Printer Friendly

Effect of photoinitiator concentration on the properties of polyethylene glycol based hydrogels for potential regenerative medicine applications.


Hydrogels are a key group of biomaterials that have recently come to the forefront of tissue engineering research. They have been utilised to support and assist restoration of a range of tissues, such as bone [1], cartilage [2], nerves [3], vessels [4] and skin [5]. Many hydrogels are biocompatible in nature [6]. Photopolymerised hydrogels have recently gained increased attention in biomedical applications because aqueous macromer solutions can be delivered in a minimally invasive manner and photocrosslinked rapidly within seconds in situ following brief exposure to ultraviolet (UV) light [7, 8]. UV-photopolymerisation is the most commonly applied method due to its distinct advantages of rapid cure, low curing temperature, in-line production, the absence of organic/inorganic solvents and low energy requirement. Photopolymerisation is favoured because hydrogels can be synthesised at temperatures and pH conditions near physiological conditions and even in the presence of biologically active materials. Furthermore, photopolymerisation can be easily controlled by adjusting the dosage and intensity of UV light, the curing temperature and the curing time [9].

Previous work undertaken in our laboratory involved the synthesis of a variety of hydrogels from monomeric precursors [6, 10, 11, 12]. However, little research has been conducted on the effect of the photoinitiator concentration on the thermal behaviour and mechanical performance of the hydrogel scaffolds.

Photopolymerisation schemes generally use a photoinitiator that has high absorption at a specific wavelength of light to produce radical initiating species. Other factors that should be considered in selecting the photoinitiator include its biocompatibility, solubility in water, stability and cytotoxicity [13]. Biocompatible photopolymerising polymers for biomedical and tissue engineering applications have the potential to reduce the invasiveness and cost of biomaterial implants designed to repair or augment tissues. Irgacure 2959 was chosen as a photoinitiator due to its enhanced biocompatibility [14] and the potential tissue engineering application of these hydrogels. Irgacure 2959 is hydrophobic due to its polar hydroxyl end groups and minimises oxygen inhibition which also promotes biocompatibility and was previously reported to have minimal toxicity over a broad range of mammalian cell types and species ranging from human foetal osteoblasts to bovine chondrocytes [1]. In this study the photoinitiator concentration was varied from 0.01-1wt% to ascertain if the concentration of photoinitiator would have any impact on the thermal behaviour and mechanical properties of the hydrogels.


Material selection:

The materials and methods used in this study were selected whilst keeping in mind the end application being biomedical; therefore the choice of biocompatible materials was of paramount importance. The macromolecular monomer, poly(ethylene glycol) dimethacrylate (PEGDMA), with molecular weights (Mw) 400, 600 and 1000was obtained from PolySciences. The photoinitiator used was 4- (2-hydroxyethoxy)phenyl-(2-hydroxy-2- propyl)ketone (Irgacure 2959) supplied by Ciba Specialty Chemicals. All materials were used as received.

Hydrogel synthesis:

UV light can interact with light-sensitive compounds called photoinitiators to create free radicals that can then initiate polymerisation to form crosslinked hydrogels [13]. Initial studies were carried out to determine the effect of the disparity in mechanical properties and thermal behaviour of these hydrogels by varying the photoinitiator concentration and molecular weight of PEGDMA whilst keeping the polymer concentration constant (50wt%). The precursor compositions can be seen in Table 1; they were photopolymerised using a UV curing system (Dr. Grobel UV-Electronik GmbH). This particular irradiation chamber is a controlled radiation source with twenty UV-tubes that provide a spectral range of between 315-400nm at an average intensity of 10-13.5mW/[cm.sup.2] The pre-polymerised mixtures were prepared by combining specified concentrations of the macromolecular monomer and distilled water with varying concentrations (0.01- 1wt%) of photoinitiator (Table 1). These parameters are within the cytocompatible initiating conditions [15]. The batches were placed in a 50mL beaker, mixed using a magnetic stirrer and sonicated for one hour until a homogenous mixture was achieved. The solutions were pipetted into silicone moulds and photopolymerisation was carried out for 10-mins, after which time gelation had occurred.

Characterisation techniques:

Attenuated total reflectance Fourier transform infrared spectroscopy:

Attenuated total reflectance Fourier transform infrared spectroscopy (ATR-FTIR) was carried out on round samples to determine the bonds present in the material on a Perkin Elmer Spectrum One fitted with a universal ATR sampling accessory. All data was recorded at 21[degrees]C, in the spectral range of 4000-650 [cm.sup.-1], utilising a 16 scan per sample cycle and a fixed universal compression force of 80 N. Subsequent analysis was carried out using Spectrum software.

Preparation of aqueous salts and pH buffer solutions:

Potassium chloride 0.2M (KCl) and monobasic potassium phosphate 0.2M (K[H.sub.2]P[O.sub.4]) were used to prepare the buffer pH 7.4. Hydrochloric acid and sodium hydroxide solutions were used to adjust the ionic strength of the solutions to 0.2M. The buffer solutions were prepared in the laboratory and filtered under vacuum using a Millipore filtration apparatus. The pH of the buffer solutions was measured using a Jenway 3520 pH meter.

Swelling studies:

Swelling experiments were performed on samples in buffer solution (pH 7.4). After photopolymerisation, samples with an average weight of 0.8[+ or -]0.1g were placed into a Petri dish. The Petri dish was filled with 30ml of pH 7.4 buffer solution and left at room temperature. The percentage swelling of the samples was calculated using the formulain Equation 1 [16]:

Swelling (%) =([W.sub.s]-[W.sub.f]/[W.sub.d]) * 100, (Eq. 1)

where [W.sub.s] and [W.sub.d] are the weights of the hydrogels in the swelling state and the dried state, respectively. Tests were carried out in triplicate and data is presented as mean [+ or -] SD.

Gel fraction measurement:

The gel fraction of all batches was measured in triplicate using circular discs with an average mass of 0.8 [+ or -] 0.1g. After photopolymerisation the hydrogel samples were dried under vacuum at 200mmHg for 24 hours at 80[degrees]C to a consistent weight and their apparent dry weights ([W.sub.d]) were measured. The samples were then allowed to swell in a sealed Petri dish with 30mL of buffer solution pH 7.4 for 72 hours at 21[degrees]C until equilibrium swelling was achieved. The solution was replaced with fresh buffer solution daily to ensure that the soluble fraction was completely extracted. Once equilibrium swelling was attained, samples were again dried in the vacuum oven at 80[degrees]C in the absence of water until no change in weight was observed. Gel fraction (%) was calculated using Equation 2:

Gel fraction (%) = ([W.sub.ex]/[W.sub.o]) * 100, (Eq. 2)

where [W.sub.o] and [W.sub.ex] are the weight of the dried hydrogel after photopolymerisation and the dried weight of the sample after extraction of soluble parts (i.e. the swelling state and the dried state) respectively. Data is presented as mean [+ or -] SD.

Differential scanning calorimetry:

Differential scanning calorimetry (DSC) was carried out using a TA Instruments 2010 DSC. Xerogels (dehydrated form of hydrogel) of between 8-10mg were weighed out using a Sartorius scale at a resolution of 1 x [10.sup.-5]g. These xerogels were dried in a vacuum oven (24 hours, 80[degrees]C) prior to testing. All measurements were conducted in sealed non-hermetic aluminium pans, which were crimped before testing, with an empty aluminium pan being used as the reference cell. Samples were cooled to - 110[degrees]C using liquid nitrogen and heated to 300[degrees]C at a rate of 10[degrees]C/min. The glass transition temperature was considered as the mid-point temperature of the endothermic drift in the heating curves, in all cases. Volatiles were removed from the purging head using nitrogen at a rate of 30ml/min. Calibration of the instrument was performed using indium as standard. Analysis was carried out in triplicate and data is presented as mean [+ or -] SD.

Rheological measurements:

In all cases throughout this study, rheological measurements were performed using an Advanced Rheometer AR1000 (TA Instruments), to investigate the comparative strength of the samples. Dynamic strain sweep test experiments were performed at a constant frequency of 1Hz with percentage strain ranging from 1.80 x [10.sup.-4] to 1.0 x [10.sup.-3]. Samples were tested in triplicate (using individual samples) within 72 hours of preparation at a temperature of 37[degrees]C using a 40cm parallel steel plate where the samples were in the equilibrium swollen state. Prior to testing, all samples were blotted free of water using filter paper in an attempt to minimise slippage. A compression load of 5 [+ or -] 0.2N was exerted on the samples during testing and the mean [+ or -] SD was reported. Rheological test parameters, storage/elasticity (G') and loss (G") modulus were obtained under dynamic conditions for these non-destructive oscillatory tests.

Uniaxial tensile testing:

Tensile strain to failure tests were performed on a Lloyd Lr10K Plus Series twin column materials testing apparatus using a 3kN load cell. The machine was connected to a control computer with Nexygen[TM] software, to run appropriate testing criteria and for analysis of results. Hydrogel samples were moulded into dumbbell shape in accordance to ASTM Standard Method D638-04, and the length, breath and thickness of each sample recorded. After photopolymerisation, hydrogels were swelled to equilibrium and vacuum dried for 2 hours before testing. This swelling and dehydration step was carried out to remove unreacted low molecular weight polymer from the surface so as to improve the adhesion between the sample and grippers. Fixed grips were mounted onto the tensile testing machine and a crosshead speed of 20mm/min was used. Sticky tape was placed between the hydrogel samples and the grip surface to prevent slippage during loading. Tensile testing was performed on 10 separate test specimens for each batch and mean stress at break; Young's modulus and strain at break values were calculated and data is presented as mean [+ or -] SD.

In vitro cytotoxicity assessment:

The in vitro biocompatibility assessment of hydrogel samples was performed using a mouse preosteoblastic (MC3T3-E1) cell line. Cytotoxicity tests were performed in accordance with ISO 10993-5 (2009) guidelines for extract dilution with MTT as an endpoint. Samples were incubated in cell media at physiological temperature for 24 hours and resultant extracts tested for cytotoxicity using MTT as an endpoint. All cultures and manipulations of cells were performed within a Class II laminar flow unit in a dedicated cell culture laboratory. To ensure an adequate level of sterility, aseptic technique was strictly adhered to in accordance with standard procedures. Tissue culture plastic ware was utilised at all times to ensure sterility while carbon dioxide (C[O.sub.2]) incubators and laminar flow units were periodically cleaned with 70% ethanol to minimise the risk of contamination.


Hydrogel formulation:

In this study chemically crosslinked PEGDMA hydrogels were synthesised by means of photopolymerisation. The pre-polymerised mixtures were prepared by combining a specified amount of PEGDMA (Mw 400, 600, 1000) and distilled water whilst varying the concetration of photoinitiator (0.011wt%), as shown in Table 1. On observation all the pre-polymerised mixtures appeared homogeneously mixed after 1 hour. The hydrogels once cured appeared colourless and transparent. Additionally varying the photoinitiator concentration had no observed impact on the appearance of the cured hydrogel samples.

Attenuated total reflectance Fourier transform infrared spectroscopy:

In this study PEGDMA based hydrogels with varying photoinitiator concentration were characterised using ATR-FTIR. Only samples in the dehydrated state were analysed; prior to testing samples were dried in a vacuum oven at 80[degrees]C for 24 hours. Figure 1 illustrates that the characteristic bonds for the dimethacrylate groups (C-H, C=C and C-O) have disappeared in the ATR-FTIR spectra for the P600-0.01, P600-0.1 and P600-1 hydrogels respectively when compared with the macromolecular monomer PEGDMA600. Killion et al. [6] reported that polymerisation had occurred with similar materials by the disappearance of peaks at 815cm\1167cm-1and 1637cm-1. For biomedical applications this confirmation of polymerisation is essential as most of the problems associated with hydrogel biocompatibility are as a consequence of unreacted monomers, oligomers and other chemical agents that leach out during application. The cytotoxicity results further confirmed there was no unreached monomer after photopolymerisation of the hydrogels as cell viability was at an acceptable level for biomedical applications as discussed in Section 3.7.1. The ATR-FTIRs illustrated in Figure 1 substantiate that varying the photoinitiator concentration has little effect on the disappearance of peaks on the surface of the hydrogels. The spectra for the other molecular weight samples exhibited a comparable trend (data not shown).

Differential scanning calorimetry:

In this study a single glass transition temperature ([T.sub.g]) was observed for each xerogel irrespective of the formulation used. Figure 2 illustrates the thermograms for P600-0.01 to P600-1 which shows minimal difference in [T.sub.g] as the photoinitiator concentration increased. However as the molecular weight of the macromolecular monomer increased so too did the [T.sub.g](Figure 3). The trend which illustrated considerable changes in [T.sub.g] values resulting from the different molecular weight PEGDMAs used during the photopolymerisation were expected, as with increasing molecular weight the length of the polymer chains are increased resulting in a more flexible material [17]. A lower [T.sub.g] allows for greater polymer chain mobility thus producing more flexible material [18] and subsequently a weaker material, hence the PEGDMA based hydrogels prepared with the 1000 molecular weight macromolecular monomer would be expected to be weaker than the samples prepared using a 400 molecular weight macromolecular monomer.

Uniaxial tensile testing:

The effect of varying photoinitiator concentrations on the mechanical properties were compared using three different molecular weights of the PEGDMA monomer. For all three molecular weight macromolecular monomers the tensile strength decreased with an increase in photoinitiator concentration as illustrated in Figure 4. The hydrogels characterised in this study had a tensile strength of 1.21- 2.98MPa. Such hydrogels would not be suitable for load bearing bone applicationswhere atensile strength range of 14-40MPa is required [19]. However they would be suitable for applications where a mechanical strength range of 1-3MPa is required such as human blood vessels [20], as barriers following tissue injury in order to improve healing response [13], as the basement membrane which is positioned between epithelial cells and connective tissue [21] and potentially as scaffolds for development and regeneration of soft tissues such as cartilagewhere a tensile strength of 0.5-1MPa is required [22].

These results corroborate that as the photoinitiator concentration increased the tensile strength decreased which is well aligned with swelling results (Section 3.5) where the equilibrium swelling increased with an increase in photoinitiator concentration. The samples containing the lowest molecular weight macromolecular monomer, PEGDMA400, revealed the highest tensile strength and conversely the samples containing the highest molecular weight macromolecular monomer, PEGDMA1000, had the lowest tensile strength. These results highlight how manipulation of the mechanical properties of the hydrogels is possible through variation of the photoinitiator concentration and molecular weight of the macromolecular monomer, thus expanding their suitability for a wide range of tissue engineering applications.


When a hydrogel is brought into contact with solvent, the solvent diffuses into the material and the hydrogel swells. Diffusion involves the migration of solvent into pre-existing spaces among hydrogel chains. The sample continues to absorb the solvent until it reaches a state of equilibrium. Equilibrium is accomplished when the osmotic pressure from the swelling and the elasticity of the polymer network are equal [17]. Swelling analysis was selected because if any variation in efficiency of formation of the network has occurred the swelling ratio will most likely be affected. The uptake of solvent into the hydrogel matrix occurred very rapidly during the initial stages of the swelling profile with equilibrium swelling being reached within 48 hours for all hydrogel batches. The effect of photoinitiator concentration on the swelling percentage of the PEGDMA600 hydrogels is reported in Figure 5.

Overall the hydrogels investigated in this study showed an increase in swelling percentage overtime for all photoinitiator concentrations. This sequence of results followed the same trend for each of the three molecular weight PEGDMA samples analysed. For P400-0.01 to P400-1 samples the maximum equilibrium swelling reached was 24.27%; for P600-0.01 to P600-1 samples it was 26.75% whereas for P1000-0.01 to P1000-1 samples it was 46.48%. The chemical structure of the polymer may greatly affect the swelling ratio of the hydrogels. Hydrogels containing hydrophilic groups swell to a higher degree compared to those containing hydrophobic groups. Hydrophobic groups collapse in the presence of water, thus minimising their exposure to the water molecule. As a result, the hydrogels will swell much less compared to hydrogels containing hydrophilic groups [23]. The higher the molecular weight of the macromolecular monomer the greater the swelling ratio due to the longer polymer chains and hence larger pore size that subsequently increased swelling. The equilibrium swelling increased with an increase in photoinitiator concentration, which can be linked to the crosslink density as the lower the crosslink density the lower the photoinitiator concentration hence the lower the swelling percentage. This sequence of results followed the same trend for the three batches of PEGDMA hydrogels analysed. Figure 6 illustrates the increase in swelling percentagefor the three molecular weights of the macromolecular monomer containing 0.1% photoinitiator. This highlights a consistent trend across the three molecular weights of the macromolecular monomer, with a gradual increase in swelling. Ideally a photoinitiator which will give optimum curing with the least impact on the biocompatibility of the hydrogel is required.


Parallel plate rheometry was carried out at 37[degrees]C on equilibrium swollen crosslinked hydrogels to investigate the comparative strength of the samples. The AR 1000 rheometer was used to measure the stress-strain relationship of the hydrogels. For all tests, shear storage modulus (G') values were greater than the shear loss modulus (G") due to the elastic response dominating, which is commonplace for hydrogels and solid like materials [6]. The storage modulus follows the thermodynamic behaviour and is governed by the entropy of the system [24].

Table 2 shows the rheological results for all hydrogel batches analysed. For all hydrogels analysed the storage modulus increased with an increase in photoinitiator concentration. This finding was more pronounced with the lowest molecular weight monomer (PEGDMA400) resulting in the highest G' and widest range from roughly 10,000 to 66,000Pa, which can be related to the lower water update of the gels as seen in the swelling results. The gel strength for the PEGDMA600 hydrogels ranged from approximately 7,000 to 15,000Pa (as shown in Figure 6) whereas the PEGDMA1000 hydrogels had a storage modulus range of about 10,000 to 34,000Pa. Similar findings were reported by Killion et al. [17] for variation of molecular weight. An increase in G' can be related to an increase in photoinitiator concentration resulting in hydrogels with reduced flexibility and hence an increase in brittleness. Additionally increasing the degree of crosslinking of the system will result in a stronger gel [23]. However a higher degree of crosslinking creates a more brittle structure [25].

The gels containing the lowest photoinitiator concentration exhibit poorer mechanical strength than those with higher photoinitiator concentration which is in direct contradiction with the tensile results previously discussed. The trend observed in this study leads to a hypothesis that as the photoinitiator concentration increased the level of crosslinks in the gel increased proportionally resulting in a stronger but more brittle (i.e. less elastic) structure, as proposed by Lyons et al. [25]. The tensile and swelling results have already shown how changing the photoinitiator concentration and molecular weight of the macromolecular monomer can be utilised as a means of achieving different mechanical properties of the hydrogels. The storage modulus results further corroborate this. These values are in line with the gel strength reported for other synthesised hydrogels for potential tissue engineering applications. Wang et al. [26] used a gelatin scaffold and it obtained gel strength of approximately 15,000Pa. Other PEG hydrogel based composites exhibit similar values ranging between 8,000-14,000Pa [27]. Huang et al. [28] used a chitosan/nHAP/collagen gel for potential use in bone tissue engineering and it obtained gel strength of approximately 10,000Pa.

Cytotoxicity testing:

Cytotoxicity tests are performed to act as reliable and reproducible screening methods and represent the initial phase in testing biocompatibility of biomaterials [29]. Preceding in vivo preclinical testing on animals, the materials must exhibit good biocompatibility. Therefore the evaluation of biocompatibility and cytotoxicity of hydrogels is a critical step in the development of a material for tissue engineering applications. Since any unreacted residues and by-products (unreacted monomers and initiators) from the radical polymerisation reaction can be critical to cell viability, the cytotoxicity of any substance leached from the crosslinked hydrogels was determined. Hydrogel materials intended for implantation will first exert their effect at a cellular level; results from in vitro toxicity tests can therefore predict immediate toxic response which can then be extrapolated for prediction of potential systemic response [30].

Even though polyethylene glycol is biocompatible and FDA approved, toxicological assessment is still essential to ensure that the formulation and photopolymerisation process does not alter the properties of the material. The purpose of the MTT assay system is to obtain knowledge concerning the cytotoxic potential of the polymers. This test detects the level of toxicity the sample exerts by assessing mitochondrial activity.

MTTAssay(3-(4, 5-dimethylthiazol-2yl)-2, 5-diphenyl-2 tetrazoliumbromid):

Many hydrogels are biocompatible in nature, resulting in minimal inflammatory responses and tissue damage [31]. Toxicological analysis however has shown that such hydrogels require a number of time consuming washing steps to ensure cytocompatibility. This is mainly a result of the high photoinitiator concentration and the presence of minute levels of unreacted monomer following the photopolymerisation process [32]. The yellow tetrazolium salt (MTT) is reduced by mitochondrial dehydrogenases in living cells to a blue-magenta coloured formazan precipitate [33]. The absorption of dissolved formazan in the visible region correlates with the number of viable metabolically active cells [34].

In this study, chemically crosslinked hydrogel samples were assessed and the results obtained from the absorbance readings were transformed as percentage of untreated control cell values. The mean percentage viability of replicate results and the SEM were calculated for each cell line. Based on this testing the results for the MTT viability assessment MC3T3-E1 cells following 24 hour exposure to diluted and undiluted extracts of hydrogel samples are shown in Figure 8. These hydrogels showed minimal adverse effects on the MC3T3-E1 cells. Cell viability was greater than 75% for the hydrogels tested at the different extract concentrations. The results indicate that even at the highest photoinitiator concentration, 76% of MC3T3-E1 cells remained viable at 24 hours when compared to the blank control. Incorporation of reduced concentration of photoinitiator also led to a slight decrease in viability. However 80% of the cells remained viable indicating that these hydrogels exert minimal toxicity to MT3T3-E1 cells and therefore are good candidates to be used as base materials for implants and may find utilisation in various biomedical applications.


Hydrogels are attractive biomaterials for numerous medical applications where the mechanical performance and biocompatibility of the material are essential. During this study the photoinitiator concentration was varied to determine if this would impact on the thermal and mechanical properties of the hydrogel. The hydrogel properties were analysed by varying the photoinitiator concentration from 0.01- 1wt%. The mechanical properties of the PEGDMA hydrogels were controlled by varying the macromolecular monomer molecular weights prior to photopolymerisation. The results illustrate that as the photoinitiator concentration increased the tensile strength decreased as the swelling and rheological values increased. Additionally as the molecular weight of the macromolecular monomer increased there was an increase in the glass transition temperature with an increase in swelling, increase in storage modulus and decrease in tensile strength. This was expected, as with an increase in the molecular weight of the macromolecular monomer the length of the polymer chains are increased resulting in larger pore size, which subsequently resulted in a more flexible material but weaker hydrogel.

From the cytotoxicity testing, although the hydrogels with the lowest photoinitiator concentration have a higher cell viability than those hydrogels with the highest photoinitiator concentration, all hydrogel batches analysed resulted in favourable MC3T3-E1 cell viability and may hold promising potential for further applications in tissue engineering and regenerative medicine.


Article history:

Received 25 September 2014

Received in revised form 26 October 2014

Accepted 25 November 2014

Available online 31 December 2014


[1] Williams, J.M., A. Adewunmi, R.M. Schek, C.L. Flanagan, P.H. Krebsbach, S.E. Feinberg, S.J. Hollister, S. Das, 2005. Bone tissue engineering using polycaprolactone scaffolds fabricated via selective laser sintering, Biomaterials, 26(23): 4817-4827.

[2] Mercier, N.R., H.R. Costantino, M.A. Tracy, L.J. Bonassar, 2005. Poly(lactide-co-glycolide) microspheres as a moldable scaffold for cartilage tissue engineering, Biomaterials, 26: 1945- 1952.

[3] Lietz, M., L. Dreesmann, M. Hoss, S. Oberhoffner, B. Schlosshauer, 2006. Neuro tissue engineering of glial nerve guides and the impact of different cell types, Biomaterials, 27: 1425-1436.

[4] Vaz, C.M., S. van Tuijl, C.V.C. Bouten, F.P.T. Baaijens, 2005. Design of scaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique, Acta Biomaterialia, 1: 575-582.

[5] Dai, N.T., M.R. Williamson, N. Khammo, E.F. Adams, A.G.A. Coombes, 2004. Composite cell support membranes based on collagen and polycaprolactone for tissue engineering of skin, Biomaterials, 25: 4263-4271.

[6] Killion, J.A., L.M. Geever, D.M. Devine, J.E. Kennedy, C.L. Higginbotham, 2011. Mechanical properties and thermal behaviour of PEGDMA hydrogels for potential bone regeneration application, Mechanical behaviour of biomedical materials, 4: 1219-1227.

[7] Schmedlen, R.H., K.S. Masters, J.L. West, 2002. Photocrosslinkable polyvinyl alcohol hydrogels that can be modified with cell adhesion peptides for use in tissue engineering, Biomaterials, 23(22): 4325-4332.

[8] Stalling, S.S., S.O. Akintoye, S.B. Nicoll, Development of photocrosslinked methylcellulose hydrogels for soft tissue reconstruction. Acta Biomaterialia, 5: 1911-1918.

[9] He, H., L. Li, L.J. Lee, Photopolymerization and structure formation of methacrylic acid based hydrogels: The effect of light intensity, Reactive and Functional Polymers, 68: 103-113.

[10] Geever, L.M., D.M. Devine, M.J.D. Nugent, J.E. Kennedy, J.G. Lyons, A. Hanley, C.L. Higginbotham, 2006. Lower critical solution temperature control and swelling behaviour of physically crosslinked thermosensitive copolymers based on N-isopropylacrylamide, European Polymer Journal, 42: 2540-2548.

[11] Devine, D.M., S.M. Devery, J.G. Lyons, L.M. Geever, J.E. Kennedy, C.L. Higginbotham, 2006. Multifunctional polyvinylpyrrolidinone-polyacrylic acid copolymer hydrogels for biomedical applications, International Journal of Pharmaceutics, 326: 50-59.

[12] Kennedy, J.E., C.L. Higginbotham, 2011. Synthesis and characterisation of styrene butadiene styrene-g-Nvinyl-2- pyrrolidinonefor use in biomedical applications, Materials Science and Engineering: C, 31: 246- 251.

[13] Nguyen, K.T., J.L. West, 2002. Photopolymerisable hydrogels for tissue engineering applications, Biomaterials, 23: 4307-4314.

[14] Karaca, N., G. Temel, D. Karaca Balta, M. Aydin, N. Arsu, 2010. Preparation of hydrogels by photopolymerisation of acrylates in the presence of Type I and one-component Type II photoinitiators, Photochemistry and Photobiology A: Chemistry, 209(1): 1-6.

[15] Bryant, S., C. Nuttelman, K. Anseth, 2000. Cytocompatibility of UV and visible light photoinitiating systems on cultured NIH/3T3 fibroblasts in vitro, Biomaterial Science Polymer Edition, 11(5): 439-457.

[16] Tang, Q., X Sun, Q. Li, J. Wu, J. Lin, M. Huang, 2009. A simple route to high-strength hydrogel with an interpenetrating polymer network, e-Polymers, no. 090 (2009) http://www.e- 778-789.

[17] Killion, J.A., L.M. Geever, D.M. Devine, L. Grehan, J.E. Kennedy, C.L. Higginbotham, 2012. Modulating the mechanical properties of photopolymerised polyethylene glycol-polypropylene glycol hydrogels for bone regeneration, Journal of Materials Sciences, 47(18): 6577-6585.

[18] Graham, B.S., D.W. Jones, E.J. Sutow, 1991. An in vivo and in vitro study of the loss of plasticizer from soft polymer-gel materials, Journal of Dental Restoration, 70(5): 870-3.

[19] Murugan, R., S. Ramakrishna, 2005. Development of nanocomposites for bone grafting. Composites Science and Technology, 65(15-16): 2385-2406.

[20] Uchida, T., S. Ikeda, H. Oura, M. Tada, T. Nakano, T. Fukuda, T. Matsuda, M. Negoro, F. Arai, 2008. Development of biodegradable scaffolds based on patient-specific arterial configuration, Journal Biotechnology, 20(133)(2): 213-8.

[21] Kang, E., J. Ryoo, G.S. Jeong, Y.Y. Choi, S.M. Jeong, J. Ju, S. Chung, S. Takayama, S.H. Lee, 2013. Large-Scale, Ultrapliable, and Free-Standing Nanomembranes, Advanced materials, 25(15): 2167-2173.

[22] Park, S., S.B. Nicoll, R.L. Mauck, G.A. Ateshian, 2008. Cartilage mechanical response under dynamic compression at physiological stress levels following collagenase digestion, Ann Biomed Engineering, 36: 425-434.

[23] Peppas, N.A., P. Bures, W. Leobandung, H. Ichikawa, 2000. Hydrogels in pharmaceutical formulations. Pharmaceutics and Biopharmaceutics, Review Article, 50: 27-46.

[24] Singh, N.K., A.L. Lesser, Mechanical and Thermo-Mechanical Studies of Double Networks Based on Thermoplastic Elastomers, Wiley InterScience, pp: 778-789.

[25] Lyons, G.L., L.M. Geever, M.J.D. Nugent, J.E. Kennedy, C.L. Higginbotham, 2009. Development and characterisation of an agar-polyvinyl alcohol blend hydrogel, Journal of the mechanical behaviour of biomedical materials, 2: 485-493.

[26] Wang, H., Q. Zou, O.C. Boerman, A.W. Nijhuis, J.A. Jansen, Y. Li, S.C. Leeuwenburgh, 2013. Combined delivery of BMP-2 and bFGF fromnanostructured colloidal gelatin gels and its effect on bone regeneration in vivo, J. Control. Release, 166(2): 172-181.

[27] Gaharwar, A.K., S.A. Dammu, J.M. Canter, C.J. Wu, G. Schmidt, Highly Extensible, 2011a. Tough and Elastomeric Nanocomposite Hydrogels from Poly (ethylene glycol) and Hydroxyapatite Nanoparticles. Biomacromolecules, 12(5): 1641-1650.

[28] Huang, Z., B. Yu, Q. Feng, S. Li, Y. Chen, L. Lou, 2011. In situ-forming chitosan/nanohydroxyapatite/collagen gel for the delivery of bone marrow mesenchymal stem cells, Carbohydrate Polymer, 85(1): 261-267.

[29] Gomes, M.E., R.L. Reis, A.M. Cunha, C.A. Blitterwijk, J.D. de Bruijin, 2000. Cytocompatability and response of osteoblastic-like cells to starch based polymers: effect of several additives and processing conditions, Biomaterials, 22: 1911-1917.

[30] Garle, M.J., J.H. Fentem, J.R. Fry, 1994. In vitro cytotoxicity tests for the prediction of acute toxicity in vivo, vitro, 8: 1303-1312.

[31] Graham, N.B., 1998a. Hydrogels: their future, Part I. Medical Device Technology, pp: 18-22.

[32] Geever, L.M., C.C. Cooney, J.G. Lyons, J.E. Kennedy, M.J.D. Nugent, S.M. Devery, C.L. Higginbotham, 2008. Characterisation and controlled drug release from novel drug-loaded hydrogels,Pharmaceutics and Biopharmaceutics, 69(3): 1147-1159.

[33] Mosmann, T., 1983. Rapid colorimetric assay for cellular growth and survival: Application to proliferation and cytotoxicity assays. Journal of Immunological Methods, 65(1-2): 55-63.

[34] Stockert, J.C., A.B. Castro, M. Canete, R.W. Horobin, A. Villanueva, 2012. MTT assay for cell viability: Intracellular localization of the formazan product is in lipid droplets, Acta Histochemica, 114: 785-796.

Tess Geever, John Killion, Laura Grehan, L.M. Geever, Edel Chadwick, Clement


Materials Research Institute, Athlone Institute of Technology, Dublin Road,

Athlone, Co. Westmeath, Ireland

Corresponding Author: Clement Higginbotham, Materials Research Institute, Athlone Institute of Technology, Dublin Road, Athlone, Co. Westmeath, Ireland


Table 1: Formulated composition of PEGDMA with
distilled water prior to photopolymerisation at
varying photoinitiator concentrations

Hydrogel     PEGDMA400   PEGDMA600   PEGDMA1000   Distilled   Irgacure
code           (wt%)       (wt%)       (wt%)        water       2959
                                                    (wt%)      (wt%)

P400-0.01       50                                   50         0.01
P400-0.05       50                                   50         0.05
P400-0.1        50                                   50         0.1
P400-0.5        50                                   50         0.5
P400-1          50                                   50          1
P600-0.01                   50                       50         0.01
P600-0.05                   50                       50         0.05
P600-0.1                    50                       50         0.1
P600-0.5                    50                       50         0.5
P600-1                      50                       50          1
P1000-0.01                               50          50         0.01
P1000-0.05                               50          50         0.05
P1000-0.1                                50          50         0.1
P1000-0.5                                50          50         0.5
P1000-1                                  50          50          1

Table 2:    Storage modulus of hydrogels of
varying photoinitiator concentration and
associated molecular weights of the macromolecular

Hydrogel code (a)   Storage modulus G'    [+ or -] Standard deviation

P400-0.01                  9,992                  0.403
P400-0.05                 24,298                  1.533
P400-0.1                  29,603                  1.056
P400-0.5                  51,273                 15.549
P400-1                    65,637                  9.789
P600-0.01                  6,713                  1.193
P600-0.05                  7,158                  1.900
P600-0.1                   8,095                  1.408
P600-0.5                  10,707                  1.435
P600-1                    15,040                  1.985
P1000-0.01                 9,994                  0.875
P1000-0.05                28,526                  3.163
P1000-0.1                 31,216                  5.875
P1000-0.5                 33,725                  5.623
P1000-1                   34,339                  3.421

(a) See Table 1
COPYRIGHT 2014 American-Eurasian Network for Scientific Information
No portion of this article can be reproduced without the express written permission from the copyright holder.
Copyright 2014 Gale, Cengage Learning. All rights reserved.

Article Details
Printer friendly Cite/link Email Feedback
Author:Geever, Tess; Killion, John; Grehan, Laura; Geever, L.M.; Chadwick, Edel; Higginbotham, Clement
Publication:Advances in Environmental Biology
Article Type:Report
Date:Dec 1, 2014
Previous Article:Smart thermosensitive poly (N-vinylcaprolactam) based hydrogels for biomedical applications.
Next Article:Spinosad effects on mortality and reproduction of Culex pipiens (Diptera; Culicidae).

Terms of use | Privacy policy | Copyright © 2021 Farlex, Inc. | Feedback | For webmasters |