Biomaterials 30 (2009) 1–7
Contents lists available at ScienceDirect
Biomaterials
journal homepage: www.elsevier.com/locate/biomaterials
The fabrication of elastin-based hydrogels using high pressure CO2
Nasim Annabi a, Suzanne M. Mithieux b, Anthony S. Weiss b, Fariba Dehghani a, *
a
b
School of Chemical and Biomolecular Engineering, University of Sydney, Sydney, NSW 2006, Australia
School of Molecular and Microbial Biosciences, University of Sydney, Sydney, NSW 2006, Australia
a r t i c l e i n f o
a b s t r a c t
Article history:
Received 13 July 2008
Accepted 2 September 2008
Available online 8 October 2008
The aim of this study was to investigate the effect of high pressure CO2 on the crosslinking of elastinbased polymers and the characteristics of the fabricated hydrogels. A hydrogel was fabricated by
chemically crosslinking a-elastin with glutaraldehyde at high pressure CO2. The effects of pressure,
reaction time, and crosslinker concentration on the characteristics of the fabricated hydrogels were
determined. The reaction time had negligible effect on either the swelling ratio or the pore size of the
fabricated hydrogels. Increasing the processing pressure from 30 bar to 150 bar resulted in a 60% increase
in the hydrogel swelling ratio. The crosslinked hydrogels displayed stimuli-responsive characteristics
towards temperature and salt concentration. The dense gas process facilitated coacervation, expedited
the crosslinking reaction, and dramatically affected the micro- and macrostructures of pores within the
sample. The results of micro-CT scan and SEM images demonstrated that pore interconnectivity was
substantially enhanced for a-elastin hydrogels fabricated using high pressure CO2. Dense gas CO2
reduced the wall thickness and size of the pores and importantly induced channels within the structure
of the a-elastin hydrogels. In vitro cell culture studies demonstrated that the channels facilitated fibroblast penetration and proliferation within a-elastin structures.
Ó 2008 Elsevier Ltd. All rights reserved.
Keywords:
Elastin
Hydrogel
Crosslinking
Glutaraldehyde
High pressure CO2
Fibroblast cells
1. Introduction
Tissue engineering seeks to replace damaged organs with viable
tissues, which can grow in the body. These constructs can be
cultivated in vitro and then transplanted into the patient or they can
be created in vivo by implanting a scaffold into the patient’s body
and allowing cellular infiltration. In both approaches the requirements are cells that proliferate and make extracellular matrix,
growth factors to enhance cellular growth, and scaffolds to deliver
and support healthy cell development in a defective organ [1].
Highly porous, interconnected, biocompatible scaffolds are particularly desirable for the regeneration of such organs as they allow
enhanced nutrient and oxygen transfer, cell migration and proliferation within the 3D structure [1,2]. Polymeric hydrogels are
highly attractive for tissue engineering and scaffold fabrication due
to their hydrophilicity and high permeability to oxygen and nutrients [3,4].
Elastin and elastin-like polypeptides (ELPs) have been explored
as hydrogels due to their capacity for self-assembly and phase
transition behaviour; they also mimic many features of the extracellular matrix and have the potential to guide the migration,
* Corresponding author. Tel.: þ61 2 9351 4794; fax: þ61 2 9351 2854.
E-mail address: [email protected] (F. Dehghani).
0142-9612/$ – see front matter Ó 2008 Elsevier Ltd. All rights reserved.
doi:10.1016/j.biomaterials.2008.09.031
growth and organization of cells during regeneration processes
[5,6]. Elastin is an extremely insoluble biopolymer due to its
massively polymerised state and consequently it is difficult to
process into new biomaterials. Therefore, a-elastin, an oxalic acidsolubilised derivative of elastin is often used to produce elastinbased hydrogels. Similar to the formation of elastin from its
precursor tropoelastin, a-elastin hydrogel fabrication involves two
steps: coacervation and crosslinking. a-Elastin molecules coacervate in an aqueous solution by intermolecular hydrophobic associations [7]. The protein molecules are then chemically crosslinked
to increase mechanical integrity and form hydrogels for tissue
engineering applications. Various cross-linkers such as glutaraldehyde (GA) [8–10], disuccinimidyl glutarate (DSG) [10], bis(sulfosuccinimidyl) suberate (BS3) [11–14], copper sulfate and
pyrroloquinoline quinone (PPQ) [15], ethylene glycol diglycidyl
ether (EGDE) [16], hexamethylene diisocyanate (HMDI) [17], trissuccinimidyl aminotriacetate (TSAT) [18], disuccinimidyl suberate
(DSS) [13,14], and b-[tris(hydroxymethyl) phosphino] propionic
acid (THPP) [19] have been used to crosslink genetically engineered
ELPs [9–11,13,14,17–19], tropoelastin [8,12] and a-elastin [16] in
aqueous solution or organic solvent. Deficiencies associated with
these crosslinking approaches include the persistence of small pore
sizes and inadequately communicating channels.
Dense gas carbon dioxide is particularly useful for biomaterial
processing as an alternative to organic solvent and a gas foaming
2
N. Annabi et al. / Biomaterials 30 (2009) 1–7
agent to fabricate porosity in polymeric matrices [20]. Dense gases
(DGs) are fluids with their pressure and temperature close to the
critical point. Beyond the critical point, the substance is called
a supercritical fluid (SCF) and has solvation strength and diffusivity
approaching that of liquids and gases, respectively. The most
commonly used DG is carbon dioxide because it has low critical
parameters (Tc ¼ 31.1 C, Pc ¼ 73.8 bar), is inexpensive and nontoxic [21]. Remarkable progress has been made in the application of
DGs to the processing of polymers over the last decade [22,23]. Gas
foaming technology [1,20,24–26], crosslinking reactions at high
pressure CO2 [27–30], and supercritical CO2-water emulsion techniques [31–33] have been used to fabricate porous hydrogels or
scaffolds from different kinds of polymers. However, limited
studies have been carried out on the preparation of porous
hydrophilic biomaterials using biopolymers in combination with
supercritical CO2 without using organic solvents or any other
additives such as surfactant [27,32].
In this study, the effect of dense gas CO2 on the crosslinking of
a-elastin using glutaraldehyde (GA) as a crosslinking agent was
investigated. Our recent study confirmed that high pressure CO2
did not disrupt a-elastin coacervation and had no detrimental
effect on the secondary structure of a-elastin [34]. Following
exposure to high pressure CO2 an a-elastin solution was maintained
in a coacervated state for a longer time due to the interaction
between protein and CO2 [34]. These results led us to use high
pressure CO2 to fabricate a-elastin hydrogels. The properties of
fabricated hydrogels using high pressure CO2 including swelling
ratio, pore morphology, and cell proliferation in 3D microstructure
were compared with materials fabricated at atmospheric
conditions.
2. Materials and methods
2.1. Materials
a-Elastin extracted from bovine ligament was obtained from Elastin Products Co.
(Missouri, USA). All aqueous solutions were prepared in MilliQ water. a-Elastin was
dissolved in PBS (phosphate-buffered saline) (10 mM sodium phosphate pH 7.4,
1.35 M NaCl). Glutaraldehyde (GA) was purchased from Sigma. Food grade carbon
dioxide (99.99% purity) was supplied by BOC. GM3348 cell line was obtained from
the Coriell Cell Repository. Cells were maintained in Dulbecco’s Modified Eagle’s
Medium (DMEM) supplemented with 10% v/v fetal bovine serum (FBS), penicillin
and streptomycin. All tissue culture reagents were obtained from Sigma.
100 mg/ml, respectively. At low concentrations of GA and a-elastin, thin films of
crosslinked a-elastin were formed at the bottom of the mould due to insufficient
degree of crosslinking, while at high concentrations non-homogenous hydrogels
were formed due to the high viscosity of the solution and inadequate mixing.
Consequently, in this study, 100 mg/ml of a-elastin solution and two different
crosslinker concentrations, 0.5 and 0.25% (v/v) GA, were used to produce hydrogels.
2.2.2. Dense gas in hydrogel formation
A schematic diagram of the apparatus used to fabricate a-elastin hydrogels using
dense gas CO2 is shown in Fig. 1. A syringe pump (ISCO, Model 500D) was used to
transfer CO2 into the high pressure vessel (Thar, 100 ml, view cell). The vessel
comprised of a temperature and pressure controller that allowed for temperature
adjustments and the monitoring of both variables accurately. A peristaltic pump
(MHRE 200) was used for cold water recirculation in the jacket of the syringe pump
to condense CO2 into liquid when the pump was filled with CO2.
a-Elastin solution containing the crosslinker was injected into a custom-made
Teflon mould placed inside the high pressure vessel. After the vessel was sealed and
approached thermal equilibrium at a specific temperature, the system was pressurised with CO2 to the desired level, isolated and maintained at these conditions for
a set period of time. The system was then depressurised and samples were collected.
Crosslinked structures were immediately washed repeatedly in MilliQ water, and
then placed in 100 mM Tris in PBS for 1 h. After Tris treatment, the hydrogels were
washed twice in MilliQ water and stored in PBS for further analysis.
All experiments were conducted at 37 C. The effects of pressure, reaction time,
and crosslinker concentration on the characteristics of the hydrogel were assessed.
Different concentrations of GA (i.e. 0.25 and 0.5% (v/v)) were mixed with 100 mg/ml
a-elastin and the solutions were pipetted into the mould, containing two individual
wells, placed inside the high pressure vessel. The system was then pressurized to
a desired pressure for a certain period of time. Preliminary experiments demonstrated that the reaction time had no significant effect on the properties of fabricated
hydrogels. Therefore, in this study reaction time was kept constant at 30 min. The
pressure was varied from 30 bar to 150 bar in order to investigate the effect of
pressure on the properties of fabricated hydrogels. Samples were prepared in triplicate at each condition.
2.2.3. Swelling property
The swelling property of hydrogels can be correlated to the degree of crosslinking through the hydrogel matrices. The crosslinked hydrogels were placed in
liquid nitrogen for 5 min and then lyophilised for 20 h using a freeze dryer. The
swelling properties were measured at four different conditions: 37 C and 4 C using
either PBS or MilliQ water. Under each set of conditions, at least three samples were
placed in the media overnight. The excess liquid was then removed from the swollen
samples and the swollen mass was recorded. The swelling ratio was calculated using
the following equation:
Swelling Ratio ¼
Weight of wet sample Weight of dry sample
Weight of dry sample
2.3. Environmental SEM (ESEM)
2.2. Hydrogel formation
2.2.1. Hydrogel fabrication at atmospheric pressure
a-Elastin solution was mixed with GA and the solution was immediately
pipetted into a custom-made Teflon mould. The mould was then placed at 37 C for
a period of time ranging from 15 min to 24 h, to fabricate a hydrogel. The crosslinked
hydrogels were washed in MilliQ water, then placed in 100 mM Tris((HOCH2)3CNH2)
in PBS for 1 h to inhibit further crosslinking and stored in PBS for characterisation.
A preliminary set of experiments were conducted to determine the required
amount of GA and a-elastin for the hydrogel fabrication. The concentration of
a-elastin solution was varied between 5 mg/ml and 200 mg/ml and GA between 0.1
and 2% (v/v). The solutions were pipetted into a 24-well plate and then placed at
37 C oven for 1 h. The results demonstrated in Table 1, shows that the hydrogels
were formed when the GA and a-elastin concentration were above 0.25% (v/v) and
Table 1
Optimization of protein and GA concentration for a-elastin hydrogel fabrication
a-Elastin concentration (mg/ml)
5
10
50
100
200
Environmental SEM analysis using an FEI Quanta 200 at 15 KV was performed on
wet hydrogels to characterise the structure of the fabricated hydrogels without the
potential impact of lyophilisation or coatings that are required for other microscopic
analyses.
2.4. Micro-CT analysis (3D image analysis)
Lyophilised a-elastin hydrogels were analysed using a Skyscan 1072 (Skyscan,
Belgium) high-resolution desktop X-ray CT scanner at 2.94 mm voxel resolution,
X-ray tube current 160 mA and voltage 62 KV to obtain 3D reconstructed images. The
samples were mounted vertically on a plastic support and rotated through 360
around the z-axis of the sample. 3D reconstruction of the samples was carried out
using axial bitmap images and analysed by VG Studio Max software (Volume
Graphics GmbH, Heidelberg, Germany).
2.5. Scanning electron microscopy (SEM)
GA concentration % (v/v)
0.1
0.25
0.5
1
2
–
–
–
–
–
–
–
–
U
U
–
–
–
U
–
–
–
–
U
–
–
–
–
U
–
U: Hydrogel was formed, –: no hydrogel formed at low concentrations (thin films
were formed) or non-homogenous gel at higher concentrations.
The SEM images of samples were obtained using a Philips XL30 at 15 KV to
determine the pore characteristics of the fabricated hydrogels and to examine
cellular infiltration and adhesion. Lyophilised a-elastin hydrogels were mounted on
aluminium stubs using conductive carbon paint then gold coated prior to SEM
analysis. Cell-seeded hydrogels were fixed with 2% GA in 0.1 M Na-cacodylate buffer
with 0.1 M sucrose for 1 h at 37 C. Samples underwent post-fixation with 1%
osmium in 0.1 M Na-cacodylate for 1 h and were then dehydrated in ethanol solutions at 70%, 80%, 90% and 3 times 100% for 10 min each. For drying, the samples
were immersed for 3 min in 100% hexamethyldisilazane (HMDS) then transferred to
a desiccator for 25 min to avoid water contamination. Finally they were mounted on
stubs and sputter coated with 10 nm gold.
N. Annabi et al. / Biomaterials 30 (2009) 1–7
3
Fig. 1. Experimental set-up for the fabrication of hydrogels using high pressure CO2.
2.6. In vitro cell culture
The ability of human skin fibroblast cells (GM3348) to grow into the hydrogels
3D structure was assessed. Following crosslinking, hydrogels were transferred into
a 24-well plate and washed twice with ethanol to sterilise the materials. The
hydrogels were then washed at least twice with culture media to remove any
residual ethanol and equilibrated in culture media (DMEM, 10% FBS, pen–strep) at
37 C overnight. The cells were then seeded onto the hydrogels at 3 105 cells/well.
An unseeded hydrogel was also kept in a well as a control sample. The cells were
cultured in a CO2 incubator for 2 days at 37 C, after which the hydrogels were fixed
to assess cell proliferation and infiltration using SEM analysis.
3. Result and discussion
In this study, the effect of high pressure CO2 on the crosslinking
of a-elastin hydrogels was investigated. a-Elastin was crosslinked
in the presence of high pressure CO2 using GA, an amine-reactive
chemical crosslinker. The effect of the variation of pressure and
crosslinker concentration on swelling properties and pore sizes of
fabricated hydrogels was assessed.
3.1. Visual comparison of hydrogels formed at high pressure CO2
and atmospheric condition
Visual comparison of fabricated a-elastin hydrogels demonstrated that gels produced at high pressure CO2 were more rigid
than those produced at atmospheric pressure. a-Elastin hydrogels
fabricated at high pressure CO2 were easily handled. However, the
hydrogels produced at atmospheric pressure were very fragile.
Increasing the processing time from 1 h to 24 h had negligible effect
on promoting the macrostructure of the elastin formed at
atmospheric condition. This may be due to an insufficient degree of
crosslinking throughout the hydrogel matrices produced at atmospheric condition. High pressure CO2 promoted the coacervation of
a-elastin and accelerated the crosslinking of a-elastin. No additional crosslinking was observed for the fabricated hydrogel at high
pressure CO2 and atmospheric condition after washing and storage
in PBS.
3.2. Hydrogel swelling
a-Elastin hydrogels fabricated in this study exhibited high
swelling ratios in the range of 21–35 g H2O/g protein. Typical
a-elastin hydrogels produced at high pressure CO2, before and after
swelling in water are shown in Fig. 2. Both hydrogels produced at
high pressure CO2 and atmospheric pressure displayed stimuliresponsive characteristic towards temperature and salt concentrations when they were swelled in PBS and water at 4 C and 37 C.
The swelling ratio of the hydrogels in water was greater than those
swelled in PBS at 4 C. The fabricated hydrogels swelled more in
both water and PBS at lower temperature as indicated in Table 2.
Hydrogels produced at 60 bar CO2 pressure with 0.5% (v/v) GA
absorbed 33.2 0.8 g and 28.6 1.6 g liquid/g protein at 4 C and
37 C, respectively, when they were hydrated in water. However,
they absorbed 18.3 4.6 g and 7 3.2 g liquid/g protein at 4 C and
37 C, respectively, when they were swelled in PBS. Elevated
temperature and the presence of salt in PBS resulted in a contraction of the material due to water expulsion. As exhibited in Table 2,
the swelling behaviour of the GA crosslinked a-elastin hydrogels
produced at high pressure CO2 was either considerably greater or
comparable with other elastin-based hydrogels fabricated using
various crosslinkers.
The variation of pressure and crosslinker concentration had
a substantial effect on the swelling behaviour of fabricated hydrogels. The swelling ratio of fabricated hydrogels was enhanced by
increasing the pressure and reduced by enhancing the crosslinker
concentration. As shown in Table 3, the swelling ratio of the
samples exposed to high pressure CO2 was increased from
21.4 0.7 g to 35.2 2.5 g H2O/g protein as pressure increased
from 30 bar to 150 bar at 0.5% (v/v) GA concentration. Using
a higher concentration of crosslinker resulted in a higher degree of
crosslinking through the hydrogel matrices which can contribute to
a reduction in the swelling ratio.
3.3. Pore structure of the a-elastin hydrogel
Fig. 2. Hydrogels produced at 60 bar CO2 before and after swelling in water at 4 C.
Pore size and interconnectivity are critical hydrogel properties
that influence the ability of cells to infiltrate and proliferate
4
N. Annabi et al. / Biomaterials 30 (2009) 1–7
Table 2
Swelling ratios of fabricated hydrogels using high pressure CO2 and conventional
methods
Protein
Crosslinker
Swelling ratio
(g liquid/g protein)
PBS
MilliQ water
a-Elastin produced at
GA
high pressure CO2
Natural elastin
33.2 0.8a
28.6 1.6b
–
This
study
[42]
Synthetic tropoelastin
BS3
GA
HMDI
THPP
a-Elastin
Engineered ELPs
EGDE
TSAT
63 5a
33 4b
7–80 at 25 C
0.37a
12.3 0.5a
3.7 0.2b
8.4–24
–
[12]
Synthetic tropoelastin
Engineered ELPs
Engineered ELPs
18.3 4.6a
7 3.2b
0.8 at 2 C
0.5 at 36 C
6.9 0.5a
3.8 0.8b
–
–
–
a
b
–
w2–3.5a
w0.2–0.6b
Reference
[8]
[17]
[19]
[16]
[18]
Swelling temperature: 4 C.
Swelling temperature: 37 C.
Table 3
Effect of pressure and crosslinker concentration on the swelling ratio of fabricated
hydrogels using high pressure CO2
Pressure (bar)
GA concentration (%) (v/v)
Swelling ratio (g H2O/g protein)
30
60
150
30
60
150
0.25
0.25
0.25
0.5
0.5
0.5
34.5 1.3
38.4 6
54 9.2
21.4 0.7
33.2 0.8
35.2 2.5
within 3D structures. In this study, ESEM, SEM, and Micro-CT
analysis were used to characterise the pore morphology of
fabricated a-elastin hydrogels. ESEM analysis demonstrated
that a-elastin hydrogels fabricated under high pressure CO2
(Fig. 3a) were highly porous yet robust structures. ESEM images
of a-elastin hydrogels produced under atmospheric pressure
could not be obtained because the wet constructs disintegrated
prior to examination. SEM images of a-elastin hydrogels fabricated under high pressure (Fig. 3b) were similar to those
obtained by ESEM which indicated that the pores observed in the
SEM images were formed during the hydrogel fabrication and
were not introduced into the matrix during lyophilisation or
sample preparation for SEM analysis. At the microscopic level,
comparison of SEM images of a-elastin hydrogels produced under
either high pressure CO2 or atmospheric conditions (Fig. 3c)
indicated that high pressure CO2 reduced the pore size of the
fabricated hydrogels. Equivalent circle diameter (ECD) of the
pores and the thickness of the walls between pores were calculated using Image J software. The average pore size of hydrogels
decreased from 14.3 3.3 mm to 4.9 4.3 mm as the pressure
increased from 1 bar to 60 bar. In hydrogels fabricated by the
dense gas CO2 the pore wall thickness was dramatically diminished to 0.5 0.1 mm compared to 4.3 5.5 mm for a-elastin
hydrogels fabricated under atmospheric conditions. The organisation of the matrix produced under high pressure is similar to
that seen in some natural elastin microstructures within the body
[35,36].
The 10-fold reduction in pore wall thickness of a-elastin
hydrogels fabricated at high pressure CO2 compared with the
samples produced at atmospheric pressure can be explained by
reference to the mechanism of pore formation in porous
membrane [37]. Various techniques such as freeze drying, phase
inversion, solvent casting-particulate leaching and gas foaming
have been used to produce porous membranes and scaffolds
[38]. In all these methods pores are fabricated when a homogeneous multi-component system is separated into a multi-phase
system in order to decrease the system free energy [39]. Two
phases are generally separated out; a polymer-rich phase (e.g.
a high polymer concentration phase) and a polymer-lean phase
(e.g. a low polymer concentration solution) [39]. The porous
structure is then formed when the polymer-lean phase is
Fig. 3. (a) ESEM image of a wet a-elastin hydrogel produced at 60 bar CO2 pressure. SEM images of an a-elastin hydrogel fabricated at (b) 60 bar CO2 pressure and (c) atmospheric
pressure.
N. Annabi et al. / Biomaterials 30 (2009) 1–7
5
Fig. 4. Skyscan images of a fabricated hydrogel using (a) high pressure CO2 and (b) atmospheric pressure.
removed. The physical structure of the pores is governed by the
kinetics of the phase separation in each system. It is possible to
form powders, closed-pore foams or open-pore foams with thick
or thin-fibrous wall structures depending on the rate of the
phase separation [39]. Generally, porous matrices fabricated by
using freeze drying and phase inversion techniques give rise to
non-homogenous pores with thick walls [40].
Crosslinked a-elastin hydrogels fabricated under atmospheric
conditions consisted of polyhedral shaped pores with thick walls.
However, highly interconnected pores with thin walls were
formed when a-elastin was crosslinked under high pressure CO2.
Similar pore microstructures comprising highly interconnected
pores with thin walls were formed using high pressure CO2 in
studies by Le et al. for chitosan [41]; Palocci et al. for dextran
[32]; Lee et al. for poly vinyl alcohol (PVA), blended PVA/poly
ethylene glycol (PEG) and chitosan [33]; and Partap et al. for
alginate [31]. Palocci et al. reported that thin wall-porous
matrices of dextran that were highly interconnected were formed
using CO2–water emulsion polymerisation templating, known as
high internal phase emulsion templating technique [32]. The
average pore sizes of the fabricated porous dextran were reported
in the range of 6–25.7 mm [32]. This highly porous structure was
fabricated through the formation of numerous CO2 bubbles
(polymer-lean phase) in the polymer-rich phase. As a result the
polymer rich phase was distributed over a large interfacial area
with a simultaneous thinning of the film of the continuous phase
during the crosslinking or polymerisation stage [32]. A similar
phenomenon to that reported by Palocci et al. is likely to have
occurred for the solution of a-elastin and GA at high pressure
CO2. However, unlike the CO2–water emulsion polymerisation
templating method a surfactant was not used to stabilise the
a-elastin structure or enhance the dissolution of CO2 in an
aqueous phase.
3.3.1. Effect of cross-linker concentration on hydrogel pore
morphology
The effect of GA concentration on the pore morphology of
a-elastin hydrogels fabricated at both atmospheric and high pressure
was assessed. Increasing the concentration of GA at both atmospheric and high pressures slightly increased the pore sizes. At
atmospheric pressure when the concentration of GA was increased
from 0.25 to 0.5% (v/v), the average pore size of the a-elastin
hydrogel slightly increased from 10 1.7 mm to 14.5 3.3 mm. This
effect was more noticeable in the CO2 system. The pore sizes in the
samples produced at 60 bar CO2 pressure increased 2.5-fold from
3.6 1.3 mm to 9 5.4 mm when the GA concentration was changed
from 0.25 to 0.5% (v/v).
3.3.2. Pore interconnectivity
The pore interconnectivity of hydrogels fabricated at atmospheric and high pressure CO2 was compared using micro-CT
scan analysis. As depicted in Fig. 4a, a highly homogenous
interconnected-porous structure was detected for a-elastin
hydrogels acquired using high pressure CO2. In comparison
a-elastin hydrogels produced at atmospheric pressure showed
an internal non-uniform porous structure covered by an
impermeable layer (Fig. 4b). This skin-like formation on the top
layer and less pores in the matrix of samples fabricated at
atmospheric condition was expected to impede cell proliferation
in 3D matrices. However, the hydrogels produced at high
pressure CO2 were expected to support cellular growth due to
porosity-facilitated oxygen and nutrient transport throughout
the structures.
A unique feature of crosslinking elastin under high pressure CO2
is the formation of channels throughout the 3D structure of the
hydrogels as demonstrated in Fig. 5. These channels were most
likely induced during the depressurisation stage. At a microscopic
Fig. 5. SEM images of GA cross-linked hydrogels produced at 60 bar (0.5% GA) demonstrating the presence of channels throughout the constructs.
6
N. Annabi et al. / Biomaterials 30 (2009) 1–7
Fig. 6. SEM images of fibroblast cells attached to a-elastin hydrogel fabricated with high pressure CO2 using 0.5% (v/v) GA. (a) top surface, (b) to (g) internal surfaces of channel
obtained by cross sectioning the sample, (h) control unseeded hydrogel. Sheets of cells are in images (a)–(d) and individual cells in images (e)–(g).
level, the size of the pores fabricated at both atmospheric and high
pressure CO2 was smaller than 15 mm as indicated in Fig. 3. Whilst
these pore sizes are suitable for the diffusion of nutrients, oxygen,
and waste from cells they are not large enough to allow for cells
such as fibroblast to penetrate and grow into the matrices.
However, the presence of communicating larger channels was
expected to allow for cell infiltration and proliferation into the 3D
microstructure.
3.4. In vitro fibroblast cell proliferation using elastin hydrogels
Cellular growth and proliferation in a-elastin hydrogels were
examined by SEM to demonstrate the feasibility of using the processed material for tissue engineering applications. Cells were able
to attach and proliferate on the surface of a-elastin hydrogels
fabricated at atmospheric pressure but could not penetrate into the
structure due to the skin formation on the top surface, lack of
channels and undesirable pore size. However, encouragingly cells
were able to colonise both the top surface (Fig. 6a) and the internal
surfaces of larger channels (Fig. 6b–g) within a-elastin hydrogels
produced under high pressure CO2. These preliminary studies lay
the foundation for future experiments to investigate the use of
these hydrogels as tissue engineering scaffolds. Extensive analysis
of cellular interaction with the scaffold including extracellular
matrix deposition is required whilst fabrication techniques that
lead to uniform distribution of channels throughout the hydrogel
would be desirable.
4. Conclusions
Our study demonstrates the feasibility of fabricating elastin
hydrogels in an aqueous solution using high pressure CO2. There
N. Annabi et al. / Biomaterials 30 (2009) 1–7
were three advantages in using high pressure CO2. First, carbon
dioxide strengthened the hydrogel, whereas hydrogels that formed
at atmospheric pressure were very fragile. Second, highly interconnected pores were formed with thin walled structures that
resemble the natural elastin and allowed for rapid nutrient and
oxygen transfer. Third, the unique features of the dense gas allowed
for the fabrication of channels within the 3D structure that
substantially promoted fibroblast infiltration and growth
throughout the matrices.
Acknowledgements
This project was supported by grants to F.D. and A.S.W. by the
Australian Research Council. The authors thank Mr. Aldric Tumilar
for his help in preparation of SEM images of cells.
References
[1] Quirk RA, France RM, Shakesheff KM, Howdle SM. Supercritical fluid technologies and tissue engineering scaffolds. Curr Opin Solid State Mater Sci
2005;8(3–4):313–21.
[2] Leong KF, Cheah CM, Chua CK. Solid freeform fabrication of three-dimensional
scaffolds for engineering replacement tissues and organs. Biomaterials
2003;24(13):2363–78.
[3] Peppas NA, Hilt JZ, Khademhosseini A, Langer R. Hydrogels in biology and
medicine: from molecular principles to bionanotechnology. Adv Mater
2006;18(11):1345–60.
[4] Hoffman AS. Hydrogels for biomedical applications. Adv Drug Deliv Rev
2002;54(1):3–12.
[5] Simnick AJ, Lim DW, Chow D, Chilkoti A. Biomedical and biotechnological
applications of elastin-like polypeptides. Polym Rev (Philadelphia, PA, United
States) 2007;47(1):121–54.
[6] Daamen WF, Veerkamp JH, van Hest JCM, van Kuppevelt TH. Elastin as
a biomaterial for tissue engineering. Biomaterials 2007;28(30):4378–98.
[7] Cox BA, Starcher BC, Urry DW. Coacervation of a-elastin results in fiber
formation. Biochim Biophys Acta 1973;317(1):209–13.
[8] Mithieux SM. Synthetic elastin: construction and properties of cross-linked
human tropoelastin. PhD dissertation. Sydney, NSW: University of Sydney;
2003.
[9] Welsh ER, Tirrell DA. Engineering the extracellular matrix: a novel approach to
polymeric biomaterials. I. Control of the physical properties of artificial protein
matrices designed to support adhesion of vascular endothelial cells. Biomacromolecules 2000;1(1):23–30.
[10] Martino M, Tamburro AM. Chemical synthesis of cross-linked poly(KGGVG), an
elastin-like biopolymer. Biopolymers 2001;59(1):29–37.
[11] McMillan RA, Caran KL, Apkarian RP, Conticello VP. High-resolution topographic imaging of environmentally responsive, elastin-mimetic hydrogels.
Macromolecules 1999;32(26):9067–70.
[12] Mithieux SM, Rasko JEJ, Weiss AS. Synthetic elastin hydrogels derived from
massive elastic assemblies of self-organized human protein monomers.
Biomaterials 2004;25(20):4921–7.
[13] Di Zio K, Tirrell DA. Mechanical properties of artificial protein matrixes engineered for control of cell and tissue behavior. Macromolecules
2003;36(5):1553–8.
[14] McMillan RA, Conticello VP. Synthesis and characterization of elastin-mimetic
protein gels derived from a well-defined polypeptide precursor. Macromolecules 2000;33(13):4809–21.
[15] Bellingham CM, Lillie MA, Gosline JM, Wright GM, Starcher BC, Bailey AJ, et al.
Recombinant human elastin polypeptides self-assemble into biomaterials
with elastin-like properties. Biopolymers 2003;70(4):445–55.
[16] Leach JB, Wolinsky JB, Stone PJ, Wong JY. Crosslinked alpha-elastin biomaterials: towards a processable elastin mimetic scaffold. Acta Biomater
2005;1(2):155–64.
7
[17] Nowatzki PJ, Tirrell DA. Physical properties of artificial extracellular matrix
protein films prepared by isocyanate crosslinking. Biomaterials 2003;25(7–8):
1261–7.
[18] Trabbic-Carlson K, Setton LA, Chilkoti A. Swelling and mechanical behaviors
of chemically cross-linked hydrogels of elastin-like polypeptides. Biomacromolecules 2003;4(3):572–80.
[19] Lim DW, Nettles DL, Setton LA, Chilkoti A. Rapid cross-linking of elastin-like
polypeptides with (hydroxymethyl)phosphines in aqueous solution. Biomacromolecules 2007;8(5):1463–70.
[20] Tai H, Popov VK, Shakesheff KM, Howdle SM. Putting the fizz into chemistry:
applications of supercritical carbon dioxide in tissue engineering, drug
delivery and synthesis of novel block copolymers. Biochem Soc Trans
2007;35(3):516–21.
[21] McHugh MA. Supercritical fluid extraction: principles and practice. Boston:
Butterworths; 1986.
[22] Kazarian SG. Polymer processing with supercritical fluids. Poly Sci
2000;42(1):78–101.
[23] Cooper AI. Polymer synthesis and processing using supercritical carbon
dioxide. J Mater Chem 2000;10(2):207–34.
[24] Barry JJA, Silva MMCG, Popov VK, Shakesheff KM, Howdle SM. Supercritical
carbon dioxide: putting the fizz into biomaterials. Phil Trans R Soc A 2006;
364(1838):249–61.
[25] Cansell F, Aymonier C, Loppinet-Serani A. Review on materials science and
supercritical fluids. Curr Opin Solid State Mater Sci 2003;7(4–5):331–40.
[26] Barry JJA, Gidda HS, Scotchford CA, Howdle SM. Porous methacrylate scaffolds: supercritical fluid fabrication and in vitro chondrocyte responses.
Biomaterials 2004;25(17):3559–68.
[27] Temtem M, Casimiro T, Mano JF, Aguiar-Ricardo A. Green synthesis of
a temperature sensitive hydrogel. Green Chem 2007;9(1):75–9.
[28] Cooper AI, Wood CD, Holmes AB. Synthesis of well-defined macroporous
polymer monoliths by sol-gel polymerization in supercritical CO2. Ind Eng
Chem Res 2000;39(12):4741–4.
[29] Cooper AI, Holmes AB. Synthesis of molded monolithic porous polymers using
supercritical carbon dioxide as the porogenic solvent. Adv Mater 1999;11(15):
1270–4.
[30] Lai H-M, Lee K-R, Tsai C-C, Shih H-H, Chang Y-C, inventors. Method for
crosslinking porous biodegradable polymers using supercrit. Fluids. US Patent
No. 6670454; 2003.
[31] Partap S, Rehman I, Jones JR, Darr JA. Supercritical carbon dioxide in water
emulsion-templated synthesis of porous calcium alginate hydrogels. Adv
Mater 2006;18(4):501–4.
[32] Palocci C, Barbetta A, La Grotta A, Dentini M. Porous biomaterials obtained
using supercritical CO2-water emulsions. Langmuir 2007;23(15):8243–51.
[33] Lee J-Y, Tan B, Cooper AI. CO2-in-water emulsion-templated poly(vinyl
alcohol) hydrogels using poly(vinyl acetate)-based surfactants. Macromolecules 2007;40(6):1955–61.
[34] Dehghani F, Annabi N, Valtchev P, Mithieux SM, Weiss AS, Kazarian SG, et al.
Effect of dense gas CO2 on the coacervation of elastin. Biomacromolecules
2008;9(4):1100–5.
[35] Goncalves CA, Figueiredo MH, Bairos VA. Three-dimensional organization of
the elastic fibres in the rat lung. Anat Rec 1995;243(1):63–70.
[36] Vesely I. The role of elastin in aortic valve mechanics. J Biomech 1998;
31(2):115–23.
[37] van de Witte P, Dijkstra PJ, van den Berg JWA, Feijen J. Phase separation
processes in polymer solutions in relation to membrane formation. J Memb Sci
1996;117(1–2):1–31.
[38] Mano JF, Silva GA, Azevedo HS, Malafaya PB, Sousa RA, Silva SS, et al. Natural
origin biodegradable systems in tissue engineering and regenerative medicine: present status and some moving trends. J R Soc Interface 2007;
4(17):999–1030.
[39] Ma PX. Biomimetic materials for tissue engineering. Adv Drug Deliv Rev
2008;60(2):184–98.
[40] Kang H-W, Tabata Y, Ikada Y. Effect of porous structure on the degradation of
freeze-dried gelatin hydrogels. J Bioact Compat Polym 1999;14(4):331–43.
[41] Le M, Dehghani F, Poole-Warren LA, Foster NR. A Novel method for fabrication of
porous chitosan for biomedical or pharmaceutical applications. Proceeding of International Symposium of Supercritical Fluids, Orlando, Florida, USA, May 1–4, 2005.
[42] Gosline JM, French CJ. Dynamic mechanical properties of elastin. Biopolymers
1979;18(8):2091–103.