Swelling via Impedimetry Using Specifically Adhered Hydrogels on Co‐Planar Microfabricated Electrodes

Responsive hydrogels adhered to microfabricated electrodes find applicability as chemical or biological sensors and in electro‐stimulated drug delivery. A well‐defined method of cleaning, surface modification, and surface functionalization of microlithographically‐fabricated biochips composed of heterogeneous, abio surfaces (gold and glass) is presented for the reproducible adhesion of responsive hydrogels. The method uses cleaning approaches adapted from the semiconductor electronics industry and combines these with reactive organosilane chemistry to achieve the specific (covalent) attachment of UV cross‐linked, poly(HEMA‐co‐PEGMA‐co‐HMMA)‐based hydrogels. Specific attachment of hydrogels via acryloyl‐poly(ethylene glycol)‐3500 n‐hydroxysuccinimide (APNHS)‐functionalized surfaces and subsequent hydrogel hydration resulted in the strongest adhesive force as determined by centrifugal adhesion testing. Comparison with substrates functionalized via hydroxyl‐poly(ethylene glycol)‐3500 n‐hydroxysuccinimide (PNHS) confirmed the superiority of adhesion involving covalent bonding (APNHS) (4.48 kPa) versus hydrogen bonding (PNHS) (1.29 kPa). Adhered, fully hydrated and dehydrated hydrogels are characterized by Electrochemical Impedance Spectroscopy (EIS) and their hydration kinetics determined using impedimetry at a rationalized frequency. Impedimetry confirmed that p(HEMA‐co‐PEGMA‐co‐HMMA) hydrogels have an equilibration time of ≈30 min, a diffusion‐dependent rate coefficient k1 = 0.311 s−0.5 and relaxation‐dependent coefficient k2 = −0.022 s−1. Hydrogel swelling may be studied by impedimetry to fashion biomedical devices for co‐joined, real‐time biosensing with electro‐stimulated drug delivery.


Introduction
Biosensors based on responsive hydrogels have gained remarkable popularity within the field of biomedicine because of their high water content, tunable biotechnical properties, and potential for approval by the US FDA for indwelling performance. [1,2][12][13] We have fashioned responsive hydrogels with the responsive sensitivity of enzymes [14,15] and possessing redox mediators for use with enzymes [16] in generation-2 amperometric biosensors.21] Minimally invasive and intramuscular biosensors are beginning to revolutionize diagnosis and treatment within modern medicine -allowing for real-time, sustained monitoring of single or multiple biomarkers that could yield substantial temporal data of value to physicians and emergency medicine professionals [22,23] while providing for data-driven response, such as controlled drug release -a form of theragnostics.However, long-term biosensor implantation is not always feasible due to materials failures [24] and the foreign body response it elicits, during which the body's immune system detects and defends against foreign and/or abio materials. [25,26],26] To resolve issues of compromised biocompatibility and function, implantable biosensors ought to be encapsulated with stronglyadhered, biocompatible hydrogel coatings, which, being highly hydrophilic and molecularly engineered for biocompatibility, are able to mitigate biofouling. [25]Moreover, impedimetric biosensors require stable electrode-hydrogel interfaces that are well adhered. [12]dhesion is the assembly of an adherend (substrate), an adhesive and the associated interface in which dissimilar atoms or molecules of a liquid and solid can "stick" together, forming an adhesive joint. [30]The work of adhesion corresponds to the adherend's degree of wettability as described by the Young-Dupre equation, where W A is the sum of liquid and solid surface energies,  L and  S (measured in units of either mJ m −2 or its equivalent of mN m −1 ), excluding the energy of the interface,  SL [27,31]   (Equation 1).These terms can be applied to arrive at the equation for the equilibrium spreading pressure,  e (Equation 2).
At  e , the free surface energy of the solid is fully reduced, meaning that the maximum amount of interfacial bonding has been achieved and adhesion has been optimized.When polar forces are also present, such as at the interface of a high-energy substrate (e.g., surface modified glass) and polar liquids (e.g., hydrogel pre-polymer and water-hydrated hydrogels), Equation 2simply becomes Equation 3, as explained by Fowkes et al. [32]  e = 2 In addition to intermolecular interactions (London dispersion forces, dipole-dipole forces, and hydrogen bonds), intramolecular interactions (covalent, ionic, or metallic bonds) also affect the strength of adhesion. [33,30]Due to the difficulty in accounting for all these interactions when calculating adhesion, several experimental approaches have been developed to assess the adhesion of an adherend to various substrates. [30,34,35]However, few test methods are feasible for testing adhesion of hydrogel membranes to abio surfaces since many of these methods require the adherend and substrate to be held together with fixtures (e.g., by clamping).The use of fixtures can cause undesired shearing and fracture within the hydrogel, resulting in measurements of cohesion (hydrogel integrity) instead of adhesion to the substrate.In this paper, a centrifugal force method [36,37,34] was used to investigate the adhesive strength (a force, measured in Newtons) of 2-hydroxyethyl methacrylate (HEMA)-co-polyethyleneglycol methacrylate (PEGMA)-co-N-[Tris(hydroxymethyl)methyl] acrylamide (HMMA) [p(HEMA-co-PEGMA-co-HMMA)] based hydrogels to glass following a meticulous process of glass cleaning, surface modification, and functionalization.By optimizing surface cleaning and functionalization protocols, adhesion between hydrogels and heterogenous, abio substrates could enable the use of hydrogels as coatings for implantable medical devices for anti-biofouling and biosensing applications.
Adequate hydrogel-substrate adhesion is also of significance for characterizing and employing hydrogel swelling kinetics in electrostimulated drug release, bioelectronic biosensors, [38] and wearable health IoMT devices.The most common method of determining hydrogel swelling kinetics is gravimetric analysis, in which a swelling hydrogel structure of defied shape (slab, sphere, disc, etc.) is removed from solution at various time intervals and temporal changes in its mass measured. [39,40]ith repeated removal of the hydrogel from solution and the potential for errors in weighing, this system is not well-suited for obtaining an accurate, continuous picture of swelling kinetics consistent with bioelectronic applications.Thus, a reliable method was developed in this work to allow for continuous monitoring of swelling behavior over an extended period without sacrificing the simplicity and sample preservation offered by gravimetric analysis. [39,41]This method employed Electrical Impedance Spectroscopy (EIS) [42] to fully characterize dehydrated and equilibrium hydrated hydrogel membranes followed by impedimetry, which monitored impedance changes within swelling hydrogels that were covalently bonded to microlithographically fabricated coplanar Independently Addressable Microband Electrode-co-Interdigitated Microsensor Electrodes (IAME-co-IME) biochips. [43]Coplanar microfabricated electrodes have been widely used in generator-collector amplified electrochemical sensing [44] and in conductimetric biosensing. [45]ased on the principle of impedimetric transduction, functionalized microelectrodes are able to convert changes in analyte concentration (e.g., DNA, enzyme, ion) to a detectable, electronic signal. [46,47]Recently, the authors demonstrated how two such enzyme-conductometric biosensors could be combined as discreate inputs within a Wein oscillator to produce a single ratiometric frequency output. [38]Likewise, microfabricated electrodes can be exploited for the measurement of ion concentration changes within adhered hydrogels during swelling.The coplanar, interdigitated nature of the electrodes used herein generates a complex, fringing electrical field that specifically makes such measurements possible.The fringe field allows for interrogation of the near-surface region, [48] compared to the bulk that is sampled by electric fields produced by parallel, opposing electrodes, making coplanar, interdigitated electrodes highly suitable for thoroughly detecting bioelectronic biosensor responses. [49]mpedance data collected during hydrogel hydration could then be related to swelling kinetics based on: i) associations between impedance and swelling kinetics shown in the literature, [50][51][52] ii) well-established transport models for Fickian diffusion and Case II transport, [53] Işık, [54,55] and iii) evidence for non-Fickian transport resulting from polymer chains inhibiting water penetration within the hydrogel. [53]This simple, continuous method of obtaining hydrogel swelling kinetics could improve the implementation of hydrogels in drug delivery by allowing various hydrogel compositions to be quickly and accurately screened for the swelling kinetics required in a specific application.Likewise, knowledge of swelling profiles could be of significance in bioelectronic biosensing, wherein an electrode encapsulated in a stimuli-responsive hydrogel membrane must be calibrated to correlate changes in hydrogel swelling with linked biochemical reactions.Electrical impedance and its temporal equivalent, impedimetry, could then be the basis for monitoring fluctuations in bioanalyte concentrations (e.g., pH/acidosis using a pH-responsive hydrogel [56] ) in order to detect aberrations in patient health, [57] Wilson and Guiseppi-Elie . [2] Experimental Section

Cleaning, Surface Modification, and Functionalization of Glass Slides
With the intent of establishing specific attachment via covalent bonds to the surface, surface modification to introduce reactive, terminal acryloyl groups onto glass for the purpose of subsequent hydrogel covalent bonding was adapted from previous reports as shown in Figure 1. [60,58,59,61]Briefly, glass slides (Gold Seal Plain Slides 3011-002, Thermo Fisher Scientific, Waltham, MA) were first solvent cleaned by sonication in acetone (solvent cleaned), UV Ozone treatment (Boekel Industries Inc., Feasterville, PA), followed by piranha (3:1 sulfuric acid (H 2 SO 4 ) and 30% hydrogen peroxide (H 2 O 2 ) cleaning to remove organic residue and hydroxylate the silica (activated).Slides were then dried completely within a vacuum oven at (>110°C) and allowed to cool to room temperature prior to standardizing hydroxyl group density with plasma treatment (Basic Plasma Cleaner PDC-32G, Harrick Plasma, Ithaca, NY) (Figure 1a).Activated glass slides were incubated in 0.01 m solution of ϒ-APS in toluene for 2 hours in the dark to promote condensation to silanol and physical adsorption of ϒ-APS molecules to hydroxyl functionalities (ϒ-APS Physisorbed) (Figure 1b).After gradual transition from toluenebased solution to ethanol, slides were heated at 40°C for 20 min and 110°C for 10 min in a vacuum oven to force condensation of adsorbed silanol to the hydroxyl functionalities and so produce covalent bonding between ϒ-APS and hydroxyl molecules (ϒ-APS Chemisorbed) [60,58,59] (Figure 1c).Slides with covalentlyattached ϒ-APS groups were submerged in deionized water for 24 h to facilitate the extension of hydrophilic amine groups from the silica surface and the hydrolysis of unbound ethoxy functional groups (ϒ-APS Chemisorbed | DI) [62] (Figure 1d).Glass slides were submerged in a 0.1 mm solution of APNHS in HEPES buffer (pH 8.5) to cause covalent conjugation of the primary amine group of ϒ-APS to the n-hydroxysuccinimide group of APNHS (Figure 1e).Finally, hydrogels were cast onto the various substrates, photo-cross-linked with UV, and hydrated in DI water, PBS 7.4, or cell culture media (Figure 1f).

Measurement of Contact Angles and Hydrogel Attachment
Contact angles were measured by sessile drop method using a Remé Hart Model 500 Advanced Contact Angle Goniometer / Tensiometer (Ramé-Hart instrument co.).Contact angles were measured at room temperature after 30 s following application of a 14 μL drop of DI water and are the results of n = 12 replicates.Hydrogel attachment (distinct from hydrogel adhesion, next section) was measured following printing of hydrogel dope onto various prepared borosilicate glass slides.Hydrogel dope was printed onto glass slides using a Biodot AD 3400 development to pilot production system in non-contact aspirate and dispense mode (BioDot, Irvine, CA).Drops were 0.1 μL and were 24, 48, or 72 features.Slides of cured p(HEMA-co-PEGMA-co-HMMA) spots were vertically arranged within resealable polypropylene slide mailers while immersed in PBS 7.4 buffer at RT for up to 7.5 days.At periodic intervals, slides were removed and the number of spots remaining attached on each slide was counted.

Measurement of Interfacial Adhesion by Centrifugal Method
Adhesion was measured on cleaned (-OH), physisorbed ϒ-APS, chemisorbed ϒ-APS, ϒ-APS that was incubated in DI water, and ϒ-APS that was covalently conjugated and thus APNHS functionalized glass slides.Formulated hydrogel cocktails were cast as discs onto each type of slide with the aid of 24-well silicone isolators (diameter ϕ = 2.0 mm, thickness t = 1.6 mm) (JTR24R-2.0,Grace Biolabs, Bend, OR) that were positioned at defined radii, r, from the center of the slide.The cocktail was then UV-crosslinked for 5 min and gels were hydrated overnight in HEPES buffer (pH = 7.4).Adhered and hydrated hydrogels were blotted with a Kimwipe® and the center of the glass slide was securely vacuum sealed upon a spin coater (Spin Coater WS650S-6NPP/Lite, Laurel Technologies Corporation, North Wales, PA).Each slide was rotated with an initial acceleration of 600 RPM s −1 and starting speed of 600 RPM, which was increased by 100 RPM consecutively until all hydrogel discs were dislodged from the treated glass slide within the spin coater chamber.Documented RPM values were then used to calculate the dislodging force, F (Newtons), of hydrogel-slide adhesion using the centripetal force equation, where m represents the mass (grams) of the dislodged hydrogel disc, r represents the radius (mm) corresponding to the center of the hydrogel disc to the center of the glass slide, and  represents the angular speed (radian s −1 ).This procedure was repeated for replicates (n = 3) of the five distinct slides following -OH activation (Figure 1a), physical adsorption of ϒ-APS (Figure 1b), covalent attachment of ϒ-APS (Figure 1c), extension of ϒ-APS with DI immersion (Figure 1d) and, finally, functionalization by APNHS (Figure 1e).As a direct comparison of the adhesion force due to covalent bonding enabled by APNHS, the procedure was also executed with final functionalization by PNHS that possesses a hydroxyl end group for hydrogel hydrogen bonding but was not capable of covalent bonding.

Surface Modification and Hydrogel Attachment to IAME-co-IME Biochip
The IAME-co-IME  slides with the following exceptions and/or additions: (1) to minimize undercutting of the adhesion-promoting titanium tungsten (TiW) layer, the piranha cleaning step was eliminated, and (2) to remove adsorbed siloxane from the gold digits of the chip, a cathodic cleaning step was added.Following silanization and before programmed temperature increase, chips were cathodically cleaned by the application of 40 cyclic voltammetric sweep cycles at 100 mVs −1 over the range −1.2 V to 0 V versus Ag/AgCl in 0.1 m PBS (pH = 7.5) buffer at RT. [65] Chips then resumed the modification process described above for glass slides.Formulated hydrogel discs were cast onto the regions of interdigitation or the microband regions using the same method as stated above.

Electrochemical Impedance Spectroscopy and Impedimetry
EIS and impedimetry were carried out using a VersaStat 4 (Princeton Applied Research, AMETEK, Inc., Oak Ridge, TN).EIS was studied over the range 0.01 Hz -1.0 MHz using a 10 mV p-t-p interrogating sine wave.Measurements were taken in phenol red-free DMEM cell culture media at RT on a hydrogel with thickness exceeding the consideration /3 -the depth of electrical signal penetration from a biochip into media [51] which is ≈2.
where the real impedance, Re(Z T ), relates to resistance and Im(Z T ) represents the imaginary impedance that results from reactance -the component of impedance caused by both capacitance and inductance. [66]With the membrane resistance only apparent in the real part of impedance, the Im(Z T ) portion was ignored and R M was solved from Re(Z T ) (Equation 6).
The R M vector was solved using Re(Z T ) data output by Versastat software and with vectors, or "known variables," of the same size for R G , C G , R CT , and C DL containing incremented values between dehydrated and equilibrium hydrated base values shown in Table 1.Taking the inverse of membrane resistance to generate conductance values, [67,68] conductance was graphed and modeled with the common equation for anomalous swelling to indicate swelling kinetics of the p(HEMA-co-PEGMA-co-HMMA) hydrogel membranes.

Results and Discussion
The use of organo-silanes such as -APS (also called APTMS) to modify, and PEG derivatives to functionalize, glass surfaces are not new. [7]Intended to exploit the transducer-active responses of electroactive polymers, the authors and others have previously developed interfacial engineering methods based on similar chemistries for the specific attachment of polyanilines, [69,70] polypyrroles, [7] various hydrogels, [10,18,71] and electroconductive hydrogels [72,64,58,59,73] to the surfaces of heterogeneous biochip device surfaces.[76][77][78] Additionally, the authors have likewise applied hydrogels to similarly fashioned microlithographically fabricated array electrodes [10,49] in the development of sensors.][81] However, the actual force or work of adhesion of hydrogels to device surfaces has not been rationalized and measured at each step in the multi-step fabrication process.Delineating the work of adhesion at each step in the process is reported here for the first time.

Confirmation of Reactivity of Surface Amines
To confirm reactivity of the surface amines of Y-APS, NHS fluorescein (6-Fluorescein N-hydroxysuccinimide ester) (Ex/Em = 491/516 nm), a reactive derivative of fluorescein, was used to react with and form covalent conjugates with the free amines on the surface of amino-silanized chips.The resulting amide bonds that attach fluorescein to the surface served as surrogates for the similarly fashioned reactions of acrylatepoly(ethylene glycol)−3500 n-hydroxysuccinimide (APNHS) and hydroxyl-poly(ethylene glycol)−3500 n-hydroxysuccinimide (OH-PNHS).Figure 3 illustrates the steps in this reaction and shows the results following cathodic cleaning of the gold electrodes.Cathodic cleaning removes the adsorbed silanols from the metallic surface thereby increasing the contrast between the metallic digits and the glass.Image J (v1.54 h) analysis showed a five-fold increase in contrast following cathodic cleaning.

Quantification of Hydrogel-Substrate Attachment
Hydrogel spots that were printed onto engineered surfaces and incubated in PBS 7.4 at room temperature were examined periodically over a period of 7.5 days.Figure 4 shows the percentage of printed spots remaining attached to each of the various surfaces over the course of the 7.5-day immersion.This attachment study simulates the response of the hydrogel/glass interface as the hydrogel swells to equilibrium and remains in contact with a physiological-like fluid.As expected, the solvent cleaned and UV-Piranha activated glass surface lost 100% of its hydrogel spots by the first inspection (12 h).Interestingly, solvent cleaned surfaces resulted in 10% spot retention for up to 40 h, suggesting that solvent cleaning only may not be adequate to remove mineral residues that contribute to poorly cleaned surfaces.The glass surface, which was functionalized with 3-TPMA or functionalized with ϒ-APS and subsequently derivatized with Acryl-PEG(3500)-NHS showed perfect attachment, retaining 100% of printed hydrogel spots for the duration of the 7.5 days of testing.In both cases the terminal -acryloyl group enabled covalent coupling of the chemisorbed silane within the polymer network.The presence of continuous covalent bonding from the polymer to the surface assured 100 percent retention of hydrogel spots.Interestingly, surfaces functionalized with ϒ-APS and subsequently derivatized with OH-PEG(3500)-NHS also showed good attachment, gradually falling but retaining 80% of printed hydrogel spots for the duration of the 7.5 days of testing.The concerted hydrogen bonding interactions along the reptating PEG chain allowed for this attachment.For the OTS treated group, there were 17% of swollen hydrogel spots remaining after 180 hr.The OTS surface had a water contact angle of 106˚and is therefore highly hydrophobic, allowing for purely dispersive interaction.It is noteworthy that this dispersive interaction produced better hydrogel attachment results than the highly hydrophilic, UV-Piranha activated glass surface (contact angle = 9˚).In a confirmatory experiment, glass slides that were functionalized with ϒ-APS and subsequently derivatized with Acryl-PEG(3500)-NHS, were each half-covered with aluminum foil and the slides exposed to 5 min of UV irradiation to promote crosslinking.Each zone was subsequently printed with 24 spots and immersed in PBS 7.4 as indicated.After 180 h, all spots in the foil-covered zone remained attached while 60% of spots in the UV-exposed zone were dislodged.The covalent coupling enabled by the -acryloyl group of the chemisorbed silane was necessary for 100% attachment.

Quantification of Hydrogel-Substrate Adhesion
Contact angles of DI water on each of the molecularly engineered surfaces are shown in Figure 5a.The solvent cleaned and UV-Piranha activated surface had the smallest contact angle of 9.0°±0.7°reflectingthe enhanced wettability the comes with an abundance of surface -OH groups, typically 4.59 -OH per nm 2 on fused silica. [82]The contact angle of physisorbed -APS-P (22.0°±2.6°)was markedly different than the chemisorbed -APS-C (34.0°±5.1°).The additional condensation of silanols (≡Si-OH) during the heating step serves to reduce the abundance of hydrolyzed but not condensed silanols, rendering the surface overall more hydrophobic because of the formation of silyl ether (Si-O-Si) These silyl ethers are formed among  molecules within the organic layer and with -OH groups of the glass surface.Incubating the -APS-C surface in DI water did not alter its contact angle with DI water.The conjugation with acryl-PNHS and OH-PNHS produced similar contact angles of 19.0°±2.3°and18.0°±1.5°,respectively.The wettability of these derivatized surfaces is likely dominated by the PEG 3500 chain and is largely agnostic to the -terminal functionality.Centrifugal force testing of adhesive force has been used to quantify cellular attachment [83,84] and hydrogel coatings. [85]In all cases, the adhesive strength is measured as the resistance to disruption in the presence of a measured centrifugal force.Centrifugal force testing of hydrogel membrane adhesion to various surface modification and functionalized surfaces, confirmed that modification followed by chemical functionalization with APNHS to promote covalent bonding during surface modification of glass slides resulted in significantly improved hydrogel-substrate adhesion compared to hydrogel-substrate adhesion at all prior steps of the surface modification process (p < 0.05) (Figure 5b).APNHS functionalization yielded an average adhesion of ≈4.5 kPa, nearly a threefold adhesion advantage relative to adhesion to cleaned glass slides of 1.6 kPa.Treatment with plasma-induced hydroxyl groups, physisorbed ϒ-APS molecules and chemisorbed, extended ϒ-APS molecules did not produce significantly different adhesion values, wherein adhesion in these treatment groups was the result of hydrogen bonding between hydroxyl or unprotonated amino end groups, respectively, and functional groups within the p(HEMA-co-PEGMA-co-HMMA) hydrogel (i.e., primarily free HEMA hydroxyl groups).ϒ-APS molecules received from Sigma were assumed to be in protonated form and unaffected by toluene considering its purely organic, water-immiscible nature that is unable to form or dissociate hydronium ions.When comparing physisorbed to chemisorbed ϒ-APS molecule treatments, it was noted that the covalent bonding of ϒ-APS ethoxy functionalities to surface hydroxyls from cleaning steps had a considerable effect on adhesion.Subsequent slide treatment by extended immersion in DI water was intended to orient the amine groups of ϒ-APS out-ward, preventing downward attraction to the silica surface, and to hydrolyze residual ethoxy functionalities as indicated in previous literature. [62]However, DI immersion seemed to not have a significant effect on adhesion, yielding adhesion values less than those of the previous treatment step.One potential explanation for this phenomenon is the change in amino group protonation between ϒ-APS solution and DI water.Amino end groups considered unprotonated during ϒ-APS solution incubation and heat curing likely experienced significant protonation when submerged in DI water because of the pH of the surrounding solution (pH<7.0)being less than that of the amino group acid dissociation constant (pKa = 9.6).Finally, treatment with APNHS culminated in the highest degree of hydrogel-substrate adhesion due to covalent bonding of the acryloyl end group of APNHS within the hydrogel as well as the flexibility and length of the PEG-3500 chain.While the utility of covalent bonding is readily apparent, the PEG-3500 chain included within APNHS also served a crucial role of ensuring that the hydrogel's extreme swellability during hydration did not force breakage of covalent bonds.This was possible as the flexibility and length of the PEG chain allowed it to uncoil and stretch between the substrate surface and hydrogel membrane during hydrogel swelling.Because of this and the large quantity of oxygen atoms in PEG that could hydrogen bond within the hydrogel membrane, the PEG-3500 unit was an important component for achieving and preserving strong adhesion.
Other methods, such as scratch and delamination tests, [86] do not provide information on the force of initial adhesion and so do not quantify the strength of the adhesive force.Furthermore, the centrifugal force applied to the hydrogel is representative of the shearing forces that may be applied to the system in-vivo as tissue layers move independently. [87]There is appreciable variation in the results of adhesion force testing, evident in the standard deviation, that could be caused by multiple sources including, i) reproducibility of the surface coverage during surface modification, ii) variation in the contact surface area between the glass substrate and hydrogel cocktail (assumed to be constant) (e.g., if the cocktail did not fill the well of the silicone mold completely or leakage under the well wall), and iii) titration errors resulting in variations of mass of hydrogel cocktail transferred.Given the large value of replicates, n = 38, statistical analysis strongly supports the conclusion that covalent bonding via surface functionalization improves adhesion force by 300%.

Comparison of Polymerizable APNHS and Non-Polymerizable -OH End Groups
To investigate whether the polymerizable acryloyl end-group of Acryl-PNHS was a primary factor in creating strong adhesion, adhesion between APNHS-treated slides was compared to that of slides treated with hydroxyl-poly(ethylene glycol)−3500 nhydroxysuccinimide (OH-PNHS) (Figure 5b).Because of identical precursor treatment of slides and the identical length of the PEG spacer in OH-PNHS to that of the PEG spacer in Acryl-PNHS, changes in adhesion were speculated to directly result from the presence of covalent bonding of the acryloyl group versus hydrogen bonding the hydroxyl group within the hydrogel.Analysis of adhesion values obtained for such slides showed Acryl-PNHS-treated slides to have significantly stronger adhesion than their OH-PNHS counterparts according to a t-test (p<0.01,n = 38), confirming that covalent bonding between hydrogel and immobilized functional groups on the glass resulted in an increased hydrogel-substrate adhesion when compared to the hydrogen bonding.Thus, while the contact angles were similar, the adhesion was significantly different.

Impedance and Impedimetric Response of p(HEMA-co-PEGMA-co-HMMA) Hydrogel
Considering the physical geometry of the co-planar, metalon-glass electrodes and the electrochemistry of the electrodeelectrolyte interface of the hydrogel-biochip system, EIS data was modeled as an R M (Q G R G )(Q DL R CT ) equivalent circuit in ZSimp-Win software, where Q G R G represented the physical geometry of the co-planar, Au/Ti-on-glass electrodes and the Q DL R CT ) represented the electrochemistry of the electrode-electrolyte interface of the hydrogel-biochip system. [49]90] The second constant phase element of the circuit model was attributed to the double layer capacitance, Q DL , that forms at the electrified gold-hydrogel interface between the layer of ions partitioned into the hydrogel and the layer of ions at the electrode surface. [42,89,91]Transfer of electrons across the Helmholtz plane of the double layer establishes the Faradic resistance represented by R CT .This dual (Q DL R CT ) behavior is followed by the purely resistive behavior of ionic transport through the hydrogel membrane represented by R M . [42,49]he impedance spectra of both dehydrated and equilibrium hydrated hydrogels show two distinct semi-circles suggesting that the same equivalent circuit model may be applied to both states.However, the equivalent circuit parameters were quite different (Table 1).The value of R M (Ω mm 2 ) of 2.73 ×10 1 and 8.34 ×1°, respectively, suggests a 327% increase in R M in going from dehydrated to hydrated state.

Impedimetric Measurement and Monitoring of Hydrogel Swelling
The impedimetric response of the hydrogel was measured as changes in membrane resistance, R M , during hydration of the covalently adhered hydrogel.Impedimetry was done at the interrogation frequency,  i .To determine the interrogation frequency, the characteristic frequencies ( c ) of the hydrogel membrane were first found in both dehydrated and equilibrium hydrated states.Characteristic frequencies are defined as the frequencies corresponding to the topmost point of a Nyquist semicircular plot (Figure 6), where the imaginary impedance due to capacitance is highest and from which the characteristic time constants,  CT = R CT C DL and  G = R G C G , corresponding to each dispersion (semi-circle) are defined. [92]At these frequencies, the partial derivative of the imaginary impedance with respect to real impedance is zero, meaning that capacitive and resistive changes to impedance are independent.Because of this, impedance changes measured and monitored at these characteristic frequencies because of hydration depend only upon changes in R CT and R M values.The characteristic frequencies for the hydrogel in dehydrated and hydrated states were found to be 0.01 Hz and 0.25 Hz, respectively, and were then used to calculate hydrogel cut-off frequencies using Equation 5.
Cut-off frequencies,  1,2 , are defined as the first point at which the phase angle of a Bode plot in a typical R(QR) system reaches 45°, and vary according to  c , R M , and R CT .At low frequencies, resistances have an additive effect wherein large R CT values overwhelm the influence of R M .In contrast, the upper cut-off frequencies for both dehydrated and hydrated states of the hydrogel, denoted as  2d and  2h , allow for the influence of R M on the system to become readily apparent and were therefore selected as the frequencies of interest for interrogation.Multiple calculations of  i were compared, including the average, weighted average, and root mean square of  2d and  2h .These were calculated as 0.13 Hz, 0.22 Hz, and 0.18 Hz, respectively.The weighted average of  2d and  2h (0.22 Hz) was selected as the interrogation frequency for its ability to best capture changes in membrane resistance over the course of membrane hydration.

Determination of Hydrogel Swelling Kinetics Through Impedimetry
[52] This relationship makes sense considering that that the resistance of hydrogel-biochip system is largely influenced by the membrane resistance, R M , which is the primary circuit element influenced by hydrogel swelling.Thus, conductance, the inverse of membrane resistance, was modeled by implementing the following two-term equation for anomalous hydrogel swelling kinetics initially proposed by Peppas, wherein the first term represents swelling by Fickian diffusion and the second term represents swelling by first-order, Case-II transport Işık. [54,55]M t and M ∞ denote fluid uptake of the hydrogel at time t and at equilibrium, while k 1 and k 2 represent the diffusion-controlled swelling rate constant and relaxationcontrolled swelling rate constant, respectively, [93,94] Işık. [54,55]Several constraints exist when applying Equation 6 in mmodelingthe swelling behavior via conductance monitoring, namely that the term representative of Fickian diffusion is only valid up to 60% swelling and that Case-II transport only remains linear until the point at which the penetration fronts (i.e., inwardly diffusing water/ion fronts) meet. [55]Figure 7 shows an example of the time dependent change in normalized conductance, G t /G ∞ , obtained for the hydration of previously dehydrated hydrogel covalently attached to the IME biochip substrate.Employing the MATLAB® Curve Fitting Toolbox, the two-term power equation for normalized conductance was generated up to 60% with R 2 = 0.99 as G t / G ∞ = 0.311t 1/2 − 0.022t, where G t and G ∞ represent the conductance at time t and final conductance at equilibrium, respectively (Figure 7).In this model fitting, the magnitude of the diffusion-dependent rate coefficient, k 1 = 0.311 s −0.5 , when compared to the relaxation-dependent coefficient, k 2 = −0.022s −1 , reveals the rapid diffusion of water molecules and ions from DMEM such as Na + , OCH 3 − , K + , Cl − , etc., into the hydrogel.Following the initial diffusion-dependent portion, relaxation of segments of the cross-linked polymer network became dominant and was described by the negative k 2 value, which initiated the decrease in conductance slope, leading to final equilibrium conductance and eventual equilibrium hydration.Diffusion and relaxation processes were affected by two key physical phenomena.First, during hydrogel swelling, osmotic pressure increases within the membrane from entrance of water and ions from solution.This is necessary for maintaining charge neutrality via Donnan partitioning, which is driven by the extent of hydrogen ion disassociation from the p(HEMA-co-PEGMA-co-HMMA) polymer network (pKa = ≈7.7)when submerged in DMEM (pH = 8.2 at 37°C). [94,95,4,55]Second, during membrane hydration, water exists in three distinct states in relation to the polymer network, namely strongly-bound water (SBW), weakly-bound water (WBW) and non-bound, or free, water. [96]Upon exposure of the initially dehydrated p(HEMA-co-PEGMA-co-HMMA) membrane to the aqueous fluid, SBW engages in hydrogen bonding with the oxygen atoms found in abundance in polymeric side chains, creating a densely packed environment.SBW molecules encaged around polymer side chains increase the time needed for restructuring and reorganizing of polymer side chains during hydration of the membrane.This prolongs the total network relaxation time during which maximum WBW and free water enter the polymer network, [96] as determined by k 2 .
With this series of equations, it can be said that the conductance trend of p(HEMA-co-PEGMA-co-HMMA) hydrogels directly correlates to the swelling kinetics of the hydrogel.The ability to relate the conductance of hydrogel membranes with hydrogel swellability provides means of understanding a hydrogel's capacity for change in different solutions, as indicated by the conductance time constant, as well as determining the full period needed for equilibrium hydration of various hydrogel compositions.Swelling kinetics parameters which describe water and ion diffusion within the hydrogel, are crucial considerations in multiple applications including biomolecule transport in drug delivery [17] and ion transport in bio-responsive hydrogels. [97]Furthermore, using impedimetry for understanding swelling kinetics could be particularly useful for biomedical applications, [98,99] such as to correlate changes in membrane resistance to fluctuations of in vivo pH in real-time (acidosis) [10,100] or the conductance response of an enzyme-linked biosensor. [64]

Conclusions
A surface chemistry protocol for modification of glass and microfabricated gold-on-glass IAME-co-IME biochips was evaluated and proved to significantly increase adhesion of hydrogel membranes to substrates suitable for biomedical devices and biosensors.Increased adhesion resulted from the progressive modification of glass substrates with hydroxyl moieties, ϒ-APS, and eventual functionalization with APNHS, whereby the -acryloyl end groups conferred an adhesion advantage between the glass substrate and the hydrogel membrane that was applied.Sufficient adhesion enabled EIS and impedimetry analysis to be conducted in cell culture media, which revealed the swelling kinetics of p(HEMA-co-PEGMA-co-HMMA) hydrogels to be in agreement with known hydrogel swelling trends following a two-term power model for anomalous swelling.More specifically, impedimetry revealed that ≈30 min were required for full equilibrium hydration of a dehydrated hydrogel corresponding to the achievement of steady state conductance and an overall time constant of 9.22 min.The association of pH, hydrogel swellability, conductance, and time could be useful in electro-stimulated drug delivery and biosensing applications.

Figure 1 .
Figure 1.Representation of cleaning, surface modification and functionalization process for hydrogel attachment to glass slides.The process was as follows: cleaning of slides by sonication, UV ozone, piranha, and plasma treatment a); physical adsorption of ϒ-APS molecules b); covalent bonding of ϒ-APS molecules following oven curing c); elevation of amine groups by immersion in DI d); covalent attachment of APNHS to amines e); finally, casting, photo-crosslinking, and hydration of hydrogel on glass slide f).

Figure 2 .
Figure 2. Representation of surface modification and hydrogel attachment for IAME-co-IME biochips.Surface cleaning by solution washes and plasma treatment a); physical adsorption of ϒ-APS molecules b); cathodic cleaning to remove adsorbed silanols c); thermal curing of ϒ-APS to promote covalent bonding to chip d); elevation of amine groups with DI water e); reaction with APNHS f); casting, crosslinking and hydration of hydrogel illustrating covalent incorporation of the -acryloyl group into the crosslinked hydrogel g).
0 μm for IAME-co-IME 2-1 Au biochips given the spacing of interdigitations.Bode plots (frequency vs |Z|, and ) and Nyquist plots (Z real vs Z img ) were obtained directly from the software provided with the instrument (VersaStudio version 1.51, AMETEK, Inc., Oak Ridge, TN).Further data analysis using ZSimpWin software version 3.60 (AMETEK, Inc., Oak Ridge, TN) employed a modified Randles equivalent circuit model of R M (Q G R G )(Q DL R CT ) to resolve the membrane resistance, R M , electrode geometry resistance, R G , electrode geometry capacitance, Q G , charge transfer resistance, R CT , and the double layer capacitance, Q DL , also termed as the constant phase element (CPE).Constant phase elements are considered imperfect capacitors.For simplicity of representation, Q G and Q DL are written as C G and C DL , respectively.Breakpoint frequencies ( = 1/R CT C DL ) of dehydrated and hydrated hydrogel membranes were extracted from Nyquist plots and averaged to define a characteristic dispersion frequency (CDF).Impedimetry of dehydrated hydrogels was performed at the CDF for 1 h in DMEM to observe impedance changes within a hydrogel during the process of hydration.DMEM was chosen to mimic physiological conditions.Equation 5 was implemented to find the total impedance of the R M

Figure 3 .
Figure 3. Representation of surface modification and covalent conjugation of NHS fluorescein to the IAME-co-IME biochips.Reaction of physically adsorbed ϒ-APS molecules with NHS fluorescein to produce the amide derivative a); cathodic cleaning to remove adsorbed silanol derivatives b).Scale bars are 1 μm.

Figure 7 .
Figure 7. Impedimetry of a p(HEMA-co-PEGMA-co-HMMA) hydrogel membrane was conducted for 1 h at the interrogation frequency of 0.22 Hz during hydrogel swelling in DMEM.Membrane resistance was extracted from impedance data and converted to conductance, which revealed the time constant () corresponding to 63.2% of maximum conductive, and therefore swelling, behavior to be 9.3 min.Maximum hydrogel conductance was achieved at 32.5 min, at which time steady ion flow and thus equilibrium hydration was achieved.

Table 1 .
Equivalent circuit parameters for R M (Q G R G )(Q DL R CT ) circuit extracted from ZSimpWin 3.60 software for both dehydrated and equilibrium hydrated samples measured in air and DMEM, respectively.
2.8.Statistical AnalysisAll data are reported as the mean ± SD of replicates as indicated.Contact angles are the result of n = 12 replicates.Attachment studies based on aspirate and dispensed drops were 24, 48, or 72 features for each level of surface preparation with each performed in triplicate (n = 3).Hydrogel adhesion studies via centrifugal method were performed as replicates (n = 38) for each level of surface preparation and were statistically evaluated using 2-sided t-tests.Where reported, significance was determined as a p-value calculated in JMP 16.2 software wherein ** and * denote groups as significantly different for p<0.01 and p<0.05, respectively.