Author's Accepted Manuscript
Hydrogels for lung tissue engineering: Biomechanical properties of thin collagen-elastin constructs Siobhán E Dunphy, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj
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S1751-6161(14)00097-6 http://dx.doi.org/10.1016/j.jmbbm.2014.04.005 JMBBM1131
To appear in: Journal of the Mechanical Behavior of Biomedical Materials
Received date:18 September 2013 Revised date: 7 April 2014 Accepted date: 9 April 2014 Cite this article as: Siobhán E Dunphy, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj, Hydrogels for lung tissue engineering: Biomechanical properties of thin collagen-elastin constructs, Journal of the Mechanical Behavior of Biomedical Materials, http://dx.doi.org/10.1016/j. jmbbm.2014.04.005 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting galley proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Hydrogels for lung tissue engineering: biomechanical properties of thin collagen-elastin constructs
Siobhán E Dunphy*, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj
Institute for Science and Technology in Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK *
Author to whom correspondence should be addressed
Contact information: Miss Siobhán E Dunphy (corresponding author), Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, [email protected]
Miss Jessica Bratt, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, [email protected]
Prof Alicia J. El Haj, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 554605, Fax: +44 (0) 1782 747319, [email protected]
Dr Khondoker M Akram, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 674388, [email protected]
Dr Nicolas R Forsyth, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 674388, [email protected]
Abstract In this study, collagen-elastin constructs were prepared with the aim of producing a material capable of mimicking the mechanical properties of a single alveolar wall. Collagen has been used in a wide range of tissue engineering applications, however, due to its low mechanical properties its use is limited to non load-bearing applications without further manipulation using methods such as cross-linking or mechanical compression. Here, it was hypothesised that the addition of soluble elastin to a collagen hydrogel could improve its mechanical properties. Hydrogels made from collagen only and collagen plus varying amounts elastin were prepared. The Young’s modulus of each membrane was measured using the combination of a non-destructive indentation and a theoretical model previously described. An increase in Young’s modulus was observed with increasing concentration of elastin. The use of non-destructive indentation allowed for online monitoring of the elastic moduli of cell-seeded constructs over 8 days. The addition of lung fibroblasts into the membrane increased the stiffness of the hydrogels further and cell-seeded collagen hydrogels were found to have a stiffness equal to the theoretical value for a single alveolar wall (≈ 5 kPa). Through provision of some of the native extracellular matrix components of the lung parenchyma these scaffolds may be able to provide an initial building block toward the regeneration of new functional lung tissue.
Keywords Collagen; Elastin; Hydrogels; Biomechanics; Lung tissue engineering
1. Introduction Lung disease is a leading cause of morbidity and mortality worldwide. Some of the main diseases include chronic obstructive pulmonary disease (COPD), cystic fibrosis (PH), pulmonary hypertension (PH), and lung cancer. At present, the only treatment with positive outcomes is a cadaveric lung transplant, however, organ transplantation is hindered by worldwide shortages of available donors, the requirement of life-long immunesuppression, and limited success rates. The lung parenchyma presents an interface between air in the lungs and blood in the cardiovascular system providing a suitable surface for efficient gas exchange. Air is inhaled through the mouth, which then flows through the trachea, into by the bronchioles, finally reaching the smallest functional units of the lung, the alveoli. This is the site of gas exchange, which comprises a monolayer of type I and type II pneumocytes, a thin layer extracellular matrix (ECM), and depending on location, the interstitium or endothelial layer of a blood vessel. Parenchymal ECM is mainly composed mechanically dominant type I collagen as well as type III collagen, which provide structural integrity [1]. Other major constituents of pulmonary ECM are elastin and proteoglycans. Elastin plays a pivotal role in the mechanical function of lung tissue. It has been shown the macroscopic elastic and dissipative properties of alveolar tissue are dominated by both collagen and elastin [2,3]. Moreover, elastin was demonstrated to be the most important factor in determining recoil for small lung volumes while as the volume increases collagen starts to take over [4]. As such, both collagen and elastin are of significant interest due to their complementary roles in the biomechanical behaviour of healthy lung tissue.
This study presents a captivating introduction to the idea of a tissue-engineered lung structure in the clinical context. The distal lung is thought to contain several totipotent stem cell populations [5-7] and stem cell therapies to treat lung disease have shown promise, however, these therapies have been hampered by low cell engraftment [8]. Thus, a material capable of supporting and transporting these stem cell populations to the lung would be of significant benefit. A number of biodegradable materials have been proposed for tissue engineering applications [9]. The requirements for a material suitable to build a tissue-engineered lung include biocompatibility, biodegradability, porosity, and mechanical integrity, preferably equalling that of native, healthy tissue. Porosity is crucial in order to allow diffusion between the air and blood, a critical function of the alveolus. The material should provide initial support and functionality to the damaged area but would eventually be remodelled by resident or transplanted cells leaving behind new, functional lung tissue.
The mechanical behaviour of the single alveolar components is difficult to assess with lab-based measurement techniques due the scale and overall complexity of the tissue. Instead, a number of computational models of the respiratory system have been used to assess the mechanical properties of different lung components [10-12]. Such mathematical models of the complex lung system have provided important insight into mechanisms of the various pulmonary components and how they contribute to its vital mechanical behaviour. Of particular interest, Cavalcante and colleagues developed a hexagonal network model to study the mechanical interactions of lung parenchymal components [10]. To this end, a theoretical value for the Young’s modulus of a single alveolar wall was calculated and estimated to be ~ 5 kPa [10].
The objective of this work was to start with the smallest functional unit of the lung, a single alveolus, and produce a material capable of mimicking its mechanical and structural characteristics as a building block toward a larger, tissue-engineered lung construct. The main components of lung ECM, collagen and elastin, were chosen as a starting point. Non-destructive mechanical characterisation has previously been used to examine the effects of corneal fibroblast contraction and cross-linking on the Young’s modulus of collagen hydrogels[1316]. These indentation techniques combined with OCT observations of properties allows an analysis of the mechanical properties online over time of tissue-engineered constructs [17,18].. In this study, the same technique has been used to examine the influence of elastin on collagen hydrogel stiffness and viscoelastic behaviour. Furthermore, the effect of incorporating lung fibroblasts into the hydrogels was examined. The effect of elastin on the mechanical properties of collagen hydrogels has not, hitherto, been examined. This work serves a prerequisite toward the aim of creating a functional lung replacement, comprised the main structural components of the lung parenchyma.
2. Methods 2.1 Preparation of hydrogel constructs Two types hydrogels were prepared, collagen only, and collagen with different ratios of soluble elastin, according to a previously described protocol[13-16]. Briefly, rat-tail high concentration type-I collagen (BD Bioscience,UK) was mixed with 10 x standard concentration of Dulbecco’s modified Eagle’s medium (DMEM, ICN Biomedicals, UK), sodium hydroxide, and distilled water. A collagen concentration of 3.5 mg/ml was used and solutions were prepared on ice. To make the collagen-elastin hydrogels the above procedure was modified by adding a solution of soluble bovine elastin (10mg/ml, Sigma Aldrich, UK) dissolved in PBS and filter-
sterilised using a 0.2 µm filter (Minisart High-Flow, Sartorius Stedim, UK). The elastin solution was added to the collagen solution before gelation to produce a hybrid material with varying collagen to elastin ratios 4:1, 2:1, and 1:1 with respective elastin concentrations of 0.9, 1.75, and 3.5 mg/ml. A collagen concentration of 3.5 mg/ml was used in all cases. Hydrogel discs were prepared by pipetting 500 µl of solution into a filter paper ring with a diameter of 20 mm. The discs were allowed to set for 1 hour at 37˚C, 5% CO2 and the newly formed hydrogels were stored in DMEM (Lonza, UK) supplemented with 10% fetal bovine serum (Sigma), 1% nonessential amino acids (Sigma Aldrich), 1% L-glutamine (Sigma Aldrich) and 1% antibiotic-antimycotic solution (Sigma Aldrich) at 37˚C, 5% CO2 overnight or until required for testing.
2.2 Culture of human lung fibroblasts Human tissue was used with approval by the local ethics research committee in accordance with the tenets of the Declaration of Helsinki following consent obtained from the donors or their relatives. Primary human lung fibroblasts were cultured in supplemented DMEM. At 80-90% confluence, cells were passaged using TrypLE™ Express (Life Technologies, UK), counted using an improved Neubauer Haemocytometer, and resuspended in cell culture medium. The cells were incorporated into the hydrogels through mixing of the cell suspension of desired concentration into the hydrogel solution prior to gelation. Hydrogels made of collagen only and with a collagen to elastin ratio of 1:1 (3.5 mg/ml per component to a total concentration of 7 mg/ml) were seeded with either a low or high cell density, 5x103 or 2.5x105 cells per hydrogel, respectively and cultured for up to 8 days at 37°C, 5% CO2. The ratio of collagen to elastin of 1:1 was chosen for cell-seeded hydrogels as this is the ratio of these components reported for human lung parenchyma [19].
2.3 Biomechanical testing of hydrogels The biomechanical properties of the hydrogels were measured using non-destructive spherical indentation, previously reported [16]. The apparatus comprised a sample holder with a spherical indenter and an image acquisition system. Briefly, the hydrogel discs were clamped between two plastic rings with an inner diameter 20mm. A polytetrafluoroethylene (PTFE) ball of mass 715 mg and 4 mm diameter acted as the spherical indenter and was placed on top of the membrane. The displacement was measured using the image acquisition system, consisting a long focal distance objective microscope (Edmund Industrial Optics, USA) with a computer-linked CCD camera (XC-ST50CE, Sony, Japan). All measurements were performed in phosphate buffered saline (Sigma Aldrich). For modulus measurements, the ball was positioned in the centre of the
hydrogel and left in place for ten minutes to allow for initial creep deformation to occur. At ten minutes, all hydrogels were assumed to have reached their saturated deformation state and the image was taken. Creep behaviour was analysed by taking images every 10 seconds after placing the ball in the centre of the hydrogel for the first minute, then every minute up to up to 10 minutes, and finally, every 5-10 minutes up to 90 minutes. The thickness of each gel, needed to perform calculations, was determined using an optical coherence tomography (OCT) system with an axial resolution of 14 µm in free space and penetration depth of 2 mm. The source of the OCT was a superluminescent diode with a central wavelength of 1310 nm and bandwidth of 52 nm. All measurements were taken in an environment that was both sterile and non-destructive.
2.4 Theoretical analysis The mechanical properties of the hydrogels were determined using a previously described mathematical model, which describes the deformation of a membrane due to the weight of a ball [20]. The amount of deformation
6w δ δ = 0.075( ) 2 + 0.78( ) , where w is the weight of the ball, h is EhR R R the thickness of the membrane, δ is the central displacement and R is the radius of the ball [20]. The
was quantified using the equation
Young’s modulus E was then be calculated by rearranging the equation. This model assumes a low ratio of thickness to radius and large deformation such that stretching dominates over bending. The equation was developed for a ball and sample with dimensional characteristics a/R = 5 and ߜ/R = 1.7, where a is the radius of the thin membrane being measured. The assumption that stretching dominates over bending 3
was validated using a dimensionless parameter defined as
λ = ⎡⎣12 (1 −ν 2 ) ⎤⎦ 2 (
wa 2 ) , where ν is the Eh 4
Poisson’s ratio. If λ > 2000 stretching will dominate over bending [21]. The effects of bending can therefore be neglected and equation 2.5.2 is validated [16]. 2.6 LIVE/DEAD® staining The viability and morphology of the lung fibroblasts within the hydrogels was assessed using the LIVE/DEAD® cell viability assay (Life Technologies). After 8 days in culture hydrogels were washed briefly with PBS and stained with a solution of 2 µM calcein AM and 4 µM EthD-1 for 40 minutes at room temperature. Samples were imaged using a laser scanning confocal microscope (UTBI90, Olympus, Japan).
2.7 Statistical analysis Six samples for each condition were tested at each time-point, for which the mean and one standard deviation were calculated. All data was analysed using Graphpad Prism® 6.0a Graphical Software (USA). Non-linear regression was used to assess linearity, where appropriate. Two-way ANOVA was used to assess significance between different groups. A confidence value of 95% (p = 0.05) was used to determine significance.
3. Results 3.1 Variation of stiffness with elastin concentration Hydrogels were left in culture medium overnight at 37°C and measurements were taken on the following day. Dramatic differences in the displacement of the ball between the different hydrogel compositions are clearly seen in the acquired images (Fig. 1). Addition of elastin into the hydrogels resulted in a corresponding increase in stiffness of the gels, shown in Fig. 2.. There was a significant increase in Young’s modulus of hydrogels with the addition of elastin for gels with a 2:1 and 1:1 ratio of collagen to elastin (p < 0.001 and p < 0.0001, respectively, two-way ANOVA with Tukey’s multiple comparisons test). A relationship between Young’s Modulus and elastin concentration was found using simple regression analysis. Although a linear relationship cannot be defined based on the sample size, it can be concluded that a correlation does exist between the modulus of elasticity and elastin concentration (non-linear regression, degrees of freedom = 22, R2 = 0.7594, y=0.5281x + 0.5366).
3.2 Creep deformation of hydrogels The time-dependent viscoelastic behaviour was analysed by measuring the central deflection of the hydrogels over time with the ball in place. Representative curves of the time-dependent deformation behaviour of each gel are shown in Fig. 3. The central deformation was found to increase over time with each hydrogel until it reached a saturated state. The collagen-only hydrogels demonstrated viscoelastic behaviour characterised by a gradual increase in deformation over time until reaching a plateau. The collagen-elastin gels showed a faster increase in deformation with time and, therefore, tended to exhibit more elastic behaviour. This was evident even at the lowest concentration of elastin and little variation in creep behaviour observed with increasing elastin concentration.
3.2 Lung fibroblasts increase the stiffness of hydrogels The method used for mechanical characterization allowed for online monitoring and non-destructive measurement that could be repeated at different time points. As such, the Young’s’ modulus of cell-seeded hydrogels was measured over 8 days in culture. The stiffness of the hydrogels increased with the addition of both elastin and lung fibroblasts separately, however, they did not provide a synergistic effect in terms of the overall Young’s modulus. No significant differences were found between the Young’s moduli of the two types
of hydrogel when seeded with lung fibroblasts (Fig. 4 a and b). There was no significant difference in stiffness between the low and high cell seeding densities.
3.4 Lung fibroblast morphology and viability To assess cell viability and examine the morphology of lung fibroblasts within the hydrogels primary lung fibroblasts were cultured for 8 days and stained with LIVE/DEAD® cell viability reagent. The presence of elastin had no adverse effect on cell viability, shown in Fig. 4. High numbers of live cells (green) were observed and very few dead cells (red) can be seen. Similar cell morphologies were seen for both type of hydrogel (Fig. 4). The fibroblasts maintained a characteristic elongated body indicating that cells were adhering completely to the hydrogel in the presence of elastin.
3.5 Cellular contraction The increase in modulus has previously been reported to rely upon contraction of cells within the hydrogels [15]. Here, however, the thickness of the hydrogels did not change between the acellular and cell-seeded hydrogels, which would indicate that the fibroblasts were not contracting within the gels (Fig. 5). This would suggest the increase in stiffness is due to the presence of the cellular cytoskeleton rather than active contractile behaviour.
3.6 Comparison of Young’s modulus of hydrogels constructs The Young’s moduli of acellular and cell-seeded hydrogels are shown in Fig. 6. It can clearly be seen that there are significant differences between the collagen and elastin hydrogels without cells. Cell-seeded collagen constructs reached a Young’s modulus value equally that of the theoretical alveolar wall value, however there was no significant difference with regards to hydrogel type between the cell-seeded constructs.
4. Discussion 4.1 Mechanical characterisation of hydrogels The use of hydrogels has become a mainstay in tissue engineering, due to their biocompatibility, ease of production, and porous diffusive properties [22]. Natural hydrogels such as collagen, fibrin, gelatin, and hyaluronic acid are being explored because they bring the added advantage of intrinsic bioactive components that promote cell adhesion, proliferation, and guided differentiation. A major enthusiasm has developed toward
the design of hydrogel materials with mechanical and viscoelastic properties that mimic tissues found in the human body, including the elasticity of load-bearing soft tissues such as the blood vessels and lung parenchyma, in the hope of producing bioartificial tissue equivalents. Therefore, the ability to measure and continuously monitor the mechanical properties of hydrogels in a non-destructive biological environment is key.
The indentation technique, developed by Liu and Ju [20] and further elaborated by Ahearne [16] can be applied to both permeable and semi-permeable membranes. It has accurate force and displacement resolution [16] and allows for online monitoring by taking non-destructive measurements that can be repeated at different time points in a sterile environment. The sample can be kept under biological surroundings i.e., fully immersed in a solution, either cell culture medium or phosphate buffered saline, and at physiological conditions, during measurements without damaging the apparatus. This is of particular importance as temperature can affect the mechanical properties of the hydrogel.
In this study, the mechanical properties of collagen-elastin hydrogels were successfully studied. Cell-seeded hydrogels were analysed in real-time without damaging the cells. An increase in stiffness resulting from the addition of elastin into the collagen hydrogels was effectively demonstrated, while the stiffness of cell-seeded hydrogels was efficiently studied over 8 days. Importantly, incorporating cells into the hydrogels meant it was possible to achieve the theoretical value for the Young’s modulus previously calculated by Calvacante and colleagues for a single alveolar wall.
4.2 Collagen-elastin composites for tissue engineering applications Extracellular matrix proteins provide unique characteristics that promote tissue structure and function in vivo. The main structural protein in most tissues is collagen and owing to this, it been studied extensively for use in tissue-engineering applications. One shortcoming, however, is the lack of mechanical strength that collagen hydrogels alone exhibit. Several strategies have been introduced to improve the stiffness of collagen hydrogels including crosslinking [23-25] and mechanical compression [26]. Another approach has been to combine collagen with other ECM proteins, which can enhance both its biomechanical and biological properties.
In the past, collagen has been combined with elastin to improve its mechanical behaviour, biological activity, and tissue specificity. An artificial connective matrix was first described using elastin solubilised peptides and
various combinations with collagens and proteoglycans for development of a vascular prosthesis [27]. Later, a dermal substitute Matriderm®, composed of porcine-derived collagen and elastin, was developed and successfully used to treat severe burns in the clinic [28,29]. The presence of elastin was essential for successful integration including accelerated vascularisation and reduction in wound contraction [28]. Membranes of soluble collagen and elastin have also been manufactured via the electrospinning technique [30,31] and porous lyophilised scaffolds of collagen, elastin, and glycosaminoglycans have been produced with adequate tensile properties (a range of stiffness values from 0.42 – 1 MPa) and good cell compatibility [32]. Finally, elastin in lyophilised porous collagen scaffolds led to increased vascularisation and stimulated the production of both elastin and collagen fibres by host cells in a preclinical model [33].
Here, the addition of soluble elastin to collagen hydrogels resulted in an increase in the measured stiffness. The stiffness of a material has been shown to have an influence on crucial cellular functions including proliferation and differentiation [34,35]. Furthermore, the viscoelastic properties of hydrogels have been shown to enhance cell differentiation [35]. Changes in ECM stiffness occur during development and similarly, after injury or disease. These physical conditions may influence optimal progenitor and stem cell differentiation. Mimicking the physical environment that occurs during natural tissue progression may be useful in aiding stem cell differentiation for tissue engineering application or in the development of in vitro model systems to study various disease states. Conditions such as emphysema and COPD result in thickening or the airways due to fibrotic tissue and a reduction in elastin fibres, which causes dramatic changes in mechanical properties and recoil ability of the tissue [36]. An in vitro model that could be tailored to mimic the protein composition of disease states could be useful in studying these conditions and unravelling the various disease mechanisms.
4.3 Tissue engineering a complex organ There is a paucity of research in the area of lung tissue engineering. The complex architecture and hierarchy of the lung makes it extremely difficult to recapitulate in vitro. Current research has focused mainly on the development of acellular lung constructs to replace diseased or damaged lung tissue [37-40]. Fewer studies have evaluated the generation of lung tissue using biomaterial scaffolds. Sugihara and colleagues reported the formation of alveolus-like structures in a collagen hydrogel [41]. Similarly, a collagen-glycosaminoglycan structure was used to produce alveolar-like structures in vitro [42]. Polyester scaffolds have been explored for use in the lung [43,44]. Cortiella and colleagues used a polyglycolic acid to grow lung progenitor cells in vitro
and in vivo [43]. The cells expressed lung-specific markers but were not successful in forming alveolar structures. Although synthetic scaffolds can provide excellent mechanical support they do not provide the biological cues necessary to support the formation of tissue-specific structures without the addition of bioactive components and growth factors.
This paper has presented a bottom up approach to the idea of building a lung structure. Collagen and elastin were chosen as they are the main components of lung ECM and primarily responsible for the stretch properties of the lung. A simplified scaffolding approach, like the one presented, used in combination with appropriate cell types may be able to provide the necessary cues to allow the tissue to regenerate itself in vivo or act as an in vitro model to study the lung. Cells were found to have dominant effects the mechanical behaviour of the hydrogels and as such, the benefits of elastin are yet to be elucidated. Functional studies in the future are needed to determine whether the addition of elastin can have a positive impact on lung cell types of interest and could be useful in a lung tissue engineering applications. This is only a small first step toward thinking about a feasible scaffold design for lung tissue engineering. Much work is still needed in this arena to realise a practical lung tissue engineering solution to instruct improved outcomes for patients with fatal lung disease. The choice of scaffolding material used to construct a bioartificial lung will be of utmost importance, with the nature and composition of the material playing a fundamental role in the success of lung tissue engineering strategies. Once identified, the appropriate material could be combined with more complex microfabrication techniques such as 3D rapid-prototyping techniques and ink-jet printing to create anatomically similar alveolar and bronchiole structures. Such techniques have had resounding success in producing organ structures suitable for transplantation.
5. Conclusion This brief study has demonstrated the significant effect of elastin addition on the stiffness of collagen hydrogels. The synergistic effect of combining collagen and elastin produced a biomaterial with superior mechanical properties. Furthermore, incorporation of lung fibroblasts into the constructs resulted in a value of Young’s modulus matching the theoretical value for a single alveolar wall. Collagen-elastin constructs may have potential in future lung tissue engineering applications. Moreover, further elaboration could lead to the development of useful tools in the context of in vitro model systems of lung parenchyma for disease modelling and drug testing. Further work should include more in depth exploration of the time-dependent viscoelastic
properties of the material, examination of diffusion through the substrate, and a more detailed analysis of the material effects on cell behaviour and phenotype, for the specific pulmonary cell types of interest.
Conflict of Interest The authors declare no conflicts of interest.
Acknowledgments The Engineering and Physical Sciences Research Council (EPSRC) Doctoral Training Centre for Regenerative Medicine funded this work (grant number EP/F500491/1).
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Figure 1 Central deformation of collagen hydrogels with varying concentrations of elastin. Differences in the amount of deformation are clearly seen in images taken using the acquisition system. Scale bar = 2 mm.
Figure 2 Variation of Young’s modulus with elastin concentration. The Young’s Modulus (E) increased with elastin concentration (eln conc.) of collagen-elastin hydrogels. Error bars represent ± one standard deviation.
Figure 3 Creep deformation of collagen-elastin hydrogels. Representative curves show the creep behaviour of collagen-elastin hydrogels with (a) collagen only, (b) 4:1, (c) 2:1, and (d) 1:1 collagen to elastin ratios.
Figure 4 Change in Young’s modulus over time of hydrogels with lung fibroblasts. Young’s modulus of hydrogels measured at day 1,3,5, and 8 with lung fibroblasts seeded into the hydrogels at a (a) low and (b) high ®
seeding density. Error bars represent ± one standard deviation. Representative LIVE/DEAD images of lung fibroblasts after 8 days in culture within (c and d) collagen only and (e and f) collagen-elastin hydrogels at (c and e) low and (d and f) high cell seeding densities.
Figure 5 Change in thickness of hydrogels over time. The contraction of lung fibroblasts was monitored over 8 days in culture by measuring the thickness of each hydrogel construct using optical coherence tomography (OCT). The thickness was measured for acellular hydrogels and seeded with low and high cell densities on day 1, 3, 5, and 8. Error bars represent ± one standard deviation.
Figure 6 Comparison of Young’s moduli between different hydrogel constructs. The variation in stiffness for each type of hydrogel on day 1 is shown. Error bars represent ± one standard deviation.
Fig. 1
Fig. 2
Fig. 3
Fig. 4
Fig. 5
Fig. 6
abstract Graphical a
Highlights •
Addition of elastin to collagen hydrogels increases the Young’s modulus.
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Elastin alters the viscoelastic behaviour of collagen hydrogels.
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Lung fibroblasts further increase the stiffness of hydrogels.
Hydrogels for lung tissue engineering: Biomechanical properties of thin collagen-elastin constructs Siobhán E Dunphy, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj
www.elsevier.com/locate/jmbbm
PII: DOI: Reference:
S1751-6161(14)00097-6 http://dx.doi.org/10.1016/j.jmbbm.2014.04.005 JMBBM1131
To appear in: Journal of the Mechanical Behavior of Biomedical Materials
Received date:18 September 2013 Revised date: 7 April 2014 Accepted date: 9 April 2014 Cite this article as: Siobhán E Dunphy, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj, Hydrogels for lung tissue engineering: Biomechanical properties of thin collagen-elastin constructs, Journal of the Mechanical Behavior of Biomedical Materials, http://dx.doi.org/10.1016/j. jmbbm.2014.04.005 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting galley proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
Hydrogels for lung tissue engineering: biomechanical properties of thin collagen-elastin constructs
Siobhán E Dunphy*, Jessica Bratt, Khondoker M Akram, Nicholas R Forsyth, Alicia J El Haj
Institute for Science and Technology in Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK *
Author to whom correspondence should be addressed
Contact information: Miss Siobhán E Dunphy (corresponding author), Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, [email protected]
Miss Jessica Bratt, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, [email protected]
Prof Alicia J. El Haj, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 554605, Fax: +44 (0) 1782 747319, [email protected]
Dr Khondoker M Akram, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 674388, [email protected]
Dr Nicolas R Forsyth, Institute for Science and Technology in Medicine, School of Medicine, Keele University, Stoke-on-Trent, ST4 7QB, UK, Tel: +44 (0) 1782 674388, [email protected]
Abstract In this study, collagen-elastin constructs were prepared with the aim of producing a material capable of mimicking the mechanical properties of a single alveolar wall. Collagen has been used in a wide range of tissue engineering applications, however, due to its low mechanical properties its use is limited to non load-bearing applications without further manipulation using methods such as cross-linking or mechanical compression. Here, it was hypothesised that the addition of soluble elastin to a collagen hydrogel could improve its mechanical properties. Hydrogels made from collagen only and collagen plus varying amounts elastin were prepared. The Young’s modulus of each membrane was measured using the combination of a non-destructive indentation and a theoretical model previously described. An increase in Young’s modulus was observed with increasing concentration of elastin. The use of non-destructive indentation allowed for online monitoring of the elastic moduli of cell-seeded constructs over 8 days. The addition of lung fibroblasts into the membrane increased the stiffness of the hydrogels further and cell-seeded collagen hydrogels were found to have a stiffness equal to the theoretical value for a single alveolar wall (≈ 5 kPa). Through provision of some of the native extracellular matrix components of the lung parenchyma these scaffolds may be able to provide an initial building block toward the regeneration of new functional lung tissue.
Keywords Collagen; Elastin; Hydrogels; Biomechanics; Lung tissue engineering
1. Introduction Lung disease is a leading cause of morbidity and mortality worldwide. Some of the main diseases include chronic obstructive pulmonary disease (COPD), cystic fibrosis (PH), pulmonary hypertension (PH), and lung cancer. At present, the only treatment with positive outcomes is a cadaveric lung transplant, however, organ transplantation is hindered by worldwide shortages of available donors, the requirement of life-long immunesuppression, and limited success rates. The lung parenchyma presents an interface between air in the lungs and blood in the cardiovascular system providing a suitable surface for efficient gas exchange. Air is inhaled through the mouth, which then flows through the trachea, into by the bronchioles, finally reaching the smallest functional units of the lung, the alveoli. This is the site of gas exchange, which comprises a monolayer of type I and type II pneumocytes, a thin layer extracellular matrix (ECM), and depending on location, the interstitium or endothelial layer of a blood vessel. Parenchymal ECM is mainly composed mechanically dominant type I collagen as well as type III collagen, which provide structural integrity [1]. Other major constituents of pulmonary ECM are elastin and proteoglycans. Elastin plays a pivotal role in the mechanical function of lung tissue. It has been shown the macroscopic elastic and dissipative properties of alveolar tissue are dominated by both collagen and elastin [2,3]. Moreover, elastin was demonstrated to be the most important factor in determining recoil for small lung volumes while as the volume increases collagen starts to take over [4]. As such, both collagen and elastin are of significant interest due to their complementary roles in the biomechanical behaviour of healthy lung tissue.
This study presents a captivating introduction to the idea of a tissue-engineered lung structure in the clinical context. The distal lung is thought to contain several totipotent stem cell populations [5-7] and stem cell therapies to treat lung disease have shown promise, however, these therapies have been hampered by low cell engraftment [8]. Thus, a material capable of supporting and transporting these stem cell populations to the lung would be of significant benefit. A number of biodegradable materials have been proposed for tissue engineering applications [9]. The requirements for a material suitable to build a tissue-engineered lung include biocompatibility, biodegradability, porosity, and mechanical integrity, preferably equalling that of native, healthy tissue. Porosity is crucial in order to allow diffusion between the air and blood, a critical function of the alveolus. The material should provide initial support and functionality to the damaged area but would eventually be remodelled by resident or transplanted cells leaving behind new, functional lung tissue.
The mechanical behaviour of the single alveolar components is difficult to assess with lab-based measurement techniques due the scale and overall complexity of the tissue. Instead, a number of computational models of the respiratory system have been used to assess the mechanical properties of different lung components [10-12]. Such mathematical models of the complex lung system have provided important insight into mechanisms of the various pulmonary components and how they contribute to its vital mechanical behaviour. Of particular interest, Cavalcante and colleagues developed a hexagonal network model to study the mechanical interactions of lung parenchymal components [10]. To this end, a theoretical value for the Young’s modulus of a single alveolar wall was calculated and estimated to be ~ 5 kPa [10].
The objective of this work was to start with the smallest functional unit of the lung, a single alveolus, and produce a material capable of mimicking its mechanical and structural characteristics as a building block toward a larger, tissue-engineered lung construct. The main components of lung ECM, collagen and elastin, were chosen as a starting point. Non-destructive mechanical characterisation has previously been used to examine the effects of corneal fibroblast contraction and cross-linking on the Young’s modulus of collagen hydrogels[1316]. These indentation techniques combined with OCT observations of properties allows an analysis of the mechanical properties online over time of tissue-engineered constructs [17,18].. In this study, the same technique has been used to examine the influence of elastin on collagen hydrogel stiffness and viscoelastic behaviour. Furthermore, the effect of incorporating lung fibroblasts into the hydrogels was examined. The effect of elastin on the mechanical properties of collagen hydrogels has not, hitherto, been examined. This work serves a prerequisite toward the aim of creating a functional lung replacement, comprised the main structural components of the lung parenchyma.
2. Methods 2.1 Preparation of hydrogel constructs Two types hydrogels were prepared, collagen only, and collagen with different ratios of soluble elastin, according to a previously described protocol[13-16]. Briefly, rat-tail high concentration type-I collagen (BD Bioscience,UK) was mixed with 10 x standard concentration of Dulbecco’s modified Eagle’s medium (DMEM, ICN Biomedicals, UK), sodium hydroxide, and distilled water. A collagen concentration of 3.5 mg/ml was used and solutions were prepared on ice. To make the collagen-elastin hydrogels the above procedure was modified by adding a solution of soluble bovine elastin (10mg/ml, Sigma Aldrich, UK) dissolved in PBS and filter-
sterilised using a 0.2 µm filter (Minisart High-Flow, Sartorius Stedim, UK). The elastin solution was added to the collagen solution before gelation to produce a hybrid material with varying collagen to elastin ratios 4:1, 2:1, and 1:1 with respective elastin concentrations of 0.9, 1.75, and 3.5 mg/ml. A collagen concentration of 3.5 mg/ml was used in all cases. Hydrogel discs were prepared by pipetting 500 µl of solution into a filter paper ring with a diameter of 20 mm. The discs were allowed to set for 1 hour at 37˚C, 5% CO2 and the newly formed hydrogels were stored in DMEM (Lonza, UK) supplemented with 10% fetal bovine serum (Sigma), 1% nonessential amino acids (Sigma Aldrich), 1% L-glutamine (Sigma Aldrich) and 1% antibiotic-antimycotic solution (Sigma Aldrich) at 37˚C, 5% CO2 overnight or until required for testing.
2.2 Culture of human lung fibroblasts Human tissue was used with approval by the local ethics research committee in accordance with the tenets of the Declaration of Helsinki following consent obtained from the donors or their relatives. Primary human lung fibroblasts were cultured in supplemented DMEM. At 80-90% confluence, cells were passaged using TrypLE™ Express (Life Technologies, UK), counted using an improved Neubauer Haemocytometer, and resuspended in cell culture medium. The cells were incorporated into the hydrogels through mixing of the cell suspension of desired concentration into the hydrogel solution prior to gelation. Hydrogels made of collagen only and with a collagen to elastin ratio of 1:1 (3.5 mg/ml per component to a total concentration of 7 mg/ml) were seeded with either a low or high cell density, 5x103 or 2.5x105 cells per hydrogel, respectively and cultured for up to 8 days at 37°C, 5% CO2. The ratio of collagen to elastin of 1:1 was chosen for cell-seeded hydrogels as this is the ratio of these components reported for human lung parenchyma [19].
2.3 Biomechanical testing of hydrogels The biomechanical properties of the hydrogels were measured using non-destructive spherical indentation, previously reported [16]. The apparatus comprised a sample holder with a spherical indenter and an image acquisition system. Briefly, the hydrogel discs were clamped between two plastic rings with an inner diameter 20mm. A polytetrafluoroethylene (PTFE) ball of mass 715 mg and 4 mm diameter acted as the spherical indenter and was placed on top of the membrane. The displacement was measured using the image acquisition system, consisting a long focal distance objective microscope (Edmund Industrial Optics, USA) with a computer-linked CCD camera (XC-ST50CE, Sony, Japan). All measurements were performed in phosphate buffered saline (Sigma Aldrich). For modulus measurements, the ball was positioned in the centre of the
hydrogel and left in place for ten minutes to allow for initial creep deformation to occur. At ten minutes, all hydrogels were assumed to have reached their saturated deformation state and the image was taken. Creep behaviour was analysed by taking images every 10 seconds after placing the ball in the centre of the hydrogel for the first minute, then every minute up to up to 10 minutes, and finally, every 5-10 minutes up to 90 minutes. The thickness of each gel, needed to perform calculations, was determined using an optical coherence tomography (OCT) system with an axial resolution of 14 µm in free space and penetration depth of 2 mm. The source of the OCT was a superluminescent diode with a central wavelength of 1310 nm and bandwidth of 52 nm. All measurements were taken in an environment that was both sterile and non-destructive.
2.4 Theoretical analysis The mechanical properties of the hydrogels were determined using a previously described mathematical model, which describes the deformation of a membrane due to the weight of a ball [20]. The amount of deformation
6w δ δ = 0.075( ) 2 + 0.78( ) , where w is the weight of the ball, h is EhR R R the thickness of the membrane, δ is the central displacement and R is the radius of the ball [20]. The
was quantified using the equation
Young’s modulus E was then be calculated by rearranging the equation. This model assumes a low ratio of thickness to radius and large deformation such that stretching dominates over bending. The equation was developed for a ball and sample with dimensional characteristics a/R = 5 and ߜ/R = 1.7, where a is the radius of the thin membrane being measured. The assumption that stretching dominates over bending 3
was validated using a dimensionless parameter defined as
λ = ⎡⎣12 (1 −ν 2 ) ⎤⎦ 2 (
wa 2 ) , where ν is the Eh 4
Poisson’s ratio. If λ > 2000 stretching will dominate over bending [21]. The effects of bending can therefore be neglected and equation 2.5.2 is validated [16]. 2.6 LIVE/DEAD® staining The viability and morphology of the lung fibroblasts within the hydrogels was assessed using the LIVE/DEAD® cell viability assay (Life Technologies). After 8 days in culture hydrogels were washed briefly with PBS and stained with a solution of 2 µM calcein AM and 4 µM EthD-1 for 40 minutes at room temperature. Samples were imaged using a laser scanning confocal microscope (UTBI90, Olympus, Japan).
2.7 Statistical analysis Six samples for each condition were tested at each time-point, for which the mean and one standard deviation were calculated. All data was analysed using Graphpad Prism® 6.0a Graphical Software (USA). Non-linear regression was used to assess linearity, where appropriate. Two-way ANOVA was used to assess significance between different groups. A confidence value of 95% (p = 0.05) was used to determine significance.
3. Results 3.1 Variation of stiffness with elastin concentration Hydrogels were left in culture medium overnight at 37°C and measurements were taken on the following day. Dramatic differences in the displacement of the ball between the different hydrogel compositions are clearly seen in the acquired images (Fig. 1). Addition of elastin into the hydrogels resulted in a corresponding increase in stiffness of the gels, shown in Fig. 2.. There was a significant increase in Young’s modulus of hydrogels with the addition of elastin for gels with a 2:1 and 1:1 ratio of collagen to elastin (p < 0.001 and p < 0.0001, respectively, two-way ANOVA with Tukey’s multiple comparisons test). A relationship between Young’s Modulus and elastin concentration was found using simple regression analysis. Although a linear relationship cannot be defined based on the sample size, it can be concluded that a correlation does exist between the modulus of elasticity and elastin concentration (non-linear regression, degrees of freedom = 22, R2 = 0.7594, y=0.5281x + 0.5366).
3.2 Creep deformation of hydrogels The time-dependent viscoelastic behaviour was analysed by measuring the central deflection of the hydrogels over time with the ball in place. Representative curves of the time-dependent deformation behaviour of each gel are shown in Fig. 3. The central deformation was found to increase over time with each hydrogel until it reached a saturated state. The collagen-only hydrogels demonstrated viscoelastic behaviour characterised by a gradual increase in deformation over time until reaching a plateau. The collagen-elastin gels showed a faster increase in deformation with time and, therefore, tended to exhibit more elastic behaviour. This was evident even at the lowest concentration of elastin and little variation in creep behaviour observed with increasing elastin concentration.
3.2 Lung fibroblasts increase the stiffness of hydrogels The method used for mechanical characterization allowed for online monitoring and non-destructive measurement that could be repeated at different time points. As such, the Young’s’ modulus of cell-seeded hydrogels was measured over 8 days in culture. The stiffness of the hydrogels increased with the addition of both elastin and lung fibroblasts separately, however, they did not provide a synergistic effect in terms of the overall Young’s modulus. No significant differences were found between the Young’s moduli of the two types
of hydrogel when seeded with lung fibroblasts (Fig. 4 a and b). There was no significant difference in stiffness between the low and high cell seeding densities.
3.4 Lung fibroblast morphology and viability To assess cell viability and examine the morphology of lung fibroblasts within the hydrogels primary lung fibroblasts were cultured for 8 days and stained with LIVE/DEAD® cell viability reagent. The presence of elastin had no adverse effect on cell viability, shown in Fig. 4. High numbers of live cells (green) were observed and very few dead cells (red) can be seen. Similar cell morphologies were seen for both type of hydrogel (Fig. 4). The fibroblasts maintained a characteristic elongated body indicating that cells were adhering completely to the hydrogel in the presence of elastin.
3.5 Cellular contraction The increase in modulus has previously been reported to rely upon contraction of cells within the hydrogels [15]. Here, however, the thickness of the hydrogels did not change between the acellular and cell-seeded hydrogels, which would indicate that the fibroblasts were not contracting within the gels (Fig. 5). This would suggest the increase in stiffness is due to the presence of the cellular cytoskeleton rather than active contractile behaviour.
3.6 Comparison of Young’s modulus of hydrogels constructs The Young’s moduli of acellular and cell-seeded hydrogels are shown in Fig. 6. It can clearly be seen that there are significant differences between the collagen and elastin hydrogels without cells. Cell-seeded collagen constructs reached a Young’s modulus value equally that of the theoretical alveolar wall value, however there was no significant difference with regards to hydrogel type between the cell-seeded constructs.
4. Discussion 4.1 Mechanical characterisation of hydrogels The use of hydrogels has become a mainstay in tissue engineering, due to their biocompatibility, ease of production, and porous diffusive properties [22]. Natural hydrogels such as collagen, fibrin, gelatin, and hyaluronic acid are being explored because they bring the added advantage of intrinsic bioactive components that promote cell adhesion, proliferation, and guided differentiation. A major enthusiasm has developed toward
the design of hydrogel materials with mechanical and viscoelastic properties that mimic tissues found in the human body, including the elasticity of load-bearing soft tissues such as the blood vessels and lung parenchyma, in the hope of producing bioartificial tissue equivalents. Therefore, the ability to measure and continuously monitor the mechanical properties of hydrogels in a non-destructive biological environment is key.
The indentation technique, developed by Liu and Ju [20] and further elaborated by Ahearne [16] can be applied to both permeable and semi-permeable membranes. It has accurate force and displacement resolution [16] and allows for online monitoring by taking non-destructive measurements that can be repeated at different time points in a sterile environment. The sample can be kept under biological surroundings i.e., fully immersed in a solution, either cell culture medium or phosphate buffered saline, and at physiological conditions, during measurements without damaging the apparatus. This is of particular importance as temperature can affect the mechanical properties of the hydrogel.
In this study, the mechanical properties of collagen-elastin hydrogels were successfully studied. Cell-seeded hydrogels were analysed in real-time without damaging the cells. An increase in stiffness resulting from the addition of elastin into the collagen hydrogels was effectively demonstrated, while the stiffness of cell-seeded hydrogels was efficiently studied over 8 days. Importantly, incorporating cells into the hydrogels meant it was possible to achieve the theoretical value for the Young’s modulus previously calculated by Calvacante and colleagues for a single alveolar wall.
4.2 Collagen-elastin composites for tissue engineering applications Extracellular matrix proteins provide unique characteristics that promote tissue structure and function in vivo. The main structural protein in most tissues is collagen and owing to this, it been studied extensively for use in tissue-engineering applications. One shortcoming, however, is the lack of mechanical strength that collagen hydrogels alone exhibit. Several strategies have been introduced to improve the stiffness of collagen hydrogels including crosslinking [23-25] and mechanical compression [26]. Another approach has been to combine collagen with other ECM proteins, which can enhance both its biomechanical and biological properties.
In the past, collagen has been combined with elastin to improve its mechanical behaviour, biological activity, and tissue specificity. An artificial connective matrix was first described using elastin solubilised peptides and
various combinations with collagens and proteoglycans for development of a vascular prosthesis [27]. Later, a dermal substitute Matriderm®, composed of porcine-derived collagen and elastin, was developed and successfully used to treat severe burns in the clinic [28,29]. The presence of elastin was essential for successful integration including accelerated vascularisation and reduction in wound contraction [28]. Membranes of soluble collagen and elastin have also been manufactured via the electrospinning technique [30,31] and porous lyophilised scaffolds of collagen, elastin, and glycosaminoglycans have been produced with adequate tensile properties (a range of stiffness values from 0.42 – 1 MPa) and good cell compatibility [32]. Finally, elastin in lyophilised porous collagen scaffolds led to increased vascularisation and stimulated the production of both elastin and collagen fibres by host cells in a preclinical model [33].
Here, the addition of soluble elastin to collagen hydrogels resulted in an increase in the measured stiffness. The stiffness of a material has been shown to have an influence on crucial cellular functions including proliferation and differentiation [34,35]. Furthermore, the viscoelastic properties of hydrogels have been shown to enhance cell differentiation [35]. Changes in ECM stiffness occur during development and similarly, after injury or disease. These physical conditions may influence optimal progenitor and stem cell differentiation. Mimicking the physical environment that occurs during natural tissue progression may be useful in aiding stem cell differentiation for tissue engineering application or in the development of in vitro model systems to study various disease states. Conditions such as emphysema and COPD result in thickening or the airways due to fibrotic tissue and a reduction in elastin fibres, which causes dramatic changes in mechanical properties and recoil ability of the tissue [36]. An in vitro model that could be tailored to mimic the protein composition of disease states could be useful in studying these conditions and unravelling the various disease mechanisms.
4.3 Tissue engineering a complex organ There is a paucity of research in the area of lung tissue engineering. The complex architecture and hierarchy of the lung makes it extremely difficult to recapitulate in vitro. Current research has focused mainly on the development of acellular lung constructs to replace diseased or damaged lung tissue [37-40]. Fewer studies have evaluated the generation of lung tissue using biomaterial scaffolds. Sugihara and colleagues reported the formation of alveolus-like structures in a collagen hydrogel [41]. Similarly, a collagen-glycosaminoglycan structure was used to produce alveolar-like structures in vitro [42]. Polyester scaffolds have been explored for use in the lung [43,44]. Cortiella and colleagues used a polyglycolic acid to grow lung progenitor cells in vitro
and in vivo [43]. The cells expressed lung-specific markers but were not successful in forming alveolar structures. Although synthetic scaffolds can provide excellent mechanical support they do not provide the biological cues necessary to support the formation of tissue-specific structures without the addition of bioactive components and growth factors.
This paper has presented a bottom up approach to the idea of building a lung structure. Collagen and elastin were chosen as they are the main components of lung ECM and primarily responsible for the stretch properties of the lung. A simplified scaffolding approach, like the one presented, used in combination with appropriate cell types may be able to provide the necessary cues to allow the tissue to regenerate itself in vivo or act as an in vitro model to study the lung. Cells were found to have dominant effects the mechanical behaviour of the hydrogels and as such, the benefits of elastin are yet to be elucidated. Functional studies in the future are needed to determine whether the addition of elastin can have a positive impact on lung cell types of interest and could be useful in a lung tissue engineering applications. This is only a small first step toward thinking about a feasible scaffold design for lung tissue engineering. Much work is still needed in this arena to realise a practical lung tissue engineering solution to instruct improved outcomes for patients with fatal lung disease. The choice of scaffolding material used to construct a bioartificial lung will be of utmost importance, with the nature and composition of the material playing a fundamental role in the success of lung tissue engineering strategies. Once identified, the appropriate material could be combined with more complex microfabrication techniques such as 3D rapid-prototyping techniques and ink-jet printing to create anatomically similar alveolar and bronchiole structures. Such techniques have had resounding success in producing organ structures suitable for transplantation.
5. Conclusion This brief study has demonstrated the significant effect of elastin addition on the stiffness of collagen hydrogels. The synergistic effect of combining collagen and elastin produced a biomaterial with superior mechanical properties. Furthermore, incorporation of lung fibroblasts into the constructs resulted in a value of Young’s modulus matching the theoretical value for a single alveolar wall. Collagen-elastin constructs may have potential in future lung tissue engineering applications. Moreover, further elaboration could lead to the development of useful tools in the context of in vitro model systems of lung parenchyma for disease modelling and drug testing. Further work should include more in depth exploration of the time-dependent viscoelastic
properties of the material, examination of diffusion through the substrate, and a more detailed analysis of the material effects on cell behaviour and phenotype, for the specific pulmonary cell types of interest.
Conflict of Interest The authors declare no conflicts of interest.
Acknowledgments The Engineering and Physical Sciences Research Council (EPSRC) Doctoral Training Centre for Regenerative Medicine funded this work (grant number EP/F500491/1).
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Figure 1 Central deformation of collagen hydrogels with varying concentrations of elastin. Differences in the amount of deformation are clearly seen in images taken using the acquisition system. Scale bar = 2 mm.
Figure 2 Variation of Young’s modulus with elastin concentration. The Young’s Modulus (E) increased with elastin concentration (eln conc.) of collagen-elastin hydrogels. Error bars represent ± one standard deviation.
Figure 3 Creep deformation of collagen-elastin hydrogels. Representative curves show the creep behaviour of collagen-elastin hydrogels with (a) collagen only, (b) 4:1, (c) 2:1, and (d) 1:1 collagen to elastin ratios.
Figure 4 Change in Young’s modulus over time of hydrogels with lung fibroblasts. Young’s modulus of hydrogels measured at day 1,3,5, and 8 with lung fibroblasts seeded into the hydrogels at a (a) low and (b) high ®
seeding density. Error bars represent ± one standard deviation. Representative LIVE/DEAD images of lung fibroblasts after 8 days in culture within (c and d) collagen only and (e and f) collagen-elastin hydrogels at (c and e) low and (d and f) high cell seeding densities.
Figure 5 Change in thickness of hydrogels over time. The contraction of lung fibroblasts was monitored over 8 days in culture by measuring the thickness of each hydrogel construct using optical coherence tomography (OCT). The thickness was measured for acellular hydrogels and seeded with low and high cell densities on day 1, 3, 5, and 8. Error bars represent ± one standard deviation.
Figure 6 Comparison of Young’s moduli between different hydrogel constructs. The variation in stiffness for each type of hydrogel on day 1 is shown. Error bars represent ± one standard deviation.
Fig. 1
Fig. 2
Fig. 3
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Fig. 6
abstract Graphical a
Highlights •
Addition of elastin to collagen hydrogels increases the Young’s modulus.
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Elastin alters the viscoelastic behaviour of collagen hydrogels.
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Lung fibroblasts further increase the stiffness of hydrogels.
