Injectable Coacervate Hydrogel for Delivery of Anticancer Drug

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Biological and Medical Applications of Materials and Interfaces

Injectable coacervate hydrogel for delivery of anticancer drug-loaded nanoparticles in vivo Ashlynn L.Z. Lee, Zhi Xiang Voo, Willy Chin, Robert J Ono, Chuan Yang, Shujun Gao, James L Hedrick, and Yi Yan Yang ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b14319 • Publication Date (Web): 29 Mar 2018 Downloaded from http://pubs.acs.org on March 29, 2018

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ACS Applied Materials & Interfaces

Injectable coacervate hydrogel for delivery of anticancer drug-loaded nanoparticles in vivo Ashlynn L. Z. Lee,‡,§ Zhi Xiang Voo,‡, § Willy Chin,‡ Robert J. Ono,† Chuan Yang, ‡ Shujun Gao,‡ James L. Hedrick†,*, and Yi Yan Yang‡,* ‡

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

†

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

These authors contributed equally to the study.

Keywords: Injectable coacervate hydrogels; Nanoparticles/hydrogel composite; Delivery of anticancer drug-loaded nanoparticles; pH-sensitive; In vivo anticancer efficacy

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Abstract In this study, Bortezomib (BTZ, a cytotoxic water-insoluble anticancer drug) was encapsulated in micellar nanoparticles having a catechol-functionalized polycarbonate core through a pHsensitive covalent bond between phenylboronic acid in BTZ and catechol, and these drug-loaded micelles were incorporated into hydrogels to form micelle/hydrogel composite. A series of injectable, biodegradable hydrogels with readily tunable mechanical properties were formed and optimized for sustained delivery of the BTZ-loaded micelles through ionic coacervation between phenylboronic acid-functionalized polycarbonate/poly(ethylene glycol) (PEG) ‘ABA’ triblock copolymer and a cationic one having guanidinium- or thiouronium-functionalized polycarbonate as ‘A’ block. In vitro release study showed pH dependence in BTZ release. At pH 7.4, BTZ release from the micelle/hydrogel composite remained low at 7%, while in an acidic environment, ~ 85% of BTZ was released gradually over 9 days . In vivo studies performed in a multiple myeloma MM.1S xenograft mouse model showed that the tumor progression of mice treated with BTZ-loaded micelle solution was similar to that of the control group, while those treated with the BTZ-loaded micelle/hydrogel composite resulted in significant delay in tumor progression. The results show that this hydrogel has great potential for use in subcutaneous and sustained delivery of drug-loaded micelles with superior therapeutic efficacy.

2


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1. Introduction Bortezomib is a novel proteasome inhibitor used in clinic to treat relapsed multiple myeloma and mantle cell lymphoma. Bortezomib can be used either alone or in combination to circumvent resistance to other anti-cancer drugs1. Recently, bortezomib has been investigated for the treatment of other cancer types2-4. Like most anti-cancer drugs, bortezomib causes a number of adverse side effects, specifically peripheral neuropathy when it is administered by standard intravenous bolus or subcutaneous injection5. A phase I study comparing the pharmacokinetics and pharmacodynamics of the two administration methods showed that subcutaneous administration of bortezomib resulted in lower mean maximum plasma concentration of the drug , and also took longer time to reach (2 h vs. 3 min)6-9. However, overall, the plasma concentration of bortezomib was similar between the two administration routes.

To overcome the clinical limitations of anticancer drugs such as bortezomib, which range from low aqueous solubility (0.6 mg/mL)10,

unstable properties to off-target side effects,

nanoparticles have been used to entrap and deliver drugs to tumor sites through leaky blood vessels in tumor tissues (i.e. enhanced permeability and retention (EPR) effect)6-9. With the aim to enhance therapeutic efficacy and patients’ compliance to adhere to the treatment regimen, we investigate the delivery of bortezomib by localized administration within a drug-loaded nanoparticle/hydrogel composite matrix that serves as both a drug reservoir and diffusion barrier. This may allow for sustained release of drug-loaded nanoparticles and prolonged blood circulation,

hence

improving

anti-tumor

efficacy

and

lowering

side

effects.

Nanoparticle/hydrogel composites, which were constructed from doxorubicin (DOX)-loaded

3



poly(l-glutamic

acid)-b-poly(propylene

oxide)-b-poly(l-glutamic

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acid)

(GPG)

micellar

nanoparticles and poly(vinyl alcohol) (PVA) or chitosan/PVA hydrogel, were reported to deliver DOX in a sustained manner11. One of the disadvantages of this micelle/hydrogel composite is that the hydrogel precursor polymer PVA is not biodegradable as it might not be able to clear from the body. In addition, no in vivo efficacy study was performed.

Another type of hydrogels, which have been used in the biomedical field, involve the use of phenylboronic acid (PBA). PBA exists in equilibrium between the neutral and ionized boronic acid when pH is approximately similar to its pKa. This mixture of uncharged and tetrahedral anionic forms intensifies the intermolecular hydrogen bonding, thus driving the gelation of nonbiodegradable poly(vinylphenyl boronic acid)12. The most common type of boronic acidcontaining hydrogels was made from non-biodegradable polyacrylamides that were synthesized by free radical polymerization of PBA-functionalized monomers using a crosslinker. These polymers were typically used for sugar sensing13-18 and insulin release19-23.

In this study, biodegradable ‘ABA’ triblock copolymers of poly(ethylene glycol) (PEG) (‘B’ block) and PBA-functionalized polycarbonate (‘A’ block) were synthesized through organocatalytic ring-opening polymerization (OROP) using PEG with double ended diols as macroinitiator, which were used to form hydrogels with PBA groups as crosslinker for sustained delivery of the bortezomib-loaded micelles. PBA-functionalized polycarbonate was chosen as it is biodegradable, and the degraded products are non-toxic24-25. Similarly, a biodegradable amphiphilic diblock copolymer of PEG and catechol-functionalized polycarbonate was synthesized through OROP of PBA-functionalized cyclic carbonate with MPEG with one diol 4


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serving as the macroinitiator. Bortezomib was encapsulated into micelles using the polymer through pH-sensitive boronate ester bond formed between boronic acid in bortezomib and catechol. Boronate ester bond is stable when the micelles circulate in the blood, while it dissociates in the endolysosomes (pH 4.5-6.0) following endocytosis26. This pH-sensitivity may prevent premature drug release in the blood stream while increasing bioavailability inside cancer cells. To modulate mechanical strength of hydrogels for injectability, PBA-functionalized triblock copolymers were mixed with oppositely charged ‘ABA’ triblock copolymers containing cationic A-block (e.g., guanidine- or thiouronium-functionalized polycarbonate) through coacervation27. Coacervates are known to have lower moduli. Ionic coacervation allows for the formation of moduli and rheological variant hydrogels by simple mixing at room temperature using low polymer concentrations28-29. In addition, network defects or “dangling chain-ends” were deliberately introduced to improve processability by mixing of ‘ABA’ triblock copolymers with ‘AB’ diblock copolymers of phenylboronic acid moieties. The bortezomib-loaded micelles were incorporated into the hydrogel network to create a drug reservoir for localized and sustained delivery of the drug-loaded micelles. Human multiple myeloma MM.1S xenografts in mice were used as an in vivo model to investigate anticancer efficacy of the bortezomib-loaded micelles/hydrogel composites.

2. Materials and methods 2.1 . Materials PEG with two diol groups (HO-PEG-OH) and methyl-PEG with one diol group (MPEG-OH) (Mn 10,000 Da, PDI 1.05 and 1.10, respectively) were purchased from Polymer SourceTM, 5



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lyophilized, and placed in a glove-box before usage. 1,8-Diazabicyclo[5,4,0]undec-7-ene (DBU) was stirred together with CaH2 over 18 h, and dried after vacuum distillation. Thiourea (TU) was synthesized using a previously published protocol30. Bortezomib was bought from LC Laboratories (U.S.A.). All other chemicals were obtained from Sigma-Aldrich. Protected catechol-functionalized cyclic carbonate (MTC-ProtCat) monomer was synthesized according to our published protocols31, and used to synthesize diblock copolymer of catechol-functionalized polycarbonate and PEG [PEG-P(Cat)13] as described previously.32 Human multiple myeloma MM.1S and human dermal fibroblasts (HDFs) were cultured in RPMI1640 and DMEM, respectively. Culture medium contains 10% fetal calf serum, 100U/ml penicillin and 100 µg/ml streptomycin (HyClone, U.S.A.). 2.2. Synthesis 2.2.1. Synthesis of phenylboronic acid-containing triblock and diblock copolymers (Figure 1A) Protected phenylboronic acid-functionalized cyclic carbonate (MTC-ProtBor) monomer was synthesized as described previously.33-34 Details of the metal-free OROP for synthesis of phenylboronic acid-functionalized polycarbonate and PEG ‘ABA’ triblock copolymer [P(Bor)5PEG-P(Bor)5] are given as an example. The synthesis was performed in a glove box. The initiator OH-PEG-OH (0.556 g, i.e. 0.056 mmol; 10 kDa) and 0.376 g (1 mmol) of MTC-ProtBor were placed in a 5 mL glass vial equipped with a stir bar, followed by addition of dichloromethane to ensure all solids were dissolved. The concentration of monomer was calibrated to 2 M, followed by addition of 8.3 µL (0.06 mmol) of DBU to begin polymerization. After stirring at room temperature for 3.5 h, the reaction was quenched with 30 mg of benzoic acid. Subsequently, the polymer intermediate P(ProtBor)5-PEG-P(ProtBor)5 was precipitated 6


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twice in diethyl ether at room temperature for purification, and dried overnight until a constant weight was obtained. 1H NMR (400 MHz, CDCl3, 22 ºC): δ 7.81-7.28 (m, 40H, -C6H4B=), 5.195.11 (m, 20H, - CH2C6H4-), 4.45-4.20 (m, 40H, -COOCH2-), 3.84.-3.43 (m, 909H, -OCH2CH2from 10kDa PEG), 1.35-1.31 (m, 120H, -OC(CH3)4CO-), 1.26-1.16 (m, 30H, -CH3). Subsequently, the protected copolymer was dissolved in equal volume of methanol and THF (7 ml each) containing 10 equivalents (with respect to moles of protected phenylboronic pinacol pendant groups) of benzene-1,4-diboronic acid and trifluoroacetic acid (TFA). The vial was subjected to heating at 50 °C overnight with constant stirring. The solvents were evaporated and the sticky mixture was re-dissolved in a mixture of isopropanol and acetonitrile mixture (1:1) for dialysis using 1000 MW cut-off membrane over 2 days against a mixture of isopropanol and acetonitrile (1:1). The solvents were dried under vacuum and the polymer was freeze-dried, giving an off-white product [P(Bor)5-PEG-P(Bor)5]. 1H NMR (400 MHz, DMSOd, 22 ºC): δ 8.08-8.00 (m, 40H, -B(OH)2), 7.80-7.18 (m, 40H, -C6H4B=), 5.19-5.03 (m, 20H, - CH2C6H4-), 4.37-4.04 (m, 40H, -COOCH2-), 3.79.-3.41 (m, 909H, -OCH2CH2- from 10kDa PEG), 1.30-1.03 (m, 30H, -CH3). Diblock copolymer of phenylboronic acid-functionalized polycarbonate and PEG [PEGP(Bor)] was synthesized and worked up in similar fashion in the glove box, with the exception of the initiator. The initiator MPEG-OH (0.22 g, i.e. 0.022 mmol; 10 kDa) and 0.376 g (1 mmol) of MTC-ProtBor were placed in a 5 mL glass vial containing a stir bar. Subsequently, the synthetic protocols were carried out similarly to those for the synthesis of P(Bor)5-PEG-P(Bor)5. 2.2.2. Synthesis of guanidinium-functionalized polycarbonate and PEG triblock copolymers [P(Gu)-PEG-P(Gu)] (Figure 1B) 7



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Using a glove box, HO-PEG-OH was added to a 5 mL glass vial that contained 1 mL of anhydrous DCM, TU (18.5 mg, 0.05 mmol) and DBU (7.47 µL, 0.05 mmol), and the mixture stirred

for

10

min.

Protected

guanidinium-functionalized

cyclic

carbonate

(MTC-

OCH2BnBocGu) (261 mg, 0.5 mmol, see Supplementary Information for the detailed synthetic protocols) was then added to the mixture, and stirred at room temperature for another 30 min. The catalyst was quenched by adding benzoic acid (10 mg, 0.08 mmol). The crude reaction mixture was purified by precipitation in diethyl ether (3 times, 40 mL each) and drying under reduced pressure to afford the desired polymer P(BocGu)-PEG-P(BocGu) that appeared as a white solid (90% yield). 1H-NMR (400 MHz, CDCl3, 22° C): δ 11.54 (bs, 9H, NH), 8.66 (bs, 9H, NH), 7.30 (m, 36H, phenyl -CH), 5.13 (m, 18H, -CH2-), 4.64 (m, 18H, -CH2-), 4.27 (m, 36H, CH2-), 3.65 (m, 909H, PEG -CH2-), 1.53 – 1.43 (m, 173H, Boc -CH3), 1.27 – 1.21 (m, 24H, CH3). The pendant Boc protecting groups in the polymer were removed using TFA. In a 20-mL reaction vial, P(BocGu)-PEG-P(BocGu) (700 mg) was dissolved in 9 mL of DCM and 1 mL of TFA. The mixture was stirred at room temperature for 18 h, followed by removal of solvent under vacuum. A light yellow waxy solid was obtained, which was subsequently dissolved in water and freeze-dried, yielding a white transparent solid [P(Gu)-PEG-P(Gu)]. Full deprotection was confirmed by 1H-NMR analysis. Yield: 87%; 1H-NMR (400 MHz, DMSO, 22° C): δ 8.18 – 7.92 (m, 9H, NH), 7.34 (m, 52H, phenyl –CH and NH), 5.12 (m, 19H, -CH2-), 4.40 – 4.14 (m, 56H, -CH2-), 3.51 (bs, 909H, PEG -CH2-), 1.24 – 1.07 (m, 28H, -CH3). 2.2.3. Synthesis of thiouronium-functionalized polycarbonate and PEG triblock copolymers [P(Th)-PEG-P(Th)] (Figure 1C) 8


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Triblock copolymer [P(BnCl)-PEG-P(BnCl)] was first synthesized as described previously.32 White solid (1.94 g, 95% yield). 1H NMR (400 MHz, CDCl3): δ 7.36-7.28 (br d, 40H, Ar-H), 5.12 (s, 20H, -OCH2-BnCl), 4.55 (s, 20H, -CH2-Cl) 4.26 (br, 40H, -OCOOCH2- and OCH2CCH3-), 3.64 (s, 909H, PEG -OCH2CH2-), 1.23 (s, 30H, -CH3). GPC: Mn=22236, Mw=23592, D=1.06. P(BnCl)-PEG-P(BnCl) (1.94 g, 1.8 mmol BnCl groups), TU (0.418 g, 5.5 mmol), and DMF (5 mL) were added to a 20 mL vial, and the reaction mixture was stirred continuously for 18 h at 40 °C. The crude product was purified by dialysis in 1000 Da molecular weight cut-off membrane against water for 18 h. The desired polymer P(Th)-PEG-P(Th) was obtained as a white solid after lyophilization (1.88 g, 90% yield). 1H NMR (400 MHz, DMSO-d6): 9.22 (br, 40H, -NH and -NH2), 7.44-7.30 (m, 40 H, Ar-H), 5.11 (s, 20H, -O-CH2-Ar), 4.50 (s, 20H, -CH2S-), 4.25 (m, 40H, -OCOOCH2- and -OCH2CCH3-), 3.49 (s, 909H, PEG -OCH2CH2-), 1.18 (s, 30H, -CCH3).

2.3. Measurement of critical micelle concentrations (CMCs) The CMCs of the polymers in de-ionized (DI) water were determined by fluorescence spectroscopy as described previously35.

2.4. Preparation and characterization of BTZ-loaded PEG-P(Cat)13 micelles 9



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Thin film-hydration method was used to prepare BTZ-loaded PEG-P(Cat)13 micellar nanoparticles. Briefly, 5 mg of BTZ and 10 mg of PEG-P(Cat)13 were dissolved in 7.5 mL of methanol via ultrasonication. The mixture was then evaporated under reduced pressure using a rotatory evaporator at 50 °C to remove methanol, and a thin film was formed. HPLC grade water (5 mL) was added at 50 °C and sonicated for 30 min. Subsequently, any insoluble residual drug was removed by 5-min centrifugation at 4000 rpm, 25°C and filtration (0.22 µm nylon syringe filters). The BTZ content was analyzed at 280 nm by high performance liquid chromatography (HPLC, Waters 996 PDA detector, U.S.A.). Specifically, the freeze-dried BTZ-loaded micelles were dissolved in mobile phase (acetonitrile, water and THF in the volume ratio of 65: 30: 5). BTZ loading level was obtained from the weight ratio of drug encapsulated in the micelles to the drug-loaded micelles.

2.5. Formation of BTZ-loaded micelle/hydrogel composite BTZ-loaded micelle/hydrogel composite was prepared by adding equivalent volumes of sodium bicarbonate buffer (600 mM, pH 8.1) and fresh bortezomib-loaded micelle solution to hydrogel precursor polymers P(Bor)5-PEG-P(Bor)5 (7 wt.%) and P(Gu)5-PEG-P(Gu)5 (3 wt.%) and mixed through gentle vortexing.

2.6. Rheological analysis of hydrogels

10


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The rheological analysis of the hydrogels was performed on an ARES-G2 rheometer (TA Instruments, USA) as described in our previous work36 under controlled strain of 2.0% and a frequency scan of 1.0 to 100 rad/s.

2.7. Scanning electron microscopy (SEM) imaging of hydrogels Prior to imaging, the hydrogels were cryo-fixed by placing the gels in a chamber filled with liquid nitrogen so as to minimize any interference to the morphological structure. The samples were then freeze-dried for 24 h. Cross-sectional morphology of the gel was imaged using a scanning electron microscope (SEM) (Jeol JSM-7400F, Japan)37.

2.8. In vitro drug release BTZ release from the BTZ-loaded micelle/hydrogel composite was investigated in a dialysis membrane tube (MWCO of 1000 Da) (Spectrum Laboratories, U.S.A.) with 500 µL of gel. The tubes were submerged in 10 mL of PBS buffer (100 mM, pH 7.4 or 5.8), and placed on an orbital shaker at 100 rpm, 37 oC. At different time intervals, the release medium was taken out and replaced with fresh medium38-39. To analyze BTZ content, samples of the release medium were collected and diluted 10 times using the HPLC mobile phase and the absorbance of the samples was measured using HPLC at 280 nm.

11



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2.9. Cytotoxicity test Human multiple myeloma MM.1S and HDF cells were seeded on a 96 well plate (density: 10 × 103 cells per well) and incubated overnight at 37 °C. The spent medium was then removed and 90 µL of fresh medium was added to each well. Subsequently, 10 µL of the hydrogels or micelles-containing hydrogels were added to each well, and the cells incubated at 37 °C for 48 h. Reagents from the CellTiter 96® AQueous One Solution Cell Proliferation Assay Kit (Promega, USA) and cell culture medium were mixed at 1:4 volume ratio. This mixture (100 µL) was then added to each well and incubated with the cells for 2 h at 37 °C in the dark. The absorbance of each well was recorded at 490 nm. A percentage of the viability of the treated cells was calculated against that of the untreated control cells.

2.10. Animal studies All animal studies were permitted by the Biological Resource Centre of Singapore with approved protocol from the Institutional Animal Care and Use Committee (IACUC). CB-17 severe combined immunodeficient (SCID) mice were inoculated with 5 × 106 MM.1S cells in 200 µL of 1:1 mixture of serum-free RPMI 1640 medium and matrigel by subcutaneous injection.

2.11. In vivo anti-cancer efficacy studies When the tumor volume reached around 300 mm3, the tumor-bearing mice were randomly divided into five groups with 7–10 mice per group, (1) untreated control, (2) blank PEG-P(Cat)13 12


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micelle solution (150 µL), (3) blank hydrogel (150 µL), (4) bortezomib-loaded micelle solution (0.8 mg/kg bortezomib, 150 µL) and (5) bortezomib-loaded micelle/hydrogel composite (0.8 mg/kg bortezomib, 150 µL). The formulations were injected subcutaneously at ~1 cm away from the tumor sites. Treatment was done only once on the first day of treatment (Day 0). Body weight of the mice was monitored, and tumor volume recorded twice a week by caliper measurements. Tumor volume was estimated using the following formula: L×W2 /2, where L and W are the major and minor diameters, respectively. Two-tailed Student’s t test was performed to study statistical differences in the data obtained and P≤0.05 was considered to show a statistical significance.

3. Results and discussion 3.1. Monomer and polymer synthesis The synthesis of pinacol-protected phenylboronic acid-functionalized cyclic carbonate monomer (MTC-ProtBor) was a straightforward reaction that involves the direct coupling of MTC-Cl with 4-hydroxymethyl phenylboronic acid pinacol in the presence of triethylamine. The polymers were synthesized through OROP of MTC-PPB using PEG or MPEG of 10 kDa as initiator together with the co-catalysts DBU and TU (Figure 1a). Proton NMR analysis of the polymers displayed peaks from both initiator and monomers. 1H NMR integration values of monomers relative to PEG correlated well (Figure S1), demonstrating controlled ROP and predictable molecular weights by using specific monomer to initiator feed ratio. Triblock (P(ProtBor)-PEG-P(ProtBor)) and diblock (PEG-P(ProtBor)) copolymers of pinacol-protected 13



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phenylboronic acid-functionalized polycarbonate and PEG had narrow molecular weight distribution (polydispersity index-PDI: 1.15 and 1.10, respectively) (Figure S2). The protecting groups in the polymers were subsequently deprotected to yield P(Bor)-PEG-P(Bor) and PEGP(Bor), and 1H NMR showed the disappearance of the methyl protons from the pinacol protecting groups (Figure S3). In addition, a new distinct peak appeared at 8.05 ppm, confirming the presence of hydroxyl groups, which correlate to the deprotected phenylboronic acid pendant groups. The degree of polymerization (DP) was identified as 10 and 12 for the triblock and diblock copolymers, respectively, based on 1H proton ratio of PEG to pinacol-protected phenylboronic acid-functionalized cyclic carbonate (Table 1). Similarly, PEG-P(Bor)6 was also synthesized.

Boc-protected guanidinium-functionalized cyclic carbonate monomer (MTC-OCH2BnBocGu) was synthesized by appending (4-aminomethyl-phenyl)-methanol onto MTC-OH in the presence of N,N-diisopropylethylamine (DIPEA) (Figure 1b), and the monomer was purified by column chromatography that provided a yield of 91%. P(Gua)-PEG-P(Gua) was synthesized through OROP of MTC-OCH2BnBocGua using PEG 10 kDa as the macroinitiator, which had relatively narrow molecular weight distribution with PDI of 1.32 (Figure S2). The polymer was completely deprotected with TFA based on 1H NMR analysis upon disappearance of peaks between the region of 1.3 to 1.5 ppm (Figure S4, Figure S5). The degree of polymerization (DP) was determined to be 9 based on 1H proton ratio of PEG 10kDa to Boc-protected guanidinefunctionalized cyclic carbonate, attributing to a DP of 4.5 (~5) on each side of the ‘ABA’ triblock copolymer (Table 1). 14


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Thiouronium-functionalized polycarbonate and PEG triblock copolymer (P(Th)-PEG-P(Th)) was synthesized with narrow molecular weight distribution (PDI: 1.06, Figure S2) through OROP of the benzyl chloride-functionalized cyclic carbonate monomer (MTC-OBnCl) using PEG 10 kDa as the macroinitiator, followed by subsequent quaternization of the polymer with thiourea (Figure 1c). 1H NMR analysis revealed no polymer degradation after the polymer was stirred and heat overnight with thiourea, followed by dialysis to achieve the purified polymer. The degree of polymerization (DP) was determined to be 10 based on 1H proton ratio of PEG to benzyl chloride functionalized carbonate.

15



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A

B

C

16


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D

+ BTZ-loaded Micelles

P(Bor)5-PEG-(Bor)5

P(Gu)5-PEG-P(Gu)5

Figure 1. Synthesis schemes of (A) pinacol-protected boronic acid monomer (MTC-ProtBor) and subsequent OROP of MTC-ProtBor to form triblock and diblock copolymers, followed by deprotection to offer P(Bor)-PEG-P(Bor) and PEG-P(Bor); (B) Boc-protected guanidiniumfunctionalized cyclic carbonate monomer (MTC-OCH2BnBocGu) and subsequent OROP of MTC-OCH2BnBocGu to offer triblock copolymer, followed by deprotection to offer P(Gua)PEG-P(Gua). (C) OROP of MTC-OBnCl to offer triblock copolymer, followed by quaternization with thiourea to yield P(Th)-PEG-P(Th). (D) Scheme showing the formation of BTZ-loaded micelle/hydrogel composite.

3.2. Bortezomib-loaded PEG-P(Cat)13 micelles PEG-P(Cat)13 self-assembled into micelles at concentrations above its critical micelle concentration (CMC) (12.0 µg/mL). Bortezomib was effectively loaded into PEG-P(Cat)13 micelles through the formation of the pH-sensitive boronate ester bonds, and the encapsulation efficiency and loading level were 57 ± 6% and 50 ± 2%, respectively. The particle size and polydispersity index (PDI) of the blank and drug-loaded micelles were determined to be (42 nm, 17



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0.26) and (62 nm, 0.11), respectively. The PDI of drug-loaded micelles was much smaller than that of the blank micelles, suggesting formation of a more compact micellar core after drug loading.

3.3. Physical properties of hydrogels The CMC values of the amphiphilic ABA-type triblock copolymers P(Bor)5-PEG-P(Bor)5, P(Th)5-PEG-P(Th)5 and P(Gu)5-PEG-P(Gu)5 were determined to be 14.0, 66.0 and 490 mg/L in water, respectively (Table S1). Boronic acid-functionalized polymers exhibited greater propensity for micelle formation and this is attributed to the strong interactions between the boronic acid groups. The CMC value for P(Th)5-PEG-P(Th)5 is lower than that of P(Gu)5-PEGP(Gu)5 due to the difference in hydrophobicity between the two polymers. The strong hydrogen bonding between boronic acid moieties in P(Bor)5-PEG-P(Bor)5 enabled it to form hydrogel at ≥ 10 wt.%. However, these hydrogels were not injectable. From Table 1, the polymer concentration significantly affected the storage modulus G’ of the gel. For example, a 20% concentration increment (10 to 12 wt.%) of P(Bor)5-PEG-P(Bor)5 resulted in 50% increase in G’ value from 4203 (Table 1, entry 2) to 6608 Pa (Table 1, entry 1). These hydrogels were too stiff to function as an injectable drug matrix through a 22G needle. To reduce the storage modulus, P(Bor)5-PEG-P(Bor)5 was mixed with the boronic acid-functionalized diblock copolymers PEG-P(Bor)6 or PEG-P(Bor)12. The addition of the diblock copolymers significantly reduced the storage modulus (Table 1, entry 4 and 5), likely by the disruption in network formation by the addition of “dangling chain-ends”.

However, these gels were still not

injectable. Mixing P(Bor)5-PEG-P(Bor)5 with the cationic guanidinium- or thiouronium18


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containing triblock copolymers P(Gua)5-PEG-P(Gua)5 and P(Th)5-PEG-P(Th)5 led to great reduction in gel stiffness, conferring injectability to the hydrogels (Table 1, entry 6 to 11).

The addition of P(Gu)5-PEG-P(Gu)5 to P(Bor)5-PEG-P(Bor)5 (Table 1, entry 6, 8 and 10) resulted in considerably lower storage modulus as compared to P(Th)5-PEG-P(Th)5 (Table 1, entry 7, 9 and 11). Reduced stiffness was also observed with increasing P(Gu)5-PEG-P(Gu)5 or P(Th)5-PEG-P(Th)5 content. The mixed gel formation was accomplished in aqueous solution under basic conditions (pH 8.1) where the boronic acid resides ∼50% uncharged, and the tetrahedral anionic form intensifies intermolecular hydrogen bonding as well as forms a coacervate with the cationic guanidinium or thiouronium moieties. The hydrogels with G’ of 1800 Pa and above were not readily injectable. Among the injectable hydrogels, the stiffness of the hydrogel prepared from a mixture of P(Bor)5-PEG-P(Bor)5 (7 wt.%) and P(Gu)5-PEG-P(Gu)5 (3 wt.%) (Table 1, entry 8) was found to be suitable for quick and easy injection, and was selected for subsequent studies. From Figure 2A, a more drastic decrease of viscosity with shear rate increase was observed for this mixture as compared to P(Bor)5-PEG-P(Bor)5 (10 wt.%) without P(Gu)5-PEG-P(Gu)5, which is desirable for injectable formulation.

Although the incorporation of the drug-loaded PEG-(Cat)13 micelles into the optimized hydrogels

increased storage moduli significantly (Table 1, entry 6* and 8*), the

micelle/hydrogel composite formed from P(Bor)5-PEG-P(Bor)5 (7 wt.%) and P(Gu)5-PEGP(Gu)5 (3 wt.%) (Table 1, entry 8*) was injectable. When the composition of P(Bor)5-PEG19



Page 20 of 31

P(Bor)5 was increased beyond 7 wt.% in the mixtures, the hydrogel became more stiff and less easy to operate for injection purpose. To use it as an injectable delivery matrix, it is imperative for the disrupted hydrogel network to recover its viscoelastic property after application of shear force (after injection). To study this property, a dynamic step strain amplitude test (γ = 2 or 100%), which mimics a typical injection procedure, was performed on the drug-loaded PEG(Cat)13 micelles-containing hydrogel formed from 7 wt.% P(Bor)5-PEG-P(Bor)5 and 3 wt.% P(Gu)5-PEG-P(Gu)5 [drug-loaded micelle/(B7+G3) hydrogel composite]. Figure 2B showed that the initial G’ of the hydrogels was ~1400 Pa at a low strain (γ = 2%) at 25°C. The hydrogel exhibited a predominantly solid-like behavior, as indicated by G’ > G’’. When the strain was increased to 100% to simulate the pushing of hydrogel through the syringe barrel, the G’ value was instantly reduced by ~ 7 folds to ~ 200 Pa. With G’ value reaching similar values to G’’, the matrix behaved more liquid-like. After 200 s of continuous stress, the strain was reverted to a low value (γ = 2%) at 37°C to simulate human body temperature. The G’ value immediately increased to 600 Pa and exceeded G” by 5 folds, showing the recovery to the solid-like behavior of the hydrogel40. The recovery of mechanical property is tremendously beneficial for the hydrogel to serve as an injectable delivery matrix.. The viscoelastic property of the hydrogel was observed to be highly dependent on temperature as the G’ and G” returned to initial values (~1400 and 100 Pa respectively) when the temperature was lowered from 37°C to 25°C. Given that the drug-loaded micelles/(B7+G3) hydrogel composite was injectable (Figure S6), and had reversible mechanical property, the composite gel was therefore further studied in vitro and in vivo.

20


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Table 1. Rheological properties of hydrogels in various compositions. Amt. of Sample

Polymer 1

Amt. of

Polymer 1

Polymer 2

(wt.%)

Polymer 2

G' (Pa)

G'' (Pa)

(wt.%)

1

P(Bor)5-PEG-P(Bor)5

12

-

-

6608 ± 139

687 ± 73

2

P(Bor)5-PEG-P(Bor)5

10

-

-

4203 ± 163

629 ± 63

3

P(Bor)5-PEG-P(Bor)5

7

-

-

4

P(Bor)5-PEG-P(Bor)5

10

PEG-P(Bor)6

2

3215 ± 139

558 ± 39

5

P(Bor)5-PEG-P(Bor)5

10

PEG-P(Bor)12

2

2875 ± 119

463 ± 45

6

P(Bor)5-PEG-P(Bor)5

8

P(Gu)5-PEG-P(Gu)5

2

1804 ± 68

187 ± 17

6*

P(Bor)5-PEG-P(Bor)5

8

P(Gu)5-PEG-P(Gu)5

2

2252 ± 78

335 ± 70

7

P(Bor)5-PEG-P(Bor)5

8

P(Th)5-PEG-P(Th)5

2

3672 ± 120

332 ± 29

8

P(Bor)5-PEG-P(Bor)5

7

P(Gu)5-PEG-P(Gu)5

3

695 ± 33

92 ± 7

8*

P(Bor)5-PEG-P(Bor)5

7

P(Gu)5-PEG-P(Gu)5

3

1415 ± 29

108 ± 3

9

P(Bor)5-PEG-P(Bor)5

7

P(Th)5-PEG-P(Th)5

3

2071 ±66

255 ± 11

10

P(Bor)5-PEG-P(Bor)5

6

P(Gu)5-PEG-P(Gu)5

4

54 ± 4

10 ± 5

11

P(Bor)5-PEG-P(Bor)5

6

P(Th)5-PEG-P(Th)5

4

871 ± 39

144 ± 10

Does not gel

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

A

B

Gel Liquid

Gel

Gel

Figure 2. Shear thinning property and phase transition between liquid-like and solid-like of (B7+G3) hydrogel. (A) Flow sweep of hydrogels with different polymer concentrations at 25oC. More drastic decrease of viscosity with increasing shear rate was observed for (B7+G3) hydrogel as compared to P(Bor)5-PEG-P(Bor)5 (10 wt.%) without P(Gu)5-PEG-P(Gu)5, which is desirable 21



Page 22 of 31

for injectable formulation. (B) Dynamic step strain amplitude test (γ = 2 or 100%) for BTZloaded micelle/(B7+G3) hydrogel composite under different strain and temperature conditions. The results demonstrated reversible recovery of the drug-loaded micelles/(B7+G3) hydrogel composite from liquid-like phase to solid-like phase after injection.

3.4. SEM imaging of hydrogels The cross-section morphology of the hydrogels was examined through SEM imaging after freeze-drying (to preserve the morphology of the hydrogels). From Figure 3A, blank (B7+G3) hydrogel exhibited porous structure (pore size: 100 - 300 µm). On the other hand, in Figure 3B, the addition of drug-loaded micelles formed a more compact/dense hydrogel, with lower porosity and smaller pore size (between 2 and 10 µm). Such phenomenon coincides with the observed higher storage modulus value of drug-loaded micelle/(B7+G3) hydrogel composite compared to the blank (B7+G3) hydrogel as shown in Table 1. Nonetheless, the pores are large enough to allow for permeation of the drug-loaded micelles.

Figure 3. SEM images of cryo-fixed (A) blank (B7+G3) hydrogel and (B) BTZ-loaded micelle/(B7+G3) hydrogel composite.

22


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3.5. In vitro drug release BTZ release from the BTZ-loaded micelle/(B7+G3) hydrogel composite was studied in PBS at pH 7.4 (mimicking the extracellular environment) and pH 5.8 (simulating the endolysosomal environment). At pH 7.4, the release of BTZ from the micelle/hydrogel composite remained low at 7% throughout the experiment, while the release was significantly higher at pH 5.8 (Figure 4). In the acidic environment, ~ 85% of the loaded BTZ was released gradually over 9 days. Several studies have reported the pH-sensitive dissociation of boronate esters 26, 41-42, where at acidic pH, the boronate ester bond is hydrolyzed, thereby resulting in the release of the encapsulated drug. Such property confers a good advantage for the micelle/hydrogel composite to function as a drug reservoir for cancer treatment as the release of bortezomib from the micelles is limited at the normal physiological pH, while the dissociation of boronate ester and bortezomib release is facilitated at the endolysosomal pH following endocytosis of the drug-loaded micelles by cancer cells.

Figure 4. Release profiles of bortezomib from bortezomib-loaded micelle/(B7+G3) hydrogel composite under different pH conditions. Each experiment was conducted in triplicates. Each error bar represents the average ± standard deviation of the replicates. 23



Page 24 of 31

3.6. In vitro cytotoxicity of drug-loaded micelle/hydrogel composite In vitro cytotoxicity of various BTZ formulations was evaluated using human multiple myeloma MM.1S cell line as model cells. In free BTZ formulation, BTZ was dissolved in DMSO. In the case of free BTZ-loaded hydrogel, BTZ was dissolved in DMSO and then mixed with the hydrogel precursor for hydrogel formation. Blank (B7+G3) hydrogel and PEG-P(Cat)13 micelles showed almost negligible toxicity on MM.1S cells with more than 90% cell viability. The reduction in cell viability was similar between BTZ-loaded micelles and free BTZ or between BTZ-loaded micelle/hydrogel composite and free BTZ-containing hydrogel, showing that BTZ was released from the micelles and the drug remained functional after the encapsulation procedure. In addition, the blank (B7+G3) hydrogel and blank micelles/(B7+G3) hydrogel were not cytotoxic against normal healthy human HDF cells (Figure S7).

Sample

Component 1

Component 2

1

Gel

-

2

Gel

Blank Micelles

3

Gel

BTZ-loaded Micelles

4

Gel

Free BTZ

5

-

Blank Micelles

6

-

BTZ-loaded Micelles

7

-

Free BTZ

Figure 5. Viability of MM.1S cells after 48-h incubation with different BTZ, micelle and hydrogel formulations. Equivalent BTZ concentration: 0.23 mg/mL.

24


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3.7. In vivo anti-cancer efficacy Biodegradability of the hydrogel was first evaluated. The hydrogel was injected into balb/c mice, and the mice sacrificed at 7th and 14th day post injection. On Day 0, 150 µL (= 150 mm3) of blank micelles/(B7+G3) hydrogel was injected into the subcutaneous region of the mice, and by Day 7 and 14, the volume of hydrogel decreased to 36 mm3 and 15 mm3, respectively (Figure S8). This demonstrates that the hydrogel degraded over more than 14 days. The therapeutic efficacy of bortezomib delivery using different formulations was then investigated by comparing the subcutaneous injection of bortezomib-loaded micelle/(B7+G3) hydrogel composite with bortezomib-loaded micelle solution, blank PEG-P(Cat)13 micelle solution and blank (B7+G3) hydrogel in mice bearing subcutaneous MM.1S tumors. The mice did not display weight loss, reduced appetite or lethargy during the course of treatment, thereby showing good tolerance to all formulations (Figure 6A). The mice were sacrificed on Day 21 from the start of treatment as the tumors in most groups, except those in the group treated with bortezomib-loaded micelle/(B7+G3) hydrogel composite, grew to more than 2000 mm3.

Large differences in tumor growth rates were observed in mice that were treated with the solution and hydrogel formulations. Similar tumor progression was observed for mice that were treated with bortezomib-loaded micelle solution and the control group, demonstrating that the solution formulation was unable to exert therapeutic activity at the given dose. The group that was treated with bortezomib-loaded micelle/(B7+G3) hydrogel composite did not show noticeable tumor growth for the first 8 days post treatment (P > 0.01) and only showed tumor size increase from 11 days post treatment. Nevertheless, treatment with the bortezomib-loaded micelle/(B7+G3) hydrogel resulted in significant impediment to tumor progression (P < 0.001) 25



Page 26 of 31

compared to other treatments (Figure 6B), and tumors that were removed from mice treated with BTZ-loaded micelle/(B7+G3) hydrogel composite were much smaller compared to other treatment groups (Figure 6C).. This was likely due to the diffusion barrier that the hydrogel matrix provided which helped to extend the release of BTZ-loaded micelles into the circulation and tumor site.

A

B

C

Figure 6. In vivo studies of bortezomib-loaded micelle/(B7+G3) hydrogel composite. Changes in (A) body weight and (B) tumor size of MM.1S-tumor bearing mice after treatment using various bortezomib formulations. P

Injectable Coacervate Hydrogel for Delivery of Anticancer Drug

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