epithelial-like cells or cells showing a polarity (e.g., osteoblast-like cells), macroporous microcarriers could be promising. A variety of porous microcarriers or microspheres are mentioned in the literature: microcarriers based on collagen (Cultispher S, Cultispher G, Cellagen), reconstituted collagen and hydroxyapatite, poly(lactide-co-glycolide), and bioceramics.21-28 The purpose of these studies was mainly to improve the expansion of primary cells or to inject the cell seeded microcarriers immediately in the site of injury.21-28 These microcarriers were seeded with different cell types (e.g., calvaria cells, chondrocytes, and mesenchymal cells) and their expansion and differentiation on the microcarriers were studied. In another study, the ability of chondrocytes to redifferentiate after expansion on the microcarriers was evaluated. We have chosen for a microcarrier Cultispher S (based on cross-linked gelatin) which is highly macroporous and has a good mechanical stability. In our experiments, both UMR-106 and primary rat bone marrow cells attached and proliferated well. Within 14 days, the microcarriers were completely colonized. Moreover, rat bone marrow cells were able to differentiate. An abundant extracellular matrix was formed and cells secreted osteocalcin. Also Doctor et al. described a high alkaline phosphatase activity and the deposition of calcium when human mesenchymal stem cells were cultured on Cultispher S microcarriers.28 When the cell seeded microcarriers were cultured for prolonged incubation periods, the microcarriers started to form aggregates which has also been described by others.22,24 As this could lead to a lower distribution of the microcarriers in the threedimensional scaffold, we decided to incorporate the UMR106 seeded microcarriers after 14 days of incubation. Encapsulation of cell seeded microcarriers into in situ, photopolymerizable 3-D polymer scaffolds was not yet described elsewhere. Photopolymerization reactions are driven by chemicals that produce free radicals when exposed to specific wavelengths of light. A photon from a light source excites the photoinitiator into a high-energy radical state, and this radical then induces the polymerization of a macromer solution. Photopolymerization has several advantages over conventional polymerization techniques. These include spatial and temporal control over polymerization, fast curing rates (less than a second to a few minutes) at room temperature or physiological temperatures, and minimal heat production.7 Polymers can be cross-linked in the presence of photoinitiators using visible or ultraviolet (UV) light. As ease of handling is of high importance for clinical use, the viscous properties must be balanced between the need for the material to remain at the site of injection and the need for the surgeon to easily manipulate its placement.5 The copolymer based on CL and LA with monomer molar ratio 50/50 and molecular weight of 2700 g/mol is viscous-liquid and possesses a glass transition temperature of -32 °C because of the internal plasticizing effect of the aliphatic chains derived from CL units and the copolymer amorphous


structure. Moreover the addition of 2-hydroxyethyl methacrylate (HEMA) as the co-cross-linker significantly reduces the polymerization mixtures viscosity. Hence, an injectable system with ideal viscous properties can be obtained. The viscous polymer allowed us to mix the microcarriers properly with the polymer, the catalyst/initiator, and the porosigen gelatin. An injectable material should set in situ in several minutes to minimize the length of the procedure while allowing surgeons enough time for placement before hardening of the material. In our experiments, the cross-linking of the 3-D scaffold was completed within 20 s. As gelatin melts at a temperature of 37 °C, a porous 3-D network can be obtained within 4-7 h. 20% microcarriers is a good quantity to encapsulate as the microcarriers remain in the scaffold after the gelatin already has leached out. Increasing the microcarrier concentration leads to scaffolds with a reduced cross-linking density and some microcarriers leached out. Although leaching out of the microcarriers should not be considered inferior in the light of the clinical application, an increased concentration of microcarriers did result in inferior mechanical characteristics. The mechanical characteristics of 3-D scaffolds without encapsulated microcarriers are described by Gorski et al.18 3-D scaffolds with encapsulated microcarriers were not tested mechanically, but we observed, by handling the scaffolds, that the mechanical strength was not optimal. Parts of the polymer were not crosslinked and leached out of the scaffold. The creation of a high-energy radical species creates the potential for oxidative damage to the cell populations that are photoencapsulated. Free radicals can cause damage to cell membranes, nucleic acids, and proteins, which can lead to cell death.6 UMR-106 cells seeded on microcarriers and encapsulated in the 3-D scaffold showed a good viability. The irradiation with the visible light source (400-520 nm) did not decrease the viability as shown by the MTT-assay. Only one study describes a similar system: osteoblast-like cells were encapsulated in an alginate gel and subsequently mixed with a calcium phosphate paste that hardens in situ.20 The cells remained viable, although a comparison cannot be made with our results, because the cells are encapsulated in a hydrogel and the hardening of the calcium phosphate paste does not involve a photopolymerization process. Some groups studied the cytocompatibility of visible light photoinitiating systems. They found that the cell-damaging effects with the camphorquinone system were rather mild and at low concentrations, camphorquinone is the most promising cytocompatible visible light initiating system.29,30 During encapsulation of cells into 3-D scaffolds, reactive polymer groups (methacrylate groups) would react with free radicals, thus decreasing the deleterious effect of the radicals on the encapsulated cells.6 The cytotoxicity of compounds including photoinitiators has been thought to be due in part to their hydrophobicity since permeability through phospholipid bilayers of cellular membranes increases with the hydrophobicity of a compound. Also the cellular proliferation rate may play a role in cellular toxicity of radical photoinitiators.6 The polymerization of methacrylate groups is an exothermic process. The most commonly used injectable material, PMMA-based bone cement, suffers from high curing tem-

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peratures. In the case of this particular copolymer used in our work, no significant temperature increase was observed during the photocuring process and no negative influence of temperature on the encapsulated cells was observed. For methacrylate-endcapped polymers with low molecular weight (less than 2700 g/mol), the number of methacrylate endgroups is higher. As the cross-linking involves an exothermic process, this increased amount of methacrylate side-chains would lead to a negative influence on the encapsulated cells. A copolymer with high molecular weight (e.g., 2700 g/mol) leads to lower heat release during the polymerization process. 5. Conclusions Osteoblast-like cells (UMR-106 and rat bone marrow) attached, proliferated, and differentiated well on the highly macroporous microcarriers (Cultispher S). The microcarriers were completely colonized after 14 days in dynamic culture. As microcarrier aggregates were formed starting from day 14, the UMR-106 seeded microcarriers were encapsulated after 14 days in the three-dimensional scaffold. The microcarriers and the porosigen gelatin are evenly distributed throughout the scaffold. Within a few hours, the gelatin leached out, and a porous scaffold was obtained. The cell seeded microcarriers remained in the scaffold even after a culture period of 7 days. The present study demonstrates that the photopolymerization system is sufficiently mild (low light intensity, short irradiation time, physiological temperature, and low organic solvent levels) and can be carried out in the presence of cells. Acknowledgment. The work was supported by a fund of the Ghent University (GOA project 2001, No 12050701). The authors thank M. Debruycker and L. Pieters for technical assistance. References and Notes (1) Hutmacker, D. W. Biomaterials 2000, 21, 2529-25433. (2) Hench, L. L.; Polak, J. M. Science 2002, 295, 5557, 1014-1016. (3) Rose, F. R. A. J.; Oreffo, R. O. C. Biochem. Biophys. Res. Com. 2002, 292, 1-7. (4) Burdick, J. A.; Anseth, K. S. Biomaterials 2002, 23, 4315-4323. (5) Temenoff, J. S.; Mikos, A. G. Biomaterials 2000, 21, 2405-2412. (6) Williams, C. G.; Malik, A. N.; Kim, T. K.; Manson, P. N.; Elisseeff, J. H. Biomaterials 2005, 26, 1211-1218.

Declercq et al. (7) Nguyen, K. T.; West, J. L. Biomaterials 2002, 23, 4307-4314. (8) Elisseeff, J.; McIntosh, W.; Anseth, K.; Riley, S.; Ragan, P.; Langer, R. J. Biomed. Mater. Res. 2000, 51, 164-171. (9) Wang, L.; Shelton, R. M.; Cooper, P. R.; Lawson, M.; Triffitt, J. T.; Barralet, J. E. Biomaterials 2003, 24, 3475-3481. (10) Mol, A.; van Lieshout, M. I.; Dam-de Veen, C. G.; Neuenschwander, S.; Hoerstrup, S. P.; Baaijens, F. P. T.; Bouten, C. V. C. Biomaterials 2005, in press. (11) Declercq, H. A.; Verbeeck, R. M. H.; De Ridder, L. I. F. J. M.; Schacht, E. H.; Cornelissen, M. J. Biomaterials 2005, 26, 49644974. (12) Jackers, G. pHd thesis, 2002. Proefschrift voorgelegd tot het behalen van de graad: Doctor in de Wetenschappen, Scheikunde. Promotor: E. Schacht. Polymer Material Research Group, Krijgslaan 281 (S4), Ghent University, B-9000 Ghent, Belgium. (13) Schacht, E.; Gorski, T.; Jackers, G.; Cornelissen, M.; Declercq, H.; Verbeeck, R.; Van den Vreken, N.; Gasthuys, F.; Vlaminck, L. Abstract. 40th IUPAC Symposium on Macromolecules, July 4-9th, 2004, Paris. (14) Robson, F.; Storey, R. F.; Warren, S. C.; Allison, C. J.; Wiggins, J. S. Polymer 1993, 34, 4365-4372. (15) Kucybala, Z.; Pietrzak, M.; Paczkowski, J.; Linden, L. A.; Rabek, J. F. Polymer 1996, 20, 4585-4596. (16) Mikos, A. G.; Thorsen, M. J.; Czerwonka, L. A.; Bao, Y.; Langer, R.; Winslow, D. N.; Vacanti, J. P. Polymer 1994, 35, 1068-1077. (17) Thomson, R. T.; Yaszmski, M. J.; Powers, J. M.; Mikos, A. G. J. Biomater. Sci. Polym. Ed. 1995, 7, 23-28. (18) Gorski, T.; Jackers, G.; Schacht, E. 2005, submitted for publication. (19) Declercq, H. A.; Gorski, T. L.; Schacht, E. H.; Cornelissen, M. J. 2005, Submitted for publication. (20) Simon, C. G.; Guthrie, W. F.; Wang, F. W. J. Biomed. Mater. Res. 2004, 15, 68 A, 4, 628-639. (21) Malda, J.; Kreijveld, E.; Temenoff, J. S.; van Blitterswijk, C. A.; Riesle, J. Biomaterials 2003, 24, 5153-5161. (22) Overstreet, M.; Sohrabi, A.; Polotsky, A.; Hungerford, D. S.; Frondoza, C. G. In Vitro Cell DeV. Biol.-Animal 2003, 39, 228234. (23) Frondoza, C.; Sohrabi, A.; Hungerford, D. Biomaterials 1996, 17, 879-888. (24) Bouchet, B. Y.; Colon, M.; Polotsky, A.; Shikani, A. H.; Hungerford, D. S.; Frondoza, C. G. J. Biomed. Mater. Res. 2000, 52, 716-724. (25) Newman, K. D.; McBurney, M. W. Biomaterials 2004, 25, 57635771. (26) Hsu, F. Y.; Chueh, S.-C.; Wang, Y. J. Biomaterials 1999, 20, 19311936. (27) Wu, T.-J.; Huang, H.-H.; Lan, C.-W.; Lin, C.-H.; Hsu, F.-Y.; Wang, Y.-J. Biomaterials 2004, 25, 651-658. (28) Doctor, J.; Petraglia, C. A.; Loveland, A.; Dietz, M.; Minich, E.; Leung, J.; Hollinger, J. DeV. Biol. 2002, 247, 505. (29) Atsumi, T.; Murata, J.; Kamiyanagi, I.; Fujisawa, S.; Ueha, T. Arch. Oral Biol. 1998, 41, 1, 73-81. (30) Bryant, S. J.; Nuttelman, C. R.; Anseth, K. S. J. Biomater. Sci.Polym. Ed. 2000, 11, 5, 439-457.

BM050031S

Encapsulation of Osteoblast Seeded Microcarriers ... - ACS Publications

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