Hyaluronic acid and chondroitin sulfate (meth)acrylate-based hydrogels for tissue engineering: Synthesis, characteristics and pre-clinical evaluation.

Hydrogels based on photocrosslinkable Hyaluronic Acid Methacrylate (HAMA) and Chondroitin Sulfate Methacrylate (CSMA) are presently under investigation for tissue engineering applications. HAMA and CSMA gels offer tunable characteristics such as tailorable mechanical properties, swelling characteristics, and enzymatic degradability. This review gives an overview of the scientific literature published regarding the pre-clinical development of covalently crosslinked hydrogels that (partially) are based on HAMA and/or CSMA. Throughout the review, recommendations for the next steps in clinical translation of hydrogels based on HAMA or CSMA are made and potential pitfalls are defined. Specifically, a myriad of different synthetic routes to obtain polymerizable hyaluronic acid and chondroitin sulfate derivatives are described. The effects of important parameters such as degree of (meth)acrylation and molecular weight of the synthesized polymers on the formed hydrogels are discussed and useful analytical techniques for their characterization are summarized. Furthermore, the characteristics of the formed hydrogels including their enzymatic degradability are discussed. Finally, a summary of several recent applications of these hydrogels in applied fields such as cartilage and cardiac regeneration and advanced tissue modelling is presented.

Protein release from water-swellable poly(D,L-lactide-PEG)-b-poly(ϵ-caprolactone) implants.

In this study, water-swellable multiblock copolymers composed of semi-crystalline poly(ϵ-caprolactone) [PCL] blocks and amorphous blocks consisting of poly(D,L-lactide) (PDLLA) and poly(ethylene glycol) (PEG) [PDLLA-PEG] were synthesized. The block ratio of these [PDLLA-PEG]-b-[PCL] multiblock copolymers was varied and the degradation of implants prepared of these polymers by hot melt extrusion (HME) was compared with implants prepared of [PCL-PEG]-b-[PCL], a copolymer which has been described previously (Stankovic et al., 2014). It was shown that the initial degradation rate of the [PDLLA-PEG]-b-[PCL] multiblock copolymers increased with increasing the content of amorphous [PDLLA-PEG] block and that the degradation rate of these multiblock copolymers was faster than that of the [PCL-PEG]-b-[PCL] multiblock copolymers due to rapid degradation of the [PDLLA-PEG] block. Furthermore, the release of the model proteins lysozyme and bovine serum albumin from polymer implants prepared by HME was studied. It was found that the protein release from [PDLLA-PEG]-b-[PCL] copolymers was incomplete, which is not acceptable for any application of these polymers. Besides, [PCL-PEG]-b-[PCL] copolymers showed slow and continuous release. We hypothesize that the incomplete release is explained by an irreversible interaction between the proteins and polymer degradation products or by entrapment of the protein in the hydrophobic and non-swellable polymer matrix that was left after degradation and loss of the hydrophilic [PDLLA-PEG] blocks from the degrading polymer.

Tailored protein release from biodegradable poly(ε-caprolactone-PEG)-b-poly(ε-caprolactone) multiblock-copolymer implants.

In this study, the in vitro release of proteins from novel, biodegradable phase-separated poly(ε-caprolactone-PEG)-block-poly(ε-caprolactone), [PCL-PEG]-b-[PCL]) multiblock copolymers with different block ratios and with a low melting temperature (49-55°C) was studied. The effect of block ratio and PEG content of the polymers (i.e. 22.5, 37.5 and 52.5 wt%) as well as the effect of protein molecular weight (1.2, 5.8, 14, 29 and 66 kDa being goserelin, insulin, lysozyme, carbonic anhydrase and albumin, respectively) on protein release was investigated. Proteins were spray-dried with inulin as stabilizer to obtain a powder of uniform particle size. Spray-dried inulin-stabilized proteins were incorporated into polymeric implants by hot melt extrusion. All incorporated proteins fully preserved their structural integrity as determined after extraction of these proteins from the polymeric implants. In general, it was found that the release rate of the protein increased with decreasing molecular weight of the protein and with increasing the PEG content of the polymer. Swelling and degradation rate of the copolymer increased with increasing PEG content. Hence, release of proteins of various molecular weights from [PCL-PEG]-b-[PCL] multi-block copolymers can be tailored by varying the PEG content of the polymer.

Low temperature extruded implants based on novel hydrophilic multiblock copolymer for long-term protein delivery.

Parenteral protein delivery requires preservation of the integrity of proteins and control over the release kinetics. In order to preserve the integrity, parenteral protein delivery formulations typically need to be processed at low temperatures. Therefore, we synthesized a novel low melting biodegradable hydrophilic multiblock copolymer composed of poly (ethylene glycol) and poly (ε-caprolactone) to allow extrusion at relatively low temperatures. We investigated the extrusion characteristics of this polymer and explored a strategy how to control the release of the model protein lysozyme from small diameter extruded implants. It was found that the polymer could be well extruded at temperatures as low as 55 °C. Moreover, lysozyme remained active both during extrusion as well as during release. Lysozyme release kinetics could be tailored by the co-incorporation of an oligosaccharide, inulin, which functions as a pore-forming excipient. It was concluded that this hydrophilic multiblock copolymer has promising characteristics for the preparation by melt extrusion of protein delivery implants with a release profile that is sustained over a period of more than 7 months.

Probing DNA interstrand cross-link formation by an oxidized abasic site using nonnative nucleotides.

The C4'-oxidized abasic site (C4-AP) forms two types of interstrand cross-links with the adjacent nucleotides in DNA. Previous experiments revealed that dG does not react with the lesion and that formation of one type of cross-link is catalyzed by the opposing dA. iso-Guanosine·dC and 2-aminopurine·dT base pairs were used to determine why dG does not cross-link with C4-AP despite its well known reactivity with other bis-electrophiles. 7-Deaza-2'-deoxyadenosine was used to probe the role of the nucleotide opposite C4-AP in the catalysis of interstrand cross-link formation.

A newly developed chemically crosslinked dextran-poly(ethylene glycol) hydrogel for cartilage tissue engineering.

Cartilage tissue engineering, in which chondrogenic cells are combined with a scaffold, is a cell-based approach to regenerate damaged cartilage. Various scaffold materials have been investigated, among which are hydrogels. Previously, we have developed dextran-based hydrogels that form under physiological conditions via a Michael-type addition reaction. Hydrogels can be formed in situ by mixing a thiol-functionalized dextran with a tetra-acrylated star poly(ethylene glycol) solution. In this article we describe how the degradation time of dextran-poly(ethylene glycol) hydrogels can be varied from 3 to 7 weeks by changing the degree of substitution of thiol groups on dextran. The degradation times increased slightly after encapsulation of chondrocytes in the gels. The effect of the gelation reaction on cell viability and cartilage formation in the hydrogels was investigated. Chondrocytes or embryonic stem cells were mixed in the aqueous dextran solution, and we confirmed that the cells survived gelation. After a 3-week culturing period, chondrocytes and embryonic stem cell-derived embryoid bodies were still viable and both cell types produced cartilaginous tissue. Our data demonstrate the potential of dextran hydrogels for cartilage tissue engineering strategies.

Release of model proteins and basic fibroblast growth factor from in situ forming degradable dextran hydrogels.

Our previous studies showed that degradable dextran hydrogels are rapidly formed in situ upon mixing aqueous solutions of dextran vinyl sulfone (dex-VS) conjugates and tetrafunctional mercapto poly(ethylene glycol) (PEG-4-SH) by Michael addition. The hydrogel degradation time and storage modulus could be controlled by the degree of vinyl sulfone substitution (DS) and dextran molecular weight. The degradation time could further be adjusted by the spacer between the thioether and the ester bond of the dex-VS conjugates (ethyl vs. propyl, denoted as dex-Et-VS and dex-Pr-VS, respectively). In this paper, the release of three model proteins, i.e. immunoglobulin G (d(h)=10.7 nm, IgG), bovine serum albumin (BSA, d(h)=7.2 nm) and lysozyme (d(h)=4.1 nm), as well as basic fibroblast growth factor (bFGF) from these in situ forming dextran hydrogels is studied. Proteins could be easily loaded into the hydrogels by mixing protein containing solutions of dex-VS and PEG-4-SH. The release of IgG from dex-Et-VS hydrogels followed biphasic release kinetics, with a slow, close to first order release for the first 9 days followed by an accelerated release and over 80% of IgG was released in 12 to 25 days. Interestingly, the release of IgG from dex-Pr-VS hydrogels followed close to zero order kinetics, wherein approximately 95% was released in 21 days. The release of BSA from dex-Pr-VS hydrogels followed biphasic kinetics, with almost first order release followed by close to zero order release. Approximately 75% of the entrapped BSA could be released from dex-Pr-VS hydrogels in 16 days. Dex-Pr-VS hydrogels released 40% of lysozyme in 14 days, with full preservation of the enzymatic activity of the released lysozyme, as determined by bacteria lysis experiments. The release of basic fibroblast growth factor (bFGF) from dex-Pr-VS hydrogels showed first order kinetics, with quantitative release in 28 days. These results show that the in situ forming degradable dextran hydrogels can be used for the controlled release of proteins.

Rapidly in situ forming biodegradable robust hydrogels by combining stereocomplexation and photopolymerization.

Our previous studies have shown that stereocomplexed hydrogels can be rapidly formed in vitro as well as in vivo upon mixing aqueous solutions of eight-arm poly(ethylene glycol)-poly(l-lactide) (PEG-PLLA) and poly(ethylene glycol)-poly(d-lactide) (PEG-PDLA) star block copolymers. In this study, stereocomplexation and photopolymerization are combined to yield rapidly in situ forming robust hydrogels. Two types of methacrylate-functionalized PEG-PLLA and PEG-PDLA star block copolymers, PEG-PLLA-MA and PEG-PDLA-MA, which have methacrylate groups at the PLA chain ends and PEG-MA/PLLA and PEG-MA/PDLA, which have methacrylate groups at the PEG chain ends, were designed and prepared. Results showed that stereocomplexed hydrogels could be rapidly formed (within 1-2 min) in a polymer concentration range of 12.5-17.5% (w/v), in which the methacrylate group hardly interfered with the stereocomplexation. When subsequently photopolymerized, these hydrogels showed largely increased storage moduli as compared to the corresponding hydrogels that were cross-linked by stereocomplexation or photopolymerization only. Interestingly, the storage modulus of stereocomplexed-photopolymerized PEG-PLA-MA hydrogels increased linearly with increasing stereocomplexation equilibration time prior to photopolymerization (from ca. 6 to 32 kPa), indicating that stereocomplexation aids in photopolymerization. Importantly, photopolymerization of stereocomplexed hydrogels could take place at very low initiator concentrations (0.003 wt %). Swelling/degradation studies showed that combining stereocomplexation and photopolymerization yielded hydrogels with prolonged degradation times as compared to corresponding hydrogels cross-linked by photopolymerization only (3 vs 1.5 weeks). Stereocomplexed-photopolymerized PEG-MA/PLA hydrogels degraded much slower than corresponding PEG-PLA-MA hydrogels, with degradation times ranging from 7 to more than 16 weeks. Therefore, combining stereocomplexation and photopolymerization is a novel approach to obtain rapidly in situ forming robust hydrogels.

In vitro and in vivo protein delivery from in situ forming poly(ethylene glycol)-poly(lactide) hydrogels.

Previous studies have shown that stereocomplexed hydrogels are rapidly formed in situ by mixing aqueous solutions of eight-arm poly(ethylene glycol)-poly(L-lactide) and poly(ethylene glycol)-poly(D-lactide) star block copolymers (denoted as PEG-(PLLA)(8) and PEG-(PDLA)(8), respectively). In this study, in vitro and in vivo protein release from stereocomplexed hydrogels was investigated. These hydrogels were fully degradable under physiological conditions. Proteins could be easily loaded into the stereocomplexed hydrogels by mixing protein containing aqueous solutions of PEG-(PLLA)(8) and PEG-(PDLA)(8) copolymers. The release of the relatively small protein lysozyme (d(h)=4.1 nm) followed first order kinetics and approximately 90% was released in 10 days. Bacteria lysis experiments showed that the released lysozyme had retained its activity. The relatively large protein IgG (d(h)=10.7 nm) could be released from stereocomplexed hydrogels with nearly zero order kinetics, wherein up to 50% was released in 16 days. The in vitro release of the therapeutic protein rhIL-2 from stereocomplexed hydrogels also showed nearly zero order kinetics, wherein up to 45% was released in 7 days. The therapeutic efficacy of stereocomplexed hydrogels loaded with 1x10(6) IU of rhIL-2 was studied using SL2-lymphoma bearing DBA/2 mice. The PEG-(PLLA)(8)/PEG-(PDLA)(8)/rhIL-2 mixture could be easily injected intratumorally. The released rhIL-2 was therapeutically effective as the tumor size was reduced and the cure rate was 30%, whereas no therapeutic effect was achieved when no rhIL-2 was given. However, the cure rate of rhIL-2 loaded stereocomplexed hydrogels was lower, though not statistically significant, compared to that of a single injection with 1x10(6) IU of free rhIL-2 at the start of the therapy (cure rate=70%). The therapeutic effect of rhIL-2 loaded stereocomplexed hydrogels was retarded for approximately 1-2 weeks compared to free rhIL-2, most likely due to a slow, constant release of rhIL-2 from the hydrogels.

Rapidly in situ-forming degradable hydrogels from dextran thiols through Michael addition.

Thiol-functionalized dextrans (dex-SH) (M(n,dextran) = 14K or 31K) with degrees of substitution (DS) ranging from 12 to 25 were synthesized and investigated for in situ hydrogel formation via Michael type addition using poly(ethylene glycol) tetra-acrylate (PEG-4-Acr) or a dextran vinyl sulfone conjugate with DS 10 (dex-VS DS 10). Dex-SH was prepared by activation of the hydroxyl groups of dextran with 4-nitrophenyl chloroformate and subsequent reaction with cysteamine. Hydrogels were rapidly formed in situ under physiological conditions upon mixing aqueous solutions of dex-SH and either PEG-4-Acr or dex-VS DS 10 at polymer concentrations of 10 to 20 w/v%. Rheological studies showed that these hydrogels are highly elastic. By varying the DS, concentration, dextran molecular weight, and type of cross-linker, hydrogels with a broad range of storage moduli of 9 to 100 kPa could be obtained. Varying the ratio of thiol to vinyl sulfone groups from 0.9 to 1.1 did not alter the storage modulus of the hydrogels, whereas larger deviations from equimolarity (thiol to vinyl sulfone ratios of 0.75 and 1.5) considerably decreased the storage modulus. The plateau value of hydrogel storage modulus was reached much faster at pH 7.4 compared to pH 7, due to a higher concentration of the thiolate anion at higher pH. These hydrogels were degradable under physiological conditions. Degradation times were 3 to 7 weeks for dex-SH/dex-VS DS 10 hydrogels and 7 to over 21 weeks for dex-SH/PEG-4-Acr hydrogels, depending on the DS, concentration, and dextran molecular weight.

Enzyme-mediated fast in situ formation of hydrogels from dextran-tyramine conjugates.

Dextran hydrogels were formed in situ by enzymatic crosslinking of dextran-tyramine conjugates and their mechanical, swelling and degradation properties were evaluated. Two types of dextran-tyramine conjugates (M(n,dextran)=14k, M(w)/M(n)=1.45), i.e. dextran-tyramine linked by a urethane bond (denoted as Dex-TA) or by an ester-containing diglycolic group (denoted as Dex-DG-TA), with different degrees of substitution (DS) were prepared. Hydrogels were rapidly formed under physiological conditions from Dex-TA DS 10 or 15 and Dex-DG-TA DS 10 at or above a concentration of 2.5 wt% in the presence of H(2)O(2) and horseradish peroxidase (HRP). The gelation time ranged from 5s to 9 min depending on the polymer concentration and HRP/TA and H(2)O(2)/TA ratios. Rheological analysis showed that these hydrogels are highly elastic. The storage modulus (G'), which varied from 3 to 41 kPa, increased with increasing polymer concentration, increasing HRP/TA ratio and decreasing H(2)O(2)/TA ratio. The swelling/degradation studies showed that under physiological conditions, Dex-TA hydrogels are rather stable with less than 25% loss of gel weight in 5 months, whereas Dex-DG-TA hydrogels are completely degraded within 4-10d. These results demonstrate that enzymatic crosslinking is an efficient way to obtain fast in situ formation of hydrogels. These dextran-based hydrogels are promising for use as injectable systems for biomedical applications including tissue engineering and protein delivery.

In-situ formation of biodegradable hydrogels by stereocomplexation of PEG-(PLLA)8 and PEG-(PDLA)8 star block copolymers.

Eight-arm poly(ethylene glycol)-poly(L-lactide), PEG-(PLLA)(8), and poly(ethylene glycol)-poly(D-lactide), PEG-(PDLA)(8), star block copolymers were synthesized by ring-opening polymerization of either L-lactide or D-lactide at room temperature in the presence of a single-site ethylzinc complex and 8-arm PEG (M(n) = 21.8 x 10(3) or 43.5 x 10(3)) as a catalyst and initiator, respectively. High lactide conversions (>95%) and well-defined copolymers with PLLA or PDLA blocks of the desired molecular weights were obtained. Star block copolymers were water-soluble when the number of lactyl units per poly(lactide) (PLA) block did not exceed 14 and 17 for PEG21800-(PLA)(8) and PEG43500-(PLA)(8), respectively. PEG-(PLA)(8) stereocomplexed hydrogels were prepared by mixing aqueous solutions with equimolar amounts of PEG-(PLLA)(8) and PEG-(PDLA)(8) in a polymer concentration range of 5-25 w/v % for PEG21800-(PLA)(8) star block copolymers and of 6-8 w/v % for PEG43500-(PLA)(8) star block copolymers. The gelation is driven by stereocomplexation of the PLLA and PDLA blocks, as confirmed by wide-angle X-ray scattering experiments. The stereocomplexed hydrogels were stable in a range from 10 to 70 degrees C, depending on their aqueous concentration and the PLA block length. Stereocomplexed hydrogels at 10 w/v % polymer concentration showed larger hydrophilic and hydrophobic domains as compared to 10 w/v % single enantiomer solutions, as determined by cryo-TEM. Correspondingly, dynamic light scattering showed that 1 w/v % solutions containing both PEG-(PLLA)(8) and PEG-(PDLA)(8) have larger "micelles" as compared to 1 w/v % single enantiomer solutions. With increasing polymer concentration and PLLA and PDLA block length, the storage modulus of the stereocomplexed hydrogels increases and the gelation time decreases. Stereocomplexed hydrogels with high storage moduli (up to 14 kPa) could be obtained at 37 degrees C in PBS. These stereocomplexed hydrogels are promising for use in biomedical applications, including drug delivery and tissue engineering, because they are biodegradable and the in-situ formation allows for easy immobilization of drugs and cells.