Biodegradable Strain-Promoted Click Hydrogels for Encapsulation of

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Biodegradable strain-promoted click hydrogels for encapsulation of drug-loaded nanoparticles and sustained release of therapeutics Robert J. Ono, Ashlynn L.Z. Lee, Zhi Xiang Voo, Shrinivas Venkataraman, Bei Wei Koh, Yi Yan Yang, and James L Hedrick Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.7b00377 • Publication Date (Web): 06 Jul 2017 Downloaded from http://pubs.acs.org on July 7, 2017

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Biomacromolecules

Biodegradable strain-promoted click hydrogels for encapsulation of drug-loaded nanoparticles and sustained release of therapeutics Robert J. Ono,†,§ Ashlynn L. Z. Lee,‡,§ Voo Zhi Xiang,‡ Shrinivas Venkataraman, ‡ Bei Wei Koh,‡ Yi Yan Yang‡,* and James L. Hedrick†,* †

IBM Almaden Research Center, 650 Harry Road, San Jose, California 95120, United States Email: [email protected] ‡

Institute of Bioengineering and Nanotechnology, 31 Biopolis Way, Singapore 138669, Singapore Email: [email protected]

§

These authors contributed equally to the study.

Keywords: biomaterials, polycarbonates, hydrogels, nanomedicine, drug delivery, soft materials

ABSTRACT

21

Biodegradable polycarbonate-based ABA triblock copolymers were synthesized via

22

organocatalyzed ring-opening polymerization, and successfully formulated into chemically

23

crosslinked hydrogels by strain-promoted alkyne-azide cycloaddition (SPAAC). The synthesis

24

and crosslinking of these polymers is copper-free, thereby eliminating the concern over

25

metallic contaminants for biomedical applications. Gelation occurs rapidly within a span of 60

26

seconds by simple mixing of the azide- and cyclooctyne-functionalized polymer solutions.

27

The resultant hydrogels exhibited pronounced shear-thinning behavior and could be easily

28

dispensed through a 22G hypodermic needle. To demonstrate the usefulness of these gels as a

29

drug delivery matrix, doxorubicin (DOX)-loaded micelles prepared using catechol-

30

functionalized polycarbonate copolymers were incorporated into the polymer solutions to

31

eventually form micelle/hydrogel composites. Notably, the drug release rate from the

32

hydrogels was significantly more gradual compared to the solution formulation. DOX release

33

from the micelle/hydrogel composites could be sustained for 1 week while the release from

34

the micelle solution was completed rapidly within 6 h of incubation. Cellular uptake of the

35

released DOX from the micelle/hydrogel composites was observed at 3 h of incubation of 1 ACS Paragon Plus Environment

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human breast cancer MDA-MB-231 cells. A blank hydrogel containing PEG-(Cat)12 micelles

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showed almost negligible toxicity on MDA-MB-231cells where cell viability remained high

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at > 80% after treatment. When the cells were treated with the DOX-loaded micelle/hydrogel

39

composites, there was a drastic reduction in cell viability with only 25% of cells surviving the

40

treatment. In all, this study introduces a simple method of formulating hydrogel materials with

41

incorporated micelles for drug delivery applications.

42 43

INTRODUCTION

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Polymeric hydrogels have gained widespread interest of for their utility in a number of

45

growing fields, including tissue engineering, wound care, additive manufacturing, and drug

46

delivery. The physical and mechanical properties of hydrogels, which are comprised of water-

47

swollen networks of polymers, can be engineered to mimic cellular environments1-2 or to

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reinforce or augment tissues.3 Their porous structures are well-suited for encapsulation of

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both cells and therapeutics.4 While numerous hydrogel systems have been made using

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biocompatible natural polymers,5-6 hydrogels derived from synthetic polymers offer several

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advantages over their naturally-derived counterparts.7 To this end, advances in controlled

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polymerization methods enable unprecedented levels of control over polymer size, structure,

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and chemical functionality, which in turn facilitates fine tuning of hydrogel structural and

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mechanical properties.

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In addition to polymer structure, the methods by which the constituent polymers are

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crosslinked to form entangled networks can also have drastic effects on the structural,

57

mechanical, and rheological properties of synthetic hydrogels. For instance, physically

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crosslinked hydrogels, which form due to non-covalent crosslinking mechanisms such as via

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the hydrophobic interactions of amphiphilic polymers in aqueous media,8 can exhibit dynamic

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properties such as a shear-thinning response,9-11 making them useful materials as they can be 2 ACS Paragon Plus Environment

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injected through syringe needles.12-15

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This attractive property, however, comes at the expense of a relatively weaker material, owing

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to the non-covalent nature of the polymer network. Physical hydrogels are also generally very

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sensitive to changes in concentration for the same reason, and are susceptible to dissolution

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when subjected to diluting conditions. Covalently crosslinked hydrogels, on the other hand,

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are generally more mechanically robust, but as a consequence of their high elasticity are

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generally not amenable to injection via syringe for in vivo applications. Furthermore,

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concerns over the cytotoxicity of additive chemical crosslinking reagents and initiators further

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limit the biomedical application of these types of hydrogels.16

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To address limitations associated with both physically and covalently crosslinked hydrogels, a

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growing body of research focused on developing injectable chemically crosslinked hydrogel

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systems has emerged.17 Injectable hydrogels are typically comprised of two liquid components

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that contain complementarily reactive functional groups capable of forming covalent bonds

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when mixed together, resulting in gel formation. Ideally, the chemistry used for crosslinking

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of the injectable hydrogels must occur quickly, be insensitive to physiological conditions of

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pH and temperature, and be bio-orthogonal and chemically inert towards native biomolecules.

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To this end, several elegant hydrogel systems utilizing Diels–Alder,18 hydrazone,19

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disulfide,20 Michael addition,21 and azide–alkyne cycloaddition chemistries,22 among others,

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have been reported. The strain-promoted azide–alkyne cycloaddition (SPAAC)23, in particular,

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is a promising reaction for crosslinking of injectable hydrogels because of its fast reaction

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kinetics, compatibility with physiological conditions, and atom economy (i.e., does not form

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any byproducts), and also does not require a catalyst.24 Anseth and DeForest reported

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polyethylene glycol (PEG) based hydrogels crosslinked using SPAAC that could be degraded

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with spacial and temporal control using light.25-26 Adronov27 and Song28 used SPAAC to

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fabricate degradable hydrogels and evaluated their cytocompatibility. Here, we report the 3 ACS Paragon Plus Environment

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synthesis of PEG and biodegradable aliphatic polycarbonate-based azide- and cyclooctyne-

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bearing polymers and their SPAAC-promoted crosslinking to form hydrogels. The resulting

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hydrogels were embedded with doxorubicin (DOX)-loaded micelles, and evaluated as an

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injectable, slow-release drug delivery matrix.

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Demand for such controlled delivery systems exists29-30 as current drug-loaded nanoparticles

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are delivered clinically via intravenous infusion.31-36 that minimally lasts for 30 min.37-38 By

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incorporating DOX-loaded micelles into the hydrogel, a sustained release of the drug-loaded

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micelles into the circulatory system can occur for a prolonged period of time, thereby

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reducing the number of administrations required and potentially improving patients’

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convenience and compliance to the treatment.

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EXPERIMENTAL SECTION

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Materials. Doxorubicin (DOX) hydrochloride was purchased from Boryung Pharmaceutical,

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Korea. MDA-MB-231 breast cancer cell line was obtained from ATCC, U.S.A. and was

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cultured in RPMI1640 medium. Culture medium was supplemented with 10% fetal bovine

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serum

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carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium (MTS) and phenazine ethosulfate

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(PES) in MTS solution was purchased from Promega, U.S.A. Dibenzocyclooctyne-amine

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(DBCO-Amine) was purchased from Click Chemistry Tools (Scottsdale, AZ, U.S.A.)

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Synthesis

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Synthesis of P(BnCl)-PEG-P(BnCl) (Fig. 1A). The following procedure is representative. In a

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nitrogen filled glovebox, a 20 mL glass vial was charged with azeotropically dried

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poly(ethylene glycol) (Mn = 8000 Da; 1.50 g, 0.19 mmol, 1 equiv), MTC-OCH2BnCl (0.394 g,

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1.31 mmol, 7 equiv), TU (19 mg, 0.05 mmol), a Teflon-coated stir bar, and dry CH2Cl2 (1.5

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mL). The contents of the vial were allowed to dissolve, and DBU (7.6 mg, 0.05 mmol) was

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added to start the polymerization. After stirring for 30 min at room temperature, an excess of

(FBS)

and

100

U/ml

penicillin.

3-(4,5-Dimethylthiazol-2-yl-2-yl)-5-(3-

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benzoic acid (30 mg, 0.24 mmol) was added to quench the catalyst and stop the

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polymerization. The crude reaction mixture was then precipitated into diethyl ether (40 mL).

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Three cycles of trituration, centrifugation, and decantation of the diethyl ether supernatant,

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followed by drying under reduced pressure, afforded the desired polymer P(BnCl)-PEG-

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P(BnCl) as a white solid (1.80 g, 95% yield). 1H NMR (400 MHz, CDCl3): δ 7.36-7.26 (m,

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26H, Ar-H), 5.14 (s, 13H, -OCH2-BnCl), 4.57 (s, 13H, -CH2-Cl), 4.28 (br, 26H, -OCOOCH2-

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and -OCH2CCH3-), 3.64 (s, 727H, PEG -OCH2CH2-), 1.24 (s, 20H, -CH3).

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Synthesis of P(BnN3)-PEG-P(BnN3) (Fig. 1A). The following procedure is representative. To

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a 20 mL glass vial was added P(BnCl)-PEG-P(BnCl) (1.80 g, 1.3 mmol BnCl groups), sodium

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azide (0.100 g, 1.54 mmol), DMF (3 mL), and a Teflon-coated stir bar. The vial was loosely

123

sealed with a rubber septum and the reaction mixture stirred at room temperature for 4 h. The

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reaction mixture was transferred directly into a dialysis membrane (1000 MWCO) and

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dialyzed against deionized water. Lyophilization afforded the desired polymer P(BnN3)-PEG-

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P(BnN3) as a white solid (1.40 g, 77% yield). 1H NMR (400 MHz, CDCl3): δ 7.31 (br, 26H,

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Ar-H), 5.15 (s, 13H, -OCH2-BnN3), 4.33-4.28 (overlapping, 37H, -CH2N3 and -OCOOCH2-

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and -OCH2CCH3-), 3.64 (s, 727H, PEG -OCH2CH2-), 1.24 (s, 19H, -CH3).

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Synthesis of P(C6F5)-PEG-P(C6F5) (Fig. 1A). The following procedure is representative. In a

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nitrogen filled glovebox, a 20 mL glass vial was charged with azeotropically dried

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poly(ethylene glycol) (Mn = 8000 Da; 1.50 g, 0.188 mmol, 1 equiv), MTC-OC6F5 (0.427 g,

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1.31 mmol, 7 equiv), a Teflon-coated stir bar, and dry CH2Cl2 (1.5 mL). The contents of the

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vial were stirred to allow the PEG to dissolve (MTC-OC6F5 is partially soluble in CH2Cl2 and

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formed a homogeneous suspension), and triflic acid (20 mg, 0.13 mmol) was added to start

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the polymerization. After stirring for 3 days at room temperature, the reaction mixture was

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removed from the glovebox and precipitated into diethyl ether (40 mL). Three cycles of 5 ACS Paragon Plus Environment

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trituration, centrifugation, and decantation of the diethyl ether supernatant, followed by drying

139

under reduced pressure, afforded the desired polymer P(C6F5)-PEG-P(C6F5) as a white solid

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(1.39 g, 72% yield). 1H NMR (400 MHz, CDCl3): δ 4.45-4.30 (br, 26H, -OCOOCH2- and -

141

OCH2CCH3-), 3.63 (s, 727H, PEG -OCH2CH2-), 1.49 (two s, 19H, -CH3).

142 143

Synthesis of P(DBCO)-PEG-P(DBCO) (Scheme 1). The following procedure is representative.

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A 20 mL glass vial was charged with P(C6F5)-PEG-P(C6F5) (1.00 g, 0.64 mmol C6F5 groups),

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DBCO-Amine (0.18 g, 64 mmol), triethylamine (65 mg, 0.64 mmol), a Teflon-coated stir bar,

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and dry THF (4 mL). The reaction mixture was stirred at ambient temperature for 18 h, and

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transferred directly into a dialysis membrane (1000 MWCO). Two dialyses were performed

148

sequentially, first against 1:1 v/v acetonitrile:isopropanol, followed secondly against

149

deionized water. Finally, lyophilization afforded the desired polymer as an off-white solid

150

(0.94 g, 88% yield). 1H NMR (400 MHz, CDCl3): δ 7.65 (s, 9H, Ar-H), 7.35 (br, 64H, Ar-H),

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6.76-6.49 (m, 8H, amide NH), 5.10 (br s, 8H, -NCHH-), 4.25-4.13 (m, 26H, -OCOOCH2- and

152

-OCH2CCH3-), 3.64 (s, 832H, overlapping PEG -OCH2CH2- and -NCHH-), 3.30-3.21 (br,

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14H, -CONHCH2-), 2.44 (br, 7H, -COCHH-), 1.84 (br, 7H, -COCHH-), 1.00 (m, 21H, -CH3).

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Synthesis of PEG-P(Cat)12. In a glove-box, 0.556 g (0.056 mmol) of 10 kDa MPEG-OH

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initiator and 0.376 g (1 mmol) of MTC-ProtCat were charged in a 20 mL glass vial equipped

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with a stir bar. Dichloromethane was added and the monomer concentration was adjusted to 2

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M. Once the initiator and monomer were completely dissolved, 8.3 µL (0.06 mmol) of DBU

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was added to initiate the polymerization. After 1 h of stirring at room temperature, the

160

reaction was quenched with 30 mg of benzoic acid. Subsequently, the polymer intermediate

161

was purified via precipitation twice in cold diethyl ether, and was dried on a vacuum line until

162

a constant weight was achieved.

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The protected polymer was then deprotected by hydrogenation process. It was first dissolved

164

in 14 mL of methanol and THF (1:1), followed by addition of 2 spatulas of Pd/C into a 50 ml

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glass vial. The glass vial was placed under hydrogen at room temperature with overnight

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stirring. After that, the polymer was filtered using THF/methanol (1:1) solvent mixture, and

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the collected polymer was dried under vacuum, followed by reprecitation in cold diethyl ether

168

twice. Finally, the solvents were removed and the polymer was lyophilized to obtain an off-

169

white polymer (0.35g, 80% yield, PDI: 1.25). 1H NMR (400 MHz, CDCl3, 22 ºC): δ 7.67-7.27

170

(m, 156H, -C6H5 & -COOCCH2- ), 6.97-6.81 (m, 13H, -COOCH2CH-), 5.32-5.00 (m, 52H, -

171

COCH2-), 4.57-4.32 (m, 52H, -OC2H4O-), 4.31-4.10 (m, 52H, -COOCH2-), 3.89-3.42 (-

172

OCH2CH2- from 10kDa PEG), 3.38 (s, 3H, CH3-PEG-), 1.29-1.04 (m, 39H, -CH3).

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Preparation and characterization of DOX-loaded PEG-P(Cat)12 micelles (Figure 1). DOX-

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loaded PEG-P(Cat)12 micelles were prepared by a solvent evaporation technique. Briefly,

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3 mg of DOX and 10 mg of PEG-P(Cat)12 were dissolved in 5 mL of methanol via

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ultrasonication. Methanol was evaporated under reduced pressure using a rotatory evaporator

178

at 50 °C and this was followed by rapid addition of 5 mL of HPLC grade water at 50 °C and

179

sonicated for 30 min. To remove the insoluble residual drug, the mixture was centrifuged for

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5 min at 4000 rpm, 25°C, followed by filtration using 0.22 µm nylon syringe filters. Particle

181

size was measured via dynamic light scattering (scattering angle: 90°) equipped with a He-Ne

182

laser beam at 658 nm (Malvern Instruments Zetasizer Nano ZS, UK). The particles were also

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visualized via transmission electron microscopy (TEM)39.The content of DOX in PEG-

184

P(Cat)12 micelles was quantified by measuring the absorbance of the samples using a UV-Vis

185

spectrophotometer (Shimadzu UV3600, Japan) at 480 nm. The drug loading level and

186

encapsulation efficiency was calculated based on the ratio of the amount of drug encapsulated

187

in the micelles to the amount of drug-loaded micelles and the ratio of the amount of drug

188

encapsulated in the micelles to the initial added amount of drug respectively. 7 ACS Paragon Plus Environment

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Rheological Experiments. The rheological properties of the hydrogels were analyzed on an

190

ARES-G2 rheometer (TA Instruments, USA) equipped with a plate-plate geometry of 8 mm

191

diameter. Measurements were recorded by equilibrating the gels at 25 °C between the plates at

192

a gap of 1.0 mm. The data were collected under controlled strain of 2.0% and a frequency

193

scan of 1.0 to 50 rad/s. Gelation properties of the polymer suspension was monitored by

194

measuring the shear storage modulus (G’) and the loss modulus (G”). Thixotropic property of

195

the hydrogels was studied by measuring the changes in viscosity as a function of shear rate

196

from 0.1 to 25 s-1.

197 198

Scanning electron microscope (SEM) imaging of hydrogel. To minimize morphological

199

perturbations, the hydrogels were cryo-fixed by snap freezing with liquid nitrogen. A day of

200

freeze-drying process was then followed. The morphology of the gel was observed using a

201

scanning electron microscope (SEM) (Jeol JSM-7400F, Japan).

202 203

Swelling studies. To estimate structural parameters such as molecular weight between cross-

204

links, effective cross-link density and mesh size of the polymeric network, swelling studies

205

was performed. Briefly, hydrogels (100 µL) was prepared and dried in an oven at 70 °C for 24

206

h. After that, their dry mass was measured. Phosphate-buffered saline (PBS, pH 7.4, 100 µL)

207

was then added to the dried gels and incubated at 37 °C for 24 h. Subsequently, the excess

208

PBS was removed from the swollen hydrogels and the mass was recorded. The mesh size of

209

the hydrogels was then determined using the Flory-Rehner calculation method40-42.

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In vitro drug release from DOX-loaded PEG-P(Cat)12 micelle. The in vitro release of DOX

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was studied in fresh phosphate-buffered saline (PBS,pH 7.4) and acetate buffer (200 mM, pH

213

5.0). DOX-loaded micelle solution (0.5 mL) was transferred to dialysis bags with molecular

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weight cut-off of 1 kDa. The dialysis bags were then immersed into 15 mL of buffers 8 ACS Paragon Plus Environment

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respectively in plastic tubes. The tubes were shaken at 100 rpm and incubated at 37 °C. At

216

each designated time point, all the release media was withdrawn and replaced with fresh

217

buffers. The released DOX concentrations in the 1 mL solutions were measured by a UV-Vis

218

spectrophotometer (Shimadzu UV3600, Japan) at 480 nm.

219 220

In vitro drug release from DOX-loaded micelle/hydrogel composites. The DOX-loaded

221

micelle/hydrogel composite was prepared by first dissolving the alkyne- and azide-containing

222

polymers (P(DBCO)-PEG-P(DBCO) and P(BnN3)-PEG-P(BnN3)) using the DOX-loaded

223

PEG-P(Cat)12 micelle solution (Fig.

224

concentrations of the two solutions were then mixed for gelation to occur. The resulting

225

hydrogels (250 µl) were loaded into transwell inserts (8.0 µm cutoff) (Corning, U.S.A) and

226

then placed in conical tubes containing 20 mL PBS. The tubes were shaken at 100 rpm and

227

incubated at 37 °C. At each designated time point, 1 mL of the release medium was

228

withdrawn and replaced with fresh buffers. The absorbance of the release medium at 480 nm

229

was then measured using a UV-Vis spectrophotometer (Shimadzu UV3600, Japan).

1C). Subsequently, equivalent volumes and

230 231

Confocal microscopy studies. MDA-MB-231 cells were seeded onto borosilicate chambered

232

cover glass (Nunc, U.S.A.) at a density of 20 × 104 cells per well, and cultivated in 500 µL of

233

growth medium. The next day, the spent growth medium was removed from each well and

234

replaced with fresh medium. Next, DOX-loaded micelle/hydrogel composites (40 µL) were

235

then added. After 3 h of incubation at 37°C, the treatment solution was then removed. The

236

cells were rinsed thrice with 500 µL of PBS and fixed with 200 µL of formalin solution

237

(Sigma Aldrich, U.S.A.) for 15 min. The formalin solution was then removed and cells were

238

rinsed with 500 µL of PBS thrice. Thereafter, 2 drops of prolong gold with DAPI was added.

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Subsequently, the cells were imaged at 40× magnification using LSM 510 DUO confocal unit

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(Carl Zeiss, U.S.A.). 9 ACS Paragon Plus Environment

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Cytotoxicity test. MDA-MB-231 cells were seeded onto 96-well plates at a cell density of 1.5

243

× 104 cells per well, and were cultivated with 100 µl of RPMI growth medium. The cell plates

244

were placed in the incubator at 37°C for 24 h to reach a confluency of 70-80 %. When the

245

targeted cell confluency was reached, the spent growth medium was removed from the wells

246

and 100 µl of RPMI was added to each well. DOX-loaded micelle/hydrogel composites were

247

prepared at 1× (i.e. without diluting the micelle solution) and 0.2× micelle concentrations,

248

where 0.2× is used to simulate a lower amount of DOX present in the composites at a later

249

time point (e.g. after 5-day release). Following this, 10 µL of the hydrogel was added to the

250

cells and incubated for another 48 h at 37°C. To test for cell viability after treatment, MTS

251

solution (Promega, USA) and RPMI were then mixed at a volume ratio of 1:4 and 100 µl of

252

this mixture was then added to each well. Subsequently, the cells were left to incubate in the

253

dark at 37°C for two hours. Untreated cells were then used as the control. To measure the

254

amount of soluble formazan produced by reduction of MTS by viable cells, the absorbance of

255

the samples was measured at 490 nm with a micro-plate reader (Tecan, U.S.A.) and the

256

readings were then expressed as a percentage of the control. At the end of the treatment, two-

257

tailed Student’s t test was used to statistically evaluate the differences in cell viability between

258

different treatment conditions, and P≤0.05 indicates a statistically significant difference.

259 260

RESULTS AND DISCUSSION

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Synthesis

262

Azide-

263

organocatalyzed ring-opening polymerization (ROP) of the respective cyclic carbonate

264

monomers (Scheme 1). P(BnN3)-PEG-P(BnN3) was prepared by polymerization of benzyl

265

chloride-functionalized cyclic carbonate (MTC-OCH2BnCl) using PEG diol as a

266

macroinitator, followed by substitution using sodium azide.43 Cyclooctyne-functionalized

and

cyclooctyne-functionalized

triblock

copolymers

were

synthesized

by

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P(DBCO)-PEG-P(DBCO) was prepared by first polymerizing MTC-OC6F5 in the presence of

268

PEG to afford the reactive polymer intermediate (C6F5)-PEG-(C6F5),44 and displacing the

269

pentafluorophenyl ester groups with DBCO-Amine in a second step. 1H NMR spectroscopy

270

confirmed complete conversion of the reactive P(BnCl)-PEG-P(BnCl) and P(C6F5)-PEG-

271

P(C6F5) precursors to the desired azide- and DBCO-functionalized polymers, respectively. As

272

shown in Table 1, various molecular weights of PEG (i.e., 8, 18.5, and 20 kDa) were used to

273

initiate the ROP, and the block lengths (e.g., degree of polymerization (DP)) of the

274

hydrophobic polycarbonate blocks were adjusted to afford a series of triblock copolymers

275

across a range of molecular weights and hydrophobic-to-hydrophilic block ratios.

276 277

DOX loading in PEG-P(Cat)12 micelles

278

DOX is an FDA-approved anti-cancer drug effective against many types of malignancies.

279

DOX contains many chemical moieties that can interact effectively with the catechol moieties

280

present in PEG-P(Cat)12 through hydrogen bonding and/or π-π stacking and was selected as

281

the drug-of-interest to be encapsulated in the PEG-P(Cat)12 micelles. The thin film-hydration

282

preparation method was used to physically entrap DOX into the micelles. DOX and the

283

catechol moieties interact effectively and high encapsulation efficiency and loading level

284

(were achieved (99 ± 1% and 38 ± 2%, respectively).

285 286

Using dynamic light scattering technique, the particle size and PDI of the blank and DOX-

287

loaded micelles were measured to be (42 nm, 0.26) and (44 nm, 0.11) respectively. PDI of

288

DOX-loaded micelles were much lower compared to blank micelles probably due to the

289

strong interaction between DOX and catechol moieties in the micelle core. TEM was also

290

performed and it showed that both blank and dox-loaded nanoparticles were < 50 nm (Figure

291

2).

292

Physical properties of hydrogels 11 ACS Paragon Plus Environment

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293

Hydrogel formation was initiated by mixing together equivalent volumes of aqueous solutions

294

of azide- and DBCO-functionalized polymers. Notably, gelation occurred quickly, within 60 s

295

of mixing, as determined by a simple vial inversion test (Figure S1). Polymer crosslinking via

296

SPAAC was verified by taking an FTIR spectrum of a lyophilized sample of the gel, which

297

revealed complete consumption of the characteristic -N3 signal at 2100 cm-1 (Figure S2).

298

From Table 1, the gelator concentration significantly influenced the storage modulus (G’) and

299

loss modulus (G”) of the hydrogel, particularly for polymers with longer hydrophobic blocks.

300

For instance, a 20% concentration increment (10 to 12 wt.%) of the polymers resulted in a

301

more than 2 fold increase in the G’ value from 81 to 319 Pa (Table 1, entries 4 and 5,

302

respectively) due to a higher cross-linking degree. Counterintuitively, we observed that

303

polymers with longer PEG chains tend to gel at lower concentrations. For example,

304

comparing between polymers with similar DP of DBCO and azide moieties, those with longer

305

PEG length (20 kDa) were able to form hydrogels at lower concentrations than those with

306

shorter PEG length (8 kDa). We believe this is due to the ability of the longer polymer chains

307

to act as physical crosslinks between flower-like micelles across larger distances, effectively

308

lowering the concentration of polymer chains necessary to achieve network formation. This

309

might also be the reason that at the same polymer concentration (10 wt.%) the hydrogel with 8

310

kDa of PEG (Table 1 Entry 4) had G’ and G” of 81 and 49 Pa, respectively, while the

311

hydrogel with 20 kDa of PEG had higher G’ and G” (323 and 132 Pa, respectively) (Table 1,

312

Entry 13).

313 314

To investigate the effects of pre-formed micelles on the rheological properties of the

315

hydrogels, PEG-(Cat)12 micelles loaded with DOX were added to the triblock mixtures and

316

the storage moduli of the hydrogels were found to decrease (Table 1, entry 5*, 12* and 13*)

317

compared to those without pre-formed micelles. In addition, the hydrogels and DOX-loaded

318

micelle/hydrogel composites had a shear-thinning property, i.e. viscosity decreased as 12 ACS Paragon Plus Environment

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319

increasing shear rate (Figure 3A), indicating injectability. Injectability was further confirmed

320

by manually loading the hydrogels in syringes and dispensing it through a 22G needle.

321

Network mesh size, which is the average spacing between neighboring polymer chains

322

is an important parameter to describe hydrogel swelling and permeability40, 45 was calculated

323

using the Flory-Rehner calculation method40-42. Swelling studies were performed, and the

324

mesh size of the blank hydrogel, blank micelle/hydrogel and DOX-loaded micelle/hydrogel in

325

their equilibrium swollen states were estimated to be 4.2, 3.8 and 3.8 nm, respectively,

326

implying that the presence of micelles did not affect nextwork mesh size significantly.

327 328 329

SEM imaging of hydrogels

330

SEM imaging was performed to examine the cross-section morphology of the hydrogels.

331

From Figure 3B, blank hydrogel (Table 1, Entry 13) showed porous structure with pore size

332

ranging between 5 to 20 µm. Comparatively, in Figure 3C, the incorporation of DOX-loaded

333

micelles led to a hydrogel with larger pore size ranging between 20 to 200 µm. This is

334

coherent to the lower viscoelastic moduli46 as seen in Table 1, where the incorporation of

335

DOX-loaded micelles resulted in hydrogels with lower mechanical strength. The presence of

336

the micelles might reduce crosslinking density due to steric hindrance, thus leading to larger

337

pores and lower mechanical strength. Although the theoretical network mesh size of the

338

hydrogels was smaller than the micelles, the pores present in the hydrogels were large enough

339

to allow the DOX-loaded micelles to release out.

340 341 342

In vitro drug release

343

From Figure 3D, the release of DOX from micelle/hydrogel composite (Table 1, Entry 13*)

344

in PBS was gradual and could be sustained for one week. This suggests that the 13 ACS Paragon Plus Environment

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345

micelle/hydrogel composite can serve as a reservoir for sustained delivery of the drug into the

346

human body. To compare against the drug release rates of DOX-loaded micelles and to study

347

the pH-dependence of drug release rates, DOX-loaded micelles were immersed in release

348

media at pH 7.4 versus pH 5.0. Drug release from the micelles was faster than that from the

349

micelle/hydrogel composite (Figure 3E). For example, at 6 h, pH 7.4, the release from

350

hydrogel was 16% ± 1% while it was 79% ± 3% from the nanoparticles; thereby

351

demonstrating the advantage of loading the micelles into hydrogel for sustained drug delivery.

352

The release of DOX from micelles was dependent on the pH of the release media. At pH 5.0,

353

which mimics the endolysosomal microenvironment, the release of DOX from the micelle

354

was significantly higher compared to pH 7.4 at each time point throughout the experiment

355

possibly owing to increased solubility of DOX at the lower pH. Many enzymes exist in the

356

endolysosomes, which can digest drug molecules. Accelerated drug release from the

357

endolysosomes is desirable as released DOX molecules may penetrate the endolysosomal

358

membrane, enter the cytosol and then the nucleus to perform its anticancer function.

359 360

Cellular uptake and anti-cancer effects of DOX released from DOX-loaded micelle/hydrogel

361

composites

362

Confocal microscopy imaging was performed on MDA-MB-231 cells treated with DOX-

363

loaded micelle/hydrogel composites. As shown in Figure 4A-C, DOX that was released from

364

the hydrogels were effectively taken up into the cancer cells, and was able to penetrate into

365

the nuclei.

366

DOX is an FDA approved medication for the treatment of many cancer types. Since its

367

introduction in the early 1970s, DOX has been considered as one of the most active cytotoxic

368

agents for the treatment of breast cancer.47 Hence, for our study, human breast cancer MDA-

369

MB-231was selected for cytotoxicity studies. Figure 5 shows that treatment using samples

370

that do not contain DOX, i.e. blank hydrogel containing PEG-(Cat)12 micelles, showed almost 14 ACS Paragon Plus Environment

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371

negligible toxicity on MDA-MB-231cells where cell viability remained high at > 80% after

372

48 h treatment. When the cells were treated with the DOX-loaded micelle/hydrogel

373

composites (1×; 0.7 mg/mL DOX), there was a drastic reduction in cell viability with only

374

28% of cells surviving the treatment. It is worth noting that PEG-(Cat)12 micelles showed no

375

toxicity (viability ∼100%) at concentrations up to 2 mg/mL, which is higher than the

376

concentration used for the preparation of micelle/hydrogel composites. This indicates that the

377

cytotoxicity of DOX-loaded micelle/hydrogel composites was from DOX. When the

378

concentration of DOX-loaded micelles was reduced in the hydrogel to 0.2 times (0.2×; 0.14

379

mg/mL DOX), cell viability was 43%; indicating that DOX was still bioactive even at the

380

lower concentration to act on the cancer cells.

381 382

CONCLUSIONS

383

In this study, a series of biodegradable alkyne- and azide-functionalized polycarbonate-based

384

ABA triblock copolymers were synthesized via organocatalyzed ring-opening polymerization,

385

and successfully formulated into chemically crosslinked hydrogels by strain-promoted alkyne-

386

azide cycloaddition (SPAAC). This copper-free synthesis and gelation eliminates the concern

387

over adverse body reactions due to metal contaminants during pharmaceutical applications.

388

DOX-loaded micelles prepared using catechol-functionalized polycarbonate copolymers were

389

incorporated to form micelle/hydrogel composites. The drug release rate from the hydrogels

390

was markedly more gradual compared to the micelle solution formulation. The DOX-loaded

391

micelle/hydrogel composites were extremely effective in eliminating cancer cells. In all, the

392

results from this study represent a facile system of formulating hydrogel materials with

393

incorporated micelles for drug delivery applications.

394 395 396 397

Supporting Information Supporting Information is available online from the Wiley Online Library or from the author. 15 ACS Paragon Plus Environment

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398 399 400 401 402 403 404 405 406 407 408 409 410 411 412 413 414 415 416 417 418 419 420 421 422 423 424 425 426 427 428 429 430 431 432 433

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Acknowledgements This work was supported by IBM Almaden Research Center, USA and the Institute of Bioengineering and Nanotechnology (Biomedical Research Council, Agency for Science, Technology and Research), Singapore.

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434 435 436 437 438 439

Scheme 1. Synthesis of triblock copolymers P(BnN3)-PEG-P(BnN3) and P(DBCO)-PEGP(DBCO).

440

441

442 443 444 445 446 447 448 449

Figure 1. (A) Encapsulation of DOX using PEG-P(Cat)12. (B) Preparation of DOX-loaded hydrogels via SPAAC crosslinking with DOX-loaded micelles being incorporated within the polymer network.

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1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

450

Table 1. Rheological properties of hydrogels in various compositions. MW PEG

DBCO DP (m)

Azide DP (m)

Conc. (%)

Dissolve? (Y/N)

Gel? (Y/N)

G'

G''

1

8000

3.25

3.2

4

Y

N

-

-

2

8000

3.25

3.2

6

Y

N

-

-

3

8000

3.25

3.2

8

Y

N

-

-

4

8000

3.25

3.2

10

Y

Y

81 ± 13

49 ± 8

5

8000

3.25

3.2

12

Y

Y

319 ± 27

140 ± 9

5*

8000

3.25

3.2

12

Y

Y

302 ± 26

120 ± 13

6

8000

5

5.5

1

N

N

-

-

7

8000

5

5.5

2

N

N

-

-

8

8000

6

5.5

2

N

N

-

-

9

8000

6

5.5

4

N

N

-

-

10

20000

3

3.6

3

Y

N

-

-

11

20000

3

3.6

4

Y

Y

245 ± 10

38 ± 6

12

20000

3

3.6

6

Y

Y

274 ± 21

65 ± 10

12*

20000

3

3.6

6

Y

Y

119 ± 14

39 ± 12

13

20000

3

3.6

10

Y

Y

323 ± 30

132 ± 14

13*

20000

3

3.6

10

Y

Y

161 ± 25

87 ± 9

14

18500

7.3

6

1

N

N

-

-

15

18500

7.3

6

2

N

N

-

-

16

18500

7.3

6

4

N

N

-

-

Sample

451 452 453

454 455 456 457

Page 18 of 24

Average values of G' and G'' (Pa) measured between 10 to 25 rad/s. *Contains DOX-loaded micelles

Figure 2. TEM images of (A) blank micelles and (B) DOX-loaded micelles. Scale bar represents 50 nm. 18 ACS Paragon Plus Environment

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458 459 460

461 462 463 464 465 466 467 468 469 470 471

472 473 474 475 476 477

Figure 3. (A) Rheology flow sweep of blank micelle/hydrogel and DOX-loaded micelle/hydrogel composite (Table 1, Entry 13 and 13*, respectively) at 25°C. SEM images of cryo-fixed (B) blank hydrogel (Table1, Entry 13); (C) DOX-loaded micelle/hydrogel (Table1, Entry 13*). For SEM sample preparation, the hydrogels were rapidly transferred into a chamber filled with liquid nitrogen, followed by a day of freeze-drying process. Release profiles of DOX from (D) DOX-loaded micelle/hydrogel composite in PBS, pH 7.4 at 37°C and (E) DOX-loaded micelles solution in PBS (pH 7.4) and acetate buffer (pH 5.0). Each condition was tested in triplicate. Each error bar represents the average ± standard deviation of the replicates.

Figure 4. Confocal microscopy images of cells after 3-h treatment with DOX-loaded micelle/hydrogel composites: (A) DAPI staining of nuclei; (B) DOX localization; and (C) merged image of (A) and (B). 19 ACS Paragon Plus Environment

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478

479 480 481 482 483 484

Figure 5. Cell viability of MDA-MB-231 cells after 48-h treatment using the blank hydrogel and DOX-loaded micelle/hydrogel composites at different concentrations (DOX: 0.7 mg/mL and 0.14 mg/mL at 1× and 0.2×, respectively).

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485 486 487 488 489 490 491 492 493 494 495 496 497 498

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The table of contents entry should be 50−60 words long (max. 400 characters), and the first phrase should be bold. The entry should be written in the present tense and impersonal style. Keyword (see list) R. J. Ono,†,§ A. L. Z. Lee,‡,§ V. Z. Xiang,‡ B. W. Koh,‡ Shrinivas Venkataraman, ‡ Y. Y. Yang‡,* and J. L. Hedrick†,* Biodegradable strain-promoted click hydrogels nanoparticles and sustained release of therapeutics

for

encapsulation

of

drug-loaded

ToC figure ((Please choose one size: 55 mm broad × 50 mm high or 110 mm broad × 20 mm high. Please do not use any other dimensions))

499 500 501

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502

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