Self-assembled octapeptide scaffolds for in vitro chondrocyte culture

Acta Biomaterialia 9 (2013) 4609–4617

Contents lists available at SciVerse ScienceDirect

Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat

Self-assembled octapeptide scaffolds for in vitro chondrocyte culture Ayeesha Mujeeb a, Aline F. Miller b, Alberto Saiani a, Julie E. Gough a,⇑ a b

School of Materials, The University of Manchester, Grosvenor Street, Manchester M13 9PL, UK School of Chemical Engineering and Analytical Science & Manchester Institute of Biotechnology, University of Manchester, 131 Princess Street, Manchester M1 7DN, UK

a r t i c l e

i n f o

Article history: Received 12 April 2012 Received in revised form 1 August 2012 Accepted 28 August 2012 Available online 8 September 2012 Keywords: Self-assembly Octapeptides Hydrogels Cell culture Chondrocytes

a b s t r a c t Nature has evolved a variety of creative approaches to many aspects of materials synthesis and microstructural control. Molecular self-assembly is a simple and efficient way to fabricate complex nanostructures such as hydrogels. We have recently investigated the gelation properties of a series of ioniccomplementary peptides based on the alternation of non-polar hydrophobic and polar hydrophilic residues. In this work we focus on one specific octapeptide, FEFEFKFK (F, phenylalanine; E, glutamic acid; K, lysine). This peptide was shown to self-assemble in solution and form b-sheet-rich nanofibres which, above a critical gelation concentration, entangle to form a self-supporting hydrogel. The fibre morphology of the hydrogel was analysed using transmission electron microscopy and cryo-scanning electron microscopy illustrating a dense fibrillar network of nanometer size fibres. Oscillatory rheology results show that the hydrogel possesses visco-elastic properties. Bovine chondrocytes were used to assess the biocompatibility of the scaffolds over 21 days under two-dimensional (2-D) and three-dimensional (3-D) cell culture conditions, particularly looking at cell morphology, proliferation and matrix deposition. 2-D culture resulted in cell viability and collagen type I deposition. In 3-D culture the mechanically stable gel was shown to support the viability of cells, the retention of cell morphology and collagen type II deposition. Subsequently the scaffold may serve as a template for cartilage tissue engineering. Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Cartilage is avascular with a limited capacity to self-repair. Therefore, there is a need to construct tissue engineered scaffolds with a specific de novo [1,2] design that mimics the extracellular matrix (ECM) of natural cartilage tissues and promotes regeneration. Tissue regeneration can be approached in two ways, by cell therapy, which involves the use of isolated cells to replace defective cells or promote tissue growth, or, secondly, by fabricating scaffolds in vitro for tissue growth in vivo. Successful tissue engineered scaffolds must have the following characteristics: (i) be easy to handle; (ii) have mechanical properties similar to the damaged tissue to be replaced or repaired; (iii) be biocompatible with the human body to avoid inflammation; (iv) be reliable and reproducible under culture conditions; (v) serve as a template to promote cell adhesion, proliferation and ECM formation. The cellular matrix composition varies depending on the type and functional requirements of each tissue [3]. The self-assembly of macroscopic materials from molecular building blocks provides an extremely powerful method to design materials such as hydrogels [1,4–8]. Self-assembly is recognized as ⇑ Corresponding author. Tel.: +44 (0) 161 3068958. E-mail addresses: [email protected] (A. Mujeeb), j.gough@manchester. ac.uk (J. E. Gough).

a process in which small peptide molecules spontaneously organize into well-ordered structures by reversible, non-covalent interactions, including hydrogen bonding, disulphide bonds, van der Waals forces, electrostatic interactions, hydrophobic interactions and p–p stacking [5,9,10]. The resulting biological architectures are biodegradable and cytocompatible (non-toxic) [6,7]. The tailored peptide-based scaffolds are highly relevant to tissue engineering, drug delivery systems, protein therapeutics and molecular biosensing [9]. Previously peptide-based hydrogels have proved to be promising ECM mimics. Zhang and colleagues described the use of ionic complementary peptides with alternating charged (hydrophobic and hydrophilic) amino acid residues, RAD16 (RADARADARADARADA) as a peptide-based hydrogel for the culture of nerve cells and chondrocytes for in vitro tissue repair [2,11]. In another study Kisiday and co-workers demonstrated the encapsulation of chondrocytes in vitro within a self-assembled peptide hydrogel, KLD12 (AcN-KLDLKLDLKLDL-CNH2) [2], developed for cartilage repair. The hydrogels supported chondrocyte proliferation and deposited a cartilage-like ECM rich in proteoglycans and collagen type II, illustrating a stable chondrocyte phenotype [2]. Pochan and Schneider also developed a series of 20 amino acid residue peptides, e.g. MAX1 (H2-NVKVKVKVKVDPPTKVKVKVKV-CONH2), designed to self-assemble into b-hairpin based hydrogels to promote NIH 3T3 murine fibroblast proliferation [12]. Stupp and colleagues demonstrated the use of bioactive

1742-7061/$ - see front matter Ó 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.actbio.2012.08.044

4610

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

peptide amphiphile (PA) molecules that self-assembled to form nanofibrous scaffolds to support differentiation of neural progenitor cells [13–15]. These investigators mainly focused on using the self-assembling properties of longer peptides, typically containing 12–20 amino acids, which served as scaffolds for tissue engineering applications. Gazit and Ulijn studied the self-assembly behaviour of short dipeptides (with protecting groups) based on aromatic p–p stacking [16,17]. Different di-phenylalanine analogues, e.g. naphthalene (Nap), benzyloxycarbonyl (Cbz), 9-fluorenylmethoxycarbonyl (Fmoc) [16], and t-butoxycarbonyl (Boc) [18], were used and their self-assembling properties and cell culture condition characteristics were compared. The novelty of this research describes the potential use of a FEFEFKFK (with alternating polar and non-polar amino acid residues) hydrogel with excellent tunable mechanical properties to support two-dimensional (2-D) and three-dimensional (3-D) culture of bovine chondrocytes in vitro. The gel consists of eight natural amino acids (without protecting groups), much shorter than the RAD16, KLD12 and MAX1 peptide-based hydrogels, and has previously been reported to have excellent gelation properties. In this paper the cell viability, ECM production and effect of collagen deposition on the rheological properties of these gels under cell culture conditions in vitro has been demonstrated. 2. Materials and methods 2.1. Peptide synthesis Amino acids, the activator 2-(6-chloro-1H-benzotriazole-1-yl)1,1,3,3-tetramethylaminium hexafluorophosphate (HCTU) and Wang resin were purchased from Novabiochem (Merck, UK) and used as-received. All other reagents and solvents were purchased from Sigma–Aldrich (UK). The octapeptide FEFEFKFK was synthesized in the laboratory by solid phase peptide synthesis (SPPS) in a ChemTech ACT 90 peptide synthesizer (Advance ChemTech Ltd, UK), using Fmoc-Lys(Boc)-Wang resin (mesh 200, loading 0.7 mmol g–1). The product was purified by repeated precipitation in cold ether and then freeze-dried and stored at 4 °C in powder form. The peptide was characterized by high performance liquid chromatography (HPLC) and mass spectroscopy (MS) and the purity estimated to be >85%.

thoroughly using a pipette tip for a few seconds. Subsequently an opaque hydrogel formed in less than 1 min. 2.3. Transmission electron microscopy (TEM) The morphology and width of the fibres within the self-assembled hydrogel were investigated using a Tecnai 10 transmission electron microscope operating at 100 eV (calibrated magnification 43,600) with Kodak SO-163 film. Images were subsequently scanned using a UMAX2000 transmission scanner providing specimen level increments of 0.366 nm per pixel. Carbon-coated copper grids (400 mesh, Agar Scientific) were glow discharged for 5 s and placed shiny side down onto a 10 ll droplet of diluted hydrogel for 10 s. Loaded grids were immediately placed on a 10 ll droplet of double deionized water for 10 s and then blotted. Washed grids were then placed on a 10 ll droplet of freshly prepared and filtered uranyl acetate solution (4% w/v) for 60 s for negative staining and then blotted continuously against double folded Whatman 50 filter paper. 2.4. Cryo-scanning electron microscopy (cryo-SEM) Cryo-SEM was performed using a Philips XL30 ESEM-FG equipped with an Oxford Instrument Alto CT2500 cryo-transfer system. The sample prepared was placed between the rivets located on one end of the transfer rod and then slam-frozen in liquid nitrogen before being transferred to the cryostat chamber. The temperature in the chamber was maintained at 150 °C. The coated sample was then transferred to the microscope chamber. Images were captured with the microscope operating under high vacuum and with an accelerating voltage of 5 keV. 2.5. Rheology The viscoelastic properties of the hydrogel were investigated in a Bohlin C-VOR 200 digital rheometer with a 20 mm/2° parallel plate geometry. The elastic (G0 ) and viscous (G00 ) moduli of the hydrogels were recorded as a function of frequency sweeps between 0.1 and 100 Hz at 0.1% strain. A solvent trap was used to keep the hydrogels hydrated and the temperature was maintained at 25 °C. Oscillatory tests were also carried out using cell seeded 2D and 3-D gels. For this experiment all variables were kept constant and the temperature was maintained at 37 °C.

2.2. Hydrogel formation 2.6. Environmental scanning electron microscopy (ESEM) Peptide solutions were prepared in 2 ml microtubes by suspending the powder at a concentration of 20 mg ml–1 in distilled water. The peptide solutions were vortexed to obtain a homogeneous mixture. Tubes were placed in an oven at 90 °C for 3 h to aid complete powder dissolution. 500 ll of each solution was transferred to 24-well plates. The samples were neutralized to pH 7 by drop by drop addition of sterile 0.5 M NaOH solution. The contents in the well plate were stirred thoroughly using a pipette tip for a few seconds. Gelation occurred less than 1 min after adding NaOH. Following gelation the samples were sterilized for 30 min under UV and then incubated at 37 °C in a humidified atmosphere (20% O2) with 5% CO2 overnight (to equilibrate the environment to physiological pH). In the 3-D experiments the peptide solutions were prepared by suspending sterilized peptide powder at a concentration of 30 mg ml–1 in a 70:30 mixture of dH2O:1  PBS in 2 ml microtubes. The tubes were incubated in an oven at 90 °C for 3 h to aid complete powder dissolution. The tubes were removed from the oven and briefly vortexed. 500 ll per well was plated in 24-well plates using a 1000 ll Gilson pipette. 40 ll of 0.5 M NaOH solution was immediately added drop by drop on top of the samples and stirred

For ESEM analysis an FEI Quanta 200 ESEM operating in ‘‘wet’’ mode was used. The chamber was set at 100% relative humidity (Peltier cooler at 5 °C and 6.2 Torr chamber pressure) at first and then the chamber pressure was progressively decreased to enhance surface detail (as the surface dried). The samples for ESEM were prepared at a concentration of 30 mg ml–1 (3-D hydrogels with encapsulated cells). 2.7. Cell culture Bovine chondrocytes were isolated from cartilage from the proximal side of the metacarpalphalangeal joint obtained courtesy of Higginshaw abattoir (UK). Chondrocytes were cultured in Dulbecco’s modified Eagle’s medium containing 1000 mg l–1 D-glucose, Glutamax™, 10% foetal bovine serum, 1% penicillin/streptomycin (antibiotics) and 25 mg l–1 ascorbic acid. When a sub-confluent cell layer was observed in the cell culture flask it was trypsinized using a 0.05% trypsin–0.53 mM EDTA4Na solution (Gibco-Invitrogen, UK) and transferred to a Falcon tube to be centrifuged in order to remove the trypsin solution. The cell pellet at the bottom of the

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

tube was resuspended in fresh medium and adjusted to the required cell concentration before seeding on or in the peptide hydrogels. Cells were counted using a haemocytometer with a 1:1 mixture of trypan blue. 2.8. 2-D/3-D cell culture Following 2-D cell culture after overnight incubation the surface of the gel was washed very carefully three times with fresh medium. 200 ll of medium with suspended cells (2  103 cells per well) was added on top of the gel, followed by another 100 ll addition of fresh medium. Chondrocytes were evenly distributed on the gel surface by gently pipetting the medium. The medium was changed every other day. Following 3-D cell culture, after the addition of 0.5 M NaOH 100 ll of medium with 105 suspended cells per well was immediately added. The contents of the well plates were thoroughly stirred using a pipette tip for a few seconds to allow gelation. It was determined that stirring with a pipette tip did not lead to bubble formation. 150 ll of fresh medium was added on top of the gel and left in the incubator at 37 °C in a humidified atmosphere (20% O2) with 5% CO2 for 1 h. The samples were subsequently removed from the incubator and fresh medium washes were repeated three times at 15 min intervals to stabilize the gel at pH 7.2 (±0.4). The medium on top of the 3-D gel was changed every day for the first 7 days to aid cell growth within the gel. 2.9. Live/dead assay Cell viability was tested by a Live/Dead assay (Invitrogen, UK). 1 ml of PBS containing 2.5 ll ml–1 of 4 lM ethidium homodimer1 (EthD-1) assay solution and 1 ll ml–1 of 2 lM calcein AM assay solution was prepared. 100 ll of the Live/Dead solution was added on top of each sample for 15 min in an incubator at 37 °C in a humidified atmosphere (20% O2) with 5% CO2. The staining solution was removed and the samples were then viewed under a Nikon Eclipse E600 fluorescence microscope with 494 nm (green, Calcein) and 528 nm (red, EthD-1) excitation filters. Images were captured using Lucia software. For quantitative analysis a total of 250 cells were counted from each sample (mounted onto a glass slide using Prolong™ Gold anti-fade reagent (Invitrogen, UK)) over five randomly chosen areas and the viable and non-viable cells counts were recorded. 2.10. Lactate dehydrogenase (LDH) assay The LDH assay (Promega, UK) is a colorimetric assay which measures an enzyme which is released upon cell lysis. It was used to quantify the number of viable cells in each gel. At each time interval medium was removed from the wells, the gels transferred in microcentrifuge tubes and 500 ll of fresh medium added. Samples were kept in the freezer at –80 °C for 30 min and then thawed for another 30 min at 37 °C. The freeze–thaw cycles were carried out three times to lyse the cells. The microcentrifuge tubes were centrifuged at 250g for 5 min and 50 ll of the supernatant was transferred to a 96-well plate. 50 ll of the substrate mix was added. The well plates were incubated at room temperature. After 30 min the stop solution (containing acetic acid) was added to stop the reaction before readings were taken. Readings in triplicate using a Lab Systems Ascent colorimetric plate reader with absorbance at 492 nm were recorded. Viable cell numbers were determined using a standard curve (average values normalized to a standard curve of known numbers of cultured cells, i.e. 106–103 cells). Gels without cells served as a negative control and the background absorbance was subtracted from the recorded absorbance values.

4611

2.11. Cell count For each time point medium was removed from on top of the samples (prepared in 24-well plates). 1 ml of fresh medium was added on top of each sample to dissolve the hydrogels. The cell– gel suspension was transferred in 2 ml microcentrifuge tubes. A vortex was gently applied to form a homogeneous mixture. 50 ll of the cell suspension was transferred in a separate microcentrifuge tube and diluted by adding 50 ll of trypan blue. A haemocytometer was used to count viable cells. 2.12. Immunocytochemistry staining Gels were rinsed with PBS and fixed with 3% paraformaldehyde (PFA) for 15 min at room temperature. Following fixation the samples were washed using PBS before incubation with immunocytochemistry (ICC) blocking buffer for 30 min. The ICC blocking buffer consists of 1% goat serum, 0.1% Triton X-100 and 50 mg ml– 1 bovine serum albumin (BSA). The samples were incubated in ICC block solution containing the primary antibody (rabbit polyclonal antibody to collagen type I/II purchased from Abcam, UK) for 1 h at room temperature. To remove the non-bound primary antibody samples were washed three times with PBS. Subsequently TRITCconjugated secondary antibody (goat anti-rabbit IgG-Alexa FluorÒ 594) was added for 1 h in the dark. To remove the non-bound secondary antibody samples were further washed five times with PBS. The gels were mounted on the glass slides with Prolong™ Gold antifade reagent with DAPI (Invitrogen, UK). Images were captured using a Nikon Eclipse E600 fluorescence microscope. 2.13. Collagen assay The total collagen production in the scaffolds was determined using a Sircol™ collagen assay kit (Biocolor, UK). Samples were hydrolysed in 24-well plates for 16 h at 4 °C with pepsin at a concentration of 0.1 mg ml–1 in 0.5 M acetic acid, pH 3. 1 ml of Sircol™ red reagent was added (on top) to 50 ll of each sample (adjusting the samples to a total volume of 100 ll with 0.5 M acetic acid) and mixed gently using a mechanical shaker for 30 min at room temperature. The collagen–dye complex was precipitated by centrifugation at 1200 r.p.m. for 10 min, the supernatant was drained out and the collagen–dye pellet was dissolved in 1 ml of alkali reagent. The mixture was vortexed and once all the bound dye had dissolved 200 ll of each sample was transferred to individual wells of a 96-well plate. Absorbance measurements were taken at 555 nm, using a multi-well scanning spectrometer (Ascent microplate reader). Gel without cells served as a negative control and the background absorbance was subtracted from the recorded absorbance values. The assay was performed in triplicate. 2.14. Statistical analysis Statistical analysis was performed using one-way ANOVA to indicate whether any difference in the mean cell levels (measured by the LDH and collagen assays) between the three test time points (days 1, 14 and 21) were significant. A significant F value is indicative of a significant difference in cell numbers between the three days where the probability is less than 0.05, 0.01 and 0.001 (P < 0.05, 0.01 and 0.001). 3. Results and discussion 3.1. 2-D hydrogel formation We have adopted a cost-effective approach to the design of a novel peptide-based hydrogel for cell culture applications. The

4612

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

self-assembling peptide FEFEFKFK used was eight amino acids long, compared with 16, 12 and 20 for the RAD16 [2,11], KLD12 [2] and MAX1 [12] peptide systems reported in the literature. The octapeptide FEFEFKFK was shown to form transparent and self-supporting gel at low pH by simply dissolving the peptide powder at the desired concentration in distilled water at high temperatures (90 °C). Upon cooling to room temperature a stable hydrogel formed within 10 min [19]. However, for cell culture applications gels at physiological pH are required and therefore in this work a 0.5 M NaOH solution was used to adjust the pH of the gel after heating. To investigate the gelation dynamics of the hydrogel a phase diagram at physiological pH was constructed (Fig. 1A). The critical gelation concentration (CGC) of the sample was assessed using the tilting test tube method, i.e. a sample was classified as a liquid when the sample flowed freely and as a gel when the sample was self-supporting upon inversion of the vial. At room temperature the CGC concentration of the peptide under physiological conditions was found to be 5 mg ml–1, i.e. the sample is a viscous solution at concentrations lower than 5 mg ml–1 and a self-supporting hydrogel at higher concentrations. Gelation of the peptide solution was easy and occurred rapidly within less than 1 min of adding NaOH. It has been reported that the gelation of a RAD16-I (Puramatrix™, BD Biosciences, 354250) peptide hydrogel is only initiated by salt concentrations P1 mM (http:// www.bdbiosciences.com/external_files/dl/doc/manuals/live/web_ enabled/354250Lpug.pdf). The estimated time to complete gelation under cell culture conditions in vitro varies between 5 and 30 min, depending on the peptide concentration and the ionic salt concentration used (BD Biosciences) [20–22]. 3.2. Nanofibre morphology and molecular arrangement under cell culture conditions TEM (Fig. 1B) and Cryo-SEM (Fig. 1C) were used to investigate the fibre morphology of the hydrogel under cell culture conditions. T (°C)

95 85

F EF EF KF K at p H 7 wit h N aOH

A

The micrographs in Fig. 1 reveal the presence in the hydrogel of a dense fibrillar network formed by elongated interconnected nanofibres with an average diameter of 3 nm, in good agreement with the morphology described in previous works [19]. The network topology and fibrillar dimensions lie within the range found for ECM components, i.e. 5–300 nm [16]. Fig. 1D shows a schematic diagram illustrating the formation of nanofibrous hydrogels, where cells were either seeded on top of the gels (2-D culture) or encapsulated within the gels (3-D culture) to investigate the cell morphology, proliferation and ECM formation. 3.3. Oscillatory rheology Rheology studies using different concentrations of FEFEFKFK, 20, 30 and 40 mg ml–1, were conducted over the frequency range 0.1–100 Hz (Fig. 1E). The average recorded G0 values were 10– 20 Pa for the pure hydrogel at pH 3, 1–11 kPa for the hydrogel in 0.5 M NaOHat pH 7, and 6–35 kPa for the hydrogel in tissue culture medium and 0.5 M NaOHat pH 7, depending on the concentration used. These results highlight the ease with which the modulus of these gels can be tuned. From Fig. 1E it is evident that the G0 values increase with increasing peptide concentration due to an increase in fibre aggregation. We hypothesize that at pH 7, which correspond to the isoelectric point of the peptide, the repulsive electrostatic interaction between individual fibres decreases, which allows the fibres to come closer together and aggregate to form strong network junctions, resulting in an increase in gel mechanical properties. Material strength is an important mechanical characteristic that could have an impact on cell function in hydrogels [23]. Kisiday and Zhang first used RAD16 gel for chondrocyte culture, but since the gel was mechanically weak the group then designed another peptide, KLD12, for cell culture applications [24]. Previously researchers have described materials with G0 values in the range 10–100 kPa that are known to promote chondrocyte attachment

B

C

Liq uid

75 65 55

Gel

45 35 25 0

10

20

30

40

50

60 C/mg ml-1

D

E

100000

G' (Pa)

10000 1000

Pure gel Gel + NaOH

100

Gel + NaOH + medium

10 1 10

Conc mg ml-1

100

Fig. 1. Characterization of the nanofibrous hydrogel structure. (A) Phase diagram of the self-assembled hydrogel at pH 7. (B) TEM micrograph illustrating the nanofibres present in the gel matrix at pH 7. (C) Cryo-SEM image of the gel at pH 7. (D) Schematic diagram showing the protocol for developing hydrogels for 2-D and 3-D cell culture in vitro. (E) Plot of the elastic modulus (G0 ) of hydrogels with different peptide concentrations, 20, 30 and 40 mg ml–1. The graph illustrates the effects of NaOH and tissue culture medium on the G0 value of the gels at 0.1% strain. Error bars represent mean values ± standard deviation at P < 0.05, n = 9.

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

and proliferation [3]. Peptide gel FEFEFKFK is well defined and reproducible with tunable mechanical properties (G0 values ranging from 10 Pa to 35 kPa depending on the concentration used and the pH of the gel). 3.4. 2-D cell morphology and viability The gelation dynamics and mechanical properties of the gel can be controlled through the concentration and processing. The hydrogel produced at a peptide concentration of 20 mg ml–1 was subsequently used to culture bovine chondrocytes for 3 weeks. The viability of chondrocytes seeded on top of the gel surface was visually analysed using Live/Dead staining. It is well documented that chondrocytes in 2-D culture tend to adopt a fibroblast-like morphology due to dedifferentiation, but when cells are encapsulated within a 3-D scaffold they behave like chondrocytes and maintain their classic rounded morphology [2,3,25]. The images in Fig. 2A show that the majority of cells attached to tissue culture polystyrene (TCP) as a control demonstrated a highly stretched fibroblastic morphology throughout the 2 weeks culture period. In contrast, the images reveal that on day 1 chondrocytes attached to the gel surfaces retained their classic rounded morphology. However, between days 7 and 14 some chondrocytes attached to the gel surfaces adopted a slightly stretched and spindlelike morphology. Collectively the results confirmed that the scaffold design developed encouraged chondrocyte attachment with a majority of living cells present. 3.5. 2-D cell proliferation and ECM formation Successful tissue engineered scaffolds are not only cytocompatible and supportive of cell adhesion but might also promote cell proliferation [3,26]. A LDH assay and a cell count method using a haemocytometer were performed for up to 3 weeks to quantify the cell proliferation of chondrocytes. The LDH assay graphs (Fig. 2B) show an increase in cell numbers immediately after cell seeding on both TCP and the hydrogel. These findings reveal that the gels supported a 5-fold increase in cell numbers compared with TCP, which revealed a 9-fold increase in cell numbers between day 1 and 21 of culture. The slight decrease in cell numbers on day 21 on TCP may have been caused by accidental removal of cells during the medium change step or cell death due to over-confluence. The cell count method was employed to further confirm the results obtained by LDH assay. The findings revealed a gradual increase in cell numbers on day 21; the gels supported an 2-fold increase in cell numbers compared with TCP, which revealed a 4fold increase in cell numbers between day 1 and 21 in culture (Fig. 2C). Although the LDH assay and cell count method demonstrated similar trends in cell proliferation of hydrogels in 2-D, different cell numbers were recorded. As described in Sections 2.10 and 2.11, the LDH assay measurements are based on the activity of an enzyme (LDH) released upon cell lysis, which is detected using a microplate reader. The absorbance values obtained were converted to cell numbers. However, for cell counts viable cell numbers were directly counted using a haemocytometer under an optical microscope. The two different techniques employed to calculate cell numbers may have resulted in a slight discrepancy in the values recorded. Therefore, the experiments were run three times (n = 3) to confirm that the results obtained were reproducible and accurate. Previous cell proliferation studies on encapsulated chondrocytes have demonstrated that there is a clear difference in cell morphology, proliferation and cell matrix organization in 2-D and 3-D culture [3]. ECM provides the mechanical strength and stability for cells to carry out many biochemical functions. Immunocytochemical staining was performed to investigate chondrocyte

4613

activity in terms of ECM production. It was difficult to assess the phenotype of chondrocytes that had adhered to the gel surfaces from the microscopy images obtained (Fig. 2A). In some areas the cells appeared slightly spindle-like or fibroblastic, whereas in other areas they retained their rounded morphology. Experiments were carried out to investigate whether the cells were predominantly depositing collagen type I or collagen type II. Fig. 3 shows cells positively stained for collagen type I (red) during the 35 day culture period. Chondrocytes in 2-D culture were not observed to produce collagen type II. This observation is in agreement with the work of other authors who have confirmed that chondrocytes have the ability to dedifferentiate under 2-D conditions, however, in some 3-D scaffolds cells are able to retain their classic rounded morphology and behave like chondrocytes [3]. It is important for cells to deposit collagen type II as an ECM component, a crucial criterion for the regeneration of cartilagenous tissue. 3.6. 3-D hydrogel formation and cell morphology The potential of FEFEFKFK hydrogel as a 3-D scaffold encapsulating bovine chondrocytes was investigated. The 3-D fibrous gels were able to encapsulate chondrocytes without the gel disintegrating under cell culture conditions. Confocal microscopy images (Fig. 4A) show the cell viability and distribution of chondrocytes over the total depth of the sample after day 1 in culture. The micrographs reveal that the cells were homogeneously distributed across the gel, with a majority of viable cells present on day 1. Very few dead cells were detected. The gels were highly permeable to medium thus cells at the bottom of the gel remained healthy as a nutrient supply was maintained. An optical microscopy image (Fig. 4B) confirms that living chondrocytes in the 3-D cell–gel construct retained their classic rounded morphology [17] over 25 days in culture. Furthermore, an ESEM micrograph (Fig. 4C) confirms the well-defined contour of a single cell attached to the hydrogel with an average diameter of 10–20 lm. Here it is shown that in 3-D culture the gel design developed was not only able to withstand cell culture conditions for up to 3 weeks in vitro but also provided an environment for chondrocytes to live and retain their natural phenotype. 3.7. 3-D cell viability and proliferation The cell proliferation results obtained were quantitatively examined by LDH assay and a cell count method. Initially 1  105 cells per well were incorporated in the gel. Fig. 4D shows that chondrocytes demonstrated an 6-fold increase (LDH assay) and 2-fold increase (cell count method) in viable cell numbers over 21 days in culture, suggesting active proliferation inside the hydrogel (Fig. 4D). 3.8. ECM formation in three dimensions The next challenge was to investigate whether the chondrocytes encapsulated within the gels (in their natural 3-D environment) predominantly produced collagen type II, a crucial criterion for cartilage engineering. Fig. 5A shows fluorescence microscopy images revealing the deposition of both collagen type I (red) and collagen type II (red) on day 7. The nuclei of the cells were stained blue with DAPI. It was difficult to predict which type of collagen was primarily being deposited by cells. The experiment continued for another 3 weeks. On day 25 in culture cells actively deposited collagen type II in the gel matrix, with a little collagen type I deposited in some areas. These findings suggest that during the early stages of ECM production chondrocytes in 3-D culture tend to produce both collagen type I and II. However, over time collagen type II deposition levels were predominantly visible under

4614

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

A Gel

Gel

TCP

B

C Gel

TCP

200000

800000

160000

600000 400000 200000

Gel

240000

1000000 Cell numbers

Cell numbers

1200000

TCP

120000 80000 40000 0

0 Day 1

Day 14

Day 21

Day 1

2D LDH assay

Day 7

Day 14

Day 21

2D Cell count –1

Fig. 2. Chondrocyte morphology, viability and proliferation on cell seeded 2-D hydrogels (20 mg ml ). (A) Live/dead staining results after 3 weeks culture. Green spots indicate living cells present on top of the gel matrix. Scale bars represent 50 lm. Optical microscopy images suggest some chondrocytes retain their classic rounded morphology after day 1 in culture, whereas some cells dedifferentiate into spindle-like fibroblasts with a stretched morphology. (B) Quantitative analysis by LDH assay. (C) Cell proliferation measured by the 2-D cell count method using a haemocytometer. Error bars represent mean values ± standard deviation (n = 3). Statistically significant ANOVA statistics indicate that the hydrogel surfaces have significantly different cell numbers on different days.⁄P < 0.05, ⁄⁄⁄P < 0.001.

the microscope. The results obtained are in agreement with pervious literature on ECM formation by chondrocytes in 3-D culture [2,16]. Furthermore, the collagen assay graph (Fig. 5B) shows collagen deposition in both 2-D and 3-D culture over a 3 week period. The results obtained revealed relatively low collagen formation on day 1 in both systems. Interestingly, on day 14 the total amount of collagen produced by cells in 2-D culture was greater than in 3-D culture. This may be because the rate of collagen deposition by cells seeded on top of the gels is faster than the rate of collagen deposition by cells encapsulated in 3-D gel networks. On day 21 collagen production increased tremendously in the 3-D system, whereas in the 2-D system collagen levels remained relatively constant between days 14 and 21 in culture. Octapeptide gel F1 has been shown to facilitate matrix formation in vitro. Oscillatory rheology experiments were conducted to assess the mechanical properties of the gels in response to cell proliferation and matrix deposition under 3-D cell culture conditions (37 °C in a humidified atmosphere (20% O2) with 5% CO2). Fig. 5C shows a gradual increase in mechanical properties, with G0 values increas-

ing from 31 kPa on day 1 to 0.28 MPa on day 14 to 0.77 MPa on day 21. These results demonstrate that the hydrogel is capable of producing ECM molecules in 3-D culture which directly influence the mechanical properties of these peptide-based scaffolds. Cartilage is an avascular tissue with a reduced oxygen tension (oxygen levels may range from 5% at the surface to 1% in the deep zone [27]) compared with other vascularized tissues [28]. Several researchers have investigated the effect of oxygen tension on the chondrogenesis of stem cells, e.g. mesenchymal stem cells or adipose derived adult stem cells [29]. Studies revealed that a reduced oxygen tension, from 20% to 5%, encouraged chondrogenesis in alginate [30]. Further studies will be carried out to investigate the oxygen tension within peptide FEFEFKFK hydrogels and subsequent cell responses to reduced oxygen environments under cell culture conditions. 3.9. 3-D injectable hydrogel design Previously researchers demonstrated the use of injectable hydrogels for in vivo cartilage repair [31]. As a tissue engineering

4615

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

Day 14

Day 35

Collagen type I

Collagen type II

Fig. 3. Fluorescence micrographs showing collagen type I/II staining (in red) for chondrocytes seeded on top of the hydrogel. Nuclei were stained blue with DAPI. Scale bars represent 50 lm.

A1

A2 D LDH Cell count

350000 Cell numbers

300000

B

C

250000 200000 150000 100000 50000 0 Day1

Day 14

Day 21

Fig. 4. Chondrocyte morphology, viability and proliferation in cell seeded 3-D hydrogels (30 mg ml–1). (A) Confocal microscopy z-stack image: (1) viewed along the z-axis showing viable cells in green and the nuclei of dead cells in red (2); viewed along the y-axis showing the homogeneous cell distribution in the 3-D gel matrix. (B) Optical microscopy image showing that chondrocytes retained their classic rounded morphology in the gel matrix after 25 days culture. (C) ESEM micrograph showing a single cell attached to the gel. (D) Quantitative analysis of cell proliferation measured by LDH assay and the cell count method using a haemocytometer. The graph illustrates a gradual increase in cell proliferation inside the 3-D scaffold. Error bars represent mean values ± standard deviation (n = 3). Statistically significant ANOVA statistics indicate that the hydrogels had significantly different cell numbers on different days. ⁄P < 0.05; ⁄⁄P < 0.01.

scaffold the hydrogel can either be injected into a tissue defect site by itself or with specific cell types [4,32,33]. Lee and colleagues used embryonic stem cells within a RADA16-II peptide scaffold for in vivo applications [32]. The hydrogel was injected into the myocardium of 10-week-old mice. After 28 days the peptide scaffold had created a stable 3-D microenvironment which enhanced cell attachment and migration, promoting the regeneration of smooth muscle cells [32–35].

In this preliminary study the possibility of delivering encapsulated bovine chondrocytes with a peptide solution with in situ gelation properties upon injection was investigated. We developed a suitable injectable hydrogel design for cartilage engineering (Fig. 1D). Cells suspended at a concentration of 1  105 ml–1 were mixed with a 500 ll peptide solution (containing 0.5 M NaOH to adjust to the pH to 7) loaded in a 2 ml syringe with a 23G needle. The cells were homogeneously distributed within the solution

4616

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

A

Day 7

Day 25

Collagen type I

Collagen type II

B

C

900000 800000 700000 600000 500000 400000 300000 200000 100000 0

G' (Pa)

Day 1

Day 14

Day 21

Fig. 5. 3-D hydrogels promote cell viability and proliferation. (A) Collagen I/II staining (in red) for chondrocytes encapsulated inside the hydrogel. Nuclei stained blue with DAPI. The results confirm that morphological changes to chondrocytes in 2-D and 3-D cultures lead to changes in ECM deposition. Scale bars represent 50 lm. (B) Collagen production as determined by Sircol™ assay in 2-D vs. 3-D gels. Error bars represent mean values ± standard deviation (n = 3). Statistically significant ANOVA statistics indicate that the hydrogels had significantly different levels of collagen deposition on different days at ⁄P < 0.05. The level of collagen production was measured at different time points over a period of 3 weeks. Interestingly, the total amount of collagen produced in 2-D gels was greater than that in 3-D gels. This may be because the rate of collagen deposition by cells seeded on top of 2-D gels was faster than that by cells encapsulated in 3-D gels during the first 2 weeks. (C) An increase in elastic modulus (G0 ) recorded due to ECM deposition inside the hydrogel after 3 weeks culture. Error bars represent the mean ± standard deviation (⁄⁄⁄P < 0.001, n = 3).

using a pipette. The mixture was immediately syringed into a fresh 24-well plate. To avoid bubble formation slow injection is recommended as the viscous solution undergoes shear thinning. Within 10 min an opaque rigid hydrogel formed with chondrocytes embedded within the 3-D gel matrix. Our objective at this stage was primarily to investigate the effect of shear stress on cell viability under cell culture conditions at 37 °C in a humidified atmosphere (20% O2) with 5% CO2. An optical micrograph (Fig. 6A) confirms that the chondrocytes embedded within the 3-D injectable gel retained their classic rounded morphology 1 day after injection. Live/Dead staining images (Fig. 6B) show viable cells (in green) and the nuclei of

A

B1

dead cells (in red) present on days 1 (Fig. 6B1) 14 (Fig. 6B2) after injection. A slight decrease in cell viability on day 1 is inevitable as cells are forced to disengage from the 3-D viscous environment upon injection. Once the gel sets in a 24-well plate cells reattach to the gel matrix. Washes with medium were carried out three times after incubation to facilitate cell growth. On day 14 a majority of viable cells were present (Fig. 6B2). Very few dead cells were seen. Although these results demonstrate that the injectable hydrogel design introduced may be suitable for cell delivery, further investigations are required to look into cell viability, cell proliferation and matrix deposition under long-term culture conditions.

B2

Fig. 6. Preliminary injectable hydrogel design of FEFEFKFK prepared at 30 mg ml–1 in a syringe with chondrocytes encapsulated in the gel matrix to facilitate cartilage repair. (A) Optical micrograph showing that chondrocytes retained their classic rounded morphology 1 day after injection. (B) Cell viability based on Live/dead staining shows living cells (green) and the nuclei of dead cells (red) present inside the 3-D gel matrix after 1 (B1) and 14 (B2) days injection. Scale bar represents 50 lm.

A. Mujeeb et al. / Acta Biomaterialia 9 (2013) 4609–4617

4. Conclusions Here we have reported the successful exploitation of the selfassembly properties of short ionic peptides which mimic the ECM in terms of nanofibrous architecture. The novel peptide scaffold developed is capable of encapsulating chondrocytes, maintaining cell viability and proliferation for up to 35 days in vitro without the use of growth factors. Collagen type II-rich ECM deposition was observed along with an increase in gel stiffness. Acknowledgements The authors like to thank the EPSRC for financial support and Dr. Patrick Hill for assistance with the cryo-SEM and ESEM analyses. Appendix A. Figures with essential color discrimination Certain figures in this article, particularly Figs. 1–6, are difficult to interpret in black and white. The full color images can be found in the on-line version, at http://dx.doi.org/10.1016/j.actbio.2012. 08.044. References [1] Zhang S. Fabrication of novel biomaterials through molecular self assembly. Nat Biotechnol 2003;21:1171–8. [2] Kisiday J, Jin M, Kurz B, Hung H, Semino C, Zhang S, et al. Self-assembling peptide hydrogel fosters chondrocytes extracellular matrix production and cell division: implications for cartilage tissue repair. Proc Natl Acad Sci USA 2002;99:9996–10001. [3] Jayawarna V, Richardson SM, Hirst AR, Hodson NW, Saiani A, Gough JE, et al. Introducing chemical functionality in Fmoc-peptide gels for cell culture. Acta Biomater 2009;5:934–43. [4] Haines-Butterick L, Rajagopal K, Branco M, Salick D, Rughani R, Pilarz M, et al. Controlling hydrogelation kinetics by peptide design for three-dimensional encapsulation and injectable delivery of cells. Proc Natl Acad Sci USA 2007;104:7791–6. [5] Saiani A, Mohammed A, Frielinghaus H, Collins R, Hodson N, Kielty CM, et al. Self-assembly and gelation properties of a-helix versus b-sheet forming peptides. Soft Matter 2009;5:193–202. [6] Zhang S, Marini DM, Roger DK, Hwang W. Supramolecular structure of helical ribbons self-assembled from a beta-sheet peptide. J Chem Phys 2003;118:389–97. [7] Yang Z, Gu H, Zhang Y, Wang L, Xu B. Small molecules hydrogels based on a class of anti-inflammatory agents. Chem Commun 2004:208–9. [8] Langer R. Biomaterials in drug delivery and tissue engineering: one laboratory’s experience. Acc Chem Res 2000;33:94–101. [9] Mart RJ, Osborne RD, Stevens M, Ulijn RV. Peptide-based stimuli-responsive biomaterials. Soft Matter 2006;2:822–35. [10] Jun S, Hong Y, Imamura H, Ha BY, Bechhoefer J, Chen P. Self-assembly of the ionic peptide EAK16: the effect of charge distributions on self-assembly. Biophys J 2004;87:1249–59. [11] Holmes TC, Lacalle SD, Su X, Liu G, Rich A, Zhang S. Extensive neurite outgrowth and active synapse formation on self assembling peptide scaffolds. Proc Natl Acad Sci USA 2000;97:6728–33. [12] Kretsinger JK, Haines LA, Ozbas B, Pochan DJ, Schneider JP. Cytocompatibility of self-assembled b-hairpin peptide hydrogel surfaces. Biomaterials 2005;26: 5177–86.

4617

[13] Silva GA, Czeisler C, Niece KL, Beniash E, Harrington DA, Kessler JA, et al. Selective differentiation of neural progenitor cells by high-epitope density nanofibres. Science 2004;303:1352–5. [14] Beniash E, Hartgerink JD, Storrie H, Stendahl JC, Stupp SI. Self-assembling peptide amphiphile nanofiber matrices for cell entrapment. Acta Biomater 2005;1:387–97. [15] Niece KL, Hartgerink JD, Donners JJ, Stupp SI. Self-assembly combining two bioactive peptide-amphiphile molecules into nanofibers by electrostatic attraction. J Am Chem Soc 2003;125:7146–7. [16] Jayawarna V, Smith A, Gough JE, Ulijn RV. Three-dimensional cell culture of chondrocytes on modified di-phenylalanine scaffolds. Biochem Soc Trans 2007;35:535–7. [17] Jayawarna V, Ali M, Jowitt TA, Miller AF, Saiani A, Gough JE, et al. Nanostructured hydrogels for three-dimensional cell culture through self-assembly of fluorenylmethoxycarbonyl dipeptides. Adv Mater 2006;18:611–4. [18] Mahler A, Reches M, Rechter M, Cohen S, Gazit E. Rigid, self-assembled hydrogel composed of a modified aromatic dipeptide. Adv Mater 2006;18:1365–70. [19] Boothroyd S, Saiani A, Miller AF. Formation of mixed ionic complementary peptide fibrils. Macromol Symp 2008;273:139–45. [20] Genove E, Shen C, Zhang S, Semino CE. The effect of functionalized selfassembling peptide scaffolds on human aortic endothelial cell function. Biomaterials 2005;26:3341–51. [21] Liu J, Song H, Zhang L, Xu H, Zhao X. Self-assembled peptide hydrogels as tissue engineering scaffolds for three-dimensional culture of chondrocytes in vitro. Macromol Biosci 2010;10:1164–70. [22] Cunha C, Panseri S, Villa O, Silva D, Gelain F. 3D culture of adult mouse neural stem cells within functionalized self-assembling peptide scaffolds. Int J Nanomed 2011;6:943–55. [23] Georges PC, Janmey PA. Cell type-specific response to growth on soft materials. J Appl Physiol 2005;98:1547–53. [24] Zhang S, Ellis-Behnke R, Xiaojun Z, Spirio L. PuraMatrix: self-assembling peptide nanofiber scaffolds. In: Ma P, Elisseeff J, editors. Scaffolding in tissue engineering. Boca Raton, FL: CRC Press; 2006. p. 1–31. [25] Benya PD, Shaffer JD. Dedifferentiated chondrocytes re-express the differentiated collagen phenotype when cultured in agarose gels. Cell 1982;30:215–24. [26] Zhou M, Smith AM, Das AK, Hodson NW, Collins RF, Ulijn RV, et al. Selfassembled peptide-based hydrogels as scaffolds for anchorage-dependent cells. Biomaterials 2009;30:2523–30. [27] Zhou S, Cui Z, Urban JP. Factors influencing the oxygen concentration gradient from the synovial surface of articular cartilage to the cartilage–bone interface: a modeling study. Arthritis Rheum 2004;50:3915–24. [28] Fermor B, Christensen SE, Youn I, Cernanec JM, Davies CM, Weinberg JB. Oxygen, nitric oxide and articular cartilage. Eur Cell Mater 2007;13:56–65. [29] Wang DW, Fermor B, Gimble JM, Awad HA, Guilak F. Influence of oxygen on the proliferation and metabolism of adipose derived adult stem cells. J Cell Physiol 2005;204:184–91. [30] Domm C, Steinhagen J, Freitag S, Kurz B. Influence of various alginate brands on the redifferentiation of dedifferentiated bovine articular chondrocytes in alginate bead culture under high and low oxygen tension. Tissue Eng 2004;10:1796–805. [31] Lum L, Elisseeff J. Injectable hydrogels for cartilage tissue engineering. In: Ashammakhi N, Ferretti P, editors. Topics in tissue engineering. Oulu: University of Oulu; 2003. p. 1–25. [32] Gelain F, Horii A, Zhang S. Designer self-assembling peptide scaffolds for 3D tissue cell cultures and regenerative medicine. Macromol Biosci 2007;7:544–51. [33] Chung HJ, Park TG. Self-assembled and nanostructured hydrogels for drug delivery and tissue engineering. Nano Today 2009;4:429–37. [34] Pereira RC, Scaranari M, Castagnola P, Grandizio M, Azevedo HS, Reis RL, et al. Novel injectable gel (system) as a vehicle for human articular chondrocytes in cartilage tissue regeneration. J Tissue Eng Regen Med 2009;3:97–106. [35] Zhao X, Zhang S. Fabrication of molecular materials using peptide construction motifs. Trends Biotechnol 2004;22:470–6.