Hydrogels for Mucosal Drug Delivery

Abstract

Mucosal tissues are often a desirable site of drug action to treat disease and engage the immune system. However, systemically administered drugs suffer from limited bioavailability in mucosal tissues where technologies to enable direct, local delivery to these sites would prove useful. In this Spotlight on Applications article, we discuss hydrogels as an attractive means for local delivery of therapeutics to address a range of conditions affecting the eye, nose, oral cavity, gastrointestinal, urinary bladder, and vaginal tracts. Considering the barriers to effective mucosal delivery, we provide an overview of the key parameters in the use of hydrogels for these applications. Finally, we highlight recent work demonstrating their use for inflammatory and infectious diseases affecting these tissues.

Graphical Abstract

1. INTRODUCTION

The human body is covered with mucosa amounting to ~400 m² in total surface area, roughly 20 times greater in surface area than the skin. Mucosal layers keep bodily tissues hydrated and function as a protective layer by blocking foreign substances and pathogens. Mucus secreted at epithelial surfaces primarily consists of mucin glycoproteins that form a hydrogel mesh network via disulfide bonds and entanglement.^1,2 When subjected to high physiological shear stresses (e.g., blinking, coughing), mucus undergoes a transition from solid to viscous fluid-like behavior where the viscosity of mucus approaches that of water.^3–5 This allows for the rapid clearance, on the order of seconds to minutes, of mucus and its contents. Thus, drugs locally delivered as a suspension that become entrapped within the mucus gel are likely to be cleared from the body before they exert a therapeutic effect. In disease-affected tissues, further changes in mucin concentration and hydration due to inflammation can greatly affect mucus viscoelasticity and permeability toward therapeutic agents. Ointments and cream formulations can help to slow drug clearance but are often cumbersome to apply topically, particularly in the eye, nose, and oral cavity. To overcome these challenges, sustained drug delivery at mucosal sites can be achieved using hydrogels for local administration. Beyond their ability to act as vehicles for therapeutic cargo, hydrogels also can provide a protective barrier mimicking the native mucosal lining and may have a synergistic benefit in conditions where mucosal barrier integrity has been compromised such as in inflammatory bowel disease (IBD).

Over the past roughly three decades, mucosal delivery using polymers and hydrogel systems has advanced significantly owing to the pioneering work of Claus-Michael Lehr, Kinam Park, Nicholas Peppas, Joseph Robinson, and others.^6–9 Through this work, the principles of mucoadhesion were established that govern how polymeric materials associate to a mucus-coated epithelium.¹⁰ The identification of mucoadhesive biomaterials has enabled the design of advanced hydrogel drug delivery systems to enhance the bioavailability of topical or locally administered drug products.¹¹ While mucoadhesion remains an important design criterion, a new formulation strategy has emerged considering the pathological changes to mucosal epithelium, such as in IBD, where the mucosal interface is altered due to inflammation such that mucoadhesive strategies are not likely to be as effective. In addition to this, new design strategies have emerged that exploit physiological cues to enhance hydrogel mucoadhesion and trigger drug release at sites of injury. The purpose of this review is to highlight both previously known and newly discovered formulation strategies that should be considered in designing hydrogels for mucosal delivery. While hydrogel biomaterials can also be formulated as colloidal particle suspensions, we have focused our discussion on formulation strategies for bulk hydrogels and refer the reader to these recent review articles for additional considerations on microgel drug delivery.^12,13

Each mucosal tissue is unique owing to its functions and required protective mechanisms. To directly interrogate the microscopic properties that give rise to mucus barrier function, microrheology can be used to measure the viscoelasticity experienced by micro- and nanoscaled molecules and pathogens.⁵ Using this approach, the pore sizes of the mucus layer have been found to vary throughout different mucosal tissues ranging from 20–200 nm, 100–300 nm, and 300–500 nm within the gastrointestinal, respiratory, and reproductive tracts, respectively.^14–16 This in turn presents unique challenges in delivering drugs in different mucosal tissues, and understanding these differences is essential to the design of effective hydrogel delivery systems. As such, we discuss in this review the tissue-specific mucosal barrier properties most relevant to hydrogel-based drug delivery before introducing specific formulation strategies that have proven useful in enhancing their therapeutic effect. While many opportunities have been explored using hydrogels in drug delivery applications, we highlight recent studies that demonstrate how hydrogel formulation strategies can be harnessed for local therapeutic delivery in inflammatory and infectious disease applications.

2. MUCOSAL TISSUES

2.1. Ocular.

The tear film, which is about 2–6 μm in thickness, is the outermost layer of the ocular surface, acts as a protective film to underlying epithelia, and provides a smooth refractive surface for clear vision.^17,18 As depicted in Figure 1, the tear film is composed of a three-layered structure: an outermost lipid layer, middle aqueous layer, and innermost mucosal layer. The mucus layer is composed of soluble mucins and gel-forming mucins.¹⁹ The interface between tear film and cornea is composed of the glycocalyx and membrane tethered mucins, which are expressed on the apical surface of the corneal epithelium. These mucins play a vital role in pathogen recognition and clearance. Alterations in the percentages of different mucins can be observed in eye related diseases such as dry eye syndrome, ocular allergies, and conjunctivitis. In addition to defense against pathogens, the mucosal layer of the ocular surface also acts as a barrier to drug absorption. In vitro studies conducted on rose bengal permeation across corneal cell monolayer and stratified corneal cells with mucosal differentiation showed a blockage of rose bengal penetrance in the presence of mucins.^20,21 In addition to the mucosal layer, the epithelium underlying the cornea is another barrier to drug absorption due to tight junctions that prevent paracellular transport.

Figure 1.

Illustration of lipid, aqueous, and mucinous layers of the tear film. Inspired by refs 22 and 23.

A major limitation of ocular drug delivery is the rapid clearance of topically administered drugs. Tear drainage occurs as soon as eye drops are applied through the nasolacrimal duct to remove excess fluid on the ocular surface. In addition, rapid tear turnover of 0.5–2.2 μL min⁻¹ and blinking reflexes of 6 to 15 times per minute lead to drug elimination at the rate of 0.5–1.0 min⁻¹, reducing contact time with the ocular surface. Low contact time, along with permeation barriers such as the mucosal layer and underlying epithelium, leads to a low bioavailability of 0.1 to 5%. Increasing the contact time of a drug by using hydrogel formulations can help to increase absorption through the ocular surface, leading to higher bioavailability. Hydrogels offer a promising solution as a delivery system that can increase the retention time of drugs. However, ideally ocular hydrogel formulations should not blur vision. Several hydrogel delivery systems have been developed and proven effective for topical ophthalmic administration in animal models to treat a range of conditions such as cataracts, glaucoma, and bacterial conjunctivitis.^24–30

2.2. Upper Respiratory Tract.

The nasal cavity consists of three major compartments: the vestibular, respiratory, and olfactory regions. The vestibular region is the outermost region and is composed of nasal hairs, squamous epithelial cells, and ciliated cells.³¹ The middle region is the respiratory region, the most vascularized region, composed of mucus-secreting goblet cells, basal cells, ciliated, and nonciliated columnar cells. The innermost region is the olfactory region, which consists of tubular Bowman’s glands, which secrete a mucus layer, and olfactory sensory neurons. The nasal tissue has historically been used as a route for local drug delivery as it is easily accessible and also highly vascularized. Compared to the other mucosal tissues, the nasal tissue has a low enzymatic concentration, adding to its appeal as a route for drug delivery.³² However, given the high density of ciliated cells and low viscosity of mucus produced in the nostrils, rapid mucociliary clearance poses a significant challenge to intranasal drug delivery. Recently, it was shown that mucus clearance velocity on the nasal tract can be up to 4 times faster (~10–15 mm/min) than that observed in the lower respiratory tract.³³ Thus, the retention of drugs in the nasal tract may be improved using hydrogels that locally adhere and slowly release therapeutics to target cells. As such, it should be noted that hydrogels used as intranasal delivery systems must be designed such that they do not impede normal clearance mechanisms or obstruct airway passages. To reach the lower respiratory tract, micro- and nanoparticulate systems provide a suitable drug delivery approach, and given our focus here on bulk hydrogels, delivery to these regions of the lung is not discussed. Several hydrogel-based formulations have been developed for intranasal delivery of therapeutics to treat conditions primarily affecting the nasal tract such as allergic rhinitis³⁴ as well as other distal tissues for conditions such as diabetes and chronic pain.^35,36 It is also proven to be an effective route for drug delivery to the brain, allowing the blood–brain barrier (BBB) to be bypassed where formulations using thermosensitive and in situ forming hydrogels have improved drug retention and achieved highly efficient nose-to-brain delivery.^37–40 For example, a mucoadhesive chitosan-based hydrogel was intranasally administered to rabbits and demonstrated significantly improved delivery of donepezil hydrochloride, used in the treatment of Alzheimer’s, with a roughly ~2-fold increase in total drug content reaching the brain as compared to orally administered tablets.⁴¹ It has also been shown that nasal immunization produces strong IgA and IgG antibody responses in the nose and other mucosal tissues (e.g., the cervicovaginal tract), making it an attractive route for vaccination against infectious diseases.⁴² As such, hydrogel delivery systems have also proven useful for the intranasal delivery of immunotherapies and vaccines to treat allergic asthma and prevent upper respiratory infections.^43–45

2.3. Oral/Buccal Cavity.

The oral or buccal cavity includes the inner linings of the lips, cheeks, palates, and tongue.⁴⁶ These tissues are composed of multiple layers including the mucus layer, epithelium, and connective tissues.⁴⁷ Epithelial layers express increased keratinization in areas, i.e., the palates and gingiva, subject to mechanical forces from mastication or swallowing.⁴⁶ Mucosal regions of the oral cavity that require more elasticity and flexibility are nonkeratinized. Secreted mucins create a heterogeneous mucus layer and are present within the salivary film.^48,49 Mucus secretions localized within the oral cavity have different affinities for bacteria and can interact with microorganisms individually or in tandem.^48,50 Saliva secretions cover mucosal tissues in a film approximately 70–100 μm in thickness with a clearance rate of 0.3–0.4 mL/min.^51,52

The oral mucosa has a large surface area with high blood supply and vascularization, which aids in rapid drug absorption for systemic uptake. Drug delivery through the oral cavity is sometimes used as an alternative to oral ingestion to bypass drug degradation by the gastrointestinal environment. However, the barrier function of the oral mucosa can be compromised in pathological conditions. For example, keratosis leads to increases in keratinized mucosa tissues and reduces drug permeability. Lesions within the oral cavity are characterized by tissue necrosis resulting in epithelial layer thinning and keratosis/hyperkeratosis.⁵³ Current treatments include topical applications of corticosteroids and analgesics via tablets or patches to treat the affected mucosa tissue. Within the oral cavity, administered drugs are prone to degradation in a high enzymatic environment as well as degradation from experiencing constant mechanical forces. In addition, saliva clearance rapidly removes drugs from the mucosal surface. Mechanical forces from mouth movements also affect topical treatments.⁴⁸ Thus, hydrogel-based delivery systems can be useful to resist salivary clearance and degradation leading to improved bioavailability.^54–56

2.4. Gastrointestinal.

Oral administration for gastrointestinal (GI) drug delivery is the most common mode of drug administration owing to its convenience. The rectal delivery route provides another alternative for local GI drug delivery. Orally administered drugs are swallowed and absorbed in the GI tract, whereas rectally administered drugs allow for local absorption of drugs. The GI tract has a total surface area of 250–400 m² that can be divided into five compartments: the esophagus, stomach, intestine, colon, and rectum. Each compartment is unique in its structure and physiology, posing challenges and opportunities for drug delivery.

The esophagus has the shortest transit time of <1 min. This leads to poor drug absorption, making localized delivery to the esophagus quite challenging. The esophageal epithelium is further lined by a thin mucus layer that is the primary protective barrier against stomach acid reflux and ingested foreign materials including therapeutics. Similar to the change in epithelial thickness in esophagitis, an adherent mucin layer of ~90 μm is observed in Barrett’s esophagus but absent in healthy esophagus.⁵⁷ To facilitate esophageal drug absorption, drug formulations with viscous liquids or gels are given to increase esophageal surface contact time. A study reported reversal of esophagitis in a dog model of Barrett’s esophagus when treated with mucoadhesive hydrogel formulated using extracellular matrix derived from porcine esophageal mucosa.⁵⁸ Another study reported use of a hyaluronic acid-based hydrogel for treatment of esophageal fistulas.⁵⁹

Orally ingested material travels through the esophagus to the stomach. The stomach’s transit time ranges between 1 and 4 h. It is highly acidic with a pH of 1–2 with a high concentration of gastric enzymes, such as pepsin, to aid in food digestion. Despite these harsh conditions, the stomach is protected from self-digestion through a thick mucus layer. As shown in Figure 2, the stomach is covered by two layers of mucus primarily composed of mucin 5AC (Muc5AC), with an unstirred inner layer firmly attached to epithelial cells and a removable outer layer. The mucosal lining of the stomach maintains a pH gradient across its layers with pH 2 at its luminal surface and pH 7 at its epithelial surface to protect the underlying epithelium.⁶⁰ This is achieved through both epithelial cells continuously secreting bicarbonate as well as gel-forming mucins to replenish and stabilize the inner mucus layer of the stomach.⁶¹ Thus, orally administered drugs must withstand these harsh conditions to achieve biopharmaceutical efficacy where drug delivery devices (e.g., capsules) are essential. In addition, osmotic pressure along the GI wall, shear stresses due to the flow of gastric juice, and peristalsis of GI muscles can cause mechanical damage to drug carriers. If the drug withstands the harsh conditions of the lumen, the mucosal layer becomes the next barrier to drug absorption. Gastroretentive systems that prolong gastric residence time are used to allow drug absorption in the upper GI tract and to prolong drug release in a controlled manner. Examples of gastroretentive systems include floating devices,^62–65 expandable dosage forms,^66–69 and mucus or tissue adhesive systems.^70–73 Mucoadhesive systems are limited by the mucus turnover time of 1–2 days and epithelial turnover time of 4–5 days across the entire GI tract.

Figure 2.

Schematic representation of the mucus organization in the gastrointestinal tract. OLM = outer loose mucus, IAM = inner adherent mucus. The mucus thicknesses in different regions are based on previous reports in rats. Figure inspired from refs 61, 74, and 75.

Intestinal transit time typically ranges between 1 to 6 h and has a drastic pH increase to 7.0–7.4. The intestine consists of a single, loosely cross-linked layer of mucus built by mucin 2 (Muc2), as depicted in Figure 2. However, the intestine is kept sterile through continuous flushing of the mucus layer and liquid secretions. The colon has the longest transit time of 1–3 days and a pH ranging between 6–6.7. Like the stomach, the colon has two mucus layers with an inner adherent layer and outer loose layer, both composed of Muc2. Interestingly, the same mucin behaves differently in the small intestine and colon. The inner mucus layer of the colon is formed by densely cross-linked sheets of mucus and is devoid of bacteria achieved through size exclusion.⁷⁶ It is renewed at the epithelial surface by goblet cell secretions and converts to the outer mucus layer at the luminal surface. This outer mucus layer expands 4 to 5-fold in volume and is less viscous as a result of mucin network cleavage by endogenous proteases creating a habitat for commensal microbiota. In pathological conditions such as ulcerative colitis, a defective inner mucus layer can lead to bacterial contact with the epithelium, triggering an inflammatory cascade.^77,78 Intestinal and colonic drug delivery can be achieved through pH dependent, time dependent, and pressure dependent systems by considering the rapid pH change, prolonged transit time, and high-pressure environment of the colon.^79–85 The rectum comprises the last 5 to 6 in. of the large intestine and temporarily stores fecal material before excretion through the anus. Within the rectal environment, an average fluid volume of 1–3 mL is maintained at a neutral pH of 7–8 with minimal buffering capacity. The lumenal facing layer of the rectum is covered with a single layer of mucus and is in direct contact with feces.⁸⁶ Rectal drug delivery offers rapid absorption and a homeostatic environment with relatively low enzymatic activity. Thus, these ideal properties make rectal drug delivery an appealing noninvasive alternative to both local and systemic oral drug delivery. Hydrogels are a promising platform for rectal drug delivery, and several studies were reported based on mucoadhesive and in situ forming hydrogels.^87–90 To provide a target retention time for hydrogels to enable delivery to the GI tract and other mucosal tissues, Table 1 provides approximate transit times of therapeutic drugs across a mucosal layer with maximum thickness on the order of millimeters. It should be noted that the diffusion coefficient and transit times will vary at different mucosal sites depending on mucus thickness and pore size. The rate of drug release from the hydrogel system should also be considered in determining the retention time. However, overall, these values provide a target retention time for hydrogels to deliver the majority of a given therapeutic payload to mucosal sites.

Table 1.

Transit Time of Therapeutics across Mucosal Barriers^a

Molecule	Dmucus (10−7 cm2/s)	tmax, h
Fluorescein	53	0.7
Polypeptide	18	2.1
Lactalbumin	12	3.1
Ovalbumin	8.8	4.2
Fc fragment	6.7	5.6
Fab fragment	6.7	5.6
BSA	5.7	6.6
F(ab′)2 fragment	4.4	8.5
Human IgA	5.1	7.3
Human IgG	2.9	12.9
Human IgM	2.8	13.3
Human S-IgA	0.49	76.2

^aTo estimate target retention times for hydrogel-based delivery systems, we calculate the maximum transit time for a representative set of small molecules and proteins across mucosal barriers. The diffusion coefficient (D_mucus) used for each molecule was previously determined in midcycle human cervical mucus at 25 °C.⁹¹ We assume a maximum mucosal thickness (L) of 1.64 mm, which is the maximum thickness observed in the human body based on endoscopic imaging within the GI tract.⁹² Considering diffusion in the direction perpendicular to the tissue only (i.e., z-direction), we calculate the transit time (t) across the mucosal barrier as t_max = L²/2D_mucus.

2.5. Cervicovaginal.

The vaginal cavity has a large surface area with a dense network of blood vessels, making it an excellent route of drug administration for both systemic and local delivery. This administration route has been traditionally used for the local delivery of antibacterial, antifungal, spermicidal, and labor inducing agents.⁹³ It also allows uterine targeting due to the first uterine pass effect when applied in the upper third of the vagina.^94,95 This can be taken advantage of for hormone replacement therapies to maximize the desired effect and minimize systemic side effects. In recent years, many studies investigated its potential for microbicide delivery to prevent sexually transmitted infections⁹⁶ and for localized treatment of cervical cancer.^97–99 However, the vaginal route is underexplored due to gender specificity and is limited by the possibility of leakage, poor distribution across vaginal walls, and low residence time due to epithelial shedding and mucosal clearance.

Like other mucosal tissues, mucus lining the cervicovaginal canal lubricates the female genital tract along with acting as a first line of defense against the outside environment. Cervical mucus (CM) is secreted by the goblet cells in the crevices of the cervix, travels down to the vaginal cavity, and mixes with host secretions, with cellular debris becoming cervicovaginal mucus (CVM) with a lower viscosity.¹⁰⁰ CM and CVM composition changes throughout the menstrual cycle and allows the movement of sperm toward the ovule depending on the menstrual phase. A cervical mucus plug is formed during pregnancy to maintain the sterility of the uterus from the vaginal canal. CVM is enriched with lipids, salts, immune cells, antimicrobial peptides, and immunoglobulins and hosts the vaginal microbiome, which plays a major role in vaginal health.¹⁰¹ The microstructure and bulk rheology of fresh undiluted CVM exhibits relatively minor changes across a broad pH range of 1 to 9¹⁰² and can provide a similar barrier to drug permeation across pH fluctuation before and after mixing with seminal fluids, which are ~4.5 and ~7.2, respectively.¹⁰³ The average pore size of CVM is 340 ± 70 nm, which is higher than human viruses. However, viruses and bacteria are trapped in the mucus barrier through interactions with mucin glycans, antibodies, or other proteins present in the mucus network.¹⁴ They are shed out of the body during mucosal clearance. However, mucosal barrier function is compromised due to microbial dysbiosis as mucolytic enzymes are released in conditions such as bacterial vaginosis.^104,105 This allows the viruses and bacteria to pass through the mucus layer and reach the underlying epithelium to cause infections including HIV.¹⁰⁶ The vaginal route is particularly effective in such cases as the drug is present in an effective concentration at the site of pathogen entry. Different dosage forms such as gels, suppositories, rings and films are preferred by women in different regions based on individual or partner choice, cultural norms and socioeconomic status of the region.^107,108 Among those, hydrogel formulations can be designed to provide ease of application, increased drug residence time, and uniform distribution in the vaginal cavity.

2.6. Urinary Bladder.

The urinary bladder temporarily stores urine and facilitates voiding through urination.¹⁰⁹ The luminal surface of the bladder is made of a mucosal layer called the urothelium, a stratified, transitional epithelium that acts as a thick barrier to protect the bladder walls from the toxic contents of urine. The apical membrane of urothelium is covered with glycocalyx containing both adherent and soluble glycoproteins, glycolipids, galectins and proteoglycans. This surface mucus layer has nonspecific antiadherent properties to defend against invading pathogens,¹¹⁰ and its disruption has shown to increase bacterial infections.¹¹¹ Different drug delivery routes and modalities are studied to treat bladder related disorders such as bladder cancer, interstitial cystitis, and urinary tract infections. Intravesical delivery, where drugs are directly instilled into the urinary bladder via catheter, is a common route for systemically administered drugs but suffers from poor drug distribution to the bladder. In addition, intravesical delivery achieves local delivery through minimizing systemic toxicity. However, the efficacy of intravesical delivery is limited by poor urothelium permeability, continuous dilution by urine, and subsequent disposal out of the bladder.¹¹²

Hydrogel systems are increasingly reported for intravesical delivery as use of hydrogels minimizes drug dilution and offers other advantages, including mucoadhesion and floating properties, which can enhance drug absorption, retention, and sustained release in the bladder. A study reported an in situ gelling, mucoadhesive hydrogel formulated with high molecular weight chitosan and β-glycerophosphate for treatment of bladder cancer that showed superior resistance to urine washout when tested ex vivo using porcine urinary bladder.¹¹³ Another study reported floating hydrogels made of Poloxamer 407 (P407) for the treatment of bladder related diseases. This gel was prepared by simple shaking and locking of the bubbles that were formed as a result of P407 foamability. The hydrogel demonstrated prolonged residence time and continuous drug release for up to 10 h when tested in vivo using a rabbit acute bladder injury model.¹¹⁴ Several other hydrogels investigated for use as intravesical drug delivery systems and therapeutic bulking agents were reported as promising drug delivery systems for treating urinary bladder disorders.^115–118

3. FORMULATION STRATEGIES

3.1. Mucoadhesion.

Mucoadhesion has been widely exploited to enhance the retention of drug delivery systems in mucosal tissues.^119–121 For these applications, there are several factors that one should consider to achieve mucoadhesion. Here, we will briefly describe general considerations for mucoadhesive biomaterial design and would recommend several recent review articles for details on the theory and mechanisms of mucoadhesion.^11,122,123 It has been found in prior work that polymers containing mucin-like chemical moieties (e.g., NH₂, OH, SO₄) generally favor mucin-biomaterial interactions via hydrogen bonding.¹²² Cationic polymers may nonspecifically interact with mucus through charge-mediated interactions with net-negatively charged mucins.¹¹ Thiol-containing biopolymers may form covalent bonds to mucin cysteine residues.¹²⁴ The degree of swelling can also impact mucosal adhesion as hydration aids in expansion of polymer chains to expose adhesive functional groups and facilitate entanglement with mucin biopolymers.¹²⁵ Increasing polymer molecular weight also favors interpretation and chain entanglement with mucus. Conversely, increasing hydrogel cross-linking density can limit interfacial interpenetration with the mucus gel layer. However, overall, a consensus has yet to be reached on specific material properties for optimal mucoadhesion due, at least in part, to a lack of standardization in measuring mucoadhesive capacity.¹²⁶ To address this, a study was performed to directly compare mucoadhesion of different natural and synthetic polymers to excised porcine small intestin,e providing insights into how material type, functional groups (e.g., cysteine-modified), and preparation method influence performance.¹²⁷ It should also be noted that changes in pH and ionic strength can significantly alter mucin–polymer interactions, making it critical for biomaterials designed for mucoadhesion to be tested at physiological conditions.

Carbomers such as Carbopol934P, Carbopol974NF, and polycarbophil (NoveonAA1) can be used as excipients to improve mucoadhesion in polymer films. Using these excipients, a systematic study was conducted to determine formulation requirements for optimal mucoadhesion of poly(vinyl alcohol) (PVA) and polyvinylpyrrolidone (PVP) films. Formulations using PVP showed higher mucoadhesion than those utilizing PVA, likely due to more favorable intermolecular interactions between PVP and carbomer excipients. Mucoadhesive PVP films were further optimized at pH between 5.5–6.8 and carbomer concentrations between 2 and 10 wt % where polycarbophil most significantly enhanced mucoadhesive strength.¹²⁸ Chitosan is another widely used mucoadhesive polymer as it contains primary amine groups that drive mucoadhesive interactions that are further reinforced by hydrophobic interactions and hydrogen binding.¹²⁹ To simulate the different ionic environments within the GI tract, Lehr et al. showed a significant decrease in the mucoadhesivity of chitosan in artificial gastric fluid (pH 1.2) as compared to artificial intestinal fluid (pH 7.5), presumably due to chitosan’s solubility in acidic conditions. Interestingly, polycarbophil was also shown to significantly enhance the mucoadhesive capacity of chitosan under all conditions tested.¹³⁰

Mucoadhesive polymers can also be chemically modified to enhance their performance. For example, alginate, a natural biopolymer known to possess mucoadhesive properties, was modified with cysteine (SH) groups. SH-modified alginate possessed a ~4-fold higher work of adhesion when applied on freshly excised mucosa as compared to unmodified alginate.¹³¹ Similarly, hyaluronic acid (HA) functionalized with sulfhydryl cysteine groups (HA-SH) significantly outperformed unmodified HA in terms of tensile strength, total work of adhesion, and maximum detachment force in nasal, buccal, and vaginal tissue.¹³² To increase the mucoadhesive capacity of chitosan, Kim et al. utilized a mussel-inspired adhesion approach and synthesized catechol-functionalized chitosan (Chi-C).¹³³ As shown in Figure 3, catechols can oxidize to form irreversible covalent bonds with thiols and amines within mucus, leading to greater bioavailability and retention time of Chi-C in the GI tract following oral administration.

Figure 3.

Mussel-inspired bioadhesion as a strategy to enhance mucoadhesion. (a) Schematic showcasing mechanisms of binding using catechol conjugated chitosan (Chi-C). Catechol moieties increase consolidation through covalent bonding with the amines and cysteines present in the mucins. (b) In vivo imaging system (IVIS) imaging of extracted organs of mice and (c) relative fluorescence intensity in the GI tract following oral administration of fluorescent labeled chitosan (Chi-FITC), poly(acrylic acid) (PAA-FTSC), and Chi-C (Chi-C-FITC). Reproduced with permission from ref 133. Copyright 2015 Elsevier.

Additional functional groups have more recently been investigated for enhancing mucoadhesive properties. Modified polymers with conjugated functional groups have been used to take advantage of the thiol-mediated chemistries to promote mucoadhesion by covalently binding to cysteine residues on mucin glycoproteins.¹³⁴ Maleimide-functionalized PEGylated polymers such as alginate-polyethylene glycol-maleimide (Alg-PEGM) and poly(lactide-co-glycolide)-block-polyethylene glycol-maleimide (PLGA-PEG-Mal) showed a propensity to aggregate with mucins and showed increased retention to the urinary bladder mucosa, respectively.^135,136 Acrylate-chitosan and polyethylene glycol diacrylate polymers have also been explored.^137,138 Similarly, methacrylate-functionalized materials have been demonstrated to exhibit enhanced adhesion to mucosal tissues.^139,140 Furthermore, designing copolymer hydrogels containing acrylate and methacrylate groups may strengthen mucoadhesion.^141,142 In addition, aldehyde-functionalized polymers have been known to enhance mucoadhesion. More recent studies have synthesized aldehyde copolymers with mucoadhesion force to mucosa comparable to that of chitosan and other functional groups.^143,144

3.2. Viscoelasticity.

Mucus viscoelasticity is directly related to mucus barrier function. In the respiratory tract, healthy mucus has low viscoelasticity to enable clearance by cilia.¹⁴⁵ In many pathological conditions, mucus is marked by increased viscosity due to the hyperexpression of mucins, mucus dehydration, and accumulation of bacterial and cellular debris that alter mucin content and composition. In the lung, this leads to prolonged accumulation of mucus gels and provides an environment permissible to chronic bacterial infections. Similarly in drug delivery applications, hydrogel viscoelasticity can be tuned to slow their clearance from mucosal tissues and extend drug release time. Increasing hydrogel elasticity can be used to increase flow resistance, reduce clearance rate, and enhance drug contact time.^146,147 For example, the addition of alginate as a gelling agent to a hydroxypropyl methylcellulose (HPMC) hydrogel was demonstrated to increase hydrogel viscosity and thereby increase drug retention and availability for ophthalmic drug delivery.¹⁴⁸ Increased elasticity also enables hydrogels to better withstand external mechanical forces that are exerted on native tissues. Chitosan composite hydrogels showed resistance to deformation when exposed to stresses.¹⁴⁹ Conversely, decreased viscosity enhances hydrogel flow, which ideally is used for the injection at target tissues. Due to having a low viscosity, hyaluronic acid and chitosan-based hydrogels are used for submucosal injections.^150,151

3.3. Stimuli Sensitive.

Stimuli responsive hydrogels, also called smart hydrogels, respond to environmental stimuli by dramatically changing their structural and mechanical properties to allow in situ gelation or therapeutic release on demand. Environmental stimuli for hydrogel delivery systems can be from endogenous sources such as pH, temperature, enzymes, and redox conditions or exogenous sources such as light, ultrasound, electric field, and magnetic field. Hydrogels engineered to respond to specific stimuli allow precise spatiotemporal control of drug delivery to help minimize off-target effects, reduce inherent toxicity of drugs due to prolonged tissue exposure, and improve drug efficacy.

To design hydrogels with smart on-demand drug release capabilities, drugs are either covalently attached to the hydrogel backbone via a stimuli sensitive linker or physically entrapped in a hydrogel that has a stimuli sensitive backbone. For example, a plant-based antioxidant compound (rutin) was attached to the hydrogel carrier via borate ester linkage for pH and reactive oxygen species (ROS) responsive drug release for the treatment of IBD.¹⁵² In another study, an HIV fusion inhibitor (Enfuvirtide) was attached to the hydrogel backbone via a prostate specific antigen (PSA) labile linker for seminal fluid responsive delivery to prevent heterosexual HIV transmission in women.^153,154 Carboxymethyl chitosan-based hydrogels show pH responsive swelling, and a study has shown higher bovine serum albumin (BSA) release in basic pH as opposed to acidic pH where the BSA was physically entrapped in the hydrogel.¹⁵⁵ In another study, salmon calcitonin (sCT) was physically entrapped in dextran hydrogels for colon specific delivery, where higher sCT release was observed in simulated colonic fluid compared to phosphate buffer solution due to degradation of the hydrogel backbone by dextranases.¹⁵⁶ Some amount of drug leaching is observed from hydrogels with physically entrapped drugs even in the absence of stimuli, which can be a drawback for some applications.

In addition to nonstimuli dependent chemical strategies to produce hydrogels in situ,¹⁵⁷ thermosensitive hydrogels have been extensively investigated for mucosal delivery.^158–162 Solutions of polymers such as pluronic F127 and poly(N-isopropylacrylamide) transition from liquid to solid at or above body temperature to enable in situ gelation for uniform spreading and ease of application on mucosal surfaces. pH and ion concentration were also studied as triggers for in situ gelation at the ocular surface.^163–166 The difference in pH across the upper and lower GI tract can also be used for localized delivery to the intestine with pH triggered hydrogel delivery systems.^167–169 Stimuli responsive delivery can also be used to increase penetration across physiological barriers. Iontophoretic hydrogels were studied for delivering therapeutics across deep layers of the eye, which is hard to achieve noninvasively.¹⁷⁰ Another interesting application of stimuli sensitive hydrogels is gastric resident dosage forms with controlled residence time. A tough triggerable hydrogel (TTH) was developed that can be ingested orally and swells significantly after ingestion, enabling prolonged gastric residence and controlled drug release. As shown in Figure 4, TTH delivery systems are cross-linked using reversible Ca²⁺ ionic bonds and disulfide bonds, which can be triggered using biocompatible chelators or reducing agent to dissolve rapidly for quick transit through the GI tract.⁶⁸

Figure 4.

Design and function of tough triggerable hydrogel (TTH) platform for delivery to the GI tract. (a) Overview of TTH design showing hydrogel backbone synthesized by separately cross-linking alginate and polyacrylamide using triggerable Ca²⁺ ionic and disulfide bonds. (b) TTH triggered from gel to viscous solution by EDTA and GSH. (c) Photographs of initial TTH strip before administration in the pig gastric cavity and the retrieved TTH strips from control and triggered pig after 1 h residence. Reprinted with permission from ref 68 under the terms and conditions of CC-BY open access license.

3.4. Hypotonic.

Fluid absorption and secretion at mucosal surfaces can occur through gradients in tonicity and osmotic driven flow. Delivery of hypertonic formulation leads to fluid secretion into the lumen and subsequent product leakage, whereas hypotonic formulations allow rapid fluid absorption and uniform product distribution at the mucosal surface. In addition, usage of hypertonic vehicles has also shown toxicity due to disruption of epithelial integrity in human colon epithelium and vaginal epithelial explants.^171,172 Thus, hypotonic formulations may offer improved safety and better therapeutic efficacy. For example, doxorubicin administered in isotonic solution (PBS) has shown vaginal surface coverage of 23%, whereas drug administration in hypotonic solution (ultrapure water) showed 83% vaginal surface coverage.¹⁷³ However, improved distribution and rapid absorption lead to poor retention in the vagina, which is not ideal for local delivery. To address this, mucus-penetrating nanoparticles can be formulated and administered in hypotonic solution, leading to improved distribution as well as retention due to slow release of the drug following uniform spreading and accumulation of nanoparticles at the vaginal epithelium.¹⁷³ Similarly, widespread and highly uniform nanoparticle distribution was observed in the colon when delivered locally as a hypotonic enema formulation.¹⁷⁴ Hence, enhanced drug distribution, tissue absorption, and retention can be expected if drugs are mucosally administered in hypotonic hydrogels. Accordingly using thermoreversible Pluronic hydrogels, a hypotonic eye-drop hydrogel formulation was developed for drug delivery to the eye. In this work, it was found that Pluronic gels spread uniformly at the ocular surface due to osmotically driven water absorption. This comfortably fits under eyelids, holds drug at the ocular surface over an extended period, and resists clearance from blinking.¹⁷⁵

3.5. Floating Hydrogels.

Floating hydrogel systems have been used to increase residence times in the gastric and urinary bladder regions, as floating hydrogels are resistant to stomach and bladder emptying. These systems float on top of the stomach and bladder contents, allowing for controlled drug release and improved bioavailability. Several approaches have been developed to formulate floating hydrogel delivery systems. One such system involves using a mixture of Poloxamer 407 (P407) and NaHCO₃ for the intravesical delivery of adriamycin to treat bladder cancer. In this system, a liquid mixture of P407, NaHCO₃, and adriamycin is instilled into the bladder through a catheter, which then gels in situ at body temperature. When the urine is acidified, NaHCO₃ forms CO₂ bubbles at the hydrogel surface, enabling the hydrogel to float.¹⁷⁶ Perfluoropentane has also been used to produce temperature-sensitive floating hydrogels. Perfluoropentane has a lower boiling point than body temperature (29 °C) and converts to a gas state in the bladder, causing the hydrogel to swell and float.¹⁷⁷ Emulsion gelation is another technique used to make floating hydrogels. Here, oil droplets are encapsulated within hydrogels during cross-linking to decrease the hydrogel density in aqueous solution to form floating hydrogels.^178,179 Other reported floating hydrogel designs include hydrophobic hydrogels, superporous hydrogels, and hydrogel rafts.^{63,65,180–182}

3.6. Inflammation-Targeted Adhesion.

At sites of inflammation in diseases like IBD, it has been found that the mucus barrier is disrupted and inflammatory lesions are left unprotected from potentially harmful microbes in the gut.¹⁸³ At these sites of disease, the epithelium is devoid of mucus and acquires a net-positive charge primarily due to the accumulation of positively charged proteins such as transferrin and antimicrobial peptides.¹⁸⁴ As a result, an emerging design parameter has been in the use of hydrogels (and other biomaterials) that specifically target these sites of inflammation. This can be achieved with the use of net-negatively charged hydrogels that will bind to the net-positively charged inflammatory lesion. In mouse models of colitis, an ascorbyl palmitate (AP) microfiber hydrogel possessing a net-negative charge was found to be retained within the colon for up to 12 h, as presented in Figure 5.¹⁸⁵ Of note, these gels were poorly retained in healthy animals due to their lack of mucoadhesion. Thus, this presents a new strategy using hydrogel systems to localize the release of therapeutic compounds at disease sites and further minimize off-target effects. While primarily demonstrated in the GI tract to date, future studies should determine how broadly applicable this approach is to other inflamed mucosal tissues.

Figure 5.

Mode of action and in vivo retention of inflammation targeting (IT) hydrogels. (a) Schematic of ascorbyl palmitate hydrogel fiber formation and (b) adherence of negatively charged IT hydrogel fibers to positively charged inflamed epithelium. (c, d) Preferential adhesion of IT hydrogel fibers to inflamed colon epithelium ex vivo and in vivo in an induced (dextran sodium sulfate (DSS) and genetic (T-bet^−/−Rag2^−/−, TRUC) mouse model of IBD. Reproduced with permission from ref 185. Copyright 2015 The American Association for the Advancement of Science.

4. APPLICATIONS IN INFLAMMATORY AND INFECTIOUS DISEASES

4.1. Mucosal Delivery of Small Molecules.

Local delivery of anti-inflammatory drugs to mucosal tissues can reduce systemic exposure and may help avoid side effects (e.g., immunosuppression, hepatotoxicity) by reducing the dose required. Hydrogel-based delivery systems can greatly improve outcomes in IBD and other inflammatory conditions by locally concentrating the therapeutic payload within mucosal tissues. In addition to their inflammation-targeting properties, AP hydrogels are innately amphiphilic, making them effective for loading of lipophilic, anti-inflammatory compounds. This amphiphilicity enables the self-assembly of micelles into a microfiber hydrogel and encapsulation of lipophilic dexamethasone. As noted above, these hydrogels adhere to inflamed murine intestinal mucosa where sustained release of dexamethasone significantly reduces colitis severity.¹⁸⁵ A similar strategy has been used to formulate hydrogels for the encapsulation and buccal delivery of dexamethasone.¹⁸⁶ Mometasone, another lipophilic corticosteroid, has been delivered through encapsulation and release via thermosensitive copolymer hydrogel microspheres to the sinonasal cavity. Thermosensitivity of the copolymer hydrogel system allowed the hydrogel to conform to the complex architecture of the sinonasal cavity before gelation at body temperature and adherence to the mucosal layer. Treatment with mometasone-loaded hydrogels reduced inflammation in a rabbit model of chronic rhinosinusitis.¹⁸⁷ This prior work demonstrates how amphiphilic hydrogels can be effectively used to deliver lipophilic drugs at targeted sites that otherwise would have poor bioavailability and/or greater metabolic clearance.

Antibiotic misuse and overuse is a widespread issue leading to antibiotic resistant and multidrug resistant bacteria. In addition to this, systemic administration of antibiotics has many drawbacks including low bioavailability and unwanted side effects such as drowsiness, diarrhea, and nausea.¹⁸⁸ To circumvent these challenges, there has been a large body of work using hydrogels for antibiotic delivery to mucosal tissues. A general strategy employed is to load the antibiotic into a carrier such as liposomes or solid lipid nanoparticles (SLNs) and then load the carrier into a hydrogel for extended release. Chen et al. successfully loaded clarithromycin (CAM) into liposomes, which were then loaded into a 4-arm PEG hydrogel.¹⁸⁸ For in vivo studies, CAM-liposome loaded PEG hydrogels were injected into the maxillary sinus of rabbits to treat acute bacterial rhinosinusitis (ABRS). They reported a significant decrease in the burden of Streptococcus pneumoniae with minimal signs of inflammation or other side effects, as shown in Figure 6. Hydrogels have also been used to deliver antibiotics to treat oral periodontitis disease.^189–192 Interestingly, Wróblewska et al. compared the efficacy of alginate hydrogels to that of organogels (nonaqueous, lipophilic) and bigels (biphasic hydrogel and organogel) loaded with metronidazole (MZT).¹⁹² Alginate hydrogels were found to have the highest antibacterial activity, cumulative permeation through porcine buccal mucosa, and retention time with the mucosa. Hydrogels have also been formulated for delivery of antifungals to mucosal sites. To target Candida albicans in the oral cavity, Şenel et al. successfully loaded chlorhexidine gluconate (Chx) into chitosan hydrogels.¹⁹³ Chx-loaded gels achieved retention times over 3 h and showed significant in vitro antifungal activity against C. albicans. To target C. albicans in the vaginal cavity, Mirza et al. loaded itraconazole into SLNs and then into a gel composed of a mucoadhesive excipient (Carbopol 934) and Pluronic F127.¹⁹⁴ The hydrogel formulations outperformed a commercially available antimicrobial cream (Gynazole), sustaining microbicidal activity over 21 days in a rat model of C. albicans infection.

Figure 6.

Hydrogel-based mucosal drug delivery for bacterial rhinosinusitis. (a) Schematic of the hydrogel design and (b) in vivo administration process through local injection to treat acute bacterial rhinosinusitis in rabbits. (c, d) RT-PCR and Western blots showing the decrease in inflammatory markers with the use of CAM liposome loaded gels in vivo. Reproduced with permission from ref 188. Copyright 2020 John Wiley and Sons.

4.2. Mucosal Delivery of Biologics.

Biological therapeutics, such as monoclonal antibodies, have increased dramatically in recent years with rapid advances in recombinant protein production and genetic engineering. However, the bioavailability of parenterally administered biologics at the mucosal tissues is generally low and often requires administration of very high doses to treat diseases, such as IBD. Recent evidence also suggests mucosal delivery of vaccines may offer greater protective benefits where new strategies for intranasal delivery are desired. This motivates the use of hydrogels to enable low dosage treatments of glycan, protein, or nucleic acid drugs through direct delivery to the mucosal surface. Biological cargoes are also susceptible to rapid enzymatic degradation once administered where hydrogel carriers can provide additional protection.

In a previous study, hydrogels were shown to be effective in the local treatment of IBD using heparin, a naturally occurring glycosaminoglycan and anticoagulant. Specifically, heparin was loaded into an injectable protein hydrogel (HEP-Ag-BSA) and delivered rectally in an induced mouse model of colitis. Due to its net-negative charge, HEP-Ag-BSA hydrogels displayed preferential binding to inflamed tissue and local sustained release of heparin, leading to hindrance of inflammation and microthrombosis that catalyzes IBD pathogenesis, as shown in Figure 7.¹⁹⁵ For vaccine applications, an intranasal gel was developed for delivery of broadly acting subunit influenza vaccine. Intranasal gel precursors consisting of branched polyethylenimine (PEI) and oxidized dextran (Ox-DEX) were administered by dual-barrel syringe, and PEI-Ox-DEX hydrogel was formed by in situ gelation via Schiff’s base cross-linking. Antigens were loaded into the gel by mixing with gel precursor solutions. PEI-Ox-DEX hydrogel has shown prolonged mucoadhesion and nasal retention, leading to robust mucosal immune responses.¹⁹⁶ A similar study was conducted to evaluate the effect of chitosan-hydrogel vaccine against influenza infection in mice. Mice intranasally immunized with in situ gelling chitosan hydrogel loaded with influenza-specific antigens demonstrated robust mucosal immune responses and protection after challenge with influenza virus.⁴⁵ Hydrogels have also been shown to be effective in delivering gene therapy for ocular wound repair. In this study, Connexin43 antisense oligonucleotide (ASON) was delivered to rats after corneal injury using a chitosan-based in situ gelling system. This study demonstrated the effective delivery of the ASON to the wounded tissue.¹⁹⁷ Gene therapy can also be achieved by loading the hydrogel with nucleic acid-loaded nanoparticles (e.g., liposomes, polyplexes) as cargo to further enhance mucosal gene delivery.¹⁹⁸

Figure 7.

Injectable protein hydrogels for mucosal delivery of biologics in IBD. (a) Schematic showing gelation of BSA-SH, heparin, and Ag⁺ through disulfide, coordination, and ionic bonds. (b) IVIS imaging of rhodamine labeled HEP-Ag-BSA treated colons showing lesion adhesive property at inflamed site. (c) Illustration of HEP-Ag-BSA administration and its effect on mouse colon with DSS induced colitis. (d) Comparison of colon lengths following in vivo treatment with HEP-Ag-BSA, demonstrating inflammation modulation capability. Reprinted with permission from ref 195 under the terms and conditions of CC-BY open access license.

5. CONCLUSION

Delivery of therapeutics to the mucosa can have tremendous therapeutic effect for individuals with chronic inflammatory diseases and difficult-to-treat infections. A local delivery approach may also spare patients from unintended side effects and reduce overall costs by lowering dose requirements. To make mucosal delivery a viable option in the clinic, hydrogels may help to address the challenges associated with overcoming physiological barriers to effective local drug delivery. Using the formulation strategies outlined here, hydrogels can be designed to encapsulate a wide array of therapeutics, retain the payload at the mucosal site of interest, and release the drug to target cells in a sustained and/or stimuli responsive manner. For inflammatory and infectious disease applications, future work could explore the use of hydrogels for mucosal delivery of cellular,¹⁹⁹ extracellular vesicle,²⁰⁰ and microbial therapies.²⁰¹ In addition, hydrogel-based systems with intrinsic regenerative and/or immune-modulating properties could be particularly useful as mucosal therapies for these applications.^202,203 With careful consideration of the unique challenges to effective hydrogel delivery in mucosal tissues, abundant opportunities exist to create new highly efficacious treatments.