Degradable Adhesives for Surgery and Tissue ... - ACS Publications

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Degradable Adhesives for Surgery and Tissue Engineering Vrushali Bhagat† and Matthew L. Becker*,†,‡ †

Department of Polymer Science and ‡Department of Biomedical Engineering, The University of Akron, Akron, Ohio 44325, United States ABSTRACT: This review highlights the research on degradable polymeric tissue adhesives for surgery and tissue engineering. Included are a comprehensive listing of specific uses, advantages, and disadvantages of different adhesive groups. A critical evaluation of challenges affecting the development of next generation materials is also discussed, and insights into the outlook of the field are explored.

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INTRODUCTION Many surgical procedures are performed worldwide and the number continues to grow every year. In a recent study, over 300 million surgeries were performed in 2012; an increase of 33.6% over the last 8 years.1,2 Current surgical closure techniques involve the use of invasive techniques like sutures, staples, or clips, which often result in secondary tissue damage, microbial infection, fluid or air leakage, and poor cosmetic outcome.3 Despite being the current standard of care, the use of sutures is tricky during intricate and sensitive surgeries such as vascular anastomosis, nerve repair, or ocular surgeries because of the high risk factors involved. The use of metallic grafts like plates, pins, or screws is common practice in orthopedic surgeries for assistance in osseointegration. However, the modulus mismatch between the stiff graft material and the bone tissue can result in localized stress at the point of fixation and bone atrophy. The metallic grafts act as support mediums that lack chemical interaction with the bone and often suffer from aseptic loosening resulting in poor bone healing.4 Despite these shortcomings, sutures, staples, and metallic grafts still remain the gold standard3 for tissue reconstruction owing to a lack of noninvasive techniques capable of outperforming them. An appealing option to alleviate the use of these invasive techniques is the use of tissue adhesives. An adhesive spreads over the entire contact area, which eliminates stress localization facilitating load transfer between the fractured surfaces.4,5 Additionally, adhesives are easy to apply, join dissimilar materials, increase design flexibility, improve cost effectiveness, can act as sealants or hemostatic agents to prevent fluid leakage through the anastomosed site, and cause minimal or no tissue damage at the application site.6 The attractive advantages of using adhesives over traditional closure techniques have led to a steep rise in adhesive research and development, and it is estimated that adhesives currently constitute a market share of ∼$38 billion.7 © XXXX American Chemical Society

Bone and tissue adhesives have been around for centuries with one of the most ancient materials used as a tissue adhesive being “Plaster of Paris”. Since then, a number of naturally derived semisynthetic or synthetic adhesives have been developed. Donkerwolcke and Muster have discussed the evolution of adhesives from as early as the 1700s to the 90s.8 The journey from sutures or metallic implants (metallic plates, pins, and screws) to tissue adhesives,8,9 the variety of tissue adhesives currently available,3,9−11 their clinical applications,12−14 and future prospects11,15 have been extensively covered in previous review articles.16−20 On the basis of their function, tissue glues can be divided as hemostats, sealants, and adhesives. Although they are often addressed interchangeably, they are quite different from each other. A hemostat is responsible for blood clotting and fails to function in the absence of blood. A sealant develops a barrier layer to prevent leakage of fluid or gas, and an adhesive functions to bind two surfaces firmly and hold them together. Although a hemostat performs effectively in the presence of blood, sealants and adhesives often fail to perform under wet conditions.14,15 A tissue adhesive should have strong wet adhesion, stability under physiological conditions, rapid curing/ cross-linking without excessive heat generation, nontoxicity, cytocompatibility, minimum swelling, modulus comparable to the underlying tissue, biodegradability, and bioresorbability. Adhesion between two substrates is a result of a combination of adhesive and cohesive strength. For strong adhesive performance, it is imperative to attain an optimum balance between the adhesive and cohesive strengths of the material.8 Adhesive and cohesive interactions involve mechanical interlocking, intermolecular bonding, electrostatic bonding, chain entanglement, or cross-link formation.21,22 Received: July 8, 2017 Revised: August 14, 2017

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DOI: 10.1021/acs.biomac.7b00969 Biomacromolecules XXXX, XXX, XXX−XXX

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Biomacromolecules Initial attempts to develop tissue adhesives involved use of epoxy resin, polyurethane foam, poly(methyl methacrylate), and calcium/magnesium phosphate-based bone cement, lactide−methacrylate-based systems, and zinc polycarboxylateand glass ionomer-based cements. The commercial or clinical application of these materials was limited due to their lack of degradability, low bonding strengths, high infection rates, foreign body reactions, or leaching of toxic metallic ions into the body.9,11 Adhesives like fibrin glue, cyanoacrylate glue, and gelatin-resorcin formaldehyde/glutaraldehyde glues, which do have certain limitations like low bonding under wet conditions and poor cytocompatibility, have been approved for clinical use and discussed in the following sections in this paper.12,14,23 Recently, there has been considerable interest in the development of biomimetic tissue adhesives inspired from natural examples of adhesion like mussels, sandcastle worms, caddisflies, and geckos. Though the exact mechanism of adhesion in these organisms is still under debate, several attempts to mimic these adhesives have been made by different research groups and will be discussed further in this article.24,25 The current paper reviews the structure, composition, and functioning mechanism of semisynthetic and synthetic tissue adhesives, hemostats, and sealants. Because biodegradability and tissue compatibility are essential requirements for tissue adhesives, only biodegradable tissue adhesives are discussed in this article. In addition, their adhesion/bonding strengths, suggested clinical applications, and limitations are also reviewed.

Figure 1. Mechanism of clot formation in fibrin glue resembling physiological coagulation.

wound site. Simultaneously, cross-linking also occurs between the adhesive glycoproteins with collagen and other tissue proteins. The cumulative result of all the cross-links at the wound site and the presence of plasmin inhibitors is the formation of a strong, adhesive, insoluble clot resistant to fibrinolysis.28 Both autologous and homologous fibrin sealants have been developed based on whether the plasma is obtained from the same patient or another person, respectively. The fibrin glue is biocompatible, resorbable, and does not cause tissue necrosis, fibrosis, or inflammation. The degradation time of fibrin glue varies from a few days to months depending on the composition.14 Despite the use of fibrin glue as a hemostatic agent for a range of surgeries, the risk of virus transmission still prevails. The fibrin glue components are subjected to virus screening and virus inactivation or reduction treatments like pasteurization, two step vapor heat treatment, solvent-detergent cleansing, dry heat treatment, nanofiltration, precipitation, pH treatment, and some chromatographic steps. However, a particular treatment is not effective against all the viruses, and a combination of these treatments is generally required for medical application.29 An extensive review on the composition and fairly recent (year 2013−2014) medical applications of fibrin glue as a hemostat, sealant, or adhesive has been compiled by Spotnitz.14 The composition of commercially available fibrin glues from different parts of the world compiled by Jackson is presented in Table 1.29 A comparison between the adhesives emphasizes that they differ in concentration of their main components with fibrinogen and thrombin, the source of thrombin, as well as the method used for virus inactivation. The mechanical strength of the fibrin clot is a function of the fibrinogen concentration and often used as the measure of the glue quality, whereas the thrombin concentration directly relates to the speed of clotting. Therefore, an optimum concentration of both components is necessary to achieve rapid hemostasis as well as satisfactory adhesion and mechanical properties.27,30 Adhesion strength of the fibrin glue depends on the substrate, composition of the glue, method of fibrinogen preparation, presence of water, fat, or collagen, and setting time, making it redundant to deduce a universal value for the adhesion strength of fibrin glue.31

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FIBRIN GLUE A primitive, unprocessed form of fibrin glue containing fibrinogen and thrombin was first introduced in the 1940s.26 Alving and co-workers in 1995 summarized different fibrin compositions, their applications, adverse reactions, or complications arising from their use, possible new applications, and the need for controlled trials to determine their clinical efficiency. The fibrin glues synthesized in Europe were a step ahead compared to those in the USA in which their compositions involved the use of antifibrinolytic agents like aprotinin and epsilon amino caproic acid, though the efficacy of using such antifibrinolytic agents was not proven.27 As described by Martinowitz and Saltz, clot formation in fibrin glue resembles the final step in physiological coagulation.28 Briefly, fibrin sealants consist of two major components: fibrinogen with factor XIII and thrombin with Ca2+. Thrombin cleaves off fibrinopeptide A and B from α and β chains (respectively) in fibrinogen to form fibrin monomer. The asformed monomer physically cross-links via hydrogen bonding to form an unstable clot. Factor XIII is a fibrin stabilizing factor activated by thrombin using Ca2+ as a cofactor to form factor XIIIa. Factor XIIIa then acts upon the fibrin monomer or the unstable clot to form cross-links in the form of amide bonds between glutamine and lysine residues resulting in an insoluble clot resistant to proteolytic cleavage (Figure 1). The cross-linking reaction also involves attachment of plasmin inhibitors like α2 plasmin inhibitor (α2-PI), α2macroglobulin, and plasminogen activator inhibitor 2 (PAI-2) to the α chain of fibrin, which further strengthens the clot and prevents fibrinolysis. Factor XIII also acts on other adhesive glycoproteins like fibronectin, thrombospondin, vitronectin, and von Willebrand factor. A number of cross-linking steps are involved in clot formation; for example, fibrin primarily crosslinks with both collagen and adhesive glycoproteins at the

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GELATIN-RESORCINOL-FORMALDEHYDE/GLUTARALDEHYDE GLUE (GRFG) Gelatin-resorcin-formaldehyde/glutaraldehyde glue was reported as early as 1966. This adhesive is a combination of all the individual components. The resorcin-formaldehyde forms a B


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cross-linked polymer in basic conditions as shown in Figure 2.32 Adhesives based on formaldehyde have strong initial bonding, and glutaraldehyde (GA)-based adhesives show higher stability in vivo. Therefore, the adhesive formulations sometimes incorporate both formaldehyde and glutaraldehyde to improve adhesion as well as in vivo stability. Gelatin incorporation in the adhesive imparts elasticity comparable to the underlying tissue as well as degradability.32,33 Bonding strength of GRFG on dry substrate is comparable to cyanoacrylate glue and significantly stronger than fibrin glue but deteriorates under wet conditions.33 Histology studies on rat femoral vessels after GRFG application indicated that the long-term adhesive nature of the glue is not intrinsic but arises from extracellular remodeling around the blood vessels.34 GRFG has been used as a hemostatic agent, adhesive for vascular surgeries, gastrointestinal surgery, thoracoscopic operations, and lung surgeries.34−39 GRFG glue with a gelatin-resorcin:formaldehyde/ glutaraldehyde ratio of 2:1 has shown maximum adhesion strength of 170.5 ± 41.5 kPa in dry conditions and 47.8 ± 17.6 kPa under wet conditions. Although GRFG glue has shown impressive hemostatic properties and satisfactory adhesive properties, the possible carcinogenicity associated with the use of aldehydes limits the clinical use of this glue.33,35 Results of the histological studies on tissues treated with GRFG glues from different research groups have been conflicting.34 Shortterm effects were limited to minimal tissue necrosis and local tissue inflammation in most of the studies; however, more studies are required to evaluate the long-term or cytotoxic effects of the glue. A few other studies have incorporated GA as a cross-linking agent to promote tissue adhesion. Matsuda and co-workers demonstrated the benefit of adding GA to gelatin films to promote tissue adhesion. The dual role of GA cross-linking with the amine groups on gelatin as well as interactions with the amine groups on tissue leads to the impressive adhesion strength of these GA cross-linked gelatin films. The aldehyde content in the system was dependent on temperature, pH, treatment time of gelatin with GA, and GA concentration; consequently, the bonding strength increased with increasing aldehyde concentration. The bonding strength significantly decreased after aldehyde reduction, confirming the role of aldehydes in promoting adhesion. The maximum bonding strength was 250 gf/cm2 (24.5 kPa) on dry porcine skin and deteriorated to negligible bonding on wet skin.40 Addition of proteinoids to the gelatin-glutaraldehyde system resulted in improved adhesion strength along with lower toxicity. Proteinoids are synthetic copolymers of amino acids, particularly RGDKANE, which increase the functionality in the adhesive leading to improved cross-linking and bonding strength.41 Gelatin−resorcin-based adhesives cross-linked with epoxy compound (GRE), water-soluble carbodiimide (GRC), or genipin (GG) instead of formaldehyde or glutaraldehyde showed longer cross-linking time and lower bonding strength. GRC and GG glue had greater cytocompatibility while GRE glue was not deemed suitable for clinical applications.42 Jebrail and co-workers used GA as a cross-linker for bovine serum albumin (BSA) adhesive. The bonding strengths of the adhesive were studied on wood samples, and an impressive bonding strength of ∼6.74 MPa was obtained on hydrated wood. However, the mechanism of adhesion was not clearly explained. The bonding strengths on wood samples are not translational to clinical or biological adhesion, making it difficult to infer performance in physiologically relevant models.43

Bolheal (Kaketsuken Pharmaceutical, Japan) Biocol (LFB-Lille, France) VIGuard F.S. (Vitex: VI Technologies, USA)

80 127 50−95

frozen solution lyophilizate lyophilizate lyophilizate

75 11 3−5

250 558 200

none (92 mg/mL tranexamic acid) 1000 3000 none 60−100

lyophilizate

Hemaseel (APR Haemacure, Canada) (As Tisseel VH Kit Baxter-Immuno) Quixil (Omrix Biopharmaceuticals SA, Israel)

None 75−115

lyophilizate Beriplast P (Aventis Behring, Germany)

75−115 lyophilizate Tisseel (VH Kit Baxter-Immuno AG, USA)

90(65−115)

1000

3000

1000 500 (400−600)

500

3000 500

60 (40−80)

3000 500 and 4 10−50 70−110

Reprinted from Jackson, M. R. Fibrin sealants in surgical practice: An overview, 1S-7S, Am. J. Surg., 182, Copyright 2001, with permission from Elsevier.

two-step vapor heat at 60 and 80 °C two-step vapor heat at 60 and 80 °C two-step vapor heat at 60 and 80 °C pasteurization (liquid solution, 10 h at 60 °C) two-step vapor heat at 60 and 80 °C solvent-detergent treatment, nanofiltration dry heat (96 h at 65 °C) solvent-detergent treatment solvent-detergent treatment, ultraviolet C light two-step vapor heat at 60 and 80 °C two-step vapor heat at 60 and 80 °C two-step vapor heat at 60 and 80 °C pasteurization (liquid solution, 10 h at 60 °C) two-step vapor heat at 60 and 80 °C solvent-detergent treatment, pasteurization dry heat (144 h at 65 °C) solvent-detergent treatment solvent-detergent treatment, ultraviolet C light 3000 500 10−50 70−110

frozen solution lyophilizate Tisseel, Tissucol (Duo Baxter-Immuno AG, Austria) Tisseel, Tissucol (Kit Baxter-Immuno AG, Austria)

Table 1. Composition of Commercial Fibrin Glue

human fibrinogen (mg/mL)

human factor XIII (U/mL)

human thrombin (IU/mL)

bovine aprotinin (KIU/ mL)

virus-inactivated fibrinogen

virus-inactivated thrombin

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C


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Figure 2. Cross-linking reaction of resorcin and formaldehyde under basic conditions. Reproduced from Lin, C.; Ritter, J. A. Effect of synthesis pH on the structure of carbon xerogels, Carbon, 35, 1271−1278, Copyright 1997, with permission from Elsevier.

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in tissues whereas octyl α-cyanoacrylate exhibits minimal inflammation.44,48 N-Butyl-2-cyanoacrylate-based commercial glues, Histoacryl (Germany), and Glubran (Italy) are commercially used in Europe as tissue adhesives during endoscopic surgeries. 2-Octyl cyanoacrylate-based glue Dermabond (USA) is FDA approved for medical use in topical applications only. Mizrahi and co-workers replaced the alkyl side chains with alkoxy chains (ether linkages) to improve the elastic and mechanical properties of the glue. 2-Octyl cyanoacrylate (Dermabond), however, demonstrated higher adhesion compared to that of the alkoxy analogue.49 Bond strengths of modified cyanoacrylates alkyl 2-cyanoacryloyl: glycolate and 1,2-isopropylidene glyceryl 2-cyanoacrylate, was comparable to alkyl 2-cyanoacrylates and exhibited rapid degradation.50−52 Cyanoacrylates were also modified by copolymerizing with 1,1,2-trichlorobutadiene-1,3-methyl methacrylate (MMA), incorporation of elastomeric polymers, and addition of fillers or modifiers like polymeric oxalates to improve their bonding strength, flexibility, and biocompatibility.53−58 Table 2 is an extensive list of surgical and medical applications of cyanoacrylates compiled by Leggat and coworkers in addition to a few other references.59−62 Despite the attractive properties and impressive wet adhesion, the use of cyanoacrylate glue is limited to topical applications owing to the toxic nature of the degradation products. The shorter alkyl chain cyanoacrylates degrade faster, resulting in a stronger tissue inflammatory response, whereas the longer alkyl chain cyanoacrylates degrade rather slowly, preventing the buildup of toxic products and facilitating their removal from the body resulting in mild inflammation. Nevertheless, cyanoacrylates have caused chronic inflammation, tissue necrosis in vivo, intimal thickening, arterial ocular lesion, occupational asthma, dermatitis, and in vitro cytotoxicity for cells in direct contact as well as in leached solutions to list a

CYANOACRYLATE GLUE Cyanoacrylate glue is a class of synthetic adhesives made from alkyl α-cyanoacrylates for tissue adhesive application since the 1980s. Cyanoacrylates polymerize rapidly in the presence of weak basic conditions, for example, water or blood (Figure 3).

Reproduced from ref 25 with permission of The Royal Society of Chemistry. http://dx.doi.org/10.1039/ c2bm00121g.

Figure 3. Mechanism of cyanoacrylate polymerization.

The amine groups in proteins present on the tissue surface are also speculated to initiate cyanoacrylate polymerization, resulting in covalent bonding between tissues and the resulting adhesive layer likely responsible for the impressive adhesive strength of cyanoacrylates. Cyanoacrylate glue is superior to the rival adhesives in terms of strong wet adhesion, low cost, rapid curing, inherent bactericidal properties, and good cosmetic outcome.25,44−47 These properties make cyanoacrylates an attractive choice for wound closure or as hemostatic agents to be used in adjunct with the traditional closure techniques. However, the rapid polymerization of cyanoacrylate monomers is associated with significant heat dissipation at the application site resulting in formation of a hard and brittle film. The properties of cyanoacrylate glue can be tuned by changing the chain length of the ester side chain. For example, methyl α-cyanoacrylate elicits a severe inflammatory response D


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based adhesive hydrogels were developed using water-soluble carbodiimide as a cross-linker and/or N-hydroxysuccinimideactivated PLGA. These hydrogels provided strong adhesion within a short gelation time compared to that of the commercial fibrin glue on mouse soft tissue (Figure 4).65−69 In a similar work, Zilberman and co-workers replaced PLGA with an anionic polysaccharide−alginate to develop gelatinbased tissue adhesives cross-linked with carbodiimides and/or NHS-activated route. They also loaded these adhesives with anesthetic, analgesic, antibiotic drugs, hemostatic agents, or bioactive ceramics (Figure 5). Excessive swelling of the

Table 2. Surgical and Medical Applications of Cyanoacrylate Glue specialty general surgery emergency medicine and general practice endoscopy

past, present, and potential applications surgical wound repair control of hemorrhage traumatic wound repair

control of variceal bleeding and obliteration of esophagogastric varices ophthalmology temporary repair of corneal perforations arterial surgery arterial anastomoses thoracic surgery closure of pulmonary leaks neurosurgery repair of peripheral nerves microanastomosis of sciatic nerve otological surgery ossicular chain reconstruction interventional radiology and embolotherapy of various vascular cardiology abnormalities including aneurysms pediatrics wound closure in children pediatric endoscopic surgery tissue approximation and hemostasis pharmacotherapeutics drug carriers

Reproduced from ref 59 with permission from John Wiley and Sons. Copyright 2007 Royal Australasian College of Surgeons.

few.59,63,64 The full potential of cyanoacrylate glue cannot be realized in vivo or commercially unless the problem of toxicity is resolved.

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POLYSACCHARIDE, POLYPEPTIDE, OR POLYMERIC ADHESIVES Polysaccharides, polypeptides, or proteins are rich in amine, hydroxyl, or carboxylic acid functionalities. Adhesives developed from these building blocks adhere to amine groups on the tissue surface via covalent interaction by N-hydroxysuccinimide activation or Schiff base formation, Michael addition reaction, biaryl formation, imine formation, or π−π interaction. The following section is focused on gelatin-, chitosan-, dextran-, and alginate-based adhesives. In addition, adhesives developed using isocyanate and acrylate-functionalized polymers and hyperbranched polymers are also discussed. Gelatin is a thermally denatured collagen composed of polypeptides and proteins rich in amines and carboxylic acid groups. The degradability and biocompatibility of gelatin makes it one of the most popular choices for application in tissue adhesives. In the late 90s, Ikada and co-workers performed extensive research to develop gelatin-based bioresorbable adhesive hydrogels. Gelatin−poly(L-glutamic acid) (PLGA)-

Reproduced with permission from ref 73. Copyright 2014, John Wiley and Sons, Ltd.

Figure 5. N-Ethyl-N-(3-(dimethylamino)propyl)carbodiimide (EDC)and N-hydroxysuccinimide (NHS)-activated cross-linking of gelatinand alginate-based tissue adhesives (Gel, gelatin; Al, alginate).

adhesive, cytotoxicity of the cross-linker, burst release of the drugs, and their unpredictable effect on bonding strength limits the application of these adhesives. Addition of bioactive ceramics indeed improved the adhesion strength on soft tissues from 8.4 ± 2.3 to 18.1 ± 4.0 kPa, whereas on hard tissue the adhesion strengths were ∼71.4 ± 28.2 kPa.70−75 Another example of an NHS-based adhesive was given by Taguchi and co-workers. They used an N-hydroxysuccinimide derivative of citric acid to cross-link a collagen-based matrix. This adhesive demonstrated an adhesion strength of 19.9 ± 1.9 kPa on porcine soft tissue. In the following studies, collagen was replaced with human serum albumin (HSA) owing to the

(a) Reprinted from Iwata, H.; Matsuda, S.; Mitsuhashi, K.; Itoh, E.; Ikada, Y., A novel surgical glue composed of gelatin and N-hydroxysuccinimide activated poly(L-glutamic acid) and its gelation with gelatin. Biomaterials, 19, 1869−1876, Copyright 1998, with permission from Elsevier.(b) Reprinted from Otani, Y.; Tabata, Y.; Ikada, Y., Effect of additives on gelation and tissue adhesion of gelatin-poly(L-glutamic acid) mixture. Biomaterials, 19, 2167−2173, Copyright 1998, with permission from Elsevier.

Figure 4. Gelatin−poly(L-glutamic acid) (PLGA)-based hydrogels. (a) Hydrogels based on N-hydroxysuccinimide (NHS)-activated PLGA. (b) Gelatin−PLGA-based hydrogels physically cross-linked by hydrogen bonding at low temperature and chemically cross-linked in the presence of water-soluble carbodiimides (WSC). E


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cross-links with another amine-functionalized moiety to promote cross-linking as well as cohesion. This chemistry eliminates the need to use cross-linking agents, which is beneficial in terms of biocompatibility of the reagents used. On the basis of this concept, Mo and co-workers developed an adhesive hydrogel by cross-linking aldehyde-functionalized alginate with amine-functionalized gelatin via Schiff base reaction. The maximum adhesion strength on porcine skin was around 11.51 ± 1.3 kPa, slightly lower than the commercial fibrin glue (13.54 ± 3.07 kPa). The degradation products and an increase in the aldehyde content of the hydrogel lowered the cell viability as observed from the MTT assay.85 A dextranbased hydrogel comprised of aminodextran and oxidized dextran (dextran aldehyde) was capable of sealing an incision in the swine uterine horn with maximum burst pressure of 64.6 ± 9.3 mmHg in addition to being noncytotoxic, low swelling, and degradable.86 Ikada and co-workers compared the bonding and sealing ability of three different biodegradable adhesives comprised of: modified gelatin + aldehyde dextran (gel-dext), modified gelatin + oxidized (aldehyde) hydroxyethyl starch (gel-HES), and chitosan + modified dextran (chit-dext). Chitdext gels owing to high stiffness showed lower bonding strength (130 gf/cm2 (12.75 kPa)) and sealing ability compared to geldext and gel-HES (bonding strengths: 210 gf/cm2 (20.60 kPa) and 227 gf/cm2 (22.27 kPa), respectively).87 An acidic solution of non-cross-linked collagen or gelatin modified by oxidative cleavage forms a bioresorbable tissue adhesive when neutralized to a pH between 6 and 10. The collagen/gelatin reacts with the proteins in the tissue and undergoes rapid cross-linking when neutralized, forming adhesive bonds comparable to fibrin glue.88 In another study, a combination of periodate-oxidized dextran (dextran dialdehyde, DDA) and chitosan hydrochloride (amine-functionalized chitosan) resulted in rapid in situ gelation forming a Schiff base product. Jayakrishnan and coworkers elaborately demonstrated the efficacy of the gel as a tissue adhesive, hemostat, and drug delivery medium in addition to its cytocompatibility and degradability.89 Artzi and co-workers studied changes in adhesion of dextran with a polyamidoamine dendrimer upon varying the tissue chemistry. However, the adhesion strength of this adhesive was low compared to those of other dextran-based adhesives.90 Bioadhesives were also synthesized by reaction between aldehyde dextran and ε-poly(L-lysine) (LYDEX). They were shown to have high bonding strength and degradability along with low cytotoxicity. Hyon and co-workers successfully applied these adhesives in lung surgeries, laparoscopic partial nephrectomy, ocular reconstruction, cartilage or bone regeneration, cardiovascular surgery, and as a drug or gene carrier agent.91−103 High aldehyde concentrations are detrimental to tissue and may result in tissue inflammation or tissue necrosis. It is therefore necessary to monitor the aldehyde concentration in the adhesives to avoid cytocompatibility issues. The bonding strengths could also be improved by incorporation of a vinyl or photo-cross-linkable group in the adhesive structure. When irradiated with light, these photocross-linkable groups undergo cross-linking, increasing the cohesive as well as the adhesive strength of the network. Matsuda and co-workers developed a photocurable gel using vinylated protein, polysaccharide, or a combination of both, which adhered to tissue upon photo-irradiation. In vivo studies using styrene-derivatized gelatin showed satisfactory wound closure on dog thoracic aorta and rat liver tissue upon irradiation with visible light. The maximum adhesive strength in

spherical structure and negative charges, which were assumed to promote bonding. The maximum bonding strength obtained on collagen casings (model soft tissue substrate) was 760 g/cm2 (74.56 kPa), comparable to cyanoacrylate glue and ∼9 times stronger than fibrin glue. The adhesive showed complete wound closure with gradual degradation and mild inflammatory response on mouse skin. In a subsequent study, an adhesive based on HSA crosslinked with organic acid based cross-linker−disuccinimidyl tartarate (DST) reached maximum adhesion strength of 489.14 ± 93.06 kPa (on collagen casing) within 5 min of mixing; however, it resulted in mild tissue inflammation. When HSA was replaced with cholesteryl-functionalized gelatin (Figure 6),

Reprinted from Matsuda, M.; Ueno, M.; Endo, Y.; Inoue, M.; Sasaki, M.; Taguchi, T., Enhanced-tissue penetration−induced high bonding strength of a novel tissue adhesive composed of cholesteryl groupmodified gelatin and disuccinimidyl tartarate, Colloids Surf., B, 91, 48−56, Copyright 2012, with permission from Elsevier.

Figure 6. Bonding mechanism of 70% CholGltn/Gltn-DST composition on porcine arterial media. Improved tissue penetration and hydrophobic interactions between the cholesteryl groups in addition to covalent cross-linking by DST contribute to increased adhesion strength.

improved bonding and peeling strengths were obtained due to improved tissue penetration by anchoring to the lipid bilayer of cell membranes. The hydrophobic interactions between cholesteryl groups in addition to covalent cross-linking by DST increased the cohesive strength of this adhesive. In the following study, alkaline-treated gelatin was modified with hexanoyl, decanoyl, and stearoyl chloride to introduce hydrophobic groups in the polymers. The bonding strength on porcine intestinal tissue showed an improvement for shorter alkyl chain modification and the burst strength improved with the degree of functionalization of the hydrophobic moiety.76−83 In another work, Nagatomi and co-workers developed acrylate and N-hydroxysuccinimide (NHS) bifunctional tetronic hydrogels for soft tissue wound closure. The adhesive showed maximum adhesion strength of 74 kPa with low swelling but poor burst pressure compared to the native tissue.84 The cross-linker and reaction byproducts (urea) in these studies were cytotoxic and very likely to cause local tissue inflammation or tissue necrosis. For the cytotoxicity concern of urea byproducts to be overcome, gelatin−polysaccharide-based hydrogels by Schiff base formation were subsequently developed by other groups. Schiff base reaction occurs under mild conditions between an amine and an aldehyde group. An aldehyde-functionalized moiety reacts with the amine groups on the tissue surface to promote adhesion and simultaneously F


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Reproduced from ref 112 with permission of The Royal Society of Chemistry. http://dx.doi.org/10.1039/c3tb00578j.

Figure 7. Synthesis of gelatin-hydroxyphenyl propionic acid (GH) and gelatin-hydroxyphenyl propionic acid−tyramine (GHT) conjugates. Enzymatic or peroxide cross-linking yield gels with 2−3-times higher adhesion strength compared to that of fibrin glue.

Reprinted from Thirupati Kumara Raja, S.; Thiruselvi, T.; Sailakshmi, G.; Ganesh, S.; Gnanamani, A. Rejoining of cut wounds by engineered gelatin-keratin glue. Biochim. Biophys. Acta, Gen. Subj., 1830, 4030−4039, Copyright 2013, with permission from Elsevier.

Figure 8. Mechanism of periodate-based-based cross-linking between caffeic acid-functionalized gelatin and thiol-functionalized keratin.

vitro was 157.1 ± 14.6 g/cm2 (15.4 ± 1.4 kPa), and the adhesive was also capable of localized drug delivery at the tumor site following surgery. The radicals were generated during the cross-linking reaction using a radical initiator (camphorquinone), which could possibly result in tissue damage or tissue necrosis.104,105 Photopolymerization of phenolic-derivatized gelatin (conversion of lysine residues into tyrosine) yielded a low swelling hydrogel by dityrosine cross-link formation. The hydrogel demonstrated elasticity and withstood internal burst pressure up to 60 mmHg, however the actual adhesion strength was not mentioned.106 A photo-crosslinkable chitosan-based tissue adhesive was developed by Ishihara and co-workers by incorporating photoactive azide groups. Chitosan was also modified using lactose moieties to

impart water solubility. Upon irradiation with UV light, the azide groups were converted to highly reactive nitrene groups, which in turn reacted with the amines in the tissue proteins or chitosan to form azo groups resulting in tissue bonding and cross-linking of the adhesive, respectively. The maximum bonding strength on ham slices was 43 g/cm2 (4.2 kPa), and the maximum burst pressure up to 225 ± 25 mmHg was capable of sealing a pinhole in the thoracic aorta. The use of UV light employed for cross-linking and the presence of toxic functional groups (azide, nitrene, and azo groups) are detrimental to the underlying tissue, which calls for extensive biocompatibility evaluation.107−110 A bone adhesive with sustained adhesion under wet conditions was developed from a combination of photocurable poly(ethylene glycol) dimethaG


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Reprinted (adapted) from Nie, W.; Yuan, X.; Zhao, J.; Zhou, Y.; Bao, H. Rapidly in situ forming chitosan/ε-polylysine hydrogels for adhesive sealants and hemostatic materials. Carbohydr. Polym., 96, 342−348, Copyright 2013, with permission from Elsevier.

Figure 9. Thiolated chitosan (CSS) and maleimide-functionalized ε-polylysine (EPLM) undergo rapid cross-linking by Michael addition reaction.

higher than that of the fibrin glue. These in situ enzymatic cross-linking hydrogels degraded rapidly in vitro, which left insufficient time for tissue regeneration and could adversely affect wound healing (Figure 7).112 A protein-based adhesive synthesized by the cross-linking of caffeic acid-functionalized gelatin with thiol-functionalized keratin was attempted by Gnanamani and co-workers. Periodate oxidation triggers intra- and intermolecular cross-linking of the adhesive via Michael addition, biaryl formation, disulfide bridging, and imine formation (Figure 8). This adhesive demonstrated maximum adhesion strength of 16.6 kPa in vitro and 1.67 MPa in vivo with gradual degradation, excellent biocompatibility, and rapid wound healing.113 Bao and co-workers synthesized an in situ-forming polysaccharide/polypeptide hydrogel by simple Michael addition reaction between thiolated chitosan (CSS) and maleimide-functionalized ε-polylysine (Figure 9). The adhesive undergoes fairly rapid cross-linking with the least gelation time of 15 s under physiological conditions with excellent cytocompatibility and hemostatic property. The maximum adhesion strength measured was 87.5 kPa on simulated living tissue.114 Webb and co-workers developed acrylate end-functionalized poloxamine adhesives as a low swelling alternative to the commonly used PEG adhesives. These adhesives gelled by reverse thermal gelation of thermosensitive poloxamine

crylate (PEGDMA) and isocyanate end-capped hydrophilic statistical copolymer of poly(ethylene oxide) and poly(propylene oxide) (NCO-sP(EO-stat-PO)). Ceramic fillers like gypsum (CaSO4·2H2O), newberyite (MgHPO4·3H2O), and struvite (MgNH4PH4·6H2O) were added to improve the consistency of the final adhesive for easy handling, strong mechanical properties, and improved adhesion and porosity to impart degradability to the matrix in addition to cellular infiltration and angiogenesis. Addition of NCO-sP(EO-statPO) did not have a positive effect on the 3 point bending strength of the adhesives but preserved the shear adhesion strength in wet conditions.111 Photocurable adhesives occasionally require UV irradiation, which is harmful to the neighboring healthy tissue. Furthermore, radical initiators or reactive species generated during irradiation could be detrimental to the underlying tissue resulting in tissue necrosis. In certain cases, if the light does not penetrate the depth of the adhesive, only surface curing will occur, resulting in a stiff, cured top layer over a soft un-cross-linked adhesive layer with poor adherence to the tissue surface. Adhesives were also functionalized with phenolic and/or thiol groups to promote tissue interaction as well as intermolecular cross-linking using oxidizing agents like periodate, ferric ions, or enzymatic oxidation. Hydrogels based on hydroxyphenyl propionic acid and tyramine-conjugated gelatin cross-linked within 30 s with adhesion strengths considerably H


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hyperbranched polymers limits chain entanglement, which allows more functional groups to interact with the surface. The elastic modulus and adhesion strength was higher at physiological temperature than at room temperature. With wet adhesion strengths of 138 kPa, degradability, and noncytotoxicity, this adhesive has potential for use in biomedical applications.121 Other studies have also focused on isocyanate- or urethanefunctionalization for rapid cross-linking. Yang and co-workers developed tissue adhesives based on urethane-modified dextrans. This adhesive displayed an adhesion strength of 2.99 ± 0.12 MPa, which was much higher than that of commercial Tisseel glue (0.05 MPa). Urethane-functionalized oxidized dextran when cross-linked with gelatin demonstrated an increase in the adhesion strength up to 4.16 ± 0.72 MPa. However, hydrogel swelling was an unavoidable shortcoming of this adhesive.122,123 Matsuda and co-workers developed an isocyanate end-capped polyether copolymer of poly(ethylene glycol) and poly(propylene glycol) with impressive in vivo adhesion. Fluorinated aliphatic isocyanate demonstrated lower cytotoxicity compared to its aromatic counterparts, however, with slow bioresorption.124 To build upon Matsuda’s work, Ikada and co-workers synthesized isocyanate end-functionalized polyester copolymers, which cured in the presence of water. These adhesives demonstrated faster degradation compared to the polyether copolymers but also resulted in acute tissue inflammation at the application site.125 Bochyńska and coworkers developed isocyanate end-functionalized water curable block copolymer adhesives from trimethylene carbonate for meniscal tissue repair. The isocyanate groups react with amines in the tissue protein resulting in covalent attachment to tissue surface. PEG and trimethylolpropane ethoxylate were used as initiators to synthesize linear and 3-armed adhesives, respectively (Figure 12). The branched adhesive demonstrated higher adhesion strength (0.68 MPa) compared to those of the linear counterpart (0.35 MPa) and commercial Dermabond (0.4 MPa). In the following work, hyperbranched polymers were synthesized by a polycondensation reaction between citric acid and the linear copolymers to increase the number of reactive groups. However, compared to the previous work, no significant increase in the number of isocyanate groups was observed, and the adhesion strengths were much lower (20−80 kPa) in addition to longer curing times. Addition of an amine cross-linker (spermidine) or catalyst (2,2-dimorpholinodiethyl ether (DMDEE) and 1,4-diazabicyclo [2.2.2] octane (DABCO) reduced the curing time to 2 h from approximately 8−24 h. Usage of DABCO resulted in the strongest adhesion (64.4 ± 14.3 kPa) in the shortest time (2 h). The presence of collagenase did not seem to improve the adhesion strength. The cell viabilities on these adhesives were rather low, and in vivo cytocompatibility evaluation is necessary to deem these adhesives safe for internal use.126−129 A biodegradable tissue adhesive based on isocyanatefunctionalized 1,2-ethylene glycol bis(dilactic acid) (ELANCO) and a biopolymer consisting of free amine and/or hydroxyl groups as a chain elongation agent was developed by Schmitz and co-workers. Incorporation of chitosan chloride as a chain extender significantly improved the adhesion strength compared to that of fibrin glue on bovine muscle tissue.130,131 Ates and co-workers developed polyurethane-based adhesives for application in soft tissue adhesion. The authors have reported the effect of incorporation of chlorogenic acid and xylose in these polyurethanes. The xylose incorporated

copolymers and by Michael-type addition with a thiolcontaining cross-linker. The adhesives showed stronger adhesion (25 kPa) compared to that of 4-arm PEG on rat skin; however, the cell adhesion on native hydrogels was poor in the absence of a cell interaction ligand like fibronectin and RGD peptide.115 Hwang and co-workers synthesized a chitin nanofiber/gallic acid-based crystalline tunicate mimetic tissue adhesive. This hydrogel showed the strongest adhesion under wet conditions when cross-linked with periodate (∼215 kPa). Ferric ion cross-linking (∼98 kPa) demonstrated lower adhesion strengths than periodate but had the advantage of a possible self-healing mechanism (Figure 10).116

Reprinted from Oh, D. X.; Kim, S.; Lee, D.; Hwang, D. S. Tunicate-mimetic nanofibrous hydrogel adhesive with improved wet adhesion. Acta Biomater., 20, 104−112, Copyright 2015, with permission from Elsevier.

Figure 10. Gallic acid-conjugated chitin nanofiber precursors crosslink by two methods: Periodate (NaIO4)-induced covalent hydrogel formation (color change from white to brown) and FeCl3-induced noncovalent, iron complex hydrogel formation (color change from white to red via purple).

Bianco-Peled and co-workers developed a biomimetic tissue adhesive inspired by wet adhesion of the brown alga Fucus serratus. The adhesive, composed of phloroglucinol, alginate, and calcium ions, was capable of adhering to a variety of substrates including porcine tissue.117−119 Homo- and copolymers of polypeptides cross-linked with transglutaminase have also demonstrated adhesive characteristics.120 Lei and coworkers studied the effect of topological structures, side group chemistry, temperature variation, hydrophobicity or hydrophilicity, and cure time of hyperbranched poly(amino acid)s. The thermoresponsive poly(amino acid)s based on a combination of dopamine, arginine, cysteine, and lysine acrylamide (Figure 11) were cross-linked by catechol oxidation or Michael addition reactions. The hyperbranched polymers had higher wet adhesion strength compared to their linear counterparts and increasing the functionality on the copolymer also increased its adhesion strength. The globular structure of I


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Reproduced from ref 121 with permission of The Royal Society of Chemistry. http://dx.doi.org/10.1039/c5py01844g.

Figure 11. Thermoresponsive hyperbranched poly(amino acid)s functionalized with DOPA, arginine, cysteine, and/or lysine-acrylamide demonstrate strong wet adhesion.

Reproduced from ref 126 with permission from John Wiley and Sons. Copyright 2013 Wiley-VCH Verlag GmbH & Co. KGaA, Weinheim, Germany.

Figure 12. Hexamethylene diisocyanate-modified trimethylene carbonate (TMC) and trimethylolpropane ethoxylate (TMPE) hydrogels. (a) PEG(TMCm-HDI)2 and (b) TMPE-(TMCm-HDI)3 (m = 1 or 2).

J


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Reproduced from ref 145 with permission of The Royal Society of Chemistry. http://dx.doi.org/10.1039/c5tb00753d.

Figure 13. Synthetic protocol of s-triazine-based hyperbranched epoxy resin. L, linear unit; D, dendritic unit; T, terminal unit.

polymers showed stronger adhesion (94.0 ± 2.8 kPa) compared to the chlorogenic acid-incorporated polymers (40.09 ± 5.08 kPa) on soft tissue.132,133 A bone adhesive developed by reinforcing a polyurethane matrix with nanocrystals of hydroxyapatite showed improved adhesive performance on wet bone in comparison to the commercial bone cement.4 Tissue adhesives employing different techniques like laser welding, layer-by-layer assembly, and temperature-dependent hardening have also been developed. Thoroughly dried, deacetylated chitosan strips applied as tissue adhesives using laser welding demonstrated an adhesion strength of 14.7 ± 4.7 kPa on sheep intestinal tissue presumably by diffusion in the tissue. Modifications like the addition of genipin as a crosslinker, integration with small intestine submucosa, and addition of Rose Bengal dye were also attempted by the authors but did not bring about an appreciable change in adhesion strength. In addition to sealing intestinal pinhole defects, the adhesive was also suitable for sealing corneal defects and nerve anastomoses.134−139 In another study, the layer-by-layer (LbL) assembly technique was used to fabricate flexible, mechanically robust free-standing films by interdiffusion of poly(allylamine hydrochloride)-dextran and hyaluronic acid by Sun and coworkers. The LbL-assembled polyelectrolytes interact with silanol groups on the glass surface by hydrogen bonding and electrostatic interactions. Drug-loaded films demonstrated a burst release with ∼67% of the drug released within 1 h and strong adhesion strength of 2.69 ± 0.62 MPa on bovine periosteum. Adhesion studies on animal tissue are required for further application of these adhesives.140 Biomaterials like lactide and caprolactone are especially attractive because of their biocompatibility and degradable nature, which makes them one of the best options for adhesive building blocks. Cohn and Lando developed branched oligomers of lactide and caprolactone extending from a core. The combination of lactide and caprolactone imparted temperature-dependent changes in the rheological properties such that the polymer possessed low viscosity at the application temperature but hardened at the body temperature after application. The adhesive failure strengths of these block copolymers were higher (∼6 N/cm) than those of the commercial cyanoacrylate glues (∼1 N/cm) on polyamide substrate.141 Another polysaccharide-based biocompatible

adhesive composed of chondroitin sulfate (CS) cross-linked with bone marrow (BM) showed improved adhesion and biocompatibility compared to the PEG counterpart. The presence of bone marrow supplied growth factors at the wound site aided in bone or meniscal regrowth; however, it also increased the risk of virus transmission. A trade-off between CS and BM concentration resulted in either improved adhesion or cell infiltration along with tissue growth.142−144 A strong surgical sealant with low swelling, nontoxic degradation products, and inherent antimicrobial properties was developed from s-triazine-based hyperbranched epoxy with poly(amido amine) hardener (Figure 13). This sealant demonstrated strong mechanical properties, slow degradation, and minimal/no irritation to the rat tissue, but longer curing times could limit its application for minor surgeries involving anesthesia.145

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POLY(ETHYLENE GLYCOL) (PEG)-BASED HYDROGEL ADHESIVES PEG is a hydrophilic, biocompatible polymer widely used as a biomaterial in tissue engineering. However, PEG lacks biodegradability and is therefore often modified with degradable functionalities or copolymerized with degradable polymers. PEG-based tissue adhesives have gained popularity because of the ease of modification/functionalization, bioconjugation, drug delivery, nonimmunogenicity,146 and water solubility. In previous studies, modified PEG was combined with polysaccharides or protein based adhesives which have been described in the previous section the previous section. For example, Matsuda and co-workers developed a UV and visible light-curable elastomeric and degradable adhesive gel derived from photoreactive gelatin (functionalized with benzophenone or a xanthene dye) and poly(ethylene glycol) diacrylate (PEGDA). These gels demonstrated an adhesion strength of 12 kPa and a burst pressure of 150.7 ± 34.4 mmHg with satisfactory hemostatic and anastomotic performance.147,148 However, UV light is detrimental to tissue and cannot penetrate the thick adhesive layer resulting in only surface cross-linking. To overcome these issues, the authors further modified this gel structure to develop a visible light-curable tissue adhesive composed of styrenated gelatin, PEGDA, and carboxylated camphorquinone. A formulation of 35 wt% styrenated gelatin, 5 wt% PEGDA, and 0.05 wt% carboxylated camphorquinone demonstrated an adhesive strength of ∼140 K


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Reproduced from ref 157 with permission from John Wiley and Sons. Copyright 2006 Wiley Periodicals, Inc.

Figure 14. Aldehyde-terminated PEG−PLA block copolymeric micelles cross-link with polyamine and adhere to the tissue surface by Schiff base reaction.

mori (silkworm) with thiol and maleimide-functionalized 4-arm PEG to integrate mechanical strength and biocompatibility, respectively. The adhesive system showed an increase in adhesive strength (∼50% increase) compared with that of commercial sealant CoSeal upon increasing silk content.164 In another study, the silk fibroin from Bombyx mori was conjugated with PEG for water solubility and dopamine to develop water-soluble adhesives. The adhesion strengths of these conjugates were not only enhanced by cross-linking via oxidation but also by silk β-sheet structure, which reinforced the material strength.165 Chenault developed a degradable adhesive hydrogel based on polyglycerol aldehyde and waterdispersible multiarm amine with burst pressures ranging between 5.5 and 27.6 kPa.166 A two-component biocompatible and degradable hydrogel made from PEG-dimethacrylate and thiolated chitosan by Michael reaction was developed by ElNewehy and co-workers. The lap shear strength on rat skin varied between 16.32 ± 0.76 and 48.73 ± 0.56 kPa.167 A degradable, biocompatible, in situ gelling adhesive was designed by Yuan and co-workers for nerve anastomosis in particular. Thiolated chitosan (CSS) cross-links with dual-functionalized (catechol and maleimide) polylysine via Michael addition reaction (Figure 16), resulting in an adhesion strength capable of tolerating a 0.185 N load and effective nerve regeneration. In a preceding study when only maleimide was used as the crosslinker, the maximum adhesion strength obtained was 148 kPa L


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Figure 15. Cross-linking reaction between aldehyde-functionalized dextran and amine-functionalized PEG by Schiff base formation to form an adhesive hydrogel.

Reproduced with permission from ref 168. Copyright 2016 American Chemical Society.

Figure 16. Thiolated chitosan (CSS) and catechol-maleimide-functionalized polylysine cross-link within 10 s via Michael addition. Interactions like hydrogen bonding and electrostatic, π−π, and π−cation interactions contribute to the strong bulk cohesive force.

formed by simply mixing the two components at acidic pH, which makes it suitable for large-scale production. The adhesion strength of TAPE was dependent on the number of PEG arms and the end functional group (PEG-NH2 > PEG− OH > PEG-SH). The strong adhesion was attributed to hydrogen bonding present throughout the network with the highest adhesion strength of 0.18 MPa when PEG-NH2 was used.171

within a gelling time of 5 s. Catechol incorporation did not have a significant influence on gelling time, but it did contribute to a slight improvement in adhesion. These gels undergo considerable swelling, which could potentially impact the wound healing process.168,169 Adhesive nanocomposites have the advantage of improved mechanical properties. To this end, Lee and co-workers developed an injectable nanocomposite adhesive from dopamine-modified 4-arm PEG and Laponite (nanosilicate). Laponite incorporation not only improved the mechanical properties, curing rate, and adhesion strength (7.9 ± 1.8 kPa) of the composite but also demonstrated increased cell viability with minimal tissue inflammation (Figure 17).170 In another study, a degradable hemostatic tissue adhesive called TAPE (tannic acid and poly(ethylene glycol)) was synthesized using tannic acid and PEG, which showed impressive adhesion under wet conditions. The adhesive is

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BIOMIMETIC TISSUE ADHESIVES Mussel-Inspired Adhesives. The adhesive property of mussels, specifically Mytillus edulis, has been studied for decades and is of particular interest because of its ability to attach to virtually any surface via byssal threads secreted from the foot.172−174 This adhesive is reversible and capable of withstanding strong water currents as well as fluctuations in temperature and salinity. The mussel byssus usually consists of M


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Figure 17. Laponite-incorporated, dopamine-modified, 4-arm PEG showed improved adhesion and cohesion by (A) polymerization and di-DOPA cross-link formation by periodate oxidation, (B) interfacial cross-linking with amine groups on the tissue surface, and (C) reversible bonding with Laponite.

polypeptides, polydipeptide, polytripeptide,192 polyoctapeptide,193 and polydecapeptide,194 as mimics of the mussel adhesive protein and studied their bonding strengths. An important conclusion from these studies was that the mussel adhesive mimics were capable of bonding various materials and the bonding strength increased upon addition of a cross-linker like tyrosinase, hydrogen peroxide, or basic aqueous solution. The bonding strengths were also found to increase with increasing DOPA content, copolymer solution concentration, copolymer molecular weight, and curing temperatures.195 Out of all the amino acids present in the mussel adhesive proteins, lysine and tyrosine were identified to be of utmost importance for bonding and cross-linking characteristics.192,195,196 The authors also studied the in vivo adhesive properties of polytripeptide−poly(Gly-Tyr-Lys) (exhibited highest adhesion strength on pigskin in vitro) on white pigs and compared the results with a commercial Tisseel glue. No acute inflammation or tissue reactivity was observed in the area near the incision site. Histology results suggested that the poly(Gly-Tyr-Lys)tyrosinase system had less immune reactivity (58 ± 9%) than Tisseel.196 These studies on the adhesion properties of the mussel adhesive protein (MAP) and their synthetic polypeptide mimics led to a plethora of studies on synthetic polymeric adhesive mimics, which will be discussed in detail in the following section. Yamada and co-workers evaluated the efficacy of a deacetylated chitosan, dopamine, and tyrosinase solution as a water-resistant adhesive. Tyrosinase oxidizes dopamine into oquinone, which in turn triggered cross-linking within chitosan as well as promoted adhesion on the glass surface. The maximum bonding strength under dry conditions was 400 kPa and under water was ∼450 kPa. Their studies also concluded that the water-resistant adhesion was not specific to the enzyme-catalyzed oxidation because glutaraldehyde cross-linking provided similar results with slightly lower bonding

four main components: acid mucopolysaccharides acting as a primer, adhesive proteins consisting mainly of polyphenolic proteins rich in 3,4-dihydroxyphenylalnine (L-DOPA) and lysine, fibrous proteins that act as an attachment thread between mussel and the substrate, and finally, polyphenoloxidase to promote intermolecular cross-linking.175,176 Preliminary immunological studies showed that the mussel adhesive proteins are poor antigens and have great potential to be used for biomedical purposes specifically for biological tissue adhesives.177 Early efforts of mimicking mussel adhesives involved developing a protein as a nonspecific adhesive for cell attachment and growth178 followed by studies confirming the suitability of mussel adhesive proteins for cell viability, attachment, growth, and proliferation.179 The adhesive properties of mussel adhesive protein (MAP) extracted from the blue mussel Mytillus edulis showed satisfactory bonding on various substrates like stainless steel, pig duodenal mucosa, porcine small intestinal submucosa, and porcine skin.180−183 Other efforts involved the use of recombinant DNA technology, peptide synthesis, fragment condensation technique, and gene cloning to develop synthetic mimics of mussel adhesive.6,176,177 Yamamoto and co-workers synthesized a series of homopolymers and copolymers based on L-DOPA and L-glutamic acid184,185 followed by synthesis of a series of poly(amino acid)s and polypeptides to study their bonding strength on metals in aqueous and organic solvent systems. Poly(Lys)·HBr demonstrated the highest tensile strength of 123 kg/cm2 (12.06 MPa) on iron, and gelatin demonstrated the highest compressive shear strength of 21 kg/cm2 (2.06 kPa) on aluminum oxide substrates in water. For the organic solvent system, poly(DL-methionine) showed the highest tensile strength of 49 kg/cm2 (4.8 kPa) on iron and a compressive strength of 22 kg/cm2 (2.16 kPa) in dichloromethane.186 Inspired from the mussel adhesives, the group also synthesized a series of sequential187−189 or random copolymers190,191 and N


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Reprinted from Fan, C.; Fu, J.; Zhu, W.; Wang, D.-A. A mussel-inspired double-cross-linked tissue adhesive intended for internal medical use. Acta Biomater., 33, 51−63, Copyright 2016, with permission from Elsevier.

Figure 18. Double cross-linked tissue adhesive derived from dopamine-functionalized gelatin macromer cross-linked by Fe3+ (rapid cross-linker) and genipin (GP) (long-term cross-linker).

the moduli of the gels were sufficient for biomedical application. However, the phenolic nature of DOPA served as a radical scavenger, which hindered photo-cross-linking and prolonged the gelation time, resulting in lower gel moduli. Furthermore, UV cross-linked gels outperformed the visible light cross-linked gels, in which case the effect of UV irradiation on tissue is concerning.206 Another study in this group involved a thermosensitive, injectable, DOPA-modified hyaluronic acid (HA-D)/thiol end-capped pluronic F127 (Plu-SH) lightly cross-linked composite gel with a Michael-type catechol-thiol addition reaction. These gels show rapid, reversible sol−gel transition at physiological temperature and enhanced adhesion to tissue. In vivo and in vitro studies showed that the gels form robust structure and strong adhesion to the neighboring tissue owing to the unreacted and oxidized catechol groups.207 Lee et al. synthesized a series of DOPA-modified poly(ethylene glycol)s (PEG-DOPA) (Figure 19) capable of rapid

strengths. Addition of PEG prolonged the enzyme activity, which decreased the cross-linking time significantly to achieve similar bonding strengths as in the previous studies.197−199 Ziegler and co-workers synthesized a dual component bone adhesive based on deacetylated chitosan and oxidized starch/ dextran, which cross-linked via Schiff base formation. Collagen, a primary component of bone, contains amine groups that interact with aldehyde groups in the adhesive to promote bonding. Inspired by mussel adhesion, the starch/dextran was also functionalized with L-DOPA to incorporate catechol groups to further strengthen cross-linking as well as adhesion. The maximum bonding strength of this glue on bovine cortical bone was ∼0.41 MPa without L-DOPA, and the incorporation of L-DOPA groups did not contribute to an appreciable change in adhesion strength.200 An injectable, thermosensitive tissue adhesive made from catechol-functionalized chitosan and thiolfunctionalized pluronic F127 demonstrated an adhesion strength of 14.98 ± 3.53 kPa on mouse skin. The degradable adhesive possessed good mechanical integrity and sealing properties, whereas the biocompatibility was not evaluated.201,202 A dopamine-functionalized gelatin macromer dual cross-linked with Fe3+ (rapid cross-linker) and genipin (long-term cross-linker) (Figure 18) showed strong adhesion on porcine skin and cartilage along with degradability and tissue compatibility in vivo and in vitro.203 A dopamine-grafted, polysaccharide (alginate and hyaluronic acid)-based degradable membrane was developed as a tissue adhesive by Paoletti and co-workers. This membrane showed enhanced adhesion in vitro and in vivo on pig intestine tissue and was expected to demonstrate improved fibroblast activity.204 Messersmith and co-workers synthesized a range of PEGbased DOPA-functionalized copolymers for tissue adhesive applications while eliminating the use of strong oxidizing agents. Huang et al. synthesized DOPA and DOPA methyl ester (DME)-functionalized PEO-PPO-PEO block copolymers for biomedical applications like tissue adhesives and drug delivery. These polymers were capable of forming hydrogels with thermally triggered self-assembly, which eliminates the need for strong oxidizing agents. The gelling temperature was found to be dependent on the copolymer concentration and molecular weight. Viscometry measurements indicated that DOPA-modified copolymers were significantly more mucoadhesive than the unmodified copolymers.205 In another study, Lee et al. demonstrated the synthesis of N-methacrylated DOPA monomer and its copolymerization with PEG diacrylate to form a hydrogel. The photopolymerization technique again eliminated the use of strong oxidizing agents for gelation, and


Figure 19. Chemical structures of a series of PEG-DOPA adhesives synthesized by Lee et al.

gelation in the presence of cross-linking agents and optimized conditions.208 Burke et al. used these PEG-DOPA4 polymers for tissue adhesive applications. They utilized temperaturesensitive liposomes to sequester DOPA-oxidizing agents like periodate in the PEG-DOPA4 solution. These liposomes were triggered at physiological temperature to release the oxidizing agents resulting in gelation of the polymer solution. The crosslinked polymer demonstrated an adhesion strength of ∼35 ± 12.5 kPa on porcine skin.209 O


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Figure 20. Illustration of the photopolymerization of Y-PPM into an adhesive hydrogel. Y-PPM is the DOPA-functionalized copolymer of methacrylated PEG−PLA block copolymers.

Reproduced with permission from ref 218. Copyright 2011 American Chemical Society. (http://pubs.acs.org/doi/full/10.1021/bm201261d).

Figure 21. Periodate-mediated oxidative cross-linking yields rapid gelation and tissue adhesion of the enzymatically degradable cAAPEG macromonomer. In the presence of enzyme neutrophil elastase, the Ala-Ala dipeptide linker (blue) is cleaved to provide degradation sites. Black arrowheads indicate continuation of the cross-linked hydrogel matrix.

of these adhesives before considering their application as tissue adhesives.212 The in vivo adhesive performance of catechol-derivatized PEG (cPEG) was studied in a murine model of extrahepatic islet transplantation by Brubaker et al. with promising results. The adhesive invoked minimal inflammatory response and maintained an intact interface with the supporting tissue for up to one year. The application of cPEG as sealants for fetal membrane repair was also explored in vitro, ex vivo, and in vivo in a rabbit model. The mussel mimetic cPEG was noncytotoxic, accomplished leak proof closure, and conformed to the shape of the underlying membrane, and the bonding strengths were comparable to that of commercial fibrin glue. Although the degradation properties of this adhesive were not studied, it was an inception for medical application of this biomimetic tissue adhesive.213−217 An enzymatically degradable, mussel-inspired tissue adhesive hydrogel with an enzyme cleavable site was developed by Brubaker et al. A DOPA end-functionalized 4-arm PEG incorporated with an enzyme elastase peptide substrate (Ala-Ala dipeptide) in the backbone of the polymer was designed as a macromonomer (cAAPEG) (Figure 21). The catechol groups undergo intermolecular cross-linking leading to rapid gelation (20−30 s) and adhesion to tissue surfaces under oxidizing conditions. The adhesion strength of this adhesive was 30.4 ± 3.39 kPa, considerably higher than that of commercial fibrin glue. Although the adhesive demonstrated slow degradation in vivo with minimal inflammatory response, there is opportunity to improve the degradation rate by incorporating a longer elastase substrate peptide sequence.218 Following the work of Yu and Deming,195 Yin and coworkers synthesized a degradable copolypeptide of DOPA and L-lysine through ring-opening polymerization of NCA mono-

Monomethoxy-terminated PEG end-functionalized with 1−3 DOPA amino acids (mPEG-DOPA) showed rapid, irreversible adsorption on TiO2 surfaces. The strong adhesion of mPEGDOPA on TiO2 occurred by displacement of the TiO2 surface hydroxyl groups via formation of charge transfer complexes and provided excellent resistance to nonspecific protein adsorption. This is especially important for bone surgeries involving titanium implants, stents, catheters, and intraocular lenses.210 Lee et al. also synthesized a photopolymerizable triblock copolymer functionalized with DOPA. In this work, a glycinefunctionalized methacrylated PEG−PLA monomer (G-PPM) was copolymerized with DOPA-functionalized NCA monomer by ring-opening polymerization or with N-Boc-DOPA by simple carbodiimide coupling to obtain ∼83% DOPA-functionalized triblock copolymer (Figure 20). These water-soluble copolymers were photocured into hydrogels using di-tert-butyl dicarbonate (Boc2O), 2,2′-dimethoxy-2-phenylacetonephenone (DMPA) as the photoinitiator, and 1-vinyl-2-pyrrolidone as the solvent for hydrophobic DMPA. The hydrophilic nature of the polymer ensured that the DOPA groups remained in the aqueous environment without interfering with the cross-linking process, which occurred in the hydrophobic environment. The adhesion studies proved that DOPA-modified gels showed higher Wadh (work of adhesion) compared to their unmodified forms on Ti surfaces. The Wadh was found to decrease upon oxidizing the gels with NaIO4, thus proving their hypothesis that the catecholic form of DOPA is essential for adhesion to metallic surfaces.211 Similarly, DOPA-functionalized diblock (PS−PEO) and triblock (PMMA-PMAA-PMMA) copolymers showed strong underwater adhesion on TiO2 surfaces and porcine skin by membrane inflation method. It is, however, important to determine the biocompatibility and degradability P


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Reprinted by permission from Macmillan Publishers Ltd.: Nat. Nano (ref 224), Copyright 2014.

Figure 22. (a−c) Genetically engineered molecular hybrids of mussel adhesive proteins (Mfps) and amyloid-based adhesive proteins in E. coli (CsgA). (d) Adhesive molecular hybrids self-assemble with β-sheet amyloid protein forming the core and the Mfps flanking the exterior.http://www. nature.com/nnano/index.html.

Reprinted from Brennan, M.J.; Kilbride, B. F.; Wilker, J. J.; Liu, J. C. Biomaterials, 124, 116−125, Copyright 2017, with permission from Elsevier.

Figure 23. Design strategy of DOPA-modified recombinant elastin-like polypeptide (ELP) synthesized by Liu and co-workers.

recombinant expression method to obtain rfp-1 MAP (AKPSYPPTYK) for hydrogel formation by coordination (Fe3+) or covalent cross-linking (NaIO4). The hydrogel system showed maximum adhesion strengths of ∼130 and ∼200 kPa when cross-linked with Fe3+ and NaIO4 respectively. The difficulty in synthesizing the rfps in bulk limit the clinical application of these adhesive hydrogels.221 In another study, they engineered a residue-specific DOPA-incorporated recombinant mussel adhesive protein (dfp-3 and dfp-5) with DOPA content up to 23 mol%. The recombinant protein showed strong dry and underwater adhesion along with significant water resistance.222 They also developed a light-activated, mussel protein-based bioadhesive (LAMBA) hydrogel using a photo-oxidative reaction in the presence of blue light involving Ru(II)bpy32+ as the activator and sodium persulfate (SPS) as the oxidizing agent with recombinant MAP. LAMBA demonstrated strong adhesion to wet porcine skin and also promoted wound healing in addition to wound closure in a rat model.223 Lu and coworkers developed a hybrid molecular adhesive by fusing Mfps

mers. Adhesion strengths were evaluated on steel, aluminum, PS, and glass. The copolypeptide (1:4 DOPA/lysine) formed the strongest bond on steel, which varied with the type of crosslinking agent and increased with the increase in copolypeptide molecular weight and DOPA content. The average strengths on porcine bone and porcine skin under dry conditions were 0.295 and 0.208 MPa, respectively, for a cure time of 12 h. Under wet conditions, the average bonding strength on porcine bone decreased to 0.155 MPa, whereas the adhesion on porcine skin failed.219 Attempts were also made to develop recombinant mussel proteins for application as bioadhesives. Lim and co-workers developed MAP-based encapsulated coacervates as smart tissue adhesives with drug carrier ability. In this study, an adhesive was formed by complex coacervation between cationic recombinant hybrid MAPs (fp-131 or fp-151) and the anionic hyaluronic acid (HA). The bulk adhesive strengths of coacervates were twice as strong compared to the protein itself on aluminum substrates.220 Cha and co-workers used the Q


Review

Biomacromolecules


Figure 24. NHS-thiol condensation-based cross-linking of poly(AA-co-AANHS-co-MDOPA) and thiol-terminated 3-armed poly(ethylene glycol); AA, acrylic acid; AANHS, acrylic acid N-hydroxysuccinimide ester; MDOPA, N-methacryloyl-3,4-dihydroxy-L-phenylalanine.

(a) Reprinted from Mehdizadeh, M.; Weng, H.; Gyawali, D.; Tang, L.; Yang, J. Injectable citrate-based mussel-inspired tissue bioadhesives with high wet strengths for sutureless wound closure. Biomaterials, 33, 7972−7983, Copyright 2012, with permission from Elsevier. (b) Reprinted from Guo, J.; Kim, G. B.; Shan, D.; Kim, J. P.; Hu, J.; Wang, W.; Hamad, F. G.; Qian, G.; Rizk, E. B.; Yang, J. Click chemistry improved wet adhesion strength of mussel-inspired citrate-based antimicrobial bioadhesives. Biomaterials, 112, Copyright 2017, with permission from Elsevier.

Figure 25. (a) iCMBA prepolymer synthesis by polycondensation reaction between citric acid, PEG, and dopamine. b) iCMBAs further modified with azide and alkyne functionalities to facilitate dual cross-linking by catechol oxidation and click reaction.

which can be further explored for dental or orthopedic adhesive application.226 Liu and co-workers reported synthesis of an elastin-like polypeptide (Figure 23) with tunable phase transition in which the tyrosine residues were modified into DOPA for underwater adhesion applications. Importantly, this study showed that DOPA residues did not affect adhesion under dry conditions but contributed significantly toward underwater adhesion when compared to the unmodified protein. The adhesion strength of the DOPA-modified protein was >2 MPa under dry conditions and ∼0.24 MPa in humid conditions. The recombinant protein also demonstrated impressive adhesion strength of ∼3 kPa when tested underwater. The final protein yield (before tyrosinase modification) was around 220 mg/L of culture, which could potentially limit large-scale synthesis of this adhesive.7 A biodegradable mussel-inspired adhesive polymer composed of DOPA, polycaprolactone (PCL), and PEG was synthesized by Lee and co-workers. The copolymer demonstrated adhesive strengths 10-times that of commercial fibrin glue on

found in DOPA from mussel adhesives with the CsgA proteins found in the amyloid-based adhesives in E. coli (monomeric, CsgA-Mfp3; Mfp5-CsgA and copolymer constructs, (CsgAMfp3)-co-(Mfp5-CsgA)) (Figure 22). The molecular hybrid self-assembled in which the β-sheet amyloid protein formed the core, whereas the disordered Mfps were exposed on the exterior. The (CsgA-Mfp3)-co-(Mfp5-CsgA) copolymer demonstrated impressive adhesion energy of 20.9 mJ/m2, which made it a strong competitor for application in medical adhesives.224 In a comparative study of bulk adhesion, the recombinant fp3 protein demonstrated stronger adhesion compared to the recombinant fp-5 protein without an oxidant. However, the recombinant technique suffers from low modification yield when tyrosine is converted to DOPA using tyrosinase.225 Recently, Waite and co-workers extracted the mussel foot protein Mfp-3S, which is capable of liquid−liquid coacervation variable with buffer pH, ionic strength, and temperature. This protein showed strong adsorption on hydroxyapatite surfaces, R


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Biomacromolecules


Figure 26. Nitrodopamine-modified, catechol-functionalized 4-arm PEG adheres by oxidative and metal coordination cross-linking. Nitrodopamine groups are responsible for debonding upon photoirradiation.

commercial hernia repair biologic meshes. The lap shear and burst pressure performance of the adhesive polymer on bovine pericardium also outperformed the commercial fibrin glue. The lap shear adhesive strength of the adhesive-coated biologic scaffolds were 106 ± 22.9 kPa on bovine pericardium and 73.4 ± 24.4 kPa on porcine dermal tissue with possible application for Achilles tendon repair.227,228 Chung and Grubbs synthesized DOPA-functionalized terpolymer adhesives. A terpolymer of acrylic acid (AA), acrylic acid N-hydroxysuccinimide ester (AANHS), and N-methacryloyl-3,4-dihydroxy-L-phenylalanine (MDOPA) (poly(AA0.7-co-AANHS0.15-co-MDOPA0.15)) underwent rapid covalent cross-linking (∼30 s) with a thiolterminated 3-armed PEG (PEG-SH) via NHS-thiol condensation (Figure 24). The adhesion strength of poly(AA0.7-coAANHS0.15-co-MDOPA0.15) cross-linked with PEG-SH demonstrated 450% stronger adhesion compared to that of un-crosslinked poly(AA-co-AANHS), comparable to that of the commercial cyanoacrylate (Super Glue) on wet porcine skin.229 A photo-cross-linkable bioadhesive was developed from photocurable monomer (ethylene glycol acrylate methacrylate dopamine (EGAMA-DOPA)) and UV crosslinking agent (poly(vinyl alcohol) (UV-PVA)) (Figure 25). The adhesion strength of EGAMA-DOPA increased significantly upon addition of UV-PVA; however, the possibility of tissue damage upon UV irradiation still persists.230 Yang and co-workers developed a one-step synthesis of injectable citrate-based mussel-inspired bioadhesives (iCMBAs) composed of citric acid, PEG, and dopamine building blocks for wet tissue adhesion. The adhesive is synthesized by a simple one-step polycondensation reaction between the building blocks (Figure 25) and cross-linked using sodium periodate. Citric acid contributed to degradability while enhancing hemocompatibility and hydrophilicity. The lap shear adhesion strengths varied between 33.4 ± 8.9 and 123.2 ± 13.2 kPa, corresponding to variation in composition, and were at least 2fold stronger than commercial fibrin glue (15.4 ± 2.8 kPa). The

iCMBAs were noncytotoxic in vitro and were successfully applied for wound closure and instant bleeding control with tissue regeneration ability in vivo. Guo et al. further modified these iCMBAs into antibacterial and antifungal iCMBAs (AbAf iCs). Incorporation of 10-undecylenic acid contributed to antifungal properties, whereas silver nitrate (SN) or sodium (meta) periodate (PI) used as cross-linkers contributed to the antibacterial properties of the adhesives. On porcine small intestine submucosa, PI cross-linked adhesives showed higher adhesion strengths than SN cross-linked adhesives, and the adhesion strength in both cases was considerably higher than that of commercial fibrin glue. The AbAf iCs demonstrated fast degradation, excellent bacterial and fungal inhibition performance, and good cytocompatibility in vitro. In the following work, two different prepolymers functionalized with azide and alkyne functionalities were synthesized (Figure 25). This modification facilitated dual cross-linking by periodate-induced oxidation of DOPA and Cu-catalyzed azide−alkyne cycloaddition. The dual cross-linking strategy in addition to alkynefunctionalized gelatin showed marked improvement in cohesion and in turn on adhesive strength (223.11 ± 15.94 kPa) on wet tissue. The adhesive showed impressive antibacterial and antifungal properties along with degradability and cytocompatibility; however, the application of Cu catalyst directly on tissue is concerning, and further studies are necessary to evaluate the long-term effect.231−233 Zhu and coworkers modified a previously synthesized citrate-based adhesive232 by incorporating PEO blocks to impart hydrophilicity. The copolymer of citric acid, PEO, 1,8-octanediol, and dopamine showed water solubility for 0.3:0.7 octanediol:PEO (molar ratio). This copolymer was interesting from a tissue adhesive point of view because of the strong adhesion, degradability, and modulus comparable to soft tissue.234 del Campo and co-workers synthesized a biocompatible underwater self-curing, self-healing, surface reactive, and photodegradable adhesive material based on nitrodopamine S


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Biomacromolecules

(a) Reproduced with permission from ref 238. Copyright 2014 American Chemical Society. (b) Reproduced with permission from ref 239. Copyright 2016 American Chemical Society.

Figure 27. Chemical structure of catechol-modified poly(ester urea). (a) Catechol-modified copolymer based on tyrosine and leucine amino acids. (b) Catechol-modified terpolymer based on serine and leucine amino acid with poly(propylene glycol) (PPG) groups incorporated into the backbone for ethanol solubility.

Fe(acac)3 was used as a cross-linker, however, with poor yields and considerable cytotoxicity.240 Wilker and co-workers developed a degradable poly((3,4-dihydroxymandelic acid)-co(lactic acid)) adhesive with adhesion strengths of 2.6 ± 0.4 MPa under dry conditions and 1.0 ± 0.3 MPa under wet conditions after cross-linking on aluminum substrate. Adhesion studies performed on sanded steel and Teflon resulted in strengths of 1.7 ± 0.5 and 0.32 ± 0.05 MPa, respectively. This adhesive holds great promise for application in tissue adhesion; however, more studies on biocompatibility need to be performed.241 In a series of studies, Wilker and co-workers synthesized 3,4-dihydroxystyrene-based biomimetic polymers and studied the effect of molecular weight, curing conditions, and DOPA content on the adhesion strength.242−244 In a recent study, this biomimetic polymer demonstrated impressive underwater adhesion (∼3 MPa) compared to various commercial adhesives (∼1−0.5 MPa).245 In vitro cell studies have deemed this polymer cytocompatible; however, lack of degradability limits its biological application.246 Wu and coworkers reported a pH, glucose, and dopamine triple responsive, self-healing hydrogel tissue adhesive based on reversible covalent complexation between 4-arm PEG (4-arm PEG-DA) and phenyl boronic acid-functionalized 4-arm PEG (4-arm-PEG-PBA) (Figure 28). The as-formed hydrogel showed adhesion strength of 5.2 ± 0.28 kPa on porcine tissue, which remained nearly unchanged for a self-healed hydrogel (5.07 ± 0.35 kPa).247 A mussel-inspired tissue adhesive based on polyvinylpyrrolidone (PVP) backbone synthesized by Wan and co-workers is particularly noteworthy because of its stronger underwater adhesion compared to dry adhesion. The adhesion strength was strongly dependent on the cross-linker to catechol ratio, polymer molecular weight, and catechol content in the polymer. In this study, a polymer of molecular weight ∼10 kDa, catechol content ∼16 mol%, and 1:1 FeCl3:catechol demonstrated the highest underwater adhesion strength of 1.33 MPa. The authors postulated that the strong adhesion strength is a result of a combination of interaction of catechol and amide

(PEG-ND4) (Figure 26). The hydrogel could be cross-linked by both periodate (covalent) and Fe3+ (metal coordination). The photodegradability of the hydrogel, which was dependent on the light exposure and cross-linker concentration, was confirmed by quartz crystal microbalance (QCM-D). The hydrogel demonstrated strong bonding and light-activated debonding underwater with noncytotoxic properties. The Fe3+ cross-linked PEG-ND4 hydrogels showed complete self-healing behavior. Although the tissue adhesion characteristics of these gels were not studied, it is well-suited for wound closure applications.235 Wang and co-workers synthesized a hyperbranched poly(dopamine-co-acrylate) copolymer from dopamine and a triacrylate monomer. In addition to strong wet adhesion (stronger than commercial fibrin glue) on porcine skin, this hyperbranched copolymer also demonstrated degradation and noncytotoxic properties.236 A photo-cross-linkable terpolymer of dopamine acrylamide (DAM), N-isopropylacrylamide (NIPAAm), and polyethylene glycol-triacrylate (PEG-TA) showed thermoresponsive swelling, strong lap shear adhesion, and moisture resistance on the gelatin surface along with low cytotoxicity on mouse fibroblast (L929) cells.237 Becker and coworkers synthesized a catechol-functionalized poly(ester urea) copolymer for tissue adhesion application (Figure 27a). The copolymer showed adhesive strength (∼9 kPa) comparable to that of commercial fibrin glue on porcine skin with a strong potential for further development in medical adhesives. The polymer was further modified by incorporation of poly(propylene glycol) (PPG) (Figure 27b) to solubilize it in a clinically relevant solvent like ethanol without much effect on the adhesion strength (∼10.6 ± 2.1 kPa) on wet porcine skin.238,239 A degradable, catechol-functionalized polyester-based adhesive was synthesized by Agarwal and co-workers by a radical ring opening copolymerization on glycidyl methacrylate (GMA), oligo(ethylene glycol) methacrylate (OEGMA), and 2-methylene-1,3-dioxepane (MDO). The adhesive demonstrated the strongest adhesion strength of 13.13 ± 1.74 kPa when T


Review

Biomacromolecules

under humid conditions or underwater. Waite and co-workers have reasoned that water tends to form a weak boundary layer at the interface, crazes into the interface, triggers hydrolysis resulting in swelling or plasticizing the adhesive, which eventually causes adhesive failure.250 Mussels have figured out a way to form tenacious bonds with any substrate irrespective of the presence of water, but the synthetic mimics discussed here are far from the actual mussel adhesive in terms of adhesive performance. Furthermore, the use of UV irradiation, photoinitiators, and strong oxidizing agents could damage the healthy cells around the application site and are therefore currently only suitable for topical applications. A compilation of mussel-inspired mimics, their cross-linking conditions, test substrates, and the maximum reported adhesion strengths are presented in Table 3.

Reproduced from ref 247 with permission of The Royal Society of Chemistry. http://dx.doi.org/10.1039/ c7py00519a.

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Figure 28. Hydrogel formation or cross-linking between dopaminefunctionalized 4-arm PEG (4-arm PEG-DA) and phenyl boronic acidfunctionalized 4-arm PEG (4-arm-PEG-PBA).

GECKO-INSPIRED TISSUE ADHESIVES In one of the first studies, Karp and co-workers designed a nanopatterned gecko-inspired biocompatible and biodegradable elastomeric tissue adhesive by using poly(glycerol-co-sebacate acrylate) (PGSA) (Figure 30). The adhesion strengths on porcine intestinal tissue were dependent on nanopillar parameters, the polymer composition, and oxidized dextran coating thickness. The oxidized dextran promotes adhesion with tissue by Schiff base formation and at the same time forms hemiacetal with hydroxyl groups from the glycerol subunit of PGSA. The maximum adhesion strength on porcine intestinal tissue was ∼4.8 × 104 Pa.251

groups with the glass surface. Furthermore, the compatibility of the PVP backbone with water allows diffusion of water molecules, facilitating penetration of FeCl3 and resulting in complexation. The strong adhesion characteristics, however, need to be supported with studies on biological substrates for further application in medical adhesives.248 Lei and co-workers synthesized thermoresponsive polypeptide-pluronic-polypeptide triblock copolymers (Figure 29) and studied their adhesion in wet environments and hemostatic ability. The block copolymers were functionalized with different functionalities like catechol, guanidyl, sulfhydryl, and acryloyl to study their effect on adhesion strength. The copolymers were biocompatible, biodegradable, and showed impressive adhesion under wet and humid conditions on porcine skin and bone in addition to hemostatic properties and improved in vivo bone healing.249 A number of groups have adopted different strategies like extraction of mussel proteins, recombinant hybrids of proteins, or DOPA/catechol functionalization of polysaccharides and synthetic polymers to develop mimics of mussel adhesive. These adhesives work well under dry conditions; however, a significant drop in adhesive strength is generally observed

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SANDCASTLE WORM-INSPIRED TISSUE ADHESIVES An acrylate-based mimic of sandcastle worm glue was developed by Stewart and co-workers based on the principle of coacervation, electrostatic interaction, as well as oxidative curing. The adhesion strength on wet porcine bone was 1/3 of the strength of natural glue and ∼37% of the cyanoacrylate glue (control). In a continuation of this study, a two component mimic consisting of poly(MAEP85-dopamide15) and aminemodified gelatin with divalent cations (Ca2+ and Mg2+) showed temperature-dependent coacervation, which could be altered by


Figure 29. Structures of pluronic L-31-poly[(DOPA)-co-(Arg-co-Cys)] (PPDAC) and pluronic L-31-poly[(DOPA)-co-(Arg-co-Ac-Lys)] (PPDAL) adhesives. Different functional groups result in different interactions: (a, b) covalent interaction of catechol groups with tissue surface, (c) di-DOPA cross-linking and DOPA polymerization, (d) electrostatic interaction between guadinium ions (Gu+) and oxoanions on the tissue, (e) disulfide crosslink formation, (f) covalent interaction between catechol and sulfhydryl groups, and (g) thiol−ene click reaction. U


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Biomacromolecules

Table 3. Polymeric Mussel-Inspired Adhesives with Their Crosslinking Conditions and Maximum Reported Adhesion Strengths maximum adhesion strength adhesive MAPa (130 kDa) MAP MAP copoly(Tyr1 Lysx)b

adhesion test surface tensiometer end-to-end lap-shear

test substrate porcine duodenal mucosa porcine skin porcine small intestine submucosa iron

cross-linker

V5+ (250 mM)

cure condition and time

poly(Lys·HBr4-DOPA1)d deacetylated chitosan, dopamine PEG-DOPA4e HA/pluronic hydrogelf catechol-Ala-Ala-PEG (cAAPEG)g poly((Lys.HBr)x-(DOPA)y)

deacetylated chitosan; oxidized and DOPA-functionalized dextran mfp-131 + HA mfp-151 + HA PEG-dopamine-PCLh

humid, 48 h 1h

0.95 ± 0.19 MPa 462 ± 46 kPa

23 °C, 3 days, 60% RH

shear adhesive test tensile bond strength lap shear

porcine skin

tyrosinase

37 °C, 30 min

aluminum

tyrosinase

35 °C, 1 day

4.7 MPa

glass slides

tyrosinase

lap shear tensile bonding strength lap shear

porcine skin mouse skin

periodate

in air, 24 h underwater, 24 h 37 °C, 24 h RT, 5 min

400 kPa ∼450 kPa 35.1 ± 12.5 kPa 7.18 ± 0.93 kPa

RT, 2 h

30.4 ± 3.39 kPa

dry, 25 °C, 12 h

0.21 0.29 0.15 0.31

tyrosinase

lap-shear

end-to-end

decellularized porcine dermis porcine skin porcine bone porcine bone bovine cortical bone

ferric citrate

wet, 37 °C, 12 h wet, 37 °C, 3 h

aluminum

lap shear

bovine pericardium

NaIO4

(1) RT, 2 h, (2) PBS, 37 °C, 1 h

burst test lap shear

porcine skin

thiol-PEG

RT, 10 min

lap shear

porcine SIS

sodium (meta) periodate silver nitrate

humid chamber, 2 h

iCMBAj

lap shear lap shear

EGAMA-DOPA Mfp-3S

lap shear surface forces apparatus (SFA) lap shear

sodium (meta) periodate (PI) CuSO4, sodium Lascorbate; periodate UV-PVA

humid chamber, 2 h

Click iCs

porcine, acellular SIS porcine, acellular SIS glass mica

porcine skin

horseradish peroxide (HRP) PEG-TA

poly(dopamine-co-acrylate) (PDA) DAM-NIPAAmk

rfp-1 (MAP)

lap shear

lap shear

20 wt% gelatin solution

porcine skin

RT, 24 h

FeCl3 NaIO4

DOPA-incorporated recombinant MAP (dfp-3) LAMBAl

SFA

mica

lap shear

porcine skin

Mfps-CsgA hybrid protein

atomic force microscopy lap shear lap shear

mica substrate; silica probe tip aluminum aluminum porcine skin aluminum

rfp-3 PEU (poly(CA-Tyr-co-Leu))m PEU (poly(CA-Ser-co-Leu-coPPG))o

90 Pa

32 kg/cm2 (x = 5) 32 kg/cm2 (x = 2) 118 gf/cm2

tensile shear strength

lap-shear

poly(AA-co-AANHS-coMDOPA)i AbAf iCs

standardized

9 mN/cm2

alumina polytripeptide (Gly-Tyr-Lys)nc

reported

films dried under N2

lap shear

irradiation with dental curing lamp for 60 s

sodium periodate Bu4N(IO4)n Bu4N(IO4)

V

humid chamber, 2 h 25 °C, 24 h RT, 1 h

3.14 MPa

reference Schnurrer et al.181 Ninan et al.182 Ninan et al.183

Yamamoto et al.191

3.14 MPa 11.56 kPa

± 0.10 MPa ± 0.21 MPa ± 0.01 MPa MPa

Tatehata et al.192 Yu et al.195 Yamada et al.198 Burke et al.209 Lee et al.207 Brubaker et al.218 Wang et al.219

Hoffman et al.200

4.00 ± 0.53 MPa 3.17 ± 0.51 MPa 107 ± 24.7 kPa

Murphy et al.227

∼600 mmHg ∼11.8 kPa

Chung et al.229

168.15 ± 17.02 kPa ∼79.04 ± 9.28 kPa 123.2 ± 13.2 kPa 223.11 ± 15.94 kPa 0.32 MPa F/R = −20 mN/m

Lim et al.220

Guo et al.231

Mehdizade et al.232 Guo et al.233 Xue et al.230 Wei et al.226

RT, 1 day

76 ± 13.4 kPa

Zhang et al.236

dry, 6 h (0:10 NIPAAm:DAM) 80% humidity, 0 min (10:0 NIPAAm: DAM) 2 h, in buffer soln (pH 8.2) 30 min air-dry, 2 h in PBS underwater

2.27 ± 0.33 MPa

Ai et al.237

0.31 ± 0.07 MPa ∼130 kPa

Kim et al.221

∼200 kPa Yang et al.222

RT, 2 h, PBS

F/R = −9.4 mN/m 72.2 ± 3.7 kPa

Zhong et al.224

37 °C, 4 h 65 °C, 24 h RT, 4 h 65 °C, 24 h

F/R = 197.5 mN/m 2.28 MPa 2.46 ± 1.03 MPa 9 kPa 3.2 ± 0.8 MPa

Jeon et al.223

Yang et al.225 Zhou et al.238 Bhagat et al.239


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Biomacromolecules Table 3. continued maximum adhesion strength adhesive POEC-d (octanediol, PEO, citric acid, dopamine) DCTA (gelatin macromer, Fe3+, genipin)

adhesion test lap shear lap shear

dopamine-alginate and HA polysaccharide membrane DOPA-functionalized polyester

Bernkop− Schnürch lap shear

poly((3,4-dihydroxymandelic acid)-co-(lactic acid))

lap shear

4-arm-PEG-DA and 4-arm-PEGPBAq polypeptide-pluronicpolypeptide catechol-functionalized silk fibroin

test substrate

cure condition and time

cross-linker

reported

standardized

sodium periodate

wet, RT, 4 h humid chamber, 2 h

10.6 ± 2.1 kPa 33.7 kPa

porcine skin articular cartilage porcine intestine

FeCl3 + genipin

37 °C, 2 h

24.7 ± 3.3 kPa 194.4 ± 20.7 kPa >300 min

porcine skin aluminum aluminum

Fe(acac)3p

RT, 30 s

Bu4N(IO4)

RT, 30 min; 37 °C, 24 h 37 °C, 24 h, underwater, 24 h RT, 30 min; 37 °C, 24 h

porcine skin porcine skin

lap shear

sanded steel PTFE porcine skin

lap shear tensile shear lap shear

porcine skin porcine bone aluminum shims

RT, 16 h

RT, 30 min HRPr, H2O2

RT, 24−25 h

sodium periodate

RT, 24 h

reference Ji et al.234

13.13 ± 1.74 kPa 218.4 ± 16.0 kPa 2.6 ± 0.4 MPa

Fan et al.203 Scognamiglio et al.204 Shi et al.240 Jenkins et al.241

1.0 ± 0.3 MPa 1.7 ± 0.5 MPa 0.32 ± 0.05 MPa 5.2 ± 0.28 kPa 106 kPa 675 kPa 120 kPa

Shan et al.247 Lu et al.249 Burke et al.165

a MAP, mussel adhesive protein. bTyr, tyrosine; Lys, lysine. cGly, glycine. dDOPA, L-3,4-dihydroxyphenylalanine. ePEG, poly(ethylene glycol). fHA, hyaluronic acid. gAla, alanine. hPCL, polycaprolactone. iAA, acrylic acid; AANHS, acrylic acid N-hydroxysuccinimide ester, MDOPA, Nmethacryloyl-3,4-dihydroxy-L-phenylalanine. jiCMBA, injectable citrate-based mussel-inspired bioadhesives. kDAM, dopamine acrylamide; NIPAAm, N-isopropylacrylamide. lLAMBA, light-activated mussel protein-based bioadhesive. mPEU, poly(ester urea); CA, catechol. nBu4N(IO4), tetrabutylammonium periodate. oSer, serine; PPG, poly(propylene glycol). pFe(acac), iron(III) acetylacetonate. qPBA, phenyl boronic acid; DA, dopamine. rHRP, horseradish peroxide.

Kaur et al. added a second polymerizable phase of polyethylene glycol-diacrylate (PEG-da) to poly(MOEP-coDMA), poly(acrylamide-co-aminopropyl methacrylamide), and Ca2+ coacervates. The highest bonding strength was 973 ± 263 kPa for 17.7 wt% PEG-da on aluminum substrates with good underwater adhesion.255 Wan and co-workers functionalized a polyoxetane backbone with 5 mol% catechol moieties and 25 mol% bis-phosphoric acid groups to mimic sandcastle worm adhesive. A maximum bonding strength of 0.35 MPa was achieved under humid conditions when Fe3+ was used as a curing agent. This study also highlighted that an optimum concentration of adhesive groups and cross-linker is crucial to achieve strong bonding strengths.256 Miserez and co-workers synthesized a two component polypeptide mimic of the sandcastle worm. In this mimic, a lysine-containing positively charged polypeptide was functionalized with DOPA and a negatively charged polypeptide was functionalized with phosphorylated serine and tyrosine. Upon mixing, the two polypeptides formed complex coacervates under neutral charge conditions.257

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BARNACLE MIMETIC ADHESIVES Nishida and co-workers synthesized a polypeptide by ringopening polymerization of NCA monomers to mimic barnacle (Balanus hameri) adhesives. The synthetic mimic demonstrated highest tensileshear strength of 19.2 kg/cm2 (∼1.88 MPa) on iron substrates.258 In their subsequent efforts, they synthesized three different model peptides AA-5, AA-10, and AA-17 with different amino acid compositions. AA-17 cross-linked with tyrosine and α-chymotrypsin demonstrated strong bonds with iron surfaces (∼37 kg/cm2 (0.36 MPa)). All the model peptides showed poor adhesion on bovine bone with tyrosine crosslinked AA-17 showing the strongest bond ∼3.7 kg/cm2 (362.85

Reproduced with permission from ref 251, Copyright 2008 National Academy of Sciences.

Figure 30. Gecko-inspired poly(glycerol-co-sebacate acrylate) (PGSA) tissue adhesive. PGSA prepolymer is nanomolded by UV irradiation followed by spin coating with dextran aldehyde (DXTA). SEM image shows excellent pattern transfer fidelity.

changing cation ratio (Figure 31). This model adhesive was successfully applied in craniofacial reconstruction while demonstrating both in vitro and in vivo adhesion, noncytotoxicity toward cells, along with aligned tissue reconstruction without inflammatory response.252−254 W


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Biomacromolecules

Reprinted from ref 252 with permission from John Wiley and Sons. Copyright 2009 Wiley-VCH Verlag GmbH & KGaA, Weinheim, Germany.

Figure 31. Synthetic adhesive mimic of P. californica glue. 1: Structure of Pc3 analogue; 2: Structure of Pc1 analogue. Model of pH-dependent coacervation and adhesion. (a) At acidic pH 4, polyphosphates (black) and polyamines (gray) form colloidal polyelectrolyte complexes (PECs) with a net positive charge. (b) At a slightly basic pH of 8.2, the extended polyphosphates form a network with polyamines and divalent cations with a net negative charge. (c) Upon oxidation, 3,4-dihydroxyphenol (D) initiates covalent cross-linking and surface interaction via quinones (Q).

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kPa).259 In another study to mimic the barnacle adhesive, a polyacrylamide-based copolymer with hydroxyl and hexyl groups for surface interaction and tetra-alanine groups for cross-linking via hydrogen bonding (hydrophobic interaction) was developed. The copolymer gel showed strong adhesion on PMMA substrate with a tensile adhesive strength of 402 kPa.260

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CONCLUSIONS AND FUTURE PERSPECTIVES Sutures, staples, and metallic grafts are an integral part of surgery and also the gold standard for wound closure. However, the pain and discomfort caused by these invasive techniques have led to an urgent need for development of tissue adhesives for surgical settings. Fibrin glue is the only glue that can be passed off as a hemostat, sealant, as well as an adhesive and finds application in a number of different surgical procedures. Fibrin glue, although deemed biocompatible and degradable, suffers from risk of virus transmission and poor adhesion under wet conditions. Gelatin-resorcin-formaldehyde/glutaraldehyde glue has demonstrated adhesion strengths stronger than fibrin glue, but the application of aldehyde containing materials on tissues is potentially toxic. Cyanoacrylate glue and its variants have consistently shown the strongest adhesion on wet tissue among the different classes of adhesives; however, their toxicity limits their use only to topical applications. Even though these adhesives have been around for centuries, their shortcomings have driven the need to develop alternate tissue adhesives. Polysaccharide- or protein-based tissue adhesives usually covalently attach to the tissues (by Schiff base formation) resulting in strong adhesion under dry conditions. These adhesives often require photoirradiation, addition of photoinitiators, or techniques like tissue laser welding that are detrimental to the neighboring healthy tissue. Biomimetic adhesives is another class of tissue adhesives inspired from the examples of adhesion in nature and rapidly gaining momentum in the field of biological adhesives. Our knowledge of underwater adhesion is still quite limited and considerable efforts are being invested to study adhesion in natural systems. Deeper understanding of the interplay of environmental and chemical factors, chemistries, and mechanisms of natural adhesion will open numerous possibilities for

CADDISFLY-INSPIRED TISSUE ADHESIVES

Stewart and co-workers described an electrostatically driven coacervate formation by using alternating anionic and cationic block copolymers. As a mimic of caddisfly silk, these block copolymers are functionalized with amine, phosphate groups, divalent cations, and also dihydroxyl aromatic groups for oxidative cross-linking.261 Becker and co-workers developed a caddisfly mimic adhesive from phosphate-functionalized, amino acid-based poly(ester urea) copolymer (Figure 32). These adhesives demonstrated strong adhesion of ∼439 ± 203 kPa on bovine bone when cross-linked via electrostatic interactions with Ca2+.262


Figure 32. Caddisfly adhesive mimic of phosphate-functionalized poly(ester urea) copolymer based on serine and valine amino acids. X


Review

Biomacromolecules

(9) Heiss, C.; Kraus, R.; Schluckebier, D.; Stiller, A.-C.; Wenisch, S.; Schnettler, R. Bone adhesives in trauma and orthopedic surgery. European Journal of Trauma 2006, 32, 141−148. (10) Shah, N. V.; Meislin, R. Current state and use of biological adhesives in orthopedic surgery. Orthopedics 2013, 36, 945−956. (11) Farrar, D. F. Bone adhesives for trauma surgery: A review of challenges and developments. Int. J. Adhes. Adhes. 2012, 33, 89−97. (12) Petersen, B.; Barkun, A.; Carpenter, S.; Chotiprasidhi, P.; Chuttani, R.; Silverman, W.; Hussain, N.; Liu, J.; Taitelbaum, G.; Ginsberg, G. G. Tissue adhesives and fibrin glues. Gastrointestinal Endoscopy 2004, 60, 327−333. (13) Ryou, M.; Thompson, C. C. Tissue adhesives: A review. Techniques in Gastrointestinal Endoscopy 2006, 8, 33−37. (14) Spotnitz, W. D. Fibrin sealant: The only approved hemostat, sealant, and adhesive-a laboratory and clinical perspective. ISRN Surgery 2014, 2014, 203943. (15) Reece, T. B.; Maxey, T. S.; Kron, I. L. A prospectus on tissue adhesives. Am. J. Surg. 2001, 182, S40−S44. (16) Lerner, R.; Binur, N. S. Current status of surgical adhesives. J. Surg. Res. 1990, 48, 165−181. (17) Duarte, A. P.; Coelho, J. F.; Bordado, J. C.; Cidade, M. T.; Gil, M. H. Surgical adhesives: Systematic review of the main types and development forecast. Prog. Polym. Sci. 2012, 37, 1031−1050. (18) Ferreira, P.; Gil, H.; Alves, P., An overview in surgical adhesives. In Recent advances in adhesions research; Nova Science Publishers, Inc.: Hauppauge, N.Y, 2013; pp 59−85. (19) Annabi, N.; Tamayol, A.; Shin, S. R.; Ghaemmaghami, A. M.; Peppas, N. A.; Khademhosseini, A. Surgical materials: Current challenges and nano-enabled solutions. Nano Today 2014, 9, 574−589. (20) Peng, H. T.; Shek, P. N. Novel wound sealants: Biomaterials and applications. Expert Rev. Med. Devices 2010, 7, 639−659. (21) Mehdizadeh, M.; Yang, J. Design strategies and applications of tissue bioadhesives. Macromol. Biosci. 2013, 13, 271−288. (22) Modaresifar, K.; Azizian, S.; Hadjizadeh, A. Nano/biomimetic tissue adhesives development: From research to clinical application. Polym. Rev. 2016, 56, 329−361. (23) Chivers, R. A.; Wolowacz, R. G. The strength of adhesivebonded tissue joints. Int. J. Adhes. Adhes. 1997, 17, 127−132. (24) Stewart, R. J.; Wang, C. S.; Shao, H. Complex coacervates as a foundation for synthetic underwater adhesives. Adv. Colloid Interface Sci. 2011, 167, 85−93. (25) Bre, L. P.; Zheng, Y.; Pego, A. P.; Wang, W. Taking tissue adhesives to the future: From traditional synthetic to new biomimetic approaches. Biomater. Sci. 2013, 1, 239−253. (26) Cronkite, E. P.; Lozner, E. L.; Deaver, J. M. Use of thrombin and fibrinogen in skin grafting: Preliminary report. J. Am. Med. Assoc. 1944, 124, 976−978. (27) Alving, B. M.; Weinstein, M. J.; Finlayson, J. S.; Menitove, J. E.; Fratantoni, J. C. Fibrin sealant: Summary of a conference on characteristics and clinical uses. Transfusion (Malden, MA, U. S.) 1995, 35, 783−790. (28) Martinowitz, U.; Saltz, R. Fibrin sealant. Curr. Opin. Hematol. 1996, 3, 395−402. (29) Jackson, M. R. Fibrin sealants in surgical practice: An overview. Am. J. Surg. 2001, 182, 1S−7S. (30) Kjaergard, H. K.; Weis-Fogh, U. S. Important factors influencing the strength of autologous fibrin glue; the fibrin concentration and reaction time − comparison of strength with commercial fibrin glue. Eur. Surg. Res. 1994, 26, 273−276. (31) Sierra, D. H.; Feldman, D. S.; Saltz, R.; Huang, S. A method to determine shear adhesive strength of fibrin sealants. J. Appl. Biomater. 1992, 3, 147−151. (32) Lin, C.; Ritter, J. A. Effect of synthesis ph on the structure of carbon xerogels. Carbon 1997, 35, 1271−1278. (33) Albes, J. M.; Krettek, C.; Hausen, B.; Rohde, R.; Haverich, A.; Borst, H.-G. Biophysical properties of the gelatin-resorcinformaldehyde/glutaraldehyde adhesive. Annals of Thoracic Surgery 1993, 56, 910−915.

further advancement in biomimetic tissue adhesives. Currently, biomimetic adhesives involve coacervate formation or functionalization with an adhesive group like DOPA, catechol, or phosphates. Although these strategies have shown satisfactory adhesion under dry conditions, they often tend to fail under humid and/or wet conditions. Recent studies have looked at incorporating a range of cross-linking chemistries to strike a balance between the adhesive and cohesive strengths. Despite such extensive research on tissue adhesives, we have been unsuccessful in developing adhesive mimics capable of rivaling natural adhesives that are scalable, nontoxic, biocompatible, easy to use and degradable. There is still a need to explore the effect of backbone chemistry, polymer hydrophilicity/hydrophobicity, combination of different amino acids, and charged side groups on adhesion strengths. It is also necessary to widen the scope of adhesives beyond tissue adhesion toward drug delivery, tissue grafts, wound healing, and tissue reconstruction via addition of peptides. In addition, long-term studies and clinical trials are essential before these adhesives can be realized in medical or surgical applications.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. ORCID

Matthew L. Becker: 0000-0003-4089-6916 Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors gratefully acknowledge financial support from the Biomaterials Division of the National Science Foundation (DMR-1507420), the United States Department of Defense (USAMRMC-W81XWH-15-1-0718), and the W. Gerald Austen Professor of Polymer Science and Polymer Engineering endowed by the Knight Foundation.

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