Tissue Response and Biodistribution of Injectable Cellulose

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Bio-interactions and Biocompatibility

Tissue Response and Biodistribution of Injectable Cellulose Nanocrystal Composite Hydrogels Kevin J. De France, Maryam Badv, Jonathan Dorogin, Emily Siebers, Vishrut Panchal, Mouhanad Babi, Jose M. Moran-Mirabal, Michael Lawlor, Emily D. Cranston, and Todd Hoare ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.9b00522 • Publication Date (Web): 25 Apr 2019 Downloaded from http://pubs.acs.org on April 25, 2019

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ACS Biomaterials Science & Engineering

Tissue Response and Biodistribution of Injectable Cellulose Nanocrystal Composite Hydrogels Kevin J. De France1, Maryam Badv2,†, Jonathan Dorogin2,†, Emily Siebers3, Vishrut Panchal2, Mouhanad Babi4, Jose Moran-Mirabal4, Michael Lawlor3, Emily D. Cranston1,5,6, Todd Hoare1,2*

1. Department of Chemical Engineering, McMaster University, 1280 Main Street West, Hamilton, ON L8S 4L8, Canada 2. School of Biomedical Engineering, McMaster University, 1280 Main Street West, Hamilton, ON L8S 4L8, Canada 3. Department of Pathology and Laboratory Medicine and Neuroscience Research Center, Medical College of Wisconsin, 8701 Watertown Plank Road, Milwaukee, WI 53226, United States 4. Department of Chemistry and Chemical Biology, McMaster University, 1280 Main Street West, Hamilton, ON L8S 4M1, Canada 5. Department of Wood Science, University of British Columbia, 2424 Main Mall, Vancouver, BC, V6T 1Z4, Canada 6. Department of Chemical and Biological Engineering, University of British Columbia, 2360 East Mall, Vancouver, BC V6T 1Z3, Canada

* To whom correspondence should be addressed E-mail: [email protected]

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ABSTRACT Interest in cellulose nanocrystal (CNC)-based hydrogels for drug delivery, tissue engineering, and other biomedical applications has rapidly expanded despite the minimal in vivo research reported to date. Herein, we assess both in vitro protein adsorption and cell adhesion as well as in vivo subcutaneous tissue responses and CNC biodistribution of injectable CNCpoly(oligoethylene glycol methacrylate) (POEGMA) hydrogels. Hydrogels with different PEG side chain lengths, CNC loadings, and with or without in situ magnetic alignment of the CNCs are compared. CNC loading has a minimal impact on protein adsorption but significantly increases cell adhesion. In vivo, both CNC-only and CNC-POEGMA injections largely stay at their subcutaneous injection site over one month, with minimal bioaccumulation of CNCs in any typical clearance organ. CNC-POEGMA hydrogels exhibit mild acute and chronic inflammatory responses, although significant fibroblast penetration was observed with the magnetically aligned hydrogels. Collectively, these results suggest that CNC-POEGMA hydrogels offer promise in practical biomedical applications.

Keywords: cellulose nanocrystals, poly(oligoethylene glycol methacrylate), injectable hydrogels, tissue engineering, protein adsorption, cell encapsulation, tissue histology, biodistribution


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INTRODUCTION Cellulose nanocrystals (CNCs) are commercially available rigid rod-shaped nanoparticles extracted from cellulose, the most abundant natural polymer on earth.1 In particular, since CNCs are typically produced through a sulfuric acid hydrolysis process that leaves charged sulfate halfester groups on the surface,2 the resulting charged interface makes them a prime candidate for incorporation into hydrophilic polymer networks such as hydrogels.3 Typically, CNC nanocomposite hydrogels exhibit increased mechanical moduli,4–7 tunable swelling profiles,8–10 and network structuring/anisotropy based on the anisotropic nature of CNCs themselves.11–13 All these properties are attractive in terms of addressing many of the drawbacks of using hydrogels in biomedical applications such as cell substrates, drug delivery, and tissue engineering,3,14–18 in which the generally weaker mechanics and fully isotropic properties of conventional hydrogels can restrict their utility. Despite this wide-spread interest in designing new CNC-hydrogel materials for biomedical applications, and multiple studies that suggest that CNCs have minimal to no in vitro cytotoxicity,19–26 relatively few studies have explored the tissue responses and biodistributions of CNCs27–30 or CNC-hydrogel nanocomposites6,31–34 in vivo. In terms of CNCs alone, O’Connor et al. studied the toxicity associated with different routes of administration of CNC suspensions and found no acute adverse effects via ingestion or topical administration;29 however, the biodistribution of CNCs following either route of administration was not studied. Yanamala et al. demonstrated dose-dependent tissue damage and acute inflammatory responses upon inhalationbased exposure to CNC suspensions (a delivery route of particular concern due to the rod-shaped morphology of CNCs) but did not investigate possible chronic responses.30 Colombo et al. showed



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that CNCs administered via tail vein injection at a concentration of 35 µg/mL initially migrated to the liver, spleen, and kidneys followed by a progressive re-distribution away from the liver and the kidneys and toward the spleen within the first week post-administration,27 typical for splenic catabolism of nanoparticles of similar sizes over this time period.35 Interestingly, the authors also noted that CNCs demonstrated a tropism toward the limb bones (subsiding after 7 days), suggesting potential for using CNCs as a theranostic for bone cancer detection.27 In terms of CNChydrogel nanocomposites, studies have largely been focused on the incorporation of drugs/growth factors within nanocomposite hydrogels for wound healing and antitumor applications.6,32,34 In all cases, minimal inflammatory responses were observed at the site of subcutaneous implantation for up to 21 days post treatment.6,32,34 One other study has shown the use of hydroxyapatite-silk fibroin-CNC nanocomposite hydrogels for bone repair,31 showing enhanced bone regeneration after 12 weeks of culture relative to bare silk fibroin scaffolds. However, studies on the in vivo biodistribution as well as the long-term fate of CNCs within the scaffolds were not undertaken. Furthermore, all of the hydrogel studies listed here required surgical implantation, which limits future potential applications. Injectable hydrogels can be delivered without surgery and are therefore promising candidates for a wide variety of biomedical applications including drug delivery,36–39 tissue engineering,40–43 and cell encapsulation.21,44,45 Injectable hydrogels are typically formed from either shear thinning materials whose viscosity can be transiently reduced upon injection,37,46–48 or in situ via various physical49–52 or chemical53,54 crosslinking approaches. Due to the inherent limitations in the viscosity, mechanical properties, rate of degradation, and network stability of shear thinning and physically crosslinked hydrogels,37,40 extensive research interest has recently been focused on the synthesis of chemically crosslinked in situ gelling hydrogels. For most 4 ACS Paragon Plus Environment

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biomedical applications, the preferred formulation of such a hydrogel should be void-filling upon injection, use easily modifiable chemistry to allow for tunable network properties such as gelation kinetics and biodegradation, and demonstrate minimal non-specific protein adsorption.55 Recently, we have developed a platform of highly tunable injectable hydrogels based on the poly(oligoethylene glycol methacrylate) (POEGMA) family of synthetic polymers.56,57 The POEGMA platform has gained significant research interest in the field of biomedical engineering as it is readily (co)polymerized via free radical techniques, exhibits a tunable and physiologically relevant lower critical solution temperature (LCST), and offers low biofouling.58,59 By functionalizing POEGMA oligomers with kinetically bio-orthogonal hydrazide and aldehyde functional groups, we can form crosslinked hydrogel networks in situ following extrusion from a double barrel syringe.56 Furthermore, by copolymerizing different comonomers into the POEGMA backbone, these hydrogels can be made with varying gelation times,60 hydrophobized for increased drug loading,61 or charged for promoting cell adhesion.62 However, these and most other injectable hydrogels are typically limited by their weak mechanical properties and (in the case of engineering aligned tissues) their isotropic/random network structures, both of which hinder the practical applicability of such hydrogels as tissue engineering scaffolds. We have overcome these limitations by physically incorporating CNCs into POEGMA hydrogels,7,13,21,63 whereby the addition of 4.95 wt% CNCs leads to a 35-fold increase in the shear modulus of the hydrogels. Furthermore, the diamagnetic anisotropy of CNCs64 can be leveraged to align CNCs in situ within injectable POEGMA hydrogels, allowing for the differentiation of oriented myotubes for muscle tissue engineering.21 Herein, we evaluate the in vitro protein adsorption and cell adhesion as well as the in vivo biodistribution and both local and systemic tissue responses following subcutaneous injection of 5 ACS Paragon Plus Environment


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CNC suspensions and CNC-POEGMA nanocomposite hydrogels. More specifically, hydrogels prepared with different CNC contents (with and without magnetic alignment) and POEGMA with different PEG side chain lengths were tested in order to evaluate the potential clinical applicability of these materials. We demonstrate that the incorporation of CNCs has minimal influence on the non-fouling properties of the injectable POEGMA hydrogels while also exhibiting mild tissue responses, both locally at the site of injection and systemically as the hydrogels degrade over time. We anticipate these results are of substantial benefit to addressing questions regarding the suitability of CNC-hydrogel nanocomposites in biomedical applications.

EXPERIMENTAL SECTION Materials Di-(ethylene glycol) methyl ether methacrylate (M(EO)2MA, Sigma Aldrich, 95%) and oligo(ethylene glycol) methyl ether methacrylate (OEGMA500, Sigma Aldrich, 95%) were run through a column of basic aluminum oxide (Sigma Aldrich, type CG-20) to remove inhibitors and impurities before use. Co-monomer N-(2,2-dimethoxyethyl)-methacrylamide (DMAEAm) was synthesized as described previously.7,56 Acrylic acid (AA, Sigma Aldrich, 99%), 2,2azobisisobutyric acid dimethyl ester (AIBMe, Wako Chemicals, 98.5%), adipic acid dihydrazide (ADH, Alfa Aesar, 98%), N’-ethyl-N-(3-dimethylaminopropyl)-carbodiimide (EDC, Carbosynth, Compton CA, commercial grade), thioglycolic acid (TGA, Sigma Aldrich, 98%), sodium hydroxide (EMD Millipore Germany), sodium chloride (Sigma Aldrich, ≥99.5%), hydrochloric acid (LabChem Inc., 1 M), dioxane (Caledon Laboratory Chemicals, reagent grade), sulfuric acid (Sigma Aldrich, 95-98%), sulfo-Cy5 azide (Lumiprobe), bovine serum albumin (BSA, Sigma-


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Aldrich, > 96%), fibrinogen from human plasma (Sigma Aldrich), lysozyme from chicken egg white (Sigma Aldrich), fluorescein isothiocyanate (FITC, Sigma-Aldrich, 90%), and Whatman cotton ashless filter aid (CAT No. 1703-050, GE Healthcare Canada) were all used as received. 3T3 Mus musculus mouse fibroblast cells were obtained from ATCC: Cedarlane Laboratories (Burlington, ON, Canada) and cultured in high glucose Dulbecco’s modified Eagle media (DMEM) supplemented with 10% fetal bovine serum (FBS) and 1% penicillin streptomycin (PS). Trypsin-EDTA and a LIVE/DEAD assay kit were purchased from Invitrogen Canada (Burlington, ON, Canada) and used as received. Millipore Milli-Q grade distilled deionized water (DIW, 18.2 MΩ cm resistivity) was used for all experiments.

Synthesis and characterization of poly(oligoethylene glycol methacrylate) copolymers POEGMA precursor polymers were synthesized as described previously to contain either 100 mol% OEGMA500 (PO100) or 90 mol% M(EO)2MA/10 mol% OEGMA500 (PO10);7 the former is a highly hydrophilic PEG-like polymer while the latter is a thermoresponsive polymer with a phase transition temperature mimicking that of poly(N-isopropylacrylamide).65 In both cases, the precursor polymers were functionalized to contain either 30 mol% aldehyde groups (POxA30) or 30 mol% hydrazide groups (POxH30) using the polymerization and reaction methods described previously.7 Polymers were characterized via aqueous size exclusion chromatography (Waters 515 HPLC pump calibrated with narrow-dispersed PEG standards, Waters 717 Plus autosampler, three Ultrahydrogel columns (30 cm × 7.8 mm i.d. with exclusion limits of 0–3 kDa, 0–50 kDa and 2– 300 kDa) and a Waters 2414 refractive index detector) and 1H NMR (Bruker AVANCE 600 MHz spectrometer with deuterated chloroform as the solvent) to determine molecular weight and degree of functionalization (Table S1). 7 ACS Paragon Plus Environment


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Preparation of cellulose nanocrystal suspensions Suspensions of CNCs were prepared via the sulfuric acid-mediated hydrolysis of cotton Whatman ashless filter aid, as previously described.7 CNC dimensions were measured to be 100 – 200 nm in length and 5 – 12 nm in cross-section using atomic force microscopy (MFP-3D, Asylum Research an Oxford Instruments Company, Santa Barbara, CA). The anionic sulfate half-ester content on the CNC surface was determined via conductometric titration to be 0.42 wt% (≈ 0.30 charges per nm2). The electrophoretic mobility of the CNCs (0.25 wt% suspensions in 10 mM NaCl) was determined using a ZetaPlus analyzer (Brookhaven Instruments Corp.) to be –1.86 × 10–8 m2 V–1 s–1.

Preparation of injectable nanocomposite hydrogels Individual 20 or 40 wt% POEGMA copolymer solutions were mixed with CNC suspensions of concentrations ranging between 0 and 8.25 wt% to create precursor solutions with a final polymer concentration of 16 wt% and CNC concentrations ranging between 0-4.95 wt% in PBS (denoted POx-CNC, where x is either 10 or 100 and CNC is the final CNC wt%). These resulting precursor solutions were loaded into separate barrels of a double barrel syringe equipped with a static mixer (MedMix, L-System) and reactively coextruded under ambient conditions to form hydrogels. In cases in which magnetic alignment was desired, a 0.56 T rare earth magnet was used to align CNCs within the nanocomposite hydrogels according to our previously reported protocol (i.e., placing molds between the magnet poles and extruding hydrogels into molds under 8 ACS Paragon Plus Environment

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ambient conditions).27 Hydrogel nomenclature is denoted as “POx-CNC”, where x indicates the POEGMA precursor series (10 or 100) and CNC represents the overall wt% of CNC loaded.

In vitro protein adsorption assay Fluorescein-isothiocyanate (FITC)-labelled bovine serum albumin (FITC-BSA), FITC labelled fibrinogen (FITC-Fib) and FITC labelled lysozyme (FITC-Lys) were prepared by dissolving 50 mg of protein in 10 mL of 0.1 M carbonate buffer (pH 9.0). FITC (1 mg) was then added, and the solution was incubated in the dark at room temperature for at least 12 h under gentle mechanical agitation. The FITC-labeled proteins were dialyzed against deionized water for 6 × 6h cycles and lyophilized. The isolated conjugated proteins were stored at −4 °C in the dark. To determine protein adsorption, hydrogel precursor solutions were sterilized by passing the solutions through a 0.25 μm filter before use. PO100 series hydrogels were loaded into double barrel syringes, after which ~ 100 µL was added to the bottom of a well plate (just enough to cover the entire plate). Due to the extremely fast gelation time, PO10 hydrogels were instead prepared by sequentially extruding 50 μL of both precursor solutions into each well and manually mixing, again just enough to cover the entire plate. All hydrogels were subsequently covered with 200 μL of 10 mM PBS and allowed to swell for 30 h in order to reach equilibrium. Excess PBS was then removed, and 60 μL of either FITC-BSA, FITC-Fib, or FITC-Lys solutions in PBS were added at concentrations of 125, 250, 500, 1000, or 2000 μg/mL. Hydrogel samples were then incubated for 2 h at 37 °C and subsequently rinsed 3 × 10 minutes in 10 mM PBS to remove any non-adsorbed protein. A VICTOR 3 multilabel microplate reader operating at an excitation wavelength of 495 nm and an emission wavelength of 535 nm was used to measure the residual fluorescence in each well (n = 5 for each condition tested, with reported errors representing the standard deviation of 9 ACS Paragon Plus Environment


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the replicates). Protein absorption (uptake into the hydrogel network) was assessed through a similar procedure in which 2000 µg/mL protein solutions were added to hydrogels as described above, incubated, rinsed, and imaged using a confocal microscope (Zeiss LSM 510 laser scanning microscope, Oberkochen, Germany). Normalized fluorescent intensity was measured in the zdirection of the hydrogels to determine the penetration of proteins throughout the gels.

In vitro cell adhesion assay Cell adhesion to the hydrogels was assessed using NIH 3T3 mouse fibroblasts as model cells. Hydrogel samples (125 µL total volume) were prepared in 96-well culture plates from UVsterilized precursor polymers and allowed to swell overnight in sterile PBS at 37°C. After PBS removal, 3T3 cells (150,000 cells/mL DMEM, 200 µL DMEM/well) were plated on top of these hydrogels and incubated for either 48 hr or 1 week at 37°C. Plates were gently washed with PBS to remove non-adherent cells and then stained with calcein AM/ethidium homodimer-1 according to the manufacturer protocol to assess the percentage of live and dead cells remaining at the hydrogel interface. Samples were imaged with a Zeiss LSM 510 using a 488 nm laser and a BP 505−530 nm emission filter (for calcein AM) or a 543 nm laser with a LP 560 nm emission filter (for ethidium homodimer-1). Images were processed using Zeiss LSM Image Browser software (version 4.2).

Cy5 labelling and in vivo cellulose nanocrystal biodistribution study To enable tracking of the biodistribution of CNCs following injection, CNCs were tagged with the stable fluorophore Cy5. CNCs were first functionalized with dichlorotriazinyl aminopropyne (DTAP) to introduce an alkyne handle on the surface of the CNCs. DTAP (800 mg,


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synthesized in-house)66 was dissolved in 15 mL of acetone and added to a 16 mL solution of 1.6 wt% CNCs under magnetic stirring. NaOH (8 mL, 0.5 M) was subsequently added, and the reaction proceeded at room temperature overnight. The alkyne-functionalized CNCs were cleaned through repeated sequential cycles of precipitation by brine, centrifugation, decanting, and resuspension in each of dichloromethane, acetone, and water (minimum 5 cycles of precipitation/resuspension per solvent). Following, sulfo-Cy5 azide was grafted to the alkynefunctionalized CNCs via a copper-mediated click reaction by adding 2.4 mg of sulfo-Cy5 azide, 15 mg of ascorbic acid, and 5 mg of copper sulphate pentahydrate to the suspension and stirring overnight. The Cy5-CNCs were cleaned through the same procedure as indicated above using only water as the solvent and resuspended using point-probe sonication (Branson, Danbury, USA). Cy5-CNC suspensions at 1.0 wt% (no gel components) and PO100-1.0 wt% Cy5-CNC hydrogel precursors (i.e., a fluorescent CNC nanocomposite hydrogel) were prepared for subcutaneous injection into male BALB/c mice (Charles River, Montreal; 19−23 g) by filtering using a 0.2 µm syringe filter. All animals received care that complied with protocols approved by the Animal Research Ethics Board at McMaster University and the guidelines of the Canadian Council on Animal Care. Samples were injected at a volume of 0.2 mL/mouse into the scruff of the neck from a single (Cy5-CNCs only) or double (Cy5-CNC-POEGMA hydrogel) barrel syringe. In both cases, Cy5-CNCs were injected at a concentration of 10 mg/mL; the POEGMA concentration was held constant at 160 mg/mL to be consistent with the previously described experiments. Mice were euthanized by carbon dioxide asphyxiation at 2 day, 7 day, and 28 day time points (n = 3 for each condition/time point). The lungs, kidneys, spleen, liver, cecum, and tissue surrounding the site of injection were harvested, wrapped in tin foil, and frozen overnight at –80 ℃. To assess the Cy5 fluorescence in each organ/tissue samples, the tissues were weighed and 11 ACS Paragon Plus Environment


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homogenized in an appropriate volume of PBS to make 2 mg/mL solutions. These solutions were plated in quintuplet in 96-well plates and analyzed for fluorescence intensity (Tecan M1000 microplate reader, 633/670 nm excitation/emission). In a separate experiment, the lungs, kidney, spleen, liver, and tissue surrounding the site of injection were recovered from CNC-only injected mice (n = 3), fixed in formalin, stained with hematoxylin and eosin (H & E), and imaged to analyze organ histopathology. Inflammatory responses to the hydrogels were assessed by blinded observers (ML, ES) using a histological scoring system to semi-quantitatively describe the intensity of inflammation, where 0 = normal, 1 = inert/no acute inflammation, 2 = inflammation without necrosis, 3 = focal inflammation with some necrosis, 4 = moderate inflammation with past necrosis, fibroblast proliferation, 5 = massive inflammation

In vivo histology study In vivo tissue responses were assessed histopathologically in male BALB/c mice following subcutaneous injection of the nanocomposite hydrogels using double barrel syringes. Precursor POEGMA solutions (160 mg/mL) loaded with varying amounts of CNCs were filtered using a 0.2 µm syringe filter, loaded into sterile double barrel syringes under aseptic conditions, and injected subcutaneously in the scruff of the neck at a volume of 0.3 mL/mouse; note that gravimetric measurements post-filtration confirmed no significant change in the mass concentration of the filtered solution. Mice were anesthetized using isoflurane prior to injection to ensure reproducible injection sites and substantial gelation of the injected payload prior to mouse movement. Following visual behavior and health analysis during the experiment, animals were euthanized by carbon dioxide asphyxiation after acute (2 days) and chronic (30 days) time points (n = 3 for each condition/time point). Tissue samples were recovered from the injection site, fixed in formalin,


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stained with H & E, and imaged (3 tissue samples for each condition/time point). Inflammatory response was assessed using the scoring system described above.



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RESULTS & DISCUSSION Synthesis of Precursor Materials and Hydrogel Preparation POEGMA precursor copolymers were synthesized by free radical polymerization as described previously,56 and tailored to have molecular weights well below the physiological renal clearance limit (40 – 50 kDa)67 in order to allow clearance of the hydrogel degradation products (Supporting Information, Table S1). POEGMA copolymers were functionalized with ~30 mol% of kinetically bio-orthogonal hydrazide and aldehyde groups to facilitate in situ crosslinking via hydrolytically degradable hydrazone bonds. Experimentally measured values (by conductometric titration for hydrazide-functionalized polymers and

1H

NMR for aldehyde-functionalized

polymers) were nearly identical to each other but slightly higher than the stoichiometric prediction (Table S1). Precursor polymers were synthesized to either have a long PEG side chain (PO100, 8-9 EO repeat units), or a 90:10 molar ratio of short:long PEG side chains (PO10, 90% 2 EO repeat units, 10% 8-9 EO repeat units). PO10 polymers gel rapidly to form stiffer networks with limited swelling, whereas PO100 polymers gel much slower and form weaker networks with increased swelling due to the higher hydrophilicity and the steric hindrance toward crosslinking presented by the longer PEG side chains.56 Colloidally stable CNC suspensions were added to each precursor solution at concentrations varying from 0.2 – 4.95 wt% (the upper limit being associated with the self-gelling concentration of CNCs),68 after which the mixtures were added to either side of a double barrel syringe and pushed through a static mixer to form in situ gelling hydrogels with physically entrapped CNCs. Note that the CNCs are not covalently crosslinked to the hydrogel


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network, but we have shown previously that POEGMA strongly adsorbs to CNCs leading to significant entanglement and physical crosslinking within the composite network.7

In Vitro Protein Uptake To assess the in vitro performance of the nanocomposite hydrogels toward biomedical applications, 2D protein adsorption assays were performed using fluorescently tagged fibrinogen (FITC-Fib), bovine serum albumin (FITC-BSA), and lysozyme (FITC-Lys) as the model proteins (Figure 1). These proteins were chosen to cover a range of molecular weights and charges relevant to the broad spectrum of proteins found in vivo. In order to determine any significant effects of CNC loading, the average overall protein adsorption was calculated for each hydrogel series as a percent of total protein added. In general, PO100 hydrogels demonstrate low non-specific protein adsorption due to the presence of long PEG side chains.56 Adding a relatively high loading (PO1001.65) of CNCs led to slightly decreased BSA and fibrinogen adsorption yet increased lysozyme adsorption relative to POEGMA-only controls (p < 0.01 for BSA and fibrinogen, p < 0.05 for lysozyme, Figure 1). We attribute the decrease in BSA and fibrinogen adsorption upon CNC loading is attributed to the physical adsorption of the methacrylate backbone of POEGMA onto the CNC surface,7 increasing the percentage of protein-repellent PEG side chains exposed to the protein solutions. In contrast, the increase in lysozyme adsorption is attributed to electrostatic attraction introduced between the anionic CNCs and cationic lysozyme (pI ~ 9). At the highest CNC loading (PO100-4.95), slightly higher BSA uptake was observed that may be a result of its interactions with the (200) hydrophobic crystal plane of cellulose,69 although other protein adsorption was similar.



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Within the PO100 series, all changes in protein adsorption (although statistically significant) were relatively small, and all CNC-PO100 hydrogels were still highly protein-repellent relative to most biomaterials. PO10 hydrogels are less inherently protein repellent than their PO100 counterparts due to their shorter PEG side chains;65 however, similar general trends were observed when CNCs were added, with the incorporation of 1.65 wt% CNCs resulting in a significant decrease in BSA adsorption (p 0.05 for each pair-wise comparison). This result suggests that non-specific protein adsorption is not governed by CNC alignment but rather by the physicochemical properties of nanocomposite hydrogel networks.

Figure 2: Comparison of fibrinogen (A), BSA (B), and lysozyme (C) adsorption on PO100-1.65 hydrogels with isotropic (unaligned, filled bars) and magnetically aligned (striped bars) CNCs. No statistical difference was observed in overall percent protein adsorption (D) for BSA, fibrinogen, or lysozyme (p > 0.05 for each pair-wise comparison of aligned vs. unaligned).



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In Vitro Cell Adhesion Cell adhesion of 3T3 mouse fibroblasts to CNC-POEGMA hydrogels with varying compositions was assessed by preparing PO10 and PO100 nanocomposite hydrogels in a similar manner as for protein uptake experiments, plating 30,000 cells/well on top of the hydrogels, and incubating the cells at 37 ˚C. Note that no PO10-4.95 hydrogel could be assessed within this dataset since the high 4.95 wt% CNC content results in increased viscosity and almost instantaneous gelation upon mixing, making it difficult to generate uniform hydrogels for testing. For both PO10 and PO100 gels, increasing the CNC loading led to an increase in cell adhesion apparent after 24 hours of incubation and a significant increase in cell density after 7 days of incubation (Figure 3). This result is typical of CNC films as well as CNC-based nanocomposite hydrogels.70,71 In particular, while zero cell adhesion was observed at both time points tested for the PO100 hydrogel alone, introducing CNCs allowed for at least some degree of cell adhesion. We attribute this result to a combination of the increased mechanical strength of the hydrogel and inherent adhesive interactions between cells and cellulose surfaces.72 Importantly, cell viability across various CNC loadings for both PO10 and PO100 hydrogels remains statistically similar to TCPS controls (select red channel images demonstrate no adhered dead cells, Supporting Information Figure S2), suggesting that the low number of viable cells noted in these images is related to differences in cell adhesion rather than any potential materials-driven cytotoxicity. However, the adhered cells tended to clump together, indicative of relatively weak cell-substrate interactions;62,73 this phenomenon is readily seen in higher magnification confocal imaging (Supporting Information Figure S3). In addition, cell adhesion and subsequent proliferation on the CNC-POEGMA hydrogels was significantly less than that facilitated by the polystyrene control. As such, CNCPOEGMA hydrogels can support weak cell adhesion and localized cell proliferation that can be 18 ACS Paragon Plus Environment

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enhanced as the CNC loading is increased, despite the relatively low (and, in some cases, lower) non-specific protein binding of CNC-POEGMA hydrogels compared to the POEGMA-only gels that are not cell-adherent (Figure 1).

Figure 3: Cell adhesion of 3T3 fibroblasts after 2 days (A-H) and 7 days (I-P) of incubation on top of PO10 (A-C, IK) or PO100 (E-H, M-P) hydrogels. From left to right, PO10 gels are loaded with 0, 0.2, or 1.65 wt% CNC, while PO100 gels are loaded with 0, 0.2, 1.0, or 4.95 wt% CNC respectively. Samples were washed with PBS and stained with Calcein AM/ethidium homodimer-1 following manufacturer recommended protocols prior to imaging. Images were obtained relative to a tissue culture polystyrene (TCPS) control (D, L). All scale bars are 100 µm.



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The effects of CNC alignment on fibroblast adhesion were also assessed (Figure 4). Similar to the protein adsorption result (Figure 2), no significant differences in the amount or orientation of cells adhered to the hydrogel substrates were observed between the magnetically aligned (PO1001.65-A) and unaligned (PO100-1.65) hydrogels. This result suggests that fibroblast adhesion is not governed by CNC alignment but rather by the physicochemical and interfacial properties of nanocomposite hydrogel networks. The result is also consistent with our previous findings on C2C12 mouse myoblast cells, in which cell alignment on the magnetically-aligned CNCPOEGMA hydrogels was observed only upon differentiation of the myoblasts into myotubes that (unlike fibroblasts) align in their functional tissue state.21

Figure 4: Cell adhesion of 3T3 fibroblasts after 2 days (A, B) and 7 days (C, D) of incubation on top of PO100-1.65 hydrogels with unaligned (A, C) or magnetically aligned (B, D) CNCs. Here, images are relative to the same tissue culture polystyrene (TCPS) control as in Figure 3. All scale bars are 100 µm.

In Vivo CNC Biodistribution


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Prior to assessing the impacts of various hydrogel compositions on in vivo tissue responses, we sought to understand how (or if) CNCs administered alone or in hydrogels via subcutaneous injection transport throughout the body. CNCs were fluorescently labelled using a click reaction of alkyne-functionalized CNCs66 with Cy5 and injected subcutaneously into mice in the nape of the neck at a concentration of 1.0 wt% either as a free suspension in water or physically entrapped within a PO100 POEGMA hydrogel. Mice were sacrificed after 2 days, 7 days, and 28 days, after which tissues were collected, homogenized, and analyzed fluorometrically to quantify the presence of CNCs in each tissue (Figure 5). For the CNC-only injected mice (Figure 5A), subcutaneous tissue samples showed a large initial fluorescence intensity at the 2 day time point at the site of injection (SOI) that largely persisted over the 28 day experiment, albeit with some observed spreading throughout the surrounding subcutaneous space (in some cases, down the front limbs). Note that this concentration of CNCs does not self-gel in water;68 as a result, the apparent immobilization observed at the injection site is likely attributable to either ionotropic gelation by ions in vivo74 or the transport of a fraction of the water (dewatering) in the injection into the adjacent tissues to locally concentrate the CNCs beyond the critical gelation concentration. Histological scoring indicated that the injected CNCs elicited only a minimal local inflammatory response in the subcutaneous space at the acute time point (histological tissue score = 1-3) that diminished to a minimal score 0-1 inflammatory response by 7 days post-injection (Supporting Information, Figure S4). All of the normal nanoparticle clearance organs analyzed (lung, liver, kidney, spleen, cecum) showed minimal concentrations of CNCs at the acute time point that further reduce over time (Figure 5A); correspondingly, all of the analyzed organs demonstrated normal pathologies (Supporting Information, Figure S5), with inflammatory scores of 0 noted for all tissues at all time points tested. 21 ACS Paragon Plus Environment


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However, over the duration of the study, both the residual fluorescence and the mass of gel recovered upon sacrifice indicated an apparent decrease in the local mass of injected CNCs at the injection site. As such, CNCs appear to be slowly cleared from the injection site but not partitioned to typical nanoparticle clearance organs. Instead, we noted that several mice had enlarged cecums that, while largely maintaining normal tissue pathology (inflammation score 0 at all time points tested), showed some evidence of lymphoid tissue aggregates (Supporting Information, Figure S4F). In herbivores such as laboratory BALB/c mice, the cecum collects foodstuffs by which local bacteria break down any contained cellulose.75 As such, we hypothesize that CNCs are distributed towards the cecum for eventual breakdown and ultimate clearance through the intestines. Indeed, subcutaneous injection of a three-fold lower concentration of CNCs alone showed similar accumulation in the cecum, although without evidence in that case of cecum enlargement.27

Figure 5: Biodistribution of CNCs subcutaneously injected at 1.0 wt% as a CNC-Cy5 suspension (A) and within a PO100/CNC-Cy5 nanocomposite hydrogel (PO100-1.0, B). Tissues were collected after 2 days (green), 7 days (blue), and 28 days (red), from the organs listed and site of injections (SOI), homogenized, and evaluated fluorometrically. Note that fluorescent data for the cecum at 28 day time point was not collected. * = p < 0.05, ** = p < 0.01 in a pairwise comparison.


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Similar results were observed in the mice injected with the PO100-1.0 nanocomposite hydrogels (Figure 5B), with a high intensity of fluorescence initially observed at the SOI that diminished significantly at the 7 and 28 day time points. However, unlike the free CNC suspension, the fluorescence related to the CNCs did not distribute throughout the subcutaneous space, with the gel remaining highly localized at the SOI throughout the full 28 days of observation. This is consistent with the rapid physical immobilization of the CNCs in the POEGMA hydrogel preventing the subcutaneous transport observed in the CNC-only injections that require a longer period of time to gel via either ionic interactions or dehydration.74 In addition, only extremely small concentrations of CNCs were detected in the clearance organs at the 2 day time point that either remained steady or further decreased throughout the course of the study, thus indicating that CNCs do not tend to accumulate in normal clearance organs. However, the fluorescence at the injection site (related to the total number of CNCs present at the injection site) decreases significantly faster in the case of the CNC-POEGMA hydrogels relative to that observed with the CNC-only injections. We hypothesize that the strong physical entrapment of CNCs within the POEGMA hydrogel, coupled with the strong interfacial interactions between CNCs and POEGMA, collectively reduce CNC aggregation within the gel phase and thus allow CNCs to be released as “free” nanoparticles as the gel degrades, enabling their faster clearance. Conversely, for CNC-only injections, the lack of a crosslinked hydrogel network results in their rapid concentration at the injection site (due to dewatering into the native tissue) and direct exposure to the high salt physiological environment,7 increasing the potential for CNC aggregation may retard clearance. Furthermore, it is likely that dewatering occurs to a lesser extent for the injected hydrogel samples versus the CNC-only injections, as POEGMA hydrogels have good water retention/swelling. As with CNC-only injections, no significant inflammation was observed in any



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of the typical clearance organs studied, with inflammation scores of 0 noted for all organs at all time points. Thus, CNC-POEGMA hydrogels were largely retained at the injection site and did not induce any systemic inflammation, motivating additional studies on how the CNC and POEGMA compositions of the hydrogels may influence the types of in vivo responses observed.

In Vivo Histology The tissue compatibility of CNC-POEGMA nanocomposite hydrogels with different CNC loadings and POEGMA compositions (PO100 compared with PO10) was assessed via subcutaneous injection in BALB/c mice. No obvious skin irritation was observed post-injection for any of the tested hydrogel formulations. Both PO10 and PO100 samples showed moderate inflammatory responses at the acute (2 day) time point independent of CNC loading (Figure 6). Moderate leukocyte infiltration was observed at the tissue-hydrogel interface in all samples, with the average inflammation score of 2-3 indicating the presence of a focal inflammatory response with limited cell necrosis (Table 1). Of note, for the particularly protein-repellent and slower gelling PO100 hydrogels, a clear increase in inflammation score was observed as the amount of CNCs incorporated into the gel was increased, from 0-1 at 0 wt% CNCs to 3 at 4.95 wt% CNCs. While this still represents only a moderate inflammatory response even at the highest CNC content, this result suggests that CNCs are slightly more pro-inflammatory than long-PEG side chain POEGMA, a highly efficient non-fouling/non-inflammatory material.65 In contrast, at the chronic time point (28 days), the subcutaneous inflammatory response was significantly diminished in all tested samples. A general decrease in thickness of the leukocyte cell layer was observed with a corresponding decrease observed in the overall histological score, indicating tissue responses between minimal inflammation without cell necrosis (score: 2) and the full restoration of normal


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tissue histology (score: 0), as indicated in Table 1 (see Supporting Information, Figure S6 for individual slide scores). No significant change in chronic inflammatory response was observed as the CNC content was increased for both PO10 and PO100 hydrogels. Collectively, these results indicate that the nanocomposite hydrogels are well-tolerated in vivo. Table 1: Histological scoring for quantifying tissue reaction and inflammation in response to subcutaneously injected hydrogel samples unaligned or aligned in a magnetic field. Scores are assigned as 0 = normal, 1 = inert/no acute inflammation, 2 = inflammation without necrosis, 3 = focal inflammation with some necrosis, 4 = moderate inflammation with past necrosis, fibroblast proliferation, 5 = massive inflammation.

Category Unaligned

Aligned

Hydrogel PO10 – 0.2 PO10 – 1.65 PO100 – 0.2 PO100-1.0 PO100-1.65 PO100 – 4.95 PO100-1.65-A

Acute (2 days) 2-3 0-3 2-3 0-1 1-2 3 0-3

Chronic (30 days) 0-2 0-2 2 1 1 0-1 1 - 2, 4

The histological results further indicate that gel degradation in vivo is dependent on both the CNC loading and POEGMA side chain length (PO10 vs. PO100) of the nanocomposite hydrogel (Figure 6). All PO10 hydrogel samples tested and PO100-4.95 (the highest CNC content PO100 gel) appeared to remain macroscopically intact 30 days post-injection; in contrast, PO100 samples containing lower CNC loadings began to degrade over this time period, with thin layers of cells migrating between the remaining gel fragments. Indeed, PO100-0.2 is almost fully resorbed by the chronic time point. This degradation profile observed in vivo qualitatively mirrors trends previously observed in acid-catalyzed accelerated degradation assays in vitro, in which the densely crosslinked PO10 networks and the mechanically stronger PO100 networks with high CNC loadings demonstrated enhanced stability against degradation.7 Thus, the degradation half-life of the nanocomposite hydrogels in vivo can be regulated by the CNC loading in the hydrogel without 25 ACS Paragon Plus Environment


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inducing significant changes in either the magnitude of the inflammatory tissue response (Table 1) or the degree of protein adsorption (Figure 1).


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Figure 6: Histological sections of H&E stained subcutaneous tissue samples for injected PO10 and PO100 hydrogels containing various loading of CNCs at acute (2 days) and chronic (30 days) time points. The hydrogel-tissue interface is indicated by dashed lines (i.e.: leukocyte cell layer). Note that in some cases, especially chronic time points and for low CNC loadings, cells infiltrate through the gel-tissue interface as the gel degrades. Scale bars = 200 µm.



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To assess the potential effects of magnetic CNC alignment on the tissue response, the subcutaneous injection was performed in the presence of a 0.56 T magnetic field for PO100-1.65, the gel composition previously identified to allow for effective CNC alignment.21 The tissue responses (Figure 7) were similar between the aligned and unaligned gels at the acute timepoint, consistent with the similar protein adsorption (Figure 2) and cell adhesion (Figure 4) observed. However, the aligned PO100-1.65-A gel showed clear evidence of fibroblast proliferation throughout the gel at the chronic time point that was not evident in the unaligned (isotropic) gel. It should be noted that this histology was not observed in all PO100-1.65-A gel samples evaluated, with some samples showing no fibroblast infiltration (tissue score 1-2) and others showing substantial infiltration (tissue score 4, as per Table 1); nonetheless, the appearance of fibroblast infiltration in at least some samples is consistent with the anisotropic mechanical properties of the aligned gel that provide a relatively stiff overall modulus conducive for fibroblast proliferation76 yet preferential swelling/degradation in a direction parallel to CNC alignment that allows easier gel remodeling as fibroblasts spread into the gel. Note that the cell growth observed throughout the network at 30 days post injection corresponds to the time at which the hydrogel network is observed to, at least partially, degrade (Figure 7D), providing cells with a pathway to proliferate into the gel and use the adhesion/alignment cues present from the CNCs to spread.


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Figure 7: Histological sections of H&E stained subcutaneous tissue samples for injected PO100-1.65 hydrogels containing unaligned (A-C) or magnetically aligned (D-F) CNCs at acute (A, D; 2 days) and chronic (B, C, E, F; 30 days) time points. Dashed lines indicate the gel-tissue interface. Fibroblast proliferation within subcutaneous tissue of PO100-1.65-A injected mice, 30 days post injection (F). Spindle-like fibroblasts are clearly seen in the top right and bottom left of the image, and are absent throughout the site of injection of unaligned PO100-1.65 hydrogels (C). Scale bars (A – E) = 200 µm, (F) = 20 µm. Note that the gel-tissue interface is denoted with a dashed black line.

CONCLUSION In summary, the in vitro and in vivo biological responses of injectable CNC-POEGMA hydrogels have been investigated as a function of the PEG side chain length of the POEGMA networks, the degree of CNC loading, and the CNC orientation within the hydrogel networks. In all cases, CNC-POEGMA hydrogels show low nonspecific protein adsorption (on par with other PEG-based hydrogels) but enhanced cell adhesion and proliferation as more CNCs are added. Both CNCs in suspension and CNCs entrapped in POEGMA hydrogel networks administered subcutaneously are largely immobilized at the injection site and do not appear to accumulate in any clearance organ, although evidence suggests clearance of the CNCs through the cecum as is 29 ACS Paragon Plus Environment


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typical for digestion/clearance of cellulosic materials in mice. Furthermore, after subcutaneous injection, CNC-POEGMA hydrogels elicit only mild acute inflammatory responses in vivo (dissipating to nearly no response at all at the chronic time point) and tunable degradation times according to both the POEGMA side chain length and the concentration of added CNCs. Hydrogels with aligned CNCs are also able to promote fibroblast proliferation within the hydrogel in vivo, indicating their potential utility in tissue engineering and regenerative medicine applications. Overall, the results presented suggest promise for the clinical use of CNCs and CNCbased injectable hydrogels in a range of potential biomedical applications.

Supporting Information Available. The Supporting Information is available free of charge on the ACS Publications website. Properties of POEGMA copolymers, protein absorption, red (dead cell) channel cell adhesion confocal images, additional high magnification cell adhesion confocal images, tissue histology at the site of injection for CNC-only injections, representative organ tissue histology for CNCPOEGMA hydrogel injection, raw data of histological scoring

Author Information. † Authors contributed equally.

Acknowledgements. Funding from the Natural Sciences and Engineering Research Council of Canada (Discovery Grants RGPIN 356609 and 402329) is gratefully acknowledged.


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For Table of Contents Use Only

Tissue Response and Biodistribution of Injectable Cellulose Nanocrystal Composite Hydrogels Kevin J. De France1, Maryam Badv2,†, Jonathan Dorogin2,†, Emily Siebers3, Vishrut Panchal2, Mouhanad Babi4, Jose Moran-Mirabal4, Michael Lawlor3, Emily D. Cranston1,5,6, Todd Hoare1,2*


Tissue Response and Biodistribution of Injectable Cellulose

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