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Tissue Engineered Cartilage in Unconfined Compression: Biomechanical Analysis
Available online at www.sciencedirect.com
ScienceDirect Materials Today: Proceedings 2 (2015) 355 – 364
5th International conference on Advanced Nano Materials
Tissue engineered cartilage in unconfined compression: biomechanical analysis Cátia Bandeirasa,*, António Completoa, António Ramosa, Ana Teresa Rufinob,c, Madalena Ribeirob,c, João Pinto Ferreirab and Alexandrina Ferreira Mendesb,c a Deparment of Mechanical Engineering, University of Aveiro, Campus Universitário de Santiago, 3810-193 Aveiro, Portugal Centre for Neuroscience and Cell Biology, Largo Marquês de Pombal, Universidade de Coimbra, 3004-517 Coimbra, Portugal c Faculty of Pharmacy, University of Coimbra, Pólo das Ciências da Saúde, Azinhaga de Santa Comba, 3000-548 Coimbra, Portugal b
Abstract In this work we analyzed the mechanical properties of agarose and related the changes in those properties occurring in agarosechondrocyte constructs cultured under dynamic compression (72h, 1Hz, 10%) and positive metabolic outcomes of the culture with intrinsic mechanical parameters. Outcomes such as stiffening and delayed stress relaxation of the agarose gels, increased metabolic activity and gene expression of extracellular matrix proteins are related with maximum fluid velocities of 2-15 μm.s-1, maximum shear stresses of 0.02-0.18 Pa, maximum and minimum principal strains of 0.0125-0.055 and 0.0275-0.105 respectively. © 2014 The Authors. Elsevier Ltd. All rights reserved. © 2015 Elsevier Ltd. All rights reserved. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation.
Keywords: tissue engineered cartilage, agarose, stress hardening, stress relaxation, dynamic compression, metabolic activity, aggrecan, collagen, intrinsic mechanical parameters.
1. Introduction In order to better respond to the needs of treatment of injuries and degeneration of articular cartilage, tissue engineering of cartilage has had a great recent development. Tissue engineered cartilage is obtained by seeding chondrocytes or mesenchymal stem cells on a biodegradable and biocompatible scaffold and by culturing the resulting constructs in bioreactors under biochemical and mechanical stimulation [1-4]. Agarose is a natural polymer
* Corresponding author. Tel.:+351 234 370 830; fax: +351 234 370 953. E-mail address: [email protected]
2214-7853 © 2015 Elsevier Ltd. All rights reserved. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation. doi:10.1016/j.matpr.2015.04.032
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Nomenclature ECM EYE*0 E*∞ υ k0 e0 τ
extracellular matrix compressive Young’s modulus tangent dynamic modulus equilibrium dynamic modulus Poisson’s ratio initial permeability initial void ratio relaxation time
that is commonly used for tissue engineered cartilage scaffolds since it supports the maintenance of the chondrocytic phenotype and the production of extracellular matrix (ECM) [1,4]. Several difficulties still hinder the ability to produce viable tissue engineered cartilage, such as the lack of nutrient supply to the cores of the constructs, leading to inhomogeneous composition and mechanical properties, need for long culture times, cell availability and difficulty in maintenance of production of hyaline cartilage and possible production of fibrocartilage [1-2]. To overcome the current caveats, given that chondrocytes are mechanically challenged in vivo, one of the strategies to improve and accelerate the production of tissue engineered cartilage is the application of physiological values of mechanical stimuli that occur in vivo, such as cyclic compression, hydrostatic pressure and perfusion [3-7]. The application of these regimes in vitro has promoted increased metabolic activity, gene expression and synthesis of the main components of the extracellular matrix (ECM) and faster improvement of the mechanical properties when compared with free swelling culture, as well as establishment and maintenance of the chondrocytic phenotype [3-10]. Nevertheless, an ideal loading regime to produce cartilage with biomechanical properties similar to the native tissue still has to be defined. To better characterize the mechanical environments that increase the metabolic activity and mechanical properties of the tissues in response to mechanical stimulation, computational modeling and simulation is invaluable. These models treat cartilage from the mechanical standpoint as biphasic materials with a solid and fluid phase and allow to predict mechanical outputs such as stresses, strains and fluid stresses, pressure and velocities, among other intrinsic mechanical stimuli [3,4]. This information gives insight about which spatiotemporal distributions should be imposed on the tissues in order to promote favorable metabolic and biomechanical outcomes. In this work we aim to execute a biomechanical study of tissue engineered cartilage based on agarose 3% w/v scaffolds cultured in a custom-made bioreactor that allows real-time measurements of displacements and reaction forces in the constructs and extraction of related mechanical properties. We will combine these experimental measurements with a numerical model of cyclic unconfined compression culture to replicate the culture conditions and determine the values and distributions of mechanical parameters related with the increase of metabolic activity and gene expression of extracellular matrix proteins compared with free swelling culture reported in the same experiments [11]. 2. Materials and methods 2.1. Custom-made bioreactor The bioreactor used in this study (Figure 1) was developed to allow mechanical stimulation of 2D and 3D chondrocyte-scaffold constructs. The bioreactor has the ability to impose stimuli such as compression, tension, shear, bending and hydrostatic pressure or combinations of these stimuli through a digital linear-angular actuator and a perfusion pump at different amplitudes, frequencies and periods of stimulation. The mechanical parameters of the constructs are recorded through a loading cell fixed to the construct. This system provides real-time recording of the reaction forces. When in combination with the linear and angular displacements imposed by the actuator, this system allows the mechanical characterization of the constructs during the culture period.
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Figure 1 – Custom-made bioreactor and culture chamber.
2.2. Mechanical characterization of the constructs Agarose 3% w/v constructs were subjected to uniaxial compression tests to determine the Young’s modulus in compression. A cross-head speed of 72 mm/min (1.2 mm/s) was used until a final deformation of approximately 20% of the initial height was achieved. The ramp compressive Young’s moduli were determined from the stressstrain curves (n=4). In order to account for the viscoelastic effects of the material under cyclic compression, the dynamic measurements of the reaction forces of agarose constructs without cells were used to determine relaxation constants of the material under a first-order stress relaxation Prony series curve estimated by a least-squares fitting procedure in Matlab R2012a. This relaxation behavior is described for the dynamic modulus of the sample as follows:
E * (t )
1 e1 * (1 exp( t / W ))
(1)
In (1), E * (t ) is the time-dependent dynamic modulus, e1 is the ratio between the first-order Prony series modulus and the tangent dynamic modulus, E0 * (t ) , and W is the relaxation time constant. The equilibrium dynamic modulus of the agarose 3% w/v construct was determined from the final compressive reaction force and normalized to the tangent dynamic modulus. The ratio between the equilibrium and tangent dynamic moduli was compared with the dynamic modulus calculated from the cell-seeded constructs to assess if there was any significant remodeling of the stiffness of the material throughout the culture time (n=3). Statistical significance of the differences in the dynamic moduli ratio of the seeded constructs relatively to the unseeded constructs was assessed by using a one sample Wilcoxon ranked-sum test (GraphPad Prism, v5.03).
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2.3. Numerical model The intrinsic mechanical parameters associated with the metabolic outcomes were assessed by using a finite element model implemented in Abaqus v.6.12.1 (Dassault Systèmes, France). A 2D axisymmetric slice of the 8 mm diameter x 5 mm height cylinders was modeled and meshed with 500 pore pressure-stress elements (C4AXP) after a mesh convergence test. A biphasic model was implemented with a solid neo-Hookean hyperelastic compressible phase with time-dependent viscoelasticity. The hyperelastic parameters of the model were determined by using the average Young’s modulus extracted from the mechanical tests and by fixing the Poisson’s ratio to 0.35. The ratio between the equilibrium and tangent dynamic modulus, as well as the relaxation time constant estimated for the agarose gel, were used to calibrate the time-dependent Prony series for viscoelasticity. We have determined that running the cyclic compression simulation for a period of 2h is sufficient to obtain the equilibrium mechanical parameters. Strain-dependent permeability was implemented by using a widely reported power law dependent on the void ratio of the material [12]. All the biphasic mechanical properties of the material are shown in Table 1. The intrinsic mechanical parameters that were evaluated were the fluid velocity (related to increased proteoglycans syntesis) [13], fluid shear stress (implicated on cell proliferation and collagen synthesis) [3,9,14], the maximum principal strains (possible role on the synthesis and alignment of collagen fibers) [15] and minimum principal strains (presumably related to proteoglycan synthesis and to cell death at high values) [16,17]. Table 1. Mechanical properties used to define the biphasic model of agarose 3% w/v hydrogels Parameter
Value
Reference
EY- (kPa)
34.5
This work
υ
0.35
This work
E∞/E0
0.15
This work
τ (min)
17
This work
k0 (m4.N-1.s-1)
5 x 10-12
[3]
e0
4.0
[3]
3. Results 3.1. Mechanical testing results The stress-strain curves obtained for the agarose 3% w/v gels are shown in Figure 2. All the 4 cylinders tested exhibited a non-linear response up to 20% strain, with indications of stress hardening, which is a common behavior on compression of agarose gels [18]. The Young’s modulus-strain average curve confirms this behavior, with an approximately linear increase in the Young’s modulus of the samples with the applied compressive strain, represented by the following law:
EY (kPa)
281.71H 3.83
(2)
The global average Young’s modulus is of 34.5 kPa (+/- 18.8) and, for the purposes of this simulation, the compressive modulus at 10% strain is 32 kPa. These values are within the range reported for the compressive modulus of agarose at different concentrations [8,18-21]. The normalized force relaxation curves (Figure 3) obtained for the agarose cyclic compression without cells and the average of the curves with cells essay and the fitting to the first-order Prony model show that the agarose gel has a very fast relaxation (the estimated relaxation time constant is of approximately 17 minutes) and that the equilibrium dynamic reaction forces are reached after approximately 2h of continuous stimulation. This behavior with a fast relaxation time has also been verified in stress relaxation experiments on agarose 3% gels under a constant strain [22]. The reaction forces and, therefore, the dynamic modulus of the construct, has decayed to
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approximately 12.4% of the estimated initial value. The average normalized dynamic modulus of the cell-laden agarose constructs after 72h of cyclic uniaxial compression (n=3) is 33.8% of the initial value (Figure 4), which confirms growth of extracellular tissue that reduces the magnitude of stress relaxation. However, this change of the stress relaxation behavior is not statistically significant (p=0.25, one sample Wilcoxon Signed Rank test). We have also estimated that the relaxation times of the three seeded constructs is, in average, 160 minutes (2hours and 40 minutes), however this delay in the stress relaxation behavior of agarose due to the growth of cartilaginous tissue is also not statistically significant (p=0.125).
Figure 2 – Stress-strain and compressive Young’s modulus-strain curves. Upper panel: stress-strain curves obtained from four different tests. Lower panel: Young’s modulus calculated for each stress-strain pair from the four tests (red squares) and linear regression to a strain hardening law (black line, R2=0.81).
Figure 3 – Normalized reaction force of agarose 3% w/v scaffold (without cells, black) and the average normalized reaction force of the gels with cells (blue) on 72h cyclic compression (1Hz, 10%).
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Figure 4 – Comparison of estimated parameters from the stress relaxation curves of unseeded and seeded agarose gels after 72h in unconfined compression (1Hz, 10%). Upper panel: Comparison between the ratio of dynamic moduli (equilibrium vs tangent). Lower panel: Comparison of the estimated relaxation times.
3.2. Numerical analysis of intrinsic mechanical parameters After it was confirmed on the numerical simulations that the fluid viscoelasticity alone was not enough to describe the pronounced stress relaxation seen in the experimental data, the biphasic model with a visco-hyperelastic solid phase was used to evaluate the intrinsic mechanical parameters. To calculate the hyperelastic parameters of the neo Hookean hyperelastic model we used the average Young’s modulus (34.5 kPa) and determined the distributions of the fluid velocity, fluid shear stress, maximum and minimum principal strains on the first compression maximum (0.5s) and the last maximum (right before 2h of simulation). The fluid velocity radial values and estimated fluid shear stresses are plotted in Figure 5. The velocities are in the typical range estimated on cyclic uniaxial compression studies of agarose hydrogels [4]. Taking into account the fluid velocity threshold for aggrecan gene expression simulation of 0.25 μm.s -1 previously reported, on the initial compression period (blue curve) it is estimated that velocities above this threshold occur inwards by about 1.25 mm from the periphery (31% of radial distance), while for the final equilibrium period the fluid velocity is higher than the stimulatory value only in a distance of 0.6 mm from the periphery (15% of radial distance). For shear stresses below 0.1 Pa, a linear increase of cell proliferation has been reported up to 40% higher than the static case, while for stresses between 0.1 and 0.6 Pa a constant proliferation increase of 40% has been previously estimated [8,14]. The highest proliferation regime occurs in the periphery of the initial compression for about 0.25 mm, while shear stresses below 1 mPa, with an estimated negligible impact on proliferation (below 1% estimated increase) [14], are obtained on the inner radius of 2.5 mm. For the equilibrium situation, the maximum shear stress is 14% of the initial maximum and all the values are below the higher stimulation threshold. Negligible effects of shear stress are in this case estimated for an inner radius of approximately 3 mm.
Cátia Bandeiras et al. / Materials Today: Proceedings 2 (2015) 355 – 364
Figure 5 – Radial distributions of fluid velocity (upper panel) and fluid shear stress (lower panel) at the initial (0.5s, blue) and equilibrium (2h, red) stages of cyclic compression. Dashed lines: fluid velocity of 0.25 μm.s-1, the stimulatory threshold for aggrecan synthesis reported by [13] fluid shear stress of 0.1 Pa, threshold for maximum cell proliferation stimulation by mechanical cues reported by [14].
The results for the maximum and minimum principal strains (Figure 6) may provide insights on the gene expression and synthesis of collagen and aggrecan, as well as on possible cell death induction mechanisms. We considered that positive value of maximum principal strains may induce collagen synthesis [15], while minimum principal strains (compressive) above 0.25 may induce cartilage degeneration and cell death [17]. The maximum principal strains for the initial and equilibrium loading phase is radially homogeneous (2.4% variation centerperiphery initially and 0.5% after 2h) but the absolute average value decreases 4.5-fold from the initial loading to equilibrium, preasumably inducing a lesser degree of stimulation to collagen synthesis. Regarding the minimum principal strains, the distribution is also radially homogeneous and the maximum absolute value of 0.105 for the initial stimulation is much lower than the value estimated for degradation of ECM and cell death. These values drop by about 3.8-fold in average at equilibrium.
Figure 6 – Radial distributions of maximum (upper panel) and minimum (lower panel) principal strains for the initial (0.5s, blue) and equilibrium (2h, red) cyclic compression values.
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4. Discussion With this work we aimed to evaluate the mechanical properties of agarose scaffolds for tissue engineering, with and without cells, to assess through the real time evaluation of the dynamic modulus is there was significant changes in the stress relaxation behavior of agarose through production of extracellular matrix and to estimate the distributions of intrinsic mechanical stimuli that have been related to increased cell proliferation, metabolic activity and gene expression and synthesis of proteoglycans and collagen II. In previously reported work from our experiments [11], the cyclic unconfined compression during 72h (1Hz, 10%) has increased the metabolic activity of chondrocytes by 1.35 fold, aggrecan gene expression by 1.78 fold and collagen II gene expression by 2.64 fold relatively to free swelling culture controls. With this in mind, the mechanical cues that may have led to this increase will be discussed now. As reported in previous experiments for agarose gels at different concentrations [18], the unseeded scaffolds had strain hardening under unconfined uniaxial compression, with a linear increase of the compressive Young’s modulus with the applied compressive strain. The initial average Young’s modulus of the agarose scaffolds is estimated to be of 34.5 kPa in average, which is one order of magnitude below the typical compressive modulus for the native cartilage [8]. In a study conducted using agarose 2% w/v gels for cartilage tissue engineering, Kelly et al [8] demonstrated that the verified dependence of the compressive Young’s modulus on strain for the agarose gels had disappeared after 42 days of culture of the agarose-chondrocyte constructs and the compressive modulus obtained for these tissues was approximately constant with applied strain. In this work we did not perform stress-strain tests on the agarose-chondrocyte constructs. However, it was verified by using the real-time monitoring ability of our bioreactor that the magnitude of the stress relaxation of the agarose-chondrocyte constructs dynamically loaded during 72h was reduced compared to the stress relaxation of agarose only gels, with the magnitude of the decrease of the equilibrium dynamic modulus relative to the tangent dynamic modulus being reduced in average by 2.7-fold in the seeded constructs relatively to the agarose only constructs. We have also verified that the average estimated relaxation time of the seeded constructs has increased to 160 minutes, almost 10 times higher than the relaxation time estimated for the agarose only gel (17 minutes). The combination of decreased magnitude of stress relaxation and increased relaxation time in the seeded constructs supports the hypothesis that these changes are a result of the formation of extracellular matrix than increased the stiffness of the growing tissue. One of the mechanisms that may increase the rate of remodeling of the mechanical properties of the tissue under cyclic unconfined compression is the establishment of cyclic fluid flow. Fluid flow may increase the metabolic activity of the cells and synthesis of ECM by several phenomena. First, fluid flow increases the availability of nutrients and macromolecules to the construct by advective transport [23]. Second, increased fluid flow has been correlated with increased cell proliferations and proteoglycan amd protein synthesis with a possible direct mechanosensing response by the cells, eventually due to the establishment of fluid shear stresses [4,13]. The fluid velocity values are in the range of few micrometers per second in the periphery, with higher values in the initial loading phase, and vanish to negligible values in the center of the construct. Assuming the estimated stress relaxation of the agarose gel, the maximum fluid velocities decay by approximately 10 fold in 2h and the radial distance from the periphery where the fluid velocities are above the previously determined threshold for mechanically induced aggrecan gene expression stimulation (0.25 μm.s-1) [13] decays by about one half. However, a favorable aggrecan synthesis stimulation has been induced and we can report it for maximum fluid velocities between 2-15 μm.s-1 in the experimental conditions. Regarding the derived fluid shear stress, the maximum stress obtained for the initial loading is of 0.18 Pa, while it has decayed to approximately 0.02 Pa in the equilibrium loading phase. For the initial loading, a radius of 0.25 mm is subject to fluid shear stresses above 0.1 Pa (the threshold for maximal cell proliferation estimated in [14]). We have estimated in another work that the increase in cell densities due to the establishment of these fluid shear stresses is of 2% relatively to the free swelling case [24], therefore the increase in the specific metabolic activity per cell would be the main responsible for the experimental increase in metabolic activity due to the used mechanical cues. Another mechanism that is benefitial to anabolic activity of the chondrocytes is the generation of compressive and tensile strains that also induce mechanotransduction responses for ECM synthesis, as well as to generate a tensile strain-induced collagen architecture. The maximum and minimum principal strains had a radially homogeneous distribution, as expected from the applied loading regime. Mesalatti et al [4] propose that dynamic
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compression of solid agarose-chondrocyte constructs increases the accumulation of glycosaminoglycans (GAGs), main components of the matrix proteoglycans both in the periphery and center of the constructs, therefore the accumulation of proteoglycans may also be related with the minimum principal strains. The minimum principal strains ranged from 0.105 in the initial phase and 0.0275 in the equilibrium phase. These values are much lower than the threshold for ECM degradation and cell death of 0.25 compressive strain [17], therefore we can conclude that these values are related to the increased metabolic outputs. Regarding the maximum principal strains, they range from 0.055 in the initial phase to 0.0125 in the equilibrium phase. Maximum principal stresses have been related with the synthesis and orientation of collagen fibers [15], therefore we could expect a radially homogeneous distribution of collagen. However, there has also been experimental evidence for increased collagen II accumulation in areas of higher fluid flow [4]. A hypothesis for collagen synthesis is that both maximum principal strains and fluid shear stresses may act concurrently to increase collagen II gene expression and synthesis. By combining the data of the experimental and numerical analysis, we can state that the ranges of intrinsic mechanical parameters that are related to the increased metabolic activity and gene expression of aggrecan and collagen and, consequently, the relative increase in the equilibrium dynamic modulus after 72 days of cyclic compression are of 2-15 μm.s-1 for the maximum fluid velocity, 0.02-0.18 Pa for the maximum shear stress, 0.01250.055 maximum principal strain and 0.0275-0.105 minimum principal strains. These values are in the same range of the ones reported by Mesalatti et al [4], where an increase of GAGs and collagen synthesis and accumulation after 21 days of intermittent dynamic loading (1Hz, 10%) occur, with the estimation of maximum fluid velocities of 3 μm.s-1,minimum principal strains of 0.10 and maximum pore pressures of 450 Pa. In our work, maximum pore pressures are in the range of 165-1350 Pa (data not shown). With these results, we can only conclude that the three positive metabolic outcomes are a combination of the temporal distribution of these stimuli and we cannot assess which stimulus has a bigger influence on each outcome. To derive a more precise relationship between each stimulus and metabolic outcome, experiments using several amplitudes and frequencies of dynamic compression would have to be performed (for an example, see [3]). As a final remark, we have derived a range of intrinsic mechanical stimuli that have been reported to be advantageous to cell proliferation and anabolic activity and assessed that their values, both in the initial and equilibrium phase of gel loading, are within values reported in other studies to promote positive metabolic outcomes. We have also related these findings with the real-time measurement of reaction force, where chondrocyte-seeded constructs have a lower magnitude of stress relaxation and a delay of the relaxation time relatively to unseeded agarose gels. These generic ranges may be helpful to better understand the mechanical microenvironment available to the cells in the constructs in other types of mechanical stimulation. For future work, we aim to better understand the particular impact of these intrinsic mechanical stimuli values on solute transport, cell proliferation, synthesis of ECM components and, finally, the remodeling of mechanical properties using a custom-made finite element tool for solution of the spatiotemporal distributions of the biochemical and mechanical outcomes in tissue engineered cartilage. Acknowledgements This work is funded by FEDER through “Programa Operacional de Fatores de Competitividade” – COMPETE and by national funds through FCT – Fundação para a Ciência e Tecnologia, under the strategic project PEstC/EME/UI0481/2013 and also in the scope of the following projects: FCOMP-01-0124-FEDER-010248 (PTDC/EME-PME/103578/2008), FCOMP-01-0124-FEDER-015143 (PTDC/EME-PME/111305/2009) and FCOMP-01-0124-FEDER-015191 (PTDC/EME TME/113039/2009).
References [1] L. Bian, S. L. Angione, K. W. Ng, E. G. Lima, D. Y. Williams, D. Q. Mao, G. A. Ateshian, and C. T. Hung. Osteoarthritis Cartilage 17 (2009) 677–85. [2] L. Kock, C. C. van Donkelaar, and K. Ito. Cell Tissue Res. 347 (2012) 613–27.
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[3] A. Tasci, S. J. Ferguson, and P. Büchler. J. Tissue Eng. Regen. Med. 5 (2011) 210–219. [4] T. Mesallati, C. T. Buckley, T. Nagel, and D. J. Kelly. Biomech. Model. Mechanobiol. 12 (2013) 889–99. [5] R. L. Mauck, S.L. Seyhan., G. A. Ateshian, and C. T. Hung. Ann Biomed Eng. 30 (2002) 1046-1056. [6] B. D. Elder and K. A. Athanasiou. PloS One 3(6) (2008), e2341. [7] S. P. Grogan, S. Sovani, C. Pauli, J. Chen, A. Hartmann, C. W. Colwell Jr, M. K. Lotz and D. D. D’Lima. Tissue Eng A 18 (2012) 1784–92. [8] T.-A. N. Kelly, B.L. Roach, Z.D. Weldner, C.R. Mackenzie-Smith, G.D. O’Connell, E.G. Lima, A.M. Stoker, J.L. Cook, G.A. Ateshian and C.T. Hung. J Biomech 46 (2013), 1784-1791. [9] M. T. Raimondi, G. Candiani, M. Cabras, M. Cioffi, K. Laganà, M. Moretti and R. Pietrabissa. Biorheology (2008), 471-478. [10] D. Pelaez, C. Y. Huang and H. S. Cheung, Stem Cells Dev (2009), 93-102. [11] A. T. Rufino, M. Ribeiro, J. P. Ferreira, C. Bandeiras, A. Completo, A. Ramos and A. F. Mendes, ANM2014 Proceedings (2014), Aveiro, Portugal, 382. [12] M. H. Holmes and V. C. Mow. J. Biomech 23 (1990) 1145–56. [13] M. D. Buschmann, Y.-J. Kim, M. Wong, E. Frank, E. B. Hunziker and A. J. Grodzinsk, Arch Biochem Biophys 366(1) (1999), 1-7. [14] M. M. Nava, M. T. Raimondi and R. Pietrabissa. Biomech Model Mechanobiol 12 (2013) 1169-1179. [15] W. Wilson, N. J. B. Driessen, C. C. van Donkelaar and K. Ito, Osteoarthritis Cartilage 14 (2006), 1196-1202. [16] T.-A. N. Kelly, K. W. Ng, G. A. Ateshian, and C. T. Hung 17 (2009), 73-82. [17] M. Thibault, A. R. Poole and M. D. Buschmann. J Orthop Res 20(6) (2002), 1265-1273. [18] C. T. Buckley, S. D. Thorpe, F. J. O’Brien, A. J. Robinson and D. J. Kelly, J Mech Behav Biomed Mater 2(5) (2009), 512-521. [19] R. K. Tilwani, D. L. Bader and T. T. Chowdhury, J Funct Biomater 3 (2012), 23-36. [20] K. D. Costa, M. M. Y. Ho and C. T. Hung, 2003 Summer Bioengineering Conference Proceedings, Key Biscaine, FL, USA. [21] M. D. Buschmann, Y. A. Gluzband, A. J. Grodzinsky and E. B Hunziker. J Cell Sci 108 (1995), 1497-1508. [22] J. J. Chen, D. L. Bader, D. A. Lee and M. M. Knight. 8th ICCE Proceedings (2011), Dublin, Ireland. [23] R. L. Mauck, C. T. Hung and G. A. Ateshian, J Biomech Eng 125(2003), 602-614.
ScienceDirect Materials Today: Proceedings 2 (2015) 355 – 364
5th International conference on Advanced Nano Materials
Tissue engineered cartilage in unconfined compression: biomechanical analysis Cátia Bandeirasa,*, António Completoa, António Ramosa, Ana Teresa Rufinob,c, Madalena Ribeirob,c, João Pinto Ferreirab and Alexandrina Ferreira Mendesb,c a Deparment of Mechanical Engineering, University of Aveiro, Campus Universitário de Santiago, 3810-193 Aveiro, Portugal Centre for Neuroscience and Cell Biology, Largo Marquês de Pombal, Universidade de Coimbra, 3004-517 Coimbra, Portugal c Faculty of Pharmacy, University of Coimbra, Pólo das Ciências da Saúde, Azinhaga de Santa Comba, 3000-548 Coimbra, Portugal b
Abstract In this work we analyzed the mechanical properties of agarose and related the changes in those properties occurring in agarosechondrocyte constructs cultured under dynamic compression (72h, 1Hz, 10%) and positive metabolic outcomes of the culture with intrinsic mechanical parameters. Outcomes such as stiffening and delayed stress relaxation of the agarose gels, increased metabolic activity and gene expression of extracellular matrix proteins are related with maximum fluid velocities of 2-15 μm.s-1, maximum shear stresses of 0.02-0.18 Pa, maximum and minimum principal strains of 0.0125-0.055 and 0.0275-0.105 respectively. © 2014 The Authors. Elsevier Ltd. All rights reserved. © 2015 Elsevier Ltd. All rights reserved. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation.
Keywords: tissue engineered cartilage, agarose, stress hardening, stress relaxation, dynamic compression, metabolic activity, aggrecan, collagen, intrinsic mechanical parameters.
1. Introduction In order to better respond to the needs of treatment of injuries and degeneration of articular cartilage, tissue engineering of cartilage has had a great recent development. Tissue engineered cartilage is obtained by seeding chondrocytes or mesenchymal stem cells on a biodegradable and biocompatible scaffold and by culturing the resulting constructs in bioreactors under biochemical and mechanical stimulation [1-4]. Agarose is a natural polymer
* Corresponding author. Tel.:+351 234 370 830; fax: +351 234 370 953. E-mail address: [email protected]
2214-7853 © 2015 Elsevier Ltd. All rights reserved. Selection and peer-review under responsibility of TEMA - Centre for Mechanical Technology and Automation. doi:10.1016/j.matpr.2015.04.032
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Cátia Bandeiras et al. / Materials Today: Proceedings 2 (2015) 355 – 364
Nomenclature ECM EYE*0 E*∞ υ k0 e0 τ
extracellular matrix compressive Young’s modulus tangent dynamic modulus equilibrium dynamic modulus Poisson’s ratio initial permeability initial void ratio relaxation time
that is commonly used for tissue engineered cartilage scaffolds since it supports the maintenance of the chondrocytic phenotype and the production of extracellular matrix (ECM) [1,4]. Several difficulties still hinder the ability to produce viable tissue engineered cartilage, such as the lack of nutrient supply to the cores of the constructs, leading to inhomogeneous composition and mechanical properties, need for long culture times, cell availability and difficulty in maintenance of production of hyaline cartilage and possible production of fibrocartilage [1-2]. To overcome the current caveats, given that chondrocytes are mechanically challenged in vivo, one of the strategies to improve and accelerate the production of tissue engineered cartilage is the application of physiological values of mechanical stimuli that occur in vivo, such as cyclic compression, hydrostatic pressure and perfusion [3-7]. The application of these regimes in vitro has promoted increased metabolic activity, gene expression and synthesis of the main components of the extracellular matrix (ECM) and faster improvement of the mechanical properties when compared with free swelling culture, as well as establishment and maintenance of the chondrocytic phenotype [3-10]. Nevertheless, an ideal loading regime to produce cartilage with biomechanical properties similar to the native tissue still has to be defined. To better characterize the mechanical environments that increase the metabolic activity and mechanical properties of the tissues in response to mechanical stimulation, computational modeling and simulation is invaluable. These models treat cartilage from the mechanical standpoint as biphasic materials with a solid and fluid phase and allow to predict mechanical outputs such as stresses, strains and fluid stresses, pressure and velocities, among other intrinsic mechanical stimuli [3,4]. This information gives insight about which spatiotemporal distributions should be imposed on the tissues in order to promote favorable metabolic and biomechanical outcomes. In this work we aim to execute a biomechanical study of tissue engineered cartilage based on agarose 3% w/v scaffolds cultured in a custom-made bioreactor that allows real-time measurements of displacements and reaction forces in the constructs and extraction of related mechanical properties. We will combine these experimental measurements with a numerical model of cyclic unconfined compression culture to replicate the culture conditions and determine the values and distributions of mechanical parameters related with the increase of metabolic activity and gene expression of extracellular matrix proteins compared with free swelling culture reported in the same experiments [11]. 2. Materials and methods 2.1. Custom-made bioreactor The bioreactor used in this study (Figure 1) was developed to allow mechanical stimulation of 2D and 3D chondrocyte-scaffold constructs. The bioreactor has the ability to impose stimuli such as compression, tension, shear, bending and hydrostatic pressure or combinations of these stimuli through a digital linear-angular actuator and a perfusion pump at different amplitudes, frequencies and periods of stimulation. The mechanical parameters of the constructs are recorded through a loading cell fixed to the construct. This system provides real-time recording of the reaction forces. When in combination with the linear and angular displacements imposed by the actuator, this system allows the mechanical characterization of the constructs during the culture period.
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Figure 1 – Custom-made bioreactor and culture chamber.
2.2. Mechanical characterization of the constructs Agarose 3% w/v constructs were subjected to uniaxial compression tests to determine the Young’s modulus in compression. A cross-head speed of 72 mm/min (1.2 mm/s) was used until a final deformation of approximately 20% of the initial height was achieved. The ramp compressive Young’s moduli were determined from the stressstrain curves (n=4). In order to account for the viscoelastic effects of the material under cyclic compression, the dynamic measurements of the reaction forces of agarose constructs without cells were used to determine relaxation constants of the material under a first-order stress relaxation Prony series curve estimated by a least-squares fitting procedure in Matlab R2012a. This relaxation behavior is described for the dynamic modulus of the sample as follows:
E * (t )
1 e1 * (1 exp( t / W ))
(1)
In (1), E * (t ) is the time-dependent dynamic modulus, e1 is the ratio between the first-order Prony series modulus and the tangent dynamic modulus, E0 * (t ) , and W is the relaxation time constant. The equilibrium dynamic modulus of the agarose 3% w/v construct was determined from the final compressive reaction force and normalized to the tangent dynamic modulus. The ratio between the equilibrium and tangent dynamic moduli was compared with the dynamic modulus calculated from the cell-seeded constructs to assess if there was any significant remodeling of the stiffness of the material throughout the culture time (n=3). Statistical significance of the differences in the dynamic moduli ratio of the seeded constructs relatively to the unseeded constructs was assessed by using a one sample Wilcoxon ranked-sum test (GraphPad Prism, v5.03).
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2.3. Numerical model The intrinsic mechanical parameters associated with the metabolic outcomes were assessed by using a finite element model implemented in Abaqus v.6.12.1 (Dassault Systèmes, France). A 2D axisymmetric slice of the 8 mm diameter x 5 mm height cylinders was modeled and meshed with 500 pore pressure-stress elements (C4AXP) after a mesh convergence test. A biphasic model was implemented with a solid neo-Hookean hyperelastic compressible phase with time-dependent viscoelasticity. The hyperelastic parameters of the model were determined by using the average Young’s modulus extracted from the mechanical tests and by fixing the Poisson’s ratio to 0.35. The ratio between the equilibrium and tangent dynamic modulus, as well as the relaxation time constant estimated for the agarose gel, were used to calibrate the time-dependent Prony series for viscoelasticity. We have determined that running the cyclic compression simulation for a period of 2h is sufficient to obtain the equilibrium mechanical parameters. Strain-dependent permeability was implemented by using a widely reported power law dependent on the void ratio of the material [12]. All the biphasic mechanical properties of the material are shown in Table 1. The intrinsic mechanical parameters that were evaluated were the fluid velocity (related to increased proteoglycans syntesis) [13], fluid shear stress (implicated on cell proliferation and collagen synthesis) [3,9,14], the maximum principal strains (possible role on the synthesis and alignment of collagen fibers) [15] and minimum principal strains (presumably related to proteoglycan synthesis and to cell death at high values) [16,17]. Table 1. Mechanical properties used to define the biphasic model of agarose 3% w/v hydrogels Parameter
Value
Reference
EY- (kPa)
34.5
This work
υ
0.35
This work
E∞/E0
0.15
This work
τ (min)
17
This work
k0 (m4.N-1.s-1)
5 x 10-12
[3]
e0
4.0
[3]
3. Results 3.1. Mechanical testing results The stress-strain curves obtained for the agarose 3% w/v gels are shown in Figure 2. All the 4 cylinders tested exhibited a non-linear response up to 20% strain, with indications of stress hardening, which is a common behavior on compression of agarose gels [18]. The Young’s modulus-strain average curve confirms this behavior, with an approximately linear increase in the Young’s modulus of the samples with the applied compressive strain, represented by the following law:
EY (kPa)
281.71H 3.83
(2)
The global average Young’s modulus is of 34.5 kPa (+/- 18.8) and, for the purposes of this simulation, the compressive modulus at 10% strain is 32 kPa. These values are within the range reported for the compressive modulus of agarose at different concentrations [8,18-21]. The normalized force relaxation curves (Figure 3) obtained for the agarose cyclic compression without cells and the average of the curves with cells essay and the fitting to the first-order Prony model show that the agarose gel has a very fast relaxation (the estimated relaxation time constant is of approximately 17 minutes) and that the equilibrium dynamic reaction forces are reached after approximately 2h of continuous stimulation. This behavior with a fast relaxation time has also been verified in stress relaxation experiments on agarose 3% gels under a constant strain [22]. The reaction forces and, therefore, the dynamic modulus of the construct, has decayed to
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approximately 12.4% of the estimated initial value. The average normalized dynamic modulus of the cell-laden agarose constructs after 72h of cyclic uniaxial compression (n=3) is 33.8% of the initial value (Figure 4), which confirms growth of extracellular tissue that reduces the magnitude of stress relaxation. However, this change of the stress relaxation behavior is not statistically significant (p=0.25, one sample Wilcoxon Signed Rank test). We have also estimated that the relaxation times of the three seeded constructs is, in average, 160 minutes (2hours and 40 minutes), however this delay in the stress relaxation behavior of agarose due to the growth of cartilaginous tissue is also not statistically significant (p=0.125).
Figure 2 – Stress-strain and compressive Young’s modulus-strain curves. Upper panel: stress-strain curves obtained from four different tests. Lower panel: Young’s modulus calculated for each stress-strain pair from the four tests (red squares) and linear regression to a strain hardening law (black line, R2=0.81).
Figure 3 – Normalized reaction force of agarose 3% w/v scaffold (without cells, black) and the average normalized reaction force of the gels with cells (blue) on 72h cyclic compression (1Hz, 10%).
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Figure 4 – Comparison of estimated parameters from the stress relaxation curves of unseeded and seeded agarose gels after 72h in unconfined compression (1Hz, 10%). Upper panel: Comparison between the ratio of dynamic moduli (equilibrium vs tangent). Lower panel: Comparison of the estimated relaxation times.
3.2. Numerical analysis of intrinsic mechanical parameters After it was confirmed on the numerical simulations that the fluid viscoelasticity alone was not enough to describe the pronounced stress relaxation seen in the experimental data, the biphasic model with a visco-hyperelastic solid phase was used to evaluate the intrinsic mechanical parameters. To calculate the hyperelastic parameters of the neo Hookean hyperelastic model we used the average Young’s modulus (34.5 kPa) and determined the distributions of the fluid velocity, fluid shear stress, maximum and minimum principal strains on the first compression maximum (0.5s) and the last maximum (right before 2h of simulation). The fluid velocity radial values and estimated fluid shear stresses are plotted in Figure 5. The velocities are in the typical range estimated on cyclic uniaxial compression studies of agarose hydrogels [4]. Taking into account the fluid velocity threshold for aggrecan gene expression simulation of 0.25 μm.s -1 previously reported, on the initial compression period (blue curve) it is estimated that velocities above this threshold occur inwards by about 1.25 mm from the periphery (31% of radial distance), while for the final equilibrium period the fluid velocity is higher than the stimulatory value only in a distance of 0.6 mm from the periphery (15% of radial distance). For shear stresses below 0.1 Pa, a linear increase of cell proliferation has been reported up to 40% higher than the static case, while for stresses between 0.1 and 0.6 Pa a constant proliferation increase of 40% has been previously estimated [8,14]. The highest proliferation regime occurs in the periphery of the initial compression for about 0.25 mm, while shear stresses below 1 mPa, with an estimated negligible impact on proliferation (below 1% estimated increase) [14], are obtained on the inner radius of 2.5 mm. For the equilibrium situation, the maximum shear stress is 14% of the initial maximum and all the values are below the higher stimulation threshold. Negligible effects of shear stress are in this case estimated for an inner radius of approximately 3 mm.
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Figure 5 – Radial distributions of fluid velocity (upper panel) and fluid shear stress (lower panel) at the initial (0.5s, blue) and equilibrium (2h, red) stages of cyclic compression. Dashed lines: fluid velocity of 0.25 μm.s-1, the stimulatory threshold for aggrecan synthesis reported by [13] fluid shear stress of 0.1 Pa, threshold for maximum cell proliferation stimulation by mechanical cues reported by [14].
The results for the maximum and minimum principal strains (Figure 6) may provide insights on the gene expression and synthesis of collagen and aggrecan, as well as on possible cell death induction mechanisms. We considered that positive value of maximum principal strains may induce collagen synthesis [15], while minimum principal strains (compressive) above 0.25 may induce cartilage degeneration and cell death [17]. The maximum principal strains for the initial and equilibrium loading phase is radially homogeneous (2.4% variation centerperiphery initially and 0.5% after 2h) but the absolute average value decreases 4.5-fold from the initial loading to equilibrium, preasumably inducing a lesser degree of stimulation to collagen synthesis. Regarding the minimum principal strains, the distribution is also radially homogeneous and the maximum absolute value of 0.105 for the initial stimulation is much lower than the value estimated for degradation of ECM and cell death. These values drop by about 3.8-fold in average at equilibrium.
Figure 6 – Radial distributions of maximum (upper panel) and minimum (lower panel) principal strains for the initial (0.5s, blue) and equilibrium (2h, red) cyclic compression values.
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4. Discussion With this work we aimed to evaluate the mechanical properties of agarose scaffolds for tissue engineering, with and without cells, to assess through the real time evaluation of the dynamic modulus is there was significant changes in the stress relaxation behavior of agarose through production of extracellular matrix and to estimate the distributions of intrinsic mechanical stimuli that have been related to increased cell proliferation, metabolic activity and gene expression and synthesis of proteoglycans and collagen II. In previously reported work from our experiments [11], the cyclic unconfined compression during 72h (1Hz, 10%) has increased the metabolic activity of chondrocytes by 1.35 fold, aggrecan gene expression by 1.78 fold and collagen II gene expression by 2.64 fold relatively to free swelling culture controls. With this in mind, the mechanical cues that may have led to this increase will be discussed now. As reported in previous experiments for agarose gels at different concentrations [18], the unseeded scaffolds had strain hardening under unconfined uniaxial compression, with a linear increase of the compressive Young’s modulus with the applied compressive strain. The initial average Young’s modulus of the agarose scaffolds is estimated to be of 34.5 kPa in average, which is one order of magnitude below the typical compressive modulus for the native cartilage [8]. In a study conducted using agarose 2% w/v gels for cartilage tissue engineering, Kelly et al [8] demonstrated that the verified dependence of the compressive Young’s modulus on strain for the agarose gels had disappeared after 42 days of culture of the agarose-chondrocyte constructs and the compressive modulus obtained for these tissues was approximately constant with applied strain. In this work we did not perform stress-strain tests on the agarose-chondrocyte constructs. However, it was verified by using the real-time monitoring ability of our bioreactor that the magnitude of the stress relaxation of the agarose-chondrocyte constructs dynamically loaded during 72h was reduced compared to the stress relaxation of agarose only gels, with the magnitude of the decrease of the equilibrium dynamic modulus relative to the tangent dynamic modulus being reduced in average by 2.7-fold in the seeded constructs relatively to the agarose only constructs. We have also verified that the average estimated relaxation time of the seeded constructs has increased to 160 minutes, almost 10 times higher than the relaxation time estimated for the agarose only gel (17 minutes). The combination of decreased magnitude of stress relaxation and increased relaxation time in the seeded constructs supports the hypothesis that these changes are a result of the formation of extracellular matrix than increased the stiffness of the growing tissue. One of the mechanisms that may increase the rate of remodeling of the mechanical properties of the tissue under cyclic unconfined compression is the establishment of cyclic fluid flow. Fluid flow may increase the metabolic activity of the cells and synthesis of ECM by several phenomena. First, fluid flow increases the availability of nutrients and macromolecules to the construct by advective transport [23]. Second, increased fluid flow has been correlated with increased cell proliferations and proteoglycan amd protein synthesis with a possible direct mechanosensing response by the cells, eventually due to the establishment of fluid shear stresses [4,13]. The fluid velocity values are in the range of few micrometers per second in the periphery, with higher values in the initial loading phase, and vanish to negligible values in the center of the construct. Assuming the estimated stress relaxation of the agarose gel, the maximum fluid velocities decay by approximately 10 fold in 2h and the radial distance from the periphery where the fluid velocities are above the previously determined threshold for mechanically induced aggrecan gene expression stimulation (0.25 μm.s-1) [13] decays by about one half. However, a favorable aggrecan synthesis stimulation has been induced and we can report it for maximum fluid velocities between 2-15 μm.s-1 in the experimental conditions. Regarding the derived fluid shear stress, the maximum stress obtained for the initial loading is of 0.18 Pa, while it has decayed to approximately 0.02 Pa in the equilibrium loading phase. For the initial loading, a radius of 0.25 mm is subject to fluid shear stresses above 0.1 Pa (the threshold for maximal cell proliferation estimated in [14]). We have estimated in another work that the increase in cell densities due to the establishment of these fluid shear stresses is of 2% relatively to the free swelling case [24], therefore the increase in the specific metabolic activity per cell would be the main responsible for the experimental increase in metabolic activity due to the used mechanical cues. Another mechanism that is benefitial to anabolic activity of the chondrocytes is the generation of compressive and tensile strains that also induce mechanotransduction responses for ECM synthesis, as well as to generate a tensile strain-induced collagen architecture. The maximum and minimum principal strains had a radially homogeneous distribution, as expected from the applied loading regime. Mesalatti et al [4] propose that dynamic
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compression of solid agarose-chondrocyte constructs increases the accumulation of glycosaminoglycans (GAGs), main components of the matrix proteoglycans both in the periphery and center of the constructs, therefore the accumulation of proteoglycans may also be related with the minimum principal strains. The minimum principal strains ranged from 0.105 in the initial phase and 0.0275 in the equilibrium phase. These values are much lower than the threshold for ECM degradation and cell death of 0.25 compressive strain [17], therefore we can conclude that these values are related to the increased metabolic outputs. Regarding the maximum principal strains, they range from 0.055 in the initial phase to 0.0125 in the equilibrium phase. Maximum principal stresses have been related with the synthesis and orientation of collagen fibers [15], therefore we could expect a radially homogeneous distribution of collagen. However, there has also been experimental evidence for increased collagen II accumulation in areas of higher fluid flow [4]. A hypothesis for collagen synthesis is that both maximum principal strains and fluid shear stresses may act concurrently to increase collagen II gene expression and synthesis. By combining the data of the experimental and numerical analysis, we can state that the ranges of intrinsic mechanical parameters that are related to the increased metabolic activity and gene expression of aggrecan and collagen and, consequently, the relative increase in the equilibrium dynamic modulus after 72 days of cyclic compression are of 2-15 μm.s-1 for the maximum fluid velocity, 0.02-0.18 Pa for the maximum shear stress, 0.01250.055 maximum principal strain and 0.0275-0.105 minimum principal strains. These values are in the same range of the ones reported by Mesalatti et al [4], where an increase of GAGs and collagen synthesis and accumulation after 21 days of intermittent dynamic loading (1Hz, 10%) occur, with the estimation of maximum fluid velocities of 3 μm.s-1,minimum principal strains of 0.10 and maximum pore pressures of 450 Pa. In our work, maximum pore pressures are in the range of 165-1350 Pa (data not shown). With these results, we can only conclude that the three positive metabolic outcomes are a combination of the temporal distribution of these stimuli and we cannot assess which stimulus has a bigger influence on each outcome. To derive a more precise relationship between each stimulus and metabolic outcome, experiments using several amplitudes and frequencies of dynamic compression would have to be performed (for an example, see [3]). As a final remark, we have derived a range of intrinsic mechanical stimuli that have been reported to be advantageous to cell proliferation and anabolic activity and assessed that their values, both in the initial and equilibrium phase of gel loading, are within values reported in other studies to promote positive metabolic outcomes. We have also related these findings with the real-time measurement of reaction force, where chondrocyte-seeded constructs have a lower magnitude of stress relaxation and a delay of the relaxation time relatively to unseeded agarose gels. These generic ranges may be helpful to better understand the mechanical microenvironment available to the cells in the constructs in other types of mechanical stimulation. For future work, we aim to better understand the particular impact of these intrinsic mechanical stimuli values on solute transport, cell proliferation, synthesis of ECM components and, finally, the remodeling of mechanical properties using a custom-made finite element tool for solution of the spatiotemporal distributions of the biochemical and mechanical outcomes in tissue engineered cartilage. Acknowledgements This work is funded by FEDER through “Programa Operacional de Fatores de Competitividade” – COMPETE and by national funds through FCT – Fundação para a Ciência e Tecnologia, under the strategic project PEstC/EME/UI0481/2013 and also in the scope of the following projects: FCOMP-01-0124-FEDER-010248 (PTDC/EME-PME/103578/2008), FCOMP-01-0124-FEDER-015143 (PTDC/EME-PME/111305/2009) and FCOMP-01-0124-FEDER-015191 (PTDC/EME TME/113039/2009).
References [1] L. Bian, S. L. Angione, K. W. Ng, E. G. Lima, D. Y. Williams, D. Q. Mao, G. A. Ateshian, and C. T. Hung. Osteoarthritis Cartilage 17 (2009) 677–85. [2] L. Kock, C. C. van Donkelaar, and K. Ito. Cell Tissue Res. 347 (2012) 613–27.
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[3] A. Tasci, S. J. Ferguson, and P. Büchler. J. Tissue Eng. Regen. Med. 5 (2011) 210–219. [4] T. Mesallati, C. T. Buckley, T. Nagel, and D. J. Kelly. Biomech. Model. Mechanobiol. 12 (2013) 889–99. [5] R. L. Mauck, S.L. Seyhan., G. A. Ateshian, and C. T. Hung. Ann Biomed Eng. 30 (2002) 1046-1056. [6] B. D. Elder and K. A. Athanasiou. PloS One 3(6) (2008), e2341. [7] S. P. Grogan, S. Sovani, C. Pauli, J. Chen, A. Hartmann, C. W. Colwell Jr, M. K. Lotz and D. D. D’Lima. Tissue Eng A 18 (2012) 1784–92. [8] T.-A. N. Kelly, B.L. Roach, Z.D. Weldner, C.R. Mackenzie-Smith, G.D. O’Connell, E.G. Lima, A.M. Stoker, J.L. Cook, G.A. Ateshian and C.T. Hung. J Biomech 46 (2013), 1784-1791. [9] M. T. Raimondi, G. Candiani, M. Cabras, M. Cioffi, K. Laganà, M. Moretti and R. Pietrabissa. Biorheology (2008), 471-478. [10] D. Pelaez, C. Y. Huang and H. S. Cheung, Stem Cells Dev (2009), 93-102. [11] A. T. Rufino, M. Ribeiro, J. P. Ferreira, C. Bandeiras, A. Completo, A. Ramos and A. F. Mendes, ANM2014 Proceedings (2014), Aveiro, Portugal, 382. [12] M. H. Holmes and V. C. Mow. J. Biomech 23 (1990) 1145–56. [13] M. D. Buschmann, Y.-J. Kim, M. Wong, E. Frank, E. B. Hunziker and A. J. Grodzinsk, Arch Biochem Biophys 366(1) (1999), 1-7. [14] M. M. Nava, M. T. Raimondi and R. Pietrabissa. Biomech Model Mechanobiol 12 (2013) 1169-1179. [15] W. Wilson, N. J. B. Driessen, C. C. van Donkelaar and K. Ito, Osteoarthritis Cartilage 14 (2006), 1196-1202. [16] T.-A. N. Kelly, K. W. Ng, G. A. Ateshian, and C. T. Hung 17 (2009), 73-82. [17] M. Thibault, A. R. Poole and M. D. Buschmann. J Orthop Res 20(6) (2002), 1265-1273. [18] C. T. Buckley, S. D. Thorpe, F. J. O’Brien, A. J. Robinson and D. J. Kelly, J Mech Behav Biomed Mater 2(5) (2009), 512-521. [19] R. K. Tilwani, D. L. Bader and T. T. Chowdhury, J Funct Biomater 3 (2012), 23-36. [20] K. D. Costa, M. M. Y. Ho and C. T. Hung, 2003 Summer Bioengineering Conference Proceedings, Key Biscaine, FL, USA. [21] M. D. Buschmann, Y. A. Gluzband, A. J. Grodzinsky and E. B Hunziker. J Cell Sci 108 (1995), 1497-1508. [22] J. J. Chen, D. L. Bader, D. A. Lee and M. M. Knight. 8th ICCE Proceedings (2011), Dublin, Ireland. [23] R. L. Mauck, C. T. Hung and G. A. Ateshian, J Biomech Eng 125(2003), 602-614.