Synthetic Pressure Sensitive Adhesives for Biomedical Applications

Abstract

Pressure sensitive adhesives are components of everyday products found in homes, offices, industries, and hospitals. Serving the general purpose of fissure repair and object fixation, pressure sensitive adhesives indiscriminately bind surfaces, as long as contact pressure is administered at application. With that being said, the chemical and material properties of the adhesive formulation define the strength of a pressure sensitive adhesive to a particular surface. Given our increased understanding of the viscoelastic material requirements as well as the intermolecular interactions at the binding interface required for functional adhesives, pressure sensitive adhesives are now being explored for greater use. New polymer formulations impart functionality and degradability for both internal and external applications. This review highlights the structure-property relationships between polymer architecture and pressure sensitive adhesion, specifically for medicine. We discuss the rational, molecular-level design of synthetic polymers for durable, removable, and biocompatible adhesion to wet surfaces like tissue. Finally, we examine prevalent challenges in biomedical wound closure and the new, innovative strategies being employed to address them. We conclude by summarizing the progress of current research, identifying additional clinical opportunities, and discussing future prospects.

Graphical Abstract

1. Introduction

Pressure sensitive adhesives (PSAs) are ubiquitous in everyday life. From postage stamps to Duct^® Tape, the breadth of adhesive performance is extraordinarily vast. Remarkably, PSAs are durable enough to secure components of automotive interiors, removable to allow for the quick change of a baby’s diaper, and optically clear to permit visibility through a smartphone’s screen protector. PSAs for medical applications pose unique challenges over consumer goods including biocompatibility, biodegradability, breathability for topical applications, and adherence to wet surfaces like tissue. Even with these challenges, the medical adhesive market is expected to boom from 2021 to 2026 with projected growth from US$8.8 billion to 12.3 billion [1], largely supported by advancements in polymer chemistry and engineering that have yielded new materials with promising adhesive properties.

In general, biomedical adhesives are most frequently employed during wound closure or surgical repair procedures with the intent of restoring tissue structure and function. Traditional methods of intervention including staples, sutures, and adhesives provide sufficient mechanical strength for wound closure and the prevention of dehiscence. Despite their success at tissue approximation, sutures and staples have multiple shortcomings regarding surgical functionality and patient outcomes. Suture application is a time-consuming process that is not uniformly appropriate for all surgical disciplines, especially considering more challenging and highly vascularized organs such as the lung, liver, and brain. Additionally, suturing through the skin can pull bacteria from the skin surface into the wound resulting in a surgical site infection [2, 3]. Iatrogenic tissue trauma, fibrosis, and chronic pain add to the above concerns and emphasize the need for more effective surgical tools [4].

Adhesives, which coat the entire area of the wound, are less invasive, easier to use, and more effective at preventing fluid leakage and infection. By covering the wound surface, adhesives permit load transfer between fractured tissue, ultimately minimizing stress localization more effectively than other closure methods [5]. Multiple prospective studies demonstrate that surgical closure time is significantly shorter when performed with tissue adhesives instead of sutures [3, 6–8]. Adhesive application is less painful [8] while reducing the incidence of infection [9] and dehiscence [3, 10]. For the closure of topical lacerations, cosmetic appearance of the healed wound is a priority for many patients. Adhesives afford equivalent, if not improved, cosmesis over traditional closure methods [6, 7, 10–12].

Specifically, pressure sensitive adhesives (PSAs) offer many potential advantages over existing FDA-approved sealants/adhesives including ease of application and versatility in binding strength. With only light contact pressure, PSAs bind to a wide variety of surfaces such as metal, glass, and tissue. Currently, pressure sensitive adhesive use in the medical field is approved only for topical indications, most commonly as wound care dressings, medical and wearable device adhesives, and waterproof cover tapes [13–19]. As shown in Table 1, topical PSAs are typically composed of acrylic, silicone or polyurethane, and exhibit adhesion strength to stainless steel ranging from 4.6 to 17.2 N/25mm (TABLE 1) [20–24]. Similar data for adhesion to one common tissue is not available.

Table 1.

Examples of commercial pressure sensitive adhesives, indicated only for external use.

Pressure Sensitive Adhesive	Manufacturer	Approved Indication(s)	Adhesive Composition	Peel Adhesion to Stainless Steel
Medical Tape 1538	3M	wearable device fixation; wound dressing cover tape	acrylic	7.4 N/25 mm [#]
Tegaderm™	3M	wound dressing	polyurethane	N/A
MED 5078A	Avery Dennison	short-term surgery; wound dressing; ostomy flanges	acrylic	14.1 N/25 mm [#]
MED 5610SI	Avery Dennison	wearable device fixation; patient monitoring	silicone	N/A
Liveo™ MG-2404	DuPont	wearable device fixation	silicone	17.2 N/25 mm [#]
Liveo™ STD BIO-PSA	DuPont	transdermal drug delivery	silicone	17.2 N/25 mm [#]
Band-Aid®	Johnson & Johnson	wound dressing	acrylic	N/A
Mepilex® Border	MóInlycke	wound dressing	silicone	N/A
Cyabine™	ToyoChem	surgical tape; film dressing	polyurethane	4.6 N/25 mm [#]

PSAs are attractive for wound closure and, in our opinion, have significant potential for internal applications. There is a critical need for surgical adhesives that are biocompatible, biodegradable, and mechanically compliant with tissue. At a time when more effective surgical tools are in demand, pressure sensitive adhesives are uniquely positioned to fulfill the need for a more universal, rather than niche, solution to internal wound closure. While other reviews from the past five years discuss the breadth of adhesive and sealant compositions across applications [5, 25–28], we emphasize the importance of using specific clinical problems to drive the development of new pressure sensitive adhesives for biomedical applications. Herein, we discuss the rational design of PSAs for biomedical use from the perspective of a polymer chemist, biomaterials engineer, and practicing surgeon. We highlight the mechanisms of interfacial bonding with tissue, peel adhesion, and cohesion. We specifically focus on synthetic polymers for their significant potential as biomedical PSAs and describe recent examples. Finally, we discuss some of the most prevalent clinical challenges in the field and highlight potential innovative solutions employed to address them.

2. Limitations of Current Medical Adhesives

The development of adhesives or sealants for surgery is ongoing with significant successes noted in the reinforcement of traditional tissue closure mechanisms such as sutures [29–31]. By definition, an adhesive’s primary function is to bind materials together, be it a tissue/tissue or tissue/non-tissue junction. Adhesives such as Indermil^®, Histoacryl^®, and Histoacryl^® Blue demonstrate marked success at reapproximating deep dermal suture lines for topical wound closure [32]. In contrast, a sealant’s primary function is as a barrier to fluids where it performs as a glue. For example, sealants like Tisseel^™ and BioGlue^® improve hemostasis and shorten surgery time when used to prevent suture-hole bleeding from arterial anastomoses [33]. The Food and Drug Administration (FDA) and other global regulatory agencies have approved several adhesives and sealants of primarily fibrin, polyethylene glycol (PEG), or cyanoacrylate composition for internal and external clinical use [3]. However, many of these commercial tissue adhesives are restricted to external applications due to material strength and biocompatibility concerns (TABLE 2). Below, a few of the most prevalent limitations with existing biomedical adhesives are briefly discussed.

Table 2.

Examples of commercial liquid sealants/adhesives and their limitations.

Adhesive/Sealant	Approved Indication(s)	Adhesive Composition	Limitations
DuraSeal™	spinal surgery	tetra-NHS-derivatized PEG, trilysine	low cohesive strength
CoSeal™	vascular reconstruction	tetra-NHS-derivatized PEG, tetra-thiol-derivatized PEG	low cohesive strength
Adhearus®	dural repair	PEG, polyethyleneimine	low cohesive strength
Progel™	pulmonary air leaks	human serum albumin, PEG	low cohesive strength; biologically-derived components
Tisseel™	cardiopulmonary surgery; colostomy closure	fibrin	low cohesive strength; biologically-derived components
Artiss®	skin grafting; burn treatment	fibrin	biologically-derived components
BioGlue®	cardiopulmonary surgery	BSA, glutaraldehyde	biologically-derived components; toxic degradation products
Dermabond™	low tension tissue closure	octyl-2-cyanoacrylate	toxic degradation products; exothermic adhesive curing
Histoacryl®	low tension tissue closure; sclerotherapy	n-butyl-2-cyanoacrylate	exothermic adhesive curing

2.1. Biologically-derived sealants

Tisseel^™, a fibrin glue approved by the FDA for use in cardiopulmonary surgery and colostomy closure, is limited by its use of human-derived components which pose a viral transmission risk to the patient [28, 34–38]. However, the FDA ensures that biomaterials are robustly screened for viral agents before use and therefore this transmission risk to the patient is, in reality, quite low [39]. Additionally, sample processing such as pasteurization, pH treatment, and gamma irradiation further reduce the risk of viral transmission [40]. Of primary concern with fibrin, albumin, collagen, and other protein-based adhesives is the common occurrence of autoimmune reactions despite the controlled pre-processing and purification requirements. For instance, Artiss^® [41], Progel^™ [42, 43], and BioGlue^® [28, 44] all contain human- or animal-derived components and can elicit unintended inflammatory responses in patients. A comparison of hematoxylin and eosin (H&E) stained aortic tissue treated with either CoSeal^™ [45], a PEG sealant, or BioGlue^®, a bovine serum albumin (BSA) sealant, illustrated this biocompatibility concern in vivo. CoSeal^™ exhibits less inflammation in surrounding tissue than BioGlue^®, demonstrating the potential of synthetic adhesives to overcome the biocompatibility limitations of existing protein-based clinical options [46].

2.2. Toxic degradation products

Cyanoacrylate adhesives are advantageous as they are more strongly adherent than fibrin glues and are often waterproof – ideal for topical applications [39]. However, cyanoacrylates are only appropriate for external applications due to the production of toxic degradation products. For example, Dermabond^™, a 2-octyl-cyanoacrylate adhesive, has high bond strength and yet forms formaldehyde as it decomposes [25, 28, 47, 48]. Additional concerns with BioGlue^® stem from the toxic degradation product of glutaraldehyde, one of its primary components [33]. The release of glutaraldehyde from polymerized BioGlue^® can lead to a severe inflammatory response in patients and overall inhibition of tissue healing. Reports of such complications, such as the increase in incidence of bronchopleural fistula, have led to recommendations restricting its use in thoracic surgery to peripheral, rather than central, air leaks [49–52]. Promising alternatives to these products include PEG-based sealants which employ crosslinking via N-hydroxysuccinimide(NHS)-activated carboxylic acids and amine or thiol nucleophiles. PEG is biocompatible and widely used for internal wound closure and broader biomedical applications [48, 53–56].

2.3. Exothermic adhesive curing

Cyanoacrylate adhesives cure when in contact with water-containing surfaces like tissue via rapid crosslinking between isocyanate functional groups and organic tissue [57]. This anionic polymerization, in the presence of catalytic water, forms a covalent bond between cyanoacrylate pendant chains and the amine groups of tissue proteins, generating a strongly cohesive material [58–60]. Despite the desirable strength of the resulting adhesive, these exothermic-curing polymers release heat upon polymerization and therefore must be used with caution. For example, Histoacryl^®, a n-butyl-2-cyanoacrylate liquid adhesive, must be used extremely sparingly to avoid further damage to tissues through the heat produced during in situ polymerization [61]. This polymerization occurs quickly, and therefore physicians must be trained in order to apply the adhesive properly while minimizing patient discomfort [48]. Additionally, care must be taken to avoid use of cyanoacrylate adhesives subcutaneously or internally as they can elicit foreign body responses as well as generate tissue-damaging levels of heat [59]. In contrast, PEG-based sealants contain a large percentage of water in the formulation and generate minimal heat upon crosslinking.

2.4. Low cohesive strength

PEG-based sealants such as DuraSeal^™ are advantageous in the biocompatibility and tunability of their structure, however lack in cohesive strength (<10 J/m²), which can lead to adhesive failure [62, 63]. This is particularly problematic for tissue sealing applications which can continuously exceed 1,000 J/m² in toughness [64]. Progel^™, a human serum albumin/PEG hybrid sealant approved for treating pulmonary air leaks, is limited in functionality by low cohesive strength that fails to sufficiently maintain tissue closure following lung inflation. In one study, Progel^™ reinforced staple and suture lines used to close intraoperative air leaks (IOAL) following lung resection. Progel^™ successfully sealed 77% of IOALs, compared to 16% of IOALs in patients receiving sutures or staples alone [65]. While the sealant performed better than sutures alone, there is still room for improvement. Additionally, other PEG hydrogel sealants such as FocalSeal^® are limited in their practical application because of required photoactivation [39, 66]. Materials that require light-induced polymerization for adhesion are limited in the clinic by the need for an additional light source device and by possible spatial restrictions at the site of application during surgery.

It is important to note that, in their current state, many adhesives (especially PEG-based sealants) are contraindicated for high tension wounds due to their reduced tensile strength compared to sutures and staples [3, 12]. However, cyanoacrylate adhesives, such as Dermabond^™ are able to close high-tension wounds due to their higher matrix toughness [63, 67]. In a study comparing the use of Dermabond^™ to staples for the closure of high-tension wounds, Dermabond^™ prevents skin abscesses that result from mechanical closure with staples [68]. Additionally, high-tension wounds treated with either n-butyl-2-cyanoacrylate adhesive or silk sutures heal with similar long-term efficacy and wound breaking strength [9].

3. Mechanism of Pressure Sensitive Adhesion

The ability to spread and wet the substrate upon applied contact pressure is unique to pressure sensitive adhesion. A primarily physical, rather than chemical, mechanism of binding at the interface affords PSAs with substantial range in functionality and allows for adhesion independent of surface type. Surface roughness, surface energy, and polymer viscoelasticity are all critical parameters influencing the strength and durability of an adhesive bond. Adhesion increases with higher surface-to-adhesive contact area and higher substrate surface energy. In the medical field, adhesives must securely bind to rough and low surface energy materials where hair, wrinkles, folds, and pores all contribute to the non-homogeneity of the wet or moist surface. During application, the PSA functions as a viscous liquid to spread and wet the surface under an applied force (Fig. 1A). Upon removal of pressure, the PSA cures and behaves as an elastic solid, imparting cohesive strength that allows for natural movement of tissue underneath.

Fig. 1.

A. Schematic of pressure sensitive adhesion application, pressure-induced viscous spreading, and interfacial interactions at the substrate-adhesive interface; B. Surface energy measurements of various surfaces, including human skin; C. Methods of analysis for the critical viscoelastic properties of tack, peel strength, and shear resistance; D. Graphical depiction of the ideal viscoelastic window for general purpose PSAs.

3.1. Surface Roughness

Surface properties of the substrate highly influence the strength of adhesive bonds. Increased texture and surface irregularities (i.e., increased available surface area) yield more opportunities for binding interactions, resulting in a mechanical interlocking mechanism between the PSA and the adhered [58, 69]. The ability for the PSA to penetrate into surface deviations leads to stronger adhesive bonds, as illustrated by higher peel strengths for rough materials like stainless steel compared to smooth surfaces like glass [70]. Similarly, adhesion energy is directly related to a polymer’s ability to interdigitate with the substrate surface, with rough surfaces demonstrating greater surface area and higher propensities for strong bonding relationships [71].

3.2. Surface Energy

Surface energy, the energy change per unit of surface area, impacts the strength of initial bond formation between PSA and substrate [72]. High surface energy materials such as glass exhibit strong cohesive forces on a molecular level, yet also possess strong attractive forces towards dissimilar materials. Consequently, PSAs demonstrate higher tack on glass compared to lower surface energy materials like stainless steel [70]. Beharaj et al. describe the dependence of surface energy on adhesive binding through tack experiments on various surfaces using a polycarbonate PSA. The polymer, which has a similar peel strength to Duct^® Tape, demonstrates equivalent tack on glass, metal, and wood surfaces and significantly reduced tack on a polytetrafluoroethylene (PTFE) surface [73]. This observation is attributed to the low surface energy of PTFE compared to the other, higher energy surfaces. Consideration of surface energy will be critical for biomedical PSA design since skin is a low surface energy material with a critical surface energy below 40 mN/m [74–77]. (Fig. 1B)

3.3. Viscoelasticity

Tack, peel strength, and shear resistance are all critical measures of PSA performance that depend on the bulk viscoelastic properties of the polymer (Fig. 1C) [78]. Adhesion strength (T) directly relates to the work of adhesion between adhesive and substrate (W_a) and to the viscoelastic properties of the material, namely interfacial adhesion (B) and cohesive strength (D) [79–82].

$$T=W_a*B*D$$ (Equation 1)

Both adhesive and cohesive forces are therefore essential for functional PSAs. In order to successfully wet the surface and resist deformation, the material must simultaneously demonstrate liquid-like flow and solid-like strength [83, 84]. To illustrate the balance that is required between cohesive and adhesive forces, an ideal viscoelastic window for general purpose PSAs exists with both a loss modulus (G”) and storage modulus (G’) between 10⁴ and 10⁵ Pa (Fig. 1D) [84–87]. By changing the T_g or molecular weight of polymer components, the properties of a PSA can be optimized to fall within this viscoelastic window. For example, lowering the T_g decreases G’, while increasing the molecular weight affords greater cohesion and increases G”.

3.3.1. Tack

Tack, a measure of the rate of adhesive bond formation, highly depends on a polymer’s viscosity. For adhesion to tissue, PSAs typically possess a viscosity between 1-50 centipoise over a physiologically-relevant temperature range of 0-40°C [58]. Tack additionally depends on contact time, with longer periods of induced application pressure improving binding [88]. For polymers lacking inherent tack, adhesion improves by decreasing the molecular weight of the polymer or through the addition of low molecular weight plasticizers and tackifiers [58, 82]. Especially for biomedical applications, quick rates of binding are important in order to minimize fluid loss or leakage through a lesion.

3.3.2. Peel Strength

Once the initial contact is made between adhesive and substrate, intermolecular interactions at the interface strengthen the adhesive bond. The magnitude of this bonding relationship describes a material’s ability to resist removal upon peeling. Peel strength should be tuned according to the desired removability of the application. In medicine, peel strength is applicable to external applications where a force will eventually be applied with the intent of removing the adhesive and to internal applications where removal and repositioning of the adhesive may be required. Typical rates for patient removal of topical wound dressings range between 100-200 mm/min [75]. The lower the peel strength, the easier it will be to remove the PSA.

3.3.3. Shear Resistance

The ability of tissue (e.g., skin, lungs, intestines, etc.) to move naturally is essential for uninterrupted range of motion and organ functionality. To achieve physiologically-relevant movement at the adhesive-substrate interface, the elastic moduli of both the tissue and PSA should be similar so that they have analogous deformable properties [4, 5]. For example, in designing PSAs for adhesion to skin, the elastic modulus of both the PSA and skin should be approximately 0.1-0.3 MPa [85, 89]. Similarly, PSAs need sufficient cohesive strength to resist flow upon an applied shear force so that they may remain securely bound to the adhered. Increasing the T_g, via selection of monomer(s) or higher molecular weight, elevates the storage modulus and improves cohesivity [70]. Strongly cohesive PSAs are also advantageous for applications where debonding is required. Especially for topical bandages and tapes, high elastic shear materials which resist cohesive failure during removal are beneficial since they do not leave lingering adhesive residue on the skin [85].

3.4. Molecular Weight

Molecular weight is highly influential on the material properties of the polymer, namely regarding thermal transitions (T_g and T_m), viscosity, and toughness. These thermal transitions and viscosity will increase with molecular weight, with a higher number of chain entanglements and less free volume contributing to the changes in property [90]. Additionally, tack will increase with increasing molecular weight, as longer chains will favor substrate wetting and the subsequent dissipation of energy [91]. In general, intermediate molecular weights between 20-100 kg/mol are likely optimal due to the achievable balance between the desired properties of tack (high molecular weight) and viscous flow (low molecular weight).

3.5. Polymer Architecture

Whereas bulk properties of a polymer significantly impact adhesion, as discussed below in Section 4, polymer viscoelasticity can be fine-tuned through the use of additives such as plasticizers and tackifiers. These additives are frequently employed to break-up the adhesive’s dense polymer network and improve its wetting properties. In fact, traditional polymer networks typically require up to 50 wt% of additives to functional as optimal pressure sensitive adhesives [92]. However, due to their small size, additives have the potential to leach out of the adhesive, leaving undesirable residue on the substrate following removal. In an effort to combat this effect, efforts are underway to modulate the viscoelastic properties of bulk polymers without additives by leveraging polymer architecture. These additive-free pressure sensitive adhesives utilize bottle-brush and microneedle polymer geometries to dilute chain entanglements and improve surface contact with the adhered.

Comb and bottle-brush polymers are characterized by short, pendant polymer chains covalently bonded to a polymer backbone (Fig. 2A). The grafted architecture modulates the viscoelastic properties of the bulk adhesive through both bonding and debonding phases of use. During application, the bottlebrush ends reduce chain entanglements and viscosity. Additionally, they have more chain ends than the linear alternatives, increasing contact area with the surface and strengthening adhesion. During removal of the pressure sensitive adhesion, steric repulsion between the pendant graft polymers contributes to strain-stiffening and improved bulk elasticity [92]. Poly(isobutylene), poly(butyl acrylate), and poly(4-mehtylcaprolactone), in particular, have demonstrated success as bottle-brush pressure sensitive adhesives in recent years [92–94]. The grafting density and chain length of the bottle-brush polymer significantly impacts adhesion in both covalently and physically (i.e. hydrogen bonding) crosslinked networks [93, 95].

Fig. 2.

A. Cartoon of bottle-brush polymer architecture. B. Microneedle configuration with swellable hydrogel tip [97], Copyright 2013. Adapted with permission from Macmillan Publishers Limited. C. Ex vivo performance of swellable microneedle adhesive patch (60% DL-MN patch) compared to suture and non-microneedle (flat) patch at closure of luminal leak [98], Copyright 2019. Reproduced with permission from Elsevier Limited.

Desirable pressure sensitive adhesion can be also achieved by modifying the microtopography of the bulk adhesive. Specifically, microneedle architecture improves adhesion through increased contact area and surface penetration. Microneedle geometries, 25-1000 μm in height, are particularly effective at penetrating the stratum corneum of skin and are consequentially a useful strategy at modulating pressure sensitive adhesives for medical applications [96]. Yang et al. found inspiration from endoparasitic worms, which use proboscis to penetrate the tissue of target organisms [97]. Adhesion of a poly(styrene)-b-poly(acrylic acid) PSA patch improves with a microneedle configuration that allows for mechanical interlocking with the tissue surface (Fig. 2B). This proboscis-inspired mechanism utilizes a Velcro^®-like approach, and yet takes the binding interaction one step further by incorporating a tip that swells upon tissue penetration, thereby strengthening the PSA’s resistance to debonding. Similarly, Jeon et al. report a swellable adhesive microneedle with a silk fibroin solid core and mussel adhesive protein outer shell [98]. In an ex vivo porcine model of luminal leak closure, the microneedle patch performed with similar efficacy as sutures, the current standard of care, after 30 minutes of application times (Fig. 2C).

4. Synthetic Polymers for Pressure Sensitive Adhesion

Synthetic polymers are uniquely suited to fulfill the viscoelastic material and design requirements of biomedical pressure sensitive adhesives. Comprised of repeating units of one or more substructures, polymers can be tailored to specifically function as PSAs. Variation in backbone or pendant character, crosslinking, or molecular weight all intimately influence the thermal and material properties of a polymer. Polyacrylics, silicones, polyurethanes, polyesters, polysaccharides, and polystyrenes are some of the most common pressure sensitive adhesive formulations, and will be discussed here for their potential application in medicine. Figure 3 illustrates comparative viscoelastic properties for three of the most common pressure sensitive adhesive formulations (Fig. 3A). It is important to note that these ranges are generalizations and are heavily influenced by adhesive composition, surface character, temperature, and testing procedure.

Fig. 3.

A. General comparative viscoelastic properties of polyacrylic, silicone, and polyurethane pressure sensitive adhesives. B. General synthesis of polyacrylics, including examples common commercial pressure sensitive adhesives. C. General synthesis of silicone polymers and ideal structure of MQ resin [116], Copyright 2017. Adapted with permission from Multidisciplinary Digital Publishing Institute; D. Cartoon of intercalation of silicone/polyacrylic adhesive with substrate surface. Upon peeling, higher affinity for silicone strands to polyacrylate facilitate clean removal [121], Copyright 2019. Adapted with permission from Elsevier Science Ltd. E. Synthesis of polyurethane from diisocyanate and polyol and illustration phase separated polyurethane macromolecules with amorphous soft regions and crystalline regions generated from hydrogen bonding between isocyanate segments.

4.1. Polyacrylics

The use of polyacrylics as sealing and binding materials originates in the early twentieth century as liquid glues. Polymethyl methacrylate (PMMA) sealants mark one of the earliest uses of polyacrylic adhesives, patented for dental use in the 1930s [99, 100]. During World War II, rubber shortages necessitated the transition from natural rubber and pine gum bandages to synthetic polymer alternatives [74, 101]. Innovation boomed with new adhesives coming to market such as 2-methyl cyanoacrylate “super glue” in the early 1950s and cyanoacrylates as hemostatic bandages for field combat in the Vietnam War [102]. By the 1980s, acrylics dominated the PSA market due to their preferable tack, stability, and versatility [101] and today, growth continues with acrylics comprising approximately 50% of annual PSA production [103].

The transition of polyacrylics from consumer product to biomaterial was not without its challenges. Initial polymer structures with short alkyl pendant chains degrade into toxic byproducts such as formaldehyde, ultimately truncating their in vivo potential. Longer butyl- and octyl- acrylic polymers demonstrate reduced histocytotoxicity compared to shorter alkyl chain derivatives with increased malleability and rubber-like consistency, but still generate formaldehyde upon degradation [5, 104]. In 1964, the FDA reviewed a butyl-2-cyanoacrylate tissue adhesive formulation for in-human use [105]. However, approval was delayed decades due to in vivo experiments and clinical trials indicating potential adverse effects [104]. By the 1980s, Europe, Asia, and Canada had approval for cyanoacrylate tissue adhesives, but it was not until 1998 that similar approval was obtained in the United States for topical use [102, 104]. Key to this success was synthesis of highly pure monomers, for example by Closure Medical Corporation who developed Dermabond^®, an octocyanoacrylate topical adhesive widely-used in clinics and at home.

Beyond tissue adhesives and sealants, polyacrylics are particularly well-suited to perform in a clinical setting as pressure sensitive adhesives given their unique viscoelastic properties. Polyacrylics, composed of a hydrocarbon backbone with carbonyl-functionalized pendant chains (Fig. 3B), inherently possess low glass transition temperatures due to rotational flexibility within the backbone and therefore function as viscous liquids at room temperature to wet the substrate [84]. Additionally, polyacrylics are highly tunable. Typical polyacrylic PSA formulations consist of 70-90 mol% of low T_g polymers, such as poly(n-butyl acrylate) or poly(ethyl acrylate), designed to impart flexibility and tack while a smaller proportion is composed of high T_g polymers, such as poly(vinyl acetate) or poly(methyl methacrylate), which improve material stiffness and shear strength through enhanced cohesion properties [106, 107]. Today, compared to alternatives such as silicones and polyurethanes, acrylic polymers dominate PSA composition due to their ease of application, material stability, and low costs. Acrylic PSAs are clear, colorless, and durable against photo, thermal, and chemical stressors [26, 78, 107, 108]. These promising material traits are largely responsible for the rapid growth of polyacrylic PSAs, particularly within the last few decades.

Modern polyacrylic PSA materials include wound coverings, surgical drapes, electrode mounts, and transdermal drug delivery systems for external use. The viscoelastic properties, material stability, and structural tunability of polyacrylics allow for them to ideally address many of the key challenges surrounding biomedical adhesion: wet surface application and mechanical durability.

4.2. Silicones

Silicone adhesives, originally introduced in the 1960s, emerged as pressure sensitive adhesives in the twenty-first century [109]. These semi-inorganic polymers are products of condensation reactions between silanol-functional MQ siloxane resins and high molecular weight silanol-functional silicone polymers (Fig. 3C) [109–116]. The MQ resin, referencing trimethylsiloxy pendant groups on a silsequioxane core moiety, dominates the performance characteristics of silicone PSAs. As MQ content increases, the PSA demonstrates a proportional increase in peel adhesion and lap shear [110]. Regarding tack performance, the structure of both the resin and polymer components is influential since binding between them generates a branched, adhesive network. Crosslinking further improves PSA performance by increasing the cohesive strength of the bulk material.

Silicone PSAs are advantageous in that they exhibit excellent thermal stability due to the backbone flexibility of the polymer’s Si-O-Si bond that permits material flow [26, 110, 117, 118]. Silicone PSAs are most frequently used when the thermal performance requirements deviate from the operating conditions of organic PSAs [110, 119]. At high temperatures (100-250°C) silicone PSAs perform better than acrylic PSAs which decrease in stickiness as temperature rises [120]. Adding to their versatility, silicone PSAs function at very cold temperatures as well.

Silicone polymers demonstrate excellent removability due to strong cohesive strength that stems from the presence of three-dimensional silicone resin within the crosslinked polymer network. In fact, silicone is frequently used as a release liner for pressure sensitive adhesive backings. As reported by Arellano et al., upon applied pressure, the silicone strands intercalate with the acrylic polymer (Fig. 3D) [121]. When the backing is peeled off, most of the silicone is removed with it. Silicone PSAs are also used in bio-analytical applications as adhesive cover slips for multi-well plates, functioning as a removable sealant for biological sample assays [122].

Silicone PSAs are well-suited for biomedical applications due to their excellent adhesion to low surface energy materials, demonstrated tack and peel strength, and good biocompatibility [26, 113, 118]. Silicone PSAs are chemically inert and therefore should not interfere with innate wound healing processes. However, silicones are hydrophobic materials that struggle to adhere in moist conditions. To improve wet adhesion, hydrocolloid fillers are added to absorb water, however this reduces dry condition peel strength and tack. An alternative approach is to introduce hydrophilic additives, such as vinyl ethyl ether and maleic anhydride, which can improve moist condition adhesion and moisture vapor permeability in the bulk silicone polymer [113].

The use of silicone polymers as medical PSAs emerged in the 1980s with use of silicone gel sheeting for the treatment of hypertropic and keloid scars [123]. These polymers were composed of poly(dimethylsiloxane) (PDMS) and MQ resin, establishing what is considered to be the composition of standard biomedical silicone PSAs [110]. Additives such as magnesium stearate or ethyl cellulose improve cold flow and ease of application through disruption of the polymer’s crystallinity [110]. Commonly, silicone PSAs are used in transdermal drug delivery devices. Since the polymer is hydrophobic, it is ideal for delivery of lipophilic payloads to patients. Drug release from silicone adhesives is faster than release from polyacrylic PSAs [124]. As wound care bandages, silicone PSAs such as Mepitel^® are low-adherent, tacky dressings chosen by medical professionals specifically for patients with sensitive skin [125]. A similar adhesive, Mepilex^® Border, demonstrates less patient-experienced discomfort upon removal than cellulose-based or hydrocolloid adhesives removed with the same peel force [126]. Today, PSAs composed of silicone polymers are used exclusively topically.

4.3. Polyurethanes

Comprised of a urethane group (-NHC(O)O-) within the polymer backbone, polyurethanes were first synthesized in 1937 by Otto Bayer, building on previous advancements in the alcoholysis of isocyanates [127–130]. Formed from an addition polymerization reaction between isocyanate and polyol, polyurethanes contain a segmented structure of hard and soft regions, the ratio of which is essential to controlling the viscoelastic properties of the polymer (Fig. 3E) [129, 131–133]. The hard segment, commonly made of 23-48% isocyanate and chain extender (low molecular weight diol or diamine) units, raises the T_g of the polymer and dictates its mechanical properties [129, 132, 134]. The hydrogen bonding capability between isocyanate and chain extender segments generates crystalline regions of the polymer and imparts rigidity and cohesivity [129, 135]. The presence of the soft segment - a high molecular weight polyol, commonly polypropylene glycol or polyethylene glycol – lowers the viscosity of polyurethanes by introduction of backbone flexibility [129, 132, 134, 136]. The presence of both hard and soft segments leads to microphase separations yielding both amorphous and crystalline regions of the polymer. Ultimately, manipulation of the hard-to-soft segment ratio allows for precise control of the elastic modulus and strength, permitting tunability for specific applications, including adhesion.

Polyurethanes are excellent adhesives due to good material stability and inherent tack and peel strength [132, 137]. The low viscosity of polyurethanes allows for spreading and wetting of the adhered surface. Challenges with low peel strength or clean cohesive removal are reported [138], however those concerns can be addressed through careful consideration of polymer architecture (i.e., microneedle arrays) and formulation (i.e., additives). Limiting the degree of polymer grafting and subsequent crosslinking density, such as the g-polyurethane-co-polyacrylate polymers reported by Baron et al. [139] and Degrandi-Contraires et al. [140], improves peel adhesion. For polymers that do contain grafting, restricting the molecular weight of the pendant polymer improves initial tack performance due to less chain entanglements that can interfere with interfacial adhesion [141]. Addition of 2-10 wt% of polyol to the polyurethane polymer further improves tack via enhancing chain mobility [142]. The presence of a small amount of chain extender with hydrogen-bonding capability (less than 10 wt%) also increases initial bond formation between adhesive and surface [143].

Hydroxyl-terminated polyols are particularly effective at increasing microphase separation within the polymer and improving the balance between tack and cohesion [135]. On one hand, reducing the hydrogen-bonding capability within the polyol, or soft segment, lowers viscosity and improves adhesion. On the other hand, increasing the degree of hydrogen bonding increases peel resistance and shear strength [137, 144, 145]. Increasing the polymer’s molecular weight increases the degree of chain entanglements, chain aggregation, and overall shear resistance [131]. Perhaps less well-studied is the design of organic-inorganic composites, where the viscoelastic properties of the adhesive are tuned through the flexibility of polyurethane and the rigidity of inorganic additives such as silica [146]. Addition of silica into the polymer matrix restricts chain mobility and increases crystallinity through hydrogen-bonding with hydroxyl groups on the silica surface [147, 148]. For adhesives, which require a certain degree of amorphous character for substrate adhesion, any addition of inorganic plasticizers must be carefully balanced with tack performance requirements.

As materials for medicine, polyurethanes demonstrate good biocompatibility and are commonly used in devices such as catheters and ostomy gaskets [136, 149–151]. Polyurethanes demonstrate superior breathability, which is critical for topical wound care. Their porous structure facilitates water absorption and moisture transfer at moisture vapor transmission rates (MVTR) greater than 1000 g/m² per day, which is well within the range of a moisture permeable dressing [106, 152, 153]. Of additional importance to surgical dressings is the ability of the underlying tissue to retain normal function upon application. Polyurethanes demonstrate mechanical properties similar to tissue (elastic modulus of polyurethane: 2.0-2.2 MPa [154]; elastic modulus of soft tissue: 1.0 MPa [155]) allowing for natural movement of the adhered area. Finally, polyurethanes are suitable for contact with bodily fluids due to low platelet adhesion and low in vitro protein adsorption [128]. A limited number of polyurethane PSAs for medical applications are commercially available, such as Cyabine^™ (Toyochem Co., Ltd.): a solvent-borne, removable polyurethane film which is restricted to topical application such as surgical drapes and dressing films.

4.4. Polyesters

Polyesters are a promising new class of pressure sensitive adhesives due to their low cost of manufacture, potential biodegradability, and possibility for utilization of bio-derived monomers [156–158]. Chen et al. report the synthesis of polyester pressure sensitive adhesives synthesized through the ring opening polymerization of limonene oxide, tricyclic anhydride, and poly(ε-decalatone) (Fig. 4A) [159]. The block structure of the resulting polymer, alternating between stiff, limonene oxide, and soft, poly(ε-decalatone), regions demonstrates tunable adhesion comparable to commercial adhesives such as Post-It^® Notes and Scotch^® Tape. Limonene oxide, a main component in citrus peels, is a high T_g polymer that improves mechanical strength while the more flexible poly(ε-decalatone) block gives the polymer sufficient tack performance for adhesion. Daristotle et al. report poly(lactide-co-caprolactone) pressure sensitive adhesives capable of bandaging and wound healing with similar efficacy as a commercial polyurethane adhesive [160]. The described adhesive is facilely formulated into a bandage by spraying a thin film of adhesive on a flexible, cloth backing or applied via airbrush directly to the wound.

Fig. 4.

Pressure sensitive adhesives derived from bio-based sources. A. ABA block polyester adhesive from limonene oxide [159], Copyright 2020. Adapted with permission from John Wiley & Sons Inc.; B. ABA block polyacrylics utilizing glucose and isosorbide pendant moieties [161], Copyright 2018. Adapted with permission from American Chemical Society; C. Alginate-polyacrylamide hydrogel adhesive and pH-dependent crosslinking between chitosan chains to facilitate physical entrapment with tissue surface [162], Copyright 2022. Adapted with permission from John Wiley & Sons Inc.

4.5. Polysaccharides

Biologically-sourced synthetic molecules/macromolecules also play a key role in pressure sensitive adhesive composition, particularly for medical applications where toxicity is of utmost importance. When incorporated into polymer architecture, essential oils, simple sugars, and polysaccharides influence the adhesive properties of the resulting macromolecule. For example, Nasiri et al. report the synthesis of an ABA triblock polymer that utilizes glucose to form crosslinked pressure sensitive adhesives [161]. The synthesized thermoplastic elastomer has phase-separated morphology between glassy isosorbide and glucose acrylate tetraacetate regions and rubbery n-butyl acrylate regions (Fig. 4B). To improve the tensile and shear strength of the adhesive, deprotection of the glucose domains affords physical crosslinking between polymer chains through hydrogen bonding.

Further utility for naturally-derived carbohydrates in adhesion can be found through alginate, obtained from seaweed, and chitosan, acquired from algae and fungi. Cinturon-Cruz et al. report the synthesis of tough adhesives with high adhesion energy (> 2,000 J/m²) to tissue through quick, non-covalent interactions [162]. This alginate and polyacrylamide interpenetrating network utilizes chitosan to facilitate hydrogen bonding between deprotonated chitosan-amine groups to increase matrix toughness (Fig. 4C). Specifically, formation of this interpenetrating network is pH-dependent, transitioning past the pKa of chitosan (pH 6.5) to facilitate deprotonation in vivo. The chitosan and alginate-polyacrylamide interpenetrating network binds strongly with tissue primarily through the physical entanglement of chitosan chains with the tissue surface.

4.6. Polystyrene Block Polymers

The final adhesive composition discussed is polystyrene-based block copolymers. These ABA triblock polymers are characterized by a stiff polystyrene “A block” and a soft isoprene, ethylene, or butylene “B block”. The differences in rheology between the block regions of the polymers leads to microphase separation in which the polystyrene is able to generate physical crosslinks to improve cohesion while the center block flows more readily and is responsible for the tack performance of the adhesive. These polystyrene block copolymers are hot melt pressure sensitive adhesives since they are solid at room temperature and can become adherent upon heating. However, through the careful selection of tackifiers, resins, and other additives, these pressure sensitive adhesives are able to adhere at room temperature under light contact pressure [158, 163]. Various poly(styrene-isoprene-styrene) (SIS), poly(styrene-butadiene-styrene) (SBS), and poly(styrene-ethylene-butylene-styrene) (SEBS) materials are currently used as commercial pressure sensitive adhesives and are being investigated for their utility as various tapes, including in transdermal drug delivery devices [163–167].

5. Improving Topical Pressure Sensitive Adhesives

Today, pressure sensitive adhesives such as Band-Aid^®, Nexcare^™, and Mepilex^® function as topical wound care bandages, typically for home-use. While effective hemostats for minor injuries, these products can be improved in key areas of importance such as removability, moisture permeability, and microbial resistance. Additionally, development of topical pressure sensitive adhesives into transdermal drug delivery devices is promising for applications in which continuous therapeutic administration, patient compliance, or on-demand seizure of treatment is of concern. In the following section, recent advancements in synthetic, polymeric pressure sensitive adhesives for topical applications is discussed.

5.1. Removability

The removability of PSAs sets them apart from glue-like tissue adhesives and sealants, and is of critical importance for topical applications in the wound-care field. Currently there is a significant clinical deterrent to using PSAs such as Tegaderm^™, since patients often experience tearing of the top layers of skin, generating additional blisters. New biomedical adhesives that permit removal, repositioning, and wound inspection, ideally with minimal force so that damage to the underlying tissue is not propagated, are of interest. The allowable peel force for removal is subjective and can vary greatly between patient populations. For example, in the neonatal intensive care unit (NICU) where adhesives are used to securely fixate medical life support equipment and monitoring devices such as chest tubes, catheters, and pulse oximeters, medical adhesive-related skin injury (MARSI) occurs in 8-17% of infants within just one day [168]. Biomedical adhesives must therefore be designed to accommodate removability on a range of patient skin conditions and ages, including elderly patients with skin atrophy, loss of elasticity, and increased fragility induced via chronic steroid use [169] and infants whose epidermis is 20-40% thinner than adult skin [168].

The phenomenon of removability highlights the necessary balance between adhesion and cohesion. A strongly cohesive PSA will demonstrate adhesive failure upon removal: complete separation without residual adhesive left on the tissue surface. Clean removal such as this is non-trivial, as many novel PSAs undergo cohesive failure: separation of the bulk material during peeling [73, 88, 170, 171]. Additionally, the change in cohesion during wearing time must be considered, as the polymer may absorb liquids like sweat or secretion from the skin [75]. This may significantly change the rheology, and consequentially peeling behavior, of the PSA.

Adhesives with good removability will demonstrate low peel adhesion. Generally, PSAs with 180° peel strength between 0.4-2 N/cm demonstrate excellent removability to skin while PSAs with peel strength greater than 14 N/cm are considered permanent [172, 173]. To improve their removability, additives can be added to increase the fluidity of the bulk material. The addition of poly(butyl acrylate) oligomers (PBA) reduces the peel strength of a polyacrylic block copolymer by disrupting the intermolecular interactions between polymer chains [79]. In another example, ethoxylated amines or polyethylene oxides are added to polyacrylic PSAs with pendant chain carboxylic acids in order to reduce peel strength. These additives are most effective when they are high in hydrophilic character, most likely due to their ability to compete with hydrogen bonding between the polyacrylic PSA and surface [172].

Polyurethane PSAs demonstrate improved removability when the weight percent of flexible polyol component is increased [132]. Polyurethane elastomers such as these are easily removed from patient skin with minimal pain and without loss of tack performance [153]. Adhesive layer thickness and patterning also influence the removability of PSAs, with thinner adhesive coatings being more removable [172]. Additionally, adhesive patterning in circles, rather than in a homogenous coating, lowers the energy of absorption and eases removal [174].

Silicone wound dressings exhibit low peel forces, yielding excellent removability especially for patients with sensitive skin [110]. These soft silicone adhesives adhere only to dry skin and not the wet wound bed, making removal less painful and traumatic for the patient [125, 175]. One such silicone adhesive, Mepitel^®, is less painful than paraffin gauze dressings when treating the fingertip wounds of children [176]. Importantly, no difference in wound healing time occurs between the treatment groups. Waring et al. report that silicone PSAs are gentler to remove than polyacrylic PSAs through scanning electron microscopy (SEM) and protein assay experiments comparing Mepilex^® Border adhesive to Allevyn^® acrylic adhesive [175]. Following peel testing of each adhesive on patient skin, SEM images reveal higher amounts of matter, subsequently confirmed to be protein, deposited on polyacrylic test strips than the silicone ones. Additionally, despite comparable values for adhesion between groups, significantly less pain was associated with removal of Mepilex^® Border.

The development of polyacrylic stimuli-responsive polymers that undergo physiochemical transitions allow for on-demand removal of PSAs. Tseng et al. report the use of alkoxyphenacyl-based polyurethane PSAs with tunable failure modes of removal given changing photoirradiation times [177]. The p-hydroxyphenacyl group of the photo-responsive PSA can either undergo homolytic or heterolytic cleavage upon irradiation, leading to either adhesive or cohesive failure, respectively (Fig. 5A, 5B). Ossowicz-Rupniewska et al. report that the peel force required to remove a polyacrylate copolymer reduces upon addition of a high T_g bio-based polymer (poly(isobornyl acrylate)) and UV-initiated crosslinking agent into the PSA [178]. The hydroxyl functionality of poly(acrylic acid) segments which can hydrogen bond with the skin’s surface and the bulky nature of poly(isobornyl acrylic) segments which increase matrix toughness generate pressure sensitive adhesives with functionality comparable to commercial products. Crosslinking density, which was directly related to UV-exposure time, increased the cohesive strength and solid-like behavior of the adhesive, consequentially requiring less peel force to remove (Fig. 5C). Light-activated crosslinking can also switch off bonding following irradiation of the polymer. For example, methacrylate functionalized acrylic PSAs containing a visible light photoinitiator demonstrate a 90% reduction in peel strength following light exposure [179]. Exposure to visible light initiates free radical crosslinking between methacrylate pendant chains of the polyacrylate PSA, greatly reducing peel force with the substrate. When utilizing visible light activated crosslinking as a means of removability, care must be taken to shield the adhesive from light by means of an opaque backing until the time of intended removal [174].

Fig. 5.

Stimuli-responsive pressure sensitive adhesives engineered for removability. A. Observed changed in fibrillation following irradiation of polyurethane adhesive and B. Tunable failure modes of alkoxyphenacyl-containing polyurethane adhesives following photoirradiation. During a 180° peel test, both cohesive and adhesive failure was observed at the maximum peel strength. [177], Copyright 2021. Reproduced with permission from American Chemical Society.; C. Multi-functional polyacrylic adhesive demonstrating removability following UV-crosslinking [178], Copyright 2021. Adapted with permission from Multidisciplinary Digital Publishing Institute; D. Illustration of amorphous polymer chains melting upon heating, rendering a non-tacky and removable adhesive.

Temperature-dependent additives in PSAs are also leveraged to improve removability [180]. For example, a “Warm-Off” PSA contains a crystallizable pendant chain polymer (e.g., polyalkylene oxides, alkyl polyesters, polytetrahydrofuran) that has a melting point close to operating temperature. With a slight increase in temperature, the polymer additive passes its melting point and reduces adhesion strength (Fig. 5D) [174, 181]. Thermosensitive silicone PSAs with crystallizable alkyl pendant chains become removable upon cooling below the T_g of the polymer [120]. Other PSAs such as On/Off^®, composed of a hydrophilic base polymer and water-soluble tackifier, do not rely on temperature stimuli for removal. The On/Off^® PSA becomes a hydrogel upon the absorption of water and loses its tack performance, demonstrating a 90% reduction in peel strength after one minute of wetting [174, 182].

5.2. Moisture Permeability

Up until the 1950s, PSAs for skin applications were primarily natural or synthetic rubber and included low molecular weight tackifiers, plasticizers, and stabilizers that often caused irritation [138]. These rubber-based adhesives also exhibited low moisture permeability, leading to significant problems with wound healing. While a moist wound bed is a requirement for natural healing processes, moisture accumulation at the adhesive-skin interface must be controlled in order to avoid problems with skin maceration, odor, bacterial growth, and premature adhesive debonding. The moisture vapor transmission rate (MVTR) describes the passage of water through the polymer matrix of a pressure sensitive adhesive. Injured skin releases moisture at a rate between 279-5138 mL/m² per day [183]. In order to be considered moisture permeable, a PSA generally exhibits a MVTR of greater than 840 g/m² per day [152]. For wound dressings in particular, the recommended MVTR is between 2000-2500 mL/m² per day [184, 185]. This trafficking of moisture through the polymer network is primarily diffusion-based and depends heavily on polymer free volume and chain mobility [186]. PSAs composed of low T_g polymers with lower crosslinking density consequentially demonstrate faster MVTRs.

Acrylic polymers often have insufficient moisture vapor transmission rates compared to highly permeable silicone and polyurethane [69]. Acrylic polymers are more hydrophobic in character with less functional groups for binding and interacting with water. However, it is possible to improve the MVTR of acrylic polymers to approximately 2,000 g/m² per day through strategic monomer selection [187]. Increasing the carboxylic acid content within the pendant chains, which is hydrophilic and therefore permeable to water, is one approach to increase the MVTR [188]. Increasing tackifier concentration typically decreases MVTR due to higher crosslinking densities that subsequently inhibits the diffusion of moisture vapor [189].

The permeability of a PSA to water is also improved through the physical orientation of the adhesive layer. Specifically, introduction of pores throughout the PSA, such as in dot- or net-shaped patterns, allows for moisture to pass. For example, porous adhesive coatings increase MVTR compared to solid-layer PSA coatings [16]. When this strategy of discontinuous adhesive coating is employed, care must be taken to fabricate small pores (<0.5 μm) to block the penetration of infection-causing bacteria (1-2 μm). Another approach to improve MVTR is to combine short- and long-chain monomers within the polymer composition. The variability in pendant chain length allows for material flow at room temperature. Long-chain monomers, 2-ethylhexyl acrylate and isooctyl acrylate, while short-chain monomers, ethyl acrylate and hydroxyethyl acrylate, permit the passage of moisture through the bulk material. In one example of a PSA employing this copolymer strategy, the polyacrylic’s MVTR increased to greater than 1900 g/m² per day [187]. Importantly, this technique maintains the bacterial barrier of a continuous coating while maintaining moisture permeability.

5.3. Microbial Resistance

Impedance of bacterial penetration into the healing wound is essential for robust wound healing and the prevention of chronic wounds. Unfortunately, surgical site infection after surgery does occur. For example, staphylococcus aureus infection occurs in approximately 39% of patients following orthopedic surgery [2]. Many adhesive bandages advertised as anti-bacterial such as Dermabond^™, a 2-octyl cyanoacrylate adhesive, do not have active agents against Gram-positive or -negative bacteria. Instead, these adhesives rely on a physical barrier to microbial penetration. Even with a continuous adhesive coating functioning as a protective layer against infection, small, mobile bacteria strains may still be able to fit through pores within the polymer network [105]. If infection does occur, transparent PSAs as wound dressings would be beneficial in order to visually monitor the wound site for signs of infection and maceration, allowing for incision draining and antibiotic administration as necessary.

Iodide (I₂) is an antimicrobial agent with demonstrated success in infection treatment and potential utility in PSA formulations. Importantly, iodide is a promising antimicrobial since it is an antiseptic agent rather than an antibacterial one. Practically speaking, antiseptic agents are less likely to generate microbial resistance than antibiotics, making them effective against a broad range of infections [190, 191]. Commercial wound dressings such as Iodoflex^® and Inadine^™ incorporate iodide into starch or polyvinyl pyrrolidone (PVP) matrixes, respectively [192]. Complexing iodide with PVP is beneficial, since unbound iodide can lead to side effects such as pain and irritation. Jalil et al. report a muscoadhesive PVP-I₂ complex that increases antimicrobial efficacy by extending iodide’s residence time in the mucosal layer [190]. Thiolation of this PVP hydrogel increases the loading capacity of iodide and viscosity, a predictor of mucoadhesion, compared to PVP alone. Non-PVP, matrixes for iodide delivery such as potassium complexes (KI-I₂) in agar matrixes have been reported [192]. With that being said, during the development of new iodine-containing materials for medical use, emphasis must be placed on patient safety in addition to antimicrobial efficacy. Frequently used in surgery through Ioban^™ (3M^™) and iodine prep, iodine can lead to significant allergic reactions in patients such as skin irritation and blistering.

Inspiration for anti-bacterial properties may also be drawn from nature. Eugenol, 4-allyl-2-methoxypehnol, is derived from essential oils of plants and has demonstrated anti-bacterial activity [193, 194]. In one example, Ji et al. report that pressure sensitive adhesives derived from MQ silicone resins functionalized with eugenol demonstrate enhanced antibacterial effect compared to unmodified resin [195]. Cinnamon oil (CO) is another essential oil that acts as an antimicrobial agent, specifically to gram-positive bacteria. Sengsuk et al. report a Bio-PSA/CO conjugate that reduces microbial counts on food when applied to packaging materials such as kraft paper, aluminum foil, nylon, and polypropylene [196]. Composed of natural rubber, hemicellulose was added to the Bio-PSA/CO to improve the tack of the PSA. The Bio-PSA/CO demonstrated higher tack to porous, rather than smooth, packaging materials due to a mechanical interlocking method of interfacial adhesion.

Silver and silver nanoparticles can also impart anti-microbial functionality into adhesives. In one example of a poly(N-isopropyl acrylamide)-alginate adhesive, the addition of silver nanoparticles prevented bacterial growth without compromising the stretchable, tough, and strongly-adhesive properties of the material [197]. Aminoglycosides or other synthetic molecules can also augment PSA formulations to achieve an anti-microbial effect. Streptomycin [198, 199], gentamicin [199], and chlorhexidine [200, 201] have all shown success at reducing microbial growth in wound dressing formulations. Utilizing these compounds directly within pressure sensitive adhesive formulations will add to the efficacy of topical wound care by preventing infection.

5.4. Transdermal Drug Delivery

Transdermal drug delivery devices (TDDD) were first developed in the 1980s and today more than 50% of TDDDs approved by the FDA are composed of acrylic PSAs (Table 3) [106]. Examples include: Vivelle^®, an estradiol-loaded acrylic adhesive for the treatment of menopause symptoms [202]; Oxytrol^®, a cast film of acrylic adhesive containing oxybutynin and triacetin for the treatment of overactive bladder [203]; and Minitran^™, polyacrylic copolymers containing glyceryl trinitrate employed for the prevention of angina pectoris [204].

Table 3.

Examples of commercial transdermal drug delivery devices and their adhesive compositions.

Trade Name (Drug)	Manufacturer	Indication	Adhesive Composition
Catapres-TTS® (clonidine)	Boehringer Ingelheim	hypertension	polyisobutylene; colloidal silicon dioxide
Vivelle-Dot® (estradiol)	Novartis Pharmaceuticals	menopause symptoms	acrylic
Duragesic® (fentanyl)	Janssen Scientific Affairs	chronic pain	acrylic
Daytrana® (methylphenidate)	Noven Therapeutics	attention-deficit/hyperactivity disorder	acrylic and silicone
Nicoderm CQ® (nicotine)	GlaxoSmithKline	smoking cessation	acrylic
Nicorette® (nicotine)	Johnson & Johnson	smoking cessation	acrylic
Minitran™ (nitroglycerin)	Valeant Pharmaceuticals	angina pectoris	acrylic
Oxytrol® (oxybutynin; triacetin)	Merck	overactive bladder	acrylic
Exelon® (rivastigmine)	Novartis Pharmaceuticals	Alzheimer’s disease; Parkinson’s disease	silicone
Emsam® (selegiline)	Viatris	major depression	acrylic
Tranderm Scop® (scopolamine)	Novartis Pharmaceuticals	motion sickness	polyisobutylene
Androderm® (testosterone)	AbbVie	hypogonadism	acrylic

Drug-loaded adhesive patches allow for the controlled release of therapeutics through the skin and into systemic circulation. This mechanism of delivery offers improved systemic bioavailability over oral or injectable variations that must contend with initial liver and digestive metabolism [106, 205]. TDDDs offer additional advantages such as facile termination of treatment, improved patient compliance, and site-specific delivery. It is still important that PSAs for TDDDs are biocompatible, however they now must also be compatible with the drug of interest. This includes considerations with solubility and miscibility, as well as drug or excipient degradation or chemical reactions. Best drug-in-adhesive solubility results from materials with similar solubility parameters [85]. The kinetics of initial drug release and subsequent penetration through the skin are both highly impacted by drug-PSA interactions and therefore are important to consider during biomaterial design.

Different pendant chain functional groups on polyacrylic PSAs can either inhibit or enhance the skin penetration of drugs, as the rate of permeation is directly proportional to the degrees of freedom within the drug-PSA matrix [206]. Strong interactions between the drug and PSA decrease thermodynamic activity, increase solubility and drug loading, slow the drug release rate, and typically increase bioavailability [207, 208]. Strong ionic bonds form between carboxylic acid pendant chains and positive moieties on many small molecule drugs, such as amine groups, yielding a more sustained drug release profile [207, 209]. Polyacrylics with hydroxyl-functionalized pendant chains, however, enhance the rate of skin penetration via hydrogen bonding with tissue proteins [207]. In these systems, the hydrogen bonding between PSA and drug is not strong enough to significantly influence drug retention within the adhesive, yet increases penetration through hydrogen bonding with the skin’s surface. Commercial adhesives such as Duro-Tak^®, an acrylic copolymer PSA, utilize a similar strategy of modifying pendant chain hydroxyl and carboxyl functionality to tune drug release profiles [210]. To further control diffusion through the TDDD, a non-medicated adhesive layer is often placed between the drug-loaded adhesive and skin as a rate-controlling mechanism [205].

The T_g of the PSA also impacts the rate of skin permeation. Thicker and more viscous PSA matrices slow permeation due to their more restricted polymer chain mobility [208, 209, 211, 212]. As the opportunity for random, thermal motion within the polymer increases, generated pockets of free volume allow for the accelerated release of payloads from the delivery system [211]. The incorporation of a drug itself can disrupt the cohesive forces of the polymer, generating a more mobile and fluid PSA [207]. For example, penetration enhancers such as polyoxyethylene alkyl esters are shown to disrupt drug-PSA interactions and increase drug release when introduced into PSA formulations [213]. Additionally, patterning of the adhesive can improve delivery of drug payloads. Chew et al. report the formulation of a drug-loaded patch with a microneedle adhesive arrangement that penetrates the stratum corneum to release its payload into the epidermis upon swelling [214].

Panchaxari et al. report a TDDD for the delivery of the non-steroidal, anti-inflammatory agent diclofenac [215]. For this payload, a blended formulation of both silicone and acrylic adhesive provides the best results, taking advantage of the higher permeation capability of silicone and the better drug loading capability of acrylics. Recently, alginate-acrylamide tough adhesive hydrogels demonstrate success at a sustained, localized delivery device for corticosteroids to tendons following surgical repair. This adhesive, which is strongly adherent to tendon tissue due to use of a chitosan bridging polymer between the hydrogel and tissue, exhibits a 16-fold increase in adhesion energy to tendon tissue compared to Tisseel^™, the indicated commercial alternative. Importantly, the alginate-acrylamide network loads substantial corticosteroid amounts – 4-times higher than the polymer content – and displays a dissolution-based method of sustained release [216].

6. Advancing Towards Internal Pressure Sensitive Adhesives

The use of adhesives in vivo poses its own unique set of challenges beyond the design requirements for topical adhesives. PSAs for internal use are required to adhere to standards of biocompatibility and safety which ensure clinical use without initiating a local or systemic adverse response in the patient. Since internal PSAs will remain in the body during their lifetime, biodegradation and the toxicity of resulting degradation products is important to consider. Internal adhesives must be able to adhere to wet tissue in a variety of contexts including situations of active blood or fluid flow. It is also imperative that the PSA is mechanically strong, permitting normal functioning of the tissue throughout the wound healing process. In this section, recent developments in synthetic polymers which address biocompatibility, biodegradation, wet surface adhesion, and/or mechanical strength are discussed. Through careful consideration of these design requirements, translation of biomedical PSAs from topical to in vivo use in the clinic is on the horizon.

6.1. Biocompatibility

Regardless of function, all biomaterials must be designed with biocompatibility at the forefront of material development. Biomaterials should have minimal allergenic potential with no elicitation of local or systemic toxicity. The biocompatibility of adhesives depends on the vascularity of tissue, with more highly vascularized regions experiencing higher inflammation than avascular ones [217]. While synthetic polymers avoid many of the toxicity concerns associated with human- and animal-derived adhesives, byproducts of polymer degradation may lead to inflammatory responses in patients and therefore must be fully evaluated. Generally, high charge density within a polymer, particularly cationic, should be avoided due to potential disruption of cell membranes [218–220]. High molecular weight polymers may also increase retention and kidney stress, prompting careful evaluation of the renal clearance of PSA polymers and their byproducts [221–223]. In general, polymers between 30-50 kDa in size cap the upper limit of glomerular filtration [224, 225].

Polymerization mechanisms often use catalysts and initiators which can be toxic if not thoroughly purified out of the formulation. Organic catalysts are valuable substitutes to traditional metal-based catalysts and should be explored further for polymer synthesis. Recently, organic catalysts have been used in ring-opening [226–228] and free radical polymerizations [229–231], indicating a promising start to this new field. Researchers are also exploring alternatives to catalysis-driven polymerization to avoid these potential cytotoxic effects. For example, electron bean curing can be used to polymerize polyurethane adhesives instead of more traditional photoinitiator methods [189]. This method is advantageous since the concern of toxic byproducts upon degradation of the photoinitiator is removed. Other methods of improving PSA biocompatibility lie within the manufacturing process. Waterborne polyurethane PSAs are gaining popularity in commercial applications compared to traditional organic solvent-based film casting methods because such PSAs exhibit reduced toxicity and a smaller environmental footprint [132].

Plasticizers, tackifiers, and other additives are frequently added to PSA formulations to optimize material performance such as tack and peel adhesion [232]. However, these small molecules can potentially leach into body fluids. The intentional selection of polymers for PSA formulations can often render these additives superfluous. For example, polyacrylics are highly suitable for medical applications since they do not require additional stabilizers due to their inherent oxidative stability [106]. However, if small molecule additives are employed in a PSA formulation, care must be taken to ensure that they are biologically inert.

6.2. Biodegradation

The ability for biomaterials to degrade allows for polymers to be broken down and cleared by the body over time. Ideally, PSAs will permit natural wound healing on a physiologically-relevant timescale, with the rate of degradation limiting the potential for a long-term toxicity [218]. This degradation will be different depending on the biological environment at the site of application. The low pH of the stomach, high lipase concentration in the pancreas, and lower temperature of the skin compared to internal tissue will all serve to reduce or accelerate biomaterial degradation in vivo [233–235].The clearance of polymers and their degradation products is highly dependent on size, with molecules smaller than 70 kDa clearing through the renal pathway and the liver metabolizing larger molecules, thus highlighting the importance of long-term safety and biocompatibility testing [223].

In vivo polymer degradation primarily occurs through hydrolytic or enzymatic hydrolysis of backbone and pendant chain linkages. For example, polyurethanes [236–239], polycarbonates [171], and polylactide [240] undergo cleavage by enzymes such as cholesterol esterase and lipase. The lower the thermodynamic barrier for hydrolysis, the faster the rate of degradation, as seen in the case of PEG-based polymers that degrade faster in aqueous media when they contain carboxy-ester linkages compared to less electrophilic thioester linkages [241]. By increasing the hydrophobicity of the polymers, such as with additional methylene groups in PEG diester dithiol and vinyl sulfone hydrogels, the rate of degradation decreases due to reduced water permeability into the bulk material [242].

Likewise, the degradability of a polymer is improved by adding hydrolyzable functional groups into the polymer backbone (Fig. 6A). Freedman et al. report the improved biodegradation of tough hydrogel adhesives through the incorporation of cleavage mechanisms in the polymer formulation [243]. Specifically, degradable moieties were introduced into various entities in a poly(N-isoproyl acrylamide) (pNIPAM) and alginate network. Importantly, by replacing the original N,N’-methylenebis(acrylamide) crosslinker with a hydrolytically degradable alternative such as poly(ethylene glycol diacrylate) (PEGDA), the peel strength, toughness, and adhesion energy of the adhesive did not significantly change. By varying the crosslinker to include either PEGDA or poloxamer diacrylate, or through the replacement of alginate with oxidized alginate methacrylate, the timescale of degradation for these adhesives tunes from weeks to months.

Fig. 6.

A. Inclusion of hydrolysable functional group into the polymer structure will improve biodegradability; B. Degradation of poly(glycerol carbonate) (PGC) pressure sensitive adhesives and the resulting pH-dependent degradation profile [244], Copyright 2019. Reproduced with permission from John Wiley & Sons Inc.

Polymer crystallinity also impacts the kinetics of biodegradation, whether it be hydrolytic or enzymatic. The structure of the polymer’s amorphous, soft segment dictates the rate of polyurethanes degradation [129]. Modifying the semi-crystalline structure of polyurethane’s hard segments improves degradation, such as by introducing hydrolyzable chain extenders into the hard segment accelerates polymer degradation [129]. Additionally, varying the block and chain lengths within the polymer disrupt hydrogen bonding alignment, and thus decrease crystallinity and increase degradation [38].

Newly developed poly(glycerol carbonate) (PGC) adhesives, degrade via hydrolytic degradation of backbone carbonate linkages into non-toxic, natural metabolites like glycerol and carbon dioxide (Fig. 6B) [73, 171, 244]. These polycarbonates are biocompatible, biodegradable, and demonstrate success as biomedical PSAs through the attachment and subsequent release of a collagen buttress to a surgical stapler in an ex vivo lung resection model [171]. Additionally, PGC adhesives degrade in a pH-dependent manner, with the fastest degradation occurring in basic conditions (pH 9) due to a higher prevalence of hydroxide ions that initiate backbone hydrolysis [244].

6.3. Adhesion to Wet Surfaces

Skin and tissue are low surface energy materials coated in oils, sweat, water, and salts that pose a significant challenge for adhesion. Further, some tissues contain lipids, fats, and polysaccharides on their surface. For example, a glycosylate layer is present on the liver and the internal wall of arteries and veins and will likely adversely affect PSA adhesion. The adsorption of water between the tissue and adhesive frequently leads to failure of contact adhesion [245]. Current medical adhesives often lack in adhesive strength to wet surfaces or employ relatively extreme measures to achieve sufficient adhesion. Sutures are sometimes required to slow wound bleeding before application of adhesives such as BioGlue^®, which are restricted in functionality in the presence of excess fluid [246]. Some electrocardiogram electrode manufacturers recommend that for adherence to extremely wet skin, such as the moisturized vernix caseosa of full-term infants, the skin be prepped with isopropanol to de-wet it and improve adhesion [168]. This irritates the patient’s skin and prompts for less invasive solutions.

Relying on non-covalent interfacial interactions such as van der Waals forces and hydrogen bonding is a safe alternative to the chemical curing of cyanoacrylate adhesives, like Dermabond^™, to wet tissue. When the PSA spreads and wets the substrate surface, the resulting adhesive bond is further strengthened by the multitude of intermolecular interactions at the interface [58, 247]. For polymers, these interactions are driven by pendant chain characteristics. Longer side chains lead to better flexibility and higher tack, since the chains are able to slide across one another and flow more easily at room temperature [72, 248]. These longer chains increase the incidence of van der Waals forces, where temporary dipole moments result in attractive forces between molecules and a strengthening of the adhesive bond, and thus, greater tack (Fig. 7A). In one experiment, copolymers with a high percentage of three-carbon pendant chains, compared to one-carbon pendant chains, exhibit adhesion strength similar to Scotch^® Tape. When the pendant chain length is shortened to include more one-carbon pendant chains, the adhesive is stronger and less tacky due to reduced flexibility among the side chains [171]. However, increasing the pendant chain length too much leads to crystallization of the amorphous polymer, stiffening it and rendering it ineffective as an adhesive because of reduced flow and wetting capability [106, 249]. Balance, therefore, is critical for polymer design to ensure retention of the PSA’s low T_g and ability to perform pressure-induced viscous flow. In fact, six- to eight-carbon length alkyl groups on polyurethane prepolymers, as reported by Wang et al., generate the ideal balance between hydrophobicity and hydrophilicity in aqueous environments like tissue [249].

Fig. 7.

Key mechanisms of wet-surface adhesion including examples of A. Van der Waals forces [171], Copyright 2021. Reproduced with permission from Royal Society of Chemistry.; Hydrogen bonding through B. acrylic acid [251], Copyright 2017. Adapted with permission from American Chemical Society, and C. A-POSS cages [252], Copyright 2019. Adapted with permission from Multidisciplinary Publishing Institute.; and D. L-DOPA covalent crosslinking [245], Copyright 2017. Adapted with permission from Nature Portfolio.

For adhesion to wet tissue, hydrogen bonding is particularly important. Electrostatic interactions between pendant chain electronegative atoms and surface hydroxyl or amine groups further strengthen the binding interaction. Low T_g polyacrylate polymers, such as the copolymer of 2-methoxyethyl acrylate and N-allylthiourean, possess good wetting properties on an underwater surface [250]. To further improve the rate of adhesive bond formation, poly(butyl acrylate) is copolymerized with acrylic acid which hydrogen bonds quickly to free hydroxyl groups on the substrate surface (Fig. 7B) [251]. The hydrogen bonding capability of the copolymer now dictates the orientation and timescale of the PSA interaction with the surface. Hydrogen bonding also increases the adhesion of PSAs to metal equipment and metal surgical tools. Kowalczyk et al. report that adhesion to steel improves for a butyl acrylate and glycidyl methacrylate copolymer when acryloxypropylheptaisobutyl-polyhedral oligomeric silsesquioxane (A-POSS) cages are incorporated into the polymer. The oxygen-silicon bond on the A-POSS cage hydrogen bonds with the hydroxyl groups on the steel surface (Fig. 7C) [252].

Inspiration for successful wet-surface adhesion is also present in the composition and hierarchical architecture of natural, biological adhesives such as those utilized by mussels, geckos, and worms. A unique feature of mussel adhesive proteins is the ability to participate in both bulk and interfacial reactions important for strong adhesion, via enrichment of 3,4-dihydroxyl-L-phenylalanine (DOPA) within the protein [253]. Oxidation of the catechol side chain of DOPA gives rise to intermolecular crosslinking reactions that solidify the mussel’s excreted protein glue [254]. Taking further inspiration from mussel adhesion, incorporation of polydopamine results in the ability to turn a variety of materials (i.e. noble metals, metal oxides, semiconductors, ceramics, synthetic polymers) into adherent surfaces [255, 256]. Zhao et al. report a methoxyethyl acrylate, adamantine, and DOPA terpolymer, inspired by mussels, which performs as a pressure sensitive adhesive underwater (Fig. 7C) [245]. Furthermore, incorporation of poly(N-isopropylacrylamide) (pNIPAM) into the formulation affords an on-demand temperature dependent PSA, which undergoes conformation changes at elevated temperatures resulting in a loss of hydrogen bonding with the surface. This loss of hydrogen bonding exposes the previously-shielded DOPA pendant chain of the acrylic adhesive, leading to stronger adhesion.

6.4. Mechanical Strength and Durability

Whereas adhesive forces are imperative for initial bond formation, cohesive forces of attraction are necessary for the PSA to resist external loads without breakage, especially for long term applications [74, 232, 257]. Resistance to flow is essential once the PSA is applied so that the adhesive does not continue to spread outside the bounds of the adhesive backing. Increasing the molecular weight and T_g of the polymer enhance its cohesive forces, however a balance must be maintained to ensure good tack and adhesion strength. For example, poly(methyl methacrylate) (PMMA) possesses an inherently high T_g of 105°C [107]. By increasing the monomer feed ratio of methyl methacrylate in a terpolymer composed of 2-ethylhexyl acrylate and acrylic acid, the holding power of the PSA improves while peel strength worsens [107]. Additives can also be introduced into the PSA formulation to improve cohesion. Cellulose nanocrystals, in small amounts, enhance shear resistance of a poly(n-butyl acrylate-co-2-ethyl hexyl acrylate) PSA without sacrificing peel strength and loop tack [258].

Crosslinking is the most common method used to improve the shear resistance of a PSA (Fig. 8A). Polymers are crosslinked physically through hydrogen bonding or chemically through covalent bond formation [78]. Some crosslinking methods require induction with visible or UV-light. While these methods will be discussed, for biomedical applications specific care must be taken to ensure the practicality and safety of these light-based systems in vivo. Particularly for internal utilization, application of light at the treatment site must be efficient and effective during the surgical procedure. The toxicity of each PSA component on the target tissue must be considered individually, in addition to the safety of the final crosslinked adhesive. It is imperative that minimal, if any, heat is generated from curing with light, as heat on blood vessels can activate the clotting cascade and lead to thrombosis. These safety considerations may not always be feasible, and therefore light-induced crosslinking will not be appropriate for all applications.

Fig. 8.

A. General schematic of crosslinking via chemical, UV-, or visible-light; B. UV-induced crosslinking of polyacrylic PSAs with perfluorophenylazide pendant chains [260], Copyright 2022. Adapted with permission from American Chemical Society.; C. Crosslinking of polyacrylic PSA with blue LED light [262], Copyright 2021. Adapted with permission from Multidisciplinary Digital Publishing Institute. D. Polyacrylic PSAs with dual hydrogen bonding and NHS-induced chemical crosslinking capability for the closure of gastrointestinal wounds [263], Copyright 2022. Adapted with permission from American Association for the Advancement of Science.

That said, crosslinking with UV-light is another popular approach considering the good solubility and temperature resistance of many photoinitiators [74, 232, 259]. Abu Baker et al. recently describe the chemical crosslinking of perfluorophenylazide groups using UV light to generate polyacrylate pressure sensitive adhesives where density of crosslinks tunes the performance (Fig. 8B) [260]. The crosslinked polymer’s adhesive properties depend on both the number of UV-curable functional groups and the method of UV-exposure. In an experiment using butyl acrylate, 2-hydroxyethyl acrylate, and glycidyl methacrylate terpolymers with various UV-curable monomers, Saiki et al. demonstrate that a higher percentage of UV-curable functional groups increases the storage modulus of the PSA, reducing viscosity and improving durability [261]. Similarly, Zhu et al. report that acrylic polymer-based PSA viscosity increases with longer UV exposure time due to increased free radical production and subsequent polymer formation [78]. Czech et al. define a relationship between the UV method and adhesive performance of a 2-ethylhexyl acrylate, ethyl acrylate, and acrylic acid terpolymer [74]. Here, higher photoinitiator concentration and longer UV-exposure time leads to a stiffer material and reduction in tack performance due to a higher degree of crosslinking. Peel adhesion changed in a dose-dependent manner between photoinitiator concentration and UV-exposure time. The less photoinitiator used, the longer the curing time necessary to reach a maximum in peel adhesion.

Visible light-induced crosslinking is perhaps a safer alternative to UV-exposure, particularly where longer exposure is required. Back et al. report the radical polymerization of a n-butyl acrylate, 4-hydroxybutyl acrylate, and isobornyl acrylate polymer that is cured under LED (light-emitting diode) light into a pressure sensitive adhesive film (Fig. 8C) [262]. The inclusion of α-haloester additives to initiate the photoredox-mediated atom transfer radical polymerization increased the rate pf polymerization and film curing. However, excess initiator reduces PSA molecular weight and consequentially peel strength and loop tack, necessitating a careful balance between initiator and photocatalyst amounts for optimal adhesive properties. The resulting polyacrylic PSA films were optically clear, which can be useful for medical applications where wound inspection following adhesive application is critical.

The pendant chain structure of polyacrylics is incredibly versatile and examples are readily available of PSAs which chemically crosslink intramolecularly or which bind specific moieties on a substrate’s surface. For instance, polyacrylic PSAs crosslinked using tert-butyl peroxypivalate improve the shear strength of the polymer, extending the time until observed cohesive failure [170]. Aziridine-based agents crosslink with the carboxyl groups from acrylic polymer pendant chains, increasing holding power and improving heat resistance [107]. N-hydroxysuccinimide (NHS) esters are particularly useful agents for in situ crosslinking in that they react with both primary amines and thiols on the adhered surface [218]. In a recently developed bioadhesive patch for the closure of gastrointestinal wounds, an interpenetrating network containing poly(acrylic acid) with NHS-ester (PAA-NHS)-pendant groups chemically crosslink to tissue by forming amide bonds with surface proteins [263]. Initially, these PAA-NHS polymers hydrogen bond with the wet surface of the colon. Within five minutes of application time, the NHS pendant chains chemically crosslink with surface amines on the tissue, further strengthening the adhesive bond (Fig. 8D). Excitingly, the PAA-NHS adhesive out-performed commercial alternatives regarding tensile strength, shear strength, interfacial toughness, and burst pressure in an ex vivo porcine colon. This crosslinking functionality extends beyond polyacrylics to other polymers as well. Polyurethane pressure sensitive adhesives crosslinked with polyisocyanate demonstrate improved peel strength up to a crosslinking density of 1 weight percent [88].

7. Future Directions and Outlook

Recent developments in synthetic polyacrylics, polyurethanes, silicones, and biologically-derived polymers are yielding promising materials for medical pressure sensitive adhesives. The purpose of this review is to provide background knowledge on PSAs, to summarize recent successes in the laboratory, to share our enthusiasm for recent advances in this field, and to encourage others to perform research so that we attain the goal of expanding PSA use from topical to internal. Through understanding of the viscoelastic material design requirements of PSAs, as well as opportunities in physical and chemical crosslinking, we will engage in the rational design of new polymers, with the goal of addressing specific shortcomings in medicine. Improvements to the removability, moisture permeability, microbial resistance, and drug delivery capability of existing topical pressure sensitive adhesives will usher in products with enhanced capabilities as bandages, drug delivery devices, and equipment fixation materials. By prioritizing efforts to create biocompatible, biodegradable, and mechanically strong polymers that adhere securely to wet tissue, PSAs may function in internal wound closure applications – an indication currently void of commercial PSA products.

There are a number of clinical applications in surgery, interventional radiology, and gastroenterology where PSAs would be particularly useful. This is particularly true given the dramatic shift towards minimally invasive procedures, whereby direct access to tissues and structures is via small port incisions, endoscopic or percutaneous access. The ability to close holes, stick patches, seal anastomoses or reinforce tissue closures without the need to suture through large exposing incisions would significantly decrease tissue trauma, risk of infection, surgical morbidity, and length of hospital stay and patient recovery. In addition, portable, storable, easily applied PSAs would increase access to surgical care and interventional procedures, even in more remote areas, with limited access to more advanced surgical instrumentation.

Beyond polymer synthesis, new post-polymerization processing of polymers will enable optimization of material properties. Recent investigations into the control of PSA properties through mechanical, rather than chemical, means allows for fine control over the shear, tack, and peel strength of the adhesive. Through the physical blending of low T_g (high tack) and high T_g (high shear resistance) polymers [264], the blending of low and high molecular weight polymers [265], or the spinning of blended polymer fibers [160], unique pressure sensitive adhesives are generated with tunable properties. Such post-polymerization processing provides avenues to generate highly tunable PSAs and warrants further investigation.

Additionally, the focus on novel PSAs for medicine will lead to formulations that are broadly applicable to other commercial industries. For example, new PSAs that emphasize hydrolytic and enzymatic degradability for medical applications are of significant interest as alternatives to PSAs used in common household products and may improve the lifecycle of single-use plastics. Today, 75% of plastics are discarded after a single use, burdening waste disposal capacity in landfills [266]. As such, while the potential impact of novel biomedical pressure sensitive adhesives to patients is far-reaching, the synthetic advances and material insights from the development of these materials will extend across industries.

Abbreviations:

A-POSS

acryloxypropylheptaisobutyl-polyhedral oligomeric silsesquiloxane

BSA

bovine serum albumin

CO

cinnamon oil

DOPA

3,4-dihydroxyl-L-phenylalanine

FDA

Food and Drug Administration

H&E

hematoxylin and eosin

IOAL

intraoperative air leaks

LED

light-emitting diode

MARSI

medical adhesive-related skin injury

MVTR

moisture vapor transmission rate

NHS

N-hydroxysuccinimide

NICU

neonatal intensive care unit

PAA

polyacrylic acid

PBA

polybutyl acrylate

PDMS

polydimethyl siloxane

PEG

polyethylene glycol

PEGDA

polyethylene glycol diacrylate

PGC

polyglycerol carbonate

PMMA

polymethyl methacrylate

pNIPAM

poly(N-isopropyl acrylamide)

PSA

pressure sensitive adhesive

PTFE

polytetrafluoroethylene

PVP

polyvinyl pyrrolidone

SEM

scanning electron microscopy

TDDD

transdermal drug delivery device

T_g

glass transition temperature

UV

ultraviolet