Abstract
Bioadhesives have found significant use in medicine and engineering, particularly for wound care, tissue engineering, and surgical applications. Compared to traditional wound closure methods such as sutures and staples, bioadhesives offer advantages, including reduced tissue damage, enhanced healing, and ease of implementation. Recent progress highlights the synergy of bioadhesives and immunoengineering strategies, leading to immunomodulatory bioadhesives capable of modulating immune responses at local sites where bioadhesives are applied. They foster favorable therapeutic outcomes such as reduced inflammation in wounds and implants or enhanced local immune responses to improve cancer therapy efficacy. The dual functionalities of bioadhesion and immunomodulation benefit wound management, tissue regeneration, implantable medical devices, and post-surgical cancer management. This review delves into the interplay between bioadhesion and immunomodulation, highlighting the mechanobiological coupling involved. Key areas of focus include the modulation of immune responses through chemical and physical strategies, as well as the application of these bioadhesives in wound healing and cancer treatment. Discussed are remaining challenges such as achieving long-term stability and effectiveness, necessitating further research to fully harness the clinical potential of immunomodulatory bioadhesives.
1. Introduction
Bioadhesives are natural or synthetic polymers engineered to adhere to biological tissues through tailored interactions such as hydrogen bonding, covalent crosslinking, or physical entanglement. [1,2]. Unlike conventional normal adhesives, which often prioritize mechanical strength and durability using synthetic polymers (e.g., silicones, acrylates) and may contain cytotoxic components such as unreacted monomers or solvents, bioadhesives are specifically designed for biocompatibility and safe interaction with tissues, organs, and body fluids [3,4]. In contrast to industrial adhesives, bioadhesives incorporate functional groups (e.g., –COOH, –NH2, –NHS, –SH) and are often biodegradable to accommodate tissue regeneration processes. Recently, bioadhesive technologies have evolved beyond traditional wound closure devices (e.g., sutures and staples) and passive barrier dressings. Innovations in delivery approaches, including hydrogel patches, sprays, and pastes, have enabled their use in various tissue environments and application settings [5]. They have been adapted for applications beyond wound closure, including drug delivery, integration with wearable medical devices, and use as biomaterial scaffolds [6–10].
To extend bioadhesive technologies to a wider spectrum of clinical applications, researchers have incorporated immunomodulatory capabilities into these materials, enabling targeted control over the immune response at the application site. Immunomodulating bioadhesives refer to adhesive materials specifically engineered to direct or modify immune activity – for example, by incorporating immunostimulatory or immunosuppressive signals within the adhesive matrix. This approach combines the inherent adhesive characteristics of bioadhesives, ensuring localized and enduring adhesion, with the ability to enhance or suppress immune reactions in either an antigen-specific or non-specific way.
In contrast, traditional bioadhesive or common wound-healing dressings primarily serve as passive barriers or structural supports to promote tissue adherence or maintain a favorable environment for healing. However, they do not actively regulate the immune response. Immunomodulatory bioadhesive go a step further by actively shaping immune cell recruitment and activity, thereby aiming to deliver localized immunotherapy or reduce unwanted inflammation at the site of application. This synergy between bioadhesion and immunomodulation presents promising opportunities for the development of targeted therapeutic interventions.
In this review, we will focus on the design rationale and applications of these immunomodulatory bioadhesives. The design, characterization and applications of general bioadhesives could be found elsewhere [1,11,12]. We begin by examining key immune processes, including both innate and adaptive responses, and how bioadhesive interact with these systems. We then discuss the fundamentals of normal wound healing trajectory before exploring emerging design strategies for immunomodulatory bioadhesives. Finally, we highlight recent applications of these materials in immunosuppression for implants and transplants, wound healing, and cancer treatments, and offer our perspective on harnessing bioadhesives to direct immune system functions for tissue repair, regenerative medicine, wound care, and targeted therapeutics.
2. Immune responses
The human immune system consists of the innate and adaptive immune system [13–15]. Innate immunity serves as the first line of defense, providing a rapid but non-specific response to pathogens, reacting similarly across multiple exposures. In contrast, adaptive immunity generates immunological memory, allowing for more targeted and efficient responses upon subsequent exposure to the same pathogen [16,17]. A mind mapping of the immune response is shown in Fig. 1A.
Fig. 1.
Schematics of immune responses. (A) Mind map illustrating the immune response, including cellular and humoral components.11 (B) Sequential events leading to the foreign body reaction following biomaterial implantation2.2
Bioadhesives can modulate the immune response through cellular activation, involving phagocytes (e.g., neutrophils, macrophages, and dendritic cells (DCs)) and lymphocytes (B and T cells), or by influencing the humoral responses that involve antibodies and cytokines. By precisely regulating these immune interactions, bioadhesives can be tailored for applications such as wound healing, immunosuppression, and cancer therapy, ensuring efficient and localized immune modulation.
2.1. Cellular and humoral factors in innate immune response
The innate immune system acts as the body’s first line of defense, utilizing physical barriers (e.g. skin and mucous membranes) alongside immune components such as phagocytes and complement proteins to rapidly neutralize infection or tissue damage. A key aspect of this response is the activation of phagocytes, including neutrophils, macrophages and dendritic cells. The detection of danger signals—pathogen-associated molecular patterns (PAMPs) and damage-associated molecular patterns (DAMPs)—initiates phagocyte activation [16]. These signals stimulate tissue-resident immune cells (macrophages and DCs) through pattern recognition receptors, such as scavenger receptors and Toll-like receptors, on the cell surface [16,17]. Upon activation, these immune cells trigger intracellular signaling cascades, induce the production of cytokines, resulting in the recruitment of additional immune cells to the site of infection or injury [14,18]. Bioadhesives can influence this process by mimicking or altering the biochemical landscape of injury sites. For instance, fibrin-based adhesives can replicate the provisional matrix of blood clots, amplifying DAMP signaling to recruit neutrophils and macrophages during the inflammatory phase of wound healing [19].
As first responders from the circulatory system, neutrophils enable pathogen clearance through phagocytosis, where they engulf and destroy pathogens. They also secrete antimicrobial substances and cytokines that promote tissue repair and immune cell recruitment [20]. Following neutrophil activation, monocytes are recruited to the injury site and differentiate into macrophages. These macrophages adopt either a pro-inflammatory (M1) or anti-inflammatory (M2) phenotype, depending on environmental cues. M1 macrophages enhance pathogen clearance and release pro-inflammatory cytokines, while M2 promote tissue repair and regeneration by producing growth factors that support cell proliferation and angiogenesis. Immunomodulatory bioadhesives can be engineered to regulate phagocyte activation and functional state, directing specific immune responses in a manner that supports the healing process. For battling bacteria or fungal infection, bioadhesives can promote the recruitment of neutrophils and macrophages to the affected site, enhancing pathogen clearance. For attenuating excessive inflammation, bioadhesives can polarize macrophage toward the M2 phenotype, which is associated with tissue repair and immune resolution. This shift could lead to increased production of anti-inflammatory cytokines and growth factors, ultimately minimizing excessive inflammation or fibrosis.
Cytokines serve as pivotal chemical messengers, orchestrating immune cell recruitment, activation, and differentiation. Key cytokines include chemokines, interleukins (IL), interferons (IFN), tumor necrosis factor (TNF), and growth factors [21–23]. Chemokines such as CXCL8 direct neutrophil trafficking, while pro-inflmmatory cytokines like IL-1β, IL-6 and TNF-α amplify inflammation. IFNs establish antiviral states, and growth factors such as vascular endothelial growth factor (VEGF) and platelet-derived growth factor (PDGF) drive tissue regeneration [24]. Bioadhesives can modulate this cytokine network by delivering cytokine antagonists to suppress excessive inflammation or by releasing growth factors to promote regeneration [25–28]. This control over cytokine signaling allows bioadhesives to balance pathogen clearance and tissue regeneration, mitigating complications in conditions such as infected burns and diabetic ulcers. The specific roles and functions of cytokines in immune response during healing process will be explored in detail in Section 3.
2.2. Cellular and humoral factors in adaptive immune response
Once innate immune mechanisms recognize and clear pathogens or damaged tissue to some extents, antigen-presenting cells (APCs), particularly primary DCs, initiate the activation of B and T cells, thereby bridging the innate and adaptive immune responses [29]. The adaptive immunity generates pathogen-specific responses, equipping the body with immunological memory to mount a targeted and efficient response upon re-exposure. When infection is resolved, the immune system enters a contraction phase, in which most antigen-specific lymphocyte, undergo apoptosis, leaving behind memory T cells or B cells that ensure long-term immunity.
The adaptive immune system can also be divided into humoral and cellular immunity. Humoral immunity, mediated by B lymphocytes, involves the production of antibodies that neutralize pathogens and facilitate their clearance. Some activated B cells differentiate into memory B cells, ensuring long-term immunity [30]. Cellular immunity is primarily executed by T lymphocytes. Among them, CD8+ cytotoxic T cells target the intracellular peptides (e.g., from viruses or cancer cells) and directly identify and eliminate infected or malignant cells, while CD4+ helper T cells target the extracellular antigens (e.g., from bacteria or other pathogens) and can further differentiate into specialized subsets, including TH1, TH2, TH17 and regulatory T cells (Treg) cells [14,15,30–32]. TH1 cells activate phagocytes, natural killer (NK) cells, and CD8+ cytotoxic T cells, thereby enhancing the defense against pathogens [30]. TH2 cells promote humoral immunity and allergic responses, facilitating antibody production and helminth expulsion. TH17 cells secrete pro-inflammatory cytokines, which contribute to autoimmune and inflammatory conditions. Treg cells produce immunosuppressive cytokines that downregulate excessive immune activation, maintaining immune homeostasis [30].
By modulating the adaptive immune system, bioadhesives provide localized and sustained treatment, offering therapeutic benefits for cancer immunotherapy, autoimmune regulation, and tissue regeneration. In cancer treatment, bioadhesives can be designed to activate T cells, potentiating the immune system to recognize and eliminate tumor cells [33]. By delivering tumor antigens or immunostimulatory cytokines, bioadhesives can activate cytotoxic T cells to direct tumor cell destruction [33,34]. Additionally, bioadhesives can stimulate APCs to improve T cell priming and immune memory, strengthening long-termed anti-tumor immunity [35,36].
Beyond cellular immunity, bioadhesive can also be designed to modulate humoral immunity by incorporate cytokines and antibodies. For instance, the controlled release of immunosuppressive cytokines (e. g., IL-10, TGF-β) can promote immune tolerance, aiding in the management of autoimmune disease and promoting graft acceptance in transplantation [37]. Furthermore, bioadheisve can serve as delivery platforms for monoclonal antibodies, enabling targeted therapy for cancer immunotherapy and infection control [38–43].
Moreover, bioadhesives can facilitate immune cell trafficking, ensuring that immune cells migrate efficiently to target sites for infection clearance, tumor destruction, or tissue repair [44–49]. By incorporating chemokines or other molecules, bioadhesives can direct immune cell movement, improving therapeutic efficacy in conditions requiring localized immune modulation [44].
2.3. Immune response with biomaterials
Any material that comes into contact with the body can potentially trigger an immune response and impact surrounding cells and tissues [50]. Bioadhesive materials, widely used in regenerative medicine, organ transplantation, and implantable wearable devices, must be carefully designed to minimize undesirable immune reactions. An uncontrolled host immune response can lead to foreign body response (FBR), characterized by chronic inflammation, fibrous encapsulation, and potential implant rejection (Fig. 1B).
Upon contact, bioadhesives immediately interact with plasma proteins and platelets, forming a provisional matrix that serves as a chemoattractant gradient for neutrophils, monocytes, and macrophages, triggering an acute inflammatory response [51] (Fig. 1B). High levels of protein adsorption typically led to increased immune cell adhesion, which in turn enhances fibrous encapsulation. Therefore, modifying the surface properties of bioadhesives to regulate protein adhesion could serve as a strategy to modulate immune response and enhance biocompatibility.
If acute inflammation is effectively controlled, fibroblasts and mesenchymal stem cells can be recruited to the site, facilitating a normal wound healing process. However, if infammation persist, it can become chronic and trigger the FBR, characterized by the fusion of macrophages into foreign body giant cells (FBGCs) [52]. This prolonged immune activation often leads to the fibrous capsule development, which can isolate the bioadhesive from surrounding tissues, impeding implant integration [51,52]. Excessive contraction of the fibrotic capsule, combined with sustained inflammation can ultimately result in implant rejection, sometimes necessitating surgical removal to prevent further complications such as infection or tissue damage. To improve implant integration and prevent failure, immunomodulatory bioadhesives can be designed to locally regulate immune responses, mitigating excessive inflammation and reducing fibrotic encapsulation. By tuning bio-adhesive properties, such as surface chemistry, mechanical property and controlled release of immunomodulatory agents, it is possible to create a more favorable microenvironment for implant acceptance and long-term functionality [50].
3. Wound healing
Bioadhesives are frequently used at wound sites to achieve hemostasis and are closely intertwined with the wound healing process. Following an injury or implantation, a well-regulated immune response is essential for restoring tissue homeostasis and ensuring implant integration. However, an uncontrolled immune response can lead to chronic wounds or excessive fibrosis, ultimately compromising healing outcomes [53]. To design immunomodulatory bioadhesives for therapeutic purposes, it is important to understand the natural progression of immune responses along a normal wound healing trajectory.
This section below provides an overview of normal wound healing processes without external interventions which encompass four stages: (1) hemostasis (minutes to hours after injury), (2) inflammation (day 1–3), (3) proliferation (days 4–21), and (4) remodeling (days 21-years) [54,55]. These stages, mediated by a cascade of immune cell and signal activities, occur in a coordinated manner. The cells and protein interactions involved in mediating these responses that occur are summarized in Fig. 2. Dysregulation of these processes could delay wound healing and lead to chronic ulcers and/or excessive scarring. These processes provide the baseline understanding for the translational applications of immunomodulatory bioadhesives discussed in Section 6.
Fig. 2.
Schematics of wound healing stages. (A) Hemostasis, (B) Inflammation, (C) Proliferation, (D) Remodeling, and (E) Additional responses if a chronic wound develops.31
Hemostasis is the first critical stage of the wound healing process, serving to quickly stop bleeding and form a stable clot, which prevents further blood loss and provides a temporary matrix for cellular infiltration (Fig. 2A) [56]. The weak mechanical properties and slow formation of natural blood clots necessitate the development of hemostatic bioadhesives to promote hemostasis and prevent rebleeding. Upon skin injury, the disruption of the epidermal barrier trigger the release of pre-stored IL-1, alerting surrounding cells to the damage [24]. When blood vessels are injured, blood extravasation and localized vasoconstriction occur at the wound site [56]. Platelets in the blood then adhere to exposed extracellular matrix (ECM) components and become activated to form a primary platelet plug, a process known as primary hemostasis. Platelet degranulation triggers the releases of key growth factors, including platelet-derived growth factor (PDGF), epidermal growth factor (EGF), and transforming growth factor-beta (TGF-β). Additionally, it releases inflammatory mediators such as IL-8 (more commonly named chemokine (C-X-C motif) ligand 8, CXCL8), which is a potent neutrophil chemoattractant [55]. These factors, along with pro-inflammatory cytokines like IL-1, play crucial roles in attracting neutrophils to the wound site [24]. The primary hemostasis temporarily seals the damaged blood vessel and helps prevent further blood loss.
To reinforce this plug, secondary hemostasis occurs through a coagulation cascade that converts fibrinogen into insoluble fibrin, forming a mesh that stabilizes the platelet plug into a more robust blood clot [56]. As the clot stabilizes, it undergoes contraction, aprocess driven by activated platelets adhering to fibrin fibres and generating traction forces through contractile proteins such as non-muscle myosin IIA and actin. This contraction reduce clot volume, expels excess serum, and compacts the fibrin network, increading clot density and stiffness. These structural changes effectively sealing the wound and preventing further bleeding [57]. Our research group has reviewed the coagulation mechanisms and the mechanics of blood clotting [58], and characterize the physical properties of the blood clot, including the viscoelasticity, fatigue model and fracture mechanics [59–61]. These blood clots not only achieve hemostasis but also serve as a scaffold for the infiltration of keratinocytes, immune cells, and fibroblasts [55,56]. Recently, bioadhesives have been increasingly utilized for hemostatic management, particularly in the case of non-compressive hemorrhage. The interplay between hemostasis and immune response in biomaterial development should also be considered during wound healing or tissue regeneration [62–64].
The inflammatory phase significantly overlaps with the initial hemostasis stage (Fig. 2B). Tissue-resident immune cells such as DCs and resident macrophages are activated by DAMPs and PAMPs, initiating the inflammatory response [55]. This activation not only directs the chemotaxis of other innate inflammatory cells, but also promotes the maturation of DCs, therefore setting the stage for the activation of adaptive immunity through B and T lymphocytes at the same time.
Neutrophils are the most abundant immune cells in blood and the first line of immune cells recruited to the injury site. Their presence is noted within minutes and peaks in the first 24 h, preventing microbial invasion and clearing wound debris [65,66]. The migration of neutrophils is driven by biochemical cues such as IL-1, PDGF, and CXCL8. Neutrophils fight invading pathogens by releasing antimicrobial agents such as reactive oxygen species (ROS), nitric oxide (iNOS), antimicrobial peptides, and proteases, and engage in phagocytosis aided by neutrophil extracellular traps (NET). Additionally, neutrophils release cytokines and growth factors, including IL-1, IL-6, Tumor necrosis factor α (TNF-α), CXCL-8, monocyte chemoattractant protein-1 (MCP-1), and vascular endothelial growth factor (VEGF) [67,68]. They attract more inflammatory cells to the site, and promote the proliferation of key cells for tissue repair, such as fibroblasts, keratinocytes, and endothelial cells [55].
Monocytes are recruited to wound sites through chemokines and cytokines released by neutrophils. These monocytes, typically arriving between 1 and 3 days post-injury, differentiate into macrophages, which are crucial for inflammatory responses. Macrophages are highly phenotypically heterogeneous and may be converted into pro-inflammatory M1 macrophages or anti-inflammatory M2 macrophages [69]. M1 macrophages are responsible for phagocytosing cellular debris, damaged matrix components, microbial pathogens, and apoptotic neutrophils, while also secreting a variety of pro-inflammatory cytokines, including IFN-γ, IL-1α, IL-1β, IL-6, IL-12, IL-23, CXCLs, and TNF-α [69,70], as well as oxidative metabolites such as iNOS and ROS. These cytokines facilitate the recruitment and activation of innate immune cells, including NK cells, additional macrophages, and helper T cells [71,72]. They also secrete matrix metallopeptidase 2 (MMP-2) and MMP-9 to break down the extracellular matrix, promoting immune cell infiltration [72]. In normal wound healing processes, inflammatory cells such as neutrophils and M1 macrophages undergo apoptosis and disperse from the wound, creating an anti-inflammatory environment conducive to M2-like phenotype [20]. However, when the M1 to M2 transition is impaired, inflammation could persist, characterized by excessive inflammatory cytokines such as TNF-α, reduced pro-healing mediators like IL-10, and the activation and degranulation of CD8+ T cells [73] (Fig. 2E). This condition could result in chronic wounds that are difficult to heal.
As the inflammation resolves, the wound healing process transitions to the proliferation phase (Fig. 2C). During proliferation, M2 macrophages produce growth factors such as PDGF, IGF-1, VEGF, and TGF-β1, supporting cellular proliferation (e.g. fibroblast, and keratinocyte), granulation tissue formation, and angiogenesis [74]. Meanwhile, they produce tissue inhibitors of metalloproteinases (TIMP1) to counteract MMPs, stabilizing the extracellular matrix [75]. T cells are also critical in the resolution of inflammation and tissue remodeling: TH2 cells contribute to anti-inflammatory responses by secreting cytokines such as IL-4, IL-5, and IL-13; Treg produce anti-inflammatory cytokines such as IL-10 and TGF-β, maintaining immune homeostasis [76].
Remodeling is the final phase of the wound healing process, in which granulation tissue might mature into healthy skin or a scar (Fig. 2D). During this phase, myofibroblasts, which develop from fibroblasts in response to mechanical tension and TGF-β signalling, exert contractile forces to bring the edges of the wound together, facilitating wound contraction [55,75]. Macrophages and myofibroblasts release various MMPs and catalyze the production of collagen, essential for local tissue integration [55,75]. This results in reduced cell density and protein synthesis, with the ECM reorganized into larger fibrils aligned along tension lines. Eventually, most endothelial cells, macrophages, and myofibroblasts either leave the wound site or undergo apoptosis, marking the completion of the healing process [75].
At wound sites, bioadhesives not only merge wound edges mechanically but also actively engaging with the local immune system. They interact directly with innate immune cells, such as macrophages and neutrophils, modulating their activation and function.
A detailed discussion on how engineered immunomodulatory bioadhesives is used to regulate these processes can be found in Section 6.
4. Design strategies of immunomodulatory bioadhesives
Immunomodulatory bioadhesives are engineered to precisely control immune responses, enabling specific outcomes depending on the application. In wound healing, the primary focus is on promoting an anti-inflammatory response, through M1 to M2 macrophage polarization. While for wearable implants, the priority is to minimize fibrosis to enhance comfort and extend adhesion. The fibrotic layer can also hinder implant functionality by blocking mechanical, electrical, chemical, or optical communication [31,77]. In cancer therapy or vaccination, bioadhesives aim to enhance immune cell recruitment, including dendritic cells, M1 macrophages, and killer T cells. The design strategies employed in immunomodulatory bioadhesives build upon approaches developed for biomaterials [50,78], which can be categorized into three main areas: (1) Tuning the chemistry and formulation, (2) delivering bioactive molecules or cells, and (3) modifying physical and mechanical properties (Fig. 3).
Fig. 3.
Schematic diagrams of immunomodulation strategies employed by bioadhesives.41
4.1. Tuning chemical compositions and gelation mechanisms
The composition and formulation of hydrogels play a pivotal role in initiating and modulating immune responses. Depending on the desired outcome, biomaterial composition can be adjusted to either enhance inflammation or suppress immune activity. For example, mesoporous silica scaffolds can intensify local inflammation, whereas alginate scaffolds are non-immunogenic and reduce inflammation and foreign body reactions [49]. Incorporating ECM components to mimic the native microenvironment can also modulate immune responses. For example, decellularized extracellular matrix promote pro-regenerative (M2) macrophage activity in skin healing, enhance angiogenesis and reduce fibrosis [79,80], and highly sulphated hyaluronic acid suppresses M1 macrophage activation [81].
The specific polymer chemistries also modulate macrophage phenotypes and control foreign body responses [82,83]. Several studies have employed high-throughput screening and combinatorial chemistry to identify polymers with immunomodulatory properties. For instance, a high-throughput screening of 141 meth(acrylate) and meth(acrylamide) monomers identified molecular fragments like alkoxy groups and methacrylamide, that promote either M1 or M2 polarization, demonstratng the potential for fine-tuning immune responses through material design [82]. Another study utilizing cellular barcoding and sequencing to screen a library of 211 chemically modified alginate analogues [83], while a seperate investigation expanded this approach to a library of 774 alginate modifications, systematically evaluating how functional group modifications influence immune response to implanted polymers [84]. All studies validate their findings through in vivo models, the first study validated the macrophage-modulating properties of selected co-polymers by coating them onto clinical-grade silicone rubber tubes and implanted in rodent model, demontrating their feasiblity for medical application. The second and third studies demonstrated the anti-fibrotic efficacy of optimized alginate formulations in rodent and non-human primate models, highlighting their protential for reducing fibrotic encapsualtion and improving implant integration.
4.2. Delivery of bioactive components
Bioadhesives can serve as carriers for bioactive molecules such as growth factors, cytokines, or cells to locally modulate immune responses, providing a versatile platform for immunodulation strategically adjacent to the targeted tissue site. This strategy is adapted from bio-materials, where scaffolds are designed for controlled, targeted delivery of immunomodulatory factors and cells. Excellent reviews on this subject could be found elsewhere [85]. In wound healing and tissue regeneration purposes, bioadhesives can deliver anti-inflammatory drugs or nanoparticles to manage pro-inflammatory responses, while promoting regeneration by delivering biologics associated with TH2 cells, M2 macrophages, or stem cells [85,86]. M1 to M2 polarization can also be achieved through metabolic reprogramming. M1 macrophages primarily rely on aerobic glycolysis to sustain their pro-inflammatory activity, whereas M2 macrophages prefer mitochondrial respiration over glycolysis. Enhancing oxidative phosphorylation or fatty acid oxidation can promote the transition from M1 to M2, fostering a pro-regenerative immune response [86]. For cancer treatment, the aim is to concentrate immune cells at the tumor site, improving the targeting efficiency of immunotherapies such as immune checkpoint inhibitors [87,88], cancer vaccines [89,90], cytokines or antibodies [91,92]. They also support adoptive T cell therapies by creating environments that promote T cell survival and function [93]. By delivering these agents directly to the tumor, bioadhesives minimize systemic side effects and enhance local immune activity, creating a microenvironment that supports immune cell survival and function [94,95].
4.3. Modification of physical properties
The physical and mechanical properties of bioadhesives—such as stiffness, viscoelasticity, architecture, and surface topology—also play a crucial role in immunomodulation [69]. However, their role in regulating immune cell behavior and wound healing has received less attention, compared to biochemical and cellular strategies. Mechanotransduction, the process by which mechanical cues are transduced into biochemical signals, has been shown to influence immune response, cellular behavior, and extracellular matrix remodeling.
Physical strategies offer advantages due to their simplicity in design, and ease of implementation, making them appealing for clinical translation. Noteworthy, they are favorable for regulatory approval because modification of mechanical or structure properties often requires no therapeutic and biologic components.
Immune cells are mechano-sensitive and respond to mechanical cues within their microenvironment. Increasing matrix stiffness has been shown to promote a pro-inflammatory macrophage phenotype, prolonging inflammation [96–99], and stiff antigen-presenting matrix have been shown to enhance T cell activation and expansion, contributing to a stronger immune response [100]. During the wound healing process, mechanical force can regulate all three phases: inflammation, proliferation, and remodeling [101]. Fibroblasts and myofibroblasts respond to mechanical cues through the TGF-β1 signaling pathway that modulates collagen production and ECM remodeling, ultimately influencing wound closure and scar formation [99,101]. Mechanical force could also regulate angiogenesis, where tensile stress influencing endothelial cell alignment and vessel formation [99].
Bioadhesives with tunable mechanical properties can leverage mechanotransduction to optimize wound healing outcomes [102]. Soft bioadhesives with mechanical properties comparable to those of biological tissues may help reduce excessive immune activation, minimizing chronic inflammation and fibrotic tissue formation [27,103]. Furthermore, bioadhesives can transmit contractile forces, influencing cell migration and tissue integration through force-dependent signaling pathways, thereby modulating the healing process to promote tissue growth and minimize scar formation. However, excessive mechanical stress may also contribute to fibrosis, disrupting normal tissue regeneration and lead to scars or keloid formation [99]. Additionally, size and shape of bioadhesives could influence the FBR. For instance, larger alginate spheres reduced cell adhesion and fibrosis in mice [104]. More about physical strategies to modulate the immune cells response and thwart FBR have been reviewed [69,105–109].
Surface topography can also affect macrophage attachment and protein conformation by alternating the hydrophobicity of the implant or adhesive matrix [110,111]. Hydrogels with zwitterionic materials such as poly(carboxybetaine) and poly(sulfobetaine) resist protein adsorption and foreign body reactions due to their super-hydrophilic surfaces [112,113]. These materials possess ultra-low biofouling properties due to charge interactions from zwitterionic groups. The solvation of their charged terminal groups, along with hydrogen bonds, creates a robust hydration layer that resists non-specific protein adsorption. This is manifested by poly(carboxybetaine methacrylate) zwitterionic hydrogels, when implanted in mice, resisting the formation of a collagenous capsule for at least three months and promoting microvessel formation in the surrounding tissue [114]. However, their hydrophilic nature limits adhesion to tissues, so a bi-layer (Janus) hydrogel design has been reported—one layer for tissue adhesion and the other for antifouling properties [115]. Zwitterionic materials and their derivatives have been reviewed elsewhere [112].
Table 1 summarize the biomaterials that used in immunomodulatory bioadhesives based on their immunogenicity mechanisms. The integration of immunomodulating strategies into bioadhesives opens a vast area of research with significant potential for improving therapeutic outcomes. By tailoring immune responses, bioadhesives can address key challenges in areas such as wound healing, cancer treatment, and interface integration. In the following sections, we will explore three specific applications of immunomodulatory bioadhesives: immunosuppression, wound healing, and cancer therapy.
Table 1.
Materials and modification strategies used for immunomodulatory bioadhesives.
| Category | Examples | Role in Bioadhesive | Mechanism of Immune Modulation | Ref. |
|---|---|---|---|---|
| Inherently Immunogenic | Cyanoacrylate based adhesive | Adhesives | May induce phlebitis, hypersensitivity reactions (Types I and IV), foreign body granulomas, and endovenous glue-induced thrombosis. | [116] |
| Silica | Reinforcement of the matrix | Intensifies inflammation via surface chemistry of silica, promotes M1 macrophage polarization. | [49,117,118] | |
| Inherently Non/Low Immunogenic | Polyacrylic acid (PAAc) | Adhesive matrix | Induces apoptosis without necrosis, but may lead to an unwanted inflammatory response and ROS production. | [119] |
| Alginate | Adhesive matrix | Generally non-immunogenic, may contain endotoxins that trigger an immune reaction if not appropriately purified. | [49,120] | |
| Chitosan | Bridging polymer or adhesive matrix | Generally non-immunogenic, Potential immune reaction due to residual endotoxins if not adequately purified. | [121] | |
| Polyethylene glycol (PEG) | Adhesive matrix | Generally non-immunogenic | [122] | |
| Hyaluronic acid (HA) | Adhesive matrix | Generally non-immunogenic, may contain endotoxins that trigger an immune reaction if not appropriately purified | [123,124] | |
| Immunomodulation via Additives | Polydopamine (PDA)-conjugated hydrogels | Adhesive motif | Scavenges ROS, promotes M2 macrophage polarization via catechol groups. | [125] |
| Immunomodulation via Drug/Nanoparticle-Loaded | Immune checkpoint inhibitor (e. g. Anti-PD-1, anti-CD47 antibodies) | Drugs in adhesive matrix for local immunomodulation | Block inhibitory signals (PD-1/PD-L1, CD47-SIRPα) to reactivate T cells and macrophages. | [126–128] |
| Anti-inflammatory drugs (e.g. steroid related drug) | Drugs in adhesive matrix for local immunomodulation | Suppress pro-inflammatory cytokines (TNF-α, IL-6); promote M2 macrophage polarization. | [26,129] | |
| Metallic Oxide nanoparticle (ZnO, TiO2) | Nanoparticle in adhesive matrix for local immunomodulation | Scavenge ROS, reduce oxidative stress; antimicrobial and anti-inflammatory effects. | [130–132] | |
| Immunomodulation vis Physical Properties | Mechanical properties (stiffness, viscoelasticity) | Adhesive matrix | Influences mechanotransduction of immune cells and fibroblasts. | [69,96–99] |
| Size and shape | Implant device requiring adhesion | Altered the surface area to affect the immune cell adhesion | [104] | |
| Hydrophilicity/Hydrophobicity | Adhesive matrix | Affects protein conformation, cell attachment,
and immune cell interactions. Resist protein adsorption and suppress foreign body response (FBR) via super-hydrophilic surfaces. |
[110–115] |
5. Designs and applications for immunosuppression
Immune suppression is essential for the long-term success of medical implants and organ transplants. Fibrotic layers at the implant-tissue interface impact performance—while mild fibrosis can provide mechanical stability, excessive fibrosis hinders function and may lead to the rejection of the implant. In Table 2, we summarize in vivo studies of bioadhesives, characterizing the thickness of fibrotic layers observed through histological staining quantified using ImageJ. Bioadhesives have the potential to limit fibrosis at their interfaces with biological tissues. This point is highlighted in a recent study where a double-network hydrogel bioadhesive, comprising covalently cross-linked poly(acrylic acid) (PAAc) carrying NHS ester and physically cross-linked poly(vinyl alcohol), demonstrates strong mechanical integration with tissues. This bioadhesive mitigates fibrous capsule formation in various animal models over 12 weeks in vivo (Fig. 4A). The reduction in fibrosis is attributed to significantly lower infiltration of inflammatory cells at the adhesive implant-tissue interface compared to non-adhesive implants, thereby minimizing fibrotic capsule formation [77].
Table 2.
Biological responses of implanted bioadhesive material.
| Main Polymer Network | Additives/Moieties | Adhesion Technique | Applications | Stiffness | Adhesion Strength | Animal Model and Implantation Site | Fibrotic Capsule: width and duration | Regulation of Cytokine | Inflammation | Ref. |
|---|---|---|---|---|---|---|---|---|---|---|
| Silicone | absorbent additives (starch, carbopol) eliminate surface water | covalent bond with tissue and silicone device
via siloxane bond hydrophobic interactions
|
adhering silicon device to wet tissue, specifically designed for long-term fixation in wet tissue environment | 65–140 kPa (interpret) |
|
rat dorsal subcutaneous implantation | 95–180 μm after 4 weeks (interpret) | N/A | Histological analysis showed the absence of CD68 cells in the rat model after the 4-week implantation period. The adhesive formed a fibrous capsule without significant granulation tissue formation, necrosis, or allergic reactions. | [136] |
| Gelatin or Chitosan | PAAc Grafted with NHS ester (PAAc-NHS ester) |
|
tissue adhesive and sealant, as well as for adhering wearable and implantable medical devices to wet tissues. |
|
|
rat dorsal subcutaneous implantation | 40 μm after 2 weeks; 240 μm after 4 weeks (interpret) | N/A | Histological analysis showed that the DST induced a host response similar to that of sutured patches, with a mild to moderate chronic inflammatory response. There was no significant necrosis, allergic reactions, or damage to surrounding tissues after implantation | [137] |
| PAAc with Chitosan | Multilayer system:
|
|
tissue sealant |
|
|
rat dorsal subcutaneous implantation | 145 ± 29 μm after 2 weeks; 135 ± 7 μm after 4 weeks | N/A | Histological analysis showed the patch with the zwitterionic layer induced significantly lower fibrous capsule formation and inflammation compared to patches with hydrophobic or hydrophilic layers. After 4 weeks, the fibrous capsule thickness around the zwitterionic layer was significantly thinner, indicating a reduced inflammatory response. The degree of inflammation at the implantation site for the zwitterionic layer-containing fall within the “very mild” to “mild” inflammation range | [115] |
| Poly (Acrylamide) (PAM) or PAAc, |
|
|
hemostatic sealant and bioelectronic interfacing Using the High-throughput analysis, the best adhesive is PAM-PDA-C2 | elastic modulus: ~ 2 kPa | wet tensile adhesion strength to porcine skin ~13 kPa after 7 days, ~12 kPa after 14 days, ~11 after 28 days | rat dorsal subcutaneous implantation | 0 μm after 2 weeks (interpret) | macrophage culture indicates significantly lower levels of pro-inflammatory (M1) markers: IL-6, TNF-α, and iNOS, and significantly higher levels of anti-inflammatory (M2) markers: CD206, IL-10, TGF-β1 | Immunofluorescent showed that the CD3 and
α-SMA intensity were significantly lower in the PDA–PAM
and PDA–PAM-C2 hydrogel groups compared to the PAM hydrogel
group. The PAM-PDA-C2 hydrogels not only promote cell adhesion and proliferation through up-regulation of cell adhesion-related genes but also exhibit immunomodulatory effects by reducing pro-inflammatory responses and promoting antiinflammatory markers. |
[138] |
| Poly(ethylene glycol) (PEG) |
|
|
post-surgical cardiac adhesion |
|
N/A | rat cardiac adhesion model porcine cardiac adhesion model | 15–50 μm after 4 weeks in rat (interpret) | N/A | Histological analysis at 4 weeks post-implantation showed thin encapsulation around some remaining gel, with no indication of chronic inflammation. There was minimal macrophage infiltration, no lymphocyte or neutrophil infiltration, suggesting the gel did not cause chronic inflammation and allowed normal wound healing. | [139] |
| Alginate-PAM and chitosan |
|
|
tissue adhesives, wound dressings, and tissue repair | ~1 MPa [140] | adhesion energy >1000Jm-2 on porcine skin | attached to rat myocardium | ~130 μm after 2 weeks (interpret) | N/A | biocompatibility tests in rats showed a lower degree of inflammation compared to cyanoacrylate (CA) and similar to COSEAL. Fall within “very mild” to “mild” degree of inflammation | [141] |
| PEG |
|
|
surgical sealant applications |
|
burst pressure tests on collagen only: ~200 mmHg max pressure | rabbit carotid artery model, and subcutaneous implantation | ~60 μm after 7 days (interpret) | N/A | Histological evaluation showed mild to moderate inflammation and granulation tissue at the treatment sites in both the carotid artery and subcutaneous implant models at 7 days, and mild at 32 days with no visible implant material, indicating good biocompatibility and healing. | [142] |
| Three-armed dopamine-modified hyaluronic acid (HA) |
|
|
cartilage repair |
|
N/A[143] | rat dorsal subcutaneous implantation | 0 μm at weeks 1,2, or 4 (interpret) | N/A | Mild inflammatory reactions were observed initially, but these subsided with no significant long-term inflammation. The material was biocompatible and showed good degradation performance in vivo. | [143] |
| Cartilage decellularized extracellular matrix (CdECM) and hyaluronic acid methacrylate (HAMA) | photoinitiator ruthenium (Ru)/sodium persulfate (SPS) | Ru/SPS photoinitiator system facilitates tyrosine cross-links between the tissue and CdECM | cartilage repair | compressive modulus (10–15 % strain) 126.50 ± 13.86 |
|
rat dorsal subcutaneous implantation | 120 μm after 2 weeks, 30 μm after 4 weeks (interpret) | N/A | Histological analysis revealed mild inflammatory responses initially (2 weeks), which subsided over time without significant long-term inflammation (4 weeks). The hybrid hydrogel reduced the inflammation induced by HAMA, making it a promising material for cartilage regeneration | [144] |
| Gelatin methacryloyl (GelMA) microneedle array | silicate nanoplatelets (SNs) to enhance hemostatic properties |
|
hemorrhage control | 100–150 kPa | 43.9 ± 6.9 kPa burst pressure | rat dorsal subcutaneous implantation | 110–150 μm after 2 weeks | N/A | Immunofluorescent showed that implanted MNAs were encapsulated in fibrous tissue infiltrated with mild inflammatory cells (CD3+ and CD68+) at week 1, transforming into loose connective tissue with reduced inflammatory cell ((CD3+ and CD68+) by week 4. | [145] |
| Swellable mussel adhesive protein (MAP)-based shell and a non-swellable silk fibroin (SF)-based core |
|
|
wound closure | fracture form of the patch is around 2.2 N per needle |
|
rat H&E-stained ileum defects | wound healing model, no fibrous tissue formation | N/A | Histological analysis showed mild inflammatory
responses initially, which subsided over time, with minimal long-term
inflammation. The patch-treated wounds exhibited better healing with reduced scarring and inflammation compared to non-treated or suture-treated wounds. The MAP/HA SL-MN patch-treated intestinal wounds showed successful serosa regeneration and mild adhesion to surrounding tissues, indicating good biocompatibility and reduced risk of post-surgical complication |
[146] |
| hyperbranched poly(amino-ester) (HPAE) and gelatin |
|
Fe3+ assistant catechol tissue adhesion | cardiac patch application | Storage modulus is 34.7 kPa |
|
rat myocardial infarction model | no capsule, fibrous tissue thickness ~80 im | N/A | Histological analysis showed strong adhesion
and seamless integration of the hydrogel with cardiac tissue. The hydrogel degraded mostly within 4 weeks, with minimal long-term inflammation. Immunofluorescence staining indicated initial neutrophil (CD11c) infiltration AT 14 days, which subsided over time (28 days), demonstrating good biocompatibility. |
[147] |
| PEGylated poly (glycerol sebacate)-based Janus adhesive patch (PEGS-based JAP) |
|
|
internal tissue adhesion and function reconstruction in wet and dynamic physiological environment |
|
|
|
80–160 μm from weeks 1–4 | N/A | Local inflammation responses were assessed
through histological analyses, including Hematoxylin-Eosin (H&E) and
Masson’s trichrome staining, as well as tumor necrosis
factor-α (TNF-α) staining. In the rat gastric perforation model, the JAP-sealed stomach almost completely healed with a clear and well-ordered contour, indicating minimal inflammatory response without detecting TNF-α staining after 21 days implantation. |
[148] |
| Four-armed PEG hydrogel |
|
|
tissue adhesive, wound healing |
|
|
rat subcutaneous implantation | 305 μm after week 4; 337 μm after week 8 (without Laponite, 253 μm followed by 498 μm) | N/A | Histological analysis at four weeks
post-implantation showed the formation of a fibrous capsule and a
cellular infiltration layer rich in fibroblasts around the
hydrogels. Laponite-containing hydrogels had a thicker fibrous capsule and higher cellular density in the infiltration layer compared to Laponite-free hydrogels. At eight weeks, the fibrous capsule thickness around Laponite-free hydrogels nearly doubled, while that around Laponite-containing hydrogels remained the same. No macrophages were observed at eight weeks, and the hydrogels were completely infiltrated with fibroblasts, indicating no prolonged inflammatory response. |
[149] |
| Methacrylate Gelatin (GelMA) and Hyaluronic Acid (HA) |
|
|
wound sealing |
|
|
rat subcutaneous implantation | ~160 μm after week 1; ~120 μm after week 2 | N/A | Histology analysis showed moderate inflammation after week 1; mild inflammation after week 2, with good integration and biocompatibility observed in tissue sample | [150] |
| HA |
|
|
tissue defect repair | storage modulus 160 Pa | lap-shear peeling test on hog casing 42 kPa | rat dorsal subcutaneous implantation | 50–70 μm after 1 week | N/A | [151] | |
| Gelatin | gelatin methacryloyl (GelMA) |
|
lung tissue sealant |
|
|
rat lung leakage model Porcine lung leakage model | 50 μm after 3 days; 110 μm after 7 days; 70 μm after 28 days | N/A |
Rat Model: Histological analysis
showed no significant inflammatory response. Initial macrophage
infiltration (CD68 positive) around the implants was observed at days 3
and 7 but disappeared by day 28. No lymphocyte infiltration was detected
at any time point. Porcine Model: Histological evaluation revealed strong collagenous defect repair tissue with no signs of lymphocyte infiltration and only mild macrophage accumulation after 14 days. |
[152] |
| poly(acrylic acid) and poly (vinyl alcohol) | PAAc functionalized with NHS |
|
anti-fibrotic interface for long-term implant | 62.7 kPa | 400 Jm−2 | rat model for intraperitoneal and intrathoracic implantation over 84 days humanized mouse model to assess fibrosis and immune response over 28 days porcine model on abdominal wall and small intestine over 7 days | no capsule, collagen layer thickness ~15–35 μm from days 3–84 (abdominal wall) |
|
Histological Analysis showed minimal
inflammatory response and no observable fibrous capsule formation at the
implant-tissue interface for the adhesive hydrogel over 84
days. Immunofluorescence Sstaining quantified the presence of fibroblasts (αSMA), neutrophils (neutrophil elastase), macrophages (CD68, vimentin, CD206, iNOS), and T cells (CD3) at the implant-tissue interface, showing significantly fewer inflammatory cells for adhesive hydrogels compared to non-adhesive implants. |
[77] |
Fig. 4.
Immunomodulatory hydrogel bioadhesives for immunosuppressive purposes. (A) Prevention of fibrotic capsule formation using a swelling bioadhesive patch: Physical swelling of the bioadhesive patch, combined with chemical bonding, provides strong mechanical integration between bioadhesive and tissue, and offers a solution to prevent fibrotic capsule formation. Reproduced from Ref. [77] with permission from Nature (Nature Portfolio). (B) Janus adhesive patch with dual-functional surfaces for tendon repair: The adhesive surface enables a strong attachment to tendon tissue, while the opposite gliding surface is engineered to be non-adhesive, to minimize friction with surfacing tissue which supports smooth tendon movement and reduce the risk of scar formation. The JTA also serves as a depot for sustained release of steroid drug, to reduce inflammation and support the healing process. Reproduced from Ref. [129] with permission from Nature Biomedical Engineering (Nature Portfolio). (C) Versatile tissue-sealing patch for wet adhesion: A multilayer tissue-sealing patch designed for minimally invasive surgeries features a microtextured bioadhesive layer, a dynamic hydrophobic fluid layer, and an antifouling zwitterionic nonadhesive layer. Reproduced from Ref. [115] with permission from Advanced Material (Wiley-VCH).
Drug-loaded immunomodulatory bioadhesives show promise in mitigating immune challenges during organ transplantation. A notable example involves “cell-particle hybrids”, which integrate pancreatic islets with a polydopamine-based bioadhesive coating loaded with the immunosuppressive drug FK506. These hybrids enable localized immune suppression at the transplant site, circumventing the need for systemic immunosuppression—a strategy often limited by adverse effects and suboptimal efficacy in clinical pancreatic islet transplantation for type I diabetes [37]. By gradually releasing low doses of FK506, the coated microspheres maintain islet viability and function while prolonging xenogeneic graft survival, demonstrating a targeted therapeutic approach [37].
Beyond transplantation, bioadhesive mechanisms can be engineered for controlled drug delivery to enhance tissue regeneration. For instance, researchers have leveraged mussel-inspired adhesion strategies, combining ion coordination and molecular crosslinking to co-deliver immunomodulatory metal ions (e.g., Zn2+) and osteoinductive growth factors like Bone Morphogenetic Protein-2 (BMP-2) [133]. In vivo studies reveal that Zn2+ alone boosts M2 macrophage recruitment by 92.5 % and elevates expression of the anti-inflammatory cytokine IL-10 by 84.5 %. When combined with BMP-2, these effects are amplified, achieving a 124.7 % increase in M2 macrophage recruitment and a 171 % upregulation of IL-10. This synergistic approach enhances bone-implant integration, yielding a mechanical fixation force of 203.3 N and new bone ingrowth of 82.1 %, far surpassing the performance of unmodified TiO2 implants (98.6 N and 45.1 %, respectively) [133].
In addition to drug delivery, bioadheisves have enabled the design of bilayered systems to address immunosuppression and other clinical needs [115,129,134,135]. Researchers have engineered bilayered bioadhesives with distinct properties on either side. For instance, Janus bioadhesive hydrogels have been developed for effective myocardial infarction repair, while preventing unwanted tissue adhesion [134]. These hydrogels feature asymmetric adhesive properties, with one side designed for strong adhesion to cardiac tissue and the other to prevent tissue adhesion. The adhesive side incorporates bioadhesive agents such as acrylic acid, polyethylenimine, aldehyde cellulose, 3-sulfonic acid propyl methyl acrylic acid potassium, and caffeic acid (containing catechol group), which enhance adhesion, mechanical strength, electrical conductivity, and ROS scavenging, respectively. This side also promotes the polarization of macrophages from the M1 phenotype to the M2 phenotype. Conversely, the anti-adhesive side consists of acrylic acid, carboxylated cellulose, and polyethylene glycol diacrylate, creating a highly hydrophilic surface that exhibits anti-cell adhesion and non-fouling properties, effectively preventing tissue adhesion and postoperative synechia [134].
Moreover, Janus adhesive patches can also be used to promote tissue regeneration. For example, Janus patches has been developed to promote tendon healing and regeneration. One such patch consists of a multifunctional electrospun hydrogel patch as the inner layer and a poly-L-lactic acid fibrous membrane as the outer layer. The inner layer contains tannic acid for bioadhesion, a polyphenol capable of forming strong hydrogen bonds with biological surfaces to promote adhesion, as well as zinc oxide (ZnO) nanoparticles for anti-inflammatory and anti-oxidation properties. The outer layer serves as a physical barrier to prevent the invasion of extrinsic cells and tissue [135]. Another example of Janus adhesive patches is designed with two functional surfaces: an adhesive surface and a gliding surface (Fig. 4B). The adhesive surface is coated with chitosan, which facilitates the strong adhesion to tendon tissue, allowing the patch to stay in place even under mechanical stress or in the presence of blood as shown in Fig. 4B. The gliding surface is engineered to be non-adhesive, supporting tendon gliding and reducing friction between the tendon and surrounding tissue, helping prevent fibrosis and scar formation [129]. The drug delivery approach is also employed with the Janus tough adhesive to provide localized and sustained steroid release to control the inflammation for tendon healing. The sustained release of the drug helps modulate inflammation by reducing the levels of inflammatory marks such as serum GRO-α shortly after implantation, while also promoting M2 macrophage polarization. The local delivery of steroid through the adhesive patch minimizes systemic side effects and provides prolonged anti-inflammatory effects directly at the site of tendon injury.
Building upon the principles established in bilayered designs, multilayered adhesive patches have been developed to further enhance functionality and performance in tissue regeneration and healing. A multilayered tissue-sealing patch has recently been developed for minimally invasive surgeries. Its multilayer composition includes a microtextured bioadhesive layer, a dynamic hydrophobic fluid layer, and an antifouling zwitterionic nonadhesive layer (Fig. 4C) [115]. The bioadhesive layer employs a dry-crosslinking mechanism through interfacial water absorption and NHS-amine reactions. Silicon oil is used as a hydrophobic fluid agent due to its chemical stability and favorable wettability with the bioadhesive material. The antifouling layer is achieved by interpenetrating zwitterionic poly(sulfobetaine methacrylate) into a thin film of hydrophilic polyurethane. This multifunctional origami patch demonstrates strong immunomodulatory performance by significantly reducing bacterial adhesion, fibrinogen adsorption, and fibrous capsule formation. In vivo studies show that the patch maintained a thin fibrous capsule (135–145 μm) after four weeks, compared to much thicker capsules around other materials, indicating a reduced inflammatory and fibrotic response and potential for minimizing immune reactions in surgical applications.
6. Designs and applications for wound healing
For applications of wound healing, immunomodulatory bioadhesives serve both mechanical and biological functions concurrently. They can not only merge wound edges to promote closure and prevent infection, but also interact with immune cells and other cell types at the wound bed for modulating the healing process. This dual functionality represents a promising avenue for clinical applications, particularly in managing chronic wounds characterized by persistent inflammation, impaired polarization of the M1 to M2 macrophage phenotype, and elevated levels of inflammatory cytokines and matrix metal-loproteinases. Addressing this dysregulation through immunomodulation of bioadhesives is critical, particularly for those patients with severe infections or compromised immune function.
A straightforward strategy for modulating immune responses involves incorporating anti-inflammatory functional groups into bioadhesives. Notable examples include catechol-derived compounds such as PDA, tannic acid, epigallocatechin gallate, and gallic acid, which not only enhance bioadhesion but also influence local immune responses through their intrinsic reactive ROS scavenging capabilities. This effect has been demonstrated in immunomodulatory bioadhesives composed of PAM or PAA, along with hydrophobic alkyl monomers of varying chain lengths and catechol-derived components (Fig. 5A) [138]. Among 36 tested formulations, the PDA-AM-ethyl acrylate formulation demonstrated the highest adhesion strength of 23.5 ± 1.9 kPa, and also polarized macrophages from the M1 to M2 phenotype, indicated by decreased M1 markers (IL-6, TNF-α, iNOS) and increased M2 markers (CD206, IL-10, and TGF-β1). Additionally, this bioadhesive reduced T-cell (CD3) and fibroblast (α-SMA) activation, which are commonly associated with chronic inflammation. These properties highlight its potential for hemostatic and wound healing applications. Moreover, bioadhesives can be engineered with multiple functional groups. For instance, the combination of the aldehyde and catechol functional groups have been explored for regulating the immune response [153]. Protocatechuic aldehyde bioadhesive achieved the M1 to M2 conversion without additional cytokines or external interventions, as the catechol and the aldehyde functional groups provide ROS scavenging function [153].
Fig. 5.
Immunomodulatory hydrogel bioadhesives for wound management. (A) Development and screening of immunomodulatory wet adhesives: A series of immunomodulatory wet adhesives were developed by incorporating hydrophobic alkyl monomers of varying chain lengths and catechol-derived components into hydrophilic hydrogel networks using a high-throughput method. After implantation in a mouse model, the best-performed hydrogel demonstrated reduced activation of T-cells (CD3 marker) and fibroblasts (α-SMA marker). Reproduced from Ref. [138] with permission from ASC Nano (American Chemical Society). (B) Thermosensitive bioadhesive matrix embedded with living algae and beneficial bacterial for wound care: The algae mitigating hypoxia while the bacteria deliver antimicrobial properties. This dual-action system reduces infection risk and transit the wound environment from pro-inflammatory to a pro-healing state. Hydrogel represents thermosensitive Pluronic F-127 and adhesive polydopamine, LAH represents living algae hydrogel, LBH represents living bacterial hydrogel, LMH represents living microecological hydrogel. Reproduced from Ref. [157] with permission from Science Advance (Science). (C) Immunomodulatory effects of miR-223*-laden nanoparticles in bioadhesives: miRNA enhance macrophage polarization from the M1 to M2 phenotype and promote the wound healing process. Reproduced from Ref. [159] with permission from Small Journal (Wiley-VCH).
Researchers have also incorporated naturally derived biomaterials or ECM into bioadhesives to regulate the pro-inflammatory or anti-inflammatory phenotype. For instance, incorporating decorin proteoglycan into a surgical glue can reduce the infiltration of inflammatory cells and lower the expression of inflammatory cytokines, while also downregulating the expression of profibrotic factors (e.g., TGF-β1) and upregulating antifibrotic factors (e.g., TGF-β3, Smad7), which indicates reduced fibrosis and scar formation [154]. In addition, proteoglycan can also regulate collagen fibril growth and reassembly, resulting in a uniform collagen architecture with properly packed fibrils. This surgical glue was shown to not only accelerate initial wound healing by promoting re-epithelialization, neovascularization, and rapid collagen synthesis but also enhance overall tissue integrity [154].
In recent developments, living materials have emerged as promising bioadhesive options for immunomodulation and tissue repair. One example is Platelet-Rich Plasma (PRP) bioadhesive, derived from the patient’s own blood. It is enriched with various growth factors, such as PDGF, VEGF, and TGF-β. These factors play a crucial role in stimulating cell proliferation, angiogenesis, and collagen synthesis, creating an optimal environment for wound healing. The adhesive nature of PRP allows it to form a stable interface with the wound bed, ensuring that growth factors remain localized at the site of injury, which accelerates the healing process and reduces recovery time. The PRP has been combine with alginate [155], chitosan and chondrotinsulfate [156] to create a double network hydrogel as an immunomodulating wound dressing. Another novel approach is the use of a living microecological hydrogel as a bioadhesive wound dressing. Living algae or beneficial bacteria have been incorporated within a thermosensitive bioadhesive matrix for wound care. The algae continuously deliver oxygen to the wound site, alleviating hypoxia, while the bacteria provide antimicrobial activity to combat infections (Fig. 5B). By combining these two living materials, this approach not only reduces the risk of infection but also modulates the immune response, promoting a shift from a pro-inflammatory state to a pro-healing state [157].
Beyond the biomaterial composition, advancements in immunomodulatory bioadhesives have focused on enabling the localized delivery of anti-inflammatory or pro-wound healing agents to promote wound healing [26,158–160]. The release of bioactive agents could be coupled with the degradation of the adhesive matrix for extended release. It is exemplified with a lipoic acid-based glue, which can gradually degrade to release lipoic acid, a B vitamin that can regulate the overexpressed pro-inflammatory cytokines and pro-fibrotic factor (e. g. TGF-β1) to a normal level [161].
Furthermore, bioadhesives can be functionalized with nanoparticles, enhancing their immunomodulatory effects. Nanoparticle-loaded bioadhesive hydrogels show promise in managing immune responses in chronic wounds, which are often characterized by an acidic pH and elevated ROS levels—both of which effectively trigger for targeted drug release [160]. Direct embedding the metal-based nanoparticles, such as ZnO in hydrogel bioadhesives have shown strong evidence of regulating immune responses through anti-inflammatory effects and by promoting macrophage polarization to the M2 phenotype on the wound site [131,132,135]. Additionally, graphene oxide-loaded bioadhesive hydrogel has demonstrated the ability to modulate macrophage behavior and promote diabetic wound healing by stimulating macrophage polarization into the M2 and even M2c phenotypes, promoting the release of angiogenic factors such as PDGF and VEGF through the paracrine pathway during the early stages of wound healing [162].
Incorporating potent biomolecules into nanoparticle-laden bioadhesives further enhances their immunomodulatory capabilities. When combined with nanoparticles carrying specific anti-inflammatory microRNAs, these bioadhesives can directly instruct cells to produce proteins that facilitate macrophage polarization from the M1 to M2 phenotype (Fig. 5C). This strategy demonstrates therapeutic potential both in vitro and in vivo by enhancing the expression of anti-inflammatory proteins, reducing pro-inflammatory markers, and promoting uniform, vascularized skin formation at the wound site (Fig. 5C) [159].
7. Designs and applications for cancer treatment
Bioadhesive technologies could benefit cancer treatments by addressing the challenges imposed by immunosuppressive tumor microenvironment (TME), which often impairs the capacity of immune system to recognize and attack cancer cells. Bioadhesives can act as localized drug depots, delivering and sustaining therapeutic agents at tumor sites, thus improving treatment efficacy and reducing systemic side effects. For instance, bioadhesive-based chronic transdermal delivery systems enable the continuous release of therapeutic agents directly to the tumor site in melanoma treatment. In addition to reducing the frequency of administration and improving patient compliance, this targeted approach enhances the concentration of drugs at the site of action while minimizing adverse effects associated with systemic therapies [34,127,163–165].
Since cancer treatment often requires a combination of therapeutics, bioadhesives designed for tumor therapy are frequently loaded with biomacromolecules such as tumor antigens, growth inhibitors, and antibodies. A key target in cancer immunotherapy is the Programmed Death-Ligand 1 (PD-L1), a protein overexpressed on tumor cells that helps tumors evade immune attack by binding to the PD-1 receptor on T cells, thus inhibiting T cell activity [126]. Delivering anti-PD-1 using bioadhesives can enhance local immune checkpoint blockade therapy. For instance, a bioadhesive composed of mussel adhesive protein conjugated with anti-PD-1 through a matrix metalloproteinase 2 (MMP2)-cleavable peptide demonstrated enhanced cancer treatment efficacy by controlling drug release within the tumor microenvironment. This bioresponsive system significantly reduced tumor growth and improved survival in a mouse melanoma model compared to free antibody treatment [127].
Besides drug delivery, bioadhesives can also be engineered for stimuli responsiveness and other modulation effects at the tumor site. For instance, an in situ sprayed bioresponsive immunotherapeutic fibrin gel was developed, incorporating pH-responsive calcium carbonate (CaCO3) nanoparticles loaded with anti-CD47 antibodies, designed for post-surgical cancer treatment (Fig. 6A). This dual-function gel not only promotes wound healing by forming a protective barrier at the surgical site, sealing the wound and serving as a retentive carrier for the drug, but also modulates TME for cancer immunotherapy. The CaCO3 nanoparticles act as proton scavengers, neutralizing the acidic environment commonly found in tumors, which typically favors the M2 macrophage phenotype that supports tumor growth. By reducing acidity, the CaCO3 nanoparticles create conditions more favorable for M1 polarization. The anti-CD47 antibodies loaded into the CaCO3 nanoparticles are gradually released into the tumor site, as the nanoparticles dissolve in the acidic environment. CD47 is known as an immune checkpoint by preventing macrophages from phagocytosing cancer cells. By blocking this signal with anti-CD47 antibodies, macrophages are activated to phagocytose cancer cells. Using the fibrin glue as a carrier, 79 % of the anti-CD47 antibodies released from the CaCO3-loaded fibrin gel remained detectable at the tumor site. In contrast, the antibody injected without fibrin glue can only be detected for 20 % at the same site. This combined modulation of the pH and the CD47-SIRPα interaction at tumor sites shifted the macrophage population from the tumor-promoting M2 phenotype to the tumor-fighting M1 phenotype, therefore revising the immunosuppressive environment [128].
Fig. 6.
Immunomodulatory hydrogel bioadhesives for cancer treatments. (A) Post-surgical cancer treatment and tumor size reduction: An innovative in situ sprayed bioresponsive immunotherapeutic fibrin gel was created, incorporating calcium carbonate (CaCO3) nanoparticles loaded with anti-CD47 antibodies. This functionalized fibrin glue demonstrated a reduction in tumor size over the same time period and increased the survival rate of mice. Reproduced from Ref. [128] with permission from Nature Nanotechnology (Nature portfolio). (B) Localized delivery of immunomodulatory drug: An injectable bioadhesive hydrogel was developed, incorporating dopamine-hybridized carbon quantum dot-supported Pd single-atom nanozymes (DA-CQD@Pd SAN). This hydrogel effectively induces adaptive immune responses, preventing tumor recurrence and metastasis. Reproduced from Ref. [163] with permission from Biomaterial (Elsevier). (C) Controlled immune cell recruitment and modulation: A macroporous gelatin-polyacrylic-based bioadhesive with NHS functional groups was designed to adhere strongly to tissue. The macropores in the hydrogel are optimized to house immune cells and enable the controlled release of tumor-specific chemokines, recruiting and modulating immune cells at the site. Reproduced from Ref. [34] with permission from Cellular and Molecular Bioengineering (Springer).
In addition to drug release, bioadhesives could offer as an in vivo “factory” to have catalytic reaction to provide bioactive agents to modulate the TME. A bioadhesive injectable hydrogel incorporated catechol groups and their oxidized products (quinone and semiquinone), often used for bioadhesion, but also serves as a catechol-based catalytic system, in conjugation with Pd single atom to catalyze the degradation of peroxide/persulfide to generate ROS (Fig. 6B). This function caused immunogenic tumor cell death and generated tumor-associated antigens in tumor lysates, thus triggering anti-tumor immune responses. In addition, the bioadhesive matrix, acting as a local niche, released CpG oligodeoxynucleotides, an adjuvant that promotes APC function, to enhance anti-tumor immune responses, while preventing systemic exposure. When combined with the checkpoint inhibitor anti-PD-L1, the hydrogel effectively induced adaptive anti-tumor immune responses, preventing tumor recurrence and metastasis [163].
Furthermore, bioadhesives could act as a scaffold for in situ recruitment and modulation of immune cells. Bioadhesives can deliver tumor-associated antigens to mature DCs, stimulating T cells. For example, an MnO2-loaded injectable adhesive hydrogel based on catechol groups was developed to act as an “antigen reservoir” after photothermal therapy, recruiting more immature DCs to the tumor site and enhancing the anti-tumor immune response. The immature DCs were recruited to the tumor site, and then transported more antigens to the nearby lymph nodes to stimulate the production of effector T cells and memory T cells, thus further enhancing the anti-tumor immune response [166]. In another example, researchers engineered macropores within a bioadhesive, which are large enough to house immune cells. The bio-adhesive carried NHS functional groups for bioadhesion, and release tumor-specific chemokines to recruit and modulate immune cells at the site. By incorporating GM-CSF and tumor extracellular vesicles (EVs) into the macroporous hydrogel, a significant number of DCs can be recruited in situ to process and present EV-encased antigens. These tumor antigen-presenting DCs can then traffic to lymphatic tissue to prime antigen-specific CD8+ T cells (Fig. 6C) [34].
8. Conclusions and future directions
In this review, we discuss the complex interplay between bioadhesives and the immune systems. Recent advancements have significantly improved the mechanical performance of bioadhesives, and growing evidence demonstrates the synergies between bioadhesives and immunomodulatory strategies. The resulting immunomodulatory adhesives offer a unique advantage by modulating local cellular and cytokine responses, effectively promoting tissue healing and enabling novel therapeutic approaches. These innovations hold tremendous potential in wound management, tissue engineering, and cancer therapy.
The future of bioadhesives, particularly those with integrated immunomodulatory functions, holds immense promise for advancing both medical and engineering applications. As research continues to uncover the interactions between biomaterials and the immune system, several critical areas of development are emerging that could shape the future of this field:
Smart bioadhesives:
The integration of smart technologies into bioadhesives is an exciting frontier. Smart bioadhesives could incorporate stimuli-responsive mechanisms that dynamically adjust their properties based on environmental cues or external stimuli, such as light [167], ultrasound [168], pH [169], temperature [170,171], or specific cytokine levels [172]. This adaptability could facilitate real-time interaction with changing immune conditions at the wound or implant site, enabling more precise control of healing processes. For example, bioadhesives that release anti-inflammatory agents in response to externally applied compression or elevated inflammatory cues could prevent the formation of fibrous encapsulation around implants [173]. Recent advances in ROS-responsive nanomaterials, also demonstrate how bio-adhesive can be engineered to dynamically response to inflammatory cues [174,175]. Similarly, bioadhesives that release growth factors in response to tissue damage could accelerate regeneration, providing tailored therapeutic responses based on the evolving needs of the tissue environment. To better achieve the precise control, in vivo and in vitro imaging and biosensing technologies can be integrated into bioadhesives to monitor their immune activity and therapeutic efficacy in real time, for example, fluorescent enhancing bioadhesive allow visualization of adhesive degradation, immune cell recruitment and inflammatory resolution [174].
Long-term stability and biodegradability:
A major limitation in current bioadhesives is achieving both long-term stability and controlled biodegradation. While most in vivo studies evaluate the immune response to bioadhesives over a short period (typically four weeks to three months), few studies investigate their long-term performance. Prolonged implantation can result in material degradation or fibrotic encapsulation, reducing their effectiveness in extended use. Moreover, biocompatibility challenges persist, particularly with materials like PAAc and PAM, which can present issues with cell attachment and remodeling. Future research will likely focus on bioadhesives that maintain both their adhesive and immunomodulatory functions over extended periods while ensuring complete and non-toxic biodegradation. The development of materials with programmable degradation rates, triggered by specific biological signals, could address these challenges for enhancing safety and efficacy in long-term medical applications.
Precision immunomodulation:
One of the remaining challenges in the application of immunomodulatory bioadhesives is achieving precise control over immune responses. Most bioadhesives aim to balance inflammatory signals by polarizing macrophages from the pro-inflammatory M1 phenotype to the anti-inflammatory M2 phenotype. However, fine-tuning this balance to promote healing while preventing excessive fibrosis or chronic inflammation remains difficult. Future advancements could enable more targeted modulation of immune pathways, tailored to individual patient profiles or specific disease conditions. Advances in molecular engineering, such as the use of antigen-specific immunomodulators, could allow bioadhesives to activate or suppress immune pathways in a more controlled manner [49], offering significant advantages in cancer treatment by enhancing localized immune responses [176], or in transplant medicine by enabling immune suppression to avoid graft rejection [177].
Integration with emerging therapies and devices:
The integration of bioadhesives with other emerging therapies, such as gene therapy, and stem cell therapy, presents a promising avenue for future research [10]. This approach extends to the combination of bioadhesives with advanced biomedical devices like smart wound patches and surgical tools, such as multifunctional sutures [178,179]. These integrated systems could enhance both the controllability and overall functionality of the therapeutic platform, offering a more dynamic and responsive treatment solution. Bioadhesives could be specifically engineered to deliver therapeutic agents—such as cytokines, growth factors, or gene-editing tools—directly to the target site, thereby improving the precision and effectiveness of treatment. For example, in cancer therapy, bioadhesives could be designed to release immunostimulatory molecules into the tumor microenvironment, enhancing the efficacy of immunotherapies by fostering a more favorable immune landscape for tumor eradication. Similarly, in regenerative medicine, bioadhesives could facilitate the localized delivery of stem cells or gene-editing constructs, accelerating tissue repair and promoting more comprehensive regeneration. Notably, the immune response at the adhesive-tissue interface, particularly in the context of bioadhesive-assisted cell delivery for immunomodulation, is a critical area that demands further investigation. These integrated approaches hold the potential to revolutionize therapeutic strategies by offering targeted, sustained, and highly controlled interventions.
Clinical translation and future directions:
Despite the immense potential of immunomodulatory bioadhesives, several barriers hinder their transition from laboratory research to clinical practice. Unlike traditional bioadhesives, those incorporating immunomodulatory agents are subject to stricter regulatory scrutiny, which can delay clinical adoption. Key challenges include scalability of production, cost-effectiveness, and regulatory approval, all of which must be addressed to accelerate clinical translation.
A potential strategy to accelerate clinical translation, particularly in cancer treatment, is to incorporate clinically approved drugs into general bioadhesive formulations. Since the drug component has already undergone regulatory approval, this approach may streamline the approval process for the bioadhesive, reducing regulatory hurdles and facilitating faster adoption in clinical settings. Such a strategy allows for leveraging existing drug safety and efficacy data, improving the feasibility of integrating immunomodulatory bioadhesives into standard therapies.
One of the fundamental challenges is that immune cells are highly sensitive to their microenvironment, and the immune system exhibits an innate response to foreign materials, even those deemed biocompatible [180–182]. Achieving consistent immune modulation across different patient populations remains a significant challenge. Because immune cell behavior varies among individuals due to differences in medical history, underlying conditions, and genetic factors, further complicating bioadhesive performance in clinical settings [182,183]. This variability also emphasizes the integration of bioadhesives with immunomodulating strategies that could stimulate robust immune responses.
Future efforts should prioritize standardizing design and testing protocols for immunomodulatory bioadhesives, particularly regarding their immune-modulating effects. Establishing clear guidelines for preclinical and clinical evaluation will be essential to facilitate regulatory approval and ensure consistency in clinical applications. Additionally, the development of cost-effective and scalable manufacturing processes will be critical in making these advanced bioadhesives more widely accessible to patients.
In conclusion, the future of immunomodulatory bioadhesives lies in their ability to not only bond tissues but also to interact with the immune system in highly controlled and beneficial ways. By addressing current challenges and exploring emerging frontiers—such as smart materials, personalized medicine, and regenerative therapies—bioadhesives with immunomodulatory capabilities have the potential to revolutionize medical treatments, from wound healing to cancer therapy. Interdisciplinary collaboration among material scientists, immunologists, and clinicians will be crucial in unlocking the full clinical potential of these innovative materials.
Acknowledgements
This work was supported by the Canadian Institutes of Health Research (grant PJT 165995) and the National Institutes of Health (grant RO1-DC021461). J. Li acknowledge the support from the Canada Research Chairs program. X. Lan acknowledges the financial support from Natural Sciences and Engineering Research Council of Canada (NSERC). E. Johnston and T. Ning acknowledge the support from the Faculty of Engineering, McGill University. T. Ning also acknowledges the support from Fonds de recherche du Québec – Nature et technologies (FRQNT).
Declaration of competing interest
The authors declare the following financial interests/personal relationships which may be considered as potential competing interests: Jianyu Li reports financial support was provided by Canada Research Chairs Program. Xiaoyi Lan reports financial support was provided by Natural Sciences and Engineering Research Council of Canada. Evan Johnston reports financial support was provided by McGill University Faculty of Engineering. Tianqin Ning reports financial support was provided by McGill University Faculty of Engineering. Tianqin Ning reports financial support was provided by Quebec Research Fund Nature and Technology. Jianyu Li reports a relationship with Canadian Institutes of Health Research that includes: funding grants. If there are other authors, they declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.
Footnotes
CRediT authorship contribution statement
Xiaoyi Lan: Writing – review & editing, Writing – original draft, Conceptualization. Evan Johnston: Writing – review & editing, Writing – original draft. Tianqin Ning: Writing – review & editing, Validation. Guojun Chen: Writing – review & editing. Lisbet Haglund: Writing – review & editing, Supervision, Funding acquisition. Jianyu Li: Writing – review & editing, Supervision, Investigation, Funding acquisition, Conceptualization.
Data availability
No data was used for the research described in the article.
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Data Availability Statement
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