Self-Assembled Hydrogels Utilising Polymer-Nanoparticle Interactions

Mouldable hydrogels that flow upon applied stress and rapidly self-heal are increasingly utilised as they afford minimally invasive delivery and conformal application. Here we report a new paradigm for the fabrication of self-assembled hydrogels with shear-thinning and self-healing properties employing rationally engineered polymer-nanoparticle interactions. Biopolymer derivatives are linked together by selective adsorption to nanoparticles. The transient and reversible interactions between biopolymers and nanoparticles enable flow under applied shear stress, followed by rapid self-healing when the stress is relaxed. We develop a physical description of polymer-nanoparticle gel formation that is utilised to design biocompatible gels for minimally-invasive drug delivery. Owing to the hierarchical structure of the gel, both hydrophilic and hydrophobic drugs can be entrapped and delivered with differential release profiles, both in vitro and in vivo. The work introduces a facile and generalizable class of mouldable hydrogels amenable to a range of biomedical and industrial applications.

Values are taken at an angular frequency of 10 rad/s and a strain of 2%. As G = nαk B T , where n is the number of particle-polymer interactions per unit volume and αk B T is the interaction strength, the storage modulus is expected to scale with the number of particle-polymer interactions as well as the strength of the interactions between polymers and particles. Moreover, the number of particle-polymer interactions is expected to scale with the number of particles. Likewise, the interaction strength is expected to scale with the size of the particles. Assuming an interaction strength between polymer and nanoparticle (ε) of approximately 5k B T , we can expect that in the case of a PNP hydrogel composed of 50 nm PSNPs (10 wt%; ∼ 1.2 x 10 18 particles) and HPMC-C 12 (1 wt%; ∼ 6.0 x 10 18 polymer chains), will be formed with G ∼ 400 Pa, an average number of interactions per particle (w ≈ 15) and an average number of interactions per polymer chain (m ≈ 3.5). These values fall well within a physically-realistic range. c. d.
Supplementary Figure 8: Cryogenic transmission electron micrographs (CryoTEM) of a. PSNPs (10 wt%), b. hydrogels prepared with HPMC-C 12 (1 wt%) and PSNPs (10 wt%), c. PEG-b-PLA NPs (10 wt%) and d. hydrogels prepared with HPMC-C 12 (1 wt%) and PEG-PLA NPs (10 wt%). In the case of PEG-b-PLA NPs, some particles are slightly larger in the hydrogels than in solution, yet there is clearly a homogeneous distribution of particles within the hydrogels structure, identifying that there is no appreciable particle agglomeration during the gel preparation. a. b. c.
Supplementary Figure 9: Representative photographs of PBS control (left) and PNP hydrogels (right) injected subcutaneously in C57BL6 mice a. immediately following and b. 24h after administration. c. A representative photograph of a PNP hydrogel harvested from a euthanized mouse after 7 days in vivo showing negligible volume decrease over this time.
Supplementary Figure 10: Histological screen for hydrogel biocompatibility at 3 days showing representative images of a. and b. the subcutaneous hydrogel injection site at low magnification (4x magnification), demonstrating retention of a continuous gel material under the skin following injection. Additionally, higher magnification images c.-h. are shown of the implantation bed. c. and d. illustrate the cell-material interface (10x), while e. is representative tissue (10x) following a control injection of saline. f. and g. illustrate the cell-material interface at higher magnification (20x), while h. is representative tissue (20x) following a control injection of saline. i. and j. illustrate the material interface (40x) where invading neutrophils are marked with green arrows. a., c., e., f., h., and i. are stained with H&E, while b., d., g., and j. are stained with Mason's Trichrome. In all images, a star indicates the position of the PNP hydrogels for orientation. These representative images illustrate our observations for the beginnings of mild infiltration into the implanted material primarily by neutrophils, with no evidence of inflammation or damage in the adjacent muscle tissue. Only the outer margins of the material have been infiltrated at this time. Additionally, no evidence of material fibrosis is present at this time.
Supplementary Figure 11: Histological screen for hydrogel biocompatibility at 7 days showing representative images of a. and b. the subcutaneous hydrogel injection site at low magnification (4x magnification), demonstrating retention of a continuous gel material under the skin following injection. Additionally, higher magnification images c.-h. are shown of the implantation bed. c. and d. illustrate the cell-material interface (10x), while e. is representative tissue (10x) following a control injection of saline. f. and g. illustrate the cell-material interface at higher magnification (20x), while h. is representative tissue (20x) following a control injection of saline. i. and j. illustrate the material interface (40x) where invading macrophages are marked with red arrows. a., c., e., f., h., and i. are stained with H&E, while b., d., g., and j. are stained with Mason's Trichrome. In all images, a star indicates the position of the PNP hydrogel material for orientation. These representative images illustrate our observations for the progression of a mild immune response within the implant bed, as neutrophils have disappeared and been replaced primarily by macrophages and some fibroblasts, with no evidence of multinucleate giant cells or lymphocytes. There is no evidence of inflammation or damage in the adjacent muscle tissue. Cell infiltration has progressed within the material, as diffuse matrix staining is now evident throughout. Macrophages have begun clearance of the PNP hydrogel materials, and the margins consist primarily of macrophages with phagocytosed material. There remains no evidence of material fibrosis at this time.

Supplementary Tables
Supplementary Rheological characterisation was performed using a TA Instruments AR-G2 stress controlled rheometer fitted with a Peltier stage set to 37 • C. Dynamic oscillatory strain amplitude sweep measurements were conducted at a frequency of 10 rad/s (unless otherwise noted). Dynamic oscillatory frequency sweep measurements were conducted at a 2% strain amplitude (unless otherwise noted). All measurements were performed using a 20 mm 4 • cone geometry and analyzed using TA Instruments TRIOS software.
Cryogenic scanning electron microscopy (CryoSEM) images of PNP gels where acquired using a Zeiss NVision 40 (Carl Zeiss SMT, Inc.) field emission scanning electron microscope at an acceleration voltage of 2 kV. To prepare samples for imaging approximately 100 µL of gel was transferred to a sample stub and then plunged into a slushy (liquid and solid) nitrogen bath. The samples where next transferred to an EM VCT100 vacuum cryo-transfer system (Leica Microsystems, Inc.) to selectively remove surface water (vitreous ice) by controlled specimen sublimation. The frozen sample were then further fractured with a sharp blade and sputter coated with a thin layer of platinum and palladium metals prior to imaging.
Cryogenic transmission electron microscopy (CryoTEM) images where acquired using a JEOL 2100 FEG microscope (Jeol Inc. Peabody, MA) equipped with an Gatan 2kx2k UltraScan CCD camera at an acceleration of 200 kV and at magnification ranges of 10,000-30,000x. To prepare samples for imaging, approximately 3 µL of nanoparticle suspensions was transferred to a lacey copper grid (coated with continuous carbon). Next, using a Gatan Cryo Plunge III the grids where blotted with great care to remove any excess liquid without causing damage to the carbon layer. The prepared grids where then mounted on a Gatan 626 cryo-holder equipped in the TEM column. The specimen and holder tips were next cooled down using liquid nitrogen, and subsequently transferred to the CryoTEM for imaging.
Monomethoxy-poly(ethylene glycol) (PEG; 5 kDa) was purchased from Aldrich and was purified by azeotropic distillation with toluene. Lactide (LA) was purchased from Aldrich and dried in a desiccator over P 2 O 5 prior to use. Carboxy-functional poly(styrene) nanoparticles (PSNPs) were purchased from Phosphorex and used as received. All other materials were purchased from Aldrich and used as received.

Synthesis of functional polymers
Synthesis of PEG-b-PLA block copolymers: PEG (0.25 g, 4.1 mmol) and 1,8-diazabicycloundec-7-ene (DBU; 10.6 mg, 10 µL, 1.0 mol% relative to LA) was dissolved in DCM (1.0 mL). LA (1.0 g, 6.9 mmol) was dissolved in DCM (3.0 mL) with mild heating. The LA solution was then added rapidly to the PEG/DBU solution and was allowed to stir rapidly for 10 min. The reaction mixture was then quenched by addition of acetone (7.0 mL) and the PEG-b-PLA copolymer was recovered by precipitation from cold diethyl ether, collected by filtration, and dried under vacuum to yield a white amorphous polymer (1.15 g, 92%). GPC (THF): M n (PDI) = 25 kDa (1.09).
General synthesis of HPMC-x polymers: HPMC (1.0 g) was dissolved in N-methylpyrrolidone (NMP; 45 mL) by stirring at 80 • C for 1 h. Once the solution had cooled to room temperature, a solution of 1-hexylisocyanate, 1-adamantylisocyanate, or 1-dodecylisocyanate (0.5 mmol) and triethylamine (2 drops) was dissolved in NMP (5 mL) and added to the reaction mixture, which was then stirred at room temperature for 16 h. This solution was this precipitated from acetone and the polymer was recovered by filtration, dried under vacuum at room temperature for 24 h and weighed, yielding the functionalised HPMC as a white amorphous powder (0.96 g, 96%). FT-IR:ν = 1685, 1601, 1367 cm −1 .

General preparation of self-assembled hydrogels
Hydrogels were prepared by first dissolving HPMC-x (30 mg) in water (1.0 mL) with stirring and mild heating, where x is C 6 , Ad, or C 12 . PSNPs (D H ∼ 50, 75, 100, 200, or 500 nm) or PEG-b-PLA NPs (D H ∼ 75 nm, prepared according to literature procedures, 1 were concentrated to 15 wt% solutions in water. If aggregation was observed in the PSNPs, the solution was sonicated to re-disperse the particles. To prepare polymer-nanoparticle hydrogels (PNP gels), HPMC-x solutions were combined with NP solutions to a finally weight fraction of HPMC (1 wt%) : NPs (1-10 wt%). PNP gels were mixed well by vortex, mild centrifugation, and agitation to enable homogenization and removal of bubbles. For rheometry, 450 µL gels were prepared.

Release studies from hydrogels
Two experiments were designed to investigate encapsulation and release of a hydrophobic molecule (Oil Red dye; OR) and a hydrophilic molecule (Bovine Serum Albumin; BSA). For the first experiment, hydrogels were prepared as mentioned above except with FITC-labeled BSA dissolved alongside the HPMC-C 12 polymer, resulting in a final concentration of BSA of 1 wt% in the hydrogel. For the second experiment, OR-loaded PEG-b-PLA NPs were prepared by co-dissolving OR with the PEG-PLA block copolymers in DMSO and co-precipitation into water. These NPs were then used to prepare hydrogels with HPMC polymers as above.
Hydrogel of either type (200 µL) was placed into a 1.5 mL centrifuge tube and deionized water (1.3 mL) was added on top of the hydrogel. This setup was placed into an incubator at 37 • C and 1 mL of the aqueous supernatant solution was replaced with fresh deionized water at predetermined time intervals. The collected aqueous solutions were analyzed for solute concentration based on calibration curves prepared using either OR or FITC-Albumin absorbance. These experiments were performed in triplicate and the mean and standard deviation are reported.
Modeling of the controlled release of drugs and other cargo from these polymeric devices has been a subject of considerable research. Many models in the past have focused on Fickian diffusion, 2 however, Peppas and coworkers [3][4][5] have derived an important and exceedingly simple exponential relationship to describe the general release behaviour of cargo from a polymeric device: M t / M ∞ = kt n , where k is the Fickian diffusional release rate and n is the release coefficient. The release coefficient is an indicator of the mode of release from the materials, where pure Fickian diffusional release is described by n = 0.5, with n > 0.5 revealing a greater contribution of "anomalous" release (caused by swelling or hydrogel erosion) to the overall release profile. We fit the BSA release data with this equation and the n and k values are reported in the manuscript.

In vivo implantation of hydrogels by injection
Hydrogels (100 µL; 1 wt% HPMC-C 12 : 10 wt% PEG-b-PLA NPs) were injected subcutaneously on the back of shaved 8-week old male C57BL6 mice using a 26G syringe. Photographic images were taken immediately following injection and 24h following injection. At 7 days following administration, mice were euthanized and the hydrogel was harvested.
In vivo biocompatibility studies and histological analysis Hydrogels (100 µL; 1 wt% HPMC-C 12 : 10 wt% PEG-b-PLA NPs) were injected subcutaneously on the back of shaved 8-week old male C57BL6 mice using a 26G syringe. At 3 and 7 days following administration, mice were euthanized and the hydrogel and surrounding tissue was harvested. A total of 3 mice were used for each time point. Tissue was fixed for 24 hours in formalin and subsequently processed for histological analysis by standard paraffin embedding and histological methods. Cross-sections of the skin and underlying material, approximately 40 µm in thickness, were stained with standard haematoxylin and eosin (H&E) or Mason's trichrome. Tissue was observed using light microscopy, and characterised for the degree and type of immune cell infiltrate, the appearance of fibrosis, and cell-based PNP hydrogel material clearance. These studies were approved by the MIT Animal Care and Use Committee.

In vivo release and intravital imaging
Hydrogels (100 µL; 1 wt% HPMC-C 12 : 10 wt% PEG-b-PLA NPs) prepared with Alexa Fluor ® 680-conjugated bovine serum albumin (BSA-AF; Life Technologies) loaded into the aqueous phase (BSA-AF concentration of ≈ 10 µM in the final gel) and Texas Red ® 1,2-dihexadecanoylsn-glycero-3-phosphoethanolamine, triethylammonium salt (TR-DPHE) encapsulated within the PEG-b-PLA NPs (TR concentration of ≈ 10 µM in the final gel). Control hydrogels, containing only one of the fluorescent compounds, were prepared similarly. For in vivo imaging, 8-week old male hairless SKH1-E mice were first maintained on an alfalfa-free diet for two weeks prior to administration to limit background fluorescence. Mice were anesthetised using inhaled isoflurane, and 100 µL of gel or bolus control was injected subcutaneously into the rear right flank of the animal using a 26G syringe. Treatment groups consisted of the hydrogel with the combined fluorophores (n=5), the control hydrogel with TR only (n=2), the control hydrogel with BSA-AF (n=2), a bolus injection of BSA-AF (n=1), a bolus injection of TR-loaded NPs (n=3), and an injection of BSA-AF and HPMC-C 12 (n=3). Imaging was conducted on an IVIS ® Spectrum in vivo imaging system with a heated stage and an inhaled isoflurane manifold. Fluorescent images were collected at several time-points over the following week, using filter sets of 570/620 (Texas Red) and 675/720 (AF-680) with a 1.5 cm subject height using small binning and an F-stop of 1. For display, the fluorescent signal was false-coloured and overlaid on a bright field image to provide spatial orientation. These studies were approved by the MIT Animal Care and Use Committee.