Release and Targeting Mechanism of Nanomedicine Delivery System in Inflammatory Bowel Disease

Abstract

Inflammatory bowel disease (IBD), encompassing Crohn’s disease (CD) and ulcerative colitis (UC), arises from multifactorial pathogenesis involving colonic barrier dysfunction, gut microbiota dysbiosis, and aberrant immune regulation. The complexity of these mechanisms poses substantial challenges to achieving curative outcomes, necessitating long-term maintenance therapies. Conventional therapeutic agents (5-aminosalicylic acids, glucocorticoids, immunosuppressants, and biologics) are constrained by inherent limitations including poor solubility, chemical instability, non-specific biodistribution, and off-target toxicity, which collectively compromise therapeutic efficacy and safety. To address these challenges, nanomaterial drug delivery systems (NDDS) have emerged as a transformative strategy to enhance drug solubility, achieve spatiotemporal control, and improve intestinal targeting while reducing systemic exposure. Building upon our prior review that established a taxonomy of NDDS, this review outlines the in vivo delivery mechanisms of nanomedicine delivery systems for IBD, with the organic integration of drug-loading strategies, drug administration routes, stimulus-responsive release, and targeting mechanisms as the core framework. By systematically integrating these multidimensional mechanisms throughout the entire drug administration lifecycle, this review constructs a theoretical framework that combines macroscopic vision with mechanistic depth, laying the foundation for the rational design of next-generation intelligent NDDS and providing systematic theoretical support for improving the treatment efficacy of IBD, minimizing systemic toxicity, and promoting the clinical translation of NDDS.

Graphical Abstract

Introduction

Inflammatory bowel disease (IBD) is a chronic non-specific intestinal disease, including ulcerative colitis (UC) and Crohn’s disease (CD).^1,2 IBD has become quite severe; the prevalence of IBD has been rapidly increasing in developing areas such as Asia, Eastern Europe, and Africa over recent years, attracting considerable attention.^3,4 Even worse, IBD has a high risk of developing into intestinal cancer.^5,6 Although the pathogenesis and etiology of IBD are still unclear, increasing evidence confirms that many influencing factors, including oxidative damage, intestinal mucosal barrier damage, gut microbiota disorders, immune responses, as well as genetic and environmental factors, are closely related to IBD.^7–10 Therefore, various strategies for treating IBD have been developed, mainly focusing on antioxidant stress, suppression of immune response, and regulation of intestinal microenvironment.¹¹

At present, a permanent cure for IBD has not been established, so lifelong medication is usually necessary to maintain a healthy quality of life. Traditional treatments for IBD are designed to alleviate inflammation by modulating the immune response through the use of medications such as 5-aminosalicylic acid (5-ASA), corticosteroids, immunosuppressants, antibiotics, and anti-TNF biological therapies.^12–14 Surgery is usually performed when medication is difficult to treat and during disease outbreaks to improve the medical condition of these IBD patients.¹⁵ However, conventional therapeutics for IBD face three major clinical limitations, including: (i) Physicochemical instability: The harsh environment of the gastrointestinal tract (low pH, digestive enzymes) leads to drug degradation.¹⁶ (ii) Poor bioavailability and rapid metabolic clearance: Insufficient absorption and hepatic metabolism necessitate high daily doses to achieve therapeutic effects, increasing treatment burden and systemic exposure.^15,17 (iii) systemic toxicity caused by nonspecific biodistribution. For example, corticosteroids exhibit short-term and long-term side effects on healthy cells, leading to hypertension, hyperglycemia, osteoporosis, glaucoma, depression, and many other diseases,¹⁸ which immunosuppressive therapy increases susceptibility to infections and malignant tumors.^19,20 These three limitations collectively define a critical clinical need: a therapeutic approach that selectively delivers medications to inflamed intestinal tissues while minimizing systemic exposure and off-target toxicity. Advanced nanomaterial drug delivery systems (NDDS) have emerged as a transformative strategy to address this clinical challenge. Research indicates that NDDS enhance IBD treatment by improving drug bioavailability and targeting, protecting drugs from gastrointestinal acidity, and concentrating drugs at inflammation sites, thus boosting therapeutic efficacy and minimizing exposure to healthy tissues.^16,21–23 From this, it can be seen that targeted NDDS improves IBD treatment by delivering drugs directly to inflamed tissues, minimizing exposure to healthy tissues and reducing side effects.^24–26

Building upon our previous review that systematically categorized various nanocarriers (eg, inorganic nanoparticles, liposomes, polysaccharide-based carriers) for IBD therapy,²⁷ this review extends the scope to subsequent formulation considerations in nanomedicine development. While many existing articles focus on specific types of new materials or isolated delivery mechanisms, this work directly links intestinal pathophysiology to carrier function. Specifically, we systematically examine the interconnected roles of drug-loading methods, administration routes, responsive release, and active targeting. We analyze the practical applicability and limitations of physical versus chemical drug loading for different therapeutic agents, and compare oral, colorectal, and injectable pathways under different disease conditions. Furthermore, we discuss how stimuli-responsive mechanisms (eg, redox and pH triggers) and targeting strategies (eg, charge- and receptor-mediated approaches) can be integrated to achieve site-specific action in the inflamed gut. By connecting these delivery steps into a cohesive sequence, this review aims to provide a practical and comprehensive approach for the rational design of mucosal nanomedicines, supporting more consistent local delivery while minimizing systemic side effects.

The Way IBD Nanomaterial Carriers Bind to Drugs

The drug binding and administration methods of NDDS for IBD have undergone significant evolution. In terms of drug binding, physical encapsulation (such as hydrophobic interactions of liposomes)²⁸ was mainly used in the early stages, followed by chemical coupling (pH/enzyme sensitive bonds)²⁹ and prodrug strategies.³⁰ In recent years, there has been a trend towards multifunctional precise delivery (biomimetic carriers, nucleic acid delivery, probiotic delivery).^31,32 The drug delivery mode has been gradually optimized from traditional oral preparations (enteric coating)³³ to local delivery (enema nano gel)³⁴ and advanced intelligent delivery systems, such as magnetic micro robots or intelligent oral systems.³⁵ Different combination methods are linked with different administration methods to achieve precise delivery.

The material types of NDDS (such as polymers, liposomes, inorganic nanoparticles, etc.) directly affect the selection of drug loading methods and final delivery efficiency due to their differences in chemical properties and structural characteristics. The combination of nanocarriers and drugs mainly includes physical drug loading and chemical drug loading. Physical drug delivery relies on non-covalent interactions (such as hydrophobic interactions, electrostatic adsorption, van der Waals forces, or encapsulation) to encapsulate or adsorb drugs onto nanocarriers (such as liposomes, polymer micelles, mesoporous silica), while chemical drug loading links drugs and nanomaterials through covalent or ionic bonds, such as esterification, amidation, disulfide bonding, etc. This section provides a detailed overview of the methods for combining nanomaterials with drugs in the following section.

Physical Bond Drug Loading Method

Physical binding utilizes non-covalent interactions, typically under mild conditions, offering easy preparation but posing a risk of drug leakage. Common physical strategies, such as encapsulation/embedment and coating, both enhance drug stability, solubility, and bioavailability. Encapsulation strategies are generally more conducive to achieving long-term sustained release. In contrast, approaches involving covalent conjugation or surface functionalization are better suited for precise controlled release or targeted delivery, respectively.

Encapsulation refers to the drug being enclosed in an independent cavity (eg, liposomes, vesicles) inside the carrier, forming a core-shell structure, where the drug is completely enveloped by a layer of carrier material (Figure 1A). This approach can significantly improve drug stability and targeting in IBD treatment. Common preparation methods include the thin-film hydration method and certain double emulsion methods. Taking lipid nanomaterials as an example, hydrophilic drugs (such as 5-ASA) are soluble in the internal aqueous phase of liposomes, while hydrophobic drugs (such as budesonide) are encapsulated within the lipid bilayer, utilizing the natural intestinal adhesion of liposomes to enhance their enrichment at inflammatory sites.³⁶ Aib et al designed pH-sensitive liposomes encapsulating mesalazine in the water core and curcumin in the lipid bilayer. The study reported high encapsulation efficiencies for both drugs, with this drug achieving 78.3 ± 4.2% for mesalazine and 84.7 ± 4.7% for curcumin. Regarding drug release kinetics, 10% mesalazine and 7% curcumin were released at pH 1.2, followed by a sustained release of 88% mesalazine and 88% curcumin at pH 7.4 after 24 hours. The findings suggest that these liposomes can reduce drug-related side effects, decrease dosage, and provide effective treatment strategies for UC.³⁷

Figure 1.

Binding mode of nanodrug carriers. (A) Schematic illustration of encapsulated drug loading, where drug molecules (yellow dots) are concentrated and encapsulated within the core of the carrier material (yellow sphere), forming a core-shell structure. (B) Schematic illustration of process through which PLGA/PLA-PEG-FA nanoparticles [NPs-PEG-FA] were fabricated using a versatile single-step surface-functionalising technique. The arrows indicate the synthesis steps. Reprinted with permission from ref.38 Copyright 2018 Oxford University Press. (C) chematic illustration of embedded drug loading, where drug molecules (green dots) are uniformly dispersed within the carrier material (yellow sphere). (D) Design of the PAA@MPDA-SAP NPs for inflammatory bowel disease treatment. The arrows indicate the functionalization and assembly steps, and bold text highlights the key components of the NPs. Asterisks (*) represent the polymer chain linkage points formed through free radical polymerization of acrylic acid monomers on the MPDA nanoparticle surface. Reprinted with permission from ref.39 Copyright 2023 American Chemical Society. (E) Schematic illustration of covalently combined drug loading, where drug molecules (blue ovals) are chemically linked to the carrier structure (green lines), forming a stable conjugate. (F) Schematic presentation of the GC-B-Que prodrug micelle structure. The arrows indicate various processes such as self-assembly, drug administration routes, and triggered release mechanisms. Reprinted with permission from.⁴⁰ Copyright 2021 American Chemical Society.

Abbreviations: AA, Acrylic Acid; APS, Ammonium Persulfate; DSS, Dextran Sulfate Sodium; FA, Folic Acid; F127, Pluronic F127; GC-B-Que, GC-B-Que, quercetin-conjugated glycol chitosan prodrug linked via an aryl boronic ester; GC, Glycol Chitosan; Que, Quercetin; IBD, Inflammatory Bowel Disease; MPDA, mesoporous polydopamine; NPs, Nanoparticles; NPs-PEG-FA, folate-functionalized PLGA/PLA-PEG nanoparticles; PAA, Poly(acrylic acid); PEG, Poly(ethylene glycol); PLA, Poly(lactic acid); PLA-PEG-FA, folic acid-conjugated poly(lactic acid)-poly(ethylene glycol); PLGA, Poly(lactic-co-glycolic acid); PLGA NPs, Poly(lactic-co-glycolic acid) Nanoparticles; PVA, Polyvinyl Alcohol; ROS, Reactive Oxygen Species.

Embedment refers to the uniform dispersion or dissolution of drug molecules within the matrix of nanocarriers, without forming distinct internal cavities (Figure 1C). Common preparation methods include nanoprecipitation, emulsion solvent evaporation, and emulsion cross-linking methods. He et al utilized a solvent-free gas method to embed TNF-α-siRNA and gallic acid-mediated graphene quantum dot (GAGQD) within BSA, subsequently applying chitosan and tannin acid (CHI/TA) multilayer armor on its surface to prevent rapid degradation in the gastrointestinal tract. The embedment efficiencies of siRNA and GAGQD in BSA nanoparticles (NPs) were verified to be 77.3% and 22.1%, respectively. Stability studies of the formulation demonstrated that siRNA-GBSA(CHI/TA)₅ NPs exhibited a degradation rate of only 3.7% after 2 hours of incubation in simulated gastric fluid; after continuous incubation for 4 hours in simulated intestinal fluid, the total weight loss was merely 9.5%, with intact morphology. After 12 hours of incubation in simulated colonic fluid, their morphology remained intact and spherical. Upon reaching the target site, these nanoparticles are internalized by cells via endocytosis, releasing therapeutic cargo such as TNF-α–siRNA and GAGQD to exert anti-inflammatory and modulatory effects.⁴¹

As one of the common drug loading methods, physical bond drug loading method primarily involves physically embedding the drug in the core of the nanomaterial drug carrier, utilizing hydrophobic interaction or hydrogen bonding forces between the core and the drug. It is noteworthy that in some cases, the distinction between “embedment” and “encapsulation” may not be absolutely clear, but emulsion cross-linking is a common method for preparing such drug-loaded nanoparticles. This approach is particularly suitable for most hydrophobic drugs.³⁹ Zhang et al prepared 6-shogaol-loaded poly(lactic acid-glycolic acid copolymer) (PLGA) /PLA-PEG-FA nanoparticles (Figure 1B) by two-emulsion solvent evaporation. The drug loading and encapsulation efficiency of 6-shogaol in the nanomedicine were 35.8 ± 7.2 µg/mg and 52 ± 3.4%, respectively. In vitro release kinetics studies showed that NPs-PEG-FA released approximately 40% of 6-shogaol within the first 24 hours, and approximately 80% cumulatively within 72 hours, indicating that 6-shogaol could migrate from the hydrophobic core of PLGA to the NP surface for sustained and controlled release. Formulation stability studies demonstrated that these nanoparticles exhibited good biocompatibility, with stable morphology, particle size distribution, and zeta potential during preparation. This controlled-release characteristic significantly alleviated proctitis.³⁸

Loading/Coating refers to the process of attaching drugs to the surface or pores of nanocarriers through physical adsorption, electrostatic interactions, hydrophobic interactions, or hydrogen bonding. It is suitable for drug delivery systems that require rapid release or surface modification. Guan et al ingeniously designed and fabricated mesoporous polydopamine nanocarriers functionalized with polyacrylic acid on their surface. Leveraging the π-π stacking and hydrophobic interactions between the anti-inflammatory drug sulfasalazine (SAP) and these nanocarriers, they successfully constructed a highly efficient SAP nanomedicine (Figure 1D). This nanomedicine demonstrated a high drug loading capacity of up to 928 μg mg⁻¹. In vitro release kinetics studies revealed a significant pH-responsive controlled-release profile, with less than 17% SAP released at pH 1.2 and pH 4.5, while approximately 82% of SAP was released within 36 hours at pH 7.0. Furthermore, stability studies indicated that the nanoparticle size remained stable after incubation for 24 hours in simulated gastric and intestinal fluids, effectively preventing premature drug degradation.³⁹

Physical drug loading preserves the natural structure and activity of drugs by utilizing non-covalent interactions, making it a versatile approach compatible with a wide range of therapeutics. Its key advantages include simple preparation, scalability, and broad applicability without the need for chemical modifications that could alter drug efficacy. However, this method has limitations, including relatively low drug-loading capacity (dictated by carrier materials) and a risk of burst release or premature leakage, particularly under physiological stress such as gastrointestinal pH variations. While surface modifications (eg, chitosan or hyaluronic acid coatings) can enhance stability, uncontrolled drug diffusion remains a challenge, potentially reducing bioavailability and necessitating more frequent dosing.

Chemical Bond Drug Loading Method

Chemical bond drug loading is to link drug molecules and the functional groups of NDDS through chemical bonds to form stable drug carrier complexes. The drugs attached to the ends of functional groups are directly wrapped into the core of the nanocarrier during the self-assembly of nanomaterials.

Covalent Conjugation

Covalent conjugation is a key strategy for achieving precise controlled release in nanomedicine carriers, which binds drugs to the carrier through chemical bonds and triggers release under specific stimuli. In the treatment of IBD and other diseases, covalent coupling enables the design of responsive bonding modes targeting the lesion microenvironment (such as pH, enzymes, redox levels) to optimize drug delivery efficiency (Figure 1E).

Shen et al developed a pH/ROS dual-responsive prodrug micelle, by covalently linking quercetin to glycol chitosan. This nanomedicine achieved an encapsulation efficiency of 23.5 ± 2.3% and a loading capacity of 6.7 ± 0.1%. It demonstrated excellent stability with <5 wt% quercetin released in simulated gastric fluid within 1 hour. In inflammatory-like conditions (pH 5.8 with H₂O₂), over 95 wt% quercetin was released within 72 hours, showing significant pH/ROS dual-responsive and sustained release capabilities, enabling targeted inflammation treatment (Figure 1F).⁴⁰

Coordination Bond

The coordination bond is a special type of chemical bond that involves a ligand providing a lone pair of electrons to fill the empty orbitals of a metal ion, thereby forming a stable complex. The formation of coordination bonds can establish stable interactions between nanocarriers (such as nanoparticles, nanocapsules, etc.) and drug molecules. For instance, Fu et al designed copper-luteolin nanocomplexes (CuL NCs) through a metal-polyphenol coordination strategy, effectively utilizing metal ions for nanometrization of natural small active molecules to address their low bioavailability. These CuL NCs were found to contain approximately 14.25 wt% luteolin and 4.14 wt% copper. Characterization by DLS and ζ potential measurements confirmed their remarkable stability in various physiological solutions, such as fetal bovine serum and DMEM. While specific release kinetics were not explicitly detailed, the strategy allows for multifaceted regulation of the IBD inflammatory microenvironment by reducing oxidative stress, promoting anti-inflammatory responses, improving intestinal barrier integrity, and positively influencing gut microbiota diversity.⁴²

Click Chemistry

Click Chemistry connects drugs and carriers through efficient and specific reactions, such as alkyne azide cycloaddition. Dong et al combined azide-modified Prussian blue nanoenzyme with dibenzocyclooctyne (DBCO) modified Lactobacillus reuteri through a “click” reaction between azide and DBCO, effectively reducing ROS levels in the complex physiological environment of IBD, improving the targeted colonization and survival efficiency of probiotics, and effectively reducing intestinal inflammation.⁴³

Chemical drug loading relies on covalent or ionic bonding between the drug and the nanocarrier, offering precise control over drug release kinetics. A key advantage is the high stability of drug-carrier conjugates, minimizing premature leakage and enabling stimuli-responsive release (eg, pH-, enzyme-, or redox-triggered cleavage in the inflamed gut). This approach enhances targeting efficiency, reducing systemic toxicity—a critical factor in IBD therapy. However, chemical conjugation often requires complex synthesis, which may compromise drug bioactivity or necessitate additional purification steps. Furthermore, the irreversible nature of some bonds (eg, amide linkages) can hinder complete drug release at the target site, potentially limiting therapeutic efficacy.

An assessment of physical encapsulation and chemical conjugation is summarized in Table 1. Physical loading (driven by hydrophobic or electrostatic interactions) is best suited for lipophilic small molecules (eg, curcumin, magnolol) and labile biologics, preserving native pharmacophore integrity to provide rapid drug availability during acute inflammatory flares;⁴⁴ however, it remains susceptible to bile-salt-mediated premature burst leakage during transit. Conversely, chemical conjugation via cleavable bonds (eg, ester, disulfide) is indispensable for unstable, ionizable, or functionalized payloads (eg, short-chain fatty acids, peptides), ensuring near-zero transit leakage and lesion-specific triggered release ideal for long-term maintenance regimens, despite potential steric impediments to bio-cleavage. To reconcile these trade-offs, recent designs increasingly adopt hybrid platforms that integrate both modalities. For example, Fan et al constructed structural nanoparticles by covalently linking butyrate through redox-responsive disulfide bonds while physically entrapping lipophilic magnolol via nanoprecipitation, achieving synchronized dual-drug release that accelerated mucosal healing in colitis models.⁴⁵

Table 1.

Comparison of Characteristics Between Physical and Chemical Binding Methods

Drug Loading Strategy	Binding Mechanism	Main Advantages	Limitations and Challenges	Applicable Scenarios
Physical binding methods	Non-covalent interactions such as hydrophobic interactions, electrostatic adsorption, and hydrogen bonding.	Preserves the natural structure and activity of drugs. Simple preparation and scalability. Broad applicability without chemical modifications. Enhances drug stability, solubility, and bioavailability.	Relatively low drug-loading capacity. Risk of burst release or premature leakage. Uncontrolled drug diffusion reduces bioavailability.	Delivery systems requiring rapid release or surface functionalization. Enhancing drug stability and targeting. Suitable for lipophilic small molecules and labile biologics. Provides rapid drug availability during acute inflammatory flares.
Chemical binding methods	Covalent bonds (eg, amide bonds, disulfide bonds, boronate ester bonds), coordination bonds, and click chemistry.	High stability of drug-carrier conjugates, minimizing premature leakage. Enables precise controlled release. Can design responsive bonding modes targeting specific microenvironments (eg, pH, enzymes, redox levels). Enhances targeting efficiency, reducing systemic toxicity. Lesion-specific triggered release, ideal for long-term maintenance regimens.	Requires complex synthesis. May compromise drug bioactivity or necessitate additional purification steps. Irreversible nature of some bonds can hinder complete drug release.	Achieving precise controlled release. Targeting specific microenvironments. Suitable for unstable, ionizable, or functionalized payloads.

Nanomedicine Treatment Pathways for IBD

Depending on the site of drug action, nanomaterial drug carriers contain a variety of targeted sites and uses during drug delivery, such as i). the ability of nanomaterial drug carriers to cross the mucus layer and deliver the loaded drug to intestinal cells;⁴⁶ ii). It can also be engulfed by macrophages through phagocytosis, thereby regulating the immune environment of the gut;⁴⁷ iii). attaches to target tissues for a long time, allowing them to be used for intestinal imaging or other therapeutic purposes.⁴⁸ When it comes to different target sites and uses, it is necessary to consider the therapeutic pathway of nanomaterial drug carriers in the treatment of IBD. In general, the routes of administration for IBD treatment include oral, rectal, and injectable administration, and the mode of administration is determined by different targeting sites and uses. We have comprehensively summarized in Table 2.

Table 2.

Therapeutic Pathways of Nano-Drug Carriers in IBD

Drugs	Model	Therapeutic Pathways	Dosage of the Drug	Frequency/Schedule	Carriers	Ref.
TNF-α siRNA	3% DSS-induced colitis in C57BL/6 mice	Oral administration	66 μg/kg TNF-α siRNA	Not clearly specified in the accessible text	GPNs	[49]
Quercetin	TNBS-induced colitis in Wistar rats	Oral gavage	9 mg/kg/day	Once daily on days 3–5 after TNBS administration	PEG-containing vesicles coated with chitosan and Nutriose	[50]
5-ASA and Bac	1.5% DSS-induced chronic colitis in male C57BL/6 mice	Oral gavage	NDF-M1/L: 1.0 mg NDFs + 0.61 mg proteins + 1.0 mg 5-ASA + 4.0 mg Bac; M and H doses were 2- and 4-fold higher, respectively	Days 6–10, 16–20, and 26–30	Alginate hydrogel microspheres containing Bac and NDF-Pro/5-ASA	[51]
Dexa	3% DSS-induced colitis in male BALB/c mice	Oral gavage	5 mg/kg Dexa equivalent	Days 7–14 after 7 days of DSS induction	Dexa-GP/ES/Pu NCs	[52]
siTNF-OMe-P	5% DSS-induced colitis in female C57BL/6 mice	Intrarectal administration	4 nmol/mouse	Days 2 and 4	Lipofectamine 2000-based lipoplex	[53]
Anti-TNF-α ASO	TNBS-induced acute colitis in BALB/c mice	Intracolonic administration	5 mg/kg ASO	Single administration after TNBS induction	gal-LMWC	[34]
INF	5% DSS-induced colitis in female Kunming mice	Rectal administration	5 mg/kg (100 μL)	Once daily for 3 consecutive days	PPNPs	[54]
Dexamethasone 21-phosphate	Patients with steroid-dependent IBD (5 UC and 5 CD)	Intravenous reinfusion	5.5 ± 2.4 mg/infusion	Every 4 weeks, for a total of 3 infusions	Autologous erythrocytes	[55]
Dexamethasone 21-phosphate	18 pediatric patients with steroid-dependent CD	Intravenous reinfusion	Mean 8.8 mg/infusion (range, 2.4–14.7 mg)	Every 4 weeks; mean treatment duration, 24 months	Autologous erythrocytes	[56]
C-dot SOD nanozyme and CD98 CRISPR/Cas9 plasmid	DSS-induced UC in mice	Intravenous injection	2 mg/kg CCZM	Days 2, 4, and 6	Macrophage membrane-camouflaged ZIF-8 MOF nanoparticles (CCZM)	[57]
Antisense MIF oligonucleotide	DSS-induced colitis in mice	Intraperitoneal injection	0.2 mg/kg antisense MIF/SPG complex	Days 0, 3, 7, and 10	SPG	[58]

Abbreviations: IBD, inflammatory bowel disease; UC, ulcerative colitis; CD, Crohn’s disease; DSS, dextran sulfate sodium; TNBS, 2,4,6-trinitrobenzenesulfonic acid; TNF-α, tumor necrosis factor-alpha; siRNA, small interfering RNA; GPNs, galactosylated chitosan-modified poly(lactic-co-glycolic acid) nanoparticles; PEG, poly(ethylene glycol); NDF, nanoscale dietary fiber; 5-ASA, 5-aminosalicylic acid; Bac, Bifidobacterium; Dexa, dexamethasone; GP/ES/Pu, galactosylated poly(lactic-co-glycolic acid)/Eudragit S100/pullulan; NCs, nanocargoes; ASO, antisense oligonucleotide; gal-LMWC, galactosylated low-molecular-weight chitosan; INF, infliximab; PPNPs, polyphenol-poloxamer nanoparticles; SOD, superoxide dismutase; CRISPR/Cas9, clustered regularly interspaced short palindromic repeats/CRISPR-associated protein 9; ZIF-8, zeolitic imidazolate framework-8; MOF, metal-organic framework; CCZM, macrophage membrane-camouflaged C-dot/CD98 CRISPR/Cas9 plasmid-loaded ZIF-8 nanoparticles; MIF, macrophage migration inhibitory factor; SPG, schizophyllan.

Oral Administration

Oral administration represents one of the most practical and patient-compliant routes for managing IBD, offering non-invasive, cost-effective delivery that allows therapeutic agents to directly engage the mucosal lesions.⁵⁹ In animal experiments, oral administration typically encompasses delivery via voluntary feeding (ie, mixed in feed or drinking water) and oral gavage. In addition to digestive enzymes and pH shifts across intestinal segments, oral nanoparticles must navigate the dense and chemically altered colonic mucus hydrogel, which frequently entraps nanocarriers via electrostatic or steric interactions and prevents them from reaching the damaged epithelial layer.⁶⁰ Furthermore, upon exposure to gastrointestinal fluids, orally administered nanoparticles may adsorb proteins, bile components, and other biomolecules to form a dynamic gastrointestinal biomolecular corona, which can alter their surface properties and subsequent interactions with mucus and intestinal cells.⁶¹ Compounding these interfacial barriers, severe intestinal transit variability between disease phenotypes—ranging from rapid diarrhea-induced clearance during active flares to delayed transit in stricturing or obstructive pathologies—drastically perturbs the release kinetics and retention window of oral formulations.⁶² Conversely, unique environments, such as the different physiological characteristics of the gastrointestinal segment in IBD, the types of digestive enzymes, pH acidity changes, drug delivery times, gut microbiota, and stress-mediated mechanisms, have all been used to design targeted nanomaterial drug carriers.^63,64 Natural polysaccharides, such as alginate and chitosan, demonstrate advantages in oral drug delivery. They not only tolerate the gastric acidic environment but also facilitate colonic mucus penetration or local bioadhesion. For instance, chitosan-encapsulated nanovesicles achieve targeted colorectal release of siRNA or quercetin, effectively alleviating inflammation.^49,50

Exploiting alterations in the luminal microflora represents another strategy for precise delivery. Microbe-activated platforms, such as nanoscale dietary fibers (NDFs), enable colonic delivery of 5-ASA via microbial fermentation.⁵¹ Furthermore, galactosylated PLGA/Eudragit S100/pullulan nanocapsules can selectively degrade through colonic pullulanase to release therapeutics.⁵²

It is noteworthy that maintaining the functional integrity of biomacromolecules during mucosal transit is crucial. This necessitates carriers engineered for high cellular uptake and efficient endosomal escape (for siRNA delivery),²⁴ as well as protective matrices that safeguard orally administered monoclonal antibodies against luminal proteolysis to prevent systemic adverse effects.^65,66

Colorectal Administration

Colorectal administration is recommended as a first-line treatment for patients with mild to moderate distal colitis, such as mesalazine, which has been developed as a traditional suppository or enema for the treatment of mild to moderate colitis.⁶⁷ Compared with oral administration versus intravenous administration, colorectal administration for the treatment of IBD has several advantages:1) Colorectal administration can avoid the pharmacokinetic problems associated with gastrointestinal motility and pH changes, which is especially beneficial for biological macromolecule drugs. 2) The systemic exposure of the drug is much lower than that given orally and intravenously, reducing the likelihood of side effects.⁶⁸ 3) Colorectal administration allows the drug to be delivered directly to the inflamed site, ensuring a high concentration of the active drug.⁶⁹ 4) Rectal administration by suppository, enema, foam, and ointment forms is superior to oral administration for patients who are comatose or unable to swallow (children or those with seizures or dysphagia) or patients with stronger local effects (eg, steroid administration for rectal 5-ASA or UC), but rectal administration is superior to oral administration in dosage forms such as suppositories, enemas, foams, and ointments.⁶⁸ However, rectal administration also has some drawbacks. 1) Patient compliance with rectal administration is not as high as with oral administration, because enema-based preparations require patients to retain the delivery system for long periods. This is obviously difficult in patients with diarrhea and limits their daily activities after administration. 2) Rectal administration can stimulate the gastrointestinal tract, especially the large intestinal mucosa, which can cause some damage to the intestinal mucosa. Therefore, the rectal delivery system needs to be further designed to selectively release drugs at the site of inflammation through drug technology, prolong the residence time of the drug at the site of inflammation, and reduce the frequency of administration, so as to reduce the burden of drug delivery and protect the healthy mucosa.

Many antisense oligonucleotides (ASOs) have been successfully delivered by rectal administration in preclinical and clinical IBD studies. As an example, Ocampo et al found that rectal administration of siRNA with 2′-O-methyl and propanediol modifications significantly improved clinical endpoints and histopathological severity in DSS-treated mice.⁵³ Zuo et al injected galactosylated low-molecular-weight chitosan (gal-LMWC)/ASO nanomedicine intracolonically into mice with TNBS-induced colitis. Intracolic injection enhances the therapeutic effect of the gal-LMWC/ASO complex due to the direct delivery of the nanocomposite to the inflammatory tissue of the colon.³⁴ Fan et al found that infliximab(INF)-loaded polyphenol-poloxamer nanoparticles were more effective than direct intravenous INF injection, with a significant reduction in systemic exposure and a similar anti-inflammatory effect to intravenous INF.⁵⁴

Injection Administration

In addition to oral administration and colorectal administration, injection administration is one of the most commonly used treatment methods in clinical practice. Currently, most biologics (anti-TNF antibodies, anti-integrin antibodies) and some corticosteroids are administered intravenously or intramuscularly. Compared with oral administration, injection administration, including intravenous, intramuscular injection, subcutaneous injection, or intraperitoneal injection, also has some advantages and disadvantages, such as 1) avoiding first-pass effects, providing immediate systemic bioavailability, and delivering drugs to various tissues and organs through the whole body fluid circulation system;⁵⁶ 2) The blood concentration is high, and the amount of drug entering the body is accurate; 3) Facilitate the administration of drugs to uncooperative or unconscious patients. Disadvantages include 1) the high systemic bioavailability brought about by injectable administration inevitably increases the likelihood of systemic side effects. 2) and because injections require specialized personnel and equipment, and are often associated with pain, they are less favored by patients; 3) It is easy to cause specific adverse reactions such as thrombophlebitis, catheter-related bloodstream infection, tissue necrosis, etc.

In view of the advantages of intravenous drug administration and the limitations of some drug properties, such as large particle size or strong hydrophobicity, nanomaterial drug carriers will adhere to a large number of blood components (such as opsonin and protein) and be rapidly eliminated by the humoral circulation system.^70,71 Therefore, the size control of nanomaterial drug carriers (100–200 nm) and hydrophilic group modification are the key parameters to improve the half-life of the drug delivery system in the blood. Tang et al designed a CD47 protein-derived enzymatic digestion-tolerant D-type Self peptide (DS) and labeled it on liposomes (DSL). After intravenous administration, DSL enters the body and is rapidly adsorbed onto phagocytes, forming a long-lasting mask that encloses the cell membranes and thus reduces interactions between phagocytes and subsequently injected nanoparticles.⁷² Exosome-loaded drugs have been partially studied for intravenous injection of nanomaterial drugs through intravenous injection. For example, Castro and Annese et al demonstrated that natural red blood cells in IBD patients used as nanomaterial drug carriers can control the sustained release of dexamethasone 21-Phosphate in the blood circulation after intravenous injection, thereby continuously slowing down IBD inflammation.^55,56 Ma et al targeted the inflammatory colon under the guidance of macrophage membranes to release carbon nanodots-SOD nanozymes and CRISPR/Cas9 systems for site-directed treatment of UC after intravenous administration of biomimetic pH-responsive metal-organic frameworks (MOFs) carriers. After intravenous administration, this delivery system can target the site of inflammation of UC through the damaged vasculature.⁵⁷ Takedatsu et al administered a 0.2 mg/kg dose of antisense macrophage-migration inhibitory factor (MIF)/ schizophyllan (SPG) complex intraperitoneally with efficient absorption by macrophages to alleviate IBD.⁵⁸

In summary, the treatment of IBD by nanomedicine depends not only on the physicochemical properties of the nanomedicine carrier and the drug carried by the nanomedicine but also on the degree of development of the individual disease. For example, when inflammation has spread to extraintestinal sites such as the eyes, joints, and skin, injectable administration is particularly useful for the treatment of IBD. When inflammation is confined to the gastrointestinal tract, oral and/or rectal administration would be a better solution to reduce unnecessary systemic distribution of the drug. Beyond route selection, however, the actual therapeutic outcome is fundamentally constrained by post-internalization intracellular trafficking within target mucosal and immune cells. Unmodified nanomedicines commonly suffer from lysosomal entrapment via the classical endo-lysosomal pathway, where cytosolic escape efficiencies typically remain below 1% to 2%.⁷³ Consequently, equipping these carriers with endosomal escape capabilities is essential across all administration routes to prevent hydrolytic payload degradation prior to intracellular drug release. To bridge this macroscopic-to-intracellular transit, rationally engineered nanomedicines must couple precise environmental-responsiveness for timely payload liberation (Section 3) with robust targeting strategies to overcome biological barriers (Targeting Mechanism of IBD Nanomaterial Drug Carriers), thereby realizing full therapeutic translation. However, repeated administration of systemic nanomedicines poses critical immunopharmacological bottlenecks that are frequently overlooked in acute preclinical assessments. Specifically, surface-adsorbed opsonins and repeated exposure to functional hydrophilic polymers (such as PEG) can trigger complement activation-related pseudoallergy (CARPA) or elicit anti-carrier antibodies that mediate accelerated blood clearance (ABC), drastically truncating systemic circulation upon subsequent dosing.^74,75 Mitigating such carrier-associated immunogenicity through fine-tuned polymer architectures, biomimetic masking, or orthogonal stealth modifications remains imperative to preserve long-term pharmacokinetic durability and biocompatibility in chronic IBD therapy.⁷⁶

Mechanism of IBD Nanodrug Release

Depending on how the drug is combined with the carriers and the route of treatment, the way the drug is released in the body is also different. When the drug is prepared by wrapping, there are two main ways to release the drug: one is the release of the drug after the decomposition of the nanomaterial, and the other is that the nanomaterial does not decompose, and the drug induces its release through other environmental stimuli. When the chemical bond loading method is prepared, there are also two main ways to release the drug: one is that the nanomaterial is degraded first and then the covalent bond between the nanomaterial and the drug is broken to achieve drug release,⁵¹ and the other is that the covalent bond between the nanomaterial and the drug is broken first and then released through diffusion.⁴⁰ According to the therapeutic pathway of nanomedicine, when the drug is administered orally. The factors that induce release are mainly concentrated in the unique environment of the gastrointestinal tract. When the drug is administered by injection, the release-inducing factors may be concentrated in the unique environment where the body fluid is circulated. Regardless of the inducing factor, it is necessary to achieve precise drug release, reduce drug loss, and achieve maximum drug targeting to the lesion site. After summarizing, it is found that two main factors induce drug release regardless of the release mode: one is endogenous stimulation, such as pH, enzymes, redox, intestinal environment, intestinal tissue environment, body fluid environment, etc. The other is exogenous stimuli, such as light, electricity, etc. These strategies include but are not limited to ROS-triggered release, pH-sensitive release, redox-triggered release, and microenvironment-triggered release.^64,77 Based on these inducing factors, various vectors can be designed to release the drug in a redox environment, release the drug at a specific pH, or require bacterial cleavage to activate, and so on, several of which are currently being studied.^78,79 The following will elaborate on the release mechanism from the three inducing factors of redox, pH, and microenvironment.

Redox-Triggered Release

Reactive oxygen species (ROS) are an inflammatory mediator that is produced in small amounts in normal tissues but is overexpressed at the site of inflammation.⁸⁰ In healthy intestinal mucosa, steady-state extracellular hydrogen peroxide (H₂O₂) concentrations are tightly regulated within the sub-micromolar range (0.1–10 μM); however, sustained activation of infiltrating neutrophils and macrophages in active IBD drives mucosal ROS levels 10- to 100-fold higher, generating localized H₂O₂ concentrations reaching 100–1000 μM alongside transient micromolar bursts of hydroxyl radicals (·OH) and superoxide anions (O₂•⁻).^81,82 To exploit this sharp oxidative gradient, polymeric nanocarriers have been engineered with diverse ROS-cleavable or oxidation-sensitive linkages, including phenylboronic esters, thioketals, oxalates, and poly(thioethers).^83–85 Therefore, ROS-induced oxidative stress is one of the commonly used factors to induce drug release in nanomaterial drug carriers.⁸⁶

Phenylborate Bonds-Containing IBD Nanomaterial Drug Carriers

Phenylborate bonds are frequently used in drug delivery systems because they are selectively degraded by ROS.^87,88 Phenylboronic esters undergo nucleophilic oxidation by H₂O₂ to yield hydrophilic phenols and boric acid, which rapidly destabilizes the hydrophobic particle core to trigger drug release. Zhao et al conjugated a phenylboronic ester-modified carboxylmethyl chitosan (OC-B) conjugated with berberine (BBR) to construct a new oxidation-responsive IBD nanomedicine. It can be selectively degraded at ROS and remodel the gut microbiota, significantly improving colitis symptoms and colonic injury⁸⁹ (Figure 2A). Fan et al constructed a kind of nanomaterial containing superoxidase dismutase-mimetic temporally conjugated β-cyclodextrin and 4-(hydroxymethyl)phenylboronic acid pinacol ester-modified genistein (GEN), which could effectively eliminate H₂O₂ in a dose-response pattern and specifically release GEN to the site of inflammation.⁹⁰

Figure 2.

Redox-triggered release of nanodrug carriers for the treatment of proctitis. (A) The schematic presentation of OC-B-BBR prodrug micelles structure and their inflamed-selective anti-inflammatory drug delivery by ROS-triggered mechanism. The arrows illustrate the process of self-assembly of OC-B-BBR, subsequent oral delivery, and the ROS-triggered drug release in the inflamed area. Reprinted with permission from ref.89 Copyright 2021 Zhao, Du, Tian, Kang, Li, Dai, Li, Zhang and Li. (B) Thioketal nanoparticles are formulated from a new ROS-sensitive polymer and release orally delivered siRNA at sites of intestinal inflammation. PPADT is a new polymer composed of ROS-sensitive thioketal linkages (circled in red). Compound 1 (1,4-benzenedimethanethiol) is a dithiol monomer containing two thiol groups (-SH) that serves as the primary structural unit. Compound 2 (2,2-dimethoxypropane) is a dimethyl acetal reagent that reacts with the thiol groups under mild acidic catalysis to form thioketal linkages through step-growth polymerization in benzene at 107 °C for 24 h. Compound 3 (PPADT) is the resulting ROS-sensitive polymer with molecular weight of approximately 9,000 Da; the red dotted circles highlight the thioketal linkages (–S–C(CH₃)₂–O–) formed during polymerization. TNFα-TKNs were prepared by first precomplexing TNFα-siRNA with the cationic lipid DOTAP. Next, these siRNA-DOTAP complexes were added to an organic solution containing PPADT. A single-emulsion procedure was then used to produce particles with mean diameters of ~600 nm. The arrows indicate the synthetic steps. SEM image shows TNFα-TKNs with scale bar representing 1.5 µm. Asterisks (*) represent the polymer chain extension points in the PPADT backbone where the thioketal monomers polymerize. Reprinted with permission from ref.82 Copyright 2010, Springer Nature Limited. (C) Synthesis and characterization of TPGS-PBTE. The arrows in the diagram illustrate the synthetic steps for Macromonomer TPGS-A and the subsequent formation of TPGS-PBTE via Michael addition polymerization. The synthesis begins with TPGS (α-tocopherol polyethylene glycol succinate; shown in blue with repeating ether units and hydroxyl groups labeled a-c, b represents methylene groups adjacent to -O-, and the remaining repeating unit marked as 3) as the macromolecular precursor. TPGS is then conjugated with acryloyl chloride under triethylamine (TEA) catalysis to generate TPGS-A (j’, k’, l’), where hydroxyl groups are substituted with acryloyl moieties to serve as a macroinitiator. DTT (dithiothreitol), containing two thiol (-SH) and two hydroxyl (-OH) groups, is subsequently introduced. The core ROS-responsive polymer PBTE is then synthesized through free radical polymerization of 1,4-butanediol diacrylate (BDD) with DTT. In the final polymer structure (shown in red), key functional groups are labeled: a represents the hydroxyl positions (highlighted in red), b and d denote the methylene carbons adjacent to sulfur atoms with characteristic NMR shifts, s indicates the thioether linkages (–S–) that serve as the primary ROS oxidation targets, e-i represent the ester and ether bonds within the polymer backbone, j-l correspond to the TPGS-conjugated regions, m marks the repeating polymer units, and n represents the degree of polymerization. The thioether bonds (s) are selectively oxidized by reactive oxygen species (ROS) at inflammation sites, converting –S– to sulfoxides and then to sulfones, thereby disrupting the hydrophobic-hydrophilic balance of the polymer and triggering accelerated ester hydrolysis for targeted drug release in the inflammatory microenvironment. Reprinted with permission from ref.91 Copyright 2022 The Authors. Published by Elsevier Ltd. (D) Synthetic procedure of CMI, ATP-CMI conjugate and ATP-CMI NPs. FT-IR spectrums of INU, CMI and ATP-CMI conjugate.1^H NMR spectra of CMI and ATP-CMI conjugate in D₂O. The red circle means the new signals appearing at 6.8 ppm and 7.2 ppm in the ¹H NMR spectrum of the ATP-CMI conjugate, which indicate the presence of a benzene ring structure in 4-ATP, thus demonstrating the successful conjugation of 4-ATP to CMI. Reprinted with permission from ref.92 Copyright 2017. Published by Elsevier Ltd. (E) Schematic illustration of the protective effects of SeNG in a colitis model. The solid arrows indicate material transformation, and signaling pathways; bold red downward arrows represent the significant reduction or inhibition of reactive oxygen species (ROS) and inflammatory factors (iNOS, TNF-α, IL-6); bold black arrows signify the enhancement of the Nrf2/HO-1 signaling pathway. Reprinted with permission from ref.93 Copyright 2022 American Chemical Society.

Abbreviations: ROS, Reactive Oxygen Species; PTSA, p-toluenesulfonic acid; DOTAP, 1,2-dioleoyl-3-trimethylammonium-propane; TEA, Triethylamine; OC-B-BBR, berberine-conjugated phenylboronic ester-modified carboxymethyl chitosan; INU, Inulin; CMI, Carboxymethyl Inulin; 4-ATP, 4-Aminothiophenol; EDC, 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide; NHS, N-Hydroxysuccinimide; ATP-CMI, 4-aminothiophenol-conjugated carboxymethyl inulin; ATP-CMI NPs, 4-aminothiophenol-conjugated carboxymethyl inulin nanoparticles; HA, Hyaluronic Acid; Se-HA, selenocystamine-conjugated hyaluronic acid; W/O, Water-in-Oil; SeNG, diselenide-bridged hyaluronic acid nanogel; Nrf2, Nuclear factor erythroid 2-related factor 2; Keap1, Kelch-like ECH-associated protein 1; ARE, Antioxidant Response Element; HO-1, Heme Oxygenase-1; SOD, Superoxide Dismutase; CAT, Catalase; GSH, Glutathione; GST, Glutathione S-transferase; iNOS, Inducible Nitric Oxide Synthase; TNF-α, Tumor Necrosis Factor-alpha; IL-6, Interleukin-6; TPGS, D-α-tocopherol polyethylene glycol succinate; TPGS-A, TPGS acrylate; TPGS-PBTE, D-α-tocopherol polyethylene glycol succinate-b-poly(β-thioester); FT-IR, Fourier-transform infrared spectroscopy; NMR, nuclear magnetic resonance.

Thioketal Bond-Containing IBD Nanomaterial Drug Carriers

Thioketal bond-containing nanomaterial drug carriers are resistant to acidic, alkaline, and proteasome degradation, but can be degraded by abundant ROS produced at the site of inflammation.^82,94 The thioketal bond (S-C-S) is the key structural element responsible for ROS-responsive drug release; under elevated ROS conditions, the sulfur atoms in thioketal bonds undergo oxidative cleavage, leading to controlled drug release at the disease site. Thus, thioketal bonds are specifically incorporated into ROS-responsive drug delivery systems to achieve disease-site-specific drug release.⁹⁵ For example, Wilson et al found that the thioketal bonds in thioketal nanoparticles (TKNs) derived from a novel polymer (poly-(1,4-phenyleneacetone dimethylene thioketal) (PPADT)) could selectively degrade under ROS conditions, thereby promoting the selective release of TNFα-siRNAs in nanomaterial drug carriers at sites of intestinal inflammation with high ROS response⁸² (Figure 2B). Li et al used aromatized thioketal (ATK) to link the anti-inflammatory drug budesonide and the antioxidant tempol to make a nanomedicine. After oral administration, these ROS-sensitive nanoparticles can be delivered to the inflamed colon and simultaneously release both drugs into colitis tissues, ensuring synergistic anti-inflammatory and antioxidant efficacy.³⁰

Thioether Bond-Containing IBD Nanomaterial Drug Carriers

Unlike bonds that break completely, thioether groups do not degrade under oxidation. Instead, ROS oxidation converts hydrophobic sulfides into water-loving sulfoxides or sulfones. This change makes the nanoparticles swell like a sponge, allowing the drug to gradually diffuse out while consuming local oxidants to relieve mucosal stress.⁹⁶ Tan et al found that nanomaterial drug carriers prepared with D-α-tocopherol polyethylene glycol succinate-b-poly(β-thioester) copolymer (TPGS-PBTE) were used for the selective release of luteolin at sites of colorectal inflammation because the thioether bonds on the polymer backbone can also selectively degrade ROS⁹¹ (Figure 2C). However, there is a safety factor to consider when developing redox nanomaterial drug carriers. For example, when the ROS concentration is at a normal level, redox nanomaterial drug carriers should be inert and not produce inflammatory responses in vivo, and future research should also focus more on their biocompatibility and safety in vivo.

Disulfide Bonds Containing IBD Nanomaterial Drug Carriers

Disulfide (-S-S-) bonds have traditionally been widely utilized as redox-responsive dynamic crosslinkers in drug delivery systems, primarily relying on elevated intracellular glutathione (GSH, 2–10 mM) levels to induce reductive thiol–disulfide exchange and subsequently release therapeutic payloads inside target cells.⁹⁷ However, under pathologically elevated ROS microenvironments, disulfide bonds can likewise undergo irreversible oxidative cleavage. In active IBD, the divalent sulfur atoms within disulfide linkages are susceptible to direct ROS attack, being progressively oxidized into hydrophilic thiosulfinates and thiosulfonates, and ultimately into negatively charged sulfonic acid residues (—SO₃H).^98,99 This oxidative transition not only destabilizes the hydrophobic core or crosslinked integrity of nanocarriers, but also endows them with unique dual redox responsiveness—enabling extracellular ROS-triggered drug release within inflamed intestinal lesions as well as further intracellular degradation mediated by GSH upon cellular internalization. Based on this, Sun et al designed nanoparticles based on an amphiphilic inulin polymer with 4-aminothiophenol (ATP) grafted onto carboxymethyl inulin (CMI), which contained disulfide bonds and showed a significant pH/redox-sensitive swelling transition⁹² (Figure 2D). Fan et al found that when the butyrate is attached to the polymer backbone by disulfide bonds, the butyrate-rich polymers can not only self-assemble into nanoparticles by introducing hydrophobic butyrate but also respond to low luminal redox potential (−400 mV) and/or high levels of glutathione, thus releasing drugs at the site of inflammation. As a result, dual-responsive butyrate-rich polymer nanoparticles can optimize pH/redox-responsive drug release behavior by adjusting the butyrate content in the polymer.⁴⁵

Diselenide Bond-Containing IBD Nanomaterial Drug Carriers

Diselenide bonds are structurally weaker and more sensitive to oxidation than disulfide bonds. This is primarily attributed to their inherently lower bond energy (approximately 172 kJ/mol) and the electronegativity of selenium (2.55 on the Pauling scale).¹⁰⁰ Xu et al synthesized diselenide-bridged hyaluronic acid nanogel (SeNG) to release drugs under redox conditions. SeNG increased the content of antioxidant enzymes, namely superoxide dismutase (SOD), catalase (CAT), glutathione (GSH), and glutathione s-transferase (GST), by upregulating the Nrf2/HO-1 signaling pathway⁹³(Figure 2E).

Although effective results have been obtained with some single stimulus-dependent drug delivery systems, there are major problems with differences in the intestinal environment of patients with different stages of IBD. Formulations that respond to only one stimulus may not provide adequate targeting. To overcome this problem, dual/multiple response nanomaterial drug carriers have been used in the treatment of IBD. NPs released by dual sensitivities (pH/time, pH/enzyme, pH/microenvironment, etc.) have also been developed to improve IBD therapy.^101,102 A novel pH/ROS dual-reaction prodrug micelle, GC-B-Que, is used as an inflammation-targeting agent to achieve targeted release at H₂O₂ and intestinal pH by using aryl borate ester and biocompatible ethylene glycol chitosan as a response agent.⁴⁰ Beyond formulation chemistry, a critical yet frequently overlooked clinical challenge is the pronounced inter-patient variability in mucosal ROS levels, which fluctuate extensively across active flare-ups, remission stages, and spatial disease anatomical locations.¹⁰³ This pathophysiological heterogeneity directly dictates the in vivo responsiveness of ROS-cleavable linkages; nanocarriers with excessively high trigger thresholds risk incomplete payload release in patients with moderate oxidative burdens, whereas overly sensitive matrices may suffer premature drug leakage in off-target tissues. This highlights the critical need for high specificity in responsive chemistry, ensuring that the nanocarrier only releases its cargo in the presence of pathological ROS levels, and not in physiological conditions, to avoid off-target effects and maintain therapeutic efficacy. Consequently, integrating non-invasive ROS imaging biomarkers or engineering adaptive nanoplatforms with tunable redox thresholds will be vital to match personalized patient oxidative profiles and ensure predictable release kinetics.^104,105

pH-Triggered Release

The change in pH is one of the unique environments of the digestive system, and it is necessary to overcome the pH change in the digestive system and safely deliver the drug to the site of colorectal inflammation for drug release. In addition, colorectal pH is more acidic (pH 2.3–5.5) in patients with IBD than normal colorectal pH (pH 7.0 ± 0.7).¹⁰⁶ Combined with the above two special conditions, pH-sensitive nanomaterial drug carriers have become one of the ideal candidates for the development of targeted nanomaterial drug carriers.¹⁰⁷

Eudragit-Containing IBD Nanomaterial Drug Carriers

Kriplani et al found that Eudragit polymers, also known as polymethacrylates, may be anionic, cationic, or non-ionic polymers of methacrylic acid, dimethylaminoethyl methacrylates, and methacrylic acid esters in varying ratios.¹⁰⁸ Among them, Lv et al prepared an Eudragit S100 nanomaterial drug carrier for colorectal drug-specific delivery.¹⁰⁹ It is effective in protecting encapsulated Ovalbumin(OVAs) at low release levels (<7%) at pH levels between pH 1.2 and 4.5, and the encapsulated OVA begins to gradually accelerate release when pH rises from 4.5 to 5.0 and reaches its highest release efficiency at colorectal pH (release rate = 100%).²⁹ Deshmukh et al also used EudragitS100 for the treatment of ulcerative proctitis¹¹⁰ (Figure 3A). In addition, Varshosaz and Naeem et al found that when silybin and budesonide were encapsulated in pH-sensitive EudragitFS30D, they not only ensured the stability of the drug at gastric acidic pH but also promoted the selective release of the encapsulated drug at colorectal pH.^101,111 Similarly, Zhang et al coated dual pH-sensitive Eudragit EPO and L100 on the surface of a nano-sized curcumin drug to prepare a core-shell nanomedicine with pH-responsive drug-releasing behavior for anti-inflammatory purposes.¹¹²

Figure 3.

pH-triggered release of nanodrug carriers for the treatment of proctitis. (A) Schematic illustration of colon-targeted delivery of sulfasalazine-loaded amidated pectin microparticles in an Eudragit S100-coated capsule. The Eudragit S100 coating dissolves at colonic pH, followed by swelling and erosion of the amidated pectin microparticles and subsequent release of sulfasalazine in the colon. Reprinted with permission from ref.110 Copyright 2022 American Chemical Society. (B) pH dependent release of Melatonin from Chitosan nanoparticle (Mel-CSNPs) or Eudragit-S100-coated Chitosan nanoparticle (Mel-EUCSNPs). Reprinted with permission from ref.113 Copyright 2022 Elsevier B.V. (C) Preparation and characterization of MTX@hCEP. Black solid arrows: Indicate material transformations, preparation steps (eg, the formation of MTX@PLGA NP, and the preparation processes of MTX@EP and MTX@CEP), or signaling pathways. Blue solid arrows: Represent the administration route of the drug MTX@hCEP or its delivery process within the colonic lumen. Bold black arrows: Indicate the restoration or transition of inflamed epithelial cells towards a normal state under the action of the drug. Red solid arrows: Indicate the downregulation of expression or levels of specific indicators (eg, MPO, TNF-α, IL-6, IL-1β, iNOS, and Th1 cell activity), signifying reduction or inhibition. Green solid arrows: Indicate the upregulation of expression or levels of specific indicators (eg, ZO-1 and Occludin expression), signifying increase or promotion. Dashed arrows: Represent potential indirect effects, hypothetical signaling pathways, or mechanisms of action that are not yet fully elucidated. Reprinted with permission from ref.109 Copyright 2023 Elsevier B.V.

Abbreviations: Mel-CSNPs, melatonin-loaded chitosan nanoparticles; Mel-EUCSNPs, Eudragit S100-coated melatonin-loaded chitosan nanoparticles; MTX, Methotrexate; PLGA NP, Poly(lactic-co-glycolic acid) Nanoparticle; CEP, chitosan/Eudragit S100/PLGA nanoparticles; EDC, 1-Ethyl-3-(3-dimethylaminopropyl)carbodiimide; NHS, N-Hydroxysuccinimide; MTX@PLGA NP, Methotrexate-loaded Poly(lactic-co-glycolic acid) Nanoparticles; MTX@CEP, Methotrexate-loaded chitosan/Eudragit S100/PLGA nanoparticles; MTX@EP, MTX-loaded Eudragit S100/PLGA nanoparticles; TNF-α, Tumor Necrosis Factor-alpha; IL-6, Interleukin-6; IL-1β, Interleukin-1 beta; iNOS, Inducible Nitric Oxide Synthase; ZO-1, Zonula Occludens-1; MPO, Myeloperoxidase; IFN-γ, Interferon-gamma; IL-10, Interleukin-10; TGF-β, Transforming Growth Factor-beta; Treg cell, regulatory T cell; CS, Chitosan; HA, Hyaluronic Acid; ES100, Eudragit S100; MTX@hCEP, MTX-loaded HA-CS/ES100/PLGA nanoparticles; hCEP, HA-CS/ES100/PLGA nanoparticles; CD44, cluster of differentiation 44; Th1, T helper 1 cell.

Polysaccharide-Containing IBD Nanomaterial Drug Carriers

In addition to Eudragit-containing nanomaterial drug carriers, polysaccharides such as cellulose, alginic acid, and chitosan also undergo physical structural changes due to changes in their own intermolecular forces when the pH value changes, which in turn plays a role in physical encapsulation and switching of drug molecules^113,114 (Figure 3B). Shin et al found that chitosan, as a highly adhesive polysaccharide, is soluble at gastric acidic pH = 2 but insoluble at small intestinal pH = 5.5, so when chitosan enters the colorectal, it can be degraded by the local intestinal flora and release drugs.¹¹⁵ On this basis, Manconi further enhances the stability of the drug in the acidic environment of gastric juice by encapsulating the chitosan nanomaterial drug carrier through lipid vesicles.¹¹⁶ In addition to chitosan, Wang also prepared a nanomaterial drug carrier formed by crosslinking sodium alginate (SA) and N-succinyl chitosan (NSC) for encapsulating quercetin and demonstrated its pH-sensitive drug release properties in vitro drug release assays.¹¹⁷

Composite IBD Nanomedicine Carriers

In addition to using a single pH-sensitive nanomaterial drug loading, Lv et al constructed a composite nanomaterial drug carrier loaded with MTX anti-inflammatory drugs using PLGA, acrylic resin S100 (ES100), chitosan, and hyaluronic acid and demonstrated its pH-sensitive drug release properties at the colorectum through in vitro and in vivo experiments¹⁰⁹(Figure 3C).

The Microenvironment-Triggered Release

Microbiota-responsive NDDS represent an emerging strategy for IBD treatment, primarily leveraging the unique enzymatic environment of the gut microbiota for precise drug release. Among these, enzyme-responsive nanocarriers utilize colonic bacterial enzymes (such as oxidoreductases, glucuronidases, glycosidases, and dextranases) to selectively degrade specific polysaccharides and polymeric materials.¹¹⁸ These systems are often constructed from sensitive polymers like chitosan, dextran, guar gum, and alginate, which remain stable in the upper gastrointestinal tract but degrade upon encountering enzymes in the colon, facilitating site-specific drug release. Pullulan is a probiotic that is degraded by pullulan enzymes produced by colonic microorganisms, imparting stimulation-sensitive properties to nanocarrier microorganisms.⁵² YPs are yeast cell wall particles that can stably pass through the gastric environment and be degraded by β-glucanase, resulting in good gastric stability and sensitive release of β-glucanase¹¹⁹(Figure 4). Aided by the abundant gut microbiota around the colon, they could be digested to short-chain fatty acids and carbon dioxide, thus releasing the loaded drugs. It is an important strategy that ensures the YPs release loaded drugs to UC lesions.

Figure 4.

The microenvironment-triggered release of nanodrug carriers for the treatment of proctitis. Schematic Diagram of HA/PEI-RH NYPs in the Treatment of UC. Thick green arrows represent the delivery process of the drug within the intestinal tract and its ultimate therapeutic effects. Red dotted circles indicate magnified regions displaying the detailed molecular mechanisms of nanoparticle uptake and intracellular signaling. The left panel shows the overall delivery pathway from oral administration through the intestinal mucosa to the inflamed colon tissue. The right panel provides a magnified view of the circled regions, illustrating the CD44 receptor-mediated cellular uptake by macrophages and the subsequent TLR4-MyD88-p65 signaling cascade leading to the transcriptional regulation of pro-inflammatory cytokines (TNF-α, IL-1β, IL-6, and iNOS). Thin green arrows indicate the direction of signaling pathways within macrophages. Red T-shaped arrows denote inhibitory effects. Black downward arrows indicate a decrease in the expression of relevant inflammatory factors. Reprinted with permission from ref.119 Copyright 2021 American Chemical Society.

Abbreviations: YPs, yeast cell wall microparticles; OVA, Ovalbumin; PEI, Polyethylenimine; HA, Hyaluronic acid; RH, Rhein; NPs, Nanoparticles; TLR4, Toll-like receptor 4; MyD88, Myeloid differentiation primary response 88; TNF-α, Tumor Necrosis Factor-alpha; IL-1β, Interleukin-1 beta; IL-6, Interleukin-6; iNOS, Inducible Nitric Oxide Synthase; NYPs, nanoparticles entrapped in yeast cell wall microparticles; HA/PEI-RH NPs, HA- and PEI-modified rhein-loaded ovalbumin nanoparticles; CD44, cluster of differentiation 44; p65, NF-κB p65 subunit.

Comparative Analysis of Different Triggered-Release Nanocarriers in IBD Therapy

In the management of IBD, stimuli-responsive nanocarriers exhibit pronounced disparities in targeting precision and therapeutic mechanisms depending on their activation modalities. Conventional pH-triggered formulations address gastric degradation but their efficacy is limited by the complex and heterogeneous pH landscape in IBD. While some intestinal segments may exhibit normal or even elevated pH, active inflamed regions, particularly in the colon, often present with significant acidosis (eg, pH 2.3–3.4 in active UC).^120,121 This localized acidosis can impede polymer dissolution and drug liberation, making therapeutic outcomes heavily reliant on passive payload entrapment. Furthermore, the microscopic resolution needed to distinguish inflamed from quiescent mucosa is inherently lacking in these conventional systems. Conversely, ROS-responsive nanoplatforms exploit the 10- to 100-fold oxidative surge within mucosal lesions to achieve cellular-level discrimination, preferentially accumulating within activated M1 macrophages and neutrophils. Beyond on-demand payload release, these redox-active matrices (eg, catalytic nanozymes and polythioketal frameworks) exert intrinsic bioactivity by scavenging intracellular ROS, downregulating NF-κB and NLRP3 inflammasome cascades, and repolarizing macrophages from an M1 to an M2 phenotype, thereby markedly augmenting epithelial barrier repair compared to inert pH-responsive monotherapies. Concurrently, enzyme- and microbiota-instructed systems achieve localized bioactivation via colonic proteases or bacterial hydrolases, though pronounced dysbiosis during active IBD can compromise degradation kinetics and release predictability. Nevertheless, these single-stimulus platforms share a pervasive vulnerability regarding premature leakage and unintended activation under complex gastrointestinal conditions prior to reaching target lesions. During transit, enteric coatings or pH-dependent micelles frequently undergo partial shedding or premature disassembly within the proximal small intestine (pH 6.5–7.0), or suffer burst release driven by bile-salt-mediated micellar disruption. Similarly, ROS-cleavable bonds (such as thioketals and boronic esters) face premature oxidation by physiological luminal oxygen and trace baseline oxidants during transit, whereas enzyme-sensitive linkages (eg, esters, azo bonds, and peptide sequences) remain susceptible to non-specific cleavage by luminal trypsin or non-pathogenic commensal microflora.^122,123 Collectively, single-stimulus platforms rarely reconcile the competing demands of physiological transit stability and lesion-specific accumulation, as overcoming premature leakage typically requires dense structural shielding that inherently blunts responsiveness at inflammatory sites.^59,64 Consequently, engineering hierarchically structured multi-stimuli nanocarriers—such as compartmentalized core-shell architectures coupling sequential pH-enteric shielding with lesion-triggered ROS degradation—represents a pivotal strategy to maximize mucosal bioavailability while preventing off-target systemic exposure. While introducing multi-stimuli gating enhances spatiotemporal precision, this approach inherently demands careful optimization to balance synthetic complexity and batch-to-batch reproducibility during clinical translation.

Targeting Mechanism of IBD Nanomaterial Drug Carriers

Since 2001, when Lamprecht et al first demonstrated that nanomedicines with a size of 100 nm can target the site of inflamed colonic mucosa in mice,^124,125 the targeting of nanomaterial drugs has become the focus and hotspot of scientific research. From a histological perspective, passive targeting relies heavily on the disrupted mucus layer, enhanced intestinal barrier permeability that permits carrier extravasation, and the infiltration of mucosal immune cells (such as neutrophils and macrophages) that ingest accumulating nanocarriers.^29,126,127 Critically, passive targeting provides foundational anatomical accumulation but remains inherently constrained by inter-individual disease heterogeneity, continuous mucosal sloughing, and inadequate cellular-level specificity, which often restricts payload delivery to non-resolving phagocytic sequestration rather than targeted epithelial restoration.¹²⁸ Conversely, active targeting introduces molecular precision through ligand-receptor or electrostatic interactions that enhance cellular internalization; however, it invariably depends on upstream passive accumulation to access target epitopes, and the complex surface functionalization often exacerbates non-specific protein corona formation and off-target immunological clearance.¹²⁹ Due to the targeted nature of nanomedicines, nanomaterial drug delivery systems have shown great potential in the treatment of IBD. However, it is not accurate enough to deliver nanomaterial drugs into IBD through the above passive targeting, and with the innovation of various targeting technologies, correlation research on active targeting of nanomaterial drugs has also emerged. In the following sections, we will separately describe the mechanisms of nanomaterial drug carriers, including charge targeting, receptor targeting, antibody targeting, and cytokine targeting.

Passive/Inflammation-Driven Targeting-IBD Nanomaterial Drug Carriers

In the treatment of IBD, a fundamental strategy for nanodrug carriers exploits the pathophysiological alterations inherent to intestinal inflammation. IBD-affected regions are characterized by increased vascular permeability, impaired lymphatic drainage, and enhanced recruitment of immune cells. These inflammatory hallmarks create a permissive microenvironment that enables nanoparticles to preferentially accumulate at lesioned sites through passive mechanisms, namely the enhanced permeability and retention (EPR) effect. By contrast, intact healthy mucosa maintains normal barrier function and efficient lymphatic clearance, which restricts nanoparticle extravasation and promotes rapid systemic elimination. This inflammation-enhanced accumulation represents a passive, non-selective process that does not require ligand-receptor interactions or surface chemical modifications. Rather, it capitalizes on the intrinsic pathological features of the disease microenvironment. However, the degree of preferential accumulation at diseased versus healthy mucosa is contingent upon the severity of vascular remodeling and local inflammation intensity. In acute active IBD with prominent vasculitic changes, EPR-mediated enrichment can achieve substantial selectivity; however, in mild or remission-phase disease, the differential vascular permeability diminishes, reducing the magnitude of preferential mucosal targeting.

Physicochemical Targeting-IBD Nanomaterial Drug Carriers

Beyond passive vascular accumulation, the physicochemical properties of nanocarriers, particularly particle size and surface charge, significantly influence their distribution and cellular interactions within IBD inflammatory sites. Nanoparticle size is a critical determinant of in vivo biodistribution and tissue penetration. Nanoparticles within specific size ranges (eg, 20–200 nm) more readily permeate compromised intestinal mucosal barriers, leading to increased accumulation in inflamed regions.¹³⁰ This size-dependent penetration reflects passive physics—nanoparticles of appropriate dimensions passively traverse disrupted epithelial tight junctions and altered mucus architecture via physicochemical diffusion, not through active molecular recognition. Once nanoparticles passively accumulate in inflamed tissues via barrier disruption, their subsequent interactions with immune cells become relevant. Mucosal macrophages possess an intrinsic capacity to internalize nanoparticles through multiple endocytic pathways—including clathrin-mediated endocytosis, caveolin-mediated endocytosis, macropinocytosis, and phagocytosis—with internalization efficiency varying according to particle size.¹³¹ Recent mechanistic reviews have established that nanoparticles with diameters between 50–200 nm are efficiently internalized by macrophages via clathrin- and caveolin-mediated pathways, whereas particles <10 nm typically passively penetrate epithelial barriers without triggering significant endocytic uptake, and particles >500 nm are rapidly cleared by the reticuloendothelial system.¹³¹ Critically, this size-dependent endocytic preference is not a targeting mechanism, but rather an inherent biophysical consequence of macrophage endocytic machinery. It reflects passive, non-specific cellular internalization driven by cell-size sensitivity and particle-membrane interactions, not molecular recognition. These uptake patterns should not be misinterpreted as evidence of “targeted delivery”—they represent the passive, non-specific outcome of nanoparticle size interacting with cellular internalization mechanisms.

Concurrently, nanocarrier surface charge characteristics, categorizable into cationic and anionic delivery systems, exploit distinct electrostatic interactions within the gastrointestinal tract. The healthy colonic mucosal barrier comprises negatively charged carbohydrates, sulfate residues, sialic acid, and mucin glycoproteins, whereas the inflamed colorectal mucus layer exhibits altered charge landscapes. Positively charged nanocarriers preferentially adhere to the healthy negatively charged mucosa via non-specific electrostatic attractions, as demonstrated by Lautenschlager et al using chitosan-based nanoparticles.^132–135 However, this high electrostatic affinity often causes excessive immobilization within the superficial mucus meshwork, resulting in premature payload leakage and rapid clearance via accelerated mucus turnover. Moreover, strong positive surface charges pose inherent cytotoxicity concerns, including disruption of epithelial plasma membranes and non-specific immune cell activation, which significantly restrict the chronic application of cationic systems in IBD. In contrast, negatively charged nanomedicines can precisely attach to specific positively charged proteins and show greater permeability to the intestinal mucosa. Antoni and Larsson found that inflammatory colitis not only has negatively charged carbohydrates but is also rich in some positively charged proteins, such as transferrin and eosinophil cationic proteins, due to damage to epithelial cells.^136,137 Thus, negatively charged nanomedicines can also show preferential adhesion to inflamed areas through electrostatic interactions with these positively charged proteins. Jubeh et al reported that anionic liposomes are 2-fold more likely to adhere to inflamed mucosa than neutral or cationic liposomes. In healthy colonic mucosa, cationic liposomes exhibit preferential binding, which is 3-fold higher than that of neutral or anionic liposomes.²⁸ Similarly, Beloqui et al reported that negatively charged NPs composed of budesonide invaded deeper into the mucosa in the colitis group while remaining only superficial in healthy mucosa.¹³⁸ Nevertheless, these electrostatic interactions must be interpreted with caution. While preferential accumulation is frequently described as electrostatic targeting, it fundamentally represents non-specific charge-mediated affinity rather than receptor-selective molecular recognition. In the dynamic luminal environment, anionic nanocarriers can interact prematurely with positively charged luminal peptides, soluble glycoproteins, and bile acids, which neutralize surface charges and diminish mucosal binding.¹³⁹ Consequently, enhanced therapeutic outcomes or superficial mucosal retention achieved by charge-based systems should not be uncritically interpreted as verified targeted delivery, as they lack direct confirmation of tissue penetration and target-cell uptake.

Ligand/Receptor Targeting-IBD Nanomaterial Drug Carriers

Receptor targeting is another commonly used targeting method of nanomedicine. When colonic inflammation occurs, distinctive receptors are markedly upregulated on inflamed enterocytes, vascular endothelial cells, and infiltrating immune cells, including CD44, folate receptors (FR), transferrin receptors (TfR), mannose receptors (MR/CD206), macrophage galactose-type lectin (MGL), CD98, and F4/80.^38,140,141 Ligand-functionalized nanocarriers exploit these pathological signatures to achieve preferential cellular internalization via receptor-mediated endocytosis, markedly enhancing drug accumulation at inflammatory foci.¹⁴² For example, hyaluronic acid (HA)-decorated nanocarriers exhibited enhanced localized accumulation and therapeutic benefits in inflamed epithelial tissues, an outcome primarily attributed to CD44 recognition;^93,143 folate-conjugated copolymers mediated increased therapeutic retention within the inflamed colorectal mucosa, presumably facilitated by elevated macrophage uptake;³⁸ and mannosylated or galactosylated systems achieved superior mucosal anti-inflammatory efficacy (eg, delivering TNF-α siRNA or dexamethasone) via presumed MR or MGL engagement.^29,49,52,144 Crucially, such improved therapeutic outcomes should not be uncritically equated with verified receptor-mediated targeting, as inflamed lesions inherently promote non-specific particle entrapment and macrophage phagocytosis. In the absence of direct quantitative biodistribution profiling, target-cell colocalization, or competitive receptor-blocking assays, these efficacy gains reflect presumed targeting mechanisms and localized drug exposure rather than experimentally demonstrated active homing.

However, notable heterogeneity of receptor expression inherently exists across IBD subtypes and different clinical stages.¹⁴⁵ Pathologically, UC presents mucosal-restricted, continuous inflammation with prominent luminal upregulation of epithelial CD44, transferrin receptor (TfR), and myeloid folate receptor-β(FR-β), which favors direct apical recognition by orally administered carriers.²⁴ In contrast, CD is characterized by transmural skip lesions and granulomas; effective targeting here relies heavily on homing-related markers (α4β7 integrin/MAdCAM-1, CCR9) and deep-seated macrophage markers (CD206) to achieve full-thickness tissue penetration.^64,146 Temporally, target expression is highly dynamic: inflammatory receptors (eg, ICAM-1, CD44) peak during acute flares but drop precipitously during clinical remission, posing a significant risk of target loss for single-ligand systems.¹⁴⁷ In chronic or stricturing stages, acute inflammatory markers are often superseded by fibrotic targets such as fibroblast activation protein (FAP).¹⁴⁸ Coupled with the harsh gastrointestinal milieu that threatens ligand stability, addressing this spatiotemporal heterogeneity remains a central challenge, necessitating the development of multi-stage adaptive or dual-ligand nanoplatforms.

Antibody Targeting-IBD Nanomaterial Drug Carriers

Beyond small-molecule and carbohydrate ligands, the high affinity and single-epitope specificity of monoclonal antibodies (mAbs) and antibody fragments provide an alternative rationale for mucosal nanomedicine. During intestinal inflammation, mucosal microvascular endothelial cells prominently upregulate cell adhesion molecules, including E-selectin, P-selectin, intercellular adhesion molecule-1 (ICAM-1), and vascular cell adhesion molecule-1 (VCAM-1). To exploit these vascular signatures, Sakhalkar et al functionalized poly(lactic acid)-poly(ethylene glycol) (PLA–PEG) nanoparticles with mAbs directed against inducible endothelial cell adhesion molecules, which was designed to promote localized particle adherence to activated endothelial cells at sites of vascular inflammation.¹⁴⁹ Concurrently, the pathological overexpression of cell-surface markers on enterocytes and infiltrating immune cells has driven the formulation of immunonanocarriers. Xiao et al incorporated anti-CD98 antibodies into hydrogel microspheres and orally administered polymeric nanoparticles to engage the elevated CD98 expression on inflamed colorectal epithelial cells and mucosal macrophages, leading to reduced tissue damage and enhanced therapeutic benefits in experimental colitis.^150,151 Similarly, preclinical investigations have explored various antibody-modified nanocarriers—including anti-CD4 mAb-functionalized PLGA nanoparticles reported to interact with pathogenic CD4⁺ T-cell subsets in rectal inflammation,¹⁵² anti-β7 integrin platforms aimed at gut-homing leukocytes,¹⁵³ anti-F4/80-decorated nanocarriers intended for mucosal macrophages,¹⁵⁴ and anti-Ly6C-conjugated formulations directing agents toward inflammatory myeloid populations¹⁵⁵ —each yielding notable symptom alleviation in experimental colitis models (Figure 5).

Figure 5.

Receptor and antibody targeting-IBD nanomaterial drug carriers. Arrows indicate the cellular targets or sites of action for the drug. Modified from ref.77 Copyright 2021 Chinese Pharmaceutical Association and Institute of Materia Medica, Chinese Academy of Medical Sciences. Production and hosting by Elsevier B.V.

Abbreviations: scCD98, single-chain CD98 antibody; FA, folic acid; PLGA, poly(lactic-co-glycolic acid); PLA, Poly (lactic acid); PEG, Poly(ethylene glycol); NPs, nanoparticles; MGL-2, macrophage galactose-type lectin 2; F4/80, F4/80 antigen (ADGRE1/EMR1), a murine macrophage marker; VCAM-1:vascular cell adhesion molecule-1; ICAM-1:intercellular adhesion molecule 1; ICAM-2:intercellular adhesion molecule-2; MadCAM-1, mucosal vascular addressin cell adhesion molecule-1; CD44, cluster of differentiation 44; CD98, cluster of differentiation 98;mAb, monoclonal antibody; MPO, myeloperoxidase; ROS, reactive oxygen species; IL-1β, interleukin-1 beta; IL-6, interleukin-6; TNF-α, tumor necrosis factor-alpha; ECAM, endothelial cell adhesion molecules.

In addition, cytokines, as secreted proteins produced by immune cells, play an important role in maintaining and regulating intestinal physiological immune homeostasis and pathophysiological processes. Cytokines can be divided into different subclasses based on their conformation, including chemokines, interferons (IFNs), interleukins (ILs), and tumor necrosis factor-α (TNF-α). With the development of immunology, studies have found that elevated expression levels of TNF-α, interleukin 1β(IL-1β), and interleukin 12/23 are common in the active stage of IBD disease.¹⁵⁶ In nanotherapeutic design, cytokines and their signaling axes serve as targets through two distinct paradigms: utilizing anti-cytokine antibodies as encapsulated bioactive payloads or as surface-anchored targeting moieties. For instance, macromolecular biologics such as IFX have been encapsulated within inorganic frameworks or ginger-derived exosome-like nanocarriers to shield the protein against luminal enzymatic degradation, demonstrating superior local tissue retention and therapeutic efficacy compared with systemic intravenous IFX administration.^157–159 Following a surface-functionalization approach, Zhang et al engineered dietary fiber-derived nanocarriers surface-conjugated with anti-IL-1β mAbs and entrapped them within alginate hydrogel microspheres, achieving enhanced localized anti-inflammatory responses in murine colitis.⁵¹

Notwithstanding these reported therapeutic improvements, a rigorous distinction must be established between demonstrated antibody-mediated molecular targeting and presumed accumulation inferred solely from therapeutic outcomes. In many published reports, augmented anti-inflammatory efficacy is uncritically attributed to active antibody guidance; however, direct experimental validation—such as competitive epitope-blocking assays, quantitative ex vivo tissue biodistribution, or flow-cytometric confirmation of cell-specific internalization—is frequently lacking. Given the compromised epithelial barrier in active IBD, non-functionalized nanoparticles also passively accumulate in inflamed mucosal lesions via impaired clearance kinetics and non-specific macrophage phagocytosis. Furthermore, antibody-functionalized nanocarriers face profound translational hurdles in the intestinal milieu: exposure to luminal proteases (eg, pepsin, trypsin, and metalloproteinases) can rapidly degrade immunoglobulin chains, random chemical conjugation often compromises antigen-binding fragment (Fab) orientation, and exposed crystallizable fragments (Fc) risk triggering non-specific Fc-receptor-mediated clearance and immunogenicity. Therefore, without rigorous verification of targeted biodistribution kinetics and cellular colocalization, superior therapeutic responses in mucosal inflammation should be characterized as enhanced localized therapeutic action rather than definitive proof of antibody-guided homing.

Microbiota-Responsive/Directly Acting on Microbiota Strategies

Microbiota-responsive and direct microbiota-targeting strategies represent emerging paradigms in IBD nanomedicine, moving beyond conventional tissue or cellular delivery. These approaches leverage the dysbiotic microbiota landscape inherent to IBD, utilizing microbial imbalance either to trigger drug release or as a direct therapeutic target to restore microbial community stability. The dysbiosis characteristic of IBD—marked by a reduction in beneficial bacteria such as Lactobacillus and Faecalibacterium, and an overgrowth of pathogenic species—is not merely a consequence of inflammation but actively perpetuates disease through altered metabolite production, impaired barrier function, and dysregulated immune responses.¹⁶⁰ Consequently, targeting the microbiota itself offers a distinct therapeutic opportunity that complements host-directed anti-inflammatory strategies.

Microbiota-responsive systems exploit the unique metabolic and enzymatic profiles of dysbiotic flora to achieve spatially and temporally controlled drug release. For instance, pathogenic bacterial enzymes or microbial metabolites present in inflamed regions can specifically trigger the release of therapeutic agents, either by breaking down drug carriers (cleaving prodrug linkers) or by promoting drug liberation independently of pH changes (pH-independent cargo liberation) at lesion sites. This strategy inherently provides selectivity without requiring specific ligand-receptor recognition on host cells.

Concurrently, direct microbiota-targeting approaches employ nanocarriers functionalized with molecules that “home” to specific microbiota—such as bacterial outer membrane proteins or bacterial extracellular vesicles (BEVs)—to deliver therapeutic payloads specifically to beneficial commensal species. Biomimetic nanoparticles coated with bacterial membranes demonstrate homologous targeting: for example, Lactobacillus rhamnosus (LGG)-derived BEV-coated lipid nanoparticles preferentially deliver miRNA cargo to their parent Lactobacillus species, while minimizing unintended uptake by pathogenic bacteria like Escherichia coli.¹⁶¹ This specific mechanism relies on inherited membrane proteins and antigens from the source bacteria, enabling bacteria-to-bacteria recognition without interference from host immune clearance.¹⁶²

When combined with conventional immunosuppressants, BEV-functionalized nanoparticles carrying microbiota-modulating payloads (eg, miRNA that enhances the production of beneficial metabolites such as indole-3-carboxaldehyde and short-chain fatty acids) have shown superior efficacy in both acute and chronic colitis models. These systems restore microbial balance, enhance epithelial barrier integrity, and downregulate pro-inflammatory cytokines (IL-6, TNF-α), simultaneously elevating protective metabolites.¹⁶¹ Thus, these microbiota-centric approaches represent a conceptually distinct targeting modality that operates orthogonally to the passive, physicochemical, and receptor-mediated active targeting strategies discussed earlier.

Cellular/Therapeutic Targets

In the realm of nanomedicine, while previous sections primarily explored targeting strategies from the perspective of molecular recognition, the ultimate therapeutic efficacy of these strategies hinges on which cells within the inflamed tissue actually process these nanocarriers. At the cellular level, drug delivery is defined by the nanocarriers’ ability to be internalized by target cells or to stably adhere to their surfaces, thereby concentrating the drug within the target cells and preventing its diffuse interstitial leakage. However, it is crucial to emphasize that while these drug formulations achieve significant intracellular therapeutic effects, their predominant in vivo distribution often results from local vascular extravasation and cellular phagocytosis, rather than from explicit molecular signals actively guiding them to specific cells.

Mucosal macrophages are among the first immune cells encountered by extravasated nanoparticles in inflamed areas. They possess inherent phagocytic capabilities that can be harnessed for drug delivery. For instance, Zhang et al demonstrated that yeast-derived nanocomposites utilize surface β-glucan domains to engage Dectin-1 receptors on mucosal macrophages, promoting cellular phagocytosis to deliver MnO₂ and H₂S intracellularly. This mechanism suppresses downstream NOX4 and p38 MAPK pathways, revitalizes mitochondrial function, and promotes M2 macrophage polarization, thereby alleviating inflammation.¹⁶³ Similarly, Zhong et al capitalized on this macrophage characteristic, designing nanoparticles that carry E3 ligase, MAP3K1, and OPTN into cells to specifically clear NLRP3 inflammasomes, thereby arresting inflammatory responses.¹⁶⁴ Beyond phagocytic cells, nanomedicines can also target mucosal lymphocytes. Tan et al engineered small-diameter polymeric micelles (TPGS-PBTE) capable of delivering luteolin to helper and regulatory T-cells, subsequently correcting the pathological imbalance among Th1, Th2, Th17, and Treg lineages.⁹¹ Furthermore, broader multi-lineage cellular delivery has been reported. Wang et al developed nanomedicines that, through local mucosal accumulation, can access infiltrating neutrophils, dendritic cells, and macrophages. This leads to the inhibition of NF-κB and MAPK pathways, reducing TNF-α and IL-6 expression. Concurrently, this approach indirectly preserves the Lgr5⁺ stem cell niche and maintains intestinal microbiota homeostasis.¹⁶⁵

Extending beyond immune cells, nanomedicine targets have also expanded to non-immune mucosal tissues, offering new avenues for tissue repair. For example, addressing neuro-immune axis disruptions caused by enteric nervous system (ENS) impairment and reduced acetylcholine (ACh) signaling, Ren et al engineered hollow mesoporous polydopamine nanoparticles (HMPDA). These nanoparticles accumulate in inflamed intestinal regions, acting as local drug depots to sustain the release of neuroprotective agents and acetylcholinesterase inhibitors. This supports neuronal survival and promotes neuroimmune coordination without requiring direct nanoparticle internalization by neurons.¹⁶⁶ At the intestinal epithelial barrier, plant-derived exosome-like nanoparticles (such as those from ginger) undergo direct cellular internalization by damaged intestinal epithelial cells via intrinsic membrane-lipid affinities, directly accelerating epithelial restitution.¹⁶⁷ Similarly, targeting the mucosal vasculature can be achieved through surface-level cell engagement: Eniola et al functionalized nanocarriers with specific carbohydrate epitopes that bind inducible selectins on inflamed endothelial cell surfaces, thereby arresting particle transport and concentrating the therapeutic payload at the vascular boundary.¹⁶⁸

In summary, the diverse targeting modalities discussed above—ranging from electrostatic interactions to receptor- and antibody-mediated recognition—demonstrate significant promise in enhancing localized mucosal accumulation. However, a mechanistic appraisal reveals that the therapeutic success of these active nanomedicines is inherently governed by deeper physicochemical and kinetic trade-offs, particularly receptor affinity (K_D), ligand surface density, and internalization dynamics. Although high-affinity antibodies against integrins or MAdCAM-1 confer rapid endothelial anchoring, excessive affinity risks a “binding-site barrier” that restricts deeper mucosal penetration.¹⁶⁹ Conversely, low-affinity carbohydrate ligands such as hyaluronic acid and mannose rely on multivalent avidity, requiring strict nanoscale ligand spacing (<10nm) to trigger cooperative receptor clustering.¹⁷⁰ Furthermore, while fast-recycling receptors like transferrin receptors facilitate cytosolic delivery within minutes (t_1/2<15min), scavenger receptors on mucosal macrophages often divert nanocarriers into degradative phagolysosomes.¹⁷¹ To overcome this entrapping bottleneck, nanocarriers must achieve efficient endosomal escape before lysosomal fusion, commonly mediated by proton-sponge osmotic swelling or transient membrane destabilization.¹⁷² Once translocated into the cytosolic compartment, these engineered carriers are poised to exploit intracellular cues—such as reductive glutathione gradients or elevated reactive oxygen species discussed in Section 3—to trigger the spatial-temporal release of their therapeutic payloads. Remarkably, these critical quantitative determinants remain largely uncharacterized across many current formulations; resolving these kinetic parameters thus represents a crucial avenue for the future rational engineering of mucosal nanocarriers. Importantly, preclinical validation of these targeting kinetics relies heavily on chemically induced acute colitis models (such as DSS and TNBS), where acute epithelial sloughing artificially exposes basolateral and endothelial receptors while exaggerating nonspecific carrier extravasation. In contrast, the chronic transmural pathology and dense submucosal fibrosis encountered in human IBD present substantial physical permeation barriers, often compromising the active targeting efficacy observed in acute animal studies (Figure 6).

Figure 6.

Cell-targeting-IBD nanomaterial drug carriers. (A) Macrophage polarization modulation by nanomedicines. (B) T-cell subtype regulation by nanomedicines. (C) Broad immune regulation and epithelial regeneration. (D) Intestinal epithelial cells and mucosal barrier repair. (E) Enteric nervous system components restoration. (F) Vascular endothelial cells targeting and protection. Bold text: Denotes specific cell types, signaling pathway components, or subheadings within the figure. Thick Blue Arrows: Specifically highlight a promoting or enhancing effect of nanomedicines. Red T-shaped Arrows: Denote an inhibitory effect. Black Arrows: Represent the direction of action or interaction. Red Downward Arrows: Signify decreased expression or inhibited activity. Blue Upward Arrows: Indicate increased expression or a promoting effect.

Abbreviations: ROS, Reactive Oxygen Species; NOX4, NADPH Oxidase 4; MAP3K1, mitogen-activated protein kinase 1; OPTN, Optineurin; NLRP3, NOD-, LRR- and pyrin domain-containing protein 3; p38 MAPK, p38 Mitogen-Activated Protein Kinase; miRGS, microRNA sponges; GSDMD, Gasdermin D; M0/M1/M2, Macrophage phenotypes (M0, resting, M1, pro-inflammatory, M2, anti-inflammatory); IECs, Intestinal Epithelial Cells; ENS, Enteric Nervous System; Th1, T helper 1 cell; Th17, T helper 17 cell; Th2, T helper 2 cell; Treg, Regulatory T cell; TNF-α, Tumor Necrosis Factor-alpha; IL-17A, Interleukin-17A; IL-22, Interleukin-22; IL-13, Interleukin-13; IL-10, Interleukin-10; IL-4, Interleukin-4; IL-5, Interleukin-5; TGF-β, Transforming Growth Factor-beta; OXPHOS, oxidative phosphorylation; mtROS, mitochondrial reactive oxygen species;IL-1β, interleukin-1 beta; IL-18, interleukin-18; IL-6, interleukin-6; DCs, dendritic cells;Lgr5, leucine-rich repeat-containing G protein-coupled receptor 5.

Discussion

To date, a variety of drug carriers have been extensively studied. The application of nanomaterial delivery systems provides an efficient and safe way to treat IBD. Although recent advances have been made in nanomaterial drug delivery systems for IBD, existing reviews have primarily focused on individual dimensions of delivery systems—whether drug-loading strategies, stimulus-responsive mechanisms, or targeting approaches. However, these dimensions do not exist in isolation within the physiological IBD microenvironment. By integrating these disparate dimensions within a single analytical framework, this review aims to provide readers with a unified perspective that facilitates understanding of how individual elements relate to and constrain one another in practical nanomedicine design for IBD. This integrated approach complements existing literature by offering a more comprehensive view of the relationships between key elements in delivery, rather than treating them as independent optimization targets. This review systematically elucidates the multidimensional mechanisms underlying nanomedicine delivery for IBD, integrating drug-loading strategies, administration routes, stimulus-responsive release, and active targeting within a unified framework.

NDDS demonstrates considerable promise in enhancing drug bioavailability, reducing systemic toxicity, and achieving inflammation-specific delivery, translating these laboratory advances into clinical therapies remains challenging. Three critical bottlenecks currently hinder the clinical translation and optimization of NDDS for IBD. 1) The first significant challenge lies in the translational limitations of disease modeling. The predominant reliance on chemically induced rodent colitis models (eg, DSS, TNBS)¹⁷³ in preclinical research presents a critical translational bottleneck. Indeed, a substantial portion of the experimental data discussed in this review is derived from such animal models. These models primarily mimic acute chemical epithelial sloughing, which diverges considerably from the chronic, relapsing, and transmural pathology characteristic of human IBD. This often leads to an overestimation of nanocarrier accumulation via the epithelial-enhanced permeability and retention (eEPR) effect, an effect less pronounced in the dense fibrotic stroma and deep fissures observed in clinical IBD patients. Furthermore, interspecies differences in gastrointestinal transit kinetics, mucosal pH gradients, luminal microbiota composition, and immune receptor expression patterns profoundly impact nanoparticle bio-fate and mucosal targeting efficiency in vivo.^174,175 To bridge this critical translational gap, it is imperative to establish standardized, chronic, and humanized multi-cellular evaluation cascades.^176–178 Transformative platforms like microfluidic gut-on-a-chip and organ-on-a-chip represent a paradigm shift. By recapitulating physiological mechanical peristalsis, dynamic fluid shear stresses, personalized patient-derived intestinal epithelium, and human-relevant host-microbiome-immune crosstalk, these biomimetic microdevices offer a physiologically authentic ex vivo testbed. This enables a more accurate de-risking of carrier mucus translocation, cellular internalization, and potential off-target cytotoxicity prior to costly and time-consuming in vivo deployment.¹⁷⁹ 2) The second critical bottleneck concerns challenges in nanomedicine formulation and manufacturing. While targeted NDDS for IBD show immense promise, their clinical translation is currently hampered by insufficient human data and safety concerns. Despite initial human studies showing promise, such as the Phase IIa clinical trial for Nanocort (targeted PEGylated liposomes for moderate to severe active UC in 2016, demonstrating significant effect with few side effects)¹⁸⁰ and early-phase studies for MSC-Exosomes (targeted exosomes for refractory Crohn’s disease perianal fistula, with Phase I in 2022 and Phase II in 2024 indicating safe and effective outcomes),^181,182 as well as an ongoing clinical trial for SinaCurcumin^® (a nano-formulation for mild-moderate UC anticipated to be effective by 2026),¹⁸³ the overall human evidence remains limited. Although novel targeted drugs like vedolizumab have demonstrated good safety and efficacy in clinical settings, and research is actively exploring emerging targets and next-generation antibody solutions, the complexity of NDDS and their pharmacokinetics and biosafety in humans require more thorough validation.¹⁸⁴ Beyond laboratory-scale formulation, a major hurdle lies in overcoming key Chemistry, Manufacturing, and Controls (CMC) bottlenecks, where transitioning from conventional batch synthesis to continuous-flow microfluidics is essential to ensure manufacturing scalability, narrow polydispersity, and batch-to-batch reproducibility under GMP standards.¹⁸⁵ Moreover, developing non-destructive terminal sterilization strategies (to avoid autoclaving- or irradiation-induced particle aggregation) and optimized lyophilization protocols using lyoprotectants are indispensable to prevent premature payload leakage and ensure long-term colloidal shelf-life stability.¹⁸⁶ Furthermore, research should explore the development of NDDS that can deliver agents beyond traditional anti-inflammatory drugs, aligning with emerging clinical strategies that focus on multifaceted IBD management. This includes targeted delivery of therapeutics for restoring gut barrier function, modulating the gut microbiota composition, and correcting metabolic dysregulation that are increasingly recognized as key contributors to IBD pathogenesis and treatment outcomes in clinical guidelines.¹⁸⁷ Additionally, there is a compelling need for NDDS designed for stage-specific intervention, recognizing the distinct therapeutic demands of acute flares versus maintaining remission in IBD. This would involve developing systems capable of rapid, high-dose delivery of anti-inflammatory or immunosuppressive agents during acute exacerbations, alongside sustained, lower-dose delivery of barrier-repairing, microbiota-modulating, or metabolic-regulating agents for long-term remission maintenance and relapse prevention. 3) The third area focuses on future directions for advanced NDDS development. Future progress in NDDS for IBD demands the standardization of animal models and the development of consensus protocols for nanoparticle characterization across institutions to enable meaningful comparison of results and accelerate knowledge integration. Integration of artificial intelligence and machine learning into the design pipeline to predict optimal physicochemical parameters for specific IBD phenotypes, moving toward patient-stratified nanomedicine approaches. Accelerating the transition toward precision nanomedicine by integrating artificial intelligence (AI) and machine learning (ML) throughout the design and translation pipeline.^188,189 Rather than relying on empirical trial-and-error, predictive machine learning models and generative neural networks can systematically decode complex nano-bio interfaces, optimize multi-component formulation ratios, and predict in vivo mucosal pharmacokinetics tailored to distinct clinical endotypes. Coupling these computational paradigms with patient-specific molecular biomarkers—such as local mucosal enzymatic fingerprints, mucosal redox thresholds, and dysbiotic microbial signatures—will enable the rational design of patient-stratified and individualized nanomedicines. Parallel efforts should validate promising candidates in large animal models and initiate early-stage human trials. Establishment of regulatory frameworks and manufacturing standards specific to nanomedicines, enabling reproducible GMP production and market approval. Additionally, mechanistic studies must better elucidate nano-bio interactions in the IBD microenvironment, including how individual variations in microbiota composition and immune status influence NDDS efficacy and patient response.

In conclusion, while NDDS holds genuine therapeutic potential for IBD management, realizing this potential requires overcoming substantial scientific, technical, and regulatory hurdles. This review provides a comprehensive mechanistic framework integrating drug-loading, administration routes, stimulus-responsive triggers, and targeting strategies—offering an approach for more rational nanoparticle design. By addressing the identified bottlenecks through interdisciplinary collaboration—combining advances in materials science, immunology, bioengineering, artificial intelligence, and regulatory science—the field can accelerate the translation of NDDS from bench to bedside. We believe this framework will serve as a valuable reference for guiding future NDDS development and advancing therapeutic options for IBD patients.

Funding Statement

This work was supported by the Key Health Talents in Gusu (GSKY20210220), the Scientific Research Project of Jiangsu Provincial Health Commission (No.H2023146), and the Suzhou Municipal Bureau of Science and Technology (No. SKY2023211).

Author Contributions

All authors made a significant contribution to the work reported, whether that is in the conception, study design, execution, acquisition of data, analysis and interpretation, or in all these areas; took part in drafting, revising or critically reviewing the article; gave final approval of the version to be published; have agreed on the journal to which the article has been submitted; and agree to be accountable for all aspects of the work.

Disclosure

The author(s) report no conflicts of interest in this work.