A pressure-sensitive bioadhesive for wet tissues and organs

Abstract

A pressure-sensitive adhesive (PSA), method of manufacturing a PSA, and various uses thereof synergistically combines the advantages of viscoelastic PSAs and bioadhesives. Enabled by a one-pot scalable copolymerization of a polyester and hydrophilic polymer, the PSA provides near-instant (~ 1 s), robust, and repeatable (over 1,000 times) adhesion to wet biological tissues (i.e., skin, lung, heart) and engineering (i.e., metals, plastics) substrates without prior surface functionalization.

Description

A Pressure-Sensitive Bioadhesive for Wet Tissues and OrgansCross Reference to Related Applications

[0001] The present application is based on, claims priority to, and incorporates herein by reference in its entirety for all purposes, US Provisional Application Serial No. 63 / 601,461, filed November 21, 2023.Statement of Government Support

[0002] N / A.Background

[0003] Pressure-sensitive adhesives (PSAs), such as Scotch Tape and VHB, have routinely been used in household and industrial applications due to their set of functionalities, including instant, robust, and repositionable adhesion on diverse materials. Beyond their current applications, the PSA’s unique combination of advantages is also highly favorable for biomedical and clinical applications, such as on-demand sealant and rapid repositioning of biomedical devices on biological tissues. However, existing PSAs are not designed for use on wet surfaces, and therefore, exhibit poor adhesion performance on wet tissues and organs.

[0004] To achieve adhesion on wet biological tissues and organs, a wide range of bioadhesives have been developed in recent decades, including nanoparticle solutions, photocurable tissue adhesives, tough hydrogel adhesives, dry bioadhesives, and hydrophobic tissue adhesives. However, existing high-performance bioadhesives mostly support single irreversible adhesion formation on surfaces with specific chemical groups, facing challenges to cope with the need for repeated re-use and repositioning of bioadhesives in diverse application scenarios such as tissue repair and integration of implantable devices. While reversible bioadhesives have also been developed based on triggered cleavage of crosslinks, dynamic covalent bonds, and gradual formation of covalent crosslinks, the existing reversible bioadhesives are limited to a single or limited number (for example, less than 20) of repositioning within a short period of time after the initial application (for example, less than 1 h) on surfaces with specific chemical groups.

[0005] Hence, despite recent advances, realizing the advantages of PSA to wet tissues and organs remains an unmet demand in the field, highlighting the need for new strategies and solutions.Summary

[0006] The present disclosure provides PSAs and methods of making and applying PSAs that overcome the aforementioned drawbacks by combining the advantages of PSAs and tissue adhesives to achieve 1) instant and stable adhesion, 2) broad applicability, and 3) atraumatic repositioning. The disclosure adopts a strategy for tuning the viscoelastic properties for a PSAbased adhesion mechanism and promoting interfacial water absorption to create a PSA for wet hydrophilic tissues.

[0007] In one aspect of the present disclosure, a pressure-sensitive adhesive (PSA) for use on a surface is presented. The adhesive comprises an amphiphilic block copolymer, configured to provide repeatable adhesion to hydrophilic surfaces. The amphiphilic block copolymer comprises a polyester and a hydrophilic polymer, wherein the molar ratio of the polyester to the hydrophilic polymer is about 4:1. The polyester comprises a dicarboxylic acid and glycerol at a molar ratio of about 2: 1.

[0008] In another aspect of the present disclosure, a method of manufacturing a pressuresensitive adhesive (PSA) is described. The method comprises adding a dicarboxylic acid to a vessel, adding a hydrophilic polymer to the vessel, wherein a molar ratio of the dicarboxylic acid to the hydrophilic polymer is about 4: 1, performing a first polymerization of the dicarboxylic acid and hydrophilic polymer via condensation polymerization to form a linear polymer, adding glycerol to the vessel, wherein a molar ratio of the dicarboxylic acid to the glycerol is about 2: 1, and performing a second polymerization of the linear polymer and the glycerol to form a crosslinked adhesive.

[0009] In another aspect of the present disclosure, a method of atraumatic application on and removal from a biological tissue of a pressure-sensitive adhesive (PSA) is described. The method comprises applying the PSA to a location on the biological tissue, removing the PSA, and optionally repositioning the PSA on the location or to a different location on the biological tissue.

[0010] These aspects are nonlimiting. Other aspects and features of the systems and methods described herein will be provided below.Brief Description of the Drawings

[0011] The foregoing features of embodiments will be more readily understood by reference to the following detailed description, taken with reference to the accompanying drawings, in which:

[0012] FIG. 1 is a schematic of the pressure sensitive adhesive (PSA), as described herein.

[0013] FIG. 2 is a method of making the PSA, as described herein.

[0014] FIG. 3 is a flowchart of the non-limiting applications of the PSA, as described herein.

[0015] FIG. 4A is the concept and design of a pressure-sensitive bioadhesive (PSB). Schematic illustration for the concept of PSB synergistically combining the advantages of tissue adhesives and pressure-sensitive adhesives for substrate-independent, instant and reversible adhesion.

[0016] FIG. 4B is a schematic illustration of the PSB enabling the near-instant, atraumatic adhesion between biomedical tools and wet tissues for applications such as tissue stabilization and fault-tolerant bioelectronic interfacing.

[0017] FIG. 4C is a plot of the range of applicable substrates without pretreatment and repositioning cycles of the PSB in comparison with previous literature.

[0018] FIG. 4D is a plot of the adhesion time and repositioning cycles of the PSB in comparison with previous literature.

[0019] FIG. 4E is a schematic illustration for independent, instant (~1 s), and reversible adhesion of the PSB. The reversible adhesion of the PSB enables atraumatic removal and instant reapplication for fault tolerant applications.

[0020] FIG. 4F is an image (left) and design (right) of the PSB. The hydrophilic PEG moiety absorbs interfacial water between the adhesive and wet surfaces. The crosslinked PGS enables optimized viscoelasticity. PGS, poly(glycerol sebacate). PEG, polyethylene glycol. PTFE, poly tetrafluoroethyl ene .

[0021] FIG. 4G is schematic illustrations for the bonding and debonding phases of the PSB on a wet tissue surface. The bonding phase demonstrates a combination of interfacial water absorptions and viscoelastic flow for physical adhesion. During the debonding phase, the viscous component of the complex modulus acts as energy dissipation for adhesion hysteresis and reversible detachment.

[0022] FIG. 5 is a synthetic scheme of the PSB. During step 1 synthesis, sebacic acid and PEG condense into a linear polymer chain to synthesize the step 1 product. In step 2 synthesis, glycerol is added to the step 1 product to create a crosslinked PGS-co-PEG network (PSB) at varying crosslinking degrees by tuning the reaction time (75 h to 90 h).

[0023] FIG. 6 is an ’H NMR spectrum of the PSB. CDCI3 (7.28 ppm peak) is used as a solvent for the PSB. 4.12-4.35 ppm multiplet peaks are observed due to methylene and methine groups of glycerol. 5.25-5.31 ppm singlet peak is attributable to the hydrogen on the hydroxyl group of glycerol. A methylene peak at 3.64 ppm is observed due to the ethylene group in PEG. Peaks at 2.35, 1.62, and 1.30 ppm are assigned to methylene groups of sebacic acid. 'H NMR (400 MHz, CDCI3): 5 5.25-5.31 (s, 1H), 84.12-4.35 (m, 5H), 8 3.64 (s, 49H), 82.35 (s, 9H), 8 1.62 (s, 9H), 8 1.30 (s, 18H).

[0024] FIG. 7 is an 'H NMR spectrum of the PSB. (5H; Blue). The 5.25-5.31 ppm singlet peak is assigned to the hydrogen atom present on the hydroxyl group of glycerol (1H; Red). Numbers marked below the peaks represent relative peak integrations between methylene and methine groups (right) and the hydroxyl group (left). The relative ratio indicates the substitution rate of the hydroxyl group. As the ratio is 5.00:0.40, 40% of the hydroxyl hydrogen remains and 60% is reacted to the ester group. *H NMR (400 MHz, CDCI3): 8 4.35-4.12 (m, 5H), 5.31-5.25 (s, 1H).

[0025] FIG. 8 shows FTIR spectra of the PSB. Panel a: FTIR spectrum of the PSB synthesized with varying reaction time (8 h to 120 h). Panels b-c: Selected parts of the FTIR spectrum of the PSB for wavelength ranges between 2,400 cm'1to 3,600 cm’1and 1,650 cm’1to 1,850 cm’1. Peaks at 1,730 cm’1correspond to ester groups in the PSB (higher for increased crosslinking). Peaks at 1,690 cm’1correspond to carboxylic groups in the PSB (lower for increased crosslinking). Peaks at 3,460 cm’1correspond to hydroxyl groups in the PSB (lower for increased crosslinking).

[0026] FIG. 9A is a plot of the glass transition temperature of the viscoelastic PSB. Differential scanning calorimetry (DSC) of the PSB indicates that the inflection point at -56.3 °C is used as the glass transition temperature.

[0027] FIG. 9B is a plot of DSC of the PEG-co-PGS material system at varying degrees of crosslinking. The vertical lines are used to mark the glass transition temperature of each sample.

[0028] FIG. 10 is a schematic of the rapid and reversible adhesion mechanism of the PSB on wet surfaces (top panel). Snapshots of the rapid and reversible adhesion process of the PSB on a wet surface (bottom panel). Water is colored with a blue food dye for visualization.

[0029] FIG. 11 A shows images of the PSB with different viscoelasticity exhibiting viscous fluid (left, reaction time ~70 h), viscoelastic PSB (middle, reaction time -76 h), and viscoelastic solid (right, reaction time -100 h) like behaviors.

[0030] FIG. 1 IB shows a series of plots of the change in viscoelastic behavior (shown through viscous modulus (G”) and elastic modulus (G’)) during reaction progression of the PSB. Frequency ranges corresponding to the bonding (blue, - 1 Hz) and debonding (red, - 100 Hz) behavior in the viscous fluid state (left), viscoelastic PSB (middle), and viscoelastic solid state (right) are shaded.

[0031] FIG. 11C shows images of interfacial water absorption of the PSB on wet hydrogel surfaces. Cross-sectional fluorescent image between 2 wt.% agarose hydrogel and the PSB (top) in comparison with VHB (bottom). The dashed lines indicate the interface between the agarose hydrogel and the PSB or VHB

[0032] FIG. 12 is a plot of PGS-co-PEG crosslinking as the copolymerization reaction progresses. Change in the storage modulus (G’) occurs as crosslinking progresses during the synthesis of the PSB.

[0033] FIG. 13A is a plot of loss modulus (G”) and storage modulus (G’) in a frequency sweep shown with the Dahlquist criteria (dashed line).

[0034] FIG. 13B is a plot of the Stress-strain curves of the PSB tack adhesion tests under various applied pressures for 1 s with a stainless steel probe.

[0035] FIG. 13C is a plot of the correlation between reaction progression (measured with loss factor) and peel test interfacial toughness (porcine skin substrate) of the PSB. Values represent the mean and the standard deviation (ri = 4).

[0036] FIG. 14A is a plot of the tensile adhesion retained on wet porcine skin with reference to dry porcine skin for PSB and viscoelastic PGS.

[0037] FIG. 14B is a plot of the self-healing stress-strain graph of two pristine PSB pieces after pressing for 5 seconds. Curve and shaded area represent mean and standard deviation respectively, (n = 3).

[0038] FIG. 14C shows a schematic and photos of robust, rapid self-healing process of two separate PSB pieces.

[0039] FIG. 15A is a plot of short-term water swelling of PSB in comparison with PGS and elastomeric PGS-co-PEG. Values represent the mean and standard deviation ( / / = 4 for a).

[0040] FIG. 15B is a plot of adhesion to wet porcine skin of PSB in comparison with PGS and elastomeric PGS-co-PEG. Values represent the mean and standard deviation (n = 3~6 for b).

[0041] FIG. 15C is a plot of rheological properties (shown through loss factor) plotted against frequency of PSB in comparison with PGS, viscoelastic PGS, and elastomeric PGS-co-PEG.

[0042] FIG. 16A is a series of images of the VHB (top) and the PSB (bottom) application and detachment process on a dry glass surface.

[0043] FIG. 16B is a series of images of the VHB (top) and the PSB (bottom) application and detachment process on a wet glass surface. Values in d represent the mean and the standard deviation (n = 3-6).

[0044] FIG. 17A is a photograph the robust adhesion on a porcine heart with PSB.

[0045] FIG. 17B is a photograph of no adhesion on a porcine heart with VHB.

[0046] FIG. 17C is a photograph of no adhesion on a porcine heart with Micropore.

[0047] FIG. 17D shows images of the robust adhesion of PSB (left) on lung tissue in comparison to the low adhesion of VHB (right).

[0048] FIG. 17E is a plot of the comparison of tensile strength with PSB and commercial and clinical pressure sensitive adhesives on lung and heart tissues. Values represent the mean and the standard deviation (n = 4).

[0049] FIG. 17F is a plot of the reversible adhesion of the PSB in comparison with commercial and clinical pressure sensitive adhesives on lung tissue. Values represent the mean and the standard deviation (n = 4).

[0050] FIG. 18 is a plot of the tensile adhesion of the PSB on blood covered porcine skin in comparison to wet porcine skin. The dashed line represents the tensile adhesion of commercial PSAs on wet porcine skin. The values and error bars represent the mean and standard deviation respectively (n = 5).

[0051] FIG. 19A is a plot of the force vs. displacement curves of cyclic tensile tests of the PSB adhered on wet porcine skin.

[0052] FIG. 19B is a plot of the tensile strength of the PSB and VHB over cyclic tensile tests on wet porcine skin. Values represent the mean and the standard deviation (n = 5).

[0053] FIG. 19C is a plot of the adhesion energy of the PSB and VHB over cyclic tensile tests on wet porcine skin. Values represent the mean and the standard deviation (p = 5).

[0054] FIG. 20 is a plot of the adhesion stability of the PSB, measured with normalized tensile adhesion force, over approximately 1 hour of continuous usage and 75 adhesion cycles. The curve and shaded area represent the mean and standard deviation respectively (n = 3).

[0055] FIG. 21A is a schematic of the reversible underwater tensile adhesion of the PSB. on porcine skin.

[0056] FIG. 2 IB is a plot of normalized tensile adhesion force of the PSB as a function of adhesion cycles in underwater environments.

[0057] FIG. 22A is a plot of interfacial toughness of the PSB on various biological tissues. Values represent the mean and the standard deviation (n = 4).

[0058] FIG. 22B is a plot of interfacial toughness of the PSB on various engineering substrates (polystyrene - PS, polydimethylsiloxane - PDMS, stainless steel - SS, polyurethane - PU). Values represent the mean and the standard deviation (zz = 4).

[0059] FIG. 23 A is a plot of the interfacial toughness of PSB, VHB, Transpore, and Micropore on engineering substrate, porcine skin, and water-covered porcine skin (For engineering substrate, PSB vs Micropore p = 0.14) (For porcine skin, PSB vs VHB p = 0.005, PSB vs Transpore p = 0.005, PSB vs Micropore p = 0.005) (For water-covered porcine skin, PSB vs VHB p = 0.002, PSB vs Transpore p = 0.002, PSB vs Micropore p = 0.002).

[0060] FIG. S23B is a plot of force vs. displacement curves of the PSB, VHB, Transpore, and Micropore on an engineering substrate.

[0061] FIG. S23C is a plot of force vs. displacement curves of the PSB, VHB, Transpore, and Micropore on porcine skin.

[0062] FIG. S23D is a plot of force vs. displacement curves of the PSB, VHB, Transpore, and Micropore on water-covered porcine skin.

[0063] FIG. 24A is an image of the PSB prepared on various backing materials. From left to right: styrene ethylene butylene styrene (SEBS), polyurethane (PU), polytetrafluoroethylene (PTFE), polycarbonate (PC), and stainless steel.

[0064] FIG. 24B is a series of images of the PSB with Ecoflex backing before stretching (left), during stretching (middle), and being twisted (right).

[0065] FIG. 25 is a plot of the tensile property of the PSB. Engineering stress vs engineering strain curve of the PSB at the tensile strain rate of 1 min1.

[0066] FIG. 26A is a plot of the interfacial toughness versus pressing force of the PSB on wet porcine skin. Values represent the mean and the standard deviation (ri = 4).

[0067] FIG. 26B is a plot of the interfacial toughness versus time after pressing of the PSB on wet porcine skin. Values represent the mean and the standard deviation n = 4).

[0068] FIG. 26C is a plot of the interfacial toughness versus pressing time of the PSB on wet porcine skin. Values represent the mean and the standard deviation n = 4).

[0069] FIG. 27A is a schematic of the usage configuration of the PSB as single-sided tape.

[0070] FIG. 27B is a schematic of the usage configuration of the PSB as double-sided tape.

[0071] FIG. 28A is a schematic illustration on the challenges of tissue stabilization in commercial devices with application of the PSB.

[0072] FIG. 28B is a schematic illustration on the challenges of tissue stabilization in commercial devices without application of the PSB.

[0073] FIG. 28C shows schematic illustrations of functional effects of the PSB for use as a tissue stabilizing application.

[0074] FIG. 28D is an image comparison of tissue stabilizers on lung tissue with (left) and without (right) the PSB.

[0075] FIG. 28E is a plot of the comparison of stabilizing pressure with and without the PSB during attachment and reattachment.

[0076] FIG. 28F is a series of photos of minimal tissue damage with (top row) the PSB during tissue stabilization and photos of tissue damage without (bottom row) the PSB during tissue stabilization.

[0077] FIG. 29A is representative LIVE / DEAD assay of HEK 293 cells and 3T3 cells cultured in the control media (DMEM) and the PSB-conditioned media. DMEM, Dulbecco’s modified eagle’s medium.

[0078] FIG. 29B is a plot of the cell viability of HEK 293 cells and 3T3 cells cultured in the control media (DMEM), the PSB-conditioned media, and the biodegraded PSB-conditioned media for 24 h. Values represent the mean and the standard deviation (n = 3~4).

[0079] FIG. 30A is a histological image stained with hematoxylin and eosin (H&E) for the PSB with SEBS backing after 1 week of subcutaneous implantation for the evaluation of in vivo biocompatibility. SEBS, Styrene-ethylene-butylene-styrene.

[0080] FIG. 30B is a histological image stained with H&E for a control SEBS backing after 1 week of subcutaneous implantation for the evaluation of in vivo biocompatibility. SEBS, Sty rene-ethy 1 ene-buty 1 ene- styrene .

[0081] FIG. 30C is a histological image stained with H&E for the biodegraded PSB after 1 week of subcutaneous implantation for the evaluation of in vivo biocompatibility. SEBS, Styrene- ethylene-butylene-styrene.

[0082] FIG. 31 is a plot of the histological evaluation of subcutaneous biocompatibility with a control SEBS backing (control) and the PSB. Lesion assessment was rated as 0: normal, 1 : minimal, 2: mild, 3: moderate, 4: marked, and 5: severe. The values and error bars represent the mean and standard deviation respectively (n = 4).

[0083] FIG. 32A is a plot of the histological evaluation of epidermal thickness with a control SEBS backing (control) and the PSB. Lesion was assessment rated as 0: normal, 1 : minimal, 2: mild, 3: moderate, 4: marked, and 5: severe. The values and error bars represent the mean and standard deviation respectively (n = 4).

[0084] FIG. 32B is a plot of the histological evaluation of epidermal biocompatibility with a control SEBS backing (control) and the PSB. Lesion was assessment rated as 0: normal, 1 : minimal, 2: mild, 3: moderate, 4: marked, and 5: severe. The values and error bars represent the mean and standard deviation respectively (n = 4).

[0085] FIG. 33A is a representative histological epidermal image stained with H&E in Sprague Dawley rats testing biocompatibility of a control SEBS backing.

[0086] FIG. 33B is a representative histological epidermal image stained with H&E in Sprague Dawley rats testing biocompatibility of the PSB on dorsal epidermis.

[0087] FIG. 34A is a series of images of the PSB after incubation periods in saline solution at 37 °C.

[0088] FIG. 34B is a plot of the biodegradation of the PSB (measured through weight loss) as a function of time after incubation periods in saline solution at 37 °C. Curve and shaded area represent mean and standard deviation (n = 5).

[0089] FIG. 35A is a Schematic illustration for the PSB-assisted integration and repositioning of an electrode on the sciatic nerve for stimulation of the hindlimb.

[0090] FIG. 35B is a plot of the measured hindlimb movements after stimulation by the electrode integrated to different locations in the sciatic nerve by the PSB. Values represent the mean and the standard deviation (n = 4).

[0091] FIG. 36A is a schematic illustration for the PSB-assisted integration and repositioning of an electrode on the epicardial surface for measurement of desired electrocardiogram (ECG) signals.

[0092] FIG. 36B shows representative images (top) and ECG signals (bottom) of the electrode adhered to the epicardial surface by the PSB in different locations after initial application and repositioning.

[0093] FIG. 36C shows snapshots of the initial application and repositioning of the electrode on the epicardial surface by the PSB.

[0094] FIG. 37A is a schematic illustration for PSB-assisted integration and repositioning of an electrode on the epicardial surface for measurement of desired electrocardiogram (ECG) signals.

[0095] FIG. 37B is an image for PSB-assisted integration and repositioning of an electrode on the epicardial surface for measurement of desired electrocardiogram (ECG) signals.

[0096] FIG. 37C is a plot of the recorded ECG signals on the porcine heart after (1) initial application and (2) after repositioning.

[0097] FIG. 37D is a plot of the average recorded ECG signals on the porcine heart for 10 min. The standard deviation of the signal is marked in gray.

[0098] FIG. 37E shows snapshots of the fault-tolerant application of the PSB on the surface of the porcine lung.

[0099] FIG. 37F is a series of representative images of the electrode integrated to the lung surface by the PSB in different locations after the first (left), second (middle), and final application (right).

[0100] FIG. 37G is series of plots of the measured impedance of the strain sensor during ventilator breathing movements at the first (left), second (middle), and final (right) applications.

[0101] FIG. 38A is a schematic illustration of the custom-made strain sensor under cyclic deformation.

[0102] FIG. 38B is a plot of the impedance measurements as a function of time (IV at 1 kHz) of the strain sensor during a cyclic tensile strain amplitude of 1 %.

[0103] FIG. 38C is a plot of the impedance measurements as a function of time (IV at 1 kHz) of the strain sensor during a cyclic tensile strain amplitude of 3 %.

[0104] FIG. 38D is a plot of the impedance measurements as a function of time (IV at 1 kHz) of the strain sensor during a cyclic tensile strain amplitude of 5 %.

[0105] FIG. 39A is a plot of the sensitivity and limit of detection (LOD, 0.13% strain) of carbon nanotube (CNT)-Ecoflex composite strain sensors (n = 3). The values and error bars represent the mean and standard deviation, respectively.

[0106] FIG. 39B is a plot of the R / R0 response of CNT-Ecoflex strain sensors followed by 1 mm of machine extension.

[0107] FIG. 39C is a plot of the cyclic performance of CNT-Ecoflex strain sensors across 700 cycles. Enlarged cyclic performance during the first couple minutes of operation (100 ~ 200 s) and after 1.5 hours (4500 ~ 4600 s). All experiments were conducted on three independent samples and showed similar results.

[0108] FIG. 40 is a plot of the tensile adhesion of varying rations of sebacic acid to PEG.

[0109] FIG. 41 is a plot of the adhesion force of a PSAs with different types of dicarboxylic acids.

[0110] FIG. 42 is a plot of the adhesion force of PSAs with varying ratios between sebacic acid and glycerol.Detailed Description[0U1] Enabled by the unique amphiphilic block copolymer-based design and systematic optimization of viscoelastic properties, the PSA described herein realizes the advantageous characteristics of PSA on wet surfaces with rapid (~ 1 s), robust (interfacial toughness > 380 J m’2), and repeatedly repositionable (> 1,000 times) adhesion while being biocompatible and biodegradable. The PSA exhibits universal applicability to a diverse range of wet surfaces including various biological tissues (i.e., skin, lung, heart) as well as engineering solids commonly used in medical devices (i.e., metals, plastics including low-surface-energy material such as polytetrafluoroethylene (PTFE)) without prior surface functionalization. When used on various biological tissues, the PSA may be referred to as a pressure-sensitive bioadhesive (PSB). We further demonstrate potential biomedical applications of the PSA as a surgical sealant in fast-paced clinically-relevant scenarios as well as its applications to fault-tolerant and in situ adaptable interactions of bioelectronics with internal organs in rat and porcine models.

[0112] Unlike many bioadhesives that require specific surface chemistry or composition to achieve wet adhesion, the disclosed PSA offers near-instant, robust, and repeatedly repositionable adhesion universally applicable to a wide range of materials including biological tissues and medical device components. In addition, to validate the PSA’s adhesive performance, we demonstrate a set of proof-of-concept applications of the PSA for fault-tolerant and in situ adaptable integration of bioelectronic devices to wet and dynamic organs in both small and large animal models. The PSA described herein may enable new opportunities for the integration of implantable devices to the human body but also for clinical uses and improved patient care.

[0113] Referring to the schematic 100 of FIG. 1, the PSA (or PSB) 102 may be reversibly adhered to a surface 104. In a non-limiting example, adhesion, removal, and re-adhesion of the PSA to the surface 104 may be performed up to 1000 times. The surface 104 may include one or more surfaces. The surface 104 may include a biological tissue, medical device, or both. For example, the biological tissue may be an external tissue such as the outer surface of the skin. Alternatively, the biological tissue may be an internal tissue, such as, but not limited to, the heart and lungs. In a non-limiting example, the surface 104 may be hydrophobic or hydrophilic. As used herein, the term “hydrophilic surface” refers to a water-retaining surface.

[0114] In a non-limiting example, the PSA 102 is an amphiphilic block copolymer 106 comprising a polyester 108 and a hydrophilic polymer 110. The polyester may comprise a dicarboxylic acid and glycerol. In a non-limiting example, the dicarboxylic acid and glycerol are at a molar ratio of about 4: 1 to 1 : 1. In a non-limiting example, the molar ratio of the dicarboxylic acid and glycerol may be about 4: 1 to 2: 1. In another example, the molar ratio of the dicarboxylic acid and glycerol may be about 4: 1 to 3:1. In yet another example, the dicarboxylic acid and glycerol may be at a molar ratio of about 2: 1.

[0115] In a non-limiting example, the dicarboxylic acid may include sebacic acid, adipic acid, pimelic acid, suberic acid, phthalic acid, dodecanedioic acid, or derivatives thereof. In another example, the dicarboxylic acid may include malonic acid, succinic acid, glutaric acid, azelaic acid, terephthalic acid, or hexanedioic acid (hexanedicarboxylic acid). Further variants and substituted dicarboxylic acids that may be utilized include itaconic acid, fumaric acid, maleic acid, isophthalic acid, cyclohexanedicarboxylic acid, or trimellitic acid. In another example,emerging and specialized dicarboxylic acids such as eicosanedioic acid, octanedioic acid (substituted variants), citraconic acid, or oxalic acid may be used.

[0116] The polyester 108 may include, but is not limited to, poly(glycerol sebacate) (PGS), poly(glycerol glutarate), poly(glycerol adipate), poly(glycerol pimelate), poly(glycerol suberate), poly(glycerol phthalic acid), or poly(glycerol dodecanedioate). As will be described in further detail below, the dicarboxylic acid and glycerol form crosslinks 112 in forming the amphiphilic block copolymer 106 PSA 102.

[0117] In an alternative non-limiting example, the polyester 108 may be diol based, such as, but not limited to poly(l,6-hexanediol sebacate), or poly(l,5-pentanediol sebacate). Alternatively, the polyester 108 may be poly(lactic acid) (PLA), poly(gly colic acid) (PGA), poly(lactic-co- glycolic acid) (PLGA), polycaprolactone (PCL), poly(hydroxybutyrate) (PHB), poly(trimethylene carbonate) (PTMC), poly(ethylene glycol-co-decalactone). In another nonlimiting example, the polyester 108 may be biodegradable polyesters such as poly(dioxanone) (PDO), poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), poly(butylene succinate) (PBS), or poly(butylene adipate-co-terephthalate) (PBAT). Other emerging polyester systems may alternatively be employed, such as polypropylene fumarate) (PPF), poly(glycerol-co-citrate), poly(alkylene succinate), or poly(alkylene glutarate). In another example, crosslinked polyesters using multifunctional alcohols and diacids or hyperbranched polyesters based on glycerol and diacids may be used such as poly(glycerol citrate).

[0118] In a non-limiting example, the hydrophilic polymer 110 and the polyester 108 are at a weight ratio of about 10 to 70 % w / w. In a non-limiting example, the weight ratio may be about 10% to 60%, about 10% to 50%, about 10%, to 40%, about 10% to 30%, or about 10% to 20%. In yet another example, the weight ratio may be about 50%. The hydrophilic polymer 110 may include synthetic polymers such as polyethylene glycol (PEG), poly(vinyl alcohol) (PVA), poly(acrylic acid) (PAA), polypropylene glycol) (PPG), poly(2-hydroxyethyl methacrylate) (PHEMA), poly(N-vinylpyrrolidone) (PVP), polypiethacrylic acid) (PMAA), poly(ethylene oxide) (PEO), poly(aspartic acid), poly(acrylamide), poly(ethylenimine) (PEI). Alternatively, the hydrophilic polymer 110 may include natural or bio-based polymers such as dextran, starch- based polymers, cellulose derivates (e.g., hydroxyethyl cellulose, carboxymethyl cellulose), agarose, chitosan, alginate, hyaluronic acid, gelatin, or pectin. In another non-limiting example, the hydrophilic polymer 110 may include copolymers with hydrophilic characteristics such aspoly(PEG-co-lactic acid), poly(PEG-co-PLGA), block copolymers of PEG and PCL, or poly(PEG-co-PPG) (e g., Pluronic block copolymers). Alternatively, the hydrophilic polymer 110 may include emerging and specialized polymers such as polysialic acid, poly(betaines) (e.g., poly(sulfobetaine methacrylate)), poly(glycerol methacrylate), poly(N- isopropyl acrylamide) (PNIPAAm, in hydrophilic-to-hydrophobic transition systems), or polyoxazoline (POX).

[0119] In a non-limited example, the PSA 102 forms an immediate adhesion to the surface 104. As used herein, the term “immediate” may be about 1 second.

[0120] In a non-limited example, the PSA 102 forms an interfacial toughness up to 1 kJ / m2.

[0121] As described in further detail in the Example below, the storage modulus G’ of the PSA 102 is less than 105Pa at a frequency of 1 Hz during boding of the PSA 102 to the surface 104. In one example, the storage modulus G’ of the PSA 102 is 100 Pa during bonding at 1 Hz. The storage modulus G’ of the PSA 102 is about 1,000 Pa to 10,000 Pa at 100 Hz during removal from the surface 104. At 1 Hz, the loss modulus G” of the PSA 102 is greater than the storage modulus G’. At 100 Hz, the loss modulus G” of the PSA 102 is less than the storage modulus G’.

[0122] In a non-limiting example, the air burst pressure and water burst pressure is greater than 120 mmHg.

[0123] As will be described in further detail in the Example below, the PSA 102 may be a single-sided PSA (sPSA) or a double-sided PSA (dPSA). In the sPSA configuration, a backing material may be adhered to the side of the PSA not adhered to the surface. In a non-limiting example, the backing material includes polycarbonate. In the dPSA configuration, the PSA 102 may adhere to two or more opposite surfaces. For example, the dPSA may adhere to a tissue on one side and a medical device on the opposite side.

[0124] Referring now to FIG. 2, a method 200 of manufacturing a PSA or PSB as described above is presented. At step 202, a dicarboxylic acid is added to a reaction vessel, and may include any of the dicarboxylic acid listed previously. At step 204 a hydrophilic polymer is added to the vessel, wherein the molar ratio of the dicarboxylic acid to the hydrophilic polymer is about 20:1 to 2: 1. The hydrophilic polymer may include any of the hydrophilic polymers previously listed. In a non-limiting example, the molar ratio of the dicarboxylic acid to the hydrophilic polymer is about 4:1.

[0125] At step 206, a first polymerization reaction is performed between the dicarboxylic acid and the hydrophilic polymer. The resulting product forms a linear copolymer.

[0126] Next at step 208, glycerol is added to the vessel. In a non-limiting example, the molar ratio of the dicarboxylic acid to glycerol is about 4: 1 to 1 : 1. In another example, the molar ratio of the dicarboxylic acid to glycerol is about 4: 1 to 2: 1 or 4: 1 to 3 : 1. In yet another example, the molar ratio of the dicarboxylic acid to glycerol may be about 2: 1. At step 210, a second polymerization is performed between the linear copolymer and glycerol to form a crosslinked adhesive. It is noted that “crosslinked adhesive” may be used interchangeably with PSA or PSB.

[0127] In a non-limiting example, the second polymerization at step 210 occurs until the crosslinked adhesive reaches a storage modulus G’ of about 100 Pa at 1 Hz. Furthermore, the second polymerization at step 210 occurs until the crosslinked adhesive reaches a storage modulus G’ of about 1000 Pa to 10000 at 100 Hz.

[0128] In a non-limiting example, the second polymerization may take place at a temperature of 120-130 °C. The second polymerization may be performed for a duration between 60 and 100 hours. Specifically, the second polymerization may be performed for a duration of 70-100 hours. More specifically, the second polymerization may be performed for a duration of 75-90 hours.

[0129] Alternatively, the second polymerization may be reacted at a higher temperature and for a shorter duration. For example, the second polymerization may take place at a temperature of 140-150 °C. At this higher temperature, the second polymerization may be performed for a minimum duration of 30 hours.

[0130] In a non-limiting example, a backing material may be placed on the crosslinked adhesive to form a sPSA as previously described. Furthermore, the crosslinked adhesive may be topped with a release liner for preserving it during storage and / or before use. In a non-limiting example, the release liner includes polytetrafluorethylene (PTFE). Upon use of the PSA, the release liner is removed from the side of the adhesive intended to adhere to the surface.

[0131] In a non-limiting example, the crosslinked adhesive may be further pressed to form a uniform surface and thickness. For example, pressing the crosslink adhesive may ensure a flat plane is created on the side of the PSA intended to adhere to a surface.

[0132] An example method 300 of various atraumatic uses of the PSA on biological tissue is shown. At step 302, any of the PSA, PSB, or crosslinked adhesive formulations previously described are applied to a biological tissue. Thereafter, the PSA may be used to stabilize thebiological tissue 304’ during a surgical procedure. Additionally or alternatively, the PSA may be used to fix one or more electrodes placed on the surface of the biological tissue thereto (step 304”). In a non-limiting example, the one or more electrodes may include a measurement electrode, a stimulation electrode, or both. Additionally or alternatively, the PSA may be used to fix one or more sensors placed on the surface of the biological tissue thereto (step 304”’). In a non-limiting example, the one or more sensors may include a strain sensor. Details of applications 304’-304”’ are provided in further detail in the Example below.

[0133] At step 306, the PSA may be removed from the biological tissue. Optionally, the PSA may be repositioned to the same location or to a different location on the biological tissue at step 308. Thereafter, the method may return to any combination of steps 304’-304’”.

[0134] The following non-limiting examples demonstrate a manufacturing process, characterization, and applications of the PSA described herein and are not intended to be limiting.

[0135] Example 1

[0136] 1. Introduction

[0137] Pressure-sensitive adhesives are routinely used in household, industrial, and medical applications due to their distinct adhesive properties. These adhesives are designed to form substrate-independent, instant, and reversible attachments to diverse surfaces, leveraging their viscoelastic properties to facilitate physical interactions. In addition, pressure-sensitive adhesives are extremely simple and intuitive to attach and detach from the attached substrate, allowing people to use them without any specific training or instructions. In the clinic, pressure-sensitive adhesives are used for a wide range of on-skin applications in the form of a medical tape or an adhesive film that can instantly adhere to medical devices (e.g., intravenous catheters, electrocardiogram sensors, and defibrillators). By allowing easy repositioning and removal of surgical equipment and analytical devices, pressure-sensitive adhesives minimize tissue damage and enhance surgical adaptability compared to sutures or non-reversible medical glues (FIG. 4A and Table 1), thereby reducing complications during complex surgical procedures.Table 1. Comparison of adhesive mechanism and target substrates with wet tissue adhesives and commercial / clinical pressure-sensitive adhesives. EDC, l-ethyl-3 -(-3 -dimethylaminopropyl) carbodiimide hydrochloride; NHS, N-Hydroxysuccinimide.Adhesion onA,rFault- ,, - wet , | ,. . , . . Adhesion , , , , •n-u-i-k tolerantA. . . Interfacial hydrophobic Reversibility , .. . Interact- .. . on3. .3a. . . demon-Adhesive , adhesion . , , surfaces (single- Repositioning , ion type . . internal ... . ' , .astration on mechanism without cycle , . . . . org aans , biological surface .... ,. tissuaes activationDry double-sided Chemical EDC-NHS 0 X X X X tape chemistryTriggerable Chemical EDC-NHS 0 X 0 X X detachment chemistryBioAdheSil Chemical Siloxane 0 0 X X X coupling agentFault-tolerant Chemical Catechol O X 0 0 0 hydrogel tapes chemistryElectroadhesion Physical Electro- 6 X 6 X X adhesionPhotodetachable Physical Topohesion O X 0 X XAdhesivesCommercial Physical Mechanical X X 0 0 XPSAs (e.g., VHB) interlocking and van derWaalsClinical PSAs Physical Mechanical X X 0 0 X(e.g., Micropore) interlocking and van derWaalsPSB Physical Mechanical 0 0 0 0 0 interlocking and van derWaals

[0138] Pressure-sensitive adhesives have been utilized in applications limited to skin or nonmedical surfaces due to their physical adhesion mechanism. On macroscopic length scales, a combination of soft and viscoelastic material properties enables bulk deformation and flow upon the application of light pressure, ensuring intimate contact with dry substrates for optimal and instant wetting. At the interface, various physical interactions (van der Waals and mechanical interlocking) acts to enable adhesion independent of surface chemistry for attachment to both hydrophilic and hydrophobic substrates. Upon detachment, these physical interactions can also be reversibly formed, enabling multiple uses if the adhesive surface is not blocked by any foreign material (e.g., dust). However, when pressure-sensitive adhesives are applied on a wethydrophilic substrate such as wet internal organs, water on the tissue surface acts as an interfacial barrier between the adhesive and tissue during application, resulting in low adhesion. This barrier impedes the physical interactions between the adhesive and underlying tissues that are crucial for pressure-sensitive adhesion.

[0139] In contrast, wet tissue adhesives have faced obstacles in adhesion to hydrophobic medical equipment without prior surface activation. Many reported tissue adhesives are formulated on the basis of hydrophilic matrices and chemical interactions, restricting the surface chemistries to hydrophilic interactions targeting tissue-specific moieties based on covalent (e g., primary amine groups via carbodiimide crosslinking chemistry) and ionic bonds. This has posed specific requirements of surface chemistry, casting limitations on the application of diverse surfaces without the conjugation of the corresponding functional groups on the bonding substrates (Table 1). For adhesion on diverse surfaces, a chemical functionalization step is needed prior to clinical use, resulting in poor shelf stability and time-consuming multi-step application procedures with specialized equipment and reagents (e.g., UV generator, specialized glue mixing syringes, photoinitiators).

[0140] In addition, many chemical interactions from covalent bonds are irreversible and impede repositioning due to tissue damage during removal. However, many surgical settings may require removal and potential repositioning of equipment and medical devices. To enable removal after use, reversible (single-cycle) and repositionable wet tissue adhesives have been developed with a variety of detachment processes, such as triggering solutions and UV light. However, these detachment processes may face difficulty in practical use on dynamic organs (long detachment times, harsh detachment conditions) and cannot be repositioned on hydrophobic surfaces due to constraints on the surface chemistry (Table 2).Table 2. Comparison of performance metrics with wet tissue adhesives and commercial / clinical pressure-sensitive adhesives for integrating medical devices and surgical equipment.Interfacial. .. . Adhesion Repositioning adhesion on Detachment Detachment timeAdhes ve ~ ° time (s) cycles porcine skin (Jm- trigger (s)2)Dry double-sided ~ 5 N.A. 710 N.D. NA tapeTriggerable ~ 5 1 - 400 0 (Sodium 300 detachment bicarbonate solution)BioAdheSil 900 N.A. 1100 N.D. N.A.Fault-tolerant < 10 ~ 20 1268 X < 10 hydrogel tapesElectroadhesion 20 N.A. N.R. 0 (DC voltage) 10. Photodetachable . 3000. X . - 75. 0 (UV) . 180.AdhesivesCommercial PSAs i - 5 N.A. 20 X 1 - 5(e.g., VHB)Clinical PSAs (e.g„ 1 - 5 N.A. 30 X 1 - 5Micropore)PSB < 1 75 380 X 1 ~ 5N.A. Not applicable.N.D. Not detachable. N.R. Not reported.

[0141] In this work, a pressure-sensitive bioadhesive (PSB) is presented that can achieve repositionable adhesion on substrates without limitations to surface chemistry (both hydrophilic and hydrophobic surfaces) via a physical, pressure-sensitive adhesion mechanism. The PSB physically bonds to wet internal organs by absorbing interfacial water and subsequently adheres using its viscoelastic properties (FIG. 4A). By enabling instant (~ls) adhesion, the PSB is highly suitable for application on dynamic internal organs, in which their physiological activity may shift the application site during the time required for adhesion formation. Furthermore, the PSB can integrate with pre-existing medical devices to act as a bridging material from surgical tools and medical devices with internal tissues (FIG. 4B). Furthermore, the PSB demonstrates a combination of high repositionability (75 cycles) and immediate adhesion (~1 s) without limitations to surface chemistry (FIG. 4C-4D). The repositionable adhesion can provide atraumatic detachment and reattachment on internal tissues, reducing the risk of potential tissue damage and further assisting the surgeon for in situ adaptable positioning during surgical procedures (FIG. 4E).

[0142] 2. Experimental Section / Methods

[0143] PGS-co-PEG and PGS synthesis'. The PGS-co-PEG was synthesized by a two-step polycondensation reaction. The first step involved the condensation polymerization of sebacic acid (133 mM, Sigma Aldrich) and 1,000 PEG (33 mM, Sigma Aldrich) at 130 °C under a nitrogen atmosphere for 8 h, followed by the reaction under a vacuum of 50 mTorr for 16 h. The reaction was conducted in a heating mantle and the vacuum was fit to the reactor vessel with a screw cap containing hose connectors. In the second step, glycerol (67 mM, Sigma Aldrich) wasadded, and the reaction was continued at 130 °C under the flow of nitrogen for 8 h followed by the reaction under a vacuum of 50 mTorr for 60 h to 100 h. The reaction was terminated when the PGS-co-PEG reached a specified rheological value which corresponded to the reaction time of 75-90 h (0 h is set when the vacuum is started in the second step of the condensation reaction). PGS was synthesized in a one-step condensation reaction with sebacic acid and glycerol in a 1 : 1 molar ratio.

[0144] PSB preparation: The PSB was prepared by removing the synthesized PGS-co-PEG from the reactor container and placing it on top of a backing material (0.125 mm thick polycarbonate; CT3O13O3, Goodfellow). Polytetrafluoroethylene (PTFE, 0.1mm thick; FP3O13OO, Goodfellow) was then placed on top of the PGS-co-PEG as a release liner. To create a flat and uniform PSB, the PSB was pressed under 10 MPa at 120 °C by using a hot press (QM900L, QMESYS) for 10 min and cooled in a freezer (-20 °C). During the hot press procedure, 1 mm thick spacers were utilized to control the final thickness of the PSB. Samples were cut to sizes of 2.5cm x 4 cm for 180-degree peel tests and 2.5cm x 2.5cm for cyclic tensile tests. The prepared PSB were stored with desiccants at -20 °C before use.

[0145] Rheological characterization: For rheological characterization of the PSB, frequency sweep tests (0.1Hz to 100 Hz, 0.1% strain) were conducted with a rheometer (Anton Paar MCR302, 25-mm parallel plate geometry). The temperature was set to 37 °C and held for 2 min to emulate biological conditions. For tack adhesion tests, an 8-mm parallel plate probe was pressed for 1 second and pulled upward at 10 mm s’1.

[0146] Chemical characterizations. Attenuated total reflection Fourier transform infrared (ATR- FTIR) spectra of the PSB were scanned with a spectrometer (Nicolet iS50, Thermo Fischer) in the range of 400 cm’1to 4,000 cm’1with a resolution of 4 cm’1. The spectra were acquired with 16 scans. Proton nuclear magnetic resonance ('ll NMR) spectroscopy was performed with a spectrometer (AVHD-400, Bruker) at 298 K. The PSB was dissolved in deuterated chloroform (CDCh) and the chemical shift was referenced to 8 7.24 ppm. The molar ratio between PEG and PGS in the PSB was calculated from the quantitative peak integration of methylene moiety of PEG at 8 1.30, 1.62, and 2.35 ppm and methylene moiety of sebacate at 8 3.64 ppm. Differential scanning calorimetry was conducted after nitrogen purging at a scan rate of 10 °C per minute from -100°C to 200°C. Crosslinking density was calculated with the Flory-Rehner equation shown below.

[0147] Vrrepresents the volume fraction of PSB after swelling. Vrwas calculated as shown below:

[0148] Vorepresents molar volume of the solvent. A value of y0= 18.07 cm^mol'1was utilized for deionized water, f represents functionality of the crosslinking point. A value of f= 3 was utilized as glycerol is trifunctional. / represents the Flory -Huggins interaction parameter. A value of / =0.426 was utilized according to references on PEG-hydrogels.

[0149] Ex vivo tissue preparation: Freshly harvested porcine skin and lung tissue were purchased from a local slaughterhouse in Daejeon. Porcine skin samples were cut into sizes of 2.5cm x 4 cm for 180-degree peel tests and 2.5cm x 2.5cm for cyclic tensile tests. A polymethylmethacrylate (PMMA) film was attached to the fatty side of porcine skin with cyanoacrylate glue (Loctite). Porcine skin samples were covered with PBS-soaked tissues and stored at -20 °C before use. Before mechanical testing, tissue samples were thawed and cleaned with 70 % isopropyl alcohol.

[0150] Mechanical testing: 180-degree peel tests (ASTM F2256-05) were performed with a universal testing machine (UTA-500N, Yeonjin). A crosshead speed of 1,500 mm min'1after pressing for 5 s at 5 N was used otherwise mentioned. For experiments that evaluated adhesion on wet porcine skin, deionized water was pipetted on the surface prior to testing. For cyclic tensile tests and self-healing measurements (LS-1E, Lloyd), a crosshead speed of 300 mm min'1after pressing for 5 s at 5 N was used, with no rest periods between consecutive cycles. Between consecutive cycles, a layer of interfacial water was maintained on the tissue surface by directing a humidifier to the surface of the porcine skin. The tissue surface was not cleaned between test cycles. When testing the duration of PSB application, a crosshead speed of 2000 mm min'1was used, and detachment was visually confirmed. For the mechanical testing of blood-covered porcine skin, defibrinated sheep blood was obtained from KisanBio. During cyclic underwater testing, a crosshead speed of 300 mm min'1after pressing for 3 s at 1 N was used.

[0151] Interfacial water absorption imaging: Interfacial water absorption was evaluated by attaching PSB or control VHB tape samples to a 2 wt% agarose hydrogel loaded with rhodamine-dye for 2 hrs. The interface was nondestructively imaged by taking consecutive z- stack images through the hydrogel-tape interface using a confocal microscope. The acquired images were projected on the x or y axis to visualize the rhodamine-dye diffusion as a measure of interfacial water transport and thresholded to the same intensity to find the hydrogel -tape interface.

[0152] In vitro biodegradability: In vitro biodegradability of the PSB was evaluated by incubation in PBS at 37 °C. To visualize the degradation, the PSB samples (10 mm in width and 10 mm in length, approximately 1 g) were attached to slide glasses and immersed in 50 ml conical tubes filled with 50 ml PBS. Then, the conical tubes were placed horizontally and incubated in an incubator at 37 °C (WB-6, Daihan Scientific Co., Ltd). Samples were collected after 0, 0.5, 1, 2, 4, 10, 25, and 50 h and dried in an oven at 65 °C for 24 hrs. The weight loss of the samples was calculated based on the following equation: Weight ratio 100where Wo was the initial weight of the PSB sample in the dry state and Wd was the degraded weight of the PSB sample in the dry state.

[0153] In vitro cell viability: In vitro cell viability was evaluated by Live / Dead assay based on the PSB-conditioned cell culture media. Control cell culture medium was prepared with 10 % of fetal bovine serum (Gibco) and 1 % of penicillin-streptomycin (Sigma-Aldrich) in Dulbecco’s modified Eagle medium (DMEM). The PSB-conditioned cell culture medium was made by incubating 80 mg of the PSB in 4 ml of the cell culture medium and neutralizing its pH to 7.HEK 293 cells and 3T3 cells were plated on 96-well plates with a control cell culture medium at 37 °C for 24 h in 5 % CO2. The plated cells were then incubated with the PSB-conditioned cell culture medium at 37 °C for 24 h in 5 % CO2. In vitro cell viability was measured by a Live / Dead viability / cytotoxicity kit for mammalian cells (L3224, Thermo Fisher Scientific). Solution with calcein-AM 5 pl and ethidium homodimer- 1 20 pl in 10 ml DPBS was used for Live / Dead cell staining. Images were collected using a fluorescence microscope. The cell viability result was calculated by a microplate reader (SpectraMax iD3, Molecular Devices)using excitation / emission wavelengths of 485 nm / 530 nm and 530 nm / 645 nm for live cells and dead cells, respectively.

[0154] Vertebrate animal subjects: Female Sprague Dawley rats (225-250 g, 12 weeks, Koatec) and minipigs (originated from Korean Jeju Island's native pig, 23-25 kg, 6 months, Cronex) were used in this work. All rat studies were reviewed and approved by the Korea Advanced Institute of Science and Technology (KAIST) Institutional Animal Care and Use Committee (KA2024- 057). All porcine studies were reviewed and approved by the CRONEX Institutional Animal Care and Use Committee.

[0155] In vivo biocompatibility: In vivo biocompatibility of the PSB was evaluated based on dorsal subcutaneous and epidermal implantation for 1 week. The PSB samples were prepared with styrene-ethylene-butylene-styrene (SEBS) or Ecoflex backing instead of polycarbonate to minimize mechanical mismatch with tissues. For subcutaneous implantation, the PSB with SEBS backing and control (SEBS sample without PSB) were cut into 10 mm x 10 mm. For epidermal implantation, the PSB with Ecoflex backing and control (Ecoflex sample without PSB) were cut into 10 mm x 10 mm. Anesthesia was conducted via an isoflurane inhalation (4 % for induction, 2-3 % for maintenance in 02). After anesthesia, hair on the dorsal region was removed. Surgery was continued on the heating pad for temperature control. Then, a 2 cm skin incision was made on the back of the animal. A subcutaneous pocket was made in the incision by blunt dissection scissors. The PSB with SEBS backing or PSB with no backing material was implanted in the subcutaneous pocket with the PSB side facing the inner tissue. The PSB with Ecoflex backing was sutured on the epidermis 2 cm distally. The same implantation procedures were conducted for the control samples. After 7 days post-implantation, the animals were euthanized and the implanted region was harvested. Harvested tissue was washed with PBS and fixed with formalin for 28 h. Fixed tissue was stained with H&E for histological analysis.

[0156] In vivo rat epicardial ECG recording: For in vivo epicardial ECG recording, a PSB with one electrode (10 pm thick stainless-steel foil, MTI Korea) was prepared in the size of 2 mm x 5 mm. The electrode acted as the recording electrode and was connected by silver paste to a tungsten wire. Anesthesia was conducted via an isoflurane inhalation (4 % for induction, 2-3 % for maintenance in O2). Surgery was continued on the heating pad for temperature control. Endotracheal intubation was performed to connect the animals to a mechanical ventilator. Then, an incision was made to expose the heart. After exposure of the heart, the pericardium wasremoved by fine forceps. The PSB with electrode was applied to the epicardial surface for ECG recordings. During the ECG recording, the device was repositioned two times until the desired ECG waveform was recorded.

[0157] In vivo rat lung stabilization.' For in vivo lung stabilization, a PSB was attached to a silicon tubing (I.D. 1.5mm, O.D. 2.5 mm, DAIHAN Scientific). The silicon tubing acted as the vacuum tubing for stabilizing tissue with negative pressure and was connected to mini pump (ZR370-02PM, adafruit). Anesthesia was conducted via an isoflurane inhalation (4 % for induction, 2-3 % for maintenance in 02). Surgery was continued on the heating pad for temperature control. Endotracheal intubation was performed to connect the animals to a mechanical ventilator. Then, an incision was made to expose the lung. After exposure of the lung, the PSB attached tubing was applied to the lung surface for stabilization.

[0158] In vivo rat sciatic nerve stimulation: For in vivo sciatic nerve stimulation, a PSB with two electrodes (10 pm thick stainless-steel foil, MTI Korea) was prepared in the size of 1 mm x 2 mm similar to the ECG recording device. Anesthesia was conducted via an isoflurane inhalation (4 % for induction, 2-3 % for maintenance in O2). Surgery was continued on the heating pad for temperature control. The sciatic nerve was exposed by dissecting the vastus lateralis muscle and biceps femoris muscle of the animal. The PSB with electrodes was applied to the sciatic nerve. During the sciatic nerve stimulation, the device was repositioned two times to different locations to stimulate diverse sciatic nerve bundles. Biphasic current pulses (1 Hz, 50 pA) were used to stimulate the sciatic nerve.

[0159] In vivo porcine epicardial ECG recording: The minipigs were intubated and anesthetized via inhalation of a 3% isoflurane-oxygen mixture and placed in a supine position. The minipigs were monitored throughout the entire procedure with ECG and oxygen saturation during entire procedures. To expose the heart and surrounding lungs of the minipig, the left ribs (numbers 6-8) were spread as far apart as possible, and the 7th rib was removed. The PSB with two electrodes (10 pm thick stainless-steel foil, 2 mm * 10 mm sized, MTI Korea) was placed on a left ventricle of the minipigs, and ECG recording and analysis were conducted and analyzed in the same method as the rat experiments.

[0160] In vivo porcine lung strain measurement: A strain sensor was fabricated based on the screen printing of an Ecoflex-CNT composite (CNT 2 wt. %). A mold was fabricated via laser patterning (Universal VLS 3.50) on a polycarbonate sheet (GFM, 0.1 mm). Ecoflex layers werecoated for the insulation of the Ecoflex-CNT composite. Contact pads were electrically connected with insulated copper wires and sealed with epoxy for mechanical stability. Strain sensor impedance was measured with an LCR meter (1 V, 1 kHz, interval 0.5s) while undergoing mechanical cycling (LS-1E, Lloyd) of various degrees of strain (1 %, 3 %, 5 %) to verify strain measurement. The PSB with a strain sensor was placed on the left lung lobe surface, and the electrode was then removed and reattached to the area around the lung lobe surface where it was first attached, measured, and then removed and reattached two times. During in vivo porcine testing, strain sensor impedance was measured with an LCR meter (1 V, 1 kHz, interval 0.5 s) and the breathing rate was controlled by a ventilator connected via endotracheal intubation.

[0161] Statistical Analysis: OriginPro was used for all statistical data analysis. Data distribution normality was confirmed with Shapiro-Wilk normality tests and homogeneity of variance was confirmed with an F-test between two sample groups. Data was tested to be normal for all data distributions. If homogeneity of variance was confirmed, a two-sample t-test was utilized to compare the two groups. Otherwise, Welch’s t-test was used to compare the two groups.

[0162] 3. Results

[0163] 3.1. Design and Synthesis of the Pressure-Sensitive Bioadhesive

[0164] The PSB utilizes an amphiphilic block copolymer entanglement network composed of biocompatible polymers (FIG. 4F). The hydrophilic polyethylene glycol (PEG) moiety was chosen to absorb the interfacial water between the adhesive and wet surfaces. Poly(glycerol sebacate) (PGS) was introduced as a hydrophobic backbone to tune the entanglement network. To synthesize a pressure-sensitive adhesive, the viscoelasticity of the PSB’s amphiphilic block copolymer was optimized by tuning the reaction time. A highly scalable, two-step polycondensation reaction was employed to form the entanglement network of the PSB (FIG. 5). The synthesis and reaction progression of the crosslinked polymers was confirmed with nuclear magnetic resonance analysis and Fourier transform infrared spectroscopy (FIGS. 6-8). The PSB demonstrated a glass transition temperature of -56.3 °C, falling within the range of previously reported PSAs (i.e., -20 to -60 °C), and a shifting glass transition temperature indicative of crosslinking (FIGS. 9A-9B). Furthermore, the crosslinking density of the PSB was measured at 6.40 E-3 mol cm'3± 0.04 E-3 mol cm'3, and indicates a reduced crosslinking density in comparison with previously reported PGS elastomers.

[0165] During the bonding phase of the PSB, the hydrophilic PEG moiety rapidly absorbs and removes the interfacial water to facilitate direct physical contact on the wet tissue surface (FIG. 4G, top panel). The viscoelasticity of the PSB then further promotes the conformal contact with the tissue surface by viscoelastic flow, maximizing the physical bond formation (e.g., mechanical interlocking and van der Waals interactions). During the debonding phase, the PSB’s toughness and viscoelastic nature enable interfacial detachment rather than cohesive failure and subsequent repositionable adhesion to the tissue surface (FIG. 4G). Notably, the physical adhesion mechanism of the PSB allows repeated re-application and repositioning to wet surfaces similar to the mechanism of pressure-sensitive adhesive on dry surfaces (FIG. 10).

[0166] 3.2. Mechanism and characterization of the wet viscoelastic adhesion

[0167] Optimal viscoelasticity is a critical requirement for the PSB to be characterized as a pressure-sensitive adhesive. As reaction time progresses, the crosslinking density of the PSB increases as seen by the increasing elastic modulus, and the macroscopic behavior of the amphiphilic block copolymer gradually transforms from viscous fluid to viscoelastic solid (FIG. HA and FIG. 12).

[0168] To assess the bonding and debonding behavior of the PSB, the rheological properties were evaluated at frequencies corresponding to the bonding phase (corresponding to the complex modulus at 1 Hz) and the debonding phase (corresponding to the complex modulus at 100 Hz) (FIG. 1 IB). The optimized viscoelasticity of the PSB demonstrates a combination of good flowability for instant adhesion (viscous modulus higher than elastic modulus during the bonding phase) and elasticity for reversible adhesion (viscous modulus lower than elastic modulus during the debonding phase) in comparison to the viscous fluid and viscoelastic solid samples. In addition, the PSB satisfies the Dahlquist criterion (elastic modulus < 1 C at 1 Hz), a critical parameter for pressure-sensitive adhesion to form robust yet reversible adhesion (FIG. 13A). Optimal adhesion performance can be achieved at a loss factor range of 1.6-2.0 at 1 Hz (FIGS. 13B-C). Due to the strict requirements in the complex modulus necessary for pressure-sensitive adhesion, only samples in this loss factor range are referred to as the PSB. Notably, the optimized PSB shows a tissue-like elastic modulus (~ 100 Pa at 1 Hz). The low elastic modulus is expected to help mitigate potential tissue damage from elastic deformation during detachment.

[0169] To clearly visualize that the PSB is absorbing the water at the adhesive-tissue interface, the PSB was adhered to a 2 wt% agarose hydrogel loaded with a fluorescent dye (FIG. 11C). Theinterfacial water absorption was visualized by the transport of the dye from the bulk of the agarose hydrogel into the adhesive layers. As commercial PSAs cannot absorb the interfacial water, the transport of the fluorescent dye is trapped at the adhesive-hydrogel interface. In contrast, the PSB demonstrates transport of the fluorescent dye into the bulk material, demonstrating its ability to absorb interfacial water. In addition, to demonstrate that the wet adhesion mechanism of the PSB on wet tissues relies on a combination of water absorption and viscoelasticity, the tensile adhesion was compared (ASTM F2258-05) of the PSB with 1) a viscoelastic polymer without water absorption capabilities (i.e., pristine PGS without the hydrophilic PEG moiety), 2) an elastomeric polymer with water absorption (elastomeric PGS-co- PEG), and 3) an elastomeric polymer with no water absorption capabilities (hydrophobic PGS). By introducing the hydrophilic PEG moiety to the PGS backbone, the water absorption capability increased to ~9 wt. % in both the PSB and elastomeric PGS-co-PEG, while the hydrophobic PGS polymer demonstrated no water absorption (FIG. 14A and FIG. 15A). The hydrophobic PGS polymer group had no adhesion to wet porcine skin during tensile testing, while the PSB demonstrated robust adhesion on wet porcine skin (FIGS. 15B-15C). Though elastomeric PGS-co-PEG absorbs water, elastomeric PGS-co-PEG also showed no adhesion to wet porcine skin due to its lack of viscous flowability. Furthermore, while viscoelastic PGS showed adhesion on dry porcine skin, there was a significant drop in adhesion for viscoelastic PGS when water was applied on the surface of the skin, unlike the PSB. This indicates that a combination of viscoelasticity and interfacial water absorption was key for achieving robust, instant adhesion on wet surfaces.

[0170] To enable reversible adhesion cycles, the adhesive should be resistant to bulk damage from permanent deformation induced during its detachment process. The self-healing capability was verified by attaching two separate pieces of PSB and measuring the force required to separate the two after self-healing (FIGS. 14B-14C). The PSB demonstrates rapid self-healing characteristics to recover from bulk damages upon the application of light pressure, allowing it to maintain mechanical integrity and adhesion performance during multiple detachment-attachment cycles.

[0171] To investigate the debonding mechanisms of the optimized PSB with that of commercially available pressure-sensitive adhesives, the debonding process is monitored on both dry and wet surfaces (FIGS. 16A-16B). On dry surfaces, both VHB (i.e., commercially availablepressure-sensitive adhesive) and PSB exhibit instant, robust, and reversible adhesion with noticeable formation of cavitation and fibrillation, which are key characteristics of pressuresensitive adhesives during the debonding process. In contrast, VHB shows virtually no adhesion on wet surfaces whereas the PSB provides nearly the same debonding characteristics as on the dry surface.

[0172] The PSB harnesses its distinct adhesive properties on internal organs. Application on an ex-vivo porcine heart further demonstrates the robust attachment and atraumatic reattachment of the PSB in comparison with commercial and clinical pressure-sensitive adhesives (FIGS. 17A- C). To verify the instant and reversible adhesion of the PSB, we evaluate its adhesion in comparison with commercial (e.g., VHB and FlexSeal) and clinical (e.g., Transpore and Micropore) pressure-sensitive adhesives on hydrophilic substrates with a cyclic tensile test (FIG. 17D). The PSB demonstrates robust adhesion to both wet lung and heart tissue on five adhesive cycles in comparison to commercial and clinical pressure-sensitive adhesives (FIGS. 17E-17F). While commercially available pressure-sensitive adhesives also exhibit consistent tensile strength and adhesion energy across test cycles, the tensile strength (2 kPa for VHB vs. 18 kPa for the PSB) and adhesion energy (0.19 J m’2for VHB vs. 97 J m’2for the PSB) of the PSB were over an order higher than VHB on wet porcine skin. The relatively low adhesive energy in lung and heart tissue, in comparison to skin tissue, results from the high energy dissipation and elongation in the tissue samples during testing. Furthermore, the PSB enables adhesion on blood- covered porcine skin (7.8kPa) in comparison with VHB (2kPa) to enable attachment on blood- covered tissues (FIG. 18). The PSB further demonstrates robust adhesion in pseudo-wet environments (i.e., when the tissue is exposed to air) for up to 25 cycles on porcine skin in comparison with VHB (FIGS. 19A-19C). To test the usage duration of the PSB, additional cyclic adhesion tests in pseudo-wet environments were conducted on porcine skin. The PSB demonstrated stable adhesion for 1 hour, displaying stable adhesion for 75 cycles on porcine skin in pseudo-wet environments (FIG. 20). To further test the effects of water absorption on cyclic adhesion as the PSB swells, underwater adhesion tests were conducted (FIGS. 21A-21B). As the PSB swells, peak tensile strength decreases due to the presence of excess water and reaches an equilibrium at 50% of the tensile strength of the initial adhesion cycle.

[0173] To verify the substrate-independent and instant adhesion of the PSB, 180-degree peel tests (ASTM F2256-05) were conducted on diverse medical substrates and internal organs with 5seconds of application. Owing to the physical bonding mechanism, the PSB achieves instant and robust adhesion to diverse hydrophilic and hydrophobic substrates such as internal organs (interfacial toughness over 110 J m‘2for heart; 100 J m'2for lung; 380 J m’2for skin) and engineering solids commonly used in medical devices (interfacial toughness over 720 J m'2for polystyrene; 630 J m'2for polydimethylsiloxane; 940 J m-2 for stainless steel; 640 J m'2for polyurethane) (FIGS. 22A-22B and FIGS. 23A-23D). This allows the use of diverse backing materials, including stretchable (styrene-ethylene-butylene-styrene (SEBS) and Ecoflex), flexible (polyurethane), and stiff (polycarbonate) substrates (FIGS. 24A-24B) Interfacial failure was validated with visual inspection and measurement of PSB tensile toughness values (FIG. 25). Notably, the interfacial adhesion of the PSB saturated as low applied pressure and pressing times during peel testing (FIGS. 26A-26C). This is seen to occur due to the low modulus of the PSB, in combination with the conformal contact induced by its good flowability.

[0174] The synthesized PSB enables versatile fabrication into various forms including singlesided or double-sided tapes, depending on the configuration of the application (FIGS. 27A-27B). When the PSB is utilized in a single-sided configuration or double-sided configuration it is subsequently referred to as the sPSB or dPSB, respectively.

[0175] 3.3. Surgical tissue stabilization on ultra-soft organs

[0176] Clinical tissue stabilizers (e.g., Octopus Tissue Stabilizers, Medtronic) have been widely utilized for the mechanical stabilization of dynamically moving tissues (e.g., heart and lungs), allowing surgeons to operate on a stable, localized area without needing to stop the moving organ. During the tissue stabilization of internal tissues, vacuum suction is utilized to mechanically stabilize tissue (FIG. 28A). The suction interface may be susceptible to detachments due to the mechanical mismatch between the tissue and the vacuum hose, making an interfacing adhesive beneficial in mechanically bridging the tissue (FIG. 28B). Moreover, the strong suction pressure required during suction or gripping force applied by forceps often results in tissue damage on soft internal organs such as the lungs.

[0177] The dPSB’s unique adhesive properties for substrate-independent, instant, and reversible adhesion enables three key functionalities as an interfacing adhesive in tissue stabilization: 1) applicability to diverse surfaces including tissue stabilization devices and tissues, 2) instant attachment on dynamic internal organs, and 3) on-demand atraumatic detachment and reattachment for tissue manipulation (FIG. 28C). The dPSB demonstrates substrate-independentadhesion capable of stabilizing a rodent lung during respiration (FIG. 28D). We compare the tissue stabilization of the dPSB by measuring the negative pressure on the suction tubing with and without the dPSB at the same suction pressure (FIG. 28E). The integration of the dPSB at the distal end of the tissue stabilizer significantly improves stabilizing pressure and enables instant attachment to the lung during lung respiratory movements. While higher suction pressure may mitigate such problems, this approach is undesirable for soft and fragile organs, such as the lungs, due to associated tissue damage and potential post-surgical complications. Furthermore, the integration of the dPSB acts as an atraumatic interface by reducing the minimum pressure required for lung stabilization and enabling atraumatic detachment and subsequent reattachment (FIG. 28F).

[0178] Prior to application on internal organs, the in vitro and in vivo biocompatibility of the PSB was evaluated. An in vitro LIVE / DEAD assay based on human embryonic kidney cells (HEK 293) and mouse embryonic fibroblast (3T3) cells show that the cytotoxicity of the PSB is comparable to that of the control group (FIGS. 29A-29B). Subsequently, the PSB was implanted in the rat dorsal subcutaneous and epidermal implantation for 7 days. H&E staining data and lesion analysis conducted by a blinded histopathologist showed that the subcutaneously- implanted PSB-SEBS induces only a mild inflammatory reaction with mild inflammatory cell infiltration and fibrosis after 7 days, comparable to that of the control group (SEBS) (FIGS. 30A- 30B, FIG. 31). During epidermal biocompatibility testing, the PSB with Ecoflex backing and the control group (Ecoflex) are applied on the dorsal epidermis. Evaluation of epidermal biocompatibility by a blinded histopathologist demonstrate that both groups do not exhibit observable signs of inflammation, through analysis on the epidermal thickness, inflammatory cell infiltration and fibrosis of the surrounding tissue (FIGS. 32A-32B, FIGS. 33A-33B).

[0179] The PSB further demonstrated hydrolytic degradation biodegradation with 75% of its mass degraded within 48 hours (FIGS. 34A-34B). The biocompatibility of the biodegraded PSB was further validated through in vitro cell viability (FIGS. 29A-29B) and in vivo subcutaneous implantation (FIG. 30C). When only the PSB is implanted, the PSB exhibits negligible foreign body response due to its biocompatibility, fast biodegradation, and tissue-like mechanical properties (FIG. 30C).

[0180] 3.4. Fault-tolerant bioelectronic interfacing on dynamic organs

[0181] To demonstrate in situ adaptable integration of implantable electrodes for bioelectronic stimulation, the sPSB was utilized to integrate an electrode on a rat sciatic nerve in vivo (FIG. 35 A). The sPSB enables probing multiple stimulation sites for precise placement with quick, atraumatic removal upon the end of the procedure. Enabled by the atraumatic and repositionable adhesion of the sPSB, the same stimulation electrode can be utilized to probe the transitions between upward- and downward-evoked hindlimb movements by repositioning the electrode to various locations in the sciatic nerve (FIG. 35B).

[0182] To demonstrate fault-tolerant integration of implantable electrodes for bioelectronic sensing, the sPSB is utilized to integrate an electrode to rat left ventricular epicardial tissue in vivo (FIG. 36A). Upon the initial application of the epicardial electrode, the recorded electrocardiogram (ECG) does not display distinct left ventricular ECG signals due to the dynamic movement of the heart and subsequent misplacement of the electrode (FIG. 36B). The sPSB allows repeated atraumatic detachment and repositioning of the electrode to the epicardial tissue until the ECG waveform with a distinct left ventricular QRS complex, corresponding to standard 12-lead ECG signals, is achieved (FIGS. 36B-36C).

[0183] The applicability of sPSB-enabled applications is further demonstrated in a large-animal porcine model (FIGS. 37A-37B). An epicardial electrode could be integrated into a porcine heart in vivo by the sPSB and repeatedly repositioned until the measured ECG waveform with the targeted left ventricular QRS complex was achieved (FIG. 37C). The integrated electrode by the sPSB exhibited stable adhesion and recording capability under the dynamic motion of the heart (FIG. 37D). Rapid and atraumatic repositioning of the sPSB to soft and fragile organs such as the porcine lungs was also demonstrated (FIG. 37E). The sPSB’s capability to effectively integrate and atraumatically reposition implantable devices on delicate organs allows in situ adaptive monitoring of strain changes in dynamic organs (FIGS. 37E-37F). A custom-made strain sensor was integrated to a porcine lung by the sPSB and subsequently repositioned to different locations to obtain high-fidelity monitoring of lung inflation and deflation at various preset inhalation rates and locations without causing any tissue damage or use of new devices (FIG. 37G, FIGS. 38A- 38D, and FIGS. 39A-39C).

[0184] 4. Discussion and Conclusion

[0185] Clinical pressure-sensitive adhesives are widely formulated from polymers (e.g., acrylics, styrene block copolymers). By crosslinking the polymers, an entangled network is formed with viscoelastic properties, which can be characterized by the complex modulus, elastic and viscous components. Under light pressure, the elastic modulus of the viscoelastic polymer lower than the Dahlquist criterion enables the deformation and conformal contact of the polymers to the target substrate. For robust, reversible adhesion in pressure-sensitive adhesives, the viscous modulus should be on the order of the elastic modulus. This complex modulus enables instant and reversible adhesion to diverse dry surfaces such as medical devices and the skin.

[0186] While these clinical pressure-sensitive adhesives can adhere to hydrophobic tissues with low water content (15-30%), they cannot adhere to wet hydrophilic tissues, such as internal organs, due to their high-water content (e g., -83% water composition for lung and -74% for heart). In addition, the surfaces of internal organs are constantly perfused with liquids from the bulk of the organ. Subsequently, a layer of water exists at the adhesive-organ interface, which has been shown to act as a critical barrier for forming chemical adhesive interactions.

[0187] The PSB as described herein introduces the unique advantages of pressure-sensitive adhesives to wet internal organs by materials design, optimization, and characterizations. Unlike many bioadhesives that require specific surface chemistry or composition to achieve wet adhesion, the PSB offers substrate-independent, instant, and reversible adhesion for surgical application. The distinct adhesive performance of the PSB makes it an ideal material for bridging internal organs and surgical equipment.

[0188] Two non-limiting applications were demonstrated herein of the PSB’s unique adhesion mechanism for atraumatic tissue stabilization and fault-tolerant integration of bioelectronic interfacing on dynamic organs in both small and large animal models. When utilized as a singlesided tape (sPSB), the PSB is composed of a bilayer structure by attaching any backing polymer to the PSB (e.g., polycarbonate, Ecoflex). The backing layer acts as a protective layer, covering the other side of the PSB. This plays a two-part role: it prevents indiscriminate adhesion to other internal organs or surgical tools, and it further allows for the application of pressure on top of the backing layer. When utilized as a double-sided tape, the PSB is able to act as an interfacing material between surgical tools and internal tissues to enable enhanced mechanical stability during interactions with internal organs.

[0189] The adhesive characteristics of the PSB demonstrate high utility in surgical and bioelectronic applications that interface with highly dynamic organs such as tissue stabilization and fault-tolerant adhesion during bioelectronic interfacing. Due to the highly wet, soft, and dynamic nature of internal organs, the precise integration of implantable devices and surgical tools on target tissues has a risk of misplacement that requires repositioning or re-application of the device to ensure desired functionality, especially in minimally invasive surgical settings for probing organs such as the heart and sciatic nerve. These applications can range from epicardial recording for identification of arrhythmias to deepening the understanding of mechanistic physiology in epicardial pacing to electrical stimulation of peripheral nerves for surgical placement of stimulators during open nerve surgery.

[0190] The PSB demonstrates promising clinical applications for dynamic, transient scenarios for immediate integration with a broad range of medical devices not limited to surface chemistry for applications such as atraumatic tissue stabilization, robust interfacing with commercial electrodes for clinical cardiology and the identification of arrhythmias and probing of stimulation sites for peripheral nerves stimulators during pain treatment. The soft tissue-like mechanical properties, in combination with biocompatibility and biodegradability, will further allow minimal long-term foreign body response if left inside the body after application.

[0191] Example 2

[0192] Methods

[0193] PGS-co-PEG andPGrS synthesis in different PEG ratios: The PGS-co-PEG was synthesized by a two-step polycondensation reaction. The first step involved the condensation polymerization of sebacic acid (133 mM, Sigma Aldrich) and molecular weight 1,000 PEG (3.3, 6.6, 19.8, 33, 39.6, or 46.2 mM for 5, 10, 30, 50, 60 or 70 % w / w to PGS, respectively, Sigma Aldrich) at 130 °C under a nitrogen atmosphere for 8 h, followed by the reaction under a vacuum of 50 mTorr for 16 h. The reaction was conducted in a heating mantle and the vacuum was fit to the reactor vessel with a screw cap containing hose connectors. In the second step, glycerol (67 mM, Sigma Aldrich) was added, and the reaction was continued at 130 °C under the flow of nitrogen for 8 h followed by the reaction under a vacuum of 50 mTorr for 60 h to 100 h. Tegaderm was used as a control commercial adhesive.

[0194] PGS-co-PEG and PGS synthesis in different ratios . The PGS-co-PEG was synthesized by a two-step polycondensation reaction. The first step involved the condensation polymerization ofSigma Aldrich) and molecular weight 1,000 PEG (33 mM Sigma Aldrich) at 130 °C under a nitrogen atmosphere for 8 h, followed by the reaction under a vacuum of 50 mTorr for 16 h. The reaction was conducted in a heating mantle and the vacuum was fit to the reactor vessel with a screw cap containing hose connectors. In the second step, glycerol (67 mM, Sigma Aldrich) was added, and the reaction was continued at 130 °C under the flow of nitrogen for 8 h followed by the reaction under a vacuum of 50 mTorr for 60 h to 100 h. Tegaderm was used as a control commercial adhesive.

[0195] PGS-co-PEG and PGS synthesis in different PEG ratios . The PGS-co-PEG was synthesized by a two-step polycondensation reaction. The first step involved the condensation polymerization of sebacic acid (133 mM, Sigma Aldrich) and molecular weight 1,000 PEG (33 mM, Sigma Aldrich) at 130 °C under a nitrogen atmosphere for 8 h, followed by the reaction under a vacuum of 50 mTorr for 16 h. The reaction was conducted in a heating mantle and the vacuum was fit to the reactor vessel with a screw cap containing hose connectors. In the second step, glycerol (33, 67, 100, or 133 mM, Sigma Aldrich) was added, and the reaction was continued at 130 °C under the flow of nitrogen for 8 h followed by the reaction under a vacuum of 50 mTorr for 60 h to 100 h. Tegaderm was used as a control commercial adhesive.

[0196] Pull-off adhesion force measurements. To compare the adhesion properties of the PGS- co-PEG with different PEG ratios, different dicarboxylic acids, and different ratios between sebacic acid and glycerol, each glue was placed on a metal probe and cured on a heart (epicardium) tissue for pull-off adhesion tests using a mechanical tester. Sliced tissue samples were fixed on a standard scanning electron microscopy (SEM) pin stub mount (012.7mm x 8mm pin height, Ted Pella Inc) by instant glue (Loctite 495, Henkel) and wetted with PBS. During the test, SEM pin stubs with tissue samples were loaded on the bottom stage of the mechanical tester (ADMET). The metal probe with each adhesive was pulled off with a constant strain of 8mm / min. The applied force was monitored, and the maximum force before the complete detachment of the patch was measured as the adhesion force.

[0197] Results

[0198] As shown in FIG. 40, PGS-co-PEG with 10-60% showed adhesion forces higher than the Tegaderm, and 70% showed adhesion forces similar to Tegaderm but significantly different in student t-test (p=0.0197).

[0199] As shown in FIG. 41, PSA with different dicarboxylic acids showed similar adhesion forces with PSA with PGS (PGS-co-PEG) (not significant in ANOVA, p=0.16).

[0200] As shown in FIG. 42, PSB with 4: 1, 2: 1, 4:3, and 1: 1 ratio (sebacic acid: glycerol) showed the adhesion force of 2.68 N / cm2, 7.18 N / cm2, 5.37 N / cm2, and 4.02 N / cm2, respectively, which are higher than the adhesion force of Tegaderm.

[0201] As used in this specification and the claims, the singular forms “a,” “an,” and “the” include plural forms unless the context clearly dictates otherwise.

[0202] As used herein, “about”, “approximately,” “substantially,” and “significantly” will be understood by persons of ordinary skill in the art and will vary to some extent on the context in which they are used. If there are uses of the term which are not clear to persons of ordinary skill in the art given the context in which it is used, “about” and “approximately” will mean up to plus or minus 10% of the particular term and “substantially” and “significantly” will mean more than plus or minus 10% of the particular term.

[0203] As used herein, the terms “include” and “including” have the same meaning as the terms “comprise” and “comprising.” The terms “comprise” and “comprising” should be interpreted as being “open” transitional terms that permit the inclusion of additional components further to those components recited in the claims. The terms “consist” and “consisting of’ should be interpreted as being “closed” transitional terms that do not permit the inclusion of additional components other than the components recited in the claims. The term “consisting essentially of’ should be interpreted to be partially closed and allowing the inclusion only of additional components that do not fundamentally alter the nature of the claimed subject matter.

[0204] The phrase “such as” should be interpreted as “for example, including.” Moreover, the use of any and all exemplary language, including but not limited to “such as”, is intended merely to better illuminate the invention and does not pose a limitation on the scope of the invention unless otherwise claimed.

[0205] Furthermore, in those instances where a convention analogous to “at least one of A, B and C, etc.” is used, in general such a construction is intended in the sense of one having ordinary skill in the art would understand the convention (e.g., “a system having at least one of A, B and C” would include but not be limited to systems that have A alone, B alone, C alone, A and B together, A and C together, B and C together, and / or A, B, and C together.). It will be further understood by those within the art that virtually any disjunctive word and / or phrasepresenting two or more alternative terms, whether in the description or figures, should be understood to contemplate the possibilities of including one of the terms, either of the terms, or both terms. For example, the phrase “A or B” will be understood to include the possibilities of “A” or “B” or “A and B.”

[0206] All language such as “up to,” “at least,” “greater than,” “less than,” and the like, include the number recited and refer to ranges which can subsequently be broken down into ranges and subranges. A range includes each individual member. Thus, for example, a group having 1-3 members refers to groups having 1, 2, or 3 members. Similarly, a group having 6 members refers to groups having 1, 2, 3, 4, or 6 members, and so forth.

[0207] The modal verb “may” refers to the preferred use or selection of one or more options or choices among the several described embodiments or features contained within the same. Where no options or choices are disclosed regarding a particular embodiment or feature contained in the same, the modal verb “may” refers to an affirmative act regarding how to make or use an aspect of a described embodiment or feature contained in the same, or a definitive decision to use a specific skill regarding a described embodiment or feature contained in the same. In this latter context, the modal verb “may” has the same meaning and connotation as the auxiliary verb “can.”

Claims

Claims1. A pressure-sensitive adhesive (PSA) for use on a surface, the adhesive comprising: an amphiphilic block copolymer, configured to provide repeatable adhesion to hydrophilic surfaces, comprising: a polyester comprising: a dicarboxylic acid and glycerol at a molar ratio of about 4: 1 to 1 :1, and a hydrophilic polymer, wherein a weight ratio of the hydrophilic polymer to the polyester is about 10 to 70 % w / w.

2. The adhesive of claim 1, wherein the polyester includes poly(glycerol sebacate) (PGS), poly(glycerol glutarate), poly(glycerol adipate), poly(glycerol pimelate), poly(glycerol suberate), poly(glycerol phthalic acid), or poly(glycerol dodecanedioate).

3. The adhesive of claim 1, wherein the hydrophilic polymer includes polyethylene glycol (PEG), poly(vinyl alcohol), poly(acrylic acid), polypropylene glycol), poly(2 -hydroxyethyl methacrylate), poly(N-vinylpyrrolidone), poly(methacrylic acid), poly(ethylene oxide), poly(aspartic acid), poly(acrylamide), poly(ethylenimine), dextran, starch-based polymers, cellulose derivates (e.g., hydroxyethyl cellulose, carboxymethyl cellulose), agarose, chitosan, alginate, hyaluronic acid gelatin, or pectin.

4. The adhesive of claim 1, wherein the adhesive forms an adhesion to a surface occurs in about 1 second.

5. The adhesive of claim 4, wherein the adhesion forms an interfacial toughness up to about 1 kJ / m26. The adhesive of claim 1, wherein the adhesive forms a reversible adhesion with a surface.

7. The adhesive of claim 6, wherein the adhesive may be reversibly adhered to the surface up to 1000 times.

8. The adhesive of claim 1, wherein the storage modulus G’ is less than 105Pa at 1 Hz.

9. The adhesive of claim 1, wherein the storage modulus G‘ is about 100 Pa at 1 Hz.

10. The adhesive of claim 9, wherein the storage modulus G‘ is about 1000 Pa to 10000 Pa at 100 Hz.

11. The adhesive of claim 10, wherein a loss modulus G“ is greater than the storage modulus G‘ at 1 Hz.

12. The adhesive of claim 10, wherein a loss modulus G“ is less than the storage modulus G‘ at 100 Hz.

13. The adhesive of claim 1, wherein an air burst pressure is greater than 120 mmHg.

14. The adhesive of claim 1, wherein an water burst pressure is greater than 120 mmHg.

15. The adhesive of claim 1, wherein the surface includes biological tissues.

16. The adhesive of claim 15, wherein the adhesive forms an adhesion to a hydrophilic surface.

17. The adhesive of claim 16, wherein the biological tissues includes a lung, a heart, or other internal tissue.

18. The adhesive of claim 17, wherein the adhesive has a tensile adhesion higher than about 0.5 kPa on the hydrophilic surface.

19. The adhesive of claim 15, wherein the adhesive includes a backing material on one side not configured to adhere to the surface.

20. The adhesive of claim 19, wherein the backing material includes polycarbonate.

21. Method of manufacturing a pressure-sensitive adhesive (PSA), the method comprising: adding a dicarboxylic acid to a vessel; adding a hydrophilic polymer to the vessel, wherein a molar ratio of the dicarboxylic acid to the hydrophilic polymer is about 20: 1 to 2: 1; performing a first polymerization of the dicarboxylic acid and hydrophilic polymer via condensation polymerization to form a linear copolymer; adding glycerol to the vessel, wherein a molar ratio of the di carboxylic acid to the glycerol is about 4: 1 to 1 : 1; and performing a second polymerization of the linear polymer and the glycerol to form a crosslinked adhesive.

22. The method of claim 21, wherein the dicarboxylic acid includes sebacic acid, adipic acid, pimelic acid, suberic acid, phthalic acid, dodecanedioic acid, or derivatives thereof.

23. The method of claim 21, wherein the hydrophilic polymer includes polyethylene glycol (PEG), poly(vinyl alcohol), poly(acrylic acid), polypropylene glycol), poly(2 -hydroxyethyl methacrylate), poly(N-vinylpyrrolidone), poly(methacrylic acid), poly(ethylene oxide), poly(aspartic acid), poly(acrylamide), poly(ethylenimine), dextran, starch-based polymers, cellulose derivates (e.g., hydroxyethyl cellulose, carboxymethyl cellulose), agarose, chitosan, alginate, hyaluronic acid gelatin, or pectin.

24. The method of claim 21, wherein the second polymerization occurs until the crosslinked adhesive reaches a storage modulus G‘ of about 100 Pa at 1 Hz.

25. The method of claim 24, wherein the second polymerization occurs until the crosslinked adhesive reaches a storage modulus G‘of about 1000 Pa to 10000 Pa at 100 Hz.

26. The method of claim 25, wherein a loss modulus G“ is greater than the storage modulus G‘ at 1 Hz.

27. The method of claim 25, wherein a loss modulus G“ is less than the storage modulus G‘ at 100 Hz.

28. The method of claim 21, wherein the second polymerization is performed for a duration between 60 and 100 hours.

29. The method of claim 21, further comprising:placing the crosslinked adhesive on top of a backing material, and topping the crosslinked adhesive with a release liner.

30. The method of claim 29, wherein the backing material includes polycarbonate.

31. The method of claim 29, wherein the release liner includes polytetrafluoroethylene (PTFE).

32. The method of claim 29, further comprising pressing the crosslinked adhesive to form a uniform surface and thickness.

33. A method of atraumatic application on and removal from a biological tissue of the pressuresensitive adhesive of claim 1, the method comprising: applying the PSA to a location on the biological tissue; removing the PSA from the biological tissue; and optionally repositioning the PSA on the location or to a different location on the biological tissue.

34. The method of claim 33, wherein the PSA may be reapplied to the biological tissue up to 75 times.

35. The method of claim 33, wherein the PSA is applied for tissue stabilization during a surgical procedure.

36. The method of claim 33, further comprising placing one or more electrodes on a surface of the biological tissue; and applying the PSA over the one or more electrodes to fix the one or more electrodes to the biological tissue.

37. The method of claim 36, wherein the one or more electrodes include a measurement electrode, a stimulation electrode, or both.

38. The method of claim 33, further comprising placing one or more sensors on a surface of the biological tissue; and applying the PSA over the one or more sensors to fix the one or more sensors to the biological tissue.

39. The method of claim 38, wherein the one or more sensors includes a strain sensor.

40. The method of claim 33, wherein the biological tissue includes a hydrophilic surface.

41. The method of claim 40, wherein the PSA has a tensile adhesion higher than about 0.5 kPa on the hydrophilic surface.

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