Hyaluronic Acid Hydrogels Formed in Situ by Transglutaminase

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Hyaluronic acid hydrogels formed in situ by transglutaminase-catalyzed reaction Adrian Ranga, Matthias P. Lutolf, Joens Hilborn, and Dmitri A. Ossipov Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.5b01587 • Publication Date (Web): 25 Mar 2016 Downloaded from http://pubs.acs.org on March 28, 2016

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Hyaluronic acid hydrogels formed in situ by transglutaminase-catalyzed reaction

Adrian Ranga1,2*, Matthias P. Lutolf 1,3, Jöns Hilborn4, Dmitri A. Ossipov4*

1

Laboratory of Stem Cell Bioengineering, Institute of Bioengineering, School of Life Sciences

and School of Engineering, Ecole Polytechnique Fédérale de Lausanne (EPFL), Lausanne, CH1015, Switzerland. 2

Biomechanics Section, Mechanical Engineering Department, KU Leuven, Leuven, 3001,

Belgium 3

Science for Life Laboratory, Division of Polymer Chemistry, Department of Chemistry-

Ångström, Uppsala University, Uppsala, SE 751 21, Sweden 4

Institute of Chemical Sciences and Engineering, School of Basic Sciences, EFFL, Lausanne, CH

1015, Switzerland.

KEYWORDS: hydrogel, hyaluronic acid, poly(ethylene glycol), transglutaminase, Factor XIII, biodegradation.

ACS Paragon Plus Environment

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ABSTRACT

Enzymatically cross-linked hydrogels can be formed in situ and permit highly versatile and selective tethering of bioactive molecules, thereby allowing for a wealth of applications in cell biology and tissue engineering. While a number of studies have reported the bioconjugation of extracellular matrix (ECM) proteins and peptides into such matrices, the site-specific incorporation of biologically highly relevant polysaccharides such as hyaluronic acid (HA) has thus far not been reported, limiting our ability to reconstruct this key feature of the in vivo ECM. Here, we demonstrate a novel strategy for transglutaminase-mediated covalent linking of HA moieties to a synthetic poly(ethylene glycol) (PEG) macromer resulting in the formation of hybrid HA-PEG hydrogels. We characterize the ensuing matrix properties, and demonstrate how these cytocompatible gels can serve to modulate the cellular phenotype of human mammary cancer epithelial cells as well as mouse myoblasts. The use of HA as a novel building block in the increasingly varied library of synthetic PEG-based artificial ECMs should have applications as a structural as well as a signaling component, and offers significant potential as an injectable matrix for regenerative medicine.

INTRODUCTION The development of biomaterials which can faithfully recapitulate key aspects of the in vivo microenvironment has been a cornerstone of tissue engineering strategies. Hydrogels are a particularly attractive class of biomaterials for tissue engineering due to their high water content, tissue-like viscoelastic properties and efficient mass transport characteristics. Additionally, the possibility of tuning their characteristics via specific chemical designs has allowed for


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increasingly modular strategies for precise tailoring to specific applications.1 For many applications in regenerative medicine, it is also advantageous to design such materials to be injectable, such that they can be used in the context of minimally invasive procedures which lead to improved patient outcomes and lower healthcare costs. Examples of such injectable hydrogels for tissue engineering include cartilage repair,2 adipose tissue formation,3 corneal wound repair,4 treatment of vocal fold insufficiency,5 and myocardial infarction.6 An in situ solution-to-gel transition just prior to injection or during the injection is characteristic of such hydrogels and imposes stringent requirements for cross-linking conditions; the presence of aqueous media, a physiologic pH and temperature, suitable kinetics, and selectivity of the cross-linking reaction need to be taken into consideration. The last factor is crucial for the in situ encapsulation and delivery of sensitive biomolecules (e.g. proteins, nucleic acids) and/or live cells. Among various chemical cross-linking methods employed for the preparation of injectable hydrogels, enzymatic reactions can be characterized as the most selective toward their unique substrates.7 Particularly, transglutaminases (TGs), a family of enzymes catalyzing the formation of amide bonds between amine group of lysine and the carboxamide group of glutamine residues of proteins, have been employed as chemical triggers in the preparation of a variety of hydrogel types. In vivo, TG-catalyzed reactions are for example responsible for the stabilization of the fibrin clot, a “biological glue” that is generated from soluble fibrinogen during blood coagulation.8 The main advantage of this reaction is that it does not require additional co-factors other than Ca2+ and occurs rapidly and reasonably selectively. Consequently, proteins such as gelatin,9,10 casein,11 and soy protein,12 as well as genetically engineered polypeptides13,14 have been converted into hydrogel materials by TG-based enzymatic cross-linking, using tissue and microbial TGs. More recently, it has been shown that


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TG-catalyzed cross-linking can also be applied to synthetic PEG macromers by linking peptides functioning as TG substrates to the PEG termini via tissue TG or factor XIII (FXIII), a circulatory form of TGs participating in the formation of fibrin15-18 Moreover, rational design of the TG substrate peptides has been demonstrated to enhance their specificity towards the enzyme by several orders of magnitude, a property of considerable importance for biomolecule delivery strategies.17,18 Using this approach, gelation time has been substantially reduced, thereby ensuring that such a system could be compatible with practical tissue engineering and surgical adhesive requirements. Previously, using glutamine acceptor substrate derived from N-terminus of α2-plasmin inhibitor, we and others have shown the use of activated factor XIII (FXIIIa) for the bioengineering of fibrin matrices20,21 as well as PEG-based hydrogels.18,19. This modular approach can be employed to augment the biochemical properties of fibrin20,21 and engineer synthetic matrices with user-defined composition of covalently immobilized bioactive factors that can be assembled with high fidelity.22,23 Surprisingly, however, no TG substrates other than peptides and proteins have been developed thus far. Indeed, in order to construct a hydrogel which could mimic native ECMs more faithfully, there is a significant need to broaden the scope of available ligands, which should include highly abundant polysaccharides due to their prominent structural and well as signaling roles in native tissues. HA is one of the major components of native ECMs and is particularly abundant in connective tissues as well as in synovial fluid. High levels of endogenous HA have been detected in remodeling and healing tissues with high cell proliferation and migration. HA represents the only non-sulfated member of the glycosaminoglycans (GAGs) family, and promotes cell migration by interacting with cell surface receptors such as CD44. Due to their biocompatibility,


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injectable HA hydrogels have been utilized in a variety of biomedical applications.24 To prepare injectable hydrogels, a number of chemoselective cross-linking reactions have been applied to HA, including azide-alkyne cycloaddition click reaction,25 aqueous Diels-Alder chemistry,26 oxime coupling,27 Michael addition of thiols to acrylates,28 and thiol-disulfide exchange reactions.29 While peroxidase-mediated cross-linking of tyramine residues has been demonstrated30, to the best of our knowledge, the use of enzymatic reactions for incorporation of HA into hydrogels has not yet been achieved. As such, by providing the chemical orthogonality of enzymatic reactions with additional “click”-type non-enzymatic reactions, the use of suitable HA-based enzyme substrates would be highly desirable in the modular construction of potent ECM-mimetic hydrogels. Here, we have employed for the first time a TG-catalyzed reaction for the in situ crosslinking of HA hydrogels. By clickable disulfide conjugation of cysteine-bearing TG substrate peptides to the backbone of HA, we were able to convert this polysaccharide into a substrate of FXIIIa. This was demonstrated in a hydrogel system containing two complementary enzyme substrates prepared from HA and PEG, respectively. Furthermore, we show that the inclusion of HA into PEG hydrogels can also be exploited to tune the hydrogel’s mechanical and biochemical properties, resulting in specific modulation of cellular phenotype.


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MATERIALS AND METHODS HA Materials. Hyaluronic acid (HA) sodium salt (MW 59 kDa) was purchased from Lifecore Biomedical.

N-hydroxybenzotriazole

(HOBt)

and

1-ethyl-3-(3-dimethylaminopropyl)

carbodiimide (EDC) were purchased from Sigma-Aldrich Chemical Co. All solvents were of analytical quality (p.a.) and were dried over 4Å molecular sieves. Dialysis membranes Spectra/ Por 6 (3500 g/mol cutoff) were purchased from VWR international. 1H-NMR spectra were recorded in D2O with a JEOL JNM-ECP Series FT NMR spectrometer at a magnetic field strength 9.4 T, operating at 400 MHz. 2-(2-Pyridinyldithio)ethylhydrazinecarboxylate 1 was synthesized according to previously published procedures.31 Synthesis of 2-dithiopyridyl modified hyaluronic acid (HA-SSPy). HA (MW 59000 Da) (400 mg, 1 mmol of disaccharide repeating units) was dissolved in de-ionized water at the concentration of 8 mg/mL. 2-(2-Pyridinyldithio)ethylhydrazinecarboxylate 131 (49 mg, 0.2 mmol) was dissolved in 4 mL of methanol and the solution was added to the solution of HA. Nhydroxybenzotriazole (153 mg, 1 mmol) in 5 mL of 1:1 (v/v) mixture of acetonitrile-water was also added to the HA solution under stirring. The pH of the resultant solution was adjusted to 4.8, after which EDC (48 mg, 0.25 mmol) was added to the mixture. The reaction mixture was stirred overnight at room temperature. After the amide coupling, the mixture was acidified to 3.5 with 1M HCl and transferred to a dialysis tube (Mw cutoff = 3500). After multiple dialysis rounds against dilute HCl (pH 3.5) containing 0.1 M NaCl, followed by dialysis against dilute HCl, pH 3.5, the solution was lyophilized to give 348 mg of HA-SSPy (87% yield). The incorporation of pyridyl group was verified by 1H-NMR. Specifically, the newly appeared peaks at 7.45, 7.92 – 8.06, and 8.45 ppm corresponding to four aromatic protons of the pyridyl group were integrated., which indicated that 15 % of the HA disaccharide units had been modified.


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Synthesis of hyaluronic acid modified with glutamine substrate peptide (HA-Gln). 2dithiopyridyl modified hyaluronic acid (HA-SSPy) (25 mg, 60.5 µmol of disaccharide monomeric units, 9.1 µmol of 2-dithiopyridyl groups) was dissolved in 6 mL PBS (Dulbecco’s Phosphate Buffered Saline). Ac-NQEQVSPLERCG-NH2 peptide (12.3 mg, 9.1 µmol) was added in solid form to the HA solution. After dissolution of the peptide, the reaction mixture was stirred at room temperature for 24 hours and subsequently dialyzed against dilute HCl (pH 3.5) followed by double dialysis against pure water (Spectra/Por® 6, 3500 g/mol cut off). The dialyzed material was freeze-dried to give 33 mg of HA-Gln (yield – 90.1%). Preparation of hydrogels. For convenience, we reproduce below methodological details originally available in refs 18,19. Briefly, the factor XIIIa substrate peptides TG-MMP-Lys (AcFKGG-GPQGIWGQ-ERCGNH2 (MMP-sensitive) and AcFKGG-GDQGIAGF-ERCG-NH2 (non-MMP sensitive)) and TGGln (NQEQVSPL-ERCG-NH2) (GLBiochem, China) were added to eight-arm PEG vinylsulfone (PEG-VS) (NOF, Japan) in a 1.2-fold molar excess over VS groups in 0.3 M triethanolamine (pH 8.0) at 37°C for 2h. The reaction solution was subsequently dialysed (Snake Skin, MWCO 10K, PIERCE) against ultrapure water for 3 days at 4°C. After dialysis, the product (8-PEG– MMP-Lys and 8-PEG-Gln, respectively) was lyophilized to obtain a white powder. Activation of factor XIII (FXIII) was achieved as described previously18. Briefly, 1ml of 200U/ml FXIII (CSL Behring) was incubated for 30min at 37°C with 100µl of 20U/ml thrombin (Sigma-Aldrich) in the presence of 2.5mM CaCl2. Activated FXIII was aliquoted and stored at −80°C until use. To prepare PEG-only gels, lyophilized PEG-Gln and PEG-Lys were dissolved in deionized water at 133.3 mg/ml. Precursor solutions to give hydrogels with a final dry mass content as specified in the text were prepared by stoichiometrically balanced ([Lys]/[Gln]=1) solutions of PEG-Gln


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and PEG-Lys in Tris buffer (50mM, pH 7.6) containing 50mM CaCl2. The cross-linking reaction was initiated by 10U/ml thrombin-activated factor XIIIa. In experiments with incorporation of the adhesion peptide, RGD-Lys (AcFKGGERCG-RDGS) peptide was incorporated into the reaction mixture prior to gelation to obtain a final 50uM peptide concentration. To prepare HA-PEG gels, lyophilized HA-Gln and PEG-Lys were dissolved in deionized water at 40 mg/mL (4% w/v) and 133.3 mg/mL concentrations respectively. To prepare 60 µL gelforming mixtures, predetermined volumes of PEG-Lys and HA-Gln stock solutions were diluted to 54 µL with 6 µL of 0.5 M Tris buffer containing 0.5 M CaCl2 (10× buffer) and water. For example, to prepare 3% w/v gels with [Gln]/[Lys] = 1, 0.672 mg/16.8 µL solution of HA-Gln and 1.128 mg/8.5 µL solution of PEG-Lys were mixed and diluted with 6 µL of 10× buffer and 22.7 µL of water. 6 µL of 100 U/mL thrombin-activated factor XIIIa was added to the mixture and all contents were homogenized by pipetting for 1 minute. HA-PEG was thus formed in the same conditions as PEG-only gels (final buffer concentration of 50 mM Tris/50 mM CaCl2, 10 U/mL factor XIIIa).

In situ rheometry. Gelation kinetics were studied by performing small strain oscillatory shear experiments on a Bohlin CVO 120 high-resolution rheometer with plate-plate geometry at ambient temperature. For this purpose, 40 µL gel-forming mixtures were prepared to form 3% w/v gels with [Gln]/[Lys] = 1.4 µL of 100 U/mL thrombin-activated factor XIIIa was added to the mixture and all contents were homogenized by pipetting for 10 seconds. 34.2 µL of the homogenized yet liquid reaction mixture was poured onto the center of the bottom plate. The upper plate (2 cm diameter) was immediately lowered to a measuring gap size of 0.1 mm, and the dynamic oscillation measurement was started. The evolution of storage (G′) and loss (G′′)


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and phase angle δ (°) at a constant frequency of 0.2 Hz and a constant strain of 0.05 was recorded as a function of time. In two other control kinetics experiments, either factor XIIIa was substituted by 4 µL of water, or HA-Gln was substituted by HA-SSPy of the same mass and concentration.

Rheometry on swollen hydrogels. Storage and loss moduli (G′ and G′′) of swollen gels was were obtained using the same instrument as for in-situ gelation kinetics rheometry, Test samples consisting of gel disks were prepared by confining 50 µL of the reaction mixture between sterile hydrophobic glass microscopy slides (obtained by treatment with SigmaCote (Sigma-Aldrich, Switzerland) separated by spacers (ca. 1 mm thickness) and clamped with binder clips. Gelation occurred within 3 to 6 minutes and the cross-linking was allowed to proceed for 30 minutes. After setting, the formed hydrogel disks were removed from the microscopy slides and immersed in 3 mL of 50 mM Tris/50 mM CaCl2 buffer for 24 hours. The swollen hydrogel disks (0.7 – 1.2 mm thickness) were sandwiched between the two plates of the rheometer, with compression between 75% and 85% of their original thickness in order to avoid slipping. Measurements were then conducted in a constant (5%) strain mode as a function of frequency (from 0.1 to 10 Hz) to obtain dynamic mechanical spectra (n = 3 per condition).

Cell culture and cell-based assays. MCF7 mammary carcinoma cells and C2C12 mouse myoblast cells were grown in DMEM/Glutamax (Life Technologies) basal medium supplemented with 15% fetal calf serum (Hyclone), 100uU/mL penicillin/streptomycin (Life Technologies), 1umM sodium pyruvate (Life Technologies) and 0.1mM not essential amino acids (Life Technologies). For gel


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encapsulation experiments, cells were trypsinized using TrypLE Express (Life Technologies) and incorporated in the precursor solution immediately prior to FXIIIa enzyme addition.


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RESULTS ANDDISCUSSION Preparation of HA-based FXIII substrates. We hypothesized that linking a peptide containing the NQEQVSPL sequence, abbreviated here as Gln, as a side chain to the backbone of HA should convert this GAG into a potent substrate for FXIIIa. Gln is a part of the α2-plasmin inhibitor, which is cross-linked via glutamine (Q) residues with the lysine residues of fibrinogen during the transamidation reaction.20 It permits the attachment of the peptide during the covalent stabilization of the fibrin matrix. In order to site-specifically attach the peptide to the HA macromer, the peptide was engineered to contain a cysteine residue, whose sulfhydryl group may participate in a series of mild chemoselective reactions with acrylates, vinyl sulfones or maleimides via Michael addition reaction; with α-haloacetates via SN2 reaction; and with 2dithiopyridyls via thiol-disulfide substitution reaction. Here, we have chosen a thiol-disulfide exchange reaction because it is rapid, readily reversible, and plays an important role in maintaining proper biological functions of living cells.32 A suitable pyridyl disulfide derivative of HA was prepared by amide coupling of native HA carboxylic groups to amino group of 2-(2pyridinyldithio) ethyl hydrazinecarboxylate 1 (Scheme 1). 15% of the carboxylate groups of HA were modified in this way, as evidenced by1H-NMR spectra of the obtained HA-SSPy product (Figure S1).


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Scheme 1. Synthesis of HA-Gln. The glutamine substrate peptide was linked to 2-dithiopyridyl modified HA (HA-SSPy) via a disulfide linkage in the course of thiol-disulfide exchange reaction. Next, we conjugated the peptide to the backbone of HA-SSPy derivative in PBS overnight. 1H-NMR spectrum of HA-SSPy showed four peaks of aromatic protons belonging to 2-thiopyridyl group (Figure 1a). The progress of the reaction was monitored by UV-vis spectrometry. The increase of intensity of the absorption peak at 340 nm indicated that 2thiopyridin was liberated during the substitution reaction and that a new disulfide bond between HA and peptide was formed. After purification by dialysis, the product was recovered by freeze-


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drying. 1H-NMR examination revealed peaks corresponding to different amino acids residues of Gln peptide in the regions before 1.9 ppm, as well as between 2.2 and 3.2 ppm and between 4.0 and 4.4 ppm (Figure 1b). Particularly, methyl groups of valine V and leucine L side chains could be detected at 0.85 and 0.9 ppm. Likewise, the peak at 1.58 – 1.68 ppm was attributed to methylene protons of arginine (R). Peaks in the region between 2.2 and 3.2 ppm should be indicative of side chains of glutamate (E), arginine (R), asparagine (N), and glutamine (Q) residues. Anomeric protons of HA were found at 4.5 ppm, and other sugar protons such as H-2’, H-3’, H-4’, H-5’, and H-5’ were clustered in 3.3 – 4.0 ppm region, and acetamide CH3CONHprotons at 2.0 ppm.

A


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B

Figure 1. 1H-NMR spectra of (a) starting derivative HA-SSPy and (b) its peptide conjugate HAGln. It is noteworthy that four signals of 2-dithiopyridyl (-SSPy) group at 7.6, 8.15, 8.25, and 8.6 ppm were not observed in the obtained conjugate. This result indicated elimination of the 2thiopyridine during thiol-disulfide exchange reaction, which happened during the coupling of Gln peptide to HA backbone. The minor peaks around 7.3 and 7.5 ppm could be attributed to the traces of the 2-thiopyridine side product still present in the product. Additionally, appearance of triplets at 2.8 and 3.0 ppm (marked a and b in Figure 1b) provided evidence for the formation of a -CH2-SS-CH2-CH- linker between HA and cysteine residue of Gln peptide. These triplets correspond to methylene protons of the linker adjacent to the disulfide bond. Taken together,


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these results clearly demonstrated that the FXIIIa acceptor substrate was linked to the HA macromolecule.

A

Ac-FKGGGPQGIWGQERCG -NH2 S

O

S

O

O n O S

S

O

O

O O

O n

O 6

O

S n

PEG-Lys S

O

B Factor XIIIa HA-Gln

+

PEG-Lys

Ca2+

Figure 2. (a) Structure of PEG-Lys. (b) Factor XIIIa-catalyzed hydrogel formation between HA and PEG macromolecules derivatized with glutamine and lysine substrate peptides respectively. The amidation reaction of glutamine (Q) amino acid residues on HA by amino groups of lysine (K) amino acid residues on PEG provides cross-links between the macromolecules (e.g. through the fused peptides, represented as red and blue ovals), resulting in the formation of hydrogels within approximately three and a half minutes (at 3% w/v) (insert)


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Transglutaminase-catalyzed cross-linking of hybrid HA-PEG hydrogels. By mixing aqueous solutions containing HA-Gln and 8-arm PEG macromers bearing the lysine donor peptide (PEGLys in Figure 2a) at equal stoichiometry of to the complementary substrates, hydrogel networks were rapidly formed in the presence of FXIIIa (Figure 2). This result confirms our hypothesis that after suitable derivatization, HA can be transformed into a substrate for the FXIIIa-mediated cross-linking. In this proof-of-principle example, the HA-Gln was cross-linked to a multi-arm PEG macromer bearing eight Lys residues, resulting in the formation of hybrid HA-PEG hydrogels. Additionally, we believe that this HA-Gln conjugate can be used to enhance the functionality of other compatible hydrogel systems, notably fibrin. In order to confirm that the transformation of the aqueous HA-PEG solution into a solid hydrogel is indeed the result of the TG-catalyzed acyl-transfer reaction, we prepared control formulations in which either the enzyme was absent, or HA-SSPy, a synthetic precursor of HAGln, replaced the HA-Gln peptide conjugate. Cross-linking kinetics were assessed by smallstrain oscillatory shear rheometry according to previously reported procedures.18 As expected, no hydrogel was formed when HA-SSPy was used in the enzymatic reaction, as was evidenced by comparable values for storage modulus (G′) and loss modulus (G′′) over the course of the experiment (Figure S1b). Similarly, a mixture of HA-Gln and PEG-Lys did not form a hydrogel without FXIIIa (Figure S1c). Conversely, when FXIIIa was added to HA-Gln + PEG-Lys mixture, a crossover of G′ and G′′, corresponding to transition into a hydrogel, occurred within 220 seconds after addition of the enzyme (Figure 3 and S1a). This gel point occurred at a very similar time point for the control formulation of PEG-Gln and PEG-Lys at the equivalent solid content of 3%.


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1000 100

G', Pa

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Elastic Modulus 3 % HA-PEG Elastic Modulus 3 % PEG Elastic Modulus 3 % HA PEG no XIIIa Elastic Modulus 3 % HASSPy PEG

10 1 0,1 0,01 0

200

400

600

800

1000

Time, sec Figure 3. Small-strain oscillatory shear rheometry to determine gelation kinetics in situ. Evolution of storage (G′) modulus (at 0.25Hz) over 1000 seconds after mixing of HA-Gln and PEG-Lys ([Gln]/[Lys] = 1) (red curve) and standard PEG-only mixture from PEG-Gln and PEGLys components (blue curve) at 3% concentration. For comparison, HA-Gln was substituted with the same amount of HA-SSPy (black curve), and factor XIIIa was omitted from the mixture of HA-Gln and PEG-Lys (green curve). The dependency of gelation time and mechanical properties of HA-PEG hydrogels on polymer concentration was then studied by conducting dynamic frequency sweep rheological


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experiments. We determined that the linear equilibrium storage modulus exhibits a plateau with respect to frequency between 0.1 and 0.8 Hz. As expected, gelation time decreased and storage modulus increased upon increasing polymer concentration from 2 to 4% (Figure 4a). At concentrations lower than 1.5%, HA-Gln did not form hydrogels with PEG-Lys at equimolar ratio of Gln and Lys peptides. Cross-linking, as defined by the gel point, was achieved in 4, 3.5, and 2 minutes for 2%, 3%, and 4% formulations, respectively. This data indicates that G′ and G′′ are nearly frequency-independent from 0.1 to at least 0.8 Hz, demonstrating largely elastic properties of the TG cross-linked HA hydrogels in this region. Indeed, the ratio of G′′ to G′ was found to be approximately 0.05, underscoring the highly elastic characteristics of these HA-PEG hydrogels.33 Notably, we found a clear linear dependence between of the storage modulus and polymer concentration between 2 and 4%, ranging from 1000 to 3000 Pa (0.25 Hz) (Figure 4b). The lowest concentration at which hydrogels could be formed was 1.5% and at this low concentration, the hydrogel was too soft to fit into the linear dependency range. Therefore, it is likely that the mechanical properties of these hybrid HA-PEG networks can be optimally tuned within the 2 – 4% concentration range, a regime of high cross-linking efficiency in which all functional groups (glutamine and lysine residues) are converted to network cross-links.


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B

A 3500

3000

4%

3000

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3%

G' at 0.25 Hz

G', G'', Pa

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4,0

Concentration, %

Figure 4. (a) Frequency sweeps of swollen, elastic HA-PEG networks formed at 2 (squares), 3 (triangles), and 4% (circles) concentration of the polymer precursors and [Gln]/[Lys] = 1. (b) Storage modulus (G′) of HA-PEG networks measured at 0.25 Hz as a function of the hydrogel concentration ([Gln]/[Lys] = 1). 3D culture of human mammary cancer cells within HA-PEG hydrogels. In order to validate the biological relevance of our HA-PEG hydrogels, we thought to encapsulate mammalian cells into these hydrogels and assess their proliferative potential and morphological features in comparison to standard PEG-based hydrogels. We first chose MCF7 cancer cells, which, notably, have been shown to express the HA receptor CD44.34 Cells were embedded within control PEG-only and HA-PEG hydrogels at a standard cell density of 0.1 million cells/mL, and cultured over a period of 48 hours, at which point they were fixed and imaged.


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Figure 5. MCF7 cells encapsulated in HA-PEG and standard PEG hydrogels. Soft hydrogels were prepared at 1.5% polymer concentration, while stiff hydrogels were prepared at 3% polymer concentration. PEG component contained at its terminals either proteolytically degradable Lys substrate peptides (i.e. including matrix metalloprotease GDQG↓IAGF sequence) or non-degradable substrate peptides. Scale bar is 100 µm. The hydrogels were prepared with different stiffnesses, designated here as ‘stiff’ (3% w/v, G′ = 2000 Pa) and ‘soft’ (1.5% w/v, G′ = 284 Pa) according to rheology measurements, based on different cross-linking densities of these hydrogels and contained either matrix metalloprotease (MMP)-sensitive (GPQG↓IWGQ) or MMP-insensitive substrate peptides (GDQG↓IAGF). This strategy allowed us to further tune the degradability of the hybrid networks. The incorporation of HA as a structural component of the hybrid HA-PEG network led to marked changes in morphology as compared to the PEG-only network (Figure 5). Notably, in both stiff and soft hydrogels, HA elicited a more invasive, and less epithelialized phenotype, along with enhanced proliferation and motility, suggesting that such a matrix may facilitate an epithelial to mesenchymal transition process.


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3D culture of myoblasts within HA-PEG hydrogels. In order to assess the versatility of this new hydrogel system, and to verify that the HA incorporation could also be relevant for supporting a mesenchymal phenotype, we encapsulated C2C12 mouse myoblasts into a range of gels formed from HA-PEG and PEG. Here, in addition to modulating hydrogel stiffness and MMP sensitivity, we also incorporated the adhesion ligand RGD, covalently tethered to the hydrogel backbone using TG chemistry.18 Images of the cells encapsulated in different hydrogels are presented in Figure 6. C2C12 cells in stiff materials, whether in HA-PEG or PEG-only hydrogels, appeared round, irrespective of the presence of RGD, suggesting that gel stiffness overrides other hydrogel properties in ensuring the maintenance of the spindle-shaped myoblast phenotype. In soft hydrogels, however, significant morphological differences were evidenced between PEG only and HA-PEG matrices. In soft PEG-only hydrogels, cellular processes extended most in MMPsensitive and RGD-containing matrices. Significant loss of cellular protrusion was detected when the matrix was MMP-insensitive, or devoid of RGD. In the soft HA-PEG hydrogels, cells spread more extensively than in PEG-only gels of the same stiffness. In particular, cellular protrusions were seen both in the MMP non-degradable matrices with adhesion ligands and in the MMP degradable analogs devoid of RGD ligands, suggesting that the addition of HA imparts some degree of adhesion and permissiveness to migration. Additionally, in all cases, cellular protrusion in HA-PEG hydrogels appeared to be longer and more pronounced than in the PEGonly counterparts (b).


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Figure 6. C2C12 cells encapsulated in HA-PEG and standard PEG hydrogels. (a) Hydrogels were prepared at 1.5% (soft) and 3% (stiff) polymer concentration, with MMP-sensitive or MMP-insensitive cross-links, with and without immobilized RGD ligand. Scale bar: 100 µm. (b) Protrusion length was quantified for all combinations of soft gels. SUMMARY AND CONCLUSIONS In this work, a hybrid polysaccharide-based hydrogel system was developed and shown to provide a permissive microenvironment for both epithelial tumor-derived and mesenchymal cells. This new HA derivative with substrate properties for TG was prepared by conjugation of NQEQVSPL peptide derived from N-terminus of α2-plasmin inhibitor (glutamine acceptor). The peptide was engineered to contain a cysteine residue for disulfide coupling with 2-dithiopyridyl modified hyaluronic acid (HA-SSPy). The resulting conjugate (HA-Gln) underwent sol-to-gel transformation upon mixing with an 8-arm PEG macromer terminated with the lysine donor (FKGG) peptide in the presence of FXIIIa. Cross-linking by FXIII-catalyzed acyl-transfer reaction between glutamine and lysine residues of the HA and PEG components was confirmed.


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We demonstrated that the mechanical properties of the HA-PEG hydrogels could be controlled by modulating the concentration of the macromolecular components. Increased adhesion and spreading of MCF7 cancer cells and C2C12 myoblasts in HA-PEG matrices as compared to purely PEG-based counterparts could be attributed to interaction of these cell types with the HA component of the hydrogel, possibly through hyaladherins, i.e. HA-binding receptors, such as CD44. These results suggest that HA-PEG hydrogels cross-linked by TG constitute an interesting class of hybrid artificial ECMs which could potentially lead to increased fidelity in mimicking the complex in-vivo extracellular microenvironment. Additionally, the incorporation of this novel derivatized HA could enhance the functionality of commonly used fibrin hydrogels. The in vivo tumor microenvironment is characterized by a high HA content, and such novel hydrogel formulations can therefore be particularly useful in the establishment of more biologically relevant in vitro tumor models. We expect that such approaches will not only allow for a more realistic in vitro assessment of anti-cancer drugs, but will also potentially enhance our understanding of the role of the cancer microenvironment in in vivo animal models. Indeed, the ability of these hydrogels to be readily administered via injection methods could allow for in vivo studies on the drug response to perturbed tumor microenvironments, and could point to novel models of drug diffusion and activity when administered systemically. Given that HA constitute a major polysaccharide component of the native ECM in numerous tissues, the HA-Gln substrate developed in this work constitutes a useful and important building block in the construction of synthetic artificial ECMs of ever-increasing fidelity to their in vivo counterpart. This novel hybrid material can be combined with an increasingly varied library of such matrices and components, which can increasingly be extended by orthogonal combination of TG chemistry to other bioconjugation schemes as well as by the


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inclusion of both growth factors and GAGs. We foresee that such advanced and tunable materials will allow for increasingly tailored and application-specific uses in regenerative medicine and disease modeling.

ASSOCIATED CONTENT Supporting Information. Figure S1. Small-strain oscillatory shear rheometry to determine gelation kinetics in situ

CORRESPONDING AUTHORS


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Adrian Ranga [email protected] Mechanical Engineering Department, Biomechanics Section KU Leuven, Celestijnenlaan 300 - box 2419, 3001 Leuven, Belgium Dmitri Ossipov [email protected] Science for Life Laboratory, Division of Polymer Chemistry, Department of ChemistryÅngström, Uppsala University, Uppsala, SE 751 21, Sweden

AUTHOR CONTRIBUTIONS A.R. and D.O. designed experiments, performed experimental work and analyzed the results. M.L. and J.H. designed experiments and contributed material support. All authors contributed to the writing of the manuscript. All authors have given approval to the final version of the manuscript.

ACKNOWLEDGMENT The research leading to these results has received funding from European Community’s Seventh Framework Programme (BIODESIGN).

REFERENCES


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