Patent Application
20230173059
Clinical limitations of TLR agonist therapies are due to severe systemic toxicity and a narrow therapeutic window. Immunostimulatory molecules, like the TLR7/8 agonist Resiquimod (R848), have shown incredible promise as adjuvants in both prophylactic vaccines and cancer immunotherapies. Unfortunately, R848 has been limited clinically due to extreme systemic toxicity that results from the lack of pharmacokinetic control of the small hydrophobic drug. As such, there is a need for a delivery platform to overcome these limitations. The present invention satisfies this need and provides related advantages as well.
In some aspects, the present disclosure provides a nanoparticle comprising a polymer and a plurality of TLR agonist moieties conjugated to the polymer, wherein the plurality of TLR agonist moieties is present on the surface of the nanoparticle.
In some embodiments, the plurality of TLR agonist moieties is a plurality of TLR7/8 agonist moieties. In some embodiments, the plurality of TLR agonist moieties comprises a 1H-imidazo[4,5-c]quinolone core structure. The plurality of TLR agonist moieties may comprise resiquimod (R848), imiquimod, gardiquimod, or mixtures thereof. In some embodiments, the plurality of TLR7/8 agonist moieties comprises a plurality of N-(4-((4-amino-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-1-yl)methyl)benzyl)-3-(prop-2-yn-1-yloxy)propanamide moieties.
In some embodiments, the plurality of TLR agonist moieties is a plurality of TLR9 agonist moieties. The plurality of TLR9 agonist moieties may comprise a plurality of cytidine-phosphate-guanosine (CpG) moieties.
In some embodiments, the polymer comprises poly(ethylene glycol)-b-poly(lactic acid) (PEG-PLA). In some embodiments, the nanoparticle is up to about 200 nm in diameter. In some instances, the nanoparticle is about 50 nm in diameter. In other instances, the nanoparticle is about 30 nm in diameter. In some embodiments, the nanoparticle further comprises a plurality of mannose moieties conjugated to the polymer. In certain instances, the polymer is conjugated to the plurality of TLR agonist moieties via a 1,2,3-triazole linkage.
In other aspects, the present disclosure provides a method of producing a nanoparticle described herein, the method comprising:
In some embodiments, the polymer comprises an azide terminal group and the plurality of TLR agonist moieties comprises an alkyne derivative thereof.
In further aspects, the present disclosure provides a hydrogel comprising a nanoparticle described herein. In some embodiments, the hydrogel comprises optionally hydrophobically-modified hydroxypropyl methylcellulose (HPMC).
In related aspects, the present disclosure provides a vaccine comprising a nanoparticle and/or hydrogel as described herein. In some embodiments, the vaccine comprises one or more subunit antigens. In some embodiments, the antigen comprises a viral antigen, a bacterial antigen, a fungal antigen, or a protozoan antigen. In certain instances, the viral antigen is an antigen from a virus such as, e.g., a coronavirus (e.g., SARS-CoV or SARS-CoV-2), an influenza virus, a human immunodeficiency virus (HIV), a human papillomavirus (HPV), a porcine circovirus (PCV), etc. In certain instances, the protozoan antigen is from a parasite that causes a disease such as malaria. In particular embodiments, the viral antigen comprises a SARS-CoV-2 subunit antigen.
Methods for inducing an antigen-specific humoral immune response in a subject comprising administering a vaccine described herein and methods for enhancing cancer immunotherapy in a subject comprising administering a nanoparticle or hydrogel described herein are also provided. In embodiments related to enhancing cancer immunotherapy, the nanoparticle or the hydrogel (i.e., a nanoparticle-loaded hydrogel) can be co-administered with an immune checkpoint inhibitor, an immunomodulatory molecule, or a combination thereof. In certain instances, the immune checkpoint inhibitor is a checkpoint antibody (e.g., PD1, PDL1, CTLA4, OX40, etc.). In certain instances, the immune checkpoint inhibitor is an antibody that prevents interactions of CTLA4/(CD80/CD86) and PD1/PD-L1. In certain instances, the immunomodulatory molecule is a cytokine (e.g., IL2, IL 12, IL13, IL 15, etc.) or a chemokine.
The present inventors have engineered a modular PEG-PLA nanoparticle platform to overcome the limitations of TLR agonist therapies such as severe systemic toxicity and a narrow therapeutic window. The ability to tune the size of the PEG-PLA nanoparticles described herein allows efficient transport to lymph nodes which further reduces toxicity while enhancing immunostimulatory effects. In certain embodiments, the nanoparticles described herein are up to about 200 nm in diameter (e.g., about 20 to about 200 nm or about 20 to about 100 nm in diameter or about 30 nm, about 40 nm, about 50 nm, about 60 nm, about 70 nm, about 80 nm, about 90 nm, or about 100 nm in diameter), allowing for passive transport to lymph nodes, thus reducing systemic toxicity and enhancing interactions with target antigen presenting cells.
In certain aspects, the present disclosure provides a PEG-PLA nanoparticle that presents TLR agonist moieties on its surface. These nanoparticles are made by synthesizing PEG-PLA from N3-PEG-OH which was subsequently functionalized with alkyne derivatives of a TLR agonist. Any alkynated moiety can be conjugated to the polymers, and the Examples also describe conjugating mannose to increase opsonization of particles by immune cells. Optimally sized and biodegradable nanoparticles were then generated through nanoprecipitation of a combination of conjugated and nonconjugated PEG-PLA polymers. The TLR agonist or other conjugated moieties are then presented on the nanoparticle surface and are accessible to target receptors. The modular nature of this platform allows changing both the type(s) of molecule(s) presented on the surface and density of molecule presentation. As such, the present disclosure provides potent, non-toxic TLR agonist nanoparticle adjuvants suitable for use in cancer immunotherapies and vaccines.
In exemplary embodiments, the present disclosure provides a PEG-PLA nanoparticle with a TLR agonist attached to its surface having a controlled size (e.g., about 30 nm or about 50 nm), which allows the nanoparticle to be retained in the lymphatic system and kept out of the systemic circulation. This targeting leads to enhanced efficacy against tumors and reduced toxicity in a mouse model. The nanoparticles can be used on their own as anti-cancer therapy or as an adjunct “add on” therapy to any other immunotherapy (including checkpoint antibodies such as PD1, PDL1, CTLA4, and OX40, and/or immunomodulatory molecules such as cytokines and chemokines) to stimulate an immune response in cancer. The nanoparticles can also be used on their own or combined with controlled release hydrogels for vaccines or anti-cancer therapy (e.g., cancer immunotherapy). Non-limiting examples of hydrogels for controlled release of immunomodulatory compounds are described in International Publication No. WO 2020/072495, the disclosure of which is incorporated herein by reference in its entirety. The present disclosure also provides methods of making the nanoparticles described herein.
In other aspects, the present disclosure provides an injectable and self healing polymer nanoparticle hydrogel platform to prolong the co-delivery of vaccine components to the immune system. In certain embodiments, the nanoparticle hydrogel platform provides an enhanced magnitude and duration of germinal center responses in the lymph nodes by creation of a local inflammatory niche with the hydrogel, coupled with sustained release of vaccine cargo. The present disclosure introduces a simple and effective vaccine delivery platform that increases the potency and durability of subunit vaccines.
The present disclosure will be described in greater detail by way of specific examples. The following examples are offered for illustrative purposes only, and are not intended to limit the disclosure in any manner. Those of skill in the art will readily recognize a variety of noncritical parameters which can be changed or modified to yield essentially the same results.
Cancer immunotherapy can be augmented with toll-like receptor agonists (TLRas), which interact with immune cells to elicit the appropriate downstream immune response. Use of these drugs is limited due to their extreme potency, and lack of pharmacokinetic control, causing systemic toxicity from unregulated systemic cytokine release. Herein, we overcome these shortcomings by generating PEG-PLA nanoparticles (NPs) that present potent TLR7/8 as on their surface. We hypothesize that the pharmacokinetic profile of the NPs minimizes systemic toxicity, localizing the TLR7/8a presentation to the tumor bed and tumor draining lymph nodes. The nanoparticle platform allows precise control of TLR7/8a valency and the resulting surface presentation. In conjunction with anti-programmed death-ligand 1 (anti-PD-L1) checkpoint blockade, peritumoral injection of TLR7/8a NPs slows tumor growth, extends survival, and decreases systemic toxicity in comparison to the free drug in a murine colon adenocarcinoma model. These NPs constitute a modular platform for controlling pharmacokinetics of toxic immune-stimulatory drugs, resulting in increased potency and decreased toxicity.
Cancer immunotherapies, such as therapeutic immune checkpoint antibodies, are being developed for clinical use at increasing rates. The most widely used cancer immunotherapies are antibodies that prevent interactions of CTLA4/(CD80/CD86) and PD1/PD-L1. Though anti-CTLA4 and anti-PD1/PD-L1 therapies have shown great efficacy in some cancers, the overall response rates are highly variable mainly due to tumor cell evasion mechanisms.1 Supplementing PD-L1 checkpoint blocking antibodies with toll-like receptor agonists (TLRas) and other innate activators like STING agonists has shown great promise towards overcoming resistance mechanisms that cause low response rates.2,3,4,5. TLR7/8a are potent mimics of ssRNA that can elicit powerful immune responses that have been shown to synergize with immune checkpoint therapies.6 Unfortunately, the applicability of TLR7/8a in cancer immunotherapy is currently limited to skin cancers and metastatic cancers presenting on the skin, as systemic distribution results in a severe systemic response.7,8,9, Pharmacokinetic control of these compounds is crucial for their translation into the clinic, emphasizing the need for optimized drug delivery approaches.
TLR7/8a primarily activate pathways in dendritic cells (DCs) by mimicking single stranded nucleic acids which are the natural ligands. Activation of TLR7/8 by ssRNA mimics (TLR7/8a) boosts antigen presentation by DCs (and macrophages) through downstream signaling and cytokine production as part of the larger immune response. Synergy with PD-L1 blockade results from co-administration with TLR7/8a because TLR7/8 as aid lymph node- and tumor-resident dendritic cells in priming naïve T-cells towards tumor antigens, resulting in a tumoricidal behavior that can then be prolonged by the addition of PD-L1 blockade.10,11 Immunosuppressive tumors experience low levels of T cell priming, which renders PD-L1 checkpoint blocking ineffective. The addition of stimulatory molecules like TLR7/8a can initiate DC activation and kick-start the downstream T-cell response.
Nanocarriers presenting covalently bound TLR7/8a can activate TLR7/8 receptors.12 This is due to the TLR7/8 presenting its ligand binding site on the endosomal lumen of DCs and macrophages. Immunogenic constructs presenting TLR7/8a are thus possible, where the pharmacokinetic properties are strictly dictated by the carrier, with the construct itself being immunogenic without release of the TLR7/8a. Analogs of Resiquimod (R848) are popular TLR7/8a for these purposes, with several amine-functionalized analogs developed, some of which show higher potency and vastly different PK/PD profiles, with exacerbated toxicity.13 Multiple approaches have been effective in modulating the pharmacokinetics of these analogues. Macromolecular constructs presenting TLR7/8a have been explored for (anti-cancer) vaccine development by Lynn et al. They found that the morphology of the constructed nanoparticles had a significant impact on both T cell induction, and antibody production against a co-presented antigen.14,15 Nuhn et al. have likewise shown that pH responsive polymeric micelles with conjugated TLR7/8a is effective in eliciting a non-toxic, local immune response that synergizes with PD-L1 antibody antagonists.16,17 Although these macromolecular constructs presenting TLRas often have lower in vitro potency compared to unbound TLRas, they have better controlled tissue distribution and pharmacokinetics.
Altering the biodistribution of TLR7/8a also alters the number and biodistribution of different cytokines as well, as it impacts which immune cells that are exposed to TLR7/8 agonism, consequently, modulating the pharmacokinetics results in vastly different pharmacodynamics of the TLRas. Based on these advances, we hypothesize that PEG-PLA NPs can exert similar pharmacokinetic modulation of TLRas, as depicted in
As the basis of a modular PEG-PLA NP platform, we synthesized azide terminated PEG-PLA from N3-PEG-OH. The polymers were subsequently functionalized with alkyne derivatives of a TLR7/8a or mannose (
#Two-stage step-up method of Benjamini, Krieger and Yekutieli on Prism
&Multiple comparisons with Bonferroni method on Prism
These results demonstrate that tethering TLR7/8a to NPs slightly decreases the potency of the molecule in vitro. Previously tested TLR7/8a NP delivery systems typically lead to decreased activation in cell assays compared to the soluble form even if NP delivery is more potent in vivo.16 In this case the difference in potency and maximum activation across densities may be due to the pattern of TLR7/8a presentation which can lead to receptor clustering that is necessary for the downstream response.20
Recent studies of TLR7/8a delivery have shown that mannose can increase NP recognition and internalization by mannose-binding C-type lectins.18 A low and high mannose dose were incorporated into the medium valency TLR7/8a NPs and the same RAW-Blue assay was conducted to quantify TLR activation. We found that the presence of mannose shifted EC50 values increasing activation by approximately 2-fold for the low mannose group and by slightly less for the high mannose group. Both mannose doses, however, led to a modest decrease in maximum activation compared to the soluble TLR7/8a group (
In vitro cell assays do not take into account biodistribution and uptake by a broad array of cells within an organism, so it is important to screen TLR7/8a density and mannose presentation in vivo. To determine in vivo immune activation, serum IFNα was quantified over time following intraperitoneal (IP) administration of TLR7/8a NPs or soluble TLR7/8a in C57BL/6 mice. IFNα is a critical cytokine produced in response to TLR7/8 activation that contributes to the anti-tumor response.2′ Soluble TLR7/8a treatment led to an early spike (3 hours) in serum IFNα followed by a rapid decrease (
MC38 cells were injected subcutaneously (SC) on the right flank of mice and series of 4 treatment doses began once tumors were measurable (
Over this period, mice that received TLR7/8a either in the soluble form or the NP form had tumors that were significantly smaller than control mouse tumors (
Overall survival was also compared between treatment groups. Neither aPD-L1 nor aPD-L1 with soluble TLR7/8a significantly extended survival compared to the control (
The primary limitation of TLR7/8a therapy is the extreme systemic toxicity.25 The goal of this work was to both increase potency and decrease toxicity of TLR7/8a by tethering it to PEG-PLA NPs to localize its effect. Toxicity was assessed by measuring mouse body mass over the course of treatment and by running Luminex analysis on mouse serum that was collected 2 hours after the initial treatment. On the final treatment day mice that received the TLR7/8a NP treatment had a significantly higher average mass than the control mice (
Luminex analysis of serum cytokine levels after the first treatment showed an overall reduction in systemic cytokines in mice that received the TLR7/8a NP treatment as compared to the soluble TLR7/8a treatment (
We hypothesized that presentation of TLR7/8a on the surface of PEG-PLA NPs would promote draining to lymph nodes and would restrict immune activity to lymph nodes and the tumor environment. TLR7/8a on its own is a small molecule that will rapidly enter circulation while the NPs are in the size regime that has been shown to drain passively to lymph nodes.27 As expected, Luminex results show that mice that received soluble TLR7/8a as compared to TLR7/8a NPs generally had higher levels of serum cytokines across four main classes of cytokines that were quantified: cytokines that inhibit growth and activation; cytokines that promote growth, activation, and differentiation; chemokines; and proinflammatory cytokines (
TLR7/8 as presented on the surface of PEG-PLA nanoparticles were shown to retain their agonism in vitro, with mannose functionalized particles showing minimal increase in potency. In vivo, the TLR7/8 presenting particles led to prolonged, elevated levels of type I IFN compared to the soluble form. In a murine cancer model, the nanoparticles were shown to effectively synergize with PD-L1 checkpoint blockade to slow tumor growth and extend survival while reducing systemic cytokine release, suggesting lower toxicity. This study supports that macromolecular presentation of TLR7/8 agonists can overcome current toxicity limitations of systemic TLR7/8a delivery and is a viable complement to PD-L 1 checkpoint therapy.
Reagents. All chemicals were purchased through Sigma-Millipore, unless stated otherwise. Synthesis of alkynated mannose, alkynated TLR7/8a, and PEGPLA block copolymers is described in the supplementary information.
Materials characterization. NMR was obtained using an Inova 300 MHz NMR spectrometer with a Varian Inova console using VNMRJ 4.2 A software. 1H NMR data for IV, C, III, and D is shown in
Nanoparticle synthesis and characterization. NPs were prepared as previously reported.33 A 1 mL solution of a combination of PEG-PLA, TLR7/8a-PEG-PLA, and Mannose-PEG-PLA (depending on the experiment as shown in Table 2) in acetonitrile (50 mg/mL) was added dropwise to 10 mL of water under a high stir rate (600 rpm). NPs were purified by ultracentrifugation over a filter (molecular weight cut-off of 10 kDa; Millipore Amicon Ultra-15) followed by resuspension in water to a final concentration of 200 mg/ml. NPs were characterized by dynamic light scattering (DLS) to determine the NP diameters, and zeta potential for the NPs (Tables 3 and 4).
In vitro RAW-Blue reporter assay. The RAW-Blue reporter cell line (InvivoGen, raw-sp) was used in this study. Cells were cultured at 37° C. with 5% CO2 in Dulbecco's modified Eagle's medium (DMEM; Thermo Fisher Scientific) supplemented with L-glutamine (2 mM), D-glucose (4.5 g/L), 10% heat inactivated fetal bovine serum (Atlanta Biologicals), and penicillin (100 U/mL)/streptomycin (100 μg) and zeocin (100 μg/mL; Invivogen). Serial dilutions of soluble TLR7/8a or one of the TLR7/8a NP formulations (20 μL) was added to a 96-well tissue culture treated plate to final concentrations ranging from 0.08-10 μg/mL. About 100,000 cells were added to each well in 180 μL of media. Cells were cultured for 24 h at 37° C. in a CO2 incubator before following manufacturer instructions for SEAP quantification (Absorbance at 655 nm). Fits were generated using the “log(agonist) vs. response—Find EC50” in GraphPad Prism with the lower bound constrained to a constant value (0.22) for all fits.
Animal studies. 8-10 week old female C57BL/6 mice were obtained from Charles River and were cared for according to Institutional Animal Care and Use guidelines. Animal studies were performed in accordance with the guidelines for the care and use of laboratory animals; all protocols were approved by the Stanford Institutional Animal Care and Use Committee.
In vivo serum IFNα quantification. Mice were injected with buffer (200 μL) containing soluble TLR7/8a or one of the TLR7/8a NP formulations (25 μg TLR7/8a dose). Mice were injected IP since this administration route resulted in quantifiable cytokine levels across treatment groups. Serum was collected at the indicated times by tail vein blood collection and stored at −80° C. Serum IFNα concentrations were determined by ELISA according to the manufacturer's instructions (PBL Assay Science). Absorbance was measured at 450 nm in a Synergy H1 Microplate Reader (BioTek). Cytokine concentrations were calculated from the standard curves and represented as ng/mL).
MC38 tumor inoculation and treatments. The MC38 colon carcinoma cell line was purchased from Kerafast and cultured using DMEM (Thermo Fisher Scientific) supplemented with L-glutamine (2 mM), 0.1 mM nonessential amino acids, 10% heat inactivated fetal bovine serum (Atlanta Biologicals), penicillin (100 U/mL)/streptomycin (100 μg), and 10 mM Hepes (Sigma-Aldrich). 5×105 MC38 cells suspended in 100 μL of PBS were injected subcutaneously on the right side of the back of C57BL/6 mice. Mice were injected IP on days 8, 10, 12, and 15 post-inoculation with either PBS or 100 μg rat monoclonal anti-mouse aPD-L1 antibody (clone 10F.9G2; Bio X Cell). At the same time as the IP injections, mice were injected subcutaneously (SC) with 50 μL soluble TLR7/8a or high valency TLR7/8a NPs. The soluble TLR7/8a treatment also included PEG-PLA NPs without any conjugated TLR7/8a to account for any effects of the polymer NPs themselves. Tumor growth was monitored by measuring tumors with digital calipers (Mitutoyo Digimatic Caliper) 3 days a week. Tumor area was calculated from a length and width measurement (Area=length×width). Mice were euthanized when tumor volume exceeded 150 mm2. On day 50, the tumor-free survivor was re-challenged by SC injection on the right flank with 5×105 MC38 cells and tumor area monitoring continued.
Statistical analysis. Statistical analysis in
MC38 tumor growth and survival statistical analysis. Mice were assigned randomly to 4 treatment groups (i) no treatment, (ii) aPD-L1, (iii) aPD-L1+soluble TLR7/8a, and (iv) aPD-L1+TLR7/8a NPs. For statistical analysis, tumor area required additional transformation using the natural logarithm to meet the assumptions of homoscedasticity. Analysis was performed in JMP Pro 14. To test if tumor growth differed between treatments, we used a restricted maximum likelihood (REML) mixed model. Mouse was included as a random effect subject. The interaction between treatment and time tested whether treatment altered tumor growth over time. Post-hoc pair-wise comparisons were done between treatment groups and a Bonferroni correction was used to adjust for multiple comparisons (alpha=0.008). P-values for pair-wise comparisons (alpha=0.008): aPD-L1 vs. control, P=0.0001; TLR7/8a vs. control, P=0.0001; NP TLR7/8a vs. control, P<0.0001; a PD-L1 vs. soluble TLR7/8a P=0.7263; NP TLR7/8a vs. a PD-L1, P<0.0001; NP TLR7/8a vs. soluble TLR7/8a, P<0.0001. Survival statistical analysis was performed in SAS Version 9.4. To test if survival differed between treatments, we used a maximum likelihood parametric regression with censored data. Least-squared means were used to compare survival time between individual treatments and Tukey-Kramer post-hoc tests were used to correct for multiple comparison.
Body mass measurements. Mouse body mass was monitored every other day for the first 10 days following the start of treatment using a digital kitchen scale with 0.1 g resolution.
Luminex. 2 hours after the first treatment blood samples were collected by tail vein bleeds. Blood was collected in serum centrifuge tubes (Sarstedt), incubated at RT for 30-60 minutes, and was spun at 10,000 RCF for 5 minutes. Serum was collected and kept frozen at −80° C. until use for Luminex analysis. Mouse 38-plex Procarta kits were purchased from eBiosciences/Affymetrix/Thermo Fisher, Santa Clara, Calif. USA, and used according to the manufacturer's recommendations with modifications as described. Briefly, beads were added to a 96 well plate and washed in a Biotek ELx405 washer. Samples were added to the plate containing the mixed antibody-linked beads and incubated at room temperature for 1 hour followed by overnight incubation at 4° C. with shaking. Cold (4° C.) and room temperature incubation steps were performed on an orbital shaker at 500-600 rpm. Following the overnight incubation, plates were washed in a Biotek ELx405 washer and then biotinylated detection antibody added for 75 minutes at room temperature with shaking. The plate was washed as above, and streptavidin-PE was added. After incubation for 30 minutes at room temperature a wash was performed as above and reading buffer was added to the wells. Each sample was measured in duplicate. Plates were read using a Luminex 200 with a lower bound of 50 beads per sample per cytokine. Custom Assay Chex control beads were purchased from Radix Biosolutions, Georgetown, Tex., and are added to all wells.
Synthesis. The Benzyl amine TLR 7/8 agonist was synthesized as described by Shukla et. al.34, with a modification of the final aromatic substitution and tert butyl carbamate removal on tert-butyl (4-((4-chloro-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-1-yl)methyl)benzyl)carbamate (I), which was done as a one-pot Staudinger-type reaction.35
1 and 2 equivalents of NaN3 (68 mg, 1.04 mmol) was dissolved in 1.5 mL DMSO, and was heated to 90° C. for 2 hours. 2 equivalents of Triphenylphosphine (273 mg, 1.04 mmol) was added, and the temperature was increased to 120° C. and stirred for 16 hours. 0.5 mL 4M HCl (aq) was added and the temperature was lowered to 95° C. for 3 hours. 4 mL of water was added at which a precipitate formed, and the mixture was washed with EtOAc. Na2CO3 (sat) was added to the aqueous phase which was extracted with EtOAc, which was dried with with Na2SO4 and the solvent removed in vacuo to yield a crude solid. This was purified by silica column chromatography on a biotage system, using a gradient of DCM:MeOH, with 1% trimethylamine in the MeOH. This yielded II (1-(4-(aminomethyl)benzyl)-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-4-amine) (60 mg, 0.16 mmol, 32% yield)
II (25 mg, 69 μmol) and propargyl-N-hydroxysuccinimidyl ester (17 mg, 76 μmol) were dissolved in 1 mL anhydrous DCM, and stirred for 2 hours at room temperature. The reaction mixture was loaded onto a silica column, and the product was purified using a DCM:MeOH gradient. This yielded III (
1H NMR (300 MHz, CDCl33-d) δ 8.05 (d, J=8.4 Hz, 1H, H1), 7.73 (d, J=8.3 Hz, 1H, H2), 7.58 (t, J=7.9 Hz, 1H, H3), 7.29 (3H, H4, H5), 7.02 (d, J=7.9 Hz, 2H, H6), 6.43 (s, 1H, H7), 5.91 (s, 2H, H8), 4.78 (s, 2, H9), 4.46 (d, J=5.9 Hz, 2H, H10), 4.15 (d, J=2.4 Hz, 2H, H11), 3.82 (t, J=5.7 Hz, 2H, H12), 3.58 (q, J=7.2 Hz, 2H, H13), 2.55 (t, J=5.7 Hz, 2H, H14), 2.36 (t, J=2.4 Hz, 1H, H15), 1.14 (t, J=7.2 Hz, 3H, H16).
13C NMR (75 MHz, CDCl3) δ 171.35, 166.55, 151.77, 150.14, 138.70, 136.66, 136.09, 133.40, 129.84, 128.43, 125.71, 124.66, 121.24, 120.73, 112.79, 79.06, 75.01, 66.77, 66.02, 64.19, 58.44, 58.38, 58.21, 36.88, 31.89, 14.86. HRMS (ESI) calcd. for C27H29N5O3+H+: 472.2343; found 472.2335.
1′-O-propargyl-α-Mannose, IV. Alkynated mannose was synthesized as described elsewhere.36 α-Mannose pentaacetate (1.95 g, 5 mmol) and propargyl alcohol (0.34 g, 0.35 mL, 6.3 mmol) were dissolved in dry DCM (44 mL) while stirring. BF3Et2O (9.3 uL, 8 mg, 0.06 mmol) was added at 0° C., and the mixture was stirred for 16 h. The solution was washed with 1M NaOH, and EtOAc was added, and the organic phase was washed with brine and dried over Na2SO4. The solution was reduced to approximately 5 mL in vacuo, and diluted with 5 mL hexanes, and purified by silica chromatography with a EtOAc:Hexanes gradient. The product was dissolved in dry MeOH, and 100 μL 2% NaOMe in MeOH was added. The solution was left for one hour, quenched with a drop of acetic acid, and volatiles removed in vacuo to yield IV. 1H NMR spectrum is shown in
N3-PEG-PLA, MeO-PEG-PLA: PEG-PLA was prepared as previously reported.37 In short, 0.500 g N3-PEG(114)-OH or MeO-PEG(114)-OH in dry DCM (2 mL) with equimolar 1,8-Diazabicyclo[5.4.0]undec-7-ene (15 μL, 15 mg, 0.1 mmol) was rapidly added to a solution of 2.0 g D,L-lactide in 8 mL dry DCM. The mixture was stirred for 8 min, and quenched with the addition of 1 mL acetone with 200 μL acetic acid. The solution was precipitated into 35 mL of a 1:1 mixture of hexanes and diethyl ether in a 50 mL centrifuge tube. The supernatant was discarded, and 35 mL diethyl ether was added to fully precipitate the polymer. This was dissolved in a minimal amount of ethyl acetate, and 35 mL diethyl ether was added to re-precipitate the polymer.
TLR7/8a-PEG-PLA, Mannose-PEG-PLA: Conjugations were performed analogous to the procedure previously reported.37 A 20 mL scintillation vial was charged with III or IV (1.5 eq) and azido poly(ethylene oxide)-b-poly(D,L-lactide) 5 kDa-20 kDa (0.5 g, 20 umol) was dissolved in 4 mL of NMP and sparged with nitrogen for 10 min. 0.1 mL of a degassed CuBr (3.7 mg/mL) and THPTA (16 mg/mL) was added. The reaction mixture was further sparged with nitrogen for 10 min. The reaction mixture is incubated for 16 h at RT, and precipitated into diethyl ether in a 50 mL centrifuge tube to recover the polymer. The polymer was then dissolved in ethyl acetate and precipitated into diethyl ether and dried in vacuo. Synthetic schemes and 1H NMR data are shown in
Vaccines aim to elicit a robust, yet targeted, immune response. Failure of a vaccine to elicit such a response arises in part from inappropriate temporal control over antigen and adjuvant presentation to the immune system. In this work, we sought to exploit the immune system's natural response to extended pathogen exposure during infection by designing an easily administered slow-delivery vaccine platform. We utilized an injectable and self-healing polymer-nanoparticle (PNP) hydrogel platform to prolong the co-delivery of vaccine components to the immune system. We demonstrated that these hydrogels exhibit unique delivery characteristics whereby physicochemically distinct compounds (such as antigen and adjuvant) could be co-delivered over the course of weeks. When administered in mice, hydrogel-based sustained vaccine exposure enhanced the magnitude, duration, and quality of the humoral immune response compared to standard PBS bolus administration of the same model vaccine. We report that the creation of a local inflammatory niche within the hydrogel, coupled with sustained release of vaccine cargo, enhanced the magnitude and duration of germinal center responses in the lymph nodes. This strengthened germinal center response promoted greater antibody affinity maturation, resulting in a more than 1000-fold increase in antigen-specific antibody affinity in comparison to bolus immunization. Furthermore, our technology improved the efficacy of an influenza hemagglutinin subunit vaccine compared to the most potent adjuvant system used clinically for influenza vaccination, by increasing antibody titers and cross-reactivity against other hemagglutinin variants. In summary, this work introduces a simple and effective vaccine delivery platform that increases the potency and durability of subunit vaccines.
The challenges in designing vaccines towards rapidly mutating pathogens, such as Influenza and HIV, are complex and highly interdisciplinary. To elicit a protective antibody response, vaccines must interact with the appropriate cell types at the right time and place while also providing the necessary cues to guide the immune system1. With regards to temporal dynamics, natural infections expose the immune system to antigen and inflammatory signals for 1-2 weeks1. Conversely, the short-term presentation of subunit vaccines from a single bolus administration persists for only 1-2 days1. The sustained release of soluble antigen provides essential signals to germinal centers (GCs) in the lymph nodes—the structures responsible for the affinity maturation of B cells-leading to high affinity antibody production and humoral immune memory formation2-4. Recent work demonstrates that slow delivery of an HIV vaccine using osmotic pumps in nonhuman primates resulted in higher neutralizing titers5. Similarly, sustained release of a stabilized HIV antigen from a microneedle patch led to increased serum titers and bone marrow plasma cells in mice6,7. It has been reported that the mechanism of traditional adjuvant systems such as aluminum hydroxide (Alum) and MF59 rely on increased antigen persistence in the body8-10. These studies, along with others, demonstrate that the kinetics of antigen presentation to the immune system dramatically influences the adaptive immune responses5,6,8,9,11-22. However, these technologies are limited by several factors: (i) the inability to deliver chemically diverse cargo, (ii) restricted time-frame tunability for prolonged release, and/or (iii) low yielding and, oftentimes, challenging multi-step syntheses-. Therefore, there is a significant need for the development of new materials that allow for the sustained exposure of vaccine components to the immune system.
Hydrogel materials are well-suited for drug delivery applications due to their high water content and mechanical tunability that make them similar to many biological tissues24. However, traditional covalently cross-linked hydrogels are often not appropriate for vaccine applications as they cannot be easily administered and have limitations in their delivery kinetics for diverse cargo24. Polymer-nanoparticle (PNP) hydrogels are a type of supramolecular hydrogel where the polymeric constituents are held together by dynamic, multivalent non-covalent interactions between polymers and nanoparticles25-28. These materials exhibit many favorable characteristics such as high drug loading capacity, gentle conditions for encapsulation of biologic cargo, sustained delivery of cargo, and mechanical tunability24,28,29. In addition, PNP hydrogels are easily administered due to their shear thinning and self-healing properties27,30,31. The fabrication process, which is easily scaled and therefore highly translatable, involves straightforward mixing of the polymer, nanoparticles (NPs), and an aqueous solution of cargo28,32. This combination of unique properties makes PNP hydrogels ideal candidates for use as a vaccine delivery platform.
Here, we report the use of PNP hydrogels as a vaccine delivery platform that affords simple encapsulation of vaccine components, provides sustained co-delivery of physicochemically distinct vaccine cargo, and creates a transient inflammatory niche upon administration for activation of the immune system. We show that these materials are easily injected and initiate the vaccine response locally through recruitment of antigen presenting cells (APCs), while also providing sustained release of vaccine cargo (
We designed a PNP hydrogel material that can load vaccine components with high efficiency, is injectable, and can be tuned to co-deliver subunit vaccine components over prolonged timeframes. Our PNP hydrogels form rapidly when aqueous solutions of hydroxypropyl methylcellulose derivatives (HPMC-C12) are mixed with biodegradable polymeric NPs composed of poly(ethylene glycol)-b-poly(lactic acid) (PEG-PLA) (
The rheological properties of the hydrogel formulations of interest were then measured. Frequency-dependent oscillatory shear experiments, performed in the linear viscoelastic regime, demonstrated a formulation-dependent frequency response (
Both gel formulations were shown to be injectable, a critical consideration for future translation of our material as a vaccine platform. Injectability was tested by measuring the viscosity of the gels when alternating between a high shear rate (100 s−1) reminiscent of the injection process, and a low shear rate (0.5 s−1) reminiscent of the working conditions upon implantation26. The viscosity of both gel formulations decreased by over two orders of magnitude with the applied high shear rate, and rapidly recovered (<5 s) to their original viscosity when returned to a low shear rate (
Vaccine Cargo Dynamics
Subunit vaccines are composed of two main components: (i) antigen, which directs the antibody response to a specific substance, and (ii) adjuvant, which enhances the innate immune response34. These biomolecules can have vastly different sizes and physicochemical properties34, providing a challenge for controlled co-delivery because these characteristics typically dictate the diffusivity of these compounds and their release kinetics from hydrogel materials in vitro and in vivo35. To understand how PNP hydrogels can be used to deliver a wide range of cargo, we chose a model vaccine that incorporated an antigen and adjuvant with distinct physicochemical properties and very different molecular weights. We used the protein ovalbumin (OVA; MW=43 kDa) as a model antigen and Poly(I:C) (a toll-like receptor 3 agonist (TLR3); MW>1 MDa)36,37 as the adjuvant. Poly(I:C), a double-stranded RNA mimic that activates endosomal TLR3, is a potent type I interferon producer and has been used successfully as an adjuvant in preclinical vaccine studies, though it has not been approved for clinical use to date37-39. Importantly, this highly tunable materials platform enables these components to be easily replaced with other antigens and adjuvants for future applications.
To characterize the dynamics of vaccine exposure expected from the vaccine-loaded gels, we assessed the diffusivities of both vaccine cargo in each of the two gel formulations of interest using fluorescence recovery after photobleaching (FRAP) experiments (
Interestingly, the ratio of OVA to Poly(I:C) diffusivities (DOVA/DPoly(I:C)) was 3.4 in the 1:5 gel, and only 1.4 in the 2:10 gel (
The in vivo cargo dynamics were characterized by assessing the timeframes of OVA retention in the subcutaneous (SC) space following in vivo administration. The 1:5 and 2:10 gels were loaded with Alexa Fluor 647 conjugated OVA and explanted to measure gel mass and cargo fluorescence at multiple time-points up to 4 weeks after vaccine administration. The 1:5 gels reached half of their initial mass after 2 weeks whereas the 2:10 gels retained half of their initial mass past 4 weeks (
To investigate if sustained vaccine exposure from a single administration of the gel-based vaccines (i.e., gel carrying OVA and Poly(I:C)) could enhance the magnitude and duration of the humoral response, we quantified antigen specific antibodies in the serum after SC injection (
Most vaccines are administered with booster shots to increase the immune memory and ensure long-term protective antibodies are produced34. To study the prime-boost response, we administered OVA and Poly(I:C) in PBS for all groups at 40 or 90 days post-prime in separate mouse cohorts (
To investigate the quality of the humoral response, we also determined the affinities of the serum antibodies after the 90-day boost towards OVA using a competitive binding enzyme-linked immunosorbent assay (ELISA) (
Local Inflammatory Niche within the Hydrogel Depot
We evaluated the cells infiltrating the gel depots to understand if the gel was creating an inflammatory niche in addition to providing sustained release of the cargo (
We hypothesized that the enhanced humoral response observed for gel-based sustained-exposure vaccines was due to the prolonged lifetime of GCs in the lymph nodes. GCs are dynamic sites that form after the activation of germinal center B cells (GCBCs) and are responsible for producing memory B cells and high-affinity antibodies34. The GCs of the 2:10 and bolus groups 15 days after vaccination with OVA and Poly(I:C) were qualitatively visualized with immunohistochemistry (
To demonstrate that our gel platform can be optimized for real viral antigens, we used the influenza antigen hemagglutinin (HA), which is the most commonly used antigen in influenza subunit vaccines. In these studies, we used a TLR 7/8 agonist adjuvant because it has been previously shown to elicit strong titers against HA and has demonstrated promise for clinical translation48. To ensure sustained co-administration of HA and the TLR 7/8 agonist, which is a small molecule (314 Da), we tethered the TLR 7/8 agonist to the PEG-PLA NPs that form the PNP hydrogel network together with HPMC-C12 (
Next, to interrogate our material's ability to improve humoral immunity towards the clinically relevant HA antigen, we quantified the HA-specific IgG antibody titers after single administration of HA delivered in the TLR 7/8 NP-based 2:10 gel compared with clinically relevant adjuvants MF59, which is the most potent influenza vaccine adjuvant used clinically, and Alum. We showed that the TLR 7/8 NP-based gels led to 3.7-fold and 100-fold higher antibody titers compared to MF59 and Alum, respectively. Additionally, mice receiving the gel-based vaccine maintained antibody titers above the MF59 group's peak titer for over 140 days (
To evaluate cross reactivity to HA variants not included in the vaccine formulation, total IgG titers against HA from A/California/07/2009(H1N1) (Cal09) and A/Michigan/45/2015(H1N1) (Mich15) were quantified (
The diagrams and diffusivity values in
Germinal centers (GCs) are sites within lymphoid organs where mature B cells undergo somatic hypermutation (SHM) leading to higher affinity antibodies. Within the GCs, mature B cells go through cycles of proliferation, mutation, and selection in order to create B cells with the highest affinity B cell receptors (BCRs). The GC is divided into two anatomical compartments: the light zone (LZ) and the dark zone (DZ). In the LZ, GC B cells compete to capture antigen from follicular dendritic cells (FDCs) and present antigen peptides on MHCII2. The GC B cells with the highest affinity BCRs have increased antigen presentation to the T follicular helper (Tfh) cells and therefore will receive more positive signaling from Tfh cells. Tfh cell signaling is critical for positive selection of B cells with BCRs which have high affinity for antigen. The B cells that successfully interact with Tfh either mature into plasma cells or return the DZ for another round of proliferation and additional SHM57,58. For this reason, prolonged Tfh cell presence in GCs is indicative of prolonged SHM. Similarly, a shift towards LZ B cell markers indicates an increase in the process of affinity selection compared to the processes of expansion and SHM which is necessary to continuing B cell cycling through the GC4.
The polymer-nanoparticle (PNP) gels described herein are formed by simple mixing of two aqueous solutions comprising (i) “polymer” (HPMC-C12) and (ii) “nanoparticles” (PEG-PLA NPs). Upon mixing, a hydrogel rapidly forms as multivalent and dynamic non-covalent interactions between the HPMC-C12 polymer and the PEG-PLA NPs generate physical cross-links that give rise to a polymer network structure. This hydrogel network structure provides the sustained vaccine exposure properties described herein. These PNP hydrogels act as a simple platform for encapsulation of a diverse array of vaccine components of interest and enhance humoral immune responses by altering the timeframe of vaccine presentation to the immune system. Our study focuses on the material platform's ability to manipulate the humoral immune response, regardless of the particular vaccine, by first using a model vaccine consisting of ovalbumin (OVA) and Poly(I:C) and then using an influenza vaccine consisting of hemagglutinin (HA) and an R848 derivative. The studies with HA immunization demonstrate the ability of the PNP hydrogel to be used for slow delivery of a “real” viral antigen and reveals the broad applicability of our delivery platform as the molecular weights and physicochemical properties of OVA, Poly(I:C), R848, and HA are very diverse. It is important to note that an optimized and highly effective vaccine requires three parts: (i) advanced antigen design, (ii) potent adjuvants, and (iii) appropriate delivery characteristics.
Several adjuvant technologies have been used in commercial influenza vaccines, including MF59, Al(OH)3, and AS03. The only adjuvanted seasonal flu vaccine is FLUAD® (Seqirus Inc), which contains MF59 and is used for persons 65 years of age and older60. The remaining licensed adjuvanted flu vaccines all target the H5N1 pandemic influenza strain. These formulations include MF59 in AUDENZ (Seqirus), Al(OH)3 in Panflu (Sinovac), and AS03 in Prepandrix (GSK) and Q-Pan H5N1 (GSK)60. Yet, studies indicate that adjuvants alone may not provide sufficient immunological driving forces for promoting extensive affinity maturation, suggesting the kinetics of vaccine exposure must be carefully considered. Recently, Crotty and coworkers demonstrated that slow release of an HIV vaccine using surgically implantable pump leads to more robust germinal center responses, higher antibody titers, and the targeting of a more diverse set of epitopes than standard bolus administration of the same vaccine61. Sustained exposure can therefore enhance neutralizing antibody development by altering germinal responses and modulating the immunodominance of non-neutralizing epitopes.
For these reasons, we do not argue that we have designed an optimized influenza vaccine as we use a simple, commercially available HA antigen and compare our sustained delivery hydrogel system to current clinically used adjuvants. Our study indicates the ability to improve the humoral immune response to the HA viral antigen by choosing an appropriate adjuvant and using the PNP hydrogel platform to provide prolonged co-delivery of antigen and adjuvant.
In this study, we designed an injectable hydrogel material capable of prolonged co-delivery of physicochemically-distinct cargo on timescales relevant to the kinetics of antigen exposure typical of natural infections. The vaccine-loaded PNP hydrogels are injectable and retain their solid-like structure when under low stresses, enabling creation of a new stimulatory microenvironment within the body while facilitating sustained vaccine delivery. Additionally, the dynamic networks comprising the 1:5 and 2:10 gels provide unique cargo diffusion characteristics compared to traditional covalent hydrogel systems. Our observations suggest that the 2:10 gel mesh sufficiently constrains the diffusion of both OVA and Poly(I:C), even though Poly(I:C) is much larger than OVA, such that their diffusion and release is dictated by the self-diffusion of the dynamically cross-linked hydrogel network. These physical properties ensure co-presentation of antigen and adjuvant over prolonged timeframes to the immune system which has been a long-standing challenge within the field.
Developing effective humoral immune responses towards difficult pathogens, such as HIV-1 or malaria, requires the creation of high affinity antibodies49,50. The sustained vaccine exposure enabled by these gels was shown to increase the magnitude and duration of GC responses in the draining lymph nodes, leading to enhancements in the magnitude, persistence, and quality of the humoral immune response against the antigen. The sustained release of antigen from PNP hydrogels is able to better mimic natural infections where GCBCs continuously receive antigen-derived signaling to undergo many rounds of selection and somatic hypermutation (
These gels also act as a local stimulatory microenvironment where infiltrating cells experience high local concentrations of adjuvant and antigen. Implantable biomaterials have been shown to act as immune-stimulating niches which recruit and activate APCs that then migrate to the lymph nodes51. Herein, we observed minimal immune cell infiltration in the gels alone; however, loading the gels with vaccine cargo dramatically increased immune cell infiltration, which is likely driven by the inflammatory signals initiated by the entrapped adjuvant. The vaccine-loaded gel was able to recruit macrophages and DCs, which are professional APCs that are essential for initiating the adaptive immune response. In particular, almost all infiltrating DCs were found to be cDC2s which are known to play a critical role in activating Tfh cells and initiating the humoral immune response. We hypothesize that these cells are activated by entrapped adjuvant and migrate to the draining lymph nodes. The draining lymph nodes are therefore receiving activated immune cells from the gel as well as the soluble vaccine cargo as it is released from the gel depot over time (
In conclusion, PNP hydrogels provide a simple and effective platform for sustained delivery of subunit vaccines to increase the potency and durability of the humoral immune response. Using our platform as a tool to probe the interactions between the immune system and a vaccine depot will enable more precise material development for vaccine delivery. Our PNP hydrogel represents a highly tunable platform for effectively manipulating the humoral immune response for any subunit vaccine of interest.
Materials HPMC (meets USP testing specifications), N,N-Diisopropylethylamine (Hunig's base), hexanes, diethyl ether, N-methyl-2-pyrrolidone (NMP), dichloromethane (DCM), lactide (LA), 1-dodecylisocynate, and diazobicylcoundecene (DBU) were purchased from Sigma Aldrich and used as received. Monomethoxy-PEG (5 kDa) was purchased from Sigma Aldrich and was purified by azeotropic distillation with toluene prior to use.
Preparation of HPMC-C2: HPMC-C12 was prepared according to previously reported procedures28. HPMC (1.0 g) was dissolved in NMP (40 mL) by stirring at 80° C. for 1 h. Once the solution cooled to room temperature, 1-dodecylisocynate (105 mg, 0.5 mmol) and N,N-Diisopropylethylamine (catalyst, ˜3 drops) were dissolved in NMP (5.0 mL). This solution was added dropwise to the reaction mixture, which was then stirred at room temperature for 16 h. For rhodamine conjugated HPMC-C12, 10 mg of rhodamine B isothiocyanate (0.019 mmol) was added to the 1-dodecylisocyanate solution before adding it dropwise to the reaction mixture. This solution was then precipitated from acetone, decanted, re-dissolved in water (˜2 wt %), and placed in a dialysis tube for dialysis for 3-4 days. The polymer was lyophilized and reconstituted to a 60 mg/mL solution with sterile PBS.
Preparation of PEG-PLA NPs and TLR 7/8 NPs: PEG-PLA was prepared as previously reported28. Monomethoxy-PEG (5 kDa; 0.25 g, 4.1 mmol) and DBU (15 μL, 0.1 mmol; 1.4 mol % relative to LA) were dissolved in anhydrous dichloromethane (1.0 mL). LA (1.0 g, 6.9 mmol) was dissolved in anhydrous DCM (3.0 mL) with mild heating. The LA solution was added rapidly to the PEG/DBU solution and was allowed to stir for 10 min. The reaction mixture was quenched and precipitated by 1:1 hexane and ethyl ether solution. The synthesized PEG-PLA was collected and dried under vacuum. Gel permeation chromatography (GPC) was used to verify that the molecular weight and dispersity of polymers meet our quality control (QC) parameters.
Azide-PEG-PLA was prepared using N3-PEG-OH (0.5 g, 5 kDa, 100 μmol) in 2 mL of anhydrous DCM with 30 μL DBU (30 mg, 0.20 mmol) which was added quickly to a stirring solution of LA (2 g, 13.9 mmol) in 6 mL anhydrous DCM. The solution was stirred for 10 min, after which 2 drops of acetic acid was added, and the polymer was precipitated into a 1:1 mixture of hexanes and diethyl ether. The polymer was re-dissolved in a minimal amount of acetone, and precipitated again in diethyl ether, and dried in vacuo. GPC was used to verify that the molecular weight and dispersity of polymers meet our QC parameters.
A 20 mL scintillation vial was charged with the TLR 7/8 agonist alkyne (14 mg, 30 μmol) and azido poly(ethylene oxide)-b-poly(D,L-lactide) 5 kDa-20 kDa (0.5 g, 20 umol) was dissolved in 4 mL of NMP and sparged with nitrogen for 10 min. Next, 0.1 mL of a degassed CuBr (3.7 mg/mL) and THPTA (16 mg/mL) was added. The reaction mixture was further sparged with nitrogen gas for 10 min. The reaction mixture was incubated for 16 h at room temperature, and precipitated into diethyl ether in a 50 mL centrifuge tube to recover the polymer. The polymer was then dissolved in ethyl acetate and precipitated into diethyl ether, and dried in vacuo. GPC was used to verify that the molecular weight was not altered by conjugation.
NPs were prepared as previously reported28. A 1 mL solution of PEG-PLA in DMSO (50 mg/ml) or TLR 7/8-PEG-PLA in acetonitrile (50 mg/mL) was added dropwise to 10 mL of water at room temperature under a high stir rate (600 rpm). NPs were purified by ultracentrifugation over a filter (molecular weight cut-off of 10 kDa; Millipore Amicon Ultra-15) followed by resuspension in water to a final concentration of 200 mg/mL. NPs were characterized by dynamic light scattering (DLS) to find the NP diameters, 32±4 nm and 29±3 nm, and zeta potential, −28±7 mV and −10±7 mV, for the PEG-PLA and TLR 7/8-PEG-PLA NPs, respectively (Table 6).
PNP Hydrogel Preparation. The 2:10 formulation contained 2 wt % HPMC-C12 and 10 wt % PEG-PLA NPs in PBS. These gels were made by mixing a 2:3:1 weight ratio of 6 wt % HPMC-C12 polymer solution, 20 wt % NP solution, and PBS. The 1:5 formulation contained 1 wt % HPMC-C12 and 5 wt % PEG-PLA NPs in PBS. These gels were made by mixing a 2:3:7 weight ratio of 6 wt % HPMC-C12 polymer solution, 20 wt % NP solution, and PBS. The solutions were mixed with a spatula, mildly centrifuged to remove bubbles arising from mixing, and then loaded into a syringe.
Vaccine Formulations: The model vaccine contained a 100 μg dose of OVA (Sigma Aldrich) and 50 μg dose of Poly(I:C) (Sigma Aldrich) per 100 μL of gel or PBS. The influenza vaccine contained a 2 μg dose of Influenza A HIN1 (A/Brisbane/59/2007) hemagglutinin (HA)(Sino Biological) and an approximate TLR 7/8 agonist dose of 50 μg. For the bolus vaccines, the above vaccine concentrations were prepared in PBS and loaded into a syringe for administration. For the PNP hydrogels, the vaccine cargo was added at the appropriate concentration into the PBS component of the gel before adding the polymer and NP solutions, as described above. For MF59 (Invivogen) and Alum (Invivogen) vaccines, the formulations were prepared according to the manufacturer's instructions with a 2 μg dose of HA.
Material Characterization: Rheological characterization was performed using a TA Instruments Discovery HR-2 torque-controlled rheometer fitted with a Peltier stage. All measurements were performed using a serrated 20-mm plate geometry at 25° C.
Dynamic oscillatory frequency sweep measurements were performed with a constant torque (2 uN·m; σ=1.27 Pa) from 0.1 rad/s to 100 rad/s. Steady shear experiments were performed from 0.1 to 100 s−1. Step-shear experiments were performed by alternating between a low shear rate (0.05 s−1) and high shear rate (100 s−1) for two full cycles.
FRAP Analysis: Alexa Fluor 647 conjugated OVA (Thermo Fisher Scientific), rhodamine conjugated Poly(I:C) (Invivogen), and rhodamine conjugated HPMC-C12 were used to visualize the diffusion of the cargo and gel. Samples were photobleached with a 50 μm diameter for the region of interest (ROI). Different tests (n=5) were made for 3 different samples from the same batch at different locations of the sample. A spot was bleached with a pixel dwell time of 177.32 μs. 500 post-bleach frames were recorded at 1 frame/s to form the recovery exponential curve. The diffusion coefficient was calculated as52:
D=γD(ω2/4Tτ1/2)
Where the constant γD=τ1/2τD, with τ1/2 being the half-time of the recovery. τD the characteristic diffusion time, both yielded by the ZEN software, and ω the radius of the bleached ROI (25 μm).
The diffusivity of cargo in PBS was calculated using the Stokes-Einstein Law Equation for diffusion53 where kB is Boltzmann's constant, T is temperature in Kelvin, η is solvent viscosity, and R is solute hydrodynamic radius
The diffusivity of cargo in a model covalent PEG gel was calculated using the Multiscale Diffusion Model (MSDM) assuming 25° C., 5% volume fraction, and 35 nm mesh size54. The calculated values are comparable to experiment diffusivities of similar sized cargo found in the literature55.
Mice and Vaccination: C57BL/6 (B6) mice were purchased from Charles River and housed at Stanford University. Female mice between 6 and 10 weeks of age at the start of the experiment were used. The mice were shaved several days before vaccine administration and received a subcutaneous injection of 100 μL gel or bolus vaccine on their backs under brief isoflurane anesthesia. PBS injections used a 26-gauge needle, and gel injections used a 21-gauge needle. Mouse blood was collected from the cheek or tail vein for survival studies, or through cardiac puncture for terminal studies. The inguinal LN's and SC gels were collected for GC and cell infiltration analysis after euthanasia.
In Vivo Cargo Dynamics: Vaccine loaded gels with Alexa Fluor 647 conjugated OVA (Thermo Fisher Scientific), were explanted at 1, 7, 14, 21, or 28 days, and weighed to report the gel erosion over time. After weighing, the explanted gels were diluted in PBS and homogenized using a glass dounce homogenizer (Wheaton). The fluorescence of the homogenized gels was read with ex: 650 nm and em: 665 nm on a plate reader (Tecan Infinite M1000). Raw fluorescence values were normalized for polymer background, the natural log of these values was plotted and fit with linear equations to find the rate constant using GraphPad Prism 7.04 (GraphPad Software). This rate constant was used to calculate a half-life of cargo retention.
Antibody Concentration and Affinity: Serum antibody concentrations and affinity for the OVA model vaccine were measured using an anti-ovalbumin mouse IgG1 ELISA (Cayman Chemicals. 500830). For time course measurements the serum was diluted 1:1,000 in assay buffer and for post-challenge analysis the serum was diluted 1:100,000. The assay was performed according to the manufacturer's instructions to find concentration. The plates were analyzed using a Synergy™ H1 Microplate Reader (BioTek Instruments) at 450 nm. Serum antibody concentrations were calculated from the standard curves and represented as μg/mL or mg/mL.
Anti-OVA IgG2b and IgG2c antibody concentrations were measured using mouse anti-OVA antibody assay kits for IgG2b (chondrex, 3016) and IgG2c (Chondrex, 3029). Serum was diluted 1:1,000 and the assay was performed according to the manufacturer's instructions to find concentration. The plates were analyzed using a Synergy™ H1 Microplate Reader (BioTek Instruments) at 450 nm. Serum antibody concentrations were calculated from the standard curves and represented as μg/mL.
For affinity, dilutions of the serum were mixed with a constant 3 nM of an HRP conjugated anti-OVA antibody (BioLegend) and incubated for 2 h at room temperature. The wells were washed, incubated with TMB substrate (TMB ELISA Substrate (High Sensitivity). Abcam), and the reaction was stopped with 1 M HCl. Absorbance was read with a plate reader at 450 nm. Data was fit with a one-site competitive binding model with Graphpad Prism. The control antibody was assumed to have a KD of 1 nM based on common affinities of mAbs found in the industry. This assumption affects only the absolute KD values reported and not the relative differences between treatment groups. Statistics were performed on the log10 (KD) values. Individual binding curves can be found in
Serum IgG antibody titers for the influenza vaccine were measure using an ELISA. Ni-coated plates (Thermofisher) were coated with HA (Sino Biological) at 2.5 μg/mL in PBS for 1 h at 25° C. and blocked with PBS containing 1% BSA for 1 h at 25° C. A standard curve was created by pooling serum and completing serial dilutions (2×) before adding to the plate and serum samples were diluted 1:200 (Alum group) or 1:1,000 (Gel and MF59 groups) and added to plates. After 2 h at 25° C., goat-anti-mouse IgG Fc-HRP (1:10,000, Invitrogen, A16084) was added for 1 h at 25° C. Plates were developed with TMB substrate (TMB ELISA Substrate (High Sensitivity), Abcam). The reaction was stopped with 1 M HCl. The plates were analyzed using a Synergy™ H1 Microplate Reader (BioTek Instruments) at 450 nm. Serum antibody titers were calculated from the standard curve and represented as the dilution required to reach the detection limit.
Serum IgG1, IgG2b, and IgG2c antibody titers against A/Brisbane/59/2007 HA and IgG titers against A/California/07/2009 and A/Michigan/45/2015 HA were measure using an endpoint ELISA. Ni-coated plates (Thermofisher) were coated with HA (Sino Biological) at 2.5 μg/mL in PBS for 1 h at 25° C. and blocked with PBS containing 1% BSA for 1 h at 25° C. First, serum was diluted 1:250 and then four-fold serial dilutions were performed up to 1:4,096,000 dilution. Titrations were added to plates and after 2 h at 25° C., HRP-conjugated goat anti mouse IgG1 (Abcam, ab97240), IgG2b (Chondrex, 3016), IgG2C (Abcam, ab97255), or IgG (Invitrogen, A16084) were added at a 1:10,000 dilution for 1 h at 25° C. Plates were developed with TMB substrate (TMB ELISA Substrate (High Sensitivity), Abcam). The reaction was stopped with 1 M HCl. The plates were analyzed using a Synergy™ H1 Microplate Reader (BioTek Instruments) at 450 nm. Endpoint titers were defined as the reciprocal of the highest serum dilution that gave an optical density above 0.1.
Immunohistochemistry in LN: Draining inguinal lymph nodes were isolated and frozen in molds containing OCT medium on dry ice. Frozen lymph nodes were sectioned at 7 μm and stored at −80° C. Sections were stained with biotin labeled anti mouse IgD (eBioscience) and PE labeled anti mouse BCL6 (BD Biosciences) antibodies, and subsequently stained with A1488 conjugated Streptavidin (Thermo Fisher Scientific). Fluorescent images were captured using a 20×objectives on a fluorescence microscope (Keyence). There were fundamental limitations in the quantity and orientation of GCs in the LNs, so immunohistochemistry was not used to quantify cell types or GC features.
Immunophenotyping in LN: Inguinal lymph nodes were surgically removed from the animals after euthanasia and then were mechanically disrupted to create a cell suspension. For FACS analysis, cells were blocked with Fc receptor antibody (clone: 2.4G2, Tonbo Biosciences) and then stained with fluorochrome conjugated antibodies: CD19, GL7, CD95, CXCR4, CD86, IgG1, CD4, CXCR5, and PD1. After staining, cells were washed and fixed with 1.5% paraformaldehyde (PFA). Stained cells were analyzed on LSRII flow cytometer. Data were analyzed with FlowJo 10 (FlowJo LLC). See.
Immunophenotyping in Gel: Gels were surgically removed from the animals after euthanasia. The gels were processed using mechanical disruption to create a cell suspension which was then filtered through a 70 μm cell strainer (Celltreat). To count total cells in the gels, the single-cell suspensions were stained with anti-mouse CD45 and DAPI to count live leukocytes. The cells were mixed with CountBright Absolute Counting beads (ThermoFisher) and analyzed on a LSRFortessa X-20 flow cytometer (BD Biosciences). To analyze neutrophil and other myeloid cell counts in the gels, the single-cell suspensions were incubated with anti-CD16/CD32 (produced in house from 2.4G2 hybridoma) to block Fc receptor binding, and then stained with anti-mouse CD45, CD3, CD19, CD11c, Ly6G, CD11b, and MHCII in FACS buffer (2 mM EDTA, 2% FBS in PBS) for 30 min on ice. Dead cells were excluded by DAPI staining. Cells were acquired on a LSRFortessa X-20 and fcs files were analyzed using FlowJo 10 (FlowJo LLC). See
Animal Protocol: All animal studies were performed in accordance with National Institutes of Health guidelines, with the approval of Stanford Administrative Panel on Laboratory Animal Care.
Statistical Analysis: All results are expressed as a mean±standard deviation (s.d.) except for
FRAP Analysis: Alexa Fluor 647 conjugated OVA (Thermo Fisher Scientific), rhodamine conjugated Poly(I:C) (Invivogen), and rhodamine conjugated HPMC-C12 were used to visualize the diffusion of the cargo and gel. The samples were placed in a sterile 0.18 mm thick glass bottom dish (Ibidi). An Inverted Zeiss LSM 780 Laser Scanning Confocal Microscope (Germany) using a Plan-Apochromat 20×/0.8 M27 objective lens was used for FRAP measurements. To excite the fluorescent Alexa Fluor 647, a 5 mW 633 nm He—Ne laser was employed at 2%, and the emitted fluorescence was detected by Alexa Fluor 647 specific band pass filter (638-756 nm). To excite the fluorescent rhodamine, a 20 mW 561 nm diode pumped solid state laser was used at 2%, and the emitted fluorescence was detected by rhodamine specific band pass filter (415-638 nm). We used pixel dwell time 2.55 μs which took 390.98 ms to finish every scan. Samples were photobleached with a 405 nm diode laser, plus the 633 nm laser in the case of the Alexa Fluor, and a 514 nm argon laser and the 561 nm laser in the case of the rhodamine. All lasers were set al 100% intensity for the bleaching, with a 20-40 μm diameter for the region of interest (ROI). To avoid any extra noise, the high voltage was limited to be 700 V. Different tests (n=5) were made for 3 different samples from the same batch at different locations of the sample. For each test, 10 control pre-bleach images were captured at 1 frame/s. A spot was bleached with a pixel dwell time of 177.32 μs. 500 post-bleach frames were recorded at 1 frame/s to form the recovery exponential curve. The pixel size was set to be 1.66 μm. The diffusion coefficient was calculated as62:
D=γD(ω2/4Tτ1/2)
Where the constant γD=τ1/2/τD, with τ1/2 being the half-time of the recovery, TD the characteristic diffusion time, both yielded by the ZEN software, and ω the radius of the bleached ROI (25 μm). The software used for all FRAP tests was the ZEN lite (Zeiss). All the experiments were conducted in the Stanford University Cell Sciences Imaging Facility (CSIF) at room temperature.
Rg of Poly(I:C) was obtained from gel permeation chromatography (GPC) carried out using a Dionex Ultimate 3000 instrument (including pump, autosampler, and column compartment). Detection consisted of an Optilab TrEX (Wyatt Technology Corporation) refractive index detector operating at 658 nm and a HELEOS II light scattering detector (Wyatt Technology Corporation) operating at 659 nm. The column used was a Superose 6 increase 10/300 GL. The eluent was PBS buffer. 137 mM NaCl. 0.0027 mM KCl, 10 mM Phosphate pH 7.3, at 0.75 mL min−1 at room temperature. Analyte samples at 2 mg mL−1 were filtered through a PVDF membrane with 0.2 mm pore size prior to injection. Rg was calculated using a dndc of 0.19 mL/g and converted to RH using the following equation63: Rg/RH≈0.77.
The diffusivity of cargo in PBS was calculated using Stokes-Einstein Law Equation for diffusion64 where ka is Boltzmann's constant, T is temperature in Kelvin, η is solvent viscosity, and R is solute hydrodynamic radius:
The diffusivity of cargo in a model covalent PEG gel was calculated using the Multiscale Diffusion Model (MSDM) assuming 25° C., 5% volume fraction, and 35 nm mesh size56. The calculated values are comparable to experiment diffusivities of similar sized cargo found in the literature65.
TLR small molecule synthesis: The benzylamine TLR 7/8 agonist was synthesized as described by Shukla et. al.66, with a modification of the final aromatic substitution and tert-butyl carbamate removal on tert-butyl (4-((4-chloro-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-1-yl)methyl)benzyl)carbamate (1), which was done as a one-pot Staudinger-type reaction.67
Compound I and 2 equivalents of NaN3 (68 mg, 1.04 mmol) was dissolved in 1.5 mL DMSO, and was heated to 90° C. for 2 h. 2 equivalents of triphenylphosphine (273 mg, 1.04 mmol) was added, and the temperature was increased to 120° C. and stirred for 16 h. 4M HCl (aq; 0.5 mL) was added to the reaction flask and the temperature was lowered to 95° C. for 3 h. Water (4 mL) was added at which a precipitate formed, and the mixture was washed with EtOAc. NaCO3 (sat) was added to the aqueous phase which was extracted with EtOAc three times, which was done with Na2SO4 and the solvent removed in vacuo to yield a crude solid. This was purified by silica column chromatography on a Biotage system, using a gradient of DCM:MeOH, with 1% trimethylamine in the MeOH. This yielded compound II (1-(4-(aminomethyl)benzyl)-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-4-amine) (60 mg, 0.16 mmol, 32% yield).
Compound II (25 mg, 69 μmol) and propargyl-N-hydroxysuccinimidyl ester (17 mg, 76 μmol) were dissolved in 1 mL anhydrous DCM, and stirred for 2 hours at room temperature. The reaction mixture was loaded onto a silica column, and the product was purified using a DCM:MeOH gradient. This yielded compound III N-(4-((4-amino-2-(ethoxymethyl)-1H-imidazo[4,5-c]quinolin-1-yl)methyl)benzyl)-3-(prop-2-yn-1-yloxy)propanamide (25 mg, 53 μmol, 76% yield).
1H NMR (300 MHz, CDCl3-d): δ 8.05 (d, J=8.4 Hz, 1H, H1), 7.73 (d, J=8.3 Hz, 1H, H2), 7.58 (t, J=7.9 Hz, 1H, H3), 7.29 (3H, H4, H5), 7.02 (d, J=7.9 Hz, 2H, H6), 6.43 (s, 1H, H7), 5.91 (s, 2H, H8), 4.78 (s, 2, H9), 4.46 (d, J=5.9 Hz, 2H, H10), 4.15 (d, J=2.4 Hz, 2H, H 11), 3.82 (t, J=5.7 Hz, 2H, H12), 3.58 (q, J=7.2 Hz, 2H, H13), 2.55 (t, J=5.7 Hz, 2H, H14), 2.36 (t, J=2.4 Hz, 1H, H15), 1.14 (t, J=7.2 Hz, 3H, H16).
13C NMR (75 MHz, CDCl3): δ 171.35, 166.55, 151.77, 150.14, 138.70, 136.66, 136.09, 133.40, 129.84, 128.43, 125.71, 124.66, 121.24, 120.73, 112.79, 79.06, 75.01, 66.77, 66.02, 64.19, 58.44, 58.38, 58.21, 36.88, 31.89, 14.86.
Polymer synthesis and functionalization: N3-PEG-OH (0.5 g, 5 kDa, 100 μmol) in 2 mL anhydrous DCM with 30 μL DBU (30 mg, 0.20 mmol) was added quickly to a stirring solution of lactide (2 g, 13.9 mmol) in 6 mL anhydrous DCM. The solution was stirred for 10 min, after which 2 drops of acetic acid was added to quench the polymerization, and the polymer was precipitated into a 1:1 mixture of hexanes and diethyl ether. The polymer was re-dissolved in a minimal amount of acetone, precipitated again in diethyl ether and dried in vacuo.
A 20 mL scintillation vial was charged with compound III (14 mg, 30 μmol) and azido-poly(ethylene oxide)-b-poly(D,L-lactide) 5 kDa-20 kDa (0.5 g, 20 μmol) was dissolved in 4 mL of NMP and sparged with nitrogen for 10 min. Degassed CuBr (3.7 mg/mL: 0.1 mL) and THPTA (16 mg/mL) was added to the reaction flask. The reaction mixture was further sparged with nitrogen for 10 min. The reaction mixture was incubated for 16 h at room temperature, and precipitated into diethyl ether in a 50 mL centrifuge tube to recover the polymer. The polymer was then dissolved in ethyl acetate, precipitated into diethyl ether and dried in vacuo.
Conjugation was confirmed by 1H-NMR spectroscopy and increased UV absorption as indicated by GPC (DMF) eluagram.
NMR Spectroscopy: NMR spectra were obtained using an Inova 300 MHz NMR spectrometer with a Varian Inova console using VNMRJ 4.2 A software. Number-average (Mn) and weight-average (Mw) molar mass and dispersity (D=Mw/Mn) of polymers were obtained from gel permeation chromatography (GPC) carried out using an Dionex Ultimate 3000 instrument (including pump, autosampler, and column compartment) outfitted with an ERC Refractomax 520 refractometer. The columns were Jordi Resolve DVB 1000 Å, 5 m, 30 cm×7.8 mm and a Mixed Bed Low, 5 m, 30 cm×7.8 mm, with a Jordi Resolve DVB Guard Column, 1000 Å, 5 m, 30 cm×7.8 mm, 5 cm×7.8 mm. DMF with 10 mM LiBr was used as eluent at 1 mL min−1 at room temperature. Poly (ethylene glycol) were used to calibrate the GPC system. Analyte samples at 2 mg mL−1 were filtered through a nylon membrane with 0.2 mm pore size before injection (20 μL). Data was analyzed using Chromeleon GPC/SEC Software.
Surface Plasmon Resonance: A BIAcore X100 instrument (GE Healthcare) was used for all Surface Plasmon Resonance (SPR) kinetics measurements. Anti-mouse IgG antibodies (Cat. BR100838, GE Healthcare) were immobilized to a CM5 chip (Cat. BR100012, GE Healthcare) using ammine coupling according to the manufacturer's instructions. Serum samples were injected at a 1:100 dilution in HBS-EP+ buffer (Cat. BR 100826, GE Healthcare) for 140 seconds at a flow rate of 10 μL/min to capture the mouse antibodies onto the chip. For association analysis, 5 different concentrations of OVA in HBS-EP+ buffer (978 nM, 733 nM, 489 nM, 244 nM, 162 nM) were flowed for 90 seconds at 30 μL/min followed by 9×) seconds of buffer alone. For dissociation analysis, OVA at 489 nM was run for 90 seconds at 30 μL/min followed by 30 minutes of HBS-EP+ to allow for OVA to dissociate from the antibodies.
For each run, we had two steps of reference subtraction. First, naïve mouse serum was run in parallel to the test serum in the reference cell of the chip and subtracted from all test curves. Second, buffer was run above the captured antibodies with the same settings as the OVA solutions to subtract for temporal changes due to fluid flow.
To calculate the association rate constant, GraphPad Prism 7.04 (GraphPad Software) was used to fit a one-phase association model to all 5 concentrations of OVA. The observed k values from those fits were plotted against the concentration and fit with a linear equation. The slope from this line was defined as the association rate for that sample.
To calculate the dissociation rate constants and account for the polyclonal population of antibodies in the serum, GraphPad Prism 7.04 (GraphPad Software) was used to fit a 3-component exponential decay to the dissociation curves:
Y=Y1*exp(−K1*X)+(2*exp(−K2*X))+((Y3)*exp(−K3*X))
The initial values used for the model were:
Y1=0.25,Y2=0.25,Y3=0.5
K1=0.1,K2=0.001,K3=0.001
The component with the highest fraction and lowest dissociation rate was used for comparing the experimental groups.
In vivo passive transfer influenza challenge model: Serum was collected at day 56 after vaccination with A/Brisbane/59/2007(H1N1) hemagglutinin delivered in (i) TLR 7/8 NP-based 2:10 hydrogels, (ii) MF59, or (iii) Alum. Control serum was collected from a naïve mouse. Serum from 5 mice per group was pooled and antibodies were isolated as follows: First, a Zeba Spin Desalting Column (Thermo Scientific, 89882) was used to buffer exchange into Melon Gel Buffer, then IgG antibodies were purified using a Melon Gel IgG Spin Purification Kit (Thermo Scientific, 45206), and lastly the solution was buffer exchanged into PBS using the Zeba Spin Desalting Column. The columns were used following the manufacturer's guidelines. A mixture of the purified serum and PR8 virus was incubated at 37° C. for 30 minutes before nasal inoculation into C57/BL6 mice. The nasal inoculation contained 4 μL of purified serum, 2,000 PFU of PR8 virus, and PBS to a total volume of 20 μL. Mice were weighed daily to monitor morbidity and animals that exceeded 25% weight loss were euthanized.
Vaccines are among the most effective medical advancements in history to prevent and control infectious diseases. The eradication of smallpox, near eradication of poliomyelitis, and vast decreases in diphtheria, measles, and rubella are testaments to the ability of vaccines to mitigate disease burden worldwide1,2. It is estimated that vaccines save 2.5 million lives worldwide per year3. However, there are an abundance of infectious diseases that still do not have effective vaccines; for example, influenza, which still kills roughly 500,000 people annually worldwide despite the yearly flu vaccines. Furthermore, new viral strains continue to emerge, threatening our society and challenging our humanity. The most notable current example is the ongoing pandemic caused by severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) that emerged at the end of 2019. To date, while this is being written, more than 60 million cases and 1.4 million deaths have been confirmed; marking the COVID-19 pandemic as the most serious health crisis of modern times. The COVID-19 pandemic highlights the increasing threat of new pandemic strains of viruses, and further motivates the need for continued improvement of vaccines technologies.
Different types of vaccines, such as inactivated viruses, live-attenuated viruses and subunit vaccines4 are commonly used in the clinic. The subunit vaccine approach offers an opportunity for more precise vaccine design, improved safety, stability, and inexpensive and scalable manufacturing capabilities, as well as less extreme storage conditions5. Subunit vaccines contain specific microbial antigens, which directs the antibody response to a specific foreign substance, along with one or more immune stimulating molecule(s), which are commonly referred to as adjuvants. Since subunit vaccines, such as the spike protein in COVID-19 vaccines, elicit weaker immune responses, adjuvants are needed to promote vaccine efficacy and the body's immune response to a pathogen6. As subunit vaccines become more popular, there is a growing need for new approaches to augment the adjuvant potency and guide the immune system to enhance the quality and durability of immune responses.
Adjuvants trigger various pathways of innate immune cell activation, such as pattern recognition receptor (PRR) signaling on innate immune cells. Currently, a handful of adjuvants that have been used in the clinic include aluminum salt-based adjuvants (Alum), squalene-based oil-in-water emulsions such as MF59 and AS03, and toll-like receptor agonists (TLRa) like the synthetic oligodeoxynucleotide CpG ODN (cytidine-phosphate-guanosine oligodeoxynucleotide) 1018. CpG 1018, specifically, is the only clinically approved version of CpG, is included in the Hepatitis B vaccine Heplisav B (2017), has been shown to increase the immune response by strongly activating B cells7 and is therefore currently being tested in phase I clinical trials with numerous SARS-CoV-2 vaccine candidates.
Synthetic CpG motifs are agonists for TLR9, mimic the activity of naturally occurring motifs found in bacterial DNA and are commonly used as immunostimulatory adjuvants to increase vaccine immunogenicity. Specifically. CpG triggers intracellular signaling leading to activation of macrophages, dendritic cells (DCs) and B-cells, as well as the production of chemokines and cytokines, enhancing innate and adaptive immune responses. CpGs are composed of short synthetic single-stranded DNA molecules of approximately 20 base pairs, contain unmethylated CpG dinucleotides in a specific sequence and have a phosphorothioated backbone to allow for increased stability. Three major classes of CpGs exist, including Class A. Class B, and Class C, and they differ based on their distinct biological activity and structural characteristics. Class B (CpG-B) strongly activates B cells and TLR9-dependent nuclear factor kappa-B (NF-kB) signaling pathways but weakly stimulates interferon-alpha (IFN-α) secretion, whereas Class C (CpG-C) both stimulates B cells and induces IFN-α production8. The biggest enhancer of the humoral response is CpG-B, which has been found to be a highly effective vaccine adjuvant inducing Th1 polarized responses8-10. Numerous studies have demonstrated innate immune defenses being activated by CpG-B11-13 but fewer studies have been reported with CpG-C14-15.
To maximize the activation of antigen presenting cells (APCs) and other innate immune cells while minimizing systemic toxicities16,17, significant efforts have focused on localizing TLR agonists, among other adjuvants, to the injection site and lymph nodes (LNs) using nanoparticles (NPs). Polymer-based NPs have been heavily studied due to their modularity, ease of production, scalability and biocompatibility16-20. Furthermore, NPs have been widely employed in drug delivery applications21-26. These advantages make NPs a powerful design platform to overcome the challenges associated with conventional delivery methods to enhance the spatiotemporal responses of vaccines27. Precise control of the NPs size28,29,30,31-33 enables efficient targeting to LNs without requiring specific cell targeting ligands. NPs between 20 and 100 nm efficiently drain through the lymphatic system into the LNs28,29,34 where they are directly taken up by LN-resident APCs35. Recent reports also have discovered that covalently conjugating TLRa to polymer NPs, marks a significant increase in both antibody production against the antigen and on the induction of cytotoxic T-cells17. Similarly, other studies have leveraged particle technologies to increase the density of presented adjuvant molecules on their surface to improve the potency of the adjuvant response17,36-38. Controlling adjuvant presentation is a useful tool because many TLRs cluster upon ligand binding and this clustering effect can enhance downstream signaling, leading to a more potent response39. There is a growing body of work demonstrating that immune cells require precise spatial and temporal cues to drive specified responses, supporting the idea that spatiotemporal control of vaccines can have a profound effect on the magnitude and quality of the immune response40-49.
Recent proof-of-concept studies evaluating slow delivery, have shown that sustained release of an HIV vaccine through implantable osmotic pumps leads to greatly improved quality of vaccine responses, such as durable germinal center (GC) responses, high antibody titers, and development of better virus neutralization compared to standard bolus administration of the same vaccine46. Our group recently demonstrated the use of easily injectable physically cross-linked Polymer-NanoParticle (PNP) hydrogels for prolonged co-delivery of subunit vaccine components can greatly improve the potency, durability and breadth of the elicited immune response39. PNP hydrogels are composed of polymers physically and dynamically crosslinked by multivalent noncovalent interactions with the NPs43,50-54. This newly developed hydrogel, which can be administered in a minimally invasive fashion, is capable of providing sustained release of cargo molecules over the course of weeks, while prolonging the GC reaction and improving the affinity of the antibodies by more than 1000-fold43.
In the current study we sought to optimize the delivery of CpG, by developing a NP adjuvant platform via chemically conjugating a Class B (CpG-B) or a Class C (CpG-C) CpG adjuvant to polymeric poly(ethylene glycol)-b-poly(lactic acid) (PEG-b-PLA) NPs of approximately 50 nm in size, allowing for efficient, passive and direct transport to the LNs (
A potent vaccine response results in durable production of neutralizing antibodies (humoral immunity), and in the activation of specific T cells (cell-mediated immunity). Antibodies bind to antigens on the pathogen's surface and block them from infecting host cells by neutralizing the pathogen. Cellular immunity involves T cells that can directly kill infected cells, which in turn eliminates the pathogen.
The materials discussed in this work primarily influence APCs and downstream humoral immunity, so those mechanisms are the focus of this section. The immune system first interacts with a vaccine at the site of injection. The extent of local inflammation is controlled by both the site of administration and the presence of immune activating molecules called adjuvants in the vaccine. Innate immune cells such as macrophages and DCs immediately respond to these adjuvants by producing cytokines to recruit cells to the site of injection. After the initial inflammatory response at the site of injection, activated DCs and the vaccine components themselves travel via afferent lymph vessels to the draining LNs55,56. LNs provide a site where cells that are present at low frequencies within the body come together with precise spatial organization to enhance the cell-cell interactions necessary for generating a robust immune response. Migratory DCs present antigen directly to T cells themselves in the LN57,58. The migration of immune cells can be enhanced by the use of adjuvants59,60. In addition to antigen-presenting migratory DCs, there are also LN-resident DCs which are constantly scanning the lymph to capture any soluble antigens that reach the LNs by passive diffusion. The vaccine response therefore benefits from promoting antigen presentation and migration of tissue residents DCs as well as increasing antigen trafficking to the LN.
The overall goal of activating the innate immune system is to initiate an adaptive immune response to the vaccine, whereby APCs must transport and present antigen to B cells. Antigen-specific B cells proliferate and can differentiate into short lived antibody producing cells within a few days. Alternatively, antigen activated B cells can migrate into the B cell follicles in the LN, where they form dynamic microenvironments called germinal centers (GCs). The GC has two anatomical compartments, the dark zone where B cells proliferate and undergo somatic hypermutation (SHM) and the light zone where antigen-driven selection favors higher-affinity B cells61. It is very important for vaccines to initiate GC responses that are maintained for enough time to allow for many cycles of SHM to improve both, the breadth and the affinity of antibody responses. Finally, after the immune response is mounted, the circulation of antibodies in the blood allows them to reach infected tissues and protect the entire body.
Poly(ethylene glycol)methyl ether 5000 Da (PEG), Poly(ethylene glycol)α-hydroxy-ω-azido terminated 5000 Da (N3-PEG-OH), 3,6-Dimethyl-1,4-dioxane-2,5-dione (Lactide), 1,8-diazabicyclo(5.4.0)undec-7-ene (DBU, 98%), (Hydroxypropyl)methyl cellulose (HPMC, meets USP testing specifications), N,N-Diisopropylethylamine (Hunig's base), N-methyl-2-pyrrolidone (NMP), 1-dodecyl isocyanate (99%), mini Quick Spin Oligo columns (Sephadex G-25 Superfine packing material), Sepharose CL-6B crosslinked, bovine serum albumin (BSA), were purchased from Sigma-Aldrich, U.S. CpG-B 1826 (5′-TCCATGACGTTCCTGACGTT-3′) and CpG-C 2395 (5′-TCGTCGTTTTCGGCGCGCGCCG-3′) oligonucleotides, Aluminum hydroxide gel (Alhydrogel adjuvant 2%), zeocin were purchased from InvivoGen. Amino CpG-B 1826 (5′Amino Modifier C6, 5′-NH2-TCCATGACGTTCCTGACGTT-3′), amino CpG-C 2395 (5′Amino Modifier C6, 5′-NH2-TCGTCGTTTTCGGCGCGCGCCG-3′) were purchased from Integrated DNA Technologies (IDT); in order to allow for a direct comparison between the modified and unmodified CpG classes, unmodified CpG-B 1826 and CpG-C 2395 were also purchased from Integrated DNA Technologies. Dibenzocyclooctyne-PEG4-N-hydroxysuccinimidyl ester (DBCO-PEG4-NHS ester) and Alexa Fluor 647 DBCO (AFDye 647 DBCO) were purchased from Click Chemistry Tools. Dulbecco's modified Eagle's medium (DMEM, Gibco), phosphate buffered saline (PBS pH 7.4, Gibco) and Invitrogen E-Gel EX Agarose Gels 4% were purchased from Thermo Fisher Scientific. Heat Inactivated fetal bovine serum (HI-FBS) was purchased from Atlanta Biologicals. IFN-α cytokine enzyme-linked immunosorbent assay (ELISA) kit was purchased from PBL Assay Science, and the TNF-α cytokine ELISA kits was purchased from R&D Systems (Fisher Scientific). Goat anti-mouse IgG Fc secondary antibody (A16084) HRP (Horseradish peroxidase) was purchased from Invitrogen. 3,3′,5,5′-Tetramethylbenzidine (TMB) ELISA Substrate, high sensitivity was acquired from Abcam. HIS Lite Cy3 Bis NTA-Ni Complex was purchased from AAT Bioquest. Unless otherwise stated, all chemicals were used as received without further purification. Deionized (DI) water (Milli-Q, MilliporeSigma, 18 MΩ) was used for all the experiments. In order to remove any traces of water, all glassware and teflon-coated magnetic stir bars used for the polymer synthesis were dried overnight in the oven. Nuclear magnetic resonance (NMR) spectra of the polymeric structures were obtained on an Inova NMR spectrometer (1H: 600 MHz, 13C: 150 MHz) with a Varian Inova console and using the VNMRJ4.2 software. Chemical shifts (δ) are given in parts per million (ppm) and peak multiplicity is reported as follow: s=singlet and m=multiplet.
PEG-PLA synthesis. PEG-PLA was prepared as previously reported51. Prior to use, the commercial lactide was recrystallized in ethyl acetate and DCM was dried via cryo distillation. Under inert atmosphere (N2), PEG-methyl ether (5 kDa, 0.25 g, 4.1 mmol) and DBU (15 μL, 0.1 mmol) were dissolved in 1 mL of anhydrous DCM. Lactide (1.0 g. 6.9 mmol) was dissolved under N2 in 3 mL of anhydrous DCM. The lactide solution was then quickly added to the PEG/DBU mixture and was allowed to polymerize for 8 min. The reaction was then quenched with an acetic acid solution and the product precipitated into a 1:1 mixture of ethyl ether and hexanes solution mixture, collected by centrifugation and dried under vacuum. 1H NMR (600 MHz, CDCl3, δ in ppm in ppm): 1.48-1.62 (m, 3H, CH3 PLA), 3.64 (s, 4H, CH2 PEG), 5.1-5.24 (m, 1H, CH, PLA).
Azide PEG-PLA synthesis. Azide-PEG-PLA was synthesized similarly to PEG-PLA. Prior to use, DCM was dried using 3-4 Å molecular sieves and N3-PEG-OH was dried under vacuum overnight. Under inert atmosphere a solution of N3-PEG-OH (0.5 g, 5 kDa, 100 μmol) and DBU (30 μL, 30 mg, 9.29 mmol) in anhydrous DCM (1 mL) was rapidly added to a solution of lactide (2.0 g 13.9 mmol) in anhydrous DCM (10 mL) and stirred for 8 min. The reaction mixture was quenched with an acetic acid solution, precipitated by an ethyl ether:hexane (1:1) mixture, centrifuged and dried under vacuum overnight. 1H NMR (600 MHz, CDCl3, δ in ppm): 1.48-1.62 (m, 3H, CH3 PLA), 3.64 (s, 4H, CH2 PEG), 5.1-5.24 (m, 1H, CH, PLA). 13C NMR (150 MHz, DMSO-d6, δ in ppm): 16.4 (C-B, PLA), 68.6 (C-C, PLA), 69.8 (C-A. PEG), 169.2 (C-D, PLA).
Synthesis of DBCO-CpG intermediate. DBCO-PEG-NHS ester (3.78 mg, 5.8 μmol) was dissolved in DMSO (40 μL). The solution was then diluted with PBS in order to reach a final concentration of 10 mM. The Amino-CpG-B and Amino-CpG-C (0.58 μmol) were then reacted with the DBCO-PEG4-NHS ester solution for 4 h at RT. The solution was then directly purified by size-exclusion in PBS, using a Sephadex G-25 Superfine (mini Quick Spin Oligo) column and stored at −20° C.
NPs synthesis and conjugation. PEG-PLA NPs were synthesized as previously described43,62. A 1 mL solution of PEG-PLA in 75:25 AcCN:DMSO (50 mg/mL) was added dropwise to 10 mL of DI water stirring at 600 rpm. After NPs precipitation the particles solution was purified in centrifugal filters (Amicon Ultra, MWCO: 10 kDa) at 4500 RCF for 1 h and resuspended in PBS to reach a final concentration of 200 mg/mL. N3-PEG-PLA NPs of different valencies (10%, 20%. 30% and 50%) were synthesized by premixing PEG-PLA and N3-PEG-PLA polymer solutions at the correct volumetric ratios before precipitation. Purification and concentration of the CpG NPs were performed as above stated. For adjuvant conjugation the modified DBCO-CpG-B and DBCO-CpG-C, and N3-PEG-PLA NPs were reacted via copper-free click chemistry for 12 h at room temperature. After reaction completion, CpG-B conjugated NPs and CpG-C conjugated NPs were purified by size-exclusion chromatography using a CL-6B matrix and stored in PBS at 4° C. Successful purification of the CpG-NPs from unreacted free CpG was confirmed via aqueous GPC and via agarose gel electrophoresis (4%). The CpG concentration on the NPs was determined through absorption calibration curves at 280 nm acquired using a Synergy H1 Microplate Reader (BioTek Instruments). An individual calibration curve for each NPs valency and TLR9 agonist class was recorded.
HPMC-C12 synthesis. HPMC-C12 was prepared according to a previously reported procedure51. HPMC (1.0 g) was dissolved in NMP by stirring at 80° C. for 1 h before removing the solution from the heat. Once cooled to RT. 1-dodexylisocynate (105 mg, 0.5 mmol) and Hunig's base, acting as the catalyst (˜3 drops) were dissolved in NMP (0.5 mL). This solution was then added dropwise to the reaction mixture, which was stirred at RT for 16 h. The resulting product solution was precipitated using acetone, decanted, redissolved in Milli-Q water (˜2 wt %) and dialyzed against water for 4 days. The polymer mixture was then lyophilized and reconstituted to a 60 mg/mL solution in sterile PBS.
Polymer characterization. After polymer synthesis relative molecular weights, number-average Mn, and weight-average Mw, molar mass, and polydispersity (PDI=Mw/Mn) were obtained from gel permeation chromatography (GPC) after passing through two size exclusion chromatography columns (Resolve Mixed Bed Low DVB, ID 7.8 mm, Mw range: 200-600,000 g/mol, Jordi Labs) in a mobile phase of DMF with 10 mM LiBr at 35° C. and a flow rate of 1.0 mL/min (Dionex Ultimate 3000 pump, degasser, and autosampler, Thermo Fisher Scientific). Before injection, samples at a concentration of 5 mg/mL were filtered through a 0.22 μm nylon membrane. The data acquired was analyzed using the Chromeleon7 GPC software. Nuclear magnetic resonance (NMR) spectra of the polymeric structures were obtained on an Inova NMR spectrometer (OH: 600 MHz, 13C: 150 MHz) with a Varian Inova console and using the VNMRJ4.2 software
NPs characterization. The hydrodynamic diameter of the NPs, before and after adjuvant conjugation, was characterized by dynamic light scattering (DLS) (DynaPro II plate reader, Wyatt Technology), and the NPs surface charge was measured with a Malvern Zetasizer Nano Zs. For the CpG-NPs, successful conjugation of the TLR9 agonist to the NPs was confirmed via aqueous GPC after passing through a SEC column [5000 to 5,000,000 g/mol]; Superose 6 Increase 10/300 GL (GE Healthcare) in a mobile phase of PBS containing 300 parts per million of sodium azide and at a flow rate of 0.75 mL/min (Dionex Ultimate 3000 pump, degasser, and autosampler, Thermo Fisher Scientific). Detection consisted of an Optilab T-rEX (Wyatt Technology Corporation) refractive index detector operating at 658 nm and a diode array detector operating at 280 nm (Dionex Ultimate 3000, Thermo Fischer Scientific). Before injection, samples at a concentration of 1 mg/mL were filtered through a 0.22 μm PVDF membrane. The data acquired was analyzed using the Chromeleon7 GPC software and the ASTRA software from Wyatt Technologies.
PNP and CpG-PNP Hydrogel formation. The CpG polymer-nanoparticle (CpG-PNP) hydrogels were formed at 2 wt % HPMC-C12 and 10 wt % mixture of PEG-b-PLA and CpG-PEG-PLA NPs in PBS. The gels were prepared by mixing a 3:2:1 weight ratio of 6 wt % HPMC-C12 polymer solution, 20 wt % NPs solution, and PBS. Based on the desired adjuvant dosing, 30% CpG conjugated NPs were mixed with non-conjugated PEG-PLA NPs prior to hydrogel formation. The gels were formed by mixing the solutions using syringes connected through an elbow mixer.
Rheological Characterization of PNP gels. All rheological characterization was completed on a Discovery HR-2 Rheometer (TA Instruments). Measurements on the hydrogels were performed using a 20 mm serrated plate geometry at 25° C. and at 500 μm gap height. Dynamic oscillatory frequency sweeps were conducted at a constant 1% strain and angular frequencies from 0.1 to 100 rad/s. Amplitude sweeps were performed at a constant angular frequency of 10 rad/s from 0.5% to 10,000% strain. Flow sweep, steady shear experiments were performed at shear rates from 50 to 0.005 l/s, whereas stress-controlled flow sweep measurements were conducted at shear rates from 0.001 to 10 l/s. Step-shear experiments were performed by alternating between low shear rates (0.1 rad/s for 60 s) and high shear rates (10 rad/s for 30 s) for three full cycles. Yield stress values were extrapolated from stress-controlled flow sweep and amplitude sweep measurements.
FRAP Analysis. Fluorescence recovery after photobleaching (FRAP) was performed on the PNP hydrogel and on the CpG-PNP hydrogel formulations using a confocal LSM780 microscope. Each individual component of the hydrogel was labelled with a fluorescent component. NP-tethered AF647 (10 wt %), rhodamine-conjugated HPMC-C12, His-tagged SARS-CoV-2 spike conjugated with HIS-Lite-Cy3 Bis NTA-Ni Complex (0.27 mg per mL of gel) were used to visualize diffusion of the vaccine cargo and the gel components. Samples were imaged using low intensity lasers to collect an initial level of fluorescence. Then a high intensity laser with an aperture of 25 μm was focused on the region of interest (ROI) for 10 s in order to bleach a circular area. Subsequently, fluorescence emission data was recorded for 4 min to create an exponential fluorescence recovery curve. For each sample, replicate measurements (n=2-5) were taken at multiple locations. The diffusion coefficient D was calculated according to the following equation63:
where the constant γD=τ1/2/τD, with τ1/2 being the half-time recovery, to the characteristic diffusion time, both yielded by the ZEN software, and a) the radius of the bleached ROI. The diffusivity of the SARS-CoV-2 spike protein antigen in PBS was calculated using the Stokes-Einstein law equation for diffusion64:
with kB being the Boltzmann's constant, T the temperature in Kelvin, η the solvent viscosity, and RH the solute hydrodynamic radius. The hydrodynamic radius of the spike protein was measured via DLS to be RH=12.2 nm, whereas qi for PBS was approximated to be 0.8872 mPa s at 25° C. The measured RH agrees with the value published in literature and measured via Cryo-EM65.
Vaccine formulation. The SARS-CoV-2 spike protein vaccines were injected subcutaneously either in form of a bolus injection or in form of a gel. For the bolus injections, the vaccine formulation contained a 10 μg antigen dose of spike S1+S2 ECD (R683A, R685A, F817P, A892P, A899P, A942P, K986P, V987P)-His Recombinant Protein (Sino Biological) and a 20 μg 30% CpG-NPs adjuvant dose; boosting was performed at day 21. For the CpG-PNP hydrogels, the dose was doubled containing 20 μg of antigen and 40 μg of CpG-NPs adjuvant; for the gel groups, boosting was not performed. Control groups were composed of non-conjugated PEG-PLA NPs, Alum (100 μg, Invivogen), and soluble CpG-B and CpG-C (20 μg. IDT) vaccines. Mouse blood was collected from the tail veins each week for 5 weeks. In order to analyze the early cytokine response, blood was collected at 0 h, at 3 h and 24 h from injection and stored at −80° C. These serum samples were analyzed for IFN-α and TNF-α levels and the concentrations were determined via enzyme-linked immunosorbent assay (ELISA) according to manufacturer's instructions and were calculated from standard curves. Absorbance was measured with a Synergy H1 microplate reader (BioTek Instruments) at 450 nm.
In vitro Raw-Blue reporter assay. The Raw-Blue (NF-kB-SEAP) reporter cell line (Invivogen, raw-sp) was used to evaluate the valency effect of TLR9 agonist conjugated to PEG-PLA NPs. The cells were cultured at 37° C. with 5% CO2 in DMEM supplemented with L-glutamine (2 mM), D-glucose (4.5 g/L), 10% HI-FBS, and penicillin (100 U/mL)/streptonmcin (100 μg). Every other passage, zeocin (100 μg/mL) was added to the culture medium. Serial dilutions of soluble CpG-B and CpG-C and of the different CpG-B and CpG-C NPs formulations were added to a 96-well tissue culture treated plate to achieve final concentrations between 30 and 3.1 μg/mL of TLR9 agonist. Non-conjugated PEG-PLA NPs were used as a negative control. About 100,000 cells were added to each well in 180 μL of media and were incubated for 21 h at 37° C. in a CO2 (5%) incubator. Manufacturer instructions were followed for SEAP quantification, and absorbance levels detected at 655 nm after 3 h incubation with QUANTI-Blue Solution (Invivogen). Nonlinear regression fits were found using the “Log(agonist) vs. response—EC50” function in GraphPad Prism 8.4 software.
Animal studies. Six-to-seven weeks old female C57BL/6 (B6) mice were obtained from Charles River, housed in the animal facility at Stanford University and cared for according to Institutional Animal Care and Use guidelines. All animal studies were performed in accordance with the National Institutes of Health guidelines and the approval of Stanford Administrative Panel on Laboratory Animal Care. The day before vaccine administration mice were shaved in order to receive a subcutaneous injection of 150 μL of CpG-PNP gel or 100 μL of bolus on the right side of their backs. Mouse blood was collected from the tail veins each week for 5 weeks.
Antibody concentration. Serum Anti-spike IgG antibody endpoint titers were measured using an ELISA. Maxisorp plates (Thermofisher) were coated with SARS-CoV-2 spike protein at 2 μg/mL in PBS overnight at 4° C. and subsequently blocked with PBS containing 1 wt % BSA for 1 h at 25° C. Serum samples were serially diluted and incubated in the previously coated plates for 2 h at 25° C., and goat-anti-mouse IgG Fc-HRP (1:10,000) was added for 1 h at 25° C. Plates were developed with TMB substrate, the reaction stopped with 1 M HCl and the plates analyzed using a Synergy H1 microplate reader (BioTek Instruments) at 450 nm. End point titers were defined as the highest serum dilution at the one an optical density above 0.1 was detected.
SARS-CoV-2 Spike-pseudotyped Viral Neutralization Assay. Neutralization assays were conducted as previously described66. Briefly, SARS-CoV-2 spike pseudotyped lentivirus was produced in HEK239T cells and cells seeded at six million cells the day prior to transfection. A five-plasmid system was used for viral production. Plasmids were added to filter-sterilized water and HEPES-buffered saline was added dropwise to reach a final volume of 1 mL. In order to form transfection complexes. CaCl2 was added dropwise to the gently agitated solution. The transfection reactions were incubated for 20 min at RT and then added to plated cells. Virus-containing culture supernatants were harvested ˜72 hours after transfection via centrifugation and filtered through a 0.45 μm syringe filter. Viral stocks were stored at −80° C. For the neutralization assays, ACE2/HeLa cells were plated 1 to 2 days prior to infection and mouse serum was heat inactivated at 56° C. for 30 min prior to use. Mouse serum (1:50 dilution) and virus were diluted in cell culture medium and supplemented with polybrene at a final concentration of 5 μg/mL. Serum/virus solution at 1:50 were incubated at 37° C. for 1 h. After the 1 h incubation, the media was removed from the cells and incubated with the serum/virus solution at 37° C. for 48 h. After complete incubation, the cells were then lysed using BriteLite (Perkin Elmer) luciferase readout reagent, and luminescence was measured with a BioTek plate reader. Each plate was normalized by averaging the readout from the wells containing only the virus or only the cells.
Statistical Analysis. All results are expressed as mean±standard deviation (s.d). Comparison between two groups were conducted by a two-tailed Student's t-test. One-way ANOVA tests with a Tukey's multiple-comparisons test was used for comparison across multiple groups. Statistical analysis was performed using GraphPad Prism 8.4 (GraphPad Software). Statistical significance was considered as p<0.05.
Chemical synthesis and purification of CpG NPs. In this study, we conjugated the TLR9 adjuvant CpG to PEG-PLA NPs in order to improve its stability, targeting to LNs and uptake by APCs. CpG is a strong inducer of Th1 immune responses that have proven important in clinical translation. We have previously described the synthesis of PEG-b-PLA) and azide-terminated PEG-PLA (N3-PEG-b-PLA) block copolymers50,51,62,67 (
The synthesized polymers were used as carriers to deliver CpG. CpG functionalized PEG-PLA NPs were synthesized via a two-step reaction: i) an NHS ester reaction to chemically link CpG to a cyclooctyne derivative functionality (
CpG is a UV active molecule that absorbs UV light at a wavelength of λ=280 nm, therefore a successful conjugation to the NPs can be confirmed via an increase in UV/RI signal at the NPs elution time after passing through the SEC column of an aqueous GPC. A range of molar excesses of DBCO-CpG, from 1 to 5, were tested to optimize the conversion of the click reaction onto the differently functionalized NPs surface (
Characterization of CpG NPs. Furthermore, it was important to confirm that the conjugation of TLR9 agonist to the NPs surface did not alter the NPs physical and colloidal properties. To this end, CpG-B and CpG-C NPs were characterized via surface zeta potential measurements (
CpG NPs and plain NPs were characterized using DLS and were found to have hydrodynamic diameters of DH of 50.0±3.1 nm (sd: 2.5, 30% N3-NPs), 54.8±4.4 nm (sd: 5.8, 30% CpG-B NPs), 60.3±3.4 nm (sd: 3.2, 30% CpG-C NPs) (
In vitro evaluation of CpG NPs potency. To ensure that the CpG conjugation to the NPs did not impair the biological activity and immunogenicity of the adjuvant, RAW-Blue transgenic mouse macrophage cells were used to quantify TLR9 activation. The cells were incubated with either CpG-C NPs, or soluble CpG-C. Activation of TLR9 by the CpG agonist induces signaling pathways leading to the activation of NF-kB and AP-1 (activator protein), and the subsequent secretion of secreted alkaline phosphatase (SEAP) which is easily detectable and measurable calorimetrically using QUANTI-Blue detection medium (
In these assays, Raw-Blue cells were incubated for 21 h with either the free form, or the NPs tethered form (10%, 20%, 30%, 50%) of CpG-C at a range of CpG concentrations (3.1-29 μg/mL) in order to generate concentration dependent activation curves (
In vitro, none of the tested NPs performed as well as the free CpG, demonstrating that conjugation of adjuvants to PEG-b-PLA NPs slightly decreases the potency of the molecule. However, in vitro cell assays do not take pharmacokinetics or biodistribution into account, and therefore are not always a good proxy for in vivo studies71. Based on these results, the 30% CpG valency NP was identified as the most promising CpG NPs adjuvant to promote high magnitude and persistent immune activation and was therefore tested in an in vivo vaccine study with SARS-CoV-2 as the antigen.
Together, these data demonstrate that we were able to synthesize pure and stable CpG NPs of different valencies, that were capable of activating TLR9 and showed similar potencies in vitro to free CpG. It has recently been shown, that the density and valency of antigen and adjuvant in vaccine design, dictates the magnitude and composition of the immune response6.
Furthermore, recent studies have highlighted the importance of vaccines in initiating GC responses that are maintained for long enough to improve the breadth and affinity of antibody responses46,72. Sustained delivery technologies can promote long-lived GCs and therefore increase the magnitude and persistence of the humoral immune response. We have previously described the development of injectable PNP hydrogel materials that can encapsulate physiochemically diverse vaccine components such as antigens and adjuvants and can be tuned to provide sustained co-delivery over extended periods of time. Because of these characteristics, we expected that by loading and implementing our adjuvant NPs with improved potency into the slow and controlled release PNP hydrogels, we will further enhance the vaccine response.
Formation of CpG-PNP hydrogels and rheological characterization thereof. Physically crosslinked PNP hydrogels can be quickly and easily formed by mixing aqueous solutions of hydroxypropyl methylcellulose derivatives (HPMC-C12) and biodegradable PEG-b-PLA NPs43,50,53,54 (
To ensure that the CpG-conjugation to the NPs did not influence the mechanical properties of the PNP gels, an extensive rheological characterization was performed where we compared, PNP hydrogels loaded with the equivalent of 20 ug of CpG-C in the form of CpG-C NPs (CpG-C PNP) to PNP hydrogels loaded with non-functionalized NPs (
In order to measure the hydrogel's viscoelastic response, frequency-dependent oscillatory shear experiments were conducted within the linear viscoelastic regime (LVER). The introduction of CpG-NPs did slightly alter the PNP's frequency response. Gels with (CpG PNP gels) and without (bare PNP gels) CpG-NPs exhibited solid-like properties within the explored frequency range (0.1 rad/s-100 rad/s) where the storage (G′) modulus was greater than the loss (G″) modulus. At a representative angular frequency of ω=10 rad/s the CpG-C PNP and the PNP gels exhibited storage moduli of 440 and 430 Pa, respectively. No crossover frequency was measured within this frequency range.
The yielding response of the hydrogels was measured with both dynamic amplitude sweeps and stress-controlled flow experiments. A dynamic amplitude sweep (
η=K{dot over (γ)}n−1 (3)
with K being the consistency index, {dot over (γ)} the shear rate and n the shear thinning parameter. A shear thinning parameter below 1 is representative of a shear thinning fluid; a scaling exponent of −1 (n=0) is an evidence for a pre-yielding behavior. The flow portion of the viscosity vs. shear rate plot was fitted to the power law in order to find the shear thinning parameter. For the PNP and CpG-C PNP hydrogels n˜0.2 and n˜0 were extrapolated. From this fit we conclude that PNP hydrogels show dramatically shear thinning behaviors, whereas CpG-C PNP hydrogels show yielding. However, it has to be noted that the flow sweep measurement of the CpG-C PNP hydrogel has to be critically evaluated since the gel was partly ejected from the geometry.
The PNP hydrogels also display injectability, which is defined as the ability of the gel to flow at pressures that are clinically applied while injecting through a needle. To investigate the hydrogels' ability to self-heal after flow, step-shear experiments were performed by stepwise interchanging between low (0.1 l/s) and high (10 l/s) shear rates for 3 full cycles. At low shear rates the viscosity was of about 1000 Pa s, which decreased by 2 orders of magnitude to approximately 15 Pa s upon the application of elevated shear rates. The hydrogels are thixotropic, showing a considerable delay in restructuring their dynamic structure. This phenomenon is observable in the step-shear experiments where, once the shear is decreased to a steady-state value, the viscosity slowly increases to its equilibrium value (η˜1000 Pa s) while rebuilding its internal network (
The rheological characterization of these materials supports that the CpG-PNP hydrogels can be injected through high-gauge (e.g., 21-gauge) needles (
Vaccine cargo dynamics in PNP hydrogels. The different physiochemical properties of vaccine components pose a challenge for their controlled and sustained delivery. Vaccine components such as antigens and adjuvants can have extremely distinct polarities, molecular weights and hydrodynamic radii (RH) influencing their diffusivity in the hydrogel network43. CpG-PNP hydrogels are loaded with CpG tethered NPs (CpG NPs, RH˜30 nm) and spike protein (RH=12 nm. Mw=139 kDa). Because of its small molecular weight and hydrodynamic radius, we tethered the CpG nucleotide to the NPs. Unconjugated CpG (RH˜4 nm75, Mw˜7 kDa) is hypothesized to be smaller than the gels' mesh size64 and therefore would rapidly diffuse out of the gel, not allowing for sustained delivery. In order to tune and characterize the dynamics of vaccine exposure expected in the PNP hydrogels, we performed fluorescence after photobleaching experiments (FRAP) (
In order to compare the diffusivity of the CpG-NPs (DNP=0.68 μm2/s) and the spike protein in the gel (Dspike=0.64 μm2/s), to the self-diffusion of the PNP matrix, we calculated Dcargo/Dgel ratios close to 1 (
Humoral immune response to COVID-19 vaccine. In order to investigate if sustained delivery of newly developed CpG NPs would result in a more potent humoral immune response, we subcutaneously (s.c.) vaccinated mice against COVID-19 with spike protein and CpG delivered in different ways (i.e., CpG-NPs and CpG-PNP hydrogels). The spike protein (S) is currently one of the most promising antigens used in SARS-CoV-2 subunit vaccine research. The monomer has a molecular weight of 139 kDa and consists of two subunits, the S1 and S2, S1 contains a receptor binding domain (RBD) which recognizes and binds the cell surface receptor, whereas S2 contains basic elements required for membrane fusion. We therefore decided to fabricate different vaccines containing spike administered either with CpG in soluble form, CpG tethered to the NPs (CpG NPs) or CpG tethered to the NPs and encapsulated in a PNP gel (CpG PNP). For the two CpG NP groups, the CpG-B NPs and CpG-C NPs, mice were administered with 10 μg of spike antigen and 20 μg of CpG adjuvant. Furthermore, as a comparison we included a total of 4 control groups: 2 groups receiving spike adjuvanted with either soluble CpG-B, or CpG-C. 1 group receiving a vaccine adjuvanted with Alum (100 μg) as a clinically-relevant control, and 1 group receiving only the spike antigen with a small concentration of plain NPs added to it as a vehicle control. For the bolus injections, requiring a second dose of vaccination, s.c. priming was followed by the subsequent booster immunization at day 21 (
As previously confirmed via FRAP measurements, the size of both, the monomeric spike protein and the CpG conjugated NPs are ideal for sustained co-delivery from the PNP hydrogel. Consequently, we hypothesized that the implementation of our newly developed CpG NPs loaded in the PNP hydrogel could considerably increase the potency of a COVID-19 vaccine, eliciting a broadly neutralizing antibody response after only one single immunization. In order to investigate this, we s.c administered 2 different CpG PNP hydrogels, a CpG-B PNP and a CpG-C PNP, with double the cargo dose of the bolus groups. We argue that administering a double dose of vaccine (20 μg of CpG and 40 μg of spike protein) would provide, over the course of the 35 days of study, the same total amount of vaccine as the bolus injections, however slowly releasing the antigen and adjuvant in a controlled manner over time. To achieve the 40 μg dose of CpG in the gels, the CpG NPs were mixed with standard PEG-b-PLA NPs.
To ensure that the injected vaccines did not drive systemic cytokine responses that could pose a safety risk, we assessed the early inflammatory cytokines IFN-α (
To assess the humoral response, mice serum was collected every week for 35 days and spike-specific immunoglobulin G (IgG) endpoint antibody titers were quantified every week. Evaluation of total IgG endpoint titers over the first 21 days (
Furthermore, on day 21, before boosting, we conduced neutralization assays (
The data gathered until day 21 is premature to make any conclusions about the final efficacy of the vaccines, however it already seems to appear that the prolonged co-delivery of both the antigen and the adjuvant is critical for eliciting strong and durable immune responses and for eliciting broadly neutralizing antibodies.
In conclusion, we developed a novel adjuvant NPs platform where potent TLR9 agonists are presented on the surface of PEG-PLA NPs. These NPs allowed us to control adjuvant distribution on the NP surface while also exhibiting efficient transportation to the LNs due to their size. We also showed that the density of CpG presentation on the NP surface influenced the activation of TLR9, with 30% CpG NPs exhibiting similar potencies in vitro to free CpG. Additionally, in order to elicit a strong humoral response, improved spatiotemporal control over the vaccine delivery and distribution is needed. Therefore, we loaded our CpG-NPs in PNP hydrogels and showed through an extensive rheological characterization that the integration of CpG-NPs into the polymer network did not affect the overall mechanical properties of the PNP gels. The CpG PNP gels exhibited viscoelastic, thixotropic and shear thinning properties that allowed them to be injected and to maintain a robust solid-like structure after injection. Loading of the CpG PNP hydrogels with spike protein, as the antigen, and CpG-NPs, as the adjuvant, allowed both cargoes, despite of their physiochemical differences, to be largely immobilized by the hydrogel's network. Both vaccine components were therefore able to diffuse at similar rates out of the gel, allowing for a sustained and controlled release. In an in vivo COVID-19 vaccine study, we then showed that CpG PNP hydrogels were critical for eliciting strong immune responses. As such, we demonstrated rapid, more potent and consistent IgG responses of the CpG PNP hydrogel COVID-19 vaccines compared to bolus injections of the CpG NPs vaccines already at day 7. Furthermore, CpG PNP hydrogels led to durable APCs activation, which is necessary for the downstream humoral response, and increased antibody titers. Moreover, CpG PNP hydrogel vaccines were the only vaccine candidates in our study that elicited strong neutralizing antibody responses and protected cells from viral infection.
In sum, this study reports the development of a highly modular adjuvant NPs platform that provides for facile incorporation into PNP hydrogels for prolonged co-delivery with subunit antigens. We showed that these CpG-PNP hydrogels present a tunable platform for enhancing the humoral immune response to a COVID-19 subunit vaccine, contributing to the development of an easily administrable (injectable), stable single dose vaccine against SARS-CoV-2 that is less reliant on the cold-chain.
Exemplary embodiments provided in accordance with the presently disclosed subject matter include, but are not limited to, the claims and the following embodiments:
Although the foregoing disclosure has been described in some detail by way of illustration and example for purposes of clarity of understanding, one of skill in the art will appreciate that certain changes and modifications may be practiced within the scope of the appended claims. In addition, each reference provided herein is incorporated by reference in its entirety to the same extent as if each reference was individually incorporated by reference.
This application claims the benefit of U.S. Provisional Application No. 63/025,845, filed on May 15, 2020, and which is incorporated in its entirety herein by reference.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2021/032575 | 5/14/2021 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63025845 | May 2020 | US |
