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Article

Crosslinker-Free, Printable Alginate–Boronic Acid Hydrogel Adhesive with Enhanced Mechanical Performance for Soft Tissue Fixation

by
Anna Marszałek
1,*,
Zuzanna Kurzępa
1,
Mikołaj Gąbka
1,
Anna Ścisłowska-Czarnecka
2 and
Ewa Stodolak-Zych
1,*
1
Department of Biomaterials and Composites, AGH University of Krakow, 30-059 Kraków, Poland
2
Institute of Applied Sciences, University of Physical Culture in Kraków, 31-571 Kraków, Poland
*
Authors to whom correspondence should be addressed.
Molecules 2026, 31(5), 829; https://doi.org/10.3390/molecules31050829
Submission received: 4 December 2025 / Revised: 18 February 2026 / Accepted: 25 February 2026 / Published: 1 March 2026
(This article belongs to the Special Issue Applications of Natural Polymers in Biomedicine)
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Abstract

Tissue adhesives offer a promising alternative to traditional sutures and staples, particularly in situations requiring rapid, minimally invasive wound closure. To address the limitations of commercially available cyanoacrylate-based adhesives, numerous hydrogel adhesives have been developed. This study presents the synthesis and characterisation of an alginate–aminophenylboronic acid (Alg-APBA) hydrogel adhesive, optimised for bioprinting as a method allowing us to control the thickness of the adhesive layer. The adhesive combines the biocompatibility of alginate with the pH-responsive bonding ability of boronic acid groups, eliminating the need for oxidative crosslinkers. Successful conjugation of APBA to alginate was confirmed via 1H NMR, FTIR and UV-VIS spectroscopy, with a degree of substitution reaching approximately 46% or ~0.22 mol%. Rheological analysis demonstrated shear-thinning and self-healing properties suitable for bioprinting, achieving a high print fidelity (Pr ratio = 0.99 ± 0.08) and repeatability. Mechanical testing showed a shear strength of 19.0 ± 0.5 kPa and an interfacial toughness of 58.0 ± 2.11 J/m2, exceeding those of commercial fibrin adhesives. Additionally, the adhesive joint remained stable after one week of incubation in an acidic environment. The material demonstrated biocompatibility during in vitro testing with keratinocytes and fibroblast cells. These results indicate that Alg-APBA is a strong, biocompatible and printable hydrogel adhesive with potential applications in soft tissue implant fixation.

Graphical Abstract

1. Introduction

Tissue adhesives designed to bond soft tissues are an alternative solution to surgical sutures and staples, and they are especially useful for children and uncooperative adults [1]. Sutureless wound closure strategies improve patients’ comfort by reducing pain and scarring, and they simplify medical procedures, as they require fewer professional skills [2,3]. In addition, the technical requirements of time-consuming stitching or stapling procedures make them unsuitable for emergency situations, such as natural disasters and war [4].
Medical cyanoacrylate (CA) adhesives, such as Histoacryl (B. Braun Melsungen AG, Melsungen, Germany) [5,6], Dermabond (Ethicon Inc., Raritan, NJ, USA) [6,7], Vetbond (3M Company, Saint Paul, MN, USA) [2], and Leukosan (BSN Medical GmbH, Hamburg, Germany) [1,6], are widely used due to their strong tissue adhesion, rapid polymerisation and good cosmetic results. However, the use of such materials raises major concerns. The polymerisation of CAs is exothermic and can cause local tissue toxicity and necrosis [2]. Moreover, during degradation, toxic by-products such as cyanoacetate and formaldehyde are released, restricting internal application [4]. Cyanoacrylates also have unfavourable mechanical properties, namely low elasticity and excessive stiffness of the bond, as well as an inability to adapt to the shape of tissues and accommodate their movement [2,8]. Thus, they are unsuitable for high-tension wounds and are used only externally or temporarily [7,9,10].
The limitations of CAs have driven interest in hydrogel adhesives derived from natural biopolymers, which offer biocompatibility, biodegradability and bioactivity. These include collagen and hyaluronic acid from the mammalian extracellular matrix, fibrin from blood, as well as alginate and agarose from marine algae [11]. However, a natural origin does not ensure safety or performance: fibrin glues may transmit blood-borne viruses such as HIV or hepatitis [12,13], and they often exhibit weak adhesion to wet tissues and poorly controlled biodegradation, limiting their clinical usefulness [4,14].
Hydrogels are polymer–water networks in which water constitutes the majority of the material, similarly to soft tissues, which contain 60–90% water [8,11]. Synthetic hydrogels can closely mimic the chemical, mechanical and electrical properties of tissues and have gained attention for biomedical use due to their responsiveness to stimuli such as the temperature, pH or specific molecules. Hydrogels derived from natural polymers are biocompatible and low in toxicity, making them suitable for applications including drug delivery, tissue scaffolds and wound dressings [8,11,14,15]. They can also be engineered as bioinks for bioprinting. Their shear-thinning and self-healing behaviours ensure both easy extrusion and high shape fidelity [16,17,18].
However, the high water content of hydrogels limits their strong adhesion as they bind poorly to surfaces, including other hydrogels [8,9,10,11,12,13,14,15,16,17,18,19]. Adhesion can be enhanced by grafting polymers with polyphenols such as catechol or pyrogallol, inspired by marine organisms [4,12] and compatible with natural polymers such as alginate, hyaluronic acid and chitosan [3,12,19,20,21]. These functional groups enable strong adhesion through multiple interactions with functional groups present on tissue surfaces [8,12]. Among other compounds, boronic acid exhibits adhesive properties similar to those of catechol in mussel-inspired materials [19,22] and enables the development of pH-responsive hydrogel adhesives with adhesion strengths dependent on the acidity or basicity of the environment [15,19].
This work describes the development of an alginate–aminophenylboronic acid (Alg-APBA) hydrogel adhesive for implant fixation, a concept that has rarely been addressed in the literature. In contrast to cyanoacrylate adhesives, which have been used experimentally for mesh fixation in abdominal hernia repair, but are limited by formaldehyde release and cytotoxicity [23], the proposed Alg-APBA hydrogel offers several advantages. A distinctive feature of the Alg-APBA hydrogel is its suitability for precise, extrusion-based bioprinting. Optimised rheological behaviour allows for a controlled spatial distribution and thickness of the adhesive layer, ensuring reproducibility, uniform coverage and minimal excess material. This controlled application is crucial for achieving uniform mechanical properties and avoiding weak points or reduced joint strength caused by uneven or overly thick adhesive layers. The issue of adhesive application and a uniform layer thickness is largely overlooked in the literature, despite its significant impact on the mechanical strength of the adhesive joints. Proper adhesive distribution is necessary to obtain uniform mechanical properties and to avoid the formation of weak points where stress is concentrated. The amount of adhesive applied is equally important, as an excessively thick layer can weaken the strength of the joint [24,25,26].
The crosslinking mechanism of the proposed material is based on the formation of dynamic covalent bonds, which represent a form of chemical crosslinking. Hydrogel crosslinking mechanisms can be broadly categorised into physical (non-covalent) and chemical (covalent). Physical crosslinking mechanisms are based on reversible interactions, such as hydrogen bonding, as well as ionic, electrostatic, hydrophobic and supramolecular interactions [27,28,29]. Chemically crosslinked gels are generally more stable due to the formation of covalent bonds between polymer chains mediated by crosslinkers. Another crosslinking method is photocrosslinking, in which exposure to light (often ultraviolet) triggers chemical reactions that form covalent bonds between molecules. However, this technique requires the presence of photoreactive groups within the polymer [30,31,32].
An ideal adhesive must provide strong and durable adhesion in physiological conditions, degrade in a controlled manner without producing toxic by-products, and maintain sufficient mechanical flexibility to accommodate tissue movement without constraint. As mentioned earlier, each of the available options (CAs, fibrin adhesives) has certain disadvantages that limit its use. Meanwhile, using APBA to modify alginate allows us to obtain a non-toxic material that is chemically crosslinked by forming dynamic covalent bonds in response to the change in pH, providing strong adhesion. The pH responsiveness of boronic acid groups eliminates the need for oxidative crosslinking agents such as NaIO4, which are commonly used in catechol-based hydrogels and may generate reactive by-products such as H2O2, leading to oxidative stress in surrounding tissues [12,21,33].
We hypothesised that the Alg-APBA adhesive would provide a biocompatible and mechanically stable interface protecting the damaged tissue during its regeneration. To verify this hypothesis, the synthesis of the adhesive was confirmed by 1H NMR and FTIR spectroscopy, while UV-VIS analysis was used to quantify the degree of APBA conjugation. The optimised material was subsequently applied using a bioprinting system, and its adhesive strength and in vitro cytocompatibility were evaluated.

2. Results and Discussion

2.1. 1H NMR Spectroscopy

1H NMR spectroscopy was used to examine the structure of both pure and functionalised alginate and to identify differences indicating the conjugation of APBA to the alginate.
In the 1H NMR spectra of both materials (Figure 1a,b), the chemical shifts of protons from the alginate chain can be distinguished in the region of 3–4 ppm [34]. Chemical shifts in the range of 7–8 ppm correspond to phenyl groups, and they are present only in the Alg-APBA spectra. This observation qualitatively demonstrates the presence of aromatic groups and confirms the successful conjugation of APBA to the alginate chain [35,36]. Additional signals observed in the region of 1–3 ppm in Alg-APBA spectra correspond to the carbodiimide crosslinking by-products [37].

2.2. FTIR Spectroscopy

The second method used to examine the structure of molecules was FTIR spectroscopy, which is widely employed to identify and characterise chemical bonds by analysing their unique vibrational modes. In addition, the spectrum of Alg-APBA after crosslinking was measured to investigate the interactions responsible for the adhesive mechanism. Figure 2 presents the adhesion mechanism schematically, with dynamic covalent bonds marked in green.
The spectra for Alg and Alg-APBA (Figure 3a) show a strong band of approximately 1050 cm−1, corresponding to C-O-C stretching vibrations in the alginate chain. In addition, bands corresponding to C-O stretching and C-H bending vibrations can be distinguished; these are less intense in Alg-APBA than in Alg.
The Alg-APBA spectrum exhibits several bands that are absent in the Alg spectrum: C=O stretching at 1700 cm−1, C-N stretching at 1085 cm−1, and aromatic C-H out-of-plane bending at 700 cm−1 [38]. The presence of these bonds indicates the successful substitution of the alginate chain by APBA.
Furthermore, after calculation of the second derivative (Figure 3b), additional features can be identified in the Alg-APBA spectrum, including C=C aromatic bending (~1500–1600 cm−1 [36]), N-H bending at 1500 cm−1, and aromatic C-H out-of-plane bending of approx. 900 cm−1.
Finally, in the case of crosslinked Alg-APBA, the intensity of bands at approx. 1350 cm−1 and 1200 cm−1 decreased, indicating the formation of covalent bonds in accordance with the crosslinking mechanism (Figure 2).

2.3. UV-VIS Spectroscopy

UV-Vis spectroscopy was used to qualitatively and quantitatively assess the functionalisation of the alginate. Based on the APBA spectrum (Figure 4), a wavelength of 295 nm was selected to prepare the calibration curve for quantitative analysis. Maximum absorption for Alg-APBA was observed at this wavelength. This value is also consistent with that noted in previous studies [19,36].
Using this method, the degree of substitution (DS) was determined to be 46.16%, representing the fraction of alginate monomer units modified with APBA. In the literature, DS values are reported within the range of 10–50%, depending on reagent ratios, which can be tuned during synthesis to influence the hydrogel mechanical strength, swelling, and stimuli responsiveness [19,35,36].

2.4. ICP-OES

ICP-OES analysis was performed to measure the boron content in the modified alginate. The boron concentration was determined to be 0.12 mg B/g polymer, corresponding to a DS of ~0.22 mol% (mol B/mol alginate monomer unit). This method quantifies only the boron atoms in the APBA moiety and serves as confirmatory evidence for successful functionalisation. These values are not directly comparable to the UV-Vis DS, but collectively support the successful modification of the alginate polymer.

2.5. Printability Assessment

Although the term “printability” is widely used in the literature, there is no clear consensus on when a material can be considered “printable” [39]. It can be defined as “the ability of a material, when subjected to a certain set of printing conditions, to be printed in a way which results in printing outcomes which are desirable for a given application” [40]. In the context of extrusion bioprinting, “printability” often refers to extrudability and shape fidelity suitable to form precise, reproducible and structurally stable constructs. There are multiple established qualitative and quantitative approaches for testing printability, such as evaluating the filament, droplet formation, strand uniformity, extrudability, and shape fidelity to the intended design or in silico numerical models [39,41,42].

2.5.1. Rheological Properties

Rheological properties are used as a predictors of bioink printability and shape fidelity. Properties such as shear thinning or the thixotropic behaviour (so-called self-healing) are desirable for bioprinting [39].
An amplitude sweep (Figure 5a) was applied to investigate the viscoelastic behaviour of the material. It additionally allows the determination of the linear viscoelastic region. This test was also used to select a suitable strain value for further tests, which was γ = 1%. The limit of the LVE region, defined as the shear strain at which the value of G’ falls by 5%, was identified at 30%. The figure suggests that the material is fluid, as G” is higher than G’ across the entire range [43].
A frequency sweep (Figure 5b) was performed to further establish whether the material behaves like a viscoelastic liquid or a solid-like gel. The measurements showed the frequency dependence of G’ and G”, with a crossover at 59.2 rad/s, which is a distinctive feature of a viscoelastic liquid.
Viscosity (Figure 5d) has a significant impact on printing fidelity. In general, higher viscosity results in better printing fidelity. The non-linear shape of the rheogram (Figure 5c) indicates that Alg-APBA exhibits shear-thinning behaviour [36,39]. This property is critical for extrusion-based printing, as it allows the material to flow easily through the nozzle under shear stress while maintaining sufficient viscosity at rest to preserve the printed shape.
Shear-thinning behaviour alone is insufficient to determine the printability of a material based solely on rheological properties [41]. Another critical factor is the ability of the material to recover its structure after extrusion—the so-called self-healing capability. This is determined by thixotropy tests, which are designed to model the printing process: extrusion (high strain) and time for recovery (low strain) [39,40]. Oscillatory measurements (Figure 5e) confirmed that Alg-APBA rapidly regained its storage modulus after each high-strain cycle, demonstrating efficient recovery of its internal structure after passing through the needle during the printing process.
In summary, based on the results, it can be concluded that Alg-APBA in PBS (pH 7.4) is a viscoelastic liquid, which has been preliminarily assessed as having printing potential.

2.5.2. Shape Fidelity Assessment

In addition to rheological properties, which are responsible for the printing process, i.e., stable ink flow through the needle, an important element of assessing printability is the shape fidelity of the bioink—the ability to retain the printed shape, reflecting the designed model. The results of the rheological study presented in the previous subsection allowed us to state that the prepared bioink behaves like a viscoelastic liquid. Generally, gel-like materials allow for better shape fidelity and filament stability in bioprinting [40,43]. However, since, in this application, the printed shape is simple, planar and consists of only one layer, there was no need to investigate the printing fidelity or stability of multilayer 3D shapes. Therefore, shape fidelity was evaluated using a dimensionless printability index (Pr), which relies on the evaluation of pore geometry. For ideal printability, the interconnected filaments would demonstrate a square shape and a Pr value of 1 [44].
In order to calculate the Pr index, a grid model composed of 2 × 2 mm squares was printed (Figure 6a). The area and perimeter of four selected squares were measured and used to calculate the printability ratio (Pr, Equation (1)). The obtained value of Pr = 0.99 ± 0.08 indicates excellent shape fidelity, confirming that the adhesive maintained its geometry after deposition and that its gelation behaviour was suitable for precise bioprinting and continuous flow from the needle [44].
A 3D model designed for printing is presented in Figure 6b, and the printed sample is shown in Figure 6c. A uniform and continuous layer of the adhesive was successfully applied to the fibrous PCL substrate, demonstrating an effective coverage and adhesion to the material surface. These results confirm that the optimised rheological profile and self-healing ability of Alg-APBA ensure precise and reproducible deposition of the adhesive, while controlling the amount of applied material. This makes the Alg-APBA hydrogel suitable for extrusion-based bioprinting applications, where accurate spatial control of adhesive distribution is required.
Compared to other bioinks based on natural polymers such as gelatine methacrylate (GelMA), calcium-crosslinked alginate or catechol-modified chitosan [45,46], the Alg-APBA adhesive exhibits superior control over deposition and shape fidelity during printing. Its rheological profile, combining pronounced shear thinning with rapid post-extrusion recovery, enables accurate and repeatable application of the adhesive in predefined patterns, while preventing uncontrolled spreading typical for low-viscosity hydrogels. This controllability is particularly advantageous in biomedical applications requiring precise localisation of an adhesive, such as the fixation of implants, where the amount and spatial distribution of the material strongly influence bonding efficiency and biocompatibility. Thus, the optimised printability of Alg-APBA not only facilitates fabrication but also opens up a path towards reproducible, minimally invasive tissue fixation strategies.

2.6. Swelling Test

A swelling test was performed to evaluate the hydrated stability of Alg-APBA. In Figure 7, the initial increase in the weight of material is shown, suggesting initial swelling. However, a sharp loss of weight occurred later on, and the dissolving of the hydrogel was macroscopically observed. This behaviour can be explained by two factors. Firstly, the alginate itself, which forms the basis of the material, is soluble in water. Secondly, placing Alg-APBA in distilled water, which has a lower pH than PBS, causes partial reversal of the pH-dependent crosslinking. The DS of the material is also significant, as it is directly related to the number of bonds formed during crosslinking [19].

2.7. Mechanical Properties

The performance of a tissue adhesive is commonly assessed by two parameters: adhesion strength in tensile or shear tests, and interfacial toughness [47]. In this study, the shear strength of Alg-APBA was evaluated using the lap-shear test, which is a widely used method to quantify adhesive strength at soft tissue and biomaterial interfaces (Figure 8a,c). Meanwhile, the T-peel configuration (Figure 8b,d) highlights the interfacial toughness and failure behaviour, which are particularly relevant for thin tissue applications such as gastric wall fixation.

2.7.1. Shear Strength

The shear strength of Alg-APBA, measured via the lap-shear test, was 19.0 ± 0.5 kPa (Equation (2)). The measurement was performed on a nonwoven–adhesive–nonwoven system prepared, as described in the Section 3. This value falls within the range of values reported for boronate-based hydrogels (10–50 kPa [19,36]), and it exceeds that of fibrin adhesives (~5 kPa [12,13]). This result suggests that Alg-APBA is a promising material for use in securing implants to soft tissues subject to dynamic movements causing shear stress, such as the gastric wall.
Although its shear performance is below the strength typically observed for cyanoacrylate-based systems (~200–300 kPa [48,49,50,51]), CAs generally show superior mechanical strength, comparable to or even exceeding that of traditional sutures. The main reason for this difference lies in the bonding mechanism. For CAs, adhesion is primarily achieved through a covalent bond with the substrate. In hydrogels, adhesion is primarily governed by the dominant bonding mechanism—covalent, hydrogen or dynamic reversible interactions—as well as the network porosity and hydration level. Additionally, the majority of hydrogels are porous, providing a high surface area for interactions with the tissue or biomaterial [12]. Consequently, CAs generally demonstrate a more consistent mechanical performance than fibrin or hydrogel adhesives [52,53,54,55].
The sample retained its integrity after incubation in an acidic environment. A decrease in the shear strength was observed (15.5 ± 0.3 kPa). This value also falls within the range of values reported for boronate-based hydrogels, which demonstrates the suitability of the material for application in an acidic environment [19,36].

2.7.2. Interfacial Toughness

The T-peel test simulates the formation of a weak point in the adhesive joint by applying force to the outer edge of the adhesive bond [54]. Interfacial toughness, measured during the T-peel test, was 58.0 ± 2.1 J/m2 (Equation (3)), representing a significant improvement compared to commercial fibrin adhesives (<40 J/m2 [56]). For cyanoacrylates, interfacial toughness can exceed 2000 J/m2, although this parameter deteriorates in a wet environment [57]. As in the previous test, exposure to an acidic environment did not affect the continuity of the adhesive joint; however, it was weakened (24.62 ± 1.9 J/m2). By combining multiple adhesion mechanisms, hydrogel tissue adhesives can achieve high interfacial toughness values. The interfacial toughness of modern hydrogel adhesives falls within a wide range, from 450 J/m2 [58] to over 1500 J/m2 [59,60] on wet tissue—depending on the adhesive formulation and testing conditions.
It is important to consider the procedural limitations of these tests. Mechanical measurements and their comparability depend heavily on testing conditions, particularly the type of substrate used for testing and the thickness of the adhesive layer. In many studies, different substrates have been used for adhesion testing, such as various tissues, i.e., porcine skin tissue, colonic segments or aorta [14,54,61], thin polymeric films [62] or glass [15]. Animal tissue is generally considered the most suitable substrate for testing tissue adhesives, but tissue samples from individual animals exhibit variability in quality and properties, such as stiffness [61].
Another important factor is the thickness of the adhesive layer. Unfortunately, in few studies has the effect of the applied material quantity on joint mechanical properties been addressed, as standardised methods for testing tissue adhesives are not yet well-established [63]. In general, the optimal thickness for maximum adhesion strength depends on the adhesive type, substrate and application. Both excessively thin and excessively thick layers can compromise joint performance [24,26,63,64]. Below the optimal thickness, strength is often independent of layer thickness, whereas above it, excessive thickness can reduce strength due to factors such as crack propagation and uneven stress distribution [24,26,63]. Extremely thin (<0.05 mm) layers have also been shown to reduce strength, emphasising the need for precise application techniques to control thickness [64,65]. In this study, bioprinting allowed precise control of the layer thickness. A constant layer thickness of 0.1 mm was used, as the optimal thickness for most adhesives lies between 0.05 and 0.15 mm [26].
A controlled adhesive distribution is also critical, as it reduces the stress concentrations occurring in uneven layers, particularly at the corners of the joint [47]. In conventional hydrogel adhesive studies, a specific volume of precursor is applied to one part of the joint, and a crosslinking reagent to the other [3,33]. However, this does not guarantee a uniform distribution of the adhesive, unlike the bioprinting method implemented in this study.
For the reasons described above, it is difficult to directly compare the results of different studies. Overall, the mechanical properties of hydrogel tissue adhesives are highly tunable and depend on the specific chemistry and crosslinking strategies employed, with recent advances achieving both high strength and biocompatibility for a range of surgical and regenerative applications [8,12,33,54,55]. Hydrogels can show high shear strength and interfacial toughness, which are tunable by the type of polymer, molecular weight, pH and temperature, as their adhesion is generated through various mechanisms (e.g., covalent and non-covalent bonds) [8,12,66].

2.8. Biological Evaluation

Significant cytotoxicity concerns for alginate–APBA adhesives at concentrations used in biomedical research are not reported in the literature. Additionally, the dynamic boronic ester bonds formed between alginate and APBA generate a biocompatible, adaptable matrix. This environment does not release toxic by-products in physiological conditions [67,68].
In this study, two cell lines were used to assess the biological response to the material. For keratinocytes, the number of viable cells after seven days significantly increased for both materials compared to measurements taken on day three (Figure 9a). This was associated with an increase in cytotoxicity, i.e., the number of dead cells (Figure 9b). Combined with increased viability in both cases, this allows us to conclude that the increased number of dead cells is not due to the toxicity of the material, but rather results from the death of cells in the lower layers of the culture due to a lack of access to nutrients. While the increase in viability for Alg-APBA was lower compared to TCPS, and consequently its cytotoxicity as well, it is suggested that the material does not promote keratinocyte growth and is non-toxic.
For fibroblasts, the same trend of increased viability over time was observed for TCPS (Figure 10a), with cytotoxicity showing no statistically significant differences (Figure 10b). A decrease in cell viability was noted for Alg-APBA. However, cytotoxicity testing showed a reduction in the number of dead cells as well, suggesting that Alg-APBA is not toxic but rather inhibits fibroblast growth. These results indicate that Alg-APBA does not exhibit cytotoxicity towards keratinocytes and fibroblasts.
Tissue adhesives come into direct contact with damaged and regenerating tissues. For this reason, it is important to ensure that such materials are not cytotoxic, which is a significant concern in the case of cyanoacrylate adhesives. CAs primarily degrade through hydrolysis, especially when implanted below the epidermis or exposed to moisture. The main products of this degradation are formaldehyde and cyanoacetate, which are associated with cytotoxicity and inflammatory responses. [23,69,70,71]. Long-chain CAs degrade more slowly, resulting in lower concentrations of toxic by-products and improved biocompatibility. In vivo, complete absorption of short-chain cyanoacrylates can occur within one to two weeks, whereas longer-chain variants may persist for several weeks or even months [23,72].
In contrast, degradation of fibrin adhesives does not produce toxic or immunogenic substances in normal physiological conditions. Their decomposition is a controlled enzymatic process, mimicking the breakdown of natural blood clots by fibrinolytic enzymes as part of normal tissue remodelling and wound-healing processes. The main products of degradation are amino acids, small peptides and fibrin degradation products (FDPs). Fibrin glue is typically resorbed within 2 to 12 weeks, depending on the application site, local enzyme activity and presence of cells. This process is synchronised with tissue healing, ensuring that the scaffold is gradually replaced by new tissue [73,74,75,76,77].

3. Materials and Methods

3.1. Preparation of Alg-APBA

The concept of alginate chain modification is illustrated in Figure 11. Figure S1 presents detailed mechanism of the reaction. At this stage, 2 g of sodium alginate Protanal LF 10/60 (Alg, FMC BioPolymer, Philadelphia, PA, USA) was dissolved in 0.1M 2-(N-morpholino) ethanesulfonic acid buffer (MES, Sigma-Aldrich, Burlington, MA, USA) under constant stirring, with the pH adjusted to 5.5 to provide optimal conditions for the EDC/NHS amidation reaction. Then, 700 mg of N-(3-Dimethylaminopropyl)-N’-ethylcarbodiimide hydrochloride (EDC·HCl, Pol-Aura, Morąg, Poland) and 100 mg of N-hydroxysuccinimide (NHS, Sigma-Aldrich, Burlington, MA, USA) were added to the solution, and after 24 h, the solution was dialysed (Spectra/Por MWCO 3.5 kDa, Spectrum Laboratories Inc., Compton, CA, USA) and lyophilised.
In the second step, the obtained lyophilisate was dissolved in PBS (pH 7.4) at a concentration of 2%, and 300 mg of 3-aminophenylboronic acid hydrochloride (APBA, Sigma-Aldrich, Burlington, MA, USA) was added. After 24 h, the solution was purified and lyophilised following the same procedure as in the first step. The resulting lyophilised adhesive (Alg-APBA) was stored at room temperature.

3.2. Proton Nuclear Magnetic Resonance Spectroscopy (1H NMR)

1H NMR was used to examine the structure of Alg and Alg-APBA molecules, to determine the successful functionalisation of the polymer.
During this step, 4 mg of Alg and Alg-APBA was dissolved in 0.7 mL of D2O and analysed using liquid-state 1H NMR (FT-NMR 500 MHz spectrometer, JEOL Ltd., Tokyo, Japan).

3.3. Fourier Transform Infrared Spectroscopy (FTIR)

FTIR-ATR spectra of Alg and Alg-APBA lyophilisates were analysed in the range from 400 cm−1 to 2000 cm−1 in the solid state (Tensor 27 FT-IR Spectroscope, Bruker, Billerica, MA, USA). Additionally, the spectrum of crosslinked Alg-APBA was measured to investigate interactions between functional groups.

3.4. Ultraviolet-Visible Spectroscopy

UV-VIS spectroscopy (UV-2600i spectrophotometer, Shimadzu, Kioto, Japan) was used to qualitatively and quantitatively assess functionalisation of the alginate.
To quantify APBA directly, the spectrum of a 1 mg/mL aqueous solution of APBA was recorded within the range of 250–450 nm to identify the maximum absorbance peak. After determining the appropriate wavelength (295 nm), the absorbance of APBA solutions with increasing concentrations from 0 μg/mL to 500 μg/mL was measured to prepare a calibration curve. The absorbance of a 500 μg/mL Alg-APBA solution was then measured, and the degree of APBA substitution (DS) in the alginate was calculated based on the calibration curve.

3.5. Inductively Coupled Plasma Optical Emission Spectroscopy

The boron content in Alg-APBA hydrogel adhesive was determined using inductively coupled plasma optical emission spectroscopy (ICP-OES). Samples (~50 mg) were digested with 5 mL of concentrated nitric acid (HNO3, POCH, Gliwice, Poland) and 1 mL of hydrogen peroxide (H2O2, Chempur, Piekary Śląskie, Poland) in a closed microwave digestion system (CEM MARS6, CEM, Matthews, NC, USA) at 180 °C for 20 min to ensure complete decomposition of the organic matrix. After digestion, the solutions were diluted to a final volume of 50 mL with deionised water and filtered through a 0.45 µm membrane to remove any insoluble residues.
Boron quantification was performed using the ICP-OES spectrometer (Agilent 5110, Agilent Technologies, Santa Clara, CA, USA), at a wavelength of 249.678 nm. Calibration was carried out with standard boron solutions (boric acid, Sigma-Aldrich, Burlington, MA, USA) in the 0.01–5 mg/L range. All measurements were performed in triplicate, and the boron content was expressed as mg B per gram of polymer, which was then converted to the degree of substitution (DS, mol boron/mol alginate monomer unit).

3.6. Printability Assessment

3.6.1. Rheological Properties

All rheological measurements were carried out at 20 °C for a 4.5% Alg-APBA solution in PBS using a rheometer (Modular Compact Rheometer MCR 302, Anton Paar GmbH, Graz, Austria). A parallel plate geometry was applied, with a plate diameter of 20 mm and a gap width of 1 mm.
An amplitude sweep for shear strain γ in the 10−1–103% range was performed to determine the linear viscoelastic (LVE) region. The limit of the LVE region was established as the shear strength at which the storage modulus drops by 5% from the plateau value, in accordance with the ISO 6721-10 standard [78]. After determining the LVE region, a frequency sweep was conducted for angular frequency ω in the range of 1–100 rad/s and at a constant shear strain of 1%. The crosslinking time of the adhesive was investigated by measuring the storage modulus G’ and loss modulus G” at a constant shear strain of 1%. The crosslinking time was defined as the point where G’ and G” were equal. Shear viscosity was measured over a shear rate range of 10−3–103 1/s. The self-healing property of the adhesive was assessed by an oscillatory amplitude sweep test, in which there was a low shear strain of 1% for 180 s, followed by a high shear strain of 100% for 90 s, followed with a low shear strain of 1% for 180 s. This cycle was repeated four times.

3.6.2. Shape Fidelity Assessment

Printing (BioCloner Desktop Pro printer, BioCloner Health, Warsaw, Poland) was performed using a 22G needle (diameter 0.70 mm) at a constant extrusion rate of 4 mm/s and a temperature of 20 °C. The layer height was 0.1 mm. Printability was characterised using a semi-quantitative method by calculating a dimensionless ratio Pr [44], according to the following equation:
Pr = π/4 · 1/C = L2/(16A)
where L is the perimeter of the printed square, and A is its area. The amount of applied adhesive was calculated based on the prepared model. Two layers were printed. Images captured with a digital optical microscope (VHX-900F, Keyence International, Mechelen, Belgium) were analysed using ImageJ software (v. 2015, National Institute of Mental Health, Bethesda, MD, USA) to measure L and A values in Equation (1).

3.7. Swelling Test

Hydrated stability was investigated in the swelling test. After printing, the sample of Alg-APBA was weighed and placed in distilled water. The sample was weighed every five minutes until dissolved.

3.8. Mechanical Properties

Mechanical strength was investigated in two categories: the shear strength of the adhesive joint and the interfacial toughness between Alg-APBA and a model material, which was electrospun polycaprolactone nonwoven material. The adhesive was applied to the nonwoven strip by bioprinting using parameters described in the previous subsection, and covered with a second strip, which was then left for 12 h. This created a lap with a width of b = 10 mm and a length of l = 20 mm. Additionally, mechanical tests were performed on samples prepared in the same manner after incubation in an HCl solution (pH 2) for seven days at 37 °C. This was done to investigate the stability of adhesion in an acidic environment.
Lap-shear (based on ASTM F2255-05 standard [79]) and T-peel tests (based on ASTM F2256-05 standard [80]) were carried out using a universal testing machine (RetroLine 1435, ZwickRoell, Ulm, Germany) at a crosshead speed of 20 mm/min. Both tests were performed in triplicate at room temperature.
The shear strength of the adhesive joint was calculated using the following equation:
τ = Fmax/P [Pa]
where Fmax is the maximum registered force, and P is the adhesion area. P calculated on the basis of the model designed for printing (Figure 6b) was 103 mm2.
The interfacial toughness was calculated using the following equation:
Γ = 2F/b [J/m2]
where F is the mean registered force. According to the prepared model, the width of the adhesive joint b is not constant (Figure 6b); therefore, the mean value of 65 mm was used for calculations.

3.9. Biological Evaluation

Cell–material interactions were studied to assess the biocompatibility of the adhesive. The tests were conducted using an extract of the adhesive. Crosslinked Alg-APBA, in the amount of 1 mL, was allowed to dry completely. It was sterilised with UV radiation and then incubated for seven days in 10 mL of distilled water previously sterilised at 37 °C.
Two cell lines were used for the biological tests: fibroblasts (BJ CRL-2522, ATCC, Manassas, VA, USA) and HaCaT keratinocytes (HaCaT, ATCC, Manassas, VA, USA). Next, 1 mL of the prepared extract was added to a tissue culture polystyrene (TCPS) 24-well plate containing 104 cells, incubated at 37 °C. Four replicates were performed for each type of sample. Cell viability and cytotoxicity of the samples were evaluated after three and seven days of culturing, using the ViaLight BioAssay Kit (Lonza Bioscience, Walkersville, MD, USA) and ToxiLight BioAssay Kit (Lonza Bioscience, Walkersville, MD, USA), respectively, according to the manufacturer’s instructions. Briefly, the intensity of the luminescence in the chosen tests corresponds to the number of living (viability assay) or dead cells (cytotoxicity assay). Luminescence was measured using the FLUOstar Omega (Omega software v. 5.10, BMG Labtech, Ortenberg, Germany) reader.

3.10. Statistical Analysis

Where applicable, the results are presented as mean values ± standard deviations. One-way analysis of variance—followed by the Tukey–Kramer post hoc test—was used to identify statistically significant differences, with a p-value of less than 0.05.

4. Conclusions

The successful development of a boronate-functionalised alginate hydrogel adhesive suitable for biomedical applications was demonstrated in this study. The material exhibited the desired chemical functionality, favourable rheological behaviour, enabling extrusion-based bioprinting, and formation of uniform, reproducible adhesive layers on the implant surface.
Mechanical testing confirmed that Alg-APBA provides robust adhesion in wet physiological conditions, achieving a shear strength and interfacial toughness comparable to modern hydrogel adhesives and superior to fibrin-based systems. The material also showed good cytocompatibility towards keratinocytes and fibroblasts, with lower cytotoxicity than the reference sample. Together, these results indicate that Alg-APBA is a promising candidate for use as a printable, biocompatible tissue adhesive for implant fixation and other applications requiring flexible, conformable bonding. Overall, the findings of this study suggest that the Alg-APBA hydrogel adhesive represents a promising alternative to conventional implant fixation methods. It combines the biocompatibility and tunability of natural polymers with the strong and reversible bonding capabilities of boronic acid chemistry. Its demonstrated printability further enables integration with modern additive manufacturing techniques, providing even application of the hydrogel, which is necessary to ensure a uniform strength of the adhesive joint.

Supplementary Materials

The following supporting information can be downloaded at https://www.mdpi.com/article/10.3390/molecules31050829/s1, Figure S1: Detailed illustration of alginate functionalisation with APBA.

Author Contributions

Conceptualisation, A.M. and E.S.-Z.; methodology, A.M., M.G. and E.S.-Z.; validation, A.M.; formal analysis, A.M. and Z.K.; investigation, A.M., Z.K., A.Ś.-C. and M.G.; resources, A.M. and E.S.-Z.; writing—original draft preparation, A.M.; writing—review and editing, E.S.-Z.; visualisation, A.M.; supervision, E.S.-Z.; funding acquisition, A.M. and E.S.-Z. All authors have read and agreed to the published version of the manuscript.

Funding

This study was supported by a subvention of the Polish State Ministry of Education and Science for the AGH University of Krakow, Faculty of Materials Science and Ceramics (project no. 16.16.160.557). The research project was also supported by the IDUB “Excellence initiative—research University” programme for the AGH University of Krakow, project ID 12451 and ID 9735.

Institutional Review Board Statement

Not applicable.

Informed Consent Statement

Not applicable.

Data Availability Statement

The data presented in this study are available upon request from the corresponding author.

Conflicts of Interest

The authors declare no conflicts of interest.

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  78. ISO 6721-10; Plastics—Determination of Dynamic Mechanical Properties. ISO: Geneva, Switzerland, 2015.
  79. ASTM F2255-05; Standard Test Method for Strength Properties of Tissue Adhesives in Lap-Shear by Tension Loading. ASTM International: West Conshohocken, PA, USA, 2015.
  80. ASTM F2256-05; Standard Test Method for Strength Properties of Tissue Adhesives in T-Peel by Tension Loading. ASTM International: West Conshohocken, PA, USA, 2015.
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Figure 1. 1H NMR spectra of (a) Alg; (b) Alg-APBA.
Figure 1. 1H NMR spectra of (a) Alg; (b) Alg-APBA.
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Figure 2. Adhesion mechanism of Alg-APBA.
Figure 2. Adhesion mechanism of Alg-APBA.
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Figure 3. Results of FTIR measurements: (a) FTIR spectra of Alg, Alg-APBA lyophilisate and crosslinked Alg-APBA; (b) second derivatives of FTIR spectra.
Figure 3. Results of FTIR measurements: (a) FTIR spectra of Alg, Alg-APBA lyophilisate and crosslinked Alg-APBA; (b) second derivatives of FTIR spectra.
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Figure 4. UV-VIS spectra of Alg and Alg-APBA.
Figure 4. UV-VIS spectra of Alg and Alg-APBA.
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Figure 5. Rheological measurements of 4.5% Alg-APBA in PBS: (a) amplitude sweep; (b) frequency sweep; (c) rheogram; (d) viscosity curve; (e) self-healing of adhesive; (f) crosslinking time.
Figure 5. Rheological measurements of 4.5% Alg-APBA in PBS: (a) amplitude sweep; (b) frequency sweep; (c) rheogram; (d) viscosity curve; (e) self-healing of adhesive; (f) crosslinking time.
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Figure 6. Printability assessment of Alg-APBA: (a) grid of squares printed in order to determine shape fidelity; (b) model designed for printing; (c) model printed on PCL substrate.
Figure 6. Printability assessment of Alg-APBA: (a) grid of squares printed in order to determine shape fidelity; (b) model designed for printing; (c) model printed on PCL substrate.
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Figure 7. Weight of hydrogel over time after placement in distilled water.
Figure 7. Weight of hydrogel over time after placement in distilled water.
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Figure 8. Results of mechanical tests: (a) shear-lap test configuration; (b) T-peel test configuration; (c) exemplary force–elongation curve for lap-shear test; (d) exemplary force–elongation curve for t-test.
Figure 8. Results of mechanical tests: (a) shear-lap test configuration; (b) T-peel test configuration; (c) exemplary force–elongation curve for lap-shear test; (d) exemplary force–elongation curve for t-test.
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Figure 9. Biocompatibility tests: viability (a) and cytotoxicity (b) of Alg-APBA in contact with keratinocytes. (**) indicates p value less than 0.01; (***) indicates p value less than 0.001.
Figure 9. Biocompatibility tests: viability (a) and cytotoxicity (b) of Alg-APBA in contact with keratinocytes. (**) indicates p value less than 0.01; (***) indicates p value less than 0.001.
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Figure 10. Biocompatibility tests: viability (a) and cytotoxicity (b) of Alg-APBA in contact with fibroblasts. (**) indicates p value less than 0.01; (***) indicates p value less than 0.001.
Figure 10. Biocompatibility tests: viability (a) and cytotoxicity (b) of Alg-APBA in contact with fibroblasts. (**) indicates p value less than 0.01; (***) indicates p value less than 0.001.
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Figure 11. Functionalisation of alginate with APBA.
Figure 11. Functionalisation of alginate with APBA.
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MDPI and ACS Style

Marszałek, A.; Kurzępa, Z.; Gąbka, M.; Ścisłowska-Czarnecka, A.; Stodolak-Zych, E. Crosslinker-Free, Printable Alginate–Boronic Acid Hydrogel Adhesive with Enhanced Mechanical Performance for Soft Tissue Fixation. Molecules 2026, 31, 829. https://doi.org/10.3390/molecules31050829

AMA Style

Marszałek A, Kurzępa Z, Gąbka M, Ścisłowska-Czarnecka A, Stodolak-Zych E. Crosslinker-Free, Printable Alginate–Boronic Acid Hydrogel Adhesive with Enhanced Mechanical Performance for Soft Tissue Fixation. Molecules. 2026; 31(5):829. https://doi.org/10.3390/molecules31050829

Chicago/Turabian Style

Marszałek, Anna, Zuzanna Kurzępa, Mikołaj Gąbka, Anna Ścisłowska-Czarnecka, and Ewa Stodolak-Zych. 2026. "Crosslinker-Free, Printable Alginate–Boronic Acid Hydrogel Adhesive with Enhanced Mechanical Performance for Soft Tissue Fixation" Molecules 31, no. 5: 829. https://doi.org/10.3390/molecules31050829

APA Style

Marszałek, A., Kurzępa, Z., Gąbka, M., Ścisłowska-Czarnecka, A., & Stodolak-Zych, E. (2026). Crosslinker-Free, Printable Alginate–Boronic Acid Hydrogel Adhesive with Enhanced Mechanical Performance for Soft Tissue Fixation. Molecules, 31(5), 829. https://doi.org/10.3390/molecules31050829

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