Impacts of cross-linker chain length on the physical properties of polyampholyte hydrogels

In the biomedical field, there is a need for engineered scaffold materials that can improve and facilitate wound healing in severely damaged tissues. These scaffolds must have several important properties to ensure successful tissue recovery. First, the biomaterial must be nonfouling or resistant to nonspecific protein adsorption. It is hypothesized that this feature will minimize any foreign body response, preventing connective tissue from becoming inflamed and scarred.^1,2 Inflammation and scarring lead to the encapsulation of the implanted biomaterial, preventing the biomaterial from fully integrating and maintaining functionality.^3–7 Second, the biomaterial must also facilitate targeted cell adhesion, migration, and remodeling to promote tissue regeneration.^7,8 To do this, the material must present biochemical cues specific to the target tissue. Third, the scaffold material should be biodegradable and not break down into components that are toxic or harmful to the surrounding cells and tissues.^9,10 Ideally, the material will break down at the same rate at which tissue is reforming to prevent defects in the developing tissue.^1,11 Finally, an implant must have mechanical properties similar to the tissue it will be integrated with, to allow for faster return to functionality.^9,10,12 Mechanical properties also impact the differentiation of seeded or recruited stem cells which is important for the regeneration of the intended tissue.^13,14

There have been many approaches to develop engineered biomaterials to fulfill these criteria, but it has been difficult to design something that incorporates all of these properties. Hydrogels are the standard for tissue engineering because of their high water content and general biocompatibility.^1,15,16 They also have tunable properties and the ability to be synthesized on a larger scale.^1,8 Some common chemistries that have been explored for tissue engineering include poly(lactic acid) (PLA), poly(hydroxyethyl methacrylate) (PHEMA), and poly(caprolactone) (PCL).^1,17 PLA is biologically inert and has high degradability.^16,18 However, this material still invokes an inflammatory response as it is not nonfouling.^18,19 Additionally, its degradation rate is relatively slow.¹⁹ PHEMA has been used in tissue engineering for decades because it is biocompatible and has some ability to reduce nonspecific protein adsorption.¹⁷ Since this material is only low fouling, it still causes inflammation when implanted. While PCL is biodegradable and inexpensive, it is hydrophobic and has poor cytocompatibility.^20,21

Another common material is polyethylene glycol (PEG), which is resistant to nonspecific protein adsorption, leading to its widespread use in biomedical systems. However, PEG based materials are known to oxidize in vivo.²² This ultimately leads to protein adsorption and cell attachment as it breaks down into acids and aldehydes.²² It has also been reported that humans have developed antibodies to PEG, suggesting that there is a need to find alternative biomaterials.^23,24

Zwitterionic materials have recently become of interest for biomedical applications because they have shown great antifouling properties when compared to other nonfouling materials such as PEG.^5,25,26 These neutral systems are composed of mixed charged functional groups throughout the polymer. As a result, they are extremely hydrophilic due to the large number of charged regions.²³ The most widely studied zwitterionic materials are polycarboxybetaine (pCB), polysulfobetaine, and polyphosphorylcholine.^23,25 pCB-based materials, in particular, have gained a lot of attention because of their superior antifouling properties, biocompatibility, and functionalization capabilities.^6,7,23,25,27 Functionalization allows for the covalent attachment of biomolecules, and pCB is highly functionalizable due to its many carboxyl groups.^27,28 However, one limitation of zwitterionic polymers is that their physical properties are dependent on the underlying monomer, making their physical properties less tunable than other materials.^5,29

One family of materials that presents many of the desired properties for tissue engineering are polyampholyte hydrogels, which are a subclass of zwitterionic polymers. They are being considered for tissue engineering and drug delivery applications due to their unique characteristics including nonfouling properties, ability to deliver biomolecules to the body, and tunable mechanical properties.^30–32 Polyampholyte hydrogels are composed of equimolar mixtures of positively and negatively charged monomer subunits, resulting in an overall neutral charge.³³ Their nonfouling properties are likely derived from the formation of a strong hydration layer on the surface of these hydrogels due to electrostatic interactions with the charged terminal groups of the monomers.^31,32,34,35 Polyampholyte hydrogels also have tunable mechanical properties depending on both the underlying monomer composition and the cross-linker density.^31,32 Finally, salt concentration can have an effect on the physical properties due to the salt ions interacting with the charged groups and disrupting the electrostatic interactions between the monomer subunits.^30,33,36

While several beneficial features of polyampholytes have been demonstrated, their degradation behaviors are relatively unknown. Understanding the degradation behavior will provide insight into the in vivo behavior of these materials, allowing for targeted applications in tissue engineering. It was hypothesized that different ethylene glycol chain lengths will lead to variation in the physical and mechanical properties, without impacting the multifunctional nonfouling capability. Therefore, the goal of this investigation is to characterize the degradation and physical properties of polyampholytes with different ethylene glycol cross-linkers. The polyampholytes used in this study contain equimolar amounts of cationic [2-(acryloyloxy)ethyl] trimethylammonium chloride (TMA) and anionic 2-carboxyethyl acrylate (CAA) monomers, which are shown in Fig. 1. Hydrogels were synthesized with one of three chemical cross-linkers, di-, tri-, or tetra-ethylene glycol dimethacrylate (DEGDMA, TEGDMA, and Tetra-EGDMA), also shown in Fig. 1, to study how the cross-linker chain length impacts the properties of polyampholyte hydrogels. The multifunctionality was validated using fluorescently labeled proteins and microscopy, the compressive and tensile properties were quantified, and the degradation behavior as a function of pH was determined.

Ethylene glycol, phosphate-buffered saline (PBS, pH 7.4), fluorescein isothiocyanate bovine serum albumin (FITC-BSA), TMA, CAA, TEGDMA, DEGDMA, Tetra-EGDMA, sodium hydroxide, ammonium persulfate (APS), sodium metabisulfate (SMS), N-(3-dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), and N-hydroxysuccinimide (NHS) were purchased from Sigma-Aldrich. Potassium phosphate monobasic and sodium phosphate dibasic, anhydrous were purchased from VWR. Ethanol was purchased from Greenfield Global. All chemicals were used as received.

Hydrogels were synthesized using equimolar concentrations of TMA and CAA using procedures described previously.³¹ In short, 2 ml of a buffer solution containing 3 M NaOH, ethanol, and ethylene glycol in a 1.5:1:1.5 volume ratio was mixed with 4 mmol of TMA and CAA each and mixed thoroughly. One of the three cross-linkers was added in with monomer to cross-linker ratios of 26.3:1 (1X), 13.2:1 (2X), or 6.6:1 (4X). The reaction was initiated by adding 32 μl of 40% w/w APS and 32 μl of 15% w/w SMS and the solution was then transferred into a polyfluoroethylene (PTFE) mold. The hydrogels polymerized at room temperature for 24 h. With the exception of degradation, all hydrogels were then soaked for 24 h in PBS and used in subsequent studies.

For nonfouling and conjugation testing, hydrogels were synthesized in a flat plate mold formed from two microscope slides clamped around a 1/8 in. PTFE spacer. Following synthesis, the hydrogels were soaked in PBS for 24 h and then they were punched with a 5 mm biopsy punch.³¹ Conjugation and control samples were first soaked in a solution of 0.05 M NHS and 0.2 M EDC for 7 min, while the nonfouling samples were soaked in neutral PBS. The nonfouling and conjugation samples were then exposed to 10 μl of 1 mg/ml FITC-BSA in PBS for 15 min, while the control samples were soaked in neutral PBS. Finally, all samples were soaked in 1 ml of NaCl-PBS with a pH of 9.0 for 30 min, followed by 1 ml neutral PBS for 40 min. Nonfouling and conjugation were evaluated by comparing the fluorescence of these samples to the control sample using an inverted fluorescent Nikon microscope with a 4X objective. Images were captured using NIS Elements imaging software.

For compression testing, hydrogels were synthesized in a cylindrical PTFE mold with a 5/8 in. inner diameter. Compression samples were made by tripling the above procedures. After the samples soaked for 24 h in PBS, they were cut using a razor and mold to obtain 10 mm tall hydrogel cylinders. Compression tests were conducted using an Instron equipped with a 100 kN load cell. Samples were compressed to failure at a rate of 0.106 mm/s. The maximum stress and strain were obtained from the force displacement curve and Young's modulus was taken from the linear region (75%–85% of the maximum strain) of the stress–strain curve.

For tensile testing, hydrogels were synthesized in a PTFE dog-bone shaped mold. Tensile testing was conducted using a custom tensile load frame equipped with a 150 g load cell.³⁷ Samples were stretched at a rate of 0.1 mm/s until fractured. The maximum stress and strain were obtained from the force displacement curve and Young's modulus was taken from both the linear region (50%–100% of the maximum strain) and the toe region (0–0.1 strain values) of the stress–strain curve.

For degradation testing, hydrogels were synthesized in the flat plate mold described above. After polymerizing for 24 h, samples were immediately cut into 1 cm squares and placed in individual 50 ml centrifuge tubes containing Sorenson's buffer. Sorenson's buffer with pH values 4.5, 7.4, and 9.0 was used to induce degradation under different conditions. Degradations studies were conducted over a period of 113 days with 1X cross-linker density hydrogels. At each measurement time point, the samples were removed from the buffer and characterized before being placed in fresh buffer. Wet weight was measured using a scale, and swelling measurements were taken using calipers.

All results are presented as the average ± standard deviation from the evaluation of three samples replicated three independent times (n = 9). Statistical significance was evaluated using a one-way analysis of variance at a 95% confidence interval (p < 0.05) using OriginPro 2017 (OriginLab Corporation, MA).

Polyampholyte hydrogels are of interest for tissue engineering applications because of their demonstrated multifunctionality.³¹ Therefore, it was important to verify that their previously demonstrated nonfouling and conjugation capabilities were not impacted by the additional cross-linker species investigated in this study. Nonspecific adsorption of FITC-BSA was used to verify the nonfouling properties using fluorescent microscopy. Figures 2(a)–2(c) show representative images of TMA:CAA hydrogels following exposure to FITC-BSA and control hydrogels that have not been exposed to any protein. There are minimal to no differences in the fluorescence between these samples. Therefore, the results clearly show that the nonfouling properties are not impacted by the different cross-linker species. The protein conjugation capabilities were also verified with FITC-BSA and fluorescent microscopy. The results can be seen in Figs. 2(d)–2(f). These images show hydrogels with conjugated FITC-BSA adjacent to control samples. FITC-BSA conjugated samples showed bright fluorescence indicating that the FITC-BSA is successfully conjugated to the hydrogels, again demonstrating that the known conjugation capabilities are not impacted by the variation in cross-linker species.

The physical properties under compression were evaluated next, to identify impacts of the cross-linker length and density. Figure 3(a) shows representative stress–strain curves for 4X cross-linker density hydrogels with each cross-linker species and representative stress–strain curves for the 1X and 2X cross-linker densities can be seen in Figs. S1(a) and S1(b).⁴¹ It can be seen that decreasing the cross-linker chain length increases both stress and strain at failure. The quantitative analysis of multiple samples, in Fig. 3(b), shows that increased cross-linker density improves the stress at failure for all the cross-linker species, as expected.³¹ The maximum stress at failure value across all samples was found for the 4X cross-linker density DEGDMA hydrogels with an average stress of 860 ± 140 kPa. This value was statistically greater than the max stress at failure for both 4X cross-linker density TEGDMA and Tetra-EGDMA samples. At the 1X cross-linker density, DEGDMA samples had the statistically lowest fracture stress. DEGDMA samples also demonstrated the largest impact as cross-linker density was increased. The only statistically significant difference in the 2X cross-linker density samples was seen between the TEGDMA and Tetra-EGDMA results, with TEGDMA samples having a higher stress at failure. The maximum strain at failure results is shown in Fig. 3(c). Only the Tetra-EGDMA samples showed significant variation in the max strain at failure as a function of cross-linker density, with a statistically significant drop between the 1X and 4X cross-linker density samples. Across all cross-linker densities, DEGDMA samples presented the highest max strain at failure, with the greatest average strain of 0.51 ± 0.04 for the 2X cross-linker density, but this value was not statistically significant from the 1X and 4X DEGDMA density results. At the 4X cross-linker density, there were statistically significant differences between all of the cross-linker species. The results indicate that increasing cross-linker chain length decreases the compressive strain at failure. Finally, the elastic moduli of the various samples were evaluated, and these results are shown in Fig. 3(d). For the most part, the cross-linker species did not impact the elastic modulus under compression. The modulus increased uniformly across the cross-linker species with increasing cross-linker density and the maximum elastic modulus was seen for 4X density DEGDMA samples with a value of 3300 ± 825 kPa.

The physical properties under tension were evaluated, again to identify the influence of the cross-linker species and density. The tensile properties of this family of materials have not been previously tested. Figure 4(a) shows representative stress–strain curves for 4X cross-linker density hydrogels in tension, and representative curves for the 1X and 2X cross-linker densities can be seen in Figs. S2(a) and S2(b).⁴¹ Each curve presents a different profile with a low strain inflection point seen with the Tetra-EGDMA hydrogel curves (∼18% of total strain), no inflection point with the TEGDMA hydrogels, and a mid strain inflection point in the DEGDMA hydrogel curves (∼56% of the total strain). The TEGDMA hydrogel curves have a nearly linear stress–strain curve at a 4X cross-linker density, while the 1X and 2X cross-linker density curves have profiles similar to DEGDMA. The slope of Tetra-EGDMA stress–strain curve is similar to that of TEGDMA after the inflection point, while the slope of the DEGDMA stress–strain curve is similar to TEGDMA before its inflection point. The results in Fig. 4(b) indicate that the fracture stress at failure increases with increasing cross-linker density, again as expected. However, in contrast to the compressive fracture behaviors, Tetra-EGDMA demonstrates the highest values, reaching a maximum tensile fracture stress of 58 ± 14 kPa for the 4X cross-linker density hydrogels. Tetra-EGDMA also has the largest change as the cross-linker density is increased. Similar to the compression study, the fracture strain decreased with increased cross-linker density. However, a much more significant change was seen for all three cross-linkers species as the cross-linker concentration increased. As before, there was little significance between the strain values as a function of cross-linker species, at a given cross-linker density, though DEGDMA had a greater strain value under all conditions. Finally, Fig. 4(d) shows the average elastic modulus for these samples under tension in both the toe region (solid) and linear region (patterned). The elastic modulus was greater in the toe region than it was in the linear region and the overall elastic modulus increases with increasing cross-linker density. As with the maximum stress under tension data, the Tetra-EGDMA samples have the highest elastic modulus values and show the largest dependence on cross-linker density.

The degradation behaviors of polyampholyte hydrogels have not been characterized previously. For this degradation study, 1X cross-linker density hydrogels with the three different cross-linkers species were monitored for 113 days in Sorenson's buffer with acidic (4.5), neutral (7.4), and basic (9.0) pH conditions. The degradation was monitored by tracking the sample swelling and weight changes over time and the results are summarized in Fig. 5. There were insignificant changes in the swelling and weight change behavior for all three cross-linker species under the neutral and basic conditions, over the time period evaluated here. Degradation occurred in the acidic environment for all three cross-linker species, ultimately leading to full sample breakdown after ∼110 days for the Tetra-EGDMA cross-linked samples. The Tetra-EGDMA hydrogels increased in both weight and swelling over time. Full breakdown was determined when the samples completely dissipated into the degradation medium. This time point was used as the endpoint for all of the degradation studies even though the full sample degradation only occurred with Tetra-EGDMA samples in acidic media. Conversely, the DEGDMA and TEGDMA cross-linked hydrogels decreased in both size and weight. They also did not reach full degradation like the Tetra-EGDMA hydrogels over the time period evaluated, although surface flaking was visible during the degradation study. These paired results suggest that the Tetra-EGDMA samples degraded via bulk degradation pathways, while the DEGDMA and TEGDMA degraded via surface degradation pathways, as discussed in more detail below.

The compression samples show that increasing the cross-linker density increases the stress at failure and elastic modulus, which has been seen in previous studies.^32,38 Specifically, the physical properties under compression for TMA/CAA hydrogels with a TEGDMA cross-linker have been tested previously and the trends between the previous results and this study are similar.³¹ The results in this study show that the available range of mechanical properties can be expanded even further by varying the length of the cross-linker species, suggesting a wider range of tunability. A similar range is also demonstrated under tension in this study as well.

Polyampholyte hydrogels cross-linked with DEGDMA had the lowest stress at fracture compared to TEGDMA and Tetra-EGDMA with a 1X cross-linker density, but presented the greatest stress at fracture with a 4X cross-linker density. It is hypothesized that this is due to differences in the side-chain packing density in these various systems. The DEGDMA cross-linker has the shortest chain length and this results in the most side-chain crowding and overlap between the pendant groups of polymer chains linked by this cross-linker. The overlap is eliminated as the cross-linker chain length is increased to TEGDMA and there is the least amount of crowding between pendant groups when the length is increased to the Tetra-EGDMA cross-linker. This is shown schematically in Fig. 6. While Fig. 6 presents an example of the most tightly packed pendant group alignment, it is readily acknowledged that the pendant groups will rotate around the polymer backbone in three-dimensions to reduce the overlap and resulting steric hindrance. However, as the cross-linker density is increased, it will reduce the degrees of freedom available for this pendant group rotation. This results in more overlap between pendant groups in hydrogels cross-linked with DEGDMA, which will not be seen in the TEGDMA and Tetra-EGDMA cross-linked systems. Further complicating the discussion is the fact that these hydrogels have been shown to have ∼90 wt. % water in their fully hydrated state.³¹ The fact that the maximum stress at failure for the DEGDMA samples shows the greatest change from the 1X to 4X cross-linker density, while going from the lowest to highest value, is attributed to the packing density, increased interactions between the pendant side chains, and hydrating water.

As mentioned previously, increasing the cross-linker chain length decreased the compressive strain at failure, specifically at high cross-linker densities. Additionally, there were no significant differences in strain for each cross-linker species with changes in the cross-linker density. This indicates that the compressive strain is not dependent on the amount of cross-linker in the hydrogel. The high weight percentage of water (∼90% (Ref. 31) leads to bulk incompressible behavior, based on the incompressible nature of water. Therefore, there is no change in strain as a function of cross-linker density. The electrostatic interactions between the pendant side chains, which are influenced by the cross-linker length, could also play a role in the strain behavior. As discussed above, there is potential overlap between the pendant groups in hydrogels formed with the DEGDMA cross-linker. As the system is compressed, the pendant groups can more easily slide past each other as they are potentially already overlapping. This contrasts with the pendant group interactions in TEGDMA and Tetra-EGDMA cross-linked systems. The electrostatic interactions between intrachain pendant groups in the TEGDMA and Tetra-EGDMA systems are hypothesized to lead to more tightly bound side chains, resulting in failure when the interchain groups are compressed together.

This packing density model can also be used to explain the trends seen in the tensile testing results. When the 4X cross-linker density stress–strain curves are examined [Fig. 4(a)], it is evident that all three cross-linker species have a portion of the curve that has nearly identical slope with each other. This occurs before the inflection point in the DEGDMA cross-linked samples, across the entirety of the profile for the TEGDMA cross-linked samples, and after the inflection point in the Tetra-EGDMA cross-linked samples. It is hypothesized that this slope is indicative of the simultaneous breakage of the cross-linker species and the electrostatic interactions between pendant groups of adjacent polymer chains, along with disruptions to the significant hydrating water-functional group interactions that occur throughout the hydrogel. In Fig. 6, it can be seen that the TEGDMA samples provide the best alignment of pendant side chains from adjacent polymer chains. Therefore, there is limited rearrangement prior to breakage of the electrostatic interactions, cross-linker species, and water interactions. The DEGDMA cross-linked samples initially have the same structure disruptions occurring under tension. However, the overlap of the pendant side chains from adjacent polymer chains leads to an extended domain where only the weaker electrostatic interactions and hydrating water-functional group interactions are being broken. This occurs after the inflection point in the stress–strain curve, where there is a significantly lower elastic modulus (slope). The opposite is true in Tetra-EGDMA cross-linked samples. Prior to the inflection point, the slope of the stress–strain curve is much more significant than after, and this is attributed to the breakage of the Tetra-EGDMA cross-linker. After the inflection point, the system has very similar behavior to the TEGDMA samples. While similar trends exist in the 1X and 2X cross-linker density stress–strain curves [Figs. S2(a) and S2(b)],⁴¹ they are not as prevalent because the pendant group packing densities are lower.

The packing density model is also supported by the degradation study results. The Tetra-EGDMA cross-linked hydrogels present a different degradation behavior than the DEGDMA and TEGDMA cross-linked hydrogels under acidic conditions, as observed directly with the samples and in the quantitative analysis of the weight changes and swelling behaviors. All of the hydrogels degrade via acid hydrolysis of the ester groups present in both the cross-linkers and monomers.^1,9 The results indicate that the DEGDMA and TEGDMA cross-linked hydrogels degrade via surface degradation pathways (shrinkage and weight loss; visible surface flaking), while the Tetra-EGDMA hydrogels degrade via bulk degradation pathways (swelling and weight gain; full sample dissipation).^39,40 The strong electrostatic interactions that occur between adjacent polymer chains cross-linked with either DEGDMA or TEGDMA cross-linkers provide steric hindrance that limits accessibility to the ester groups of the cross-linker where overall degradation occurs (Fig. 6). Therefore, degradation only occurs at the surface and outmost layers of the hydrogels. In contrast, the pendant groups of adjacent polymer chains in hydrogels cross-linked with Tetra-EGDMA have wider spacing, facilitating easier penetration into the hydrogel and bulk degradation. Furthermore, full degradation was only observed with the Tetra-EGDMA cross-linked samples over the ∼110 day degradation time frame, at which point the samples fully dissipated into solution.

In this work, polyampholyte hydrogels were formed from equimolar mixtures of TMA and CAA charged monomer subunits with three different ethylene glycol based cross-linkers (di-, tri-, and tetra-ethylene glycol). Fluorescent microscopy evaluations confirmed that the multifunctional nonfouling and protein conjugation capabilities of TMA/CAA hydrogels were not impacted by the cross-linker species. The compression and tensile properties of this family of hydrogels were evaluated across a range of cross-linker densities and the results suggest that the cross-linker length influences the pendant side-chain packing density, which in turn impacts the macroscale physical characteristics. The cross-linker length and subsequent packing density also impact the acid hydrolysis degradation pathway for this family of hydrogels. Hydrogels cross-linked with DEGDMA and TEGDMA degraded via surface degradation pathways, while those cross-linked with Tetra-EGDMA degraded via bulk degradation pathways. Molecular dynamics investigations are ongoing, and they will provide further mechanistic insight into this proposed packing density model due to the complex interplay between side-chain density, electrostatic interactions, and water-functional group interactions. However, the results from this study clearly demonstrate the broad tunability of polyampholyte hydrogels for tissue engineering applications based on controlling the spacing and interactions between the underlying charged pendant groups.

The authors acknowledge Indrajit Charit for allowing us access to his Instron machine for compression testing and Nathan Schiele for assisting us with tensile testing. This work was funded by the Department of Defense, Congressionally Directed Medical Research Program through Grant No. W81XWH-15-1-0664.