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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
37
showed almost negligible toxicity on MDA-MB-231cells where cell viability remained high
38
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
44
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
48
reinforce or augment tissues.3 Their porous structures are well-suited for encapsulation of
49
both cells and therapeutics.4 While numerous hydrogel systems have been made using
50
biocompatible natural polymers,5-6 hydrogels derived from synthetic polymers offer several
51
advantages over their naturally-derived counterparts.7 To this end, advances in controlled
52
polymerization methods enable unprecedented levels of control over polymer size, structure,
53
and chemical functionality, which in turn facilitates fine tuning of hydrogel structural and
54
mechanical properties.
55
In addition to polymer structure, the methods by which the constituent polymers are
56
crosslinked to form entangled networks can also have drastic effects on the structural,
57
mechanical, and rheological properties of synthetic hydrogels. For instance, physically
58
crosslinked hydrogels, which form due to non-covalent crosslinking mechanisms such as via
59
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
64
sensitive to changes in concentration for the same reason, and are susceptible to dissolution
65
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
74
when mixed together, resulting in gel formation. Ideally, the chemistry used for crosslinking
75
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,
79
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
81
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.
90
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
93
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’
95
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,
99
Korea. MDA-MB-231 breast cancer cell line was obtained from ATCC, U.S.A. and was
100
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.)
105
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
108
poly(ethylene glycol) (Mn = 8000 Da; 1.50 g, 0.19 mmol, 1 equiv), MTC-OCH2BnCl (0.394 g,
109
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
111
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,
115
followed by drying under reduced pressure, afforded the desired polymer P(BnCl)-PEG-
116
P(BnCl) as a white solid (1.80 g, 95% yield). 1H NMR (400 MHz, CDCl3): δ 7.36-7.26 (m,
117
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).
119 120
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
124
reaction mixture was transferred directly into a dialysis membrane (1000 MWCO) and
125
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,
127
Ar-H), 5.15 (s, 13H, -OCH2-BnN3), 4.33-4.28 (overlapping, 37H, -CH2N3 and -OCOOCH2-
128
and -OCH2CCH3-), 3.64 (s, 727H, PEG -OCH2CH2-), 1.24 (s, 19H, -CH3).
129 130
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,
133
1.31 mmol, 7 equiv), a Teflon-coated stir bar, and dry CH2Cl2 (1.5 mL). The contents of the
134
vial were stirred to allow the PEG to dissolve (MTC-OC6F5 is partially soluble in CH2Cl2 and
135
formed a homogeneous suspension), and triflic acid (20 mg, 0.13 mmol) was added to start
136
the polymerization. After stirring for 3 days at room temperature, the reaction mixture was
137
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
140
(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),
145
DBCO-Amine (0.18 g, 64 mmol), triethylamine (65 mg, 0.64 mmol), a Teflon-coated stir bar,
146
and dry THF (4 mL). The reaction mixture was stirred at ambient temperature for 18 h, and
147
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),
151
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,
153
14H, -CONHCH2-), 2.44 (br, 7H, -COCHH-), 1.84 (br, 7H, -COCHH-), 1.00 (m, 21H, -CH3).
154 155
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
157
with a stir bar. Dichloromethane was added and the monomer concentration was adjusted to 2
158
M. Once the initiator and monomer were completely dissolved, 8.3 µL (0.06 mmol) of DBU
159
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
165
glass vial. The glass vial was placed under hydrogen at room temperature with overnight
166
stirring. After that, the polymer was filtered using THF/methanol (1:1) solvent mixture, and
167
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).
173 174
Preparation and characterization of DOX-loaded PEG-P(Cat)12 micelles (Figure 1). DOX-
175
loaded PEG-P(Cat)12 micelles were prepared by a solvent evaporation technique. Briefly,
176
3 mg of DOX and 10 mg of PEG-P(Cat)12 were dissolved in 5 mL of methanol via
177
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
180
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
183
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.
210 211
In vitro drug release from DOX-loaded PEG-P(Cat)12 micelle. The in vitro release of DOX
212
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
214
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.
239
Subsequently, the cells were imaged at 40× magnification using LSM 510 DUO confocal unit
240
(Carl Zeiss, U.S.A.). 9 ACS Paragon Plus Environment
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241 242
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
261
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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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
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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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