WO2009105820A1

WO2009105820A1 - Hydrogels derived from biological polymers

Info

Publication number: WO2009105820A1
Application number: PCT/AU2009/000223
Authority: WO
Inventors: Fariba Dehghani; Anthony Steven Weiss; Nasim Annabi; Suzanne Marie Mithieux
Original assignee: The University Of Sydney
Priority date: 2008-02-29
Filing date: 2009-02-27
Publication date: 2009-09-03

Abstract

The invention relates to elastin-based hydrogels that are formed under pressurised conditions

Description

Hydrogels Derived from Biological Polymers

Field of the invention

The invention relates to biocompatible hydrogels.

Background of the invention

A hydrogel is a network of polymer chains that are water-insoluble. They are superabsorbent (they can contain over 99% water) and they may be porous. Hydrogels may possess a degree of flexibility very similar to natural tissue due to significant water content.

These characteristics mean that a number of laboratory and therapeutic applications have been anticipated for hydrogels. These include applications as scaffolds for tissue engineering, sustained release delivery systems, provision of articular surfaces and as dressing for wounds.

The porous structure of hydrogels is anticipated to be particularly important in some of these applications such as tissue engineering and sustained release systems. These hydrogels have generally been formed from synthetic polymers or natural polymers such as agarose and methylcellulose.

One common feature of the therapeutic applications mentioned above is that the hydrogels are usually in contact with living tissue. This contact demands that the polymer chains be biocompatible with the tissue.

One problem to date has been that it has been difficult to obtain homogenous 3D porous hydrogels suitable for use in laboratory and therapeutic applications. In particular for therapeutic applications it has been difficult to obtain biocompatible hydrogels, especially hydrogels that are formed from the same or similar molecules present in a living tissue in which the hydrogel is to be placed.

There remains a need for hydrogels having improved biocompatibility. There also remains a need for hydrogels having a porous structure that lends them to application in one or more of the anticipated therapeutic applications mentioned above. Summary of the invention

The invention seeks to at least minimise one or more of the above problems or limitations and in one embodiment provides a method for producing a hydrogel. The method includes:

- providing a vessel including a liquid phase, the liquid phase including a coacervated elastin material;

- pressurising the vessel with a gas to dissolve the gas into the liquid phase;

-depressurising the vessel to release gas from the liquid phase, thereby producing the hydrogel.

In another embodiment there is provided a method for producing a hydrogel including one or more pores. The method includes:

- providing a vessel including a liquid phase, the liquid phase including coacervated elastin material;

- pressurising the vessel with a gas to dissolve the gas into the liquid phase;

- cross linking the coacervated elastin material with a cross linking agent to form cross linked elastin having gas dissolved in the liquid phase entrapped within it;

-depressurising the vessel to release the gas entrapped within the cross linked elastin material to form one or more pores in the cross linked elastin material; thereby producing a hydrogel including one or more pores.

In another embodiment there is provided an apparatus for forming a hydrogel, the apparatus including:

- a vessel including a liquid phase, the liquid phase including coacervated elastin material;

- pressurising means for providing the vessel with a gas to pressurise the vessel; - injection means for injection of a cross linking agent for cross linking the coacervated elastin material into the liquid phase; and optionally

- input means for input of one or more components to be provided in a hydrogel formed by the apparatus, the one or more components selected from the group consisting of a protein, a sugar, a lipid, a cell and a pharmaceutical.

In another embodiment there is provided a hydrogel, the hydrogel including:

- a scaffold of cross linked elastin material molecules

- water molecules bound to the scaffold

- the hydrogel characterised in that the scaffold of cross linked elastin material molecules are arranged to provide the hydrogel with pores that extend throughout the hydrogel.

In another embodiment there is provided a hydrogel produced by a method according to one of the embodiments described above.

Brief description of the drawings

Figure 1 shows a schematic diagram of an apparatus of one of the embodiments of the invention, which is a one-step process for simultaneous cross-linking, homogenous pore formation in the polymer matrices and removing residual of cross-linking agent.

Figure 2 shows a schematic diagram of an apparatus of another of the embodiments of the invention.

Figure 3 depicts the effect of CO₂ pressure on coacervation temperature of σ-elastin

Figure 4 depicts effect of CO₂ pressure on the coacervation of σ-elastin at 37⁰C: control ( ■ ), 180 bar( # )

Figure 5 depicts a table presenting the effect of CO₂ pressure on the time to achieve maximum coacervation for σ-elastin. Figure 6 presents SEM images of α-elastin hydrogel (a) and(b) at 100 bar; (c) and (d) at 60 bar; (e) and (f) at 150 bar; (g) at 60 bar and (h) at 1 bar.

Figure 7 presents a skyscan analysis of lyophilised σ-elastin hydrogels fabricated at 100 bar (a) and 1 bar (b)

Figure 8 shows images of fibroblast cells cultured on hydrogels produced at 60 bar CO₂ (a, b) and atmospheric pressure (c)

Figure 9 demonstrates the effect of media on swelling ratio of σ-elastin hydrogels provided in Example 4 at 4 ⁰C (0.5% (v/v) GA).

Figure 10 depicts (a) ESEM image of a wet σ-elastin hydrogel produced at 60 bar CO₂ pressure. SEM images of an σ-elastin hydrogel fabricated at (b) 60 bar CO₂ pressure, (c) 100 bar CO₂ pressure and (d) atmospheric pressure. Note that figure 10 (d) and 6(h) are equivalent; the figure is presented again for comparative purposes.

Figure 11 presents a skyscan analysis of lyophilised σ-elastin hydrogels fabricated at high pressure CO₂ (a) and 1 bar (b). Note that figure 11(b) and 7(b) are equivalent; the figure is presented again for comparative purposes.

Figure 12 depicts SEM images of fibroblast cells attached to an σ-elastin hydrogel fabricated by high pressure CO₂ using 0.5 % (v/v) GA (a) top surface, (b) to (g) internal surface of channel obtained by cross sectioning the sample, (h) control sample - an unseeded hydrogel. Sheets of cells can be seen in images (a)-(d) and individual cells in images (e)-(g).

Figure 13 depicts SEM images of σ-elastin hydrogels fabricated at (a, b) atmospheric pressure using 2 % HMDI, (c, d) 60 bar CO₂ pressure using 2 % HMDI, (e, f) 60 bar CO₂ pressure using 5 % HMDI, and (g) 100 bar CO₂ pressure using 2 % HMDI.

Figure 14 depicts unconfined compressive behaviour of HMDI crosslinked hydrogel. Cyclic stress-strain data for the sample produced at high pressure CO₂ (a) and atmospheric condition (b). Compressive modulus (c) and energy loss (d) at each strain level for both hydrogels produced at high pressure and atmospheric condition. Figure 15 depicts the swelling behaviour of fabricated hydrogel produced at high pressure and atmospheric condition in PBS, water, and DMSO.

Figure 16 depicts the relationship between the swelling ratio in PBS and compressive modulus for fabricated hydrogels at high pressure CO₂ and atmospheric condition.

Figure 17 depicts images of fibroblast cells cultured on (a-c) hydrogel produced at 60 bar CO₂ and (d and e) atmospheric pressure.

Figure 18 depicts SEM images of fibroblast cells attached to αr-elastin hydrogel fabricated at atmospheric condition (a), and high pressure CO₂ (b-f). (a), (b) and (c) depict the top surface of the hydrogel whilst (d) to (f) depict internal surfaces of hydrogel obtained by cross sectioning the sample.

Figure 19 depicts the swelling behaviour of elastin-tropoelastin hydrogels produced at high pressure CO₂ and atmospheric condition in PBS at 37°C.

Figure 20 depicts SEM images of TPE/σ-elastin hydrogels generated at atmospheric pressure using (a) 0.1 , (b) 0.25, and (c) 0.5 % (v/v) GA.

Figure 21 depicts SEM images of TPE/σ-elastin hydrogels fabricated at (a-d) 60 bar CO₂ pressure, (e-h) atmospheric pressure. Top surface of the samples are shown in images (a), (b), (g), and (h), cross sections in images (c)-(f).

Detailed description of the embodiments

Elastin is a molecule that is found in many tissues. It is essentially formed by a two step process, the first involving an alignment of tropoelastin or elastin fragments that brings functional groups into close proximity through an interaction of hydrophobic groups known as "coacervation". The second step involves the cross linking of aligned or "coacervated" elastin so as to form covalent bonds between the aligned relevant functional groups. Cross linking is generally achieved in a living system by lysyl oxidase. In the laboratory this may be achieved using a range of reagents including glutaraldehyde, amine-reactive chemical crosslinkers including BS3 and amine and carboxyl reactive crosslinkers including EDC. Coacervation is a critical step. Without this step it is basically not possible to effect a cross linking reaction which would produce elastin material in the form as is generally observed in a living system. Coacervation is a reversible step - molecules can be "unaligned" by manipulating pH, temperature or salt. The effect of pressure on coacervation at the time of this invention was basically not known.

The inventors have found that while pressure does affect coacervation, it is nonetheless possible to coacervate elastin materials where an elastin material would include tropoelastin, purified elastin such as α elastin, sub-fragments or elastin like peptides under pressures greater than normal atmospheric pressure. The inventors have recognised one application which is that hydrogels, in particular porous hydrogels can be generated by applying pressure to an elastin material coacervate together with cross linking. This enables the formation of hydrogels having biocompatible molecules therein and which may have a porous structure.

Thus in embodiment there is provided a method for producing a hydrogel. The method includes:

- pressurising the vessel with a gas to dissolve the gas into the liquid phase;

"Hydrogel" generally refers to a substance formed from, or comprised of, a network of polymer chains that are water-insoluble, in which water is the dispersion medium. Hydrogels are superabsorbent (they can contain over 99% water). Hydrogels also possess a degree of flexibility very similar to natural tissue, due to their significant water content.

The vessel including the liquid phase having coacervated elastin material may be provided by initial steps including: -providing a vessel including a liquid phase, the liquid phase in the form of a solution of elastin material and

- providing conditions to the vessel for coacervation of elastin material in the solution.

The liquid phase may be provided with a cross linking agent for cross linking elastin material in the liquid phase before or after the vessel is pressurised, or after it is depressurised.

In one embodiment the liquid phase is provided with a cross linking agent to cross link coacervated elastin material in the liquid phase before the vessel is pressurised.

In another embodiment, the liquid phase is provided with a cross linking agent to cross link coacervated elastin material in the liquid phase after the vessel has been pressurised. In this embodiment cross linking can occur either at maintained pressurisation or after some of the pressure has been released. One advantage of this embodiment is that the hydrogel formed after depressurisation is provided with one or more pores. These pores are particularly useful for increasing the water binding capacity of the hydrogel. They are also useful for providing a surface for providing the hydrogel with other components such as pharmaceutical drugs, structural and bioactive proteins, sugars and lipids. In certain embodiment the pores are useful for seeding with cells such as fibroblasts and the like. While not wanting to be bound by hypothesis, it is believed that cross linking of coacervated elastin material after gas has been dissolved in the liquid phase tends to entrap dissolved gas. The release of dissolved gas from the cross linked elastin material as the vessel is depressurised is believed to form pores in the cross linked elastin material, hence providing a hydrogel having pores. One particular advantage of this embodiment of the invention is that it is possible to control pore size, shape and distribution throughout the hydrogel formed by the process by controlling the amount of pressure provided to the vessel, the rate of depressurisation, the concentration of cross linking agent and coacervated elastin material.

Thus in certain embodiments the step of cross linking the coacervated elastin material after pressurisation of the vessel entraps gas molecules dissolved in the liquid phase within the cross linked elastin material. In other embodiments, the step of depressurisation provides conditions for release of gas entrapped within the cross linked elastin material.

In still further embodiments, the release of gas entrapped within the cross linked elastin material provides the hydrogel formed from the cross linked elastin material with one or more pores. The one or more pores preferably form one or more conduits that extend throughout the hydrogel or ramify to form networks throughout the hydrogel.

The cross linking agents may be enzymatic (such as lysyl oxidase) or chemical, examples of the latter being glutaraldehyde (GA), amine-reactive chemical cross linkers and amine and carboxyl reactive cross linkers. Examples of amine-reactive cross linkers include disuccinimidyl glutarate (DSG), bis(sulfosuccinimidyl) suberate (BS3), ethylene glycol diglycidyl ether (EGDE) under neutral conditions (pH 7), hexamethylene diisocyanate (HMDI), tris-succinimidyl aminotriacetate (TSAT), disuccinimidyl suberate (DSS) and β-[tris(hydroxymethyl)phosphino]propionic acid (THPP). Examples of carboxyl reactive cross linkers include1-ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC), and ethylene glycol diglycidyl ether (EGDE) under acidic conditions (pH<4).

Generally, the cross linker may be provided in the liquid phase in an amount of from about 0.05 to about 10 percent (v/v). In one embodiment, the cross linker is GA and is provided in an amount of from about 0.05 to about 5 percent (v/v). In another embodiment the cross linker is GA and is provided in an amount of from 0.05 to about 2 percent (v/v). In another embodiment the GA is provided in the liquid phase in an amount of from about 0.1 to about 0.5 percent (v/v).

In another embodiment, the cross linker is HMDI and is provided in the liquid phase in an amount of from about 0.25 to about 10 percent (v/v). In another embodiment, HMDI is provided in the liquid phase in an amount of from about 1 to about 5 percent (v/v). In another embodiment, the HMDI is provided in the liquid phase in an amount of from about 2 to about 5 percent (v/v).

It will be understood that the conditions provided to the vessel for coacervation of elastin material in the solution may be provided at the time that the vessel is pressurised with a gas to dissolve the gas into the liquid phase. Alternatively, the conditions for coacervation may be provided before pressurisation of the vessel so that at the time of pressurisation the elastin material contained in the liquid phase is presented in the form of a coacervate.

It will be understood that elastin material for use in producing the hydrogel may be cross linked before coacervation. An example of such an elastin material is one obtained by extraction from a tissue source, such as from bovine tissue, α-elastin is one particular example. The elastin material may not be cross linked at all before coacervation. One example of such a molecule is tropoelastin. Another is an elastin like peptide or "ELP". It will also be understood that the elastin material prior to coacervation may be a mixture of cross linked and uncross linked elastin material. Thus in certain embodiments the elastin material is selected from the group consisting of tropoelastin, an elastin such as α-elastin, sub-fragments of these molecules and ELP. In other embodiments the elastin material has been obtained by extraction from a tissue. In yet further embodiments the elastin material has been obtained from a recombinant expression system, examples of which are discussed in WO99/03886, WOOO/04043 and WO94/14958, the contents of which are disclosed herein in their entirety by reference, or from peptide synthesis such as solid phase peptide synthesis.

Elastin may be provided in the liquid phase in a range of concentrations. In one embodiment, the elastin is provided in the liquid phase in an amount of from about 5 to about 200 mg/mL.

Further to elastin, the liquid phase may further include one or more further biomolecules. Examples include proteins, sugars and lipids. Where the biomolecule is a protein, connective tissue proteins and extra cellular matrix proteins such as collagens and the like are particularly preferred. Other bioactive proteins such as hormones, growth factors or cytokines are also useful. Where the biomolecule is a sugar, molecules such as glycosaminoglycans are particularly preferred.

In other embodiments, the liquid phase further includes a pharmaceutical compound. Examples of particularly preferred compounds include anti cancer compounds, such as anti angiogenic compounds and compounds useful for tissue repair or regeneration. The pH of the liquid phase is generally selected to prevent coacervation from reversing to form soluble protein. Generally, coacervation may occur when the pH value of the liquid phase is between about pH 3 and 8.

It will be understood that the pH of the liquid phase may be influenced by the gas used to pressurise the vessel.

The liquid phase may include a buffer. In one embodiment, the liquid phase includes, or is phosphate buffered saline (PBS).

The liquid phase may include a salt. Examples of suitable salts NaCI, KCI and NH₄CI. The ionic strength of the liquid phase is generally about between about 0 and about 1.5 M.

Where a cross linking agent is to be used, the liquid phase is selected having regard to the solubility of the cross linking agent. Aqueous solutions are preferred for cross linking agents that are soluble in water. For example, water or PBS can be used where the cross linker is glutaraldehyde. Non aqueous solutions are preferred for agents that are not soluble in water. For example, DMSO can be used for HMDI cross linkers as discussed in the Examples herein. Suitable non-aqueous liquid phases include solvents that can dissolve a dense gas such as CO₂. Other examples of non-aqueous solvents include ethanol, acetone, isopropanol, dimethyl formamide, and ethyl acetate.

Where a non aqueous liquid phase is used, the method generally includes a further step of removing the non aqueous liquid phase from the elastin material after depressurisation and contacting the elastin material with water to form the hydrogel.

The gas that is used to pressurise the vessel is generally any gas that is capable of dissolving into a liquid or otherwise an aqueous solution under pressure. Examples include CO₂ and N₂ gas. Other examples include ethane, methane, 1 ,1 ,1 ,2- tetrafluoroethane, NO₂, propane and heptane gas. CO₂ gas is particularly useful.

Where the gas is CO₂, generally the pressures range from about 10-300 bar. Most preferred pressures are within the range of from about 30-100 bar. Alternatively, when the gas is CO₂, the preferred pressures are within the range of about 5 to about 180 bar.

In one embodiment the vessel is pressurised in conditions in which the gas remains below the critical point. This state is sometimes referred to as a dense gas (DG). That is, the gas is not in a state generally recognised as a supercritical fluid. In these embodiments, the pressure to be provided is to be determined having regard to desired characteristics of the hydrogel to be produced according to the method. One particular characteristic is the porosity of the hydrogel. In certain embodiments this is understood to be a function of the molar amount of gas dissolved into the liquid phase.

Where the gas is CO₂, generally the pressure to be provided to the vessel dissolves an amount of CO₂ into the liquid phase, in case of aqueous solution effective for reducing the pH of the liquid phase.

In other embodiments, the vessel is pressurised in conditions in which the gas is present as a supercritical fluid. These embodiments may be advantageous where there is a need to coacervate at lower than room temperatures. For example, as discussed below, the inventors have surprisingly found that elastin can be made to coacervate at temperatures approaching 15 ⁰C when exposed to supercritical CO₂. Examples of conditions suitable for the provision of supercritical fluid include pressures of at least about 73.8 bar and temperatures of at least about 31 ⁰C.

In certain embodiments the method further includes the step of separating the liquid phase from the hydrogel.

In other embodiments, the method includes the step of desiccating the hydrogel to form a dried form of the hydrogel. This may be then be sold as a dried form to which water may be added to reconstitute the hydrogel.

In further embodiments the method includes the further step of providing the hydrogel formed by the method with a pharmaceutical compound. Generally this can be achieved by methods known in the art. In yet further embodiments, the method includes the further step of providing the hydrogel with a cell such as a fibroblast or stem cell.

- a vessel containing including a liquid phase, the liquid phase including coacervated elastin;

- pressurising means for providing the vessel with a gas to pressurise the vessel;

- injection means for injection of a cross linking agent for cross linking the coacervated elastin material into the liquid phase; and optionally

In one particularly preferred embodiment the injection means is adapted to automatically inject the cross linking agent when the vessel reaches a pre-selected pressure.

An apparatus of one embodiment described in Figure 1. An σ-elastin solution is pipetted into the sample holder placed in the temperature controlled high pressure vessel 7. The high pressure pump 4 is then pressurised with liquid CO₂ by opening valves 2 and 3. Valves 5 and 6 are opened when the vessel approaches thermal equilibrium. Vessel 7 is then slowly pressurised by CO₂ while valve 10 is closed. The temperature of the vessel is kept constant using a temperature controller 8 during cross linking. The system is isolated by shutting down valves 2, 3, 5 and 6 and maintained at constant temperature and pressure and then depressurised by opening valve 10. Pore formation in the biopolymer matrix may occur upon depressurisation and the residual cross linking agent can be removed by passing CO₂ through the sample by opening valves 2,3, 5,6 and running pump 4 at constant pressure mode. The flow rate of CO₂ is controlled during this stage by valve 10 prior to depressurisation. In another embodiment there is provided a hydrogel, the hydrogel including:

- a scaffold of cross linked elastin material molecules

- water molecules bound to the scaffold

Typically, the one or more pores form one or more conduits that extend throughout the hydrogel or ramify forming networks throughout the hydrogel.

The scaffold is preferably formed by cross linking tropoelastin.

Example 1

The objective of this study was to investigate the effect of dense gas CO₂ on the reversible molecular association of protein biopolymers. Elastin, the major component of elastic fibers, is an insoluble extracellular matrix protein found in skin, bladder, lung, ligament, elastic cartilage and arteries. It provides elasticity and resilience to maintain the proper function of tissues that are subjected to repetitive distension and physical stress. Elastin is the most persistent protein in the body. It is particularly established by assembly of its precursor tropoelastin in utero. Elastin is an extremely insoluble biopolymer and it is difficult to process into new biomaterials so σ-elastin, an oxalic acid- solubilized derivation of elastin, is frequently used for synthesizing elastin-based materials as it undergoes reversible self-association. At low temperature both tropoelastin and σ-elastin are soluble in aqueous buffers; while, at about physiological temperature a phase separation or coacervation occurs. Coacervation plays a crucial role during elastin formation by concentrating and aligning monomers prior to crosslinking to form fibers. Coacervation of these soluble molecules is a concentration- dependent process and can be interpreted as an intermolecular hydrophobic association where solution variables such as biopolymer concentration, pH, salt and impurities can influence its efficacy.

Materials and Methods

α-elastin (molecular weight »60000) extracted from bovine ligament was purchased from Elastin Products Co. (Missouri USA). Carbon dioxide (99.99 % purity) and high purity nitrogen were supplied by BOC. All aqueous solutions were prepared in MiIIiQ water, α-elastin was dissolved in phosphate-buffered saline (10 mM sodium phosphate, 135 mM NaCI, pH 7.4; PBS) at 5 mg/mL and the solution was dissolved overnight at 4°C prior to use.

The schematic diagram of the apparatus used for investigating the effect of CO₂ pressure on α-elastin is shown in Figure 2. A high pressure pump (Thar, Model P50) was used to transfer CO₂ into the high pressure view cell (Jerguson sight gauge, Model R32). A recirculation heater (Ratek) was used to control the temperature of the water bath and the pressure of the system was monitored using a pressure transducer (Druck).

In each experiment 2 ml_ fresh solution of α-elastin, which was prepared overnight, injected into a glass tube located inside the high pressure vessel. No coacervation was observed for a 5 mg/mL α-elastin used in this study at temperatures below 32°C and atmospheric pressure. After thermal equilibrium was established in the system, the air was purged using an inert gas (e.g. CO₂) at 5 bar pressure. The system was then pressurised with the DG CO₂ to different pressures ranging from 5 to 180 bar. While the pressure was kept constant, the temperature was increased incrementally from 7⁰C to 37°C. The temperature at which the solution became turbid was monitored. The reversibility of coacervation behaviour was visually verified upon decreasing the temperature while the pressure was kept constant. The system was then depressurised slowly to minimize foaming and the sample was collected. In addition, a sample of α- elastin solution was also placed in the water bath at atmospheric pressure, as a control, and its coacervation behaviour was monitored upon increasing the temperature. Results and Discussion

The effect of CO₂ on coacervation of σ-elastin was determined at various pressures and temperatures ranging from 1 to 180 bar and 7 to 37°C. Carbon dioxide exhibited a profound effect on the coacervation temperature of σ-elastin. The temperature at which coacervation commenced was substantially diminished as the CO₂ pressure increased (Figure 3). The coacervation temperature decreased from 37⁰C to 16°C when the pressure was increased from 1 bar to 50 bar and above this pressure approached a plateau. The reversible coacervation behavior at atmospheric pressure was rapid and commonly took place in less than one minute. However, samples pressurised with CO₂ at above 30 bar required at least 45 min to achieve partial solubilisation after coacervation and decreasing the temperature to 7⁰C. The solution was not completely clear as control sample after it was kept under 30 bar for 4 hours at 7°C. However, it was converted to a very clear solution upon depressurisation of the sample to pressures below 10 bar. The CO₂ pressure therefore altered the rapid reversible coacervation that is seen at atmospheric pressure when the temperature of the solution is decreased from 37°C to 20⁰C.

The coacervation behavior of α-elastin solutions exposed to CO₂ at various pressures was monitored by UV spectrophotometry after depressurisation and compared with control samples (Figure 4). At pressures above 60 bar, CO₂ affected the coacervation profiles of σ-elastin solutions. While the control sample approached 100% coacervation within 114 seconds, samples exposed to CO₂ at pressures above 60 bar achieved this within 214 seconds. The retardation was more profound for samples exposed to higher pressures, even though the pH of all solutions after depressurisation were similar to untreated sample. The effect of pressure on the reversible coacervation profile of α- elastin exposed to nitrogen at 180 bar and CO₂ at pressures below 60 bar was negligible. Supporting data is presented in a Table in Figure 5.

The coacervation temperature of σ-elastin is governed by various factors including pH, ionic strength, and protein concentration. The coacervation temperature of tropoelastin and σ-elastin decreased as the ionic strength increased. The influence of pH on coacervation temperature of σ-elastin dissolved in pure water, in buffer solution with high and low ionic strength, and tropoelastin solution at PBS (150 mM NaCI) were studied. At pH around the isoelectric point of σ-elastin (i.e. pH= 4.8) the highest level of absorption and the lowest coacervation temperature were observed. As the pH shifted from the isoelectric point, the coacervation temperature increased and the rate of turbidity decreased. Coacervation was prevented at pH values far from the isoelectric point of σ-elastin at below 3 or above 8.

The effect of DG CO₂ on coacervation temperature and partially reversible coacervation of σ-elastin solution may be due to the lower free energy of the hydrophobic core of the coacervate, compared with the usual conformation, which then achieved a folded state. The reverse transition is, therefore, no longer thermodynamically favorable and the newly formed coacervate might have a higher stability. The main contributing factor to this transition and exact mechanism of action is currently unknown. The reduction in the coacervation temperature of σ-elastin solution in DG CO₂ system may be due to a decrease in the pH of the protein solution caused by dissolved CO₂ with a potential contribution by a depression of the melting temperature of the hydrophobic core caused by the pressure. Each of these factors has the potential to lower the activation energy of the transition by inhibiting unfavorable side chain interactions that form spontaneously during the folding process and present distinctive energy barriers that have to be overcome in order to achieve a functional final set of conformations.

The effect of pressure alone on α-elastin coacervation was assessed by using the inert gas nitrogen at 180 bar. High pressure nitrogen had no adverse effect on the coacervation profile of α-elastin which was reversible and rapid. Therefore, the partially reversible coacervation and decrease in coacervation temperature of σ-elastin at high pressure can be due to a decrease in the pH of the protein solution caused by dissolved CO₂. The pH of the σ-elastin solutions used in this study was decreased from 6 to 3 when the CO₂ pressure increased from 1 bar to 50 bar, which was comparable with previous studies on decreasing the pH of water under high pressure CO₂. As expected for a reversible system, after depressurisation the solution returned to a pH close to that of the original sample as illustrated in the Table depicted in Figure 5.

High pressure CO₂ is volatile and can dissolve in aqueous solutions to decrease the pH due to the acidification of aqueous solution and production of carbonic acid. As the solubility of CO₂ in water is a function of temperature and pressure, the drop in pH of the solution will be temperature and pressure dependant. A decrease in pH of the σ- elastin solutions exposed to CO₂ may lead to a decrease in coacervation temperature of σ-elastin.

The effect of pH on coacervation temperature of σ-elastin solution used in this study was investigated at atmospheric pressure. It was found that the coacervation temperature decreased from 37°C to 16°C as the pH lowered from 6 to 3.7. The lower coacervation temperature at pH 3.7, in our study, was due to utilising a solution with a higher ionic strength (1.3). No significant retardation was observed for coacervation behavior of σ-elastin solution at pH = 3.7 at atmospheric pressure as the solution coacervated rapidly at lower temperature. However, the rapid reversible coacervation was not observed for this solution by decreasing the temperature to 4^CC. The solution became partially clear after 25 minutes at 4°C. The partial reversible coacervation of a- elastin in DG CO₂ system may result from both the pH drop in the system caused by dissolved CO₂ and the interaction between the CO₂ and σ-elastin while the retardation in the coacervation profile of σ-elastin solution exposed to CO₂ may be associated with the interaction between CO₂ and σ-elastin that may take longer time to approach the initial condition.

Conclusions

The findings presented here show it is feasible to coacervate σ-elastin using CO₂ at high pressure. The present work demonstrates that high pressure CO₂ does not disrupt the natural ability of σ-elastin to coacervate but does decrease the temperature required for coacervation. High pressure CO₂, as a volatile acid, decreased the pH of the σ- elastin solution resulting in a reduction in coacervation temperature of σ-elastin. Following exposure to high pressure CO₂ the σ-elastin solution was also maintained in a coacervated state for a longer time. These properties may usefully provide advantages for synthesizing elastin-based biomaterials under high pressure CO₂ as the solution can be kept in coacervated state during the reaction without using any mineral acid.

Example 2

A buffer solution containing α-elastin and the cross-linking agent glutaraldehyde was prepared. The solution was then placed in a high pressure vessel and the system was connected to the high pressure pump. After purging the air from the system and thermal equilibrium the system was pressurized slowly. The sample was kept at the desired pressure and temperature for a defined time and then the system was depressurised in a controlled manner. The following microscopic images (Figure 6) compare the characteristics of hydrogels fabricated by the process described by this invention and the samples produced at atmospheric conditions.

As depicted in Figure 6 a highly homogenous-interconnected-porous structure was detected for σ-elastin hydrogel obtained at 100 bar CO2. Skyscan analysis of lyophilised σ-elastin hydrogels produced at atmospheric pressure confirmed the existence of a porous structure covered with a non-porous layer of polymer as shown in Figure 7. The skin-like formation on the top layer and less pores in the matrix of samples fabricated at atmospheric pressure could function to impede cell differentiation and proliferation in 3D matrix growth. However, the hydrogel produced at high pressure CO₂ may support cellular growth due to the porosity-facilitated higher oxygen and nutrient transport through the homogenous-interconnected-porous structures.

Example 3

We examined cell growth in GA cross-linked σ-elastin hydrogels produced at high pressure CO₂ and atmospheric pressure. The fabricated hydrogels containing the grown cells were examined using optical microscopy and SEM analysis.

Materials and Methods

The σ-elastin hydrogels were formed at atmospheric pressure using 100 mg/mL of σ- elastin and 0.5 % GA and incubating the sample at 37°C for one hour.

MG 3348 cells were maintained in Advanced Dulbecoo's Modified Eagle's Medium (DMEM) supplemented with 12.5 ml/liter antibiotic-antimycotic, 12.5 ml/liter L-glutamine (Invitrogen, Australia) and 5% v/v fetal bovine serum (FBS) (JRH biosciences, Kansas, USA). TrypLE™ express stable trypsin-like enzyme was purchased from Invitrogen. 24- well-plates were obtained from Costar Corp, Cambridge, MA.

The growth of fibroblast cells (MG 3348) cultured on GA cross-linked σ-elastin hydrogels, fabricated by the process disclosed in this invention and at atmospheric pressure, was studied. Prior to seeding the cells on fabricated scaffolds, the cells were grown in DMEM media containing 5% FBS for 3 weeks to obtain sufficient cells for seeding on scaffolds. Fibroblast cells (MG 3348) were plated on 25-cm² flasks containing 12 ml_ DMEM media with 5% FBS. The growth of cells in the media was confirmed using an optical microscope after which the cells were transferred into a 75 cm² flask).

The hydrogels were transferred into individual wells of 24-well plates, washed twice with ethanol and subsequently with cell culture media. The σ-elastin scaffolds were equilibrated with culture media by soaking in 2 ml_ DMEM media containing 5% FBS at 37°C incubator overnight. The DMEM media were then replaced with 2 mL culture media containing cells. The scaffolds that were seeded with cells were left for 2 days at 37°C in a CO₂ incubator. Cell proliferation was visually assessed by optical microscopy and SEM after fixing and staining seeded scaffolds.

The existence and growth of cells in fabricated hydrogels was confirmed using light microscopic analysis after fixing and staining of slides of cross-sections of seeded scaffolds to show cellular growth in fabricated scaffolds. The cells were fixed into the scaffolds by soaking in 10% formalin and the scaffolds were immersed in 70% ethanol. The samples were processed on an automated tissue processor on a 6 hour cycle to paraffin through a graded series of ethanol, and xylene. They were embedded in paraffin wax and 5 μm sections were taken and collected onto glass slides and dried. The slides were then deparaffinised, rehydrated, stained using a standard haematoxylin and eosin staining procedure, dehydrated, cleared in xylene and mounted in DPX. Then, the slides were monitored using light microscope connected to the camera.

Results and Discussion

Cell growth on α-elastin scaffold:

Two samples of GA cross-linked σ-elastin hydrogels fabricated either at 60 bar or at atmospheric condition (in an oven at 37°C) were transferred from sample holders into individual wells of 24-well plates. The hydrogels produced at high pressure CO₂ was easily handled and placed at the bottom of each well plate. However, the samples produced at atmospheric pressure showed good compressible elasticity but were brittle. Equal numbers of cells were seeded into both types of hydrogels and the 24-well-plates were incubated for 2 days at 37°C incubator

Microscopic analysis of adherent fibroblasts suggests that the hydrogels disclosed in this invention may be advantageous to cell culturing. The images of adherent fibroblast cells cultured on both hydrogel produced at 60 bar CO₂ and atmospheric pressure are shown in Figure 8. Haematoxylin and eosin were used for staining, as a result the cells appear as purple (dark grey) and the σ-elastin scaffolds as pink (light grey). The images (Figure 8a and 8b) show the feasibility of cell attachment and infiltration into elastin samples fabricated by the process disclosed in this invention at 60 bar. The a- elastin scaffolds could support cellular growth as the cells were capable of attaching and passing through the hydrogel scaffold. Cells showed lower attachment to the surface of elastin scaffolds fabricated at atmospheric conditions as shown in Figure 8c that can be explained by the weak mechanical properties of fabricated hydrogels and impeded pores.

The cells could infiltrate through the channels in the hydrogel matrix. The attachment of the cells to the hydrogel matrix corroborates the biocompatibility of fabricated hydrogels produced at high pressure CO₂. Engineering large channels for the cells to penetrate in the hydrogel matrix are the advantages of this novel method.

Example 4

Cross-linking σ-elastin Materials and Methods σ-elastin extracted from bovine ligament was obtained from Elastin Products Co. (Missouri USA). All aqueous solutions were prepared in MiIIiQ water, σ-elastin was dissolved in PBS (Phosphate-buffered saline) (10 mM sodium phosphate pH 7.4, 1.35 M NaCI). Gluteraldehyde (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.

Hydrogel formation at atmospheric condition

σ-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 hr, to fabricate a hydrogel. The crosslinked hydrogels were washed in MiIIiQ water, then placed in 100 mM Tris ((HOCH₂)₃CNH₂) in PBS for 1 hr to inhibit further cross-linking and stored in PBS for characterisation.

A preliminary set of experiments were conducted to determine the required amount of GA and σ-elastin for the hydrogel fabrication. The concentration of σ-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 hr. The results demonstrated in Table 1 , shows that the hydrogels were formed when the GA and σ-elastin concentration were above 0.25 % (v/v) and 100 mg/ml, respectively. At low concentrations of GA and σ-elastin, thin films of cross-linked σ-elastin were formed at the bottom of the mould due to insufficient degree of cross-linking, 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 σ-elastin solution and two different crosslinker concentrations, 0.5 and 0.25 % (v/v) GA, were used to produce hydrogels.

Table 1. Optimization of protein and GA concentration for σ-elastin hydrogel fabrication

10 . . . . .

50 . . . . .

100 - s Y s s 200 ^ -

S; Hydrogel was formed, - : no hydrogel formed at low concentrations (thin films were formed) or non-homogenous gel at higher concentrations

Dense Gas Hydrogel Formation process (DGHF)

A schematic diagram of the apparatus used to fabricate σ-elastin hydrogels using dense gas CO₂ is shown in Figure 1. A syringe pump (ISCO, Model 500D) was used to transfer CO₂ into the high pressure vessel (Thar, 100 ml_ view cell). The vessel was 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.

σ-elastin solution containing the cross-linker 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 MiIIiQ water, then placed in 100 mM Tris in PBS for 1 hr. After Tris treatment, the hydrogels were washed twice in MiIIiQ water and stored in PBS for further analysis.

All experiments were conducted at 37° C. The effect of pressure, reaction time, and crosslinker concentration on the characteristics of the hydrogel was assessed. Different concentrations of GA (i.e. 0.25 and 0.5 %( Wv)) were mixed with 100 mg/ml σ-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.

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 hr using a freeze dryer. The swelling properties were measured at four different conditions, 37 ⁰C and 4 ⁰C using either PBS or MiIIiQ 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:

~ • Weight of wet sample - Weight of dry sample

Swelling Ratio = - - - - —

Weight of dry sample

Environmental SEM (ESEM)

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.

Micro-CT analysis (3D image analysis)

Lyophilized σ-elastin hydrogels were analysed using a Skyscan 1072 (Skyscan, Belgium) high-resolution desktop X-ray CT scanner at 2.94 μm voxel resolution, X-ray tube current 160 μA 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). Scanning electron microscopy (SEM)

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 σ-elastin hydrogels were mounted on aluminium sample 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 hr at 37°C. Samples underwent post-fixation with 1 % osmium in 0.1 M Na- cacodylate for 1 hr 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.

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 37⁰C over night. 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.

Light microscopic analysis on histological samples

The growth of the cells in fabricated hydrogels was confirmed using light microscopic analysis after fixing, sectioning, and staining cross sections of cell seeded scaffold. The hydrogels containing cells were fixed by soaking in 10% formalin over night. The scaffolds were then immersed in 70% ethanol. The samples were processed on an automated tissue processor on a 6 hour cycle to paraffin through a graded series of ethanol, and xylene. They were embedded in paraffin wax and 5 μm sections were taken and collected onto glass slides and dried. The slides were then deparaffinised, rehydrated, stained using a standard haematoxylin and eosin staining procedure, dehydrated, cleared in xylene and mounted in DPX. Then, the cross-sections were examined a using light microscope connected to the camera.

Results and Discussion

In this study, the effect of high pressure CO₂ on the crosslinking of σ-elastin hydrogels was investigated, σ-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.

Visual comparison of hydrogels formed at high pressure CO₂ and atmospheric condition

Visual comparison of fabricated σ-elastin hydrogels demonstrated that gels produced at high pressure CO₂ were more rigid than those produced at atmospheric condition, σ- elastin hydrogels fabricated at high pressure CO₂ could be easily handled. However, the hydrogels produced at atmospheric pressure were very fragile. Increasing the processing time from 1 to 24 hours had negligible effect on promoting the macrostructure of the elastin formed at atmospheric condition. This may be due to an insufficient degree of cross-linking throughout the hydrogel matrices produced at atmospheric condition. High pressure CO₂ appeared to promote the coacervation of σ- elastin and accelerated the crosslinking of σ-elastin.

Hydrogel swelling

σ-elastin hydrogels fabricated in this study exhibited high swelling ratios, in the range of 21-35 g H₂OZg protein, in water. Increasing the processing pressure from 30 bar to 150 bar resulted in a 60% increase in the hydrogel swelling ratio as indicated in Figure 9. Both hydrogels produced at high pressure CO₂ and atmospheric pressure displayed stimuli-responsive characteristic toward temperature and salt concentrations when they were swelled in PBS and water at 4 ^CC and 37 ⁰C. The swelling ratio of the hydrogels in water was greater than those swelled in PBS at 4 ⁰C as shown in Figure 9. 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 and 28.6 ± 1.6 g liquid/g protein at 4 ⁰C and 37 ⁰C, respectively, when they were hydrated in water. However, they absorbed 18.3 ± 4.6 and 7 ± 3.2 g liquid/g protein at 4 ⁰C and 37 ⁰C, 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 σ-elastin hydrogels produced at high pressure CO₂ was either considerably greater or comparable with other elastin based hydrogels fabricated using various cross-linkers.

Table 2. Swelling ratios of fabricated hydrogels using high pressure CO₂ and conventional methods

Swelling Ratio (g liquid/g Protein )

Protein Crosslinker ref

PBS MiIIiQ Water σ-elastin produced 18.3± 4.6 ^a 33.2+0.8 ^a this at high pressure GA study CO₂ 7± 3.2 ^b 28.6±1.6 ^b

0.8 at 2⁰C

Natural elastin - [1] 0.5 at 36°C

Synthetic 6.9 ± 0.5 ^a 63±5 ^a

BS3 tropoelastin [2] 3.8±0.8 ^b 33+4 ^b

Synthetic 7-80 at 25°C

GA - tropoelastin [3]

Engineered ELPs HMDI - 0.37 ^a [4]

12.3±0.5 ^a

Engineered ELPs THPP - 3.7 ±0.2 ^b [5]

σ-elastin

EGDE 8.4-24 [6]

Swelling Temperature (a): 4°C and (b): 37°C

The variation of pressure and crosslinker concentration had a significant 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 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 higher concentration of cross-linker resulted in higher degree of cross-linking through the hydrogel matrices which resulted in a reduction in the swelling ratio.

Table 3: Effect of pressure and crosslinker concentration on the swelling ratio of fabricate hydrogels at high pressure CO₂.

GA

Swelling Ratio

Pressure (bar) Concentration (g H₂O/g protein)

% (v/v) 30 0.25 34.5± 1.3

60 0.25 38.4± 6

150 0.25 54 ± 9.2

30 0.5 21.4± 0. 7

60 0.5 33.2±0.8

150 0.5 35.2± 2.5

Pore structure of the a-elastin hydrogel

Pore size and interconnectivity are critical hydrogel properties that influence the ability of cells to infiltrate and proliferate within 3D structures. In this study, ESEM, SEM, and Micro-CT analysis were used to characterise the pore morphology of fabricated σ- elastin hydrogels. ESEM analysis demonstrated that σ-elastin hydrogels fabricated under high pressure CO₂ (Figure 10a) were highly porous yet robust structures. ESEM images of σ-elastin hydrogels produced under atmospheric pressure could not be obtained because the wet constructs disintegrated prior to examination. SEM images of σ-elastin hydrogels fabricated under high pressure (Figures 10b and 10c) 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 σ-elastin hydrogels produced under either high pressure CO₂ or atmospheric conditions (Figure 10d) 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 μm to 4.9 ± 4.3 μm 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 μm compared to 4.3 ± 5.5 μm for σ-elastin hydrogels fabricated under atmospheric conditions. This represents nearly a 10-fold reduction in pore wall thickness of σ-elastin hydrogels fabricated at high pressure CO₂ compared with the samples produced at atmospheric pressure. The organisation of the matrix produced under high pressure is similar to that seen in some natural elastin microstructures within the body [7, 8].

Cross-linked σ-elastin hydrogels fabricated under atmospheric conditions consisted of polyhedral shaped pores with thick walls. However, highly interconnected pores with thin walls were formed when σ-elastin was cross-linked under high pressure CO₂.

The effect of cross-linker concentration on pore morphology ofhydrogel

The effect of GA concentration on the pore morphology of σ-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 σ-elastin hydrogel slightly increased from 10 ± 1.7 μm to 14.5 ± 3.3 μm. 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 μm to 9 ± 5.4 μm when the GA concentration was changed from 0.25 to 0.5% (v/v). Increasing the crosslinker concentration is believed to result in the polymerization of GA molecules in aqueous solution. The polymerized GA can then react with the amine groups in σ- elastin molecules, resulting in the production of longer crosslink length between molecules and the formation of larger pores.

The effect of process variables on the properties ofhydrogel

The effects of processing variables including pressure, reaction time, temperature, and depressurization rate on the characteristics of the fabricated hydrogels were determined. The reaction time and depressurization rate had negligible effect on either the swelling ratio or the pore size of the fabricated hydrogels. However, increasing the processing pressure from 30 bar to 150 bar resulted in a 60% increase in the hydrogel swelling ratio. The average pore size of hydrogels decreased from 14.3 ± 3.3 μm to 4.9 ± 4.3 μm and 3.6± 1.6 μm as pressure increased from 1 bar to 60 bar and 100 bar as shown in Figure 10. The thickness of the walls between pores was also diminished from 4.3 ± 5.5 μm to 0.5 ± 0.1 μm and 0.46±0.11 μm when pressure increased from 1 bar to 60 bar and 100 bar. However, at pressures below 60 bar, dense gas CO₂ had no significant effect on the pore size of fabricated hydrogels.

In conventional methods, coacervation and crosslinking were conducted at 37⁰C. DGHF allows for the fabrication of hydrogels at lower temperatures because the elastin is able to coacervate at lower temperatures such as 16°C. It was found that the variation of processing temperature did not have any significant impact on the pore morphology and swelling ratio of fabricated hydrogels.

Pore Interconnectivity

The pore interconnectivity of hydrogels fabricated at atmospheric and high pressure CO₂ were compared using micro-CT scan analysis. As depicted in Figure 11a, a highly homogenous interconnected-porous structure was detected for σ-elastin hydrogels acquired using high pressure CO₂. In comparison σ-elastin hydrogels produced at atmospheric pressure showed an internal non-uniform porous structure covered by an impermeable layer (Figure 11b). 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, as it believed that their porosity would facilitate oxygen and nutrient transport throughout the structures.

A unique feature of cross-linking elastin under high pressure CO₂ is the formation of micro-channels throughout the 3D structure of the hydrogels as demonstrated in Figure 6(d) and Figure 6(g). These channels were most likely induced during the depressurisation stage. At a microscopic level, the size of the pores fabricated at both atmospheric and high pressure CO₂ were smaller than 15 μm as indicated in Figure 10. 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 with dimensions of 97 ± 21.4 μm was expected to allow for cell infiltration and proliferation into the 3D microstructure. In-vitro fibroblast cell proliferation on elastin hydrogels

Cellular growth and proliferation in σ-elastin hydrogels was 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 σ-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 (Figure 12a) and the internal surfaces of larger channels (Figures 12b-12g) within σ-elastin hydrogels produced under high pressure CO₂.

Conclusions

This example demonstrates the feasibility of fabricating GA crosslinked elastin hydrogels in an aqueous solution using high pressure CO₂. There 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. Consequently, the GA crosslinked σ-elastin hydrogel produced at high pressure CO₂ has high potential for in vitro applications.

Example 5

Cross-linking σ-elastin using hexamethylene diisocyanate (HMDI)

The aim of this example was to increase the mechanical properties and also pore sizes of σ-elastin hydrogel produced at high pressure CO₂ using Hexamethylene diisocyanate (HMDI) as a crosslinking agent. Hexamethylene diisocyanate, a bifunctional molecule, has been employed previously in biomaterials fixation, and its cytotoxicity is considerably lower than glutaraldehyde. In general, isocyanates may react with nucleophilic functional groups such as amines, alcohols, and protonated acids. Hexamethylene diisocyanate in particular was observed to react with the side chains of backbone-protected lysine, cysteine, and histidine, and to a lesser extent tyrosine, in water. Therefore, using HMDI as a crosslinker may increase the mechanical properties of σ-elastin hydrogels by enhancing the degree of crosslinking. HMDI may react with other amino acids available in σ-elastin structure to increase the crosslinking density and mechanical properties of fabricated hydrogels.

HMDI was not soluble in water; therefore, we used dimethylsulfoxide (DMSO) as a media to fabricate crosslinked hydrogel. Using DMSO, it is possible that larger pores may be induced in 3D structure of hydrogel fabricated at high pressure CO₂ due to the higher solubility of CO₂ in DMSO compared to aqueous solution.

Phase behavior studies on binary mixtures of CO₂ and DMSO indicate that solubility of CO₂ in DMSO is a function of temperature and pressure. Increasing the temperature decreases the CO₂ solubility in DMSO. The solubility of CO₂ in DMSO was also enhanced by increasing the pressure at constant temperature. The molar fractions of CO₂ increase from 0.097 to 0.874 when the pressure increased from 10 bar to 60 bar at 25°C [9]. These data underpin our hypothesis that by using DMSO, the solubility of CO₂ in liquid phase is dramatically increased. In our study the operating conditions of 25⁰C and 60 bar were used to achieve higher solubility of CO₂ in liquid phase. The properties of hydrogel fabricated at both high pressure and atmospheric conditions including swelling ratio, pore morphology, and mechanical properties were characterized. In vitro studies were also conducted to assess the cellular growth and proliferation in the 3D structures of fabricated hydrogels.

Hydrogel fabrication using HMDI

At atmospheric conditions 100 mg/ml σ-elastin in DMSO was mixed with HMDI and the solution was immediately pipetted into a glass Lab-Tek chamber slide. The reaction was allowed to react for 18 h at room temperature under nitrogen. The crosslinked hydrogels were then swelled in MiIIiQ water and gently agitated on a shaker for 2 hours to remove the residue of DMSO. The MiIIiQ water was exchanged with fresh MiIIiQ water every 15 min during the shaking. The hydrogels were then stored in PBS for characterisation. A preliminary set of experiments were conducted to determine the required amount of HMDI for the hydrogel fabrication, σ-elastin solutions at 100 mg/ml were mixed with various concentrations of HMDI ranging between 0.25 % (v/v) and 10 % (v/v). The solutions were then pipetted into Lab-Tek chamber slides and placed in a chamber connected to a nitrogen line for 18 hr at room temperature (25°C). The results demonstrated that hydrogels were formed at HMDI concentration above 1 %. At low concentrations of HMDI cr-elastin solution was partially crosslinked, while at high concentrations of HMDI (10%) the hydrogel was very rigid and non-elastic. Consequently, in this study, 2 % (v/v) and 5 % (v/v) HMDI were used to produce hydrogels.

The hydrogel was formed by the DGHF technique as described previously. The hydrogels were then stored in PBS for further analysis. The effects of reaction time, pressure, and crosslinker concentration on the characteristics of the hydrogel produced at high pressure CO₂ were assessed. Preliminary results demonstrated that elastin was not crosslinked at reaction time below 2 hours, when 2 % (v/v) HMDI was used. However, as the crosslinker concentration increased to 5 % (v/v), a hydrogel with desirable pore size and elasticity was obtained within one hour. As the reaction time was further increased from 1 to 2 hours, the 5 % (v/v) HMDI crosslinked hydrogel became more rigid and less elastic. These results are summarised in Table 4.

Table 4. Optimization of GA concentration and reaction time for σ-elastin hydrogel fabrication at high pressure CO2

2 S *

S; Hydrogel was formed, - : no hydrogel formed (partially crosslinked gel which dissolved in water after swelling), *: rigid or non-elastic hydrogel Mechanical characterisation: Compressive properties

Uniaxial compression tests were performed in an unconfined state using a Bose ELF3400 mechanical tester with a 50 N load cell, following a similar procedure as described elsewhere [10-12]. Prior to mechanical testing, the hydrogels produced at high pressure CO₂ and atmospheric condition were swelled for 2 hours in PBS. The thickness (3 mm) and diameter (12.5 mm) of each sample was measured using Digimatic Callipers prior to mechanical testing. The compressive properties of the samples were tested in the hydrated state, in PBS, at room temperature. The data were recorded at a cross speed of 30 μm/s and different strain levels ranging from 40% to 80%. At each of the strain levels, the samples were cyclically preconditioned for 7 cycles to minimize artefact interference. The hydrogels were subsequently subjected to another loading and unloading cycle (8th cycle) where compression (mm) and load (N) were collected. At each strain level, a stress-strain curve was constructed by converting the force-displacement data obtained from the compression test at cycle 8 using the following equation:

σ = (1 )

Δl ε = (2) Io

Where σ is the stress (MPa), F is the force (N), A is the area of the specimen (mm²), e is strain, Δl is the change in length or compression (mm) and I₀ is the original length (mm) of the sample. The compression modulus for the 8th cycle was obtained as the tangent slope of the stress-strain curve. At each strain levels, the energy loss based on 8^th cycle was also calculated as follows: area under loading curve - area under unloading curve

Energy loss = x 100% area under loading curve

Three specimens were tested for each sample type (sample produced at high pressure CO₂ or atmospheric pressure) at each strain levels.

Results and discussions

Visual observation of HMDI crosslinked σ-elastin hydrogels produced both at high pressure and atmospheric conditions demonstrated that the gels were more rigid than GA crosslinked σ-elastin hydrogels fabricated using GA as a crosslinker agent, σ-elastin hydrogels were produced at high pressure and atmospheric condition using 2 % (v/v) HMDI. Although the fabricated gels were easily handled, their water content and elasticity was lower than the GA crosslinked σ-elastin hydrogels.

Pore structure of the α-elastin hydrogel

SEM analysis demonstrated that σ-elastin hydrogels fabricated under high pressure CO₂ (Figure 13(c)-13(f)) were highly porous yet robust structures. Comparison of SEM images of σ-elastin hydrogels produced under either high pressure CO₂ or atmospheric conditions indicated that high pressure CO₂ increased the pore size of the fabricated hydrogels as shown in Figure 13. The presence of large channels in the cross section of sample fabricated at high pressure CO₂ (Figure 13(d)) could facilitate cellular penetration and growth into the 3D structures. These large channels were not present in hydrogels fabricated at atmospheric conditions as shown in Figure 13(b). Equivalent circle diameter (ECD) of the pores was calculated using Image J software. The average pore size of hydrogels fabricated by using 2 % (v/v) HMDI increased from 3.9± 0.8 μm to 79.8 ± 54.8 μm when the pressure was increased from 1 bar to 60 bar. In hydrogels fabricated by the dense gas CO₂, 53 % of the pores were above 80 μm in diameter which make these hydrogels suitable for cellular growth through the 3D structures.

The formation of large pores in the 3D structures of HMDI crosslinked hydrogels produced at high pressure CO₂ was potentially due to the high solubility of CO₂ in DMSO. The binary phase diagram of DMSO and CO₂ indicated that DMSO significantly expanded with CO₂ at 60 bar and 25°C. Therefore, the σ-elastin solution in DMSO could be expanded by CO₂ at 60 bar and 25°C before the crosslinking reaction and gel formation. However, the presence of crosslinker in σ-elastin solution may have an impact on the expansion of DMSO solution by CO₂.

Using 5 % (v/v) HMDI, the number of the large pores on the top and also in the cross section of the sample was reduced as compared to 2 % (v/v) HMDI crosslinked hydrogels (Figures 13(e) and 13(f)). However, using higher concentration of HMDI (5 % (v/v)), the hydrogels were more rigid and may have higher mechanical properties compared to 2 % (v/v) HMDI crosslinked hydrogels. Increasing the crosslinker concentration resulted in an increase in the degree of crosslinking through the hydrogel matrices, which limited the ability of σ-elastin solution to expand in CO₂ and led to the formation of smaller pores.

The effect of pressure on the pore morphology of σ-elastin hydrogels fabricated at high pressure was assessed. Increasing the pressure dramatically reduced the pore size of crosslinked hydrogel using 2 % (v/v) HMDI, as shown in Figure 13(g).

Mechanical characterisation: Compressive properties

The compressive mechanical properties of elastin hydrogel crosslinked by HMDI at high pressure CO₂ was compared with the one fabricated at atmospheric condition in Figure 14. The compression modulus of both samples, produced at high pressure and atmospheric condition, increased as the applied strain magnitude was increased. The compression modulus of the sample produced at atmospheric condition ranged from 11.25 ± 0.4 KPa to 18.8 ± 3.4 KPa at 40 % and 80% strain, respectively (Figure 14c). The stress-strain curve was nonlinear at strain levels above 40% as shown in Figure 14b. This indicated that plastic deformation occurred in the hydrogels produced at atmospheric conditions at strains greater than 40%. However, for the samples fabricated at high pressure CO₂, the strain-stress curve was still linear at 60% strain (Figure 14a). The samples produced at high pressure CO₂ were, therefore, expected to be more elastic than hydrogels produced at atmospheric conditions. The compression modulus of the HMDI crosslinked hydrogel produced by high pressure CO₂ ranged from 3.99 ± 0.53 KPa to 8.62 ± 1.75 KPa at 40% and 80% strain, respectively (Figure 14c). The compression modulus of the hydrogels produced at high pressure CO₂ was generally lower than the one produced at atmospheric condition. This means that the samples produced at atmospheric pressure were stiffer than those produced at high pressure CO₂. This was expected due to the presence of larger pores in the structures of the samples fabricated at high pressure CO₂.

High resilience of elastin provides the ability to deform reversibly without loss of energy. Energy loss is proportional to hysteresis. Using 60 % strain level, the energy loss for the fabricated hydrogel at high pressure CO₂ was 6.03 ± 1.14 %. However, for the gels produced at atmospheric condition the energy loss was 11.72 ± 4.02 %, demonstrating greater hysteresis for the samples produced at atmospheric pressure. As shown in Figure 14d, the energy loss of both hydrogels fabricated at high pressure CO₂ and atmospheric condition increased with increasing the strain levels. The energy loss for sample produced at high pressure CO₂ increased from 1.43 ± 0.86 % to 13.16 ±1.93 %, when the stain level increased from 40 % to 80 % as shown in Figure 12D. For the hydrogel produced at atmospheric pressure, it increased from 4.51 ± 1.54 % to 14.67 ± 3.31 % when the strain level increased from 40 % to 80 %.

The energy loss of elastin hydrogel produced at high pressure CO₂ was comparable with those reported for elastin-like hydrogels fabricated using various crosslinkers.

Swelling properties

The swelling behaviour of fabricated hydrogels is shown in Figure 15. Generally, the swelling ratio of the samples exposed to high pressure CO₂ was higher than the fabricated hydrogels at atmospheric condition as indicated in Figure 15. Both hydrogels produced at high pressure CO₂ and atmospheric pressure swelled more in DMSO than water and PBS. Hydrogels produced at 60 bar CO₂ pressure absorbed 6.81 ± 0.46, 10.9 ± 2.46, and 18.65 ± 0.88 g liquid / g protein when they were hydrated in PBS, water, and DMSO, respectively. However, the gels formed at atmospheric pressure absorbed 4.79 ± 0.15, 9.45 ± 0.25, and 9.82 ± 1.97 g liquid / g protein when they were swelled in PBS, water, and DMSO, respectively. Whilst the exact mechanism of an increase in the swelling ratio of HMDI crosslinked hydrogels in DMSO is unknown, it has been postulated that DMSO can destabilise the secondary structure of elastin and placticise the elastin network by increasing molecular mobility. Therefore, DMSO may plasticize the σ-elastin hydrogel, resulting in an increase in segment length between crosslinks within the HMDI crosslinked gel. The higher swelling ratio of the sample produced at high pressure CO₂ was due to the presence of larger pores through their structures compared to the hydrogel formed at atmospheric pressure.

The swelling behaviour of the HMDI crosslinked σ-elastin hydrogels produced at high pressure CO₂ was 10 fold higher than reported by Nowatzki et al. (0.37 g H₂O / g protein at 4°C).[4] This may be due to the presence of larger pores within the structures of produced hydrogels. These properties are superior for cell infiltration.

The swelling ratio of HMDI crosslinked σ-elastin using 60 bar was lower than the GA crosslinked σ-elastin hydrogel fabricated at 60 bar CO₂, which was 33.2+0.8 g H₂O / g protein as reported in Example 4. This may due to the higher degree of crosslinking through the structures of HMDI crosslinked σ-elastin hydrogel as the swelling ratio of hydrogels is correlated to the degree of crosslinking. Generally, the hydrogels with high degree of crosslinking exhibit low swelling ratio.

The swelling behaviour of the HMDI crosslinked hydrogel was correlated to the compressive modulus. In general, the samples with greater swelling have lower compressive modulus as indicated in Figure 16.

In vitro fibroblast cell proliferation using elastin hydrogels

Cellular growth and proliferation in σ-elastin hydrogels were examined by light microscopy and SEM analysis to demonstrate the feasibility of using the processed material for soft tissue engineering applications. Images of adherent fibroblast cells cultured on hydrogel produced at 60 bar CO₂ from the light microscope are shown in Figure 17. Haematoxylin and eosin was used for staining, as a result the cells appear as dark grey and the σ-elastin scaffolds as light grey. As shown in Figures 17a-17c, fibroblast cells were able to grow into the 3D structures of σ-elastin due to the presence of large pores induced by high pressure CO2. However, cells were only able to form a monolayer on surface of the hydrogel fabricated at atmospheric conditions as indicated in Figure 17d. SEM analysis also confirmed the results obtained by light microscope analysis. Cells were able to attach and proliferate on the top surface of σ-elastin hydrogels fabricated at atmospheric pressure but could not penetrate into the structure due to presence of small pores, as shown in Figure 18a. However, cells were able to colonise at the top surface (Figure 18b and 18c) and also into the 3D structure (Figure 18d-18f) of σ-elastin hydrogels produced under high pressure CO₂ due to the presence of large channel within the materials.

Conclusions

This example demonstrates the feasibility of fabricating elastin-based hydrogels with enhanced mechanical properties and pore sizes using DGHF technique and HMDI as a crosslinking agent. The fabricated hydrogels had superior characteristics compared to the one fabricated using GA crosslinker. HMDI crosslinked σ-elastin hydrogels produced by DGHF technique had larger pores compared to GA crosslinked hydrogels due to higher solubility of CO₂ in DMSO than in an aqueous solution. The fabrication of these large pores within the 3D structure substantially promoted fibroblast infiltration and growth throughout the matrices. The mechanical properties was appeared to be promoted as a result of the use of HMDI.

Example 6

Cross-linking elastin-tropoelastin composite using GA

The aim of this part of the study was to promote the mechanical properties of σ-elastin hydrogels by addition of recombinant human tropoelastin (TPE). Our previous data demonstrated that crosslinking of σ-elastin with GA using DGHF process increased the mechanical integrity of the hydrogel. As σ-elastin has a low amount of lysine residue (less than 1%) in its structure for the crosslinking with GA, our hypothesis was to add TPE with high level of lysine residues in order to increase the crosslinking density and improve the mechanical properties of the hydrogel. We used a 50:50 weight ratio of TPE/σ-elastin for crosslinking with GA and assess the feasibility of enhancing the mechanical properties of fabricated hydrogel. The effect of high pressure CO₂ on the properties of fabricated hydrogels including swelling ratio, mechanical properties, and porosity was also investigated.

TPE/a-elastin hydrogel fabrication using GA

A preliminary set of experiments were conducted to determine the required ratio of GA and TPE/σ-elastin for the hydrogel fabrication. The concentration of TPE/σ-elastin was varied between 5 mg/ml and 100 mg/ml, and GA between 0.05 and 0.5 % (v/v). The solutions were pipetted into a Lab-Tek chamber slide and then placed at 37°C oven for 24 hr. It was found that the hydrogels were formed when GA and TPE/σ-elastin concentration were above 0.1 % (v/v) and 50 mg/ml, respectively. At low concentrations of GA and σ-elastin, soft films of cross-linked TPE/σ-elastin were formed at the bottom of the Lab-Tek chamber slides due to insufficient degree of cross-linking, while at high concentrations rigid and non-elastic hydrogels were formed due to high degree of crosslinking. Consequently, in this study 100 mg/ml of protein solution and 0.25 % (v/v) GA were used to produce hydrogels. The results demonstrated that addition of TPE had a significant impact on the mechanical properties of σ-elastin hydrogel.

Swelling properties

The swelling ratio of the fabricated hydrogel fabricated from σ-elastin/TPE mixture was in the range of 5-6 g PBS / g protein as shown in Figure 19. The swelling ratio of sample exposed to high pressure CO₂ was slightly lower than the fabricated hydrogels at atmospheric condition as indicated in Figure 18. Hydrogels produced at 60 bar CO₂ pressure absorbed 5 + 1.1 g liquid / g protein when they were hydrated in PBS at 37°C. However, the gels formed at atmospheric pressure absorbed 5.9 ± 1.3 g liquid / g protein when they were swelled in PBS at 37°C. The lower swelling ratio of the samples produced at high pressure CO₂ may be due to the higher degree of crosslinking in their structures compared to the hydrogel formed at atmospheric pressure. High pressure CO₂ facilitated the coacervation and expedited the crosslinking of TPE/σ-elastin solution. The swelling behaviour of the GA crosslinked TPE/σ-elastin hydrogels produced at high pressure CO₂ was also lower than the swelling ratio of GA crosslinked σ-elastin hydrogels previously reported 7 ± 3.2 g PBS / g protein at 37°C. The presence of TPE with high level of lysine residues in protein solution could increase the crosslinking density which resulted in a reduction in the swelling ratio of TPE/σ-elastin hydrogel.

Pore structure of the TPE/a-elastin hydrogel

Effect of crosslinker concentration on hydrogel pore morphology:

The effect of GA concentration on the pore morphology of TPE/σ-elastin hydrogels fabricated at atmospheric pressure was assessed using SEM analysis. Increasing the concentration of GA at atmospheric condition slightly decreased the pore sizes. At atmospheric pressure when the concentration of GA was enhanced from 0.1 to 0.5 % (v/v), the pore size of the TPE/σ-elastin hydrogel slightly diminished as shown in Figure 20. Increasing the crosslinker concentration resulted in an increase in the degree of crosslinking which may lead to a pore-size reduction. Therefore, large pore sizes were observed when a lower concentration of crosslinker was used.

Effect of pressure on hydrogel pore morphology

SEM analysis was also used to study the effect of high pressure CO₂ on the pore morphology of the fabricated TPE/σ-elastin hydrogels. Comparison of SEM images of TPE/σ-elastin hydrogels produced under either high pressure CO₂ or atmospheric conditions indicated that high pressure CO₂ increased the pore size of the fabricated hydrogels as shown in Figure 21. The presence of large pores in both top surface (Figures 21a and 21 b) and cross section (Figures 21c and 21 d) of the sample fabricated at high pressure CO₂ could facilitate cellular penetration and growth into the 3D structures. However, these large pores were not available in the hydrogels fabricated at atmospheric conditions as shown in Figures 21 e - 21 h The skin-like formation on the top surface of TPE/σ-elastin hydrogel produced at atmospheric condition (Figures 21 g and 21 h) could also prevent the cellular growth in the 3D structures of hydrogel. High pressure CO₂ significantly increased the pore size of GA crosslinked TPE/σ-elastin hydrogel and also eliminate the skin-like formation on the top surface of the hydrogel. Conclusions

The outcome of the third part of our study demonstrates that it is feasible to promote the characteristics of GA crosslinked σ-elastin hydrogel crosslinked by addition of TPE to the system. The elasticity and mechanical properties were significantly enhanced. This effect is probably due to the presence of more lysine residues in TPE structures. Using high pressure CO₂, the pore sizes of produced hydrogels were significantly increased.

It will be understood that the invention disclosed and defined in this specification extends to all alternative combinations of two or more of the individual features mentioned or evident from the text or drawings. All of these different combinations constitute various alternative aspects of the invention.

It will also be understood that the term "comprises" (or its grammatical variants) as used in this specification is equivalent to the term "includes" and should not be taken as excluding the presence of other elements or features.

References

[1] Gosline JM, French CJ. Dynamic mechanical properties of elastin. Biopolymers 1979;18(8):2091-2103.

[2] Mithieux SM, Rasko JEJ, Weiss AS. Synthetic elastin hydrogels derived from massive elastic assemblies of self-organized human protein monomers. Biomaterials 2004;25(20):4921-4927.

[3] Mithieux SM. Synthetic elastin: construction and properties of cross-linked human tropoelastin. PhD thesis, University of Sydney, Sydeny, NSW 2003.

[4] Nowatzki PJ, Tirrell DA. Physical properties of artificial extracellular matrix protein films prepared by isocyanate crosslinking. Biomaterials 2003;25(7-8):1261 -1267.

[5] Lim DW, Nettles DL, Setton LA, Chilkoti A. Rapid Cross-Linking of Elastin-like Polypeptides with (Hydroxymethyl)phosphines in Aqueous Solution. Biomacromolecule 2007;8(5): 1463-1470.

[6] Leach JB, Wolinsky JB, Stone PJ, Wong JY. Crosslinked alpha-elastin biomaterials: towards a processable elastin mimetic scaffold. Acta Biomater 2005;1(2):155-164.

[7] Goncalves CA, Figueiredo MH, Bairos VA. Three-dimensional organization of the elastic fibres in the rat lung. Anat Rec 1995;243(1):63-70. [8] Vesely I. The role of elastin in aortic valve mechanics. J Biomech 1998;31(2):115-123.

[9] Kordikowski A, Schenk AP, Van Nielen RM, Peters CJ. Volume expansions and vapor-liquid equilibria of binary mixtures of a variety of polar solvents and certain near- critical solvents. J Supercrit Fluids 1995;8(3):205-216.

[10] Srokowski EM, Woodhouse KA. Development and characterisation of novel cross-linked bio-elastomeric materials. J Biomater Sci Polym Ed 2008;19(6):785-799.

[11] Stammen JA, Williams S, Ku DN, Guldberg RE. Mechanical properties of a novel PVA hydrogel in shear and unconfined compression. Biomaterials 2001 ;22(8):799-806.

[12] Joshi A, Fussell G, Thomas J, Hsuan A, Lowman A, Karduna A, et al. Functional compressive mechanics of a PVA/PVP nucleus pulposus replacement. Biomaterials 2006;27(2):176-184.

Claims

CLAIMSClaims defining the invention are as follows:

1. A method for producing a hydrogel including:

- providing a vessel including a liquid phase, the liquid phase including elastin material;

- pressurising the vessel with a gas to dissolve the gas into the liquid phase;

2. The method of claim 1 wherein the liquid phase having coacervated elastin material is provided by steps including:

-providing a vessel including a liquid phase, the liquid phase in the form of a solution of elastin material and

3. The method of any one of the preceding claims wherein the liquid phase is provided with a cross linking agent for cross linking elastin material in the liquid phase before or after the vessel is pressurised, or after it is depressurised.

4. The method of any one of the preceding claims wherein the liquid phase is provided with a cross linking agent to cross link coacervated elastin material in the liquid phase before the vessel is pressurised.

5. The method of any one of the preceding claims wherein the liquid phase is provided with a cross linking agent to cross link coacervated elastin material in the liquid phase after the vessel has been pressurised.

6. The method of any one of the preceding claims wherein the step of cross linking the coacervated elastin material after pressurisation of the vessel entraps gas molecules dissolved in the liquid phase within the cross linked elastin material.

7. The method of claim 6 wherein the step of depressurisation provides conditions for release of gas entrapped within the cross linked elastin material.

8. The method of claim 7 wherein the release of gas entrapped within the cross linked elastin material provides the hydrogel formed from the cross linked elastin material with one or more pores.

9. The method of claim 8 wherein the one or more pores form one or more conduits that extend throughout the hydrogel or ramify to form networks throughout the hydrogel.

10. The method of any one of the preceding claims wherein the cross linker is provided in the liquid phase in an amount of from about 0.05 to about 10 percent (v/v).

11. The method of any one of the preceding claims wherein the cross linker is an enzyme or chemical reagent.

12. The method of claim 11 wherein the chemical agent is an amine-reactive chemical cross linker or amine and carboxyl reactive cross linker.

13. The method of any one of the preceding claims wherein the cross linker is GA and is provided in an amount of from 0.05 to about 2 percent (v/v).

14. The method of any one of the preceding claims wherein the cross linker is HMDI and is provided in the liquid phase in an amount of from about 0.25 to about 10 percent (v/v).

15. The method of any one of the preceding claims wherein conditions for coacervation of elastin material in the solution are provided at the time that the vessel is pressurised with a gas to dissolve the gas into the liquid phase.

16. The method of any one of the preceding claims wherein the vessel is pressurised with CO2 or N₂ gas.

17. The method of any one of the preceding claims wherein the gas is a dense gas.

18. The method of any one of the preceding claims wherein the gas is a supercritical fluid.

19. A hydrogel including:

- a scaffold of cross linked elastin material molecules

- water molecules bound to the scaffold

20. The hydrogel of claim 19, including one or more pores forming one or more conduits that extend throughout the hydrogel or ramify forming networks throughout the hydrogel.

21. The hydrogel of any one of the preceding claims wherein the elastin material for use in producing the hydrogel is cross linked before coacervation.

22. The hydrogel of any one of the preceding claims wherein the elastin material is one obtained by extraction from a tissue source.

23. The hydrogel of any one of the preceding claims wherein the elastin material is α-elastin.

24. The hydrogel of any one of the preceding claims wherein the elastin material is not cross linked before coacervation.

25. The hydrogel of any one of the preceding claims wherein the elastin material is or includes tropoelastin.

26. The hydrogel of any one of the preceding claims wherein the liquid phase further includes one or more further biomolecules.

27. The hydrogel of claim 26 wherein the biomolecule is a proteins, sugar or lipid.

28. The hydrogel of claim 27 wherein the protein is a connective tissue protein or extra cellular matrix protein.

29. The hydrogel of claim 28 wherein the protein is a collagen.

30. The hydrogel of claims 26 wherein the biomolecule is an extracellular carbohydrate, such as such as a glycosaminoglycan, hyaluronic acid or like substance.

31. The hydrogel of any one of the preceding claims wherein the liquid phase further includes a bioactive protein such as a hormone, growth factor or cytokine.

32. The hydrogel of any one of the preceding claims wherein the liquid phase further includes a pharmaceutical compound.

33. The hydrogel of claim 32 wherein the pharmaceutical compound is an anti cancer compound, such as anti angiogenic compound or a compound useful for tissue repair or regeneration.

34. The hydrogel of any one of the preceding claims wherein the hydrogel is produced by a method of any one of the preceding claims.

35. An apparatus for forming a hydrogel, the apparatus including:

- a vessel containing including a liquid phase, the liquid phase including coacertaved elastin;

- injection means for injection of a cross linking agent for cross linking the coacervated elastin material into the liquid phase; and optionally - input means for input of one or more components to be provided in a hydrogel formed by the apparatus, the one or more components selected from the group consisting of a protein, a sugar, a lipid, a cell and a pharmaceutical.

36. The apparatus of claim 35 wherein the injection means is adapted to automatically inject the cross linking agent when the vessel reaches a pre-selected pressure.