On-Demand Cascade Release of Hydrophobic Chemotherapeutics from a Multicomponent Hydrogel System

Article abstract of DOI:10.1021/acsbiomaterials.8b00166

A multicomponent hydrogel drug-delivery platform with covalent incorporation of core-cross-linked star (CCS) polymers has been fabricated through dual cross-linking reactions. The presence of amphiphilic CCS polymers enhances the mechanical stability of the network and allows for immobilization of hydrophobic chemotherapeutics (e.g., doxorubicin: DOX) within a hydrophilic matrix. Thanks to thiol-responsive disulfide segments, the hybrid network is selectively degradable upon exposure to the glutathione-rich tumor environment, leading to an on-demand cascade release of payloads in a prolonged manner. After network degradation, the integrity of the therapeutic cargos is still well-protected by the CCS polymeric carriers. A cellular study demonstrated efficient internalization of released drug-containing CCS polymers by HeLa cells, intracellular delivery of encapsulated drugs following pH-mediated erosion of the CCS containers, and excellent cytocompatibility of this hydrogel system.

Full text of DOI:10.1021/acsbiomaterials.8b00166

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Controlled Release and Delivery Systems
On-demand Cascade Release of Hydrophobic
Chemotherapeutics from a Multi-component Hydrogel System
Dunyin Gu, Shereen Siew Ling Tan, Andrea Janet O'Connor, and Greg G. Qiao
ACS Biomater. Sci. Eng., Just Accepted Manuscript • Publication Date (Web): 15 Apr 2018
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Onꢀdemand Cascade Release of Hydrophobic
Chemotherapeutics from a Multiꢀcomponent
Hydrogel System
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Dunyin Gu, Shereen Tan, Andrea J. O’Connor, Greg G. Qiao*
Department of Chemical and Biomolecular Engineering, The University of Melbourne,
Parkville, Victoria 3010, Australia.
ABSTRACT: A multiꢀcomponent hydrogel drug delivery platform with covalent incorporation
of coreꢀcrosslinked star (CCS) polymers has been fabricated through dual crossꢀlinking
reactions. The presence of amphiphilic CCS polymers enhances the mechanical stability of the
network and allows for immobilization of hydrophobic chemotherapeutics (e.g. doxorubicin:
DOX) within a hydrophilic matrix. Thanks to thiolꢀresponsive disulfide segments, the hybrid
network is selectively degradable upon exposure to the glutathioneꢀrich tumor environment,
leading to an onꢀdemand cascade release of payloads in a prolonged manner. After network
degradation, the integrity of the therapeutic cargos is still well protected by the CCS polymeric
carriers. A cellular study demonstrated efficient internalization of released drugꢀcontaining CCS
polymers by HeLa cells, intracellular delivery of encapsulated drugs following pHꢀmediated
erosion of the CCS containers, as well as excellent cytocompatibility of this hydrogel system.
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KEYWORDS: star polymers, hydrogel, onꢀdemand degradation, cascade release, drug delivery
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INTRODUCTION
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Localized chemotherapy proves to be one of the most effective approaches for the treatment of
a number of cancerous diseases.1ꢀ3 After surgical removal of solid tumors, direct delivery of antiꢀ
cancer drugs to the affected site can suppress tumor recurrence or metastasis to a great extent and
improve the overall survival rate of patients.1, 4 Besides, a localized approach has enhanced
therapeutic efficacy and minimized side effects,5 in comparison with a systematic approach
which usually results in insufficient tumor exposure due to rapid plasma clearance of the drugs,
or elicits undesirable offꢀtarget effects with high plasma drug concentrations.6ꢀ7 In light of these
advantages, many local delivery implants have been developed in different configurations and
evaluated for clinical practice.8 These devices are predominately made from biodegradable
polymeric matrix in combination with antiꢀcancer drugs, such as Eligard® (PLGA microspheres
with Leuprorelin (PLGA: polylactideꢀcoꢀglycolide)),9 Gliadel® (polyanhydride wafer with
carmustine) and Oncogel® (PLGA–PEG–PLGA hydrogel with paclitaxel (PEG: polyethylene
glycol)).10ꢀ11 Among them, hydrogelꢀbased local delivery vehicles have attracted considerable
attention because their physicoꢀmechanical similarity to the extracellular matrix of most native
tissues tends to render them a highly biocompatible platform with minimal inflammatory
effects.12ꢀ13 However, traditional hydrogels are plagued by the limited loading capacity of
hydrophobic drugs which play increasingly important roles in cancer therapy,14 owing to their
very hydrophilic nature. To address this limitation, hydrophobic domains have been introduced
into hydrogel networks as depots for waterꢀinsoluble payloads.15ꢀ18 Although these drugs can be
efficiently encapsulated, release of hydrophobic cargos via hydrogel degradation still has issues
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related to drug stability in the blood circulation and low antiꢀtumor activity due to the multiꢀdrug
resistant efflux pumps.19
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In parallel to the development of hydrogels, a variety of nanoꢀsized particles have been
developed as alternative drug delivery platforms, including inorganic particles,20 polymeric
micelles and vesicles,21ꢀ22 liposomes and star polymers, also known as nanogels or microgels.23ꢀ24
The incorporation of nanoparticles into hydrogel networks has been attempted to form hybrid
drug delivery platforms with improved stability, less burst release, as well as increased
hydrophobic encapsulation efficiency, compared with hydrogel platforms alone. For example,
physical embedding of phyllosilicate nanoplatelet fillers in a hydrogel network of dextran
aldehyde and poly(amido amine) dendrimer, enhanced the gel modulus, and facilitated injection
and manipulation in a wide range of implantation contexts.25 Drugꢀloaded liposomes entrapped
within various networks have demonstrated extended release profiles.26ꢀ27 Hydrogels doped with
temperature responsive poly(Nꢀisopropylacylamide) (PNIPAM)ꢀbased nanoparticles have also
been explored, offering an added handle of control: drug release initiated by a specific
temperature change.28ꢀ29 Apart from the physical addition of nanoparticles to a hydrogel matrix,
specific interactions between the particles and the polymer matrix can also be employed to drive
gelation, as evidenced by several shearꢀthinning and selfꢀhealing hydrogels formed via the
interactions between hydrophobically modified natural polymer chains and the hydrophobic
surfaces of different nanoparticles.30ꢀ31 Hybrid hydrogels based on various nanoparticles have
been also fabricated via covalent interactions, for instance, a docetaxel delivery hydrogel crossꢀ
linked by PEGꢀPCL micelles (PCL: poly(εꢀcaprolactone)),32 hydrogels made of hyperbranched
polyester nanoparticles for dexamethasone delivery,33 and a hyaluronic acid hydrogel including
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nanogels for protein delivery.34 These multiꢀcomponent delivery systems have shown potential to
largely improve the drug stability and overcome the drug resistance in cancer therapy.
Despite advances in the development of nanoparticleꢀincorporated hydrogel delivery
platforms, the release mechanism upon cellular demand that can realize high anticancer efficacy
and minimal side effects is still scarcely reported. Thus, in furthering the development of this
field, we report the integration of core crossꢀlinked star (CCS) polymers within an
environmentally responsive degradable hydrogel construct to achieve onꢀdemand release of
loaded hydrophobic antiꢀcancer drugs in a sustained fashion. CCS polymers are star polymers
made via an ‘arm first’ approach and have been widely investigated as unimolecular drug
containers in recent years, thanks to their superior stability, wellꢀdefined structures obtained via
controlled polymerization techniques, high drug loading capacities of their hydrophobic cores
and ease of incorporating multiple functionalities.24, 35ꢀ36 Herein, the CCS polymers play two
important roles, (i) acting as depots for hydrophobic drugs (doxorubicin: DOX) within a
hydrophilic network, and (ii) regulating the intracellular release of the drugs after internalization
by cancer cells. The hydrogels possess high stability under physiological conditions and onꢀ
demand degradation in a simulated tumor environment in vitro, triggering the release of drugꢀ
embedded CCS polymer particles. Finally, efficient uptake of released drugꢀloaded CCS
polymers by cancer cells was also demonstrated.
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EXPERIMENTAL SECTION
Materials αꢀmaleimideꢀωꢀhydroxyl poly(ethylene glycol) (MALꢀPEGꢀOH; Mn=5 × 103 Da;
PDI=1.03) was purchased from Jenkem Technology USA Inc. αꢀamineꢀωꢀthiol poly(ethylene
glycol) (NH2ꢀPEGꢀSH; Mn=5 × 103 Da; PDI=1.03), 4ꢀarm thiolꢀpoly(ethylene glycol) (4ꢀarm
PEGꢀSH; Mn=1.0 × 104 Da; PDI=1.02), α, ωꢀdimaleimide poly(ethylene glycol) (MALꢀPEGꢀ
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MAL; Mn=5.0 × 103 Da; PDI=1.02) were purchased from Jenkem Technology USA Inc. and
used as received. α, ωꢀdiorthopyridyl disulfide poly(ethylene glycol) (OPSSꢀPEGꢀOPSS;
Mn=5.0 × 103 Da; PDI=1.02) was purchased from Creative PEGWorks (U.S.) and used as
received. Methanesulfonic acid (MSA, 98%), Tetrahydrofuran (THF, RCI Labscan, HPLC) was
distilled from sodium benzophenone ketyl before use. Toluene (Scharlau, HPLC),
dichloromethane (DCM, 99.8%, Merck), εꢀcaprolactone (99+%, Aldrich) and diethyl ether
(DEE, 99.8%, Merck) were distilled from calcium hydride (95%, Aldrich) before use. [4, 4'ꢀ
bioxepane]ꢀ7,7'ꢀdione (BOD) was synthesized according to a previously published procedure.37
Doxorubicin hydrochloride was purchased from Sigma Aldrich and deprotonated with
Triethylamine to form water insoluble doxorubicin (DOX). Phosphate buffered saline (PBS) with
pH of 7.4 (PBS 7.4) was prepared by using PBS tablets purchased from SigmaꢀAldrich. PBS
with pH of 5.5 (PBS 5.5) was prepared by mixing KH2PO4 (Merck) and Na2HPO4 (Ajax). 5, 5’ꢀ
Dithiobis(2ꢀnitrobenzoic acid) (DTNB, Ellman’s reagent) and LꢀGlutathione reduced (GSH) were
purchased Sigma Aldrich and used as received. Cell culture supplies (Dulbecco's Modified Eagle
Medium (DMEM), Fetal Bovin Serum (FBS), 100 × GlutaMax, 100 × antibioticꢀantimycotic),
AlamarBlue(R) assay reagent, paraformaldehyde, Cell Mask Deep Red Plasma Membrane Stain
and Alexa Fluor® 488 (AF488) were purchased from Life Technologies and used as received.
DAPI Fluoromount G™ was purchased from ProSciTech and used as received. Culture plates,
microscope slides and glass cover slips were purchased from Corning.
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Instrumentation and Characterization Gel permeation chromatography (GPC) was
performed on a Shimadzu liquid chromatography system fitted with a Wyatt OPTILAB DSP
interferometric refractometer (690 nm), using three Phenomenex Phenogel columns (500, 104,
and 106 Å porosity; 5 ꢁmꢀdiameter bead size) operated at 1 mL/min using THF as the mobile
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phase and with the column temperature set at 45 ºC. All sample solutions were passed through
0.45 ꢁm nylon filters prior to injection into GPC. Aqeous GPC was employed to carry out the
molecular weight characterizations, with a separate Shimadzu liquid chromatography system,
equipped with a Wyatt DAWN EOS multiangle laser light scattering (MALLS) detector (690
nm, 30 mW) and a Shimadzu RIDꢀ10 refractometer (λ = 633 nm), using three Waters
Ultrahydrogel columns in series ((i) 250 Å porosity, 6 ꢁmꢀdiameter bead size; (ii) and (iii) linear,
10 ꢁm diameter bead size), operating at room temperature. The eluent was MilliꢀQ water
containing 20% v/v acetonitrile and 0.1% w/v TFA at a flow rate of 0.5 mL/min. Astra software
(Wyatt Technology Corp.) was used to process the data to determine the molecular weights
either using known dn/dc values. All sample solutions were passed through 0.45 ꢁm nylon filters
prior to injection into GPC. Hydrogen Nuclear Magnetic Resonance (1H NMR) spectroscopic
analysis was performed on a Varian Unity Plus 400 MHz spectrometer using the deuterated
solvent as reference. Dynamic light scattering (DLS) measurements were performed using a
Malvern high performance particle sizer (HPPS) with a 3.0mWHeꢀNe laser operated at 633 nm.
Analysis was performed at an angle of 173º and a constant temperature of 25 ± 0.1 ºC.
Attenuated Total Reflectance Fourier Transform Infrared (ATRꢀFTIR) spectroscopy was carried
out using a Bruker Tensor 27 FTIR, with GladiATR ATR attachment obtained from Pike
Technologies. The FTIR was equipped with OPUS 6.5 spectroscopy software from Bruker Optik
GmBH. XꢀRay Diffraction (XRD) spectroscopy was carried out on a Bruker D8 Advance
Diffractometer, using standard Niꢀfiltered Cu Kα radiation. Rheological study was performed on
controlled stress rheometer using 40 mm parallel plate geometry (ARꢀG2, TA Instruments) with
a 52 ꢁm gap. The linear viscoelastic region of the hydrogels was determined by dynamic strain
sweeps performed from 0.1% to 100% strain at a frequency of 10 rad s−1 at 37 °C. Mechanical
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properties of the hydrogels were examined by dynamic frequency sweep at 37 °C and 2% strain,
which is within the linear viscoelastic region, as was previously determined by strain sweep
experiments. To note, as the hydrogels obtained were relatively soft (i.e., storage modulus <
4500 Pa), compression and tension measurements were not performed. In situ gelation
experiments were conducted at 37 °C with the same plate geometry via time sweep. UVꢀVis
spectrometric analysis was performed using a Shimadzu UVꢀ2101 PC UVꢀVis scanning
spectrophotometer at a scan rate of 1 nm/s, with paired quartz cuvettes (Starna Pty Ltd). Flow
cytometry characterization was conducted on an Apogee Aꢀ50 Micro Flow Cytometry.
Fluorescence intensity histograms of fluorescently labeled polymers and cells were acquired
using an excitation wavelength of 488 nm. Confocal microscope was performed on a Leica TCS
SP2 confocal microscope.
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Synthesis of CCS polymer (MAL-PEG-PCL-BOD) particles The synthesis of maleimide
functionalized PEGꢀPCL based polymer was performed according to a previous report.38 Briefly,
MALꢀPEG5kꢀOH (0.045mmol, 1 equiv.) was azeotropically dried in vacuo with toluene followed
by back purging with Ar three times prior to use. Dry DCM (2.5 mL), CL (1.125 mmol, 25
equiv.) and MSA (0.045 mmol, 1 equiv.) were subsequently added via syringe under Ar
atmosphere. The reaction solution was stirred at room temperature for 1 hour and then a 100ꢁL
1
aliquot of the reaction solution was taken for H NMR analysis. Afterwards, BOD (0.675 mmol,
15 equiv.) dissolved in dry DCM (5 mL) and MSA (0.09 mmol, 2 equiv.) were added to the
reaction solution via syringe under Ar atmosphere. The reaction mixture was then stirred for 24
1
hours at room temperature. A 100 ꢁL aliquot of the reaction solution was taken for H NMR and
GPC analysis respectively. The reaction solution was then precipitated into diethyl ether (50 mL)
1
twice to afford white solid powders, 0.46g (89.3%). H NMR (400 MHz, CDCl3): δH 1.31ꢀ1.40
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(m, ꢀCH2CH2CH2ꢀ), 1.55ꢀ1.69 (m, ꢀCH2CH2CH2ꢀ), 2.23ꢀ2.33 (m, ꢀCOCH2CH2ꢀ), 3.56ꢀ3.65 (m, ꢀ
CH2CH2Oꢀ, PEG repeat unit), 3.98ꢀ4.07 (m, ꢀCH2CH2Oꢀ), 6.68 (s, maleimide end group) ppm.
Mn = 201 kDa, PDI = 1.23. Numberꢀaverage value of arms per star polymer was calculated
according to ref 37.
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Preparation of DOX-loaded CCS polymer particles (DOX-CCS) CCS polymer (30 mg)
and DOX (5 mg) were dissolved in THF (2.5 mL). The THF solution was added dropwise to
water (5 mL) with stirring and the mixture was stirred in a fume hood overnight, allowing THF
evaporation. The trace amount of residual THF was further removed via rotatory evaporation
under room temperature. The precipitate of unloaded DOX was collected by centrifugation at
1500 ppm for 10 minutes and the amount was determined by measuring the absorbance around
480 nm using standard calibration curve in order to calculate the drug loading content (DLC) of
CCS polymer. The supernatant solution was dialyzed in a dialysis tube (MWCO = 3.5 kDa),
frozen and lyophilized to yield solid DOXꢀCCS particles for further analysis and use.
Fluorescent labelling of CCS polymer particles (AF488-CCS) CCS (30 mg) and NH2ꢀ
PEG5kꢀSH (12 mg) were added into water (2 mL) in an Eppendorf tube and the tube was then
placed onto a rotary disk to allow the reaction to proceed for 2 hours. Afterwards, AFF 488 NHS
ester solution (50 ꢁL) from a stock solution (1 mg/mL) was added, the tube was wrapped in
aluminium foil and placed onto the rotary disk to allow the further reaction overnight. The final
solution was dialyzed in a dialysis tube (MWCO = 12 kDa) against water for 2 days, with the
dialysis water changed every 12 hours. The dialyzed solution was frozen and lyophilized to yield
AF488ꢀCCS particles for further use.
Preparation of CCS-incorporated hydrogels (CCS-hydrogels) The GSHꢀresponsive CCSꢀ
incorporated hydrogels were formed via Michaelꢀtype addition between thiol groups (ꢀSH) of 4ꢀ
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arm PEGꢀSH and maleimide groups (ꢀMAL) of CCS polymer, and pyridyl disulfide reaction
between 4ꢀarm PEGꢀSH and 2ꢀpyridyldithiol groups (ꢀOPSS) of OPSSꢀPEGꢀOPSS. Briefly,
OPSSꢀPEGꢀOPSS was dissolved directly in CCS polymer solution (28.5mg/mL) at room
temperature and the mixture was vortexed for approximately 30 s to completely dissolve the
solid polymers. In parallel, PBS solution of 4ꢀarm PEGꢀSH in the same volume was also
prepared. Two solutions were mixed and purged with nitrogen then incubated at 37 °C overnight
to achieve complete crossꢀlinking. The molar ratios of reactive groups, −SH to the combination
of –MAL and –OPSS, varied from 0.8:1, 1:1, 2:1, 3:1 and 4:1, which was controlled by altering
the amount of OPSSꢀPEGꢀOPSS (38, 34, 22, 15 and 12 w.t.% , respectively) and 4ꢀarm PEGꢀSH
(33, 37, 49, 56 and 59 w.t. %, respectively) in the solid gel formula. The total solid content was
kept around 10 w.t. % for all the hydrogel samples at different crossꢀlinking ratios. The CCS
amount in the solid gel of different crossꢀlinking ratios was also kept same, at 28.5%, so that the
DLC of the dry gels could reach 3.9% (given the DLC of CCS had been calculated as 13.5%), at
the same level of a representative local delivery vehicle (Gliadel® wafer).10 The hydrogels with
incorporation of AF488ꢀCCS and DOXꢀCCS was prepared in the same way. Swelling degree
was determined as the ratio of the fully swollen divided by the dry sample mass and illustrated in
Figure S1.
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Degradation study of CCS-hydrogels The CCSꢀhydrogels were prepared as described above.
After being placed in PBS 7.4 for 1 day to reach equilibrium swelling state, the hydrogel samples
were immersed in 2mL PBS 7.4 with 0, 10 ꢁM or 10 mM GSH at 37 °C. At predetermined time
points, the mass of hydrogels after incubation was measured after the excess water on the sample
surface was blotted off. In the meantime, another set of hydrogel samples immersed in the
solutions with 10 ꢁM or 10 mM GSH were collected for mechanical testing at predetermined
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time points. At the end of the test, the solution of degradation products in 10 mM GSH solution
was collected, dialyzed in a dialysis tube (MWCO = 3.5 kDa) against water for 2 days , frozen
and lyophilized for further analysis.
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The AF488ꢀCCS incorporated hydrogels (AF488ꢀCCSꢀhydrogels) were immersed in 10 mM
GSH PBS 7.4. Over time same portion of solution was collected, then dialyzed in a dialysis tube
(MWCO = 12 kDa), and measured on flow cytometry to monitor the release of fluorescently
labelled CCS particles.
In vitro drug release study of DOX-CCS-incorporated hydrogels (DOX-CCS-hydrogels)
DOXꢀCCSꢀhydrogels were prepared as described above. Two parallel release tests were
conducted. In one test, DOXꢀCCSꢀhydrogels were directly immersed in the PBS7.4 solutions of
0, 10 ꢁM or 10 mM GSH at 37 °C. At regular intervals, 0.5 mL of solution was withdrawn for
UVꢀVis analysis and replenished. The release of DOX (or DOXꢀCCS) was determined by
measuring the absorbance of removed solution and using the standard calibration curve. In the
other test, DOXꢀCCSꢀhydrogels were immersed in the dialysis tubes (MWCO = 3.5 kDa)
containing PBS with different GSH content (0, 10 ꢁM and10 mM) and pH level (7.4 and 5.5),
and dialyzed against the corresponding PBS. At certain time intervals, 2 mL aliquots of PBS
outside the dialysis tubes were sampled for UVꢀVis analysis and replaced with the same volume
of fresh solutions.
Cytotoxicity of CCS-hydrogels and DOX-CCS-hydrogels 3T3 and HeLa cells were
maintained in cell culture medium in a humidified atmosphere containing 5% CO2 at 37 ̊C. Cells
were seeded in a T175 flask (ca. 3 × 106 cells/mL) and passaged twice a week prior to the
performance of the subsequent cell viability study. To assess the cytotoxicity of the hydrogels,
dehydrated samples (~15 mg) were placed in an 80 v/v % ethanol solution for 1 hour, then
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to the study.
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The cells were cultured into a 24ꢀwell plate at a concentration of 2000 cells/well except for the
‘medium blanks’ in which the same amount of medium was added instead. Experimental wells
received the sterilized CCSꢀhydrogels (or DOXꢀCCSꢀhydrogel) and the plate was subsequently
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incubated in a humidified atmosphere containing 5 % CO2 at 37 ̊C. After certain period of time,
100 ꢁL of alamarBlue® cell viability reagent (10% volume of cell culture media) was added to
each well (except for three wells containing medium only). After 4 hours of incubation under the
same growth conditions, the absorbance at 570 and 600 nm of each well was measured using a
Varian Cary 50 Bio UVꢀVisible spectrophotometer. The absorbance of each well was corrected
against the mediumꢀonly wells without alamarBlue® reagent and then expressed as a percentage
of the growth control.
Measurement of GSH in cell culture solutions 50 ꢁL of the cell culture solution was
sampled from the wells with confluency of 3T3 and HeLa cells respectively. The solution was
added to 150 ꢁL 5mM DTNB PBS 7.4 and mixed.
A blank sample was prepared by adding 50ꢁl fresh cell culture medium to 150 ꢁL 5mM DTNB
PBS 7.4. The absorbance at 412nm of the test sample was measured on a microplate reader
against blank. The GSH concentration was calculated according to the standard calibration curve.
Cellular uptake study of released CCS and DOX-CCS from degraded hydrogels Confocal
observation: Sterile round glass cover slips (Φ 20mm) were placed in the wells of a 6ꢀwell plate
and 3T3 or HeLa cells were seeded in culture medium at a cell density of 200,000 cells per well
in 0.5mL seeding volume before the plate was placed in a humidified incubator (95 – 100%
humidity, 5% CO2) overnight. Next day, the seeding medium was removed from each well and
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cells were gently washed with sterile PBS. Fresh culture medium was then added together with
AF488ꢀCCSꢀgels (~15 mg) in the wells containing 3T3 cells, and DOXꢀCCSꢀgels (~15 mg) in
the wells with HeLa cells. The HeLa cells cultured in the medium containing free DOX (~ 50
ꢁg/mL) were used as a comparison with those cultured in presence of DOXꢀCCSꢀgels.
For 3T3 cells, the plate was returned to the incubator for 24 hours. Afterwards, the medium was
removed from each well and cells were gently washed with PBS. Samples were fixed using
paraformaldehyde before being stained with DAPI and Cell Mask Deep Red Plasma Membrane
Stain™ and subsequently mounted onto microscopy slides for confocal microscope observation.
For HeLa cells, the plate was returned to the incubator for 6 hours. Afterwards, the medium was
removed from each well and cells were gently washed with PBS. Samples were fixed using
paraformaldehyde before being stained with DAPI and subsequently mounted onto microscopy
slides for confocal microscope observation.
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Flow cytometry measurement: 3T3 cells were seeded in culture medium at a cell density of
200,000 cells per well in 0.5mL seeding volume before the plate was placed in a humidified
incubator (95 – 100% humidity, 5% CO2) overnight. Next day, the seeding medium was
removed from each well and cells were gently washed with sterile PBS. Fresh culture medium
was then added together with AF488ꢀCCSꢀhydrogels in the wells. After further incubation for 24
hours, the medium was removed from each well and cells were gently washed with PBS,
trypsinized, fixed with paraformaldehyde and reꢀsuspended into PBS. Cell suspensions were
subsequently analyzed using flow cytometer.
Data analysis The standard deviation was calculated from three samples in each study,
presented as error bars in the plots.
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RESULTS AND DISCUSSION
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Preparation and Characterization of CCS-hydrogel Drug Delivery Platform The multiꢀ
phase drug delivery hydrogel platform under investigation was prepared in two stages, (i)
synthesis of the drug carriers (i.e., CCS polymers), and (ii) covalent incorporation of the CCS
polymers into a stimuliꢀresponsive hydrogel network consisting of 4ꢀarm PEGꢀSH and OPSSꢀ
PEGꢀOPSS (Scheme 1a). Briefly, the CCS polymer was synthesized via twoꢀstep ringꢀopening
polymerization (ROP) of εꢀcaprolactone, using maleimide (MAL) functionalized PEG as an
initiator and BOD as a crossꢀlinker.38 The hydrophobic linear PCL and crossꢀlinked poly(BOD)
segments comprise their inner core enabling the immobilization of hydrophobic drugs, while the
hydrophilic PEG blocks create a hydrophilic corona allowing for the carrier stabilization in
aqueous environments. Moreover, the particle size of this CCS polymer (DH ~ 25 nm) is in the
range of optimal nanoꢀsized delivery agents (DH: 10~200 nm), featuring a slow elimination from
the circulation and deep penetration into tumor tissues due to the enhanced permeability and
retention (EPR) effect.39 The important features of this CCS polymer are provided in Figure 1.
Notably, the MAL groups in the periphery of the CCS polymers were further utilized to cross
link with 4ꢀarm PEGꢀSH via Michaelꢀtype addition. As a result, the CCS polymers serve as
structural elements as well as hydrophobic drug loading sites within the gel network. The
remainder of the network was formed by the crosslinking between 4ꢀarm PEGꢀSH and OPSSꢀ
PEGꢀOPSS through disulfide exchange. The resulting disulfide bonds are cleavable via SHꢀ
exchange reactions in a reducing environment in the presence of glutathione (GSH).40 We
postulated that the presence of disulfide bonds would provide the network with degradability in
tumor microenvironments with an elevated concentration of GSH.41 The degradation of the bulk
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hydrogels leads to release of the CCS polymer particles along with encapsulated drug molecules,
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followed by uptake by cancer cells (Scheme 1b).
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Figure 1. Characterization of MALꢀPEGꢀPCLꢀBOD CCS polymers (a) H NMR; (b) THF GPC
progression from initiator MALꢀPEGꢀOH, to the arm intermediate (MALꢀPEGꢀPCL) then finally
to the CCS polymer ; (c) DLS measured in water.
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Scheme 1. (a) Synthesis of CCSꢀincorporated hydrogel (CCSꢀhydrogel) via a dual crossꢀlinking
mechanism; Michael addition between 4ꢀarm PEGꢀSH and the maleimide arms of the CCS
polymer (red box), and disulfide exchange between OPSSꢀPEGꢀOPSS and 4 arm PEGꢀSH (black
box); (b) Formation of DOXꢀloaded CCSꢀhydrogel (DOXꢀCCSꢀhydrogel) and subsequent
localized delivery of embedded therapeutic agents (DOX) via onꢀdemand network degradation in
GSHꢀrich tumor microenvironment.
The structure of asꢀformed hydrogels was characterized by ATRꢀFTIR and XRD spectroscopy.
In the ATRꢀFTIR spectra (Figure 2a), bands around 1100 and 550 cmꢀ1 are observed, attributed
to the CꢀOꢀC vibration stretch of a PEG backbone, and the stretch band of disulfide bond,
respectively. With the incorporation of the CCS polymers in the network (CCSꢀgels), a peak
appears at 1725 cmꢀ1. This corresponds to the C=O stretch of the PCL ester carbonyl group,
revealing the existence of the CCS polymers. The relative intensity of this PCL characteristic
peak to PEG increases as the CCS content in the gel formula is increased. Upon loading the CSSꢀ
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gel with DOX (DOXꢀCCSꢀgel), a broad band near 3500 cmꢀ1 is observed as a characteristic peak
of DOX, signaling the presence of hydrophobic drug in the gel matrix. XRD patterns (Figure
S2) show two diffraction peaks at 2θ = 19º and 23º for all the hydrogels. These two prominent
peaks are assigned to the lattice planes (203 and 015) of the PEG crystalline structure. The
absence of the characteristic PCL (2θ = 21.4º, 23.6º) or DOX peaks reflects the amorphous form
of CCS hydrophobic moieties and DOX, suggesting they are evenly distributed throughout the
PEGꢀbased hydrogel matrix,42 in line with the findings for a nanocomposite hydrogel with a
dispersion of GO nanoꢀsheets.43
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Figure 2. (a) ATRꢀFTIR spectra of control PEG gel made from 4ꢀarm PEGꢀSH and OPSSꢀPEGꢀ
OPSS without CCS, CCSꢀhydrogels with different amount of CCS in solid gel formula (14.3 and
28.5%), and DOXꢀCCSꢀgel; (b) Inꢀsitu gelation process of control PEG hydrogel and CCSꢀ
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hydrogel with the same crossꢀlinking ratio (1:1) at 37
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sweep; (c) Summary of storage moduli (G’) of CCSꢀhydrogels fabricated at different crossꢀ
linking ratios (n=3), statistically significant differences with *p < 0.0001, **p < 0.00001 or ***p
< 0.000001 vs. modulus of the CCSꢀhydrogel made at the crossꢀlinking ratio of 0.8:1.
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The gelation process and storage modulus (G’) of the CCSꢀhydrogels were examined via
rheology study. An oscillatory time sweep was performed to quantitatively monitor the inꢀsitu
gelation process of the hydrogels with or without CCS polymer particles (Figure 2b). Rapid
gelation at body temperature was observed for these two hydrogels, occurring within seconds
and becoming stable within minutes (when G’ reaches plateau). This result implies that the
gelation is mainly driven by MALꢀSH addition and disulfide exchange reaction, instead of the
disulfide formation between 4ꢀarm PEGꢀSH itself, because the reaction kinetics of SH oxidation
is significantly slower than those of Michaelꢀtype and OPSSꢀSH coupling reactions.44ꢀ45 The
slower gelation rate of the CCSꢀhydrogel in comparison with the plain PEG hydrogel (control)
could be potentially ascribed to steric hindrance or limited accessibility of the MAL groups on
the CCS polymer surface. Although the experimental studies suggest the reaction kinetics of
thiol exchange between –SH and –OPSS groups appear a bit faster than that of Michael addition,
as the gelation of PEG control gel formed purely via –SH and –OPSS crosslinking took slightly
less time to reach equilibrium than the CCSꢀhydrogel with incorporation of maleimide
functionalized CCS polymer, it is still challenging to discretely determine the crossꢀlinking point
through measurement. In addition, it is noteworthy that the modulus of the CCSꢀgel was higher
than the pure PEG hydrogel at the same crosslinking feed ratio. This suggests the CCS polymers
are covalently incorporated into the network, acting as additional crosslinking sites, analogous to
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the other studies.46ꢀ47 Featuring fast gelation at physiological temperature, mild preparation
conditions and devoid of exogenous initiators, this CCSꢀhydrogel platform holds great promise
for injection applications. Oscillatory frequency sweep shear modulus measurement was
conducted to assess the mechanical properties of the hydrogel samples. The frequency
independence and dominance of G’ over the entire frequency range (Figure S3) indicate the
formation of a stably percolated network in all cases. The impact of the crossꢀlinking degree on
the mechanical properties of the resulting hydrogels was assessed by varying the ratio of SH
groups to the total of MAL (CCS polymer) and OPSS groups. The crossꢀlinking degree was
increased from 0.8:1, indicating that 80% of both MAL and OPSS groups were consumed by SH,
to 4:1, where 4ꢀarm PEG SH was added in excess. It should be noted that the amount of CCS
polymers (2.85 w.t. %) and solid gel content (10 w.t. %) were kept the same at different crossꢀ
linking degrees for a fair comparison. Figure 2c displays the G’ of the CCSꢀhydrogels differing
in crossꢀlinking degree. As the crossꢀlinking degree increased from 0.8:1 to 4:1, G’ rose from
995 to 1922 Pa, then reached maximum (4042 Pa) at 2:1, and dropped thereafter. As would be
expected, when the SH: MAL ratio was increased to 2:1, the likelihood of the MAL arms of the
CCS polymers becoming bridged, increased, resulting in an increased G’. However, upon further
increasing the ratio (≥ 3:1), the high concentration of SH radicals results in limited polymer
chain mobility due to rapid inhomogenous gelation. This subsequently would lead to inactive
loops, unreacted chain ends and a higher chance of intramolecular crossꢀlinking.31, 48 Given the
ease with which the amount of SH or OPSS groups (rather than the MAL groups of CCS) in the
system was altered to produce different crossꢀlinking degrees, the mechanical properties of the
CCSꢀhydrogels can be facilely tailored to meet the application needs.
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Degradation Studies The stability of CCSꢀhydrogels was evaluated by mass loss studies
under physiologically relevant buffer conditions with different levels of GSH (0, 10 ꢁM, 10 mM)
analogous to different extracellular environments. GSH, a tripeptide, is highly involved in the
progression and chemoꢀresistance of various types of tumors. An elevated concentration of GSH
has been found in the tumor microenvironment (~ 10 mM),41 compared to a much lower level in
the plasma (< 10 ꢁM).49 The monitored mass loss as a function of time was fitted to pseudo firstꢀ
order kinetics, as presented in Figure 3a, assuming the sensitive disulfide linkages were the
principal source of degradation (neglecting scission of other bonds, e.g. hydrolysis of esters).
The CCSꢀhydrogels incubated in the PBS appeared stable over 11 days. Similarly, only slight
loss of mass or slow degradation was observed for those gels in 10 ꢁM GSH solution during the
same period of time, exhibiting the stability of the hydrogels under stimulated normal
extracellular environments. In contrast, a significant mass reduction was observed under GSHꢀ
abundant conditions (10 mM), where complete degradation occurred after 11 days. Further, a
control MALꢀSH only hydrogel (MALꢀPEGꢀMAL replacing OPSSꢀPEGꢀOPSS) was prepared
for comparison. Incubated in 10 mM GSH solutions, this hydrogel showed far higher stability
than its counterparts made from both MALꢀSH and OPSSꢀSH reactions, indicating the
degradation of CCSꢀhydrogels was largely induced by the GSHꢀmediated disulfide cleavage
rather than the dissociation of the more stable MALꢀSH adduct. From these degradation results it
is expected that the GSHꢀresponsive hydrogel would allow for a significantly faster drug release,
when compared to the nonꢀGSHꢀresponsive hydrogels under GSHꢀrich conditions (cancer cell
environments). Also, it should be noted, nonꢀCCS incorporated hydrogels were not included, as
drug loading would not be possible. This is because the CCS polymers provide hydrophobic
domains where the hydrophobic drug can be entrapped. In addition, a comparable modulus loss
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profile was obtained (Figure 3b) than the mass loss for the same samples under the same
conditions, suggesting the degradation was dominated by the bulk erosion mechanism. It should
be noted, the shape of the hydrogel during degradation was not specific or definable. This is
suggestively attributed to the random distribution of degradable diꢀsulfide linkages and bulk
erosion.
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Figure 3. (a) Mass loss and (b) modulus loss study of the CCSꢀhydrogels incubated in PBS (pH
= 7.4) with various concentrations of GSH (0, 10 ꢁM or 10 mM), and the hydrogels only with
MALꢀSH adduct (prepared from CCS, MALꢀPEGꢀMAL and 4ꢀarm PEGꢀSH) in 10 mM GSH
PBS (n=3).
Furthermore, the products of CCSꢀhydrogels degraded at a high concentration of GSH were
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analyzed by a series of characterization tools. H NMR spectroscopy (Figure S4) of the
degradation products (collected via dialysis to remove GSH or other impurities) revealed the
presence of intact CCS polymers after degradation, as peaks referring to PCL and PEG segments
could be seen. The absence of maleimide groups indicates the occurrence of Michael additions in
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the process of gelation and the stability of MALꢀSH adduct after degradation. Moreover, a clear
signal of SH verifies the involvement of GSHꢀinduced SHꢀexchange reaction in the degradation
process. Likewise, dynamic light scattering (DLS) was also conducted to characterize the
dialyzed supernatant. The released CCS polymer particles exhibited an average diameter of 47.5
nm, double the size of the original CCS polymer, presumably owing to the attachment of 4ꢀarm
PEG moieties to the periphery (Figure 4a), rather than the formation of CCS polymer
aggregates. To further demonstrate the release of CCS polymers from the degraded network, we
designed a hydrogel fabricated from fluorescently labeled CCS polymers (AF488ꢀCCSꢀ
hydrogels) and employed flow cytometry to monitor their release (Figure 4b & c). Over a period
of 11 days, the first few days (up to day 7) saw a gradual increase in the number of released
particles with similar fluorescent intensity and a sharp surge was witnessed in the following days
(day 9~11). The CCS release rate was slower than the hydrogel degradation rate, presumably
because the diffusion distance of the CCS polymer particles was still maintained in the process
of bulk erosion. All these results confirm the integrity of the released CCS polymers has been
well preserved after the network degradation.
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Figure 4. Characterizations of released CCS polymer particles from degraded network to
confirm their integrity (a) DLS measurement of released CCS polymer vs. original CCS; (b)
Release profile of AF488ꢀCCS polymers monitored by flow cytometry; (c) Summarized count
number of released AF488ꢀCCS polymers as a function of time.
Drug Release Studies DOX, a commonly used anthracycline drug for treatment of various
tumors, was chosen as a model hydrophobic drug for in vitro release test. The efficacy of DOX
in clinical settings is usually hindered by its poor water solubility and toxic side effects at high
cumulative dosage.7, 50 Although many particulate systems have been explored as DOX carriers,
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the shortcomings of rapid clearance and significant initial burst release still impede their
development.15, 38, 51 To demonstrate the potential of our system for controlled release of DOX,
DOXꢀCCSꢀhydrogels were exposed to the mimicry of normal (10 µM GSH) and tumor (10 mM
GSH) extracellular environments. Only about 20% of the drug was released in 10 ꢁM GSH PBS
within 11 days whereas a much faster release was observed at an elevated concentration (10
mM) of GSH over the same period (Figure 5a), which was consistent with the degradation
profiles. The fitted release curves follow the empirical Alfrey model,52 denoting the release
process relied on not only Fickian diffusion but also implant degradation. The 11ꢀday time point
was selected as the endꢀpoint for all release experiments in the series, as at Day 11 almost
complete hydrogel degradation was observed under the highest [GSH] of 10 mM. In parallel,
another set of release studies was carried out in dialysis tubes, filled with and surrounded by the
GSH PBS at different pH levels (7.4 and 5.5) (Figure S5). In dialysis settings (Figure 5b), a
notable reduction in the amount of released DOX was detected in the 10 mM GSH PBS 7.4
solution outside the dialysis tube, compared to the same solution without use of the dialysis tube,
which implies that a large portion of DOX still resided in the released CCS polymers (ca. 55%).
No such difference was observed in 10 ꢁM GSH PBS 7.4 as the majority of DOXꢀCCS was still
locked in the network. A massive fraction of DOX (nearly 70%) diffused out of the dialysis tube
at the pH of 5.5 (10 mM GSH), probably caused by the decomposition of pHꢀsensitive PCL
segments of the CCS polymers under acidic conditions, as shown in the previous study.38 pH 5.5
was purposefully selected in this release study to mimic the environment of intracellular
organelles, such as lysosomes and endosomes, since the drug is expected to be liberated from the
CCS carrier intracellularly (vide infra). Then UVꢀVis and DLS analysis were implemented to
further validate the hypothesis that the CCS polymers act as sequestration site for DOX
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molecules even after being liberated from the degraded network. The red shift in the UVꢀVis
absorbance of released DOXꢀCCS (versus the spectrum of pure DOX) (Figure 6a) and its
increased size (versus the original DOXꢀCCS particle) (Figure 6b) both signify the DOX
molecules were tightly anchored to the CCS polymer carriers. Altogether, these results
demonstrate the high stability of loaded DOX in our delivery system under normal extracellular
conditions and sustained release specifically upon demand of the tumor environment where GSH
is known to be heavily enriched. More importantly, the drugs are well protected by the CCS
polymers from burst release under physiological conditions (pH 7.4), favoring their long
retention in the plasma, and liberated in an accelerated fashion only at the pH level (5.5)
mimicking endosomal conditions, for the degradation of the CCS polymer containers following
cellular uptake.
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Figure 5. In vitro release profile of the DOXꢀCCSꢀhydrogels (a) directly immersed in 10 ꢁM or
10 mM GSH PBS; (b) in the dialysis tubes filled with the PBS with different concentrations of
GSH and pH levels.
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Figure 6. Characterization of released DOXꢀCCS to confirm the encapsulation of DOX in the
CCS polymers even after the network degradation (a) UVꢀVis absorbance of release DOXꢀCCS
vs. pure DOX; (b) Particle size comparison with original DOXꢀCCS.
Cellular Studies The cytotoxicity of the CCSꢀhydrogels was examined via a standard
alamarBlue® assay by incubating them with 3T3 fibroblast and Hela cells. Both cell lines
cultured with CCSꢀhydrogels displayed equivalent viability with the growth control of cells
alone during 1 week (Figure 7a), revealing the cytocompatibility of CCSꢀhydrogels. In addition,
the degradation products of the hydrogels (i.e., the CCS polymer and PEG derivate) are known to
be nonꢀtoxic to the cells, as tested in our previous report.38 The hydrogels bearing drugs (DOXꢀ
CCSꢀhydrogels) engendered a decline in the viability of 3T3 and HeLa cells, with a steeper
decrease in HeLa cells (only 25% of HeLa cells alive after 7 days vs. 70% of 3T3 cells). This
was attributed to a much higher level of GSH (30 times) in HeLa cell culture solution than in
3T3 cell culture (Figure S6a), as measured by Ellman’s assay (Figure S6b), which accounted
for faster network degradation resulting in more drug molecules being released. After incubation
of HeLa cells with AF488ꢀCCSꢀhydrogels, cellular uptake of released fluorescently labelled CCS
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polymer particles was subsequently probed via flow cytometry and confocal microscopy (Figure
S7). The fluorescence intensity of a substantial population of the cells cultured with AF488ꢀ
CCSꢀhydrogels was dramatically increased by nearly 100ꢀfold compared with that of untreated
ones, indicating the efficient uptake by most cells. The internalization of released AF488ꢀCCS
polymer particles in the cytoplasm was visually confirmed by the confocal image. To further
elucidate the uptake mechanism, the intracellular distribution of released DOXꢀCCS polymer
particles was studied in comparison with free DOX (Figure 7b). After 6ꢀhour incubation of cells
with free DOX, strong red fluorescence was observed in cell nuclei, evincing that free DOX
transported into cells via a passive diffusion mechanism.53 In contrast, for the cells exposed to
the DOXꢀCCSꢀhydrogel, the fluorescence was observed in not only the nuclei but also in the
cytoplasm. This observation suggested a different mechanism in which the carriers loaded with
DOX were entrapped by the cells through an endocytosis pathway, then localized in the
endocytic compartments (endosomes/lysosomes), triggering pHꢀmodulated release of DOX from
the carriers to enter nuclei and eventually kill the cells.54ꢀ55 These results are in good agreement
with the previous reports on nanomedicines and also corroborate the residence of drugs in the
CCS polymeric carriers. Overall, these findings in combination are an indicator of the nonꢀ
toxicity of the hydrogels themselves and their effective drug delivery properties.
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Figure 7. (a) Cytotoxicity study of CCSꢀhydrogels and DOXꢀCCSꢀhydrogels performed on 3T3
fibroblast and HeLa cells over 7 days, measured by a standard AlamarBlue® assay. Statistically
significant differences with *p < 0.05, **p < 0.00001 or ***p < 0.000005 vs. positive control
group of the same cell type at the same time point (n=3); (b) Confocal images of HeLa cells
incubated with free DOX or DOXꢀCCSꢀhydrogels for 6 hours. The cell nuclei were stained with
DAPI (blue) and DOX (or DOXꢀCCS) emits a red fluorescent signal (Scale bars: 20 ꢁm).
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CONCLUSIONS
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We report the design and fabrication of a hybrid, multiꢀcomponent cascade drug delivery
platform through incorporation of nanoꢀsized wellꢀdefined CCS polymer particles within a
stimuliꢀresponsive hydrogel network. The system exhibited rapid gelation and easily tunable
mechanical properties via alteration of crossꢀlinking degree. The hydrophobic segments of the
CCS polymers are exploited for sequestration of hydrophobic therapeutic agents (DOX) in the
hydrophilic matrix. The presence of disulfide linkages in the network enables GSHꢀmediated
degradation in the tumor microenvironment, leading to onꢀdemand liberation of the CCS
polymers encapsulating DOX. Cellular studies confirmed the cytocompatiblity of the drug
delivery system, and efficient uptake of released DOXꢀCCS polymer particles in the cytosolic
space.
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ASSOCIATED CONTENT
Supporting Information
The following information is available on the journal website free of charge.
Swelling degree of control and CCSꢀhydrogel (Figure S1); XRD pattern of the CCSꢀgels with
different amount of CCS and DOXꢀCCSꢀgel (Figure S2); oscillatory frequency sweep of the
1
control hydrogel and CCSꢀhydrogels with different crossꢀlinking degrees (Figure S3); H NMR
spectral comparison between the CCSꢀhydrogel building blocks (OPSSꢀPEGꢀOPSS, 4ꢀarm PEGꢀ
SH and CCS polymer) and degradation products (Figure S4); experimental setup of release study
conducted in dialysis settings (Figure S5); GSH concentration in the culture solutions with
confluency of 3T3 and HeLa cells, measured by standard Ellman’s assay (Figure S6); and
cellular uptake study of release fluorescently labelled CCS particles from CCSꢀhydrogel (Figure
S7).
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