Introduction

Adjusting to modern healthcare, biosensors currently dominate the medical field by enabling widespread applications, including accurate diagnoses of infections, pharmaceutical development, environmental monitoring, and food safety control. They also seamlessly integrate with emerging technologies such as microfluidics, nanotechnology, and electrochemical and optical sensors in various sensing platforms. While biosensors consistently rely on the same principle of detecting biological interactions, their signal output and applications vary greatly due to differences in transducer technologies1. Among the various types of biosensors, optical biosensors are valued for their remarkable sensitivity and selectivity, contributing significantly to advancements in biotechnology. Surface plasmon resonance (SPR) serves as a powerful optical detection technique for this purpose, especially for real-time bio-interaction analysis2,3. The foundational principle of this method depends on the oscillations of free electrons at the interface between a metal and a dielectric material, known as surface plasmons, which are facilitated by light interaction4,5,6,7.

For example, SPR-based biosensors are highly effective in detecting various disease biomarkers. One such critical biomarker is creatinine. Creatinine is a waste product filtered from the blood by the kidneys, and its concentration serves as a key indicator of kidney function. According to reported data, the physiological level of creatinine in normal blood ranges from 0.04 to 0.15 mM, and an increase in these blood serum levels can also be related to thyroid malfunction and muscular disorders8. While creatinine is also present in urine at significantly higher concentrations, this study specifically targets the development of a high-sensitivity SPR platform optimized for the lower concentration range found in human blood (0.05 to 0.6 mM). By focusing on this specific clinical window, the Au-ppHA sensor facilitates the high-resolution detection necessary for early clinical intervention and renal monitoring.

To achieve the required sensitivity for such clinical applications, this study utilizes a Kretschmann-based SPR (K-SPR) configuration. This setup involves the excitation of surface plasmons when p-polarized light, striking through a prism under total internal reflection, interacts with a thin metal film (e.g., gold) at a specific angle, known as the resonance angle9,10,11. This phenomenon causes a significant reduction in reflected light intensity. Because the resonance angle is highly sensitive to changes in the refractive index near the metal surface, it enables real-time, label-free detection of molecular binding events. The performance of this system, however, requires careful consideration of many factors that determine the sensitivity of the SPR sensor, as the interplay of surface chemistry, metal film thickness, refractive index contrast, and the physical properties of the transducing interface all play a role12. A key parameter that directly impacts the sensitivity and specificity of the sensor is the surface’s wettability13.

Surface wettability describes how a liquid behaves on a solid surface, specifically its tendency to spread across or adhere to that surface14,15. This characteristic is essential for optimizing biomolecule immobilization, controlling analyte distribution, and minimizing non-specific adsorption. The ability to precisely modify surface wettability is critical for sensor design16 and is often measured via the water contact angle (WCA)17. A lower WCA (< 90°) indicates a higher surface wettability, which is typically desirable for providing a better platform for biomolecular attachment and minimizing unintended binding events. In 2020, Purohit, Vernekar, Shetti and Chandra18 described how surface engineering remarkably contributes to biosensor development by providing a consistent and stable sensing surface, enhancing analyte-biorecognition element interaction, and minimizing fouling effects in biological solutions.

Plasma polymerization is an effective technique to tailor the surface’s properties and achieve desirable biological compatibility131920. This process involves the plasma-assisted activation of monomeric gases, followed by the deposition of a thin, highly cross-linked polymer film on the target substrate21. Specifically, plasma polymerization of heptylamine produces a coating with nitrogen-containing functional groups (e.g., amine) that are stable in air and aqueous environments. The deposition parameters, such as plasma power, allow for the precise control of surface wettability and a lower WCA22,23,24. The introduction of these amine-rich groups enhances the immobilization of biorecognition elements (BRE) by creating a chemically engineered surface.

Although the role of surface chemistry in biosensing has been widely proven, a thorough understanding of how plasma deposition parameters directly influence the resulting WCA and, consequently, the SPR sensitivity remains largely unexplored. Although plasma-polymerized amine-rich coatings are established in biosensing, this work provides a new mechanistic insight by quantitatively correlating plasma power (10 W to 80 W) with surface wettability (WCA) and subsequent K-SPR sensitivity (60°/RIU at 785 nm). This systematic optimization establishes a precise fabrication protocol to maximize biomolecule immobilization, which has remained largely unexplored in previous amine-modified SPR studies.

Furthermore, while the influence of plasma power on surface wettability has been explored in various material contexts, its specific application in optimizing the bio-interface for high-sensitivity creatinine detection remains largely unaddressed. This study fills this gap by establishing a quantitative link between plasma-polymerized heptylamine (ppHA) deposition parameters and the resulting enzyme-immobilization capacity. The novelty of this approach lies in the strategic optimization of the amine-rich interface to facilitate high-density covalent immobilization, resulting in a sensitivity enhancement of two orders of magnitude compared to traditional SPR creatinine sensors.

Specifically, this study aims to investigate the relationship between the ionization and deposition rate of plasma power, which directly impacts the water contact angle of heptylamine-coated gold surfaces. Furthermore, we characterized the gold plasma polymerized heptylamine sensors using chemical and physical techniques, including AFM, to confirm improved surface properties. The SPR analysis, which involves the angular sensitivity of the SPR biosensor using a Kretschmann-based configuration, is used to validate the results.

Experimental

Materials

BioNavis Ltd. supplied the sensor slides, which were coated with a 50 nm thick gold film. All chemical reagents, including N-heptylamine, ammonia (25% v/v), and hydrogen peroxide (25% v/v), were obtained from Sigma Aldrich and used as supplied. For the buffer, phosphate buffered saline (PBS) from Sigma-Aldrich was used, along with ultrapure water from a Millipore system, which had a resistivity of 18.2 MΩ.cm at room temperature (25 °C).

Preparation of sensors

K-SPR gold sensor slide was washed with distilled water in an ultrasonic cleaner, exposed to isopropanol for 10 min at temperatures below 30 °C, and then air-dried. To prevent surface contamination, Teflon-coated tweezers were used, and the sensors were handled only at their edges throughout the entire preparation process. Any remaining residue on the gold sensor was removed by heating it in a solution of distilled water, ammonia, and hydrogen peroxide (5:1:1) for 10 min at 80–90 °C. To prevent dust buildup, the gold (Au) sensor was rinsed with distilled water and then dried with nitrogen gas. To ensure the elimination of organic impurities, surface cleanliness was visually inspected and then confirmed by inserting the sensors into the SPR BioNavis equipment and running a baseline measurement with air.

Plasma polymerization process

A 13.56 MHz radio frequency (RF) generator connected to a matching network was used to deposit plasma polymer films inside a specially constructed cylindrical glass reactor, as shown in Fig. 1a. The K-SPR gold sensor slide as shown in Fig. 1b was placed gold-side up on a grounded stainless-steel platform, acting as the lower electrode.

Fig. 1
Fig. 1
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Schematic of the heptylamine plasma polymerization (ppHA) process: (a) illustrates the experimental setup and the chemical mechanism of ppHA deposition, (b) shows a gold sensor partially masked with a silicon wafer and (c) the final gold slide after the deposition process.

This platform was positioned 12.5 cm beneath a U-shaped powered electrode mounted at the top of the chamber. The upper electrode was connected to both the RF generator and a pressure gauge to monitor the internal environment. Two inlet ports enabled the simultaneous injection of vaporized heptylamine monomer and nitrogen as the carrier gas. The heptylamine monomer was delivered in vapor form from a sealed flask via a valve. On the lower side of the system, the grounded electrode was connected in series with a cold trap and vacuum pump. This setup allowed efficient evacuation of residual gases and unreacted monomer byproducts while minimizing contamination25.

Before monomer vapour was introduced into the chamber, the reactor was evacuated (< 0.01 torr) to create a steady starting pressure of 120 millitorr. Within five minutes, contaminants were removed from the sample and the chamber by air plasma26. Four millilitres of heptylamine were supplied at 25 Pa to facilitate electron release and monomer fragmentation once the chamber reached a base pressure of 6.6 Pa27. During this process, the injected heptylamine vapor underwent fragmentation due to energetic collisions within the plasma field. This led to the formation of reactive species (NH₂•, CH₂•, H•) and recombination events that formed amine-functionalized plasma polymer coatings. These coatings incorporated primary amine groups (–NH₂, –NH–R) suitable for subsequent biomolecule immobilization, while byproducts such as H₂, CH₄, and N₂ were continuously vented through the exhaust.

Unlike conventional plasma systems that rely on externally controlled gas flow rates, the precursor delivery in this configuration was governed by chamber pressure and needle valve regulation. Therefore, the deposition environment was controlled through pressure stabilization rather than direct gas flow variation, ensuring consistent monomer fragmentation and film formation. Two needle valves (Swagelok) were used to regulate the gas pressure, which was tracked using an absolute capacitance Baratron gauge (Model 627, MKS). Eight plasma powers, ranging from 10 W to 80 W, and a 60-second exposure duration were used to observe the variation in thickness and water contact angle. The treatment duration was fixed based on previously reported optimization studies demonstrating stable and reproducible plasma polymer films; consequently28, further variation of exposure time was beyond the scope of the present study, which focuses primarily on the influence of plasma power. The example of gold slide undergoing the plasma polymerization is shown in Fig. 1 (c).

Measurement techniques

The surface properties and performance of the fabricated biosensors were characterized using a suite of analytical techniques. Surface wettability was quantified using water contact angle (WCA) measurements via the sessile drop method. Atomic Force Microscopy (AFM) was employed to analyze the surface morphology and roughness at the nanoscale. Following surface characterization, the functional response of the sensors was evaluated using a Kretschmann-based Surface Plasmon Resonance to monitor changes in the resonance angle and determine key performance metrics such as refractive index sensitivity. Finally, enzyme immobilization was performed via drop-casting of creatininase onto the Au–ppHA surface, and the sensor’s performance was assessed by detecting varying concentrations of creatinine.

Water contact angle measurement

WCA measurements provide a metric for evaluating changes in hydrophilicity resulting from plasma treatments. The experiment employed variations in power, and the deposition time of 60 s was selected as it is brief yet conducive to cell adhesion and growth. This study employed the sessile drop technique29 to assess the water contact angle from two different sides with three repetitions and examined using ImageJ software (1.54 g). The sessile drop method effectively measures surface tension in phase-separated protein and polymer droplets, enabling physical characterization of these systems with limited sample volumes30. The sample deposited was put on an electrode with a white background. A 3 µL of distilled water was dropped at the side region of the gold. Images of the drops were captured immediately with an adjacent camera as shown in Fig. 2 and inserted into ImageJ software.

Fig. 2
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The water contact angle measurement on Kretschmann-based surface plasmon resonance (K-SPR) gold sensor slide.

The surface wettability was characterized by measuring the water contact angle using ImageJ software. The angle range of less than 90° indicates hydrophilic behavior, while contact angles above 90° are categorized as hydrophobic31. In SPR biosensors, enhanced hydrophilicity is particularly important at the gold–analyte interface, as it promotes improved wettability and facilitates efficient biomolecular immobilization onto the sensing surface. Stronger adsorption enhances refractive index modulation near the metal layer, thereby improving sensor sensitivity. ImageJ software was utilized to calculate the contact angle and associated uncertainty, and any image with an uncertainty value exceeding 0.2 was discarded and reanalyzed to maintain data reliability.

Atomic force microscopy (AFM)

Atomic Force Microscopy (AFM) (Park Systems, NX-10) is used to analyze the topography and determine a material’s properties at the nanoscale. This technique relies on a probing tip at the end of a cantilever to probe the surface of the substrate, which is mounted onto a piezoelectric scanner. As the cantilever probes the surface with a resonant frequency, deflections are detected via a laser beam. An incident laser beam is reflected off the cantilever and detected by a photodiode, then the light signal is converted to an electrical signal. AFM can visualize surface topography and evaluate thin film roughness using the Root Mean Square (RMS or Rq) roughness. This value considers significant changes between peaks and valleys, unlike the average roughness Ra, which is commonly employed to assess surface roughness.

SPR measurement

The SPR curves were analysed using the BioNavis Layer Solver software, which enabled identification of the refractive index (n) and thickness (d) of the ppHA layer32,33. These values were obtained through a curve-fitting process, where the software iteratively adjusts the optical parameters of the simulated multilayer model to match the experimentally observed reflectivity curve. By minimizing the difference between the simulated and experimental resonance curves, the best-fit values for the ppHA layer’s thickness and refractive index are extracted. The refractive index was expressed as a function of wavelength using Eq. 1,

$$n = a + b\lambda$$
(1)

where a and b are material-specific fitting constants and λ is the wavelength.

To correlate the change in refractive index with the amount of biomaterial adsorbed on the sensor surface, the refractive index increment was used, defined as Eq. 2,

$$\Delta n = \left( {\frac{{dC}}{{dn}}} \right)\Delta C$$
(2)

where \(\:\varDelta\:n\) is the change in refractive index due to the adsorbed biomolecular layer and is the surface concentration (typically in mg/mL). The refractive index and refractive index increment (\(\:\raisebox{1ex}{$dC$}\!\left/\:\!\raisebox{-1ex}{$dn$}\right.\)) are both used to calculate the surface mass density (Γ), which describes the amount of material bound to the sensor surface. For protein-based layers, (\(\:\raisebox{1ex}{$dC$}\!\left/\:\!\raisebox{-1ex}{$dn$}\right.\)) can be calculated theoretically if the amino acid sequence is known. However, in the absence of such detailed molecular information, the optical constants of most proteins are relatively similar, allowing the use of established generic values at each wavelength (typically (​\(\:\raisebox{1ex}{$dC$}\!\left/\:\!\raisebox{-1ex}{$dn$}\right.\)) ≈ 0.185mL/g at 670 nm).

SPR analysis workflow

After the water contact angle measurement, the distilled water droplet was gently removed without damaging the ppHA deposition on the gold sensor, and the sensor slide was immediately inserted into the BioNavis SPR instrument. The SPR Navi 200 instrument employed Kretschmann configuration where the incident light is coupled via prism so that the wave number of the incident light is enhanced to match that of the excited surface plasmon.

Fig. 3
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SPR Navi 200 instrument schematic with Kretschmann configuration measurement setup.

The optical interface between the prism and glass side of the sensor slide contains refractive index matching elastomer to match the refractive index of the prism and the sensor glass slide. Without the use of the refractive index matching elastomer, reflection and refraction might occur at the optical interface and thus the SPR excitation at the gold surface will be affected. As shown in Fig. 3, upon exposure to p-polarized light at 670 and 785 nm optical wavelengths, SPR can be excited at the sensor region. As the analyte becomes attached to the surface, the refractive index of the region changes. In this work, 785 nm will be used for the detection.

The sensing region is illustrated by the dielectric layer above the gold film. The changes in refractive index are shown in SPR response curve as shifts in resonance angles. This instrument enables real-time observation of biomolecular interactions, structural changes, and layer properties in both wet and dry states without a labelling process. Prior to sample measurements, a baseline calibration was conducted in air. The SPR angular scans were then performed sequentially in air, distilled water, and phosphate-buffered saline (PBS) to compare the refractive responses of the functionalized surface with the bare gold substrate. The scans were performed within the angular range of 35° to 80° under static conditions.

Enzyme immobilization of creatininase on Au–ppHA via drop casting method

In this part, the optimized Au-ppHA sensor slide is used. A 3 µL volume of glutaraldehyde (GA, 2.5% v/v in PBS, pH 7.4) was drop-cast onto the sensory region of the Au–ppHA surface and incubated at ambient temperature for 1 h to facilitate aldehyde activation34. Drop-casting was selected as a simple immobilization strategy, as previous studies have reported that dried sensing films formed via this method remain effective for creatinine detection in buffer systems, demonstrating the feasibility of this approach despite limited control over surface morphology35.

Following incubation, the surface was gently rinsed three times with 3 µL PBS to remove excess GA, and subsequently air-dried for 15–20 min. Creatininase enzymes (for creatinine detection) were prepared in PBS and 3 µL of the enzyme solution was drop-cast directly onto the GA-activated area. Samples were incubated under ambient conditions for 2 h to facilitate covalent attachment36, after which the modified region was rinsed again three times with 3 µL PBS to remove unbound enzyme and air-dried for 15–20 min. The prepared Au–ppHA enzyme substrates were either immediately used for sensing measurements or stored hydrated at 2–8 °C until further use37.

Detection of different creatinine concentrations

Non-enzymatic creatinine standard solution was prepared from analytical-grade powders (Sigma-Aldrich) dissolved in PBS (pH 7.4) to obtain stock solutions, from which fresh working dilutions at defined molarities were made. The non-enzymatic creatinine standards were prepared in the range of 0.05 to 0.6 mM38,39. The enzyme-functionalized Au–ppHA substrates were mounted into the SPR flow cell, and baseline stabilization was achieved with PBS buffer at room temperature. At each concentration step, 0.2 mL of the analyte solution was injected into the flow cell to cover the enzyme-modified sensing region, and the surface plasmon resonance angle shift (Δθ) was monitored in real time. After each injection, the flow cell was flushed with PBS to remove residual analyte and re-establish baseline conditions before the next concentration was tested, following established rinsing protocols for SPR sensing. The angular shifts obtained for different creatinine concentrations were plotted to construct calibration curves of Δθ against analyte concentration to evaluate sensor sensitivity, linearity, and detection range40.

For comparative purposes, creatinine detection was also performed on an unfunctionalized bare gold surface using a solution-phase enzymatic approach. Creatinine standards were prepared from analytical-grade powder, while the creatininase enzyme was dissolved separately in PBS. The enzyme was then mixed directly with the creatinine solution prior to measurement, and the mixture was introduced onto the bare Au surface for SPR analysis. This method complements the drop-casting immobilization strategy used on the ppHA-functionalized surface, allowing for a direct comparison of the trade-off between biochemical functionality and optical sensitivity.

Result and discussion

Influence of plasma power on surface wettability

Table 1 Water contact angle (WCA) measurements of gold sensor surfaces at varying plasma power levels. Values are presented as mean ± standard deviation (n = 6).
Full size table

The WCA images and corresponding mean theta E values, as shown in Table 1, illustrate the effect of varying plasma power (10–80 W) on the wettability of gold sensor surfaces, using a fixed deposition time of 60 s. These measurements, obtained via the sessile drop method and analyzed with ImageJ software, were repeated six times (n = 6), and the reported values are presented as mean ± standard deviation. The results reveal how surface treatment conditions influence hydrophilicity. The chemical functionality of the ppHA coating, specifically the presence of primary amine groups, was previously established via FTIR analysis in a related study detecting glucose28. In this study, focus is directed toward the correlation between plasma power, surface wettability, and the resulting SPR response. At lower power levels (10–20 W), the surfaces exhibit moderately hydrophilic behavior, with WCA values of 64.57° and 55.67°, respectively. This hydrophilic state is essential for biosensing, as lower WCA values (below 60°) suggest increased surface energy that promotes better molecular spreading and stronger biomolecular adhesion. While enhanced hydrophilicity is expected to improve biomolecule attachment, the direct quantitative relationship between surface wettability and creatinine sensing sensitivity was not examined in this study and represents an important avenue for future investigation.

As the plasma power increases, the WCA gradually rises, reaching 80.4° at 60 W, which indicates a transition toward more hydrophobic surface characteristics. This trend is likely driven by enhanced ion bombardment at higher plasma energies, promoting surface densification, and reducing the availability of polar functional groups. Plasma-induced surface modifications have been widely documented to influence wettability through the formation of hydrophilic functional groups and nanoscale surface restructuring, both of which contribute to significant changes in surface energy and liquid–surface interactions41. Further increases to 70 W and 80 W result in a slight decrease in WCA to 79.6° and 77.55°, respectively, which may be associated with competing plasma effects such as mild etching or surface restructuring that alter surface energy. Such variations are consistent with prior findings that prolonged plasma exposure or higher power can deteriorate the surface and modify wettability due to morphological and chemical changes42. Although these studies were conducted on polymeric substrates, the fundamental mechanisms governing plasma-induced surface modification are widely recognized to be broadly applicable across material systems.

Although these studies were performed on wood substrates, the underlying plasma-induced surface energy modifications and competing surface restructuring phenomena are relevant and consistent with trends observed in gold surfaces. From a biosensing perspective, excessive hydrophobic surfaces are less favorable for biomolecular interactions. Therefore, despite the higher WCA observed at elevated powers, the optimal treatment range is identified between 10 and 40 W, where sufficient surface activation is achieved while maintaining the hydrophilic characteristics necessary for effective biomolecule adhesion.

Controlling surface wettability via plasma power

The plasma deposition rates varied with respect to power, time, and monomer used. Figure 4 (a) depicts the WCA of five measurements from 10 W to 80 W deposition power. Increasing the power from 10 to 80 W led to the formation of thicker plasma polymers with higher refractive indices. The graph indicates the WCA rose from around 58° at 10 W to 82° at 70 W, demonstrating a transition from hydrophilic to hydrophobic surface properties as plasma power increased. As established in Sect.  3.1, this lower WCA range (below 60°) is essential for promoting molecular spreading and maximizing the interaction between the ppHA layer and subsequent biomolecular probes.

However, the WCA either plateaued or displayed a modest decline after 70 W. This is likely attributed to excessive fragmentation of monomers at high plasma powers, leading to chemically altered, crosslinked films with reduced surface polarity. Alternatively, increased surface roughness at higher deposition powers may trap air pockets, mimicking hydrophobic behavior (Cassie–Baxter state)43,44. The linear trend (R2 = 0.7584) shown in Fig. 4 (b) confirms a strong positive correlation between plasma power and WCA. The fitted trendline y = 3.1811x + 55.969 suggests a ~ 3.18° increase in WCA for every 10 W increase in power, allowing precise control of surface wettability. These results support the selection of the lower plasma powers (10–30 W) for biosensor fabrication, as this range produces surfaces with high wettability and molecular compatibility required for a predictable SPR response.

Fig. 4
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(a) WCA across plasma deposition powers (10–80 W) with six replicates and (b) the correlation between deposition power and water contact angle for Au-ppHA coated surfaces.

Thickness identification by layer solver

Fig. 5
Fig. 5
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Extracted layer thicknesses at optical wavelength of 785 nm (a) 10–80 W and (b) 10–30 W.

Figure 5 (a) presents the extracted layer thicknesses at a wavelength of 785 nm, determined via Layer Solver analysis under varying plasma powers from 10 W to 80 W. The data reveals an overall increase in thickness with increasing power, confirming the successful deposition and material accumulation on the substrate. The Layer Solver optimization, which matches theoretical reflectance models to experimental values, enabled precise estimation of optical thickness and provided insight into the deposition behavior across power variations. At lower powers (10–30 W), a progressive increase in thickness was observed, from 11.06 nm to 17.37 nm, indicating that the plasma energy was sufficient to initiate surface activation and moderate film growth.

A significant jump occurred at 50 W, reaching 27.44 nm, suggesting enhanced energy transfer and optimized precursor fragmentation, which facilitated more effective film formation. Interestingly, between 60 W and 70 W, the thickness plateaued around 19.5 nm, followed by a slight drop at 80 W (17.33 nm), before surging dramatically to 44.57 nm, possibly due to changes in plasma dynamics, redeposition effects, or non-linear growth kinetics at higher energies. Based on these results, a plasma power of 30 W was selected for further analysis due to its consistent and predictable film growth, which is critical for reproducible biosensor fabrication.

The linear trendline applied to the dataset yields a slope of 3.155, with a coefficient of determination (R²) of 0.9963 for 10 W to 30 W is shown in Fig. 5 (b). This suggests a moderate correlation between power and thickness, but also indicates the presence of nonlinear growth behavior, likely influenced by interfacial saturation, re-sputtering, or thermally induced surface roughening at higher plasma powers. These findings support the efficacy of the Layer Solver in quantifying deposition-induced thickness changes and indirectly capturing variations in refractive index due to increasing optical density. Since the refractive index and thickness are coupled in optical models, thicker regions likely correspond to higher effective refractive indices, particularly where denser or more uniform material growth is achieved. The significant increase at 80 W may reflect a densified or multi-phase layer, requiring further spectroscopic verification.

Optimization of 30 W using surface plasmon resonance

The importance of the 40°–60° WCA range lies in its strong correlation with enhanced biomolecule immobilization and SPR signal sensitivity. Surfaces within this hydrophilic range possess a higher surface energy, which facilitates stronger adsorption of polar. This is particularly advantageous for SPR sensing, as successful analyte binding on the sensor surface directly influences the magnitude and sharpness of the SPR angle shift. To further examine this, Fig. 6 highlights the WCA within this optimal zone. As the WCA approaches 40°, the surface energy increases, promoting better molecular adsorption and stronger interactions with aqueous biological media. Conversely, approaching 60° indicates a reduction in surface hydrophilicity, which may hinder biomolecule attachment and reduce SPR sensitivity.

This focused analysis confirms that controlling the plasma deposition power within this range is crucial to achieving a surface condition that supports optimal biofunctionalization and consistent sensor performance. It should be noted that the ppHA layer introduces an additional dielectric spacing between the gold substrate and the analyte, slightly reducing the evanescent plasmon field penetration into the surrounding medium. While this results in a minor decrease in refractive index sensitivity compared to bare gold, the optimized 30 W deposition produces a thin enough layer that preserves sufficient plasmon–analyte interaction while enhancing enzyme immobilization and surface bioactivity. In contrast, surfaces that become too hydrophobic (WCA > 70°) can hinder biomolecule attachment due to reduced surface wettability and poor compatibility with aqueous biological samples. This not only limits functionalization efficiency but may also introduce variability in sensor response. Therefore, maintaining the WCA at approximately 60°, as observed at 30 W plasma deposition power is essential for optimizing the surface for consistent, high-sensitivity SPR detection. Thus, for detection in K-SPR, the sensor’s slide with WCA of 60° will be used.

Fig. 6
Fig. 6
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SPR response curve for different water contact angle distribution in the hydrophilic range (40°–60°) for 30 W plasma deposition power.

Surface topography (AFM analysis)

AFM evaluation of the heptylamine plasma polymerized Au surface revealed in Fig. 7 occasionally rough shape with apparent differences throughout the scanned regions. The retrieved roughness parameters revealed that the average roughness (Ra) ranged from ~ 5 to 12 nm, with corresponding RMS roughness (Rq) values of ~ 7 to 13 nm and maximum peak-to-valley heights (Rz) of ~ 20 to 30 nm. These values are larger than those generally found on bare evaporated Au films (Ra ~ 1–4 nm)45 indicating successful polymer deposition while introducing surface heterogeneity. The AFM topography showed darker continuous regions due to consistent ppHA film coating, whereas localized brighter patches indicated thicker polymer clumps or non-uniform deposition.

Compared to ideal polymer-functionalized Au surfaces utilized in plasmonic sensing, which typically exhibit Ra values of ~ 3–10 nm and Rq of ~ 4–12 nm 46,47 the Au-ppHA-modified surface falls within the acceptable range. However, it tends to have higher roughness in some locations. This low roughness is beneficial for expanding accessible surface area and improving amine functionality for subsequent biomolecule immobilization. Overall, the AFM results demonstrate that heptylamine plasma polymerization creates an amine-functional layer on gold with adequate topographical properties for biosensing applications48.

Fig. 7
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AFM analysis for (a) surface morphology of 30 W Au-ppHA and (b) surface profile with step height of Au-ppHA.

Surface plasmon resonance analysis

Figure 8 illustrates the alterations in SPR resonance angles and fluctuations in dip depth resulting from surface modification in air and PBS on the SPR Au sensor slide and functionalized ppHA at 785 nm optical wavelength. The SPR measurements confirmed these surface modifications. Table 2 provides a summary of the resonance angle shifts for both the bare Au and Au-ppHA when measured in air and PBS. The uncoated gold surface displayed a distinct SPR dip at 42.612° in air, which red-shifted to approximately 66.411° on PBS due to the increased refractive index. The Au-ppHA surfaces exhibited broadening of the SPR dip at 50.069°, and further red-shifting to approximately 70.278° on PBS, signifying successful surface modification for biomolecule attachment. The broadening of the dip is attributed to the increased optical thickness and roughness introduced by the ppHA layer.

The sensitivity of the sensor is calculated using this following Eq. 3 and expressed as °/RIU49,

$$Sensitivity = ~\frac{{Change~in~resonance~angle~\left( {\Delta \theta _{{SPR}} } \right)}}{{Change~in~refractive~index~of~the~surrounding~medium~\left( {\Delta RI} \right)}}$$
(3)

SPR validated the surface alterations, yielding a sensitivity of ~ 69°/RIU (bare Au) and 60°/RIU (Au-ppHA). The added heptylamine layer increases the interface’s optical thickness and refractive index, raising the resonance angle. While this causes a minor reduction in inherent refractive index sensitivity, the ppHA layer supports effective plasmon coupling and improves biomolecule immobilization.

This result highlights the optimal trade-off between biochemical functionality and optical sensitivity. Consistent with the optimization in Sect.  3.1 and 3.4, the ppHA introduces abundant primary amine groups (–NH2) onto the sensor surface. These functional groups, combined with the hydrophilic nature of the Au-ppHA layer (WCA 40°–60°), are essential for maximizing enzyme loading through bio-conjugation. Although the polymer coating induces a slight attenuation of the plasmonic field (~ 13% reduction in RI sensitivity), it provides a stable, high-density interface for biorecognition. Consequently, the significant gains in immobilization efficiency and binding capacity far outweigh the minor loss in optical sensitivity, ultimately supporting the high-performance detection of creatinine demonstrated in the following section.

Fig. 8
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SPR reflectivity curves of bare gold (Au) and ppHA-coated gold (Au-ppHA) sensors in air and PBS.

Table 2 Resonance angle shift of bare gold (Au) and ppHA-coated gold (Au-ppHA) tested in air and buffer solution, PBS.
Full size table

Detection towards different concentration of creatinine

Fig. 9
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SPR reflectivity curves of Au-ppHA and Au-ppHA-GA-creatininase tested in PBS.

Before conducting the creatinine detection, the success of the enzyme immobilization process was verified through SPR angular shift analysis. Figure 9 illustrates the SPR reflectivity curves for the Au-ppHA surface in PBS buffer before and after the immobilization of creatininase. Following the covalent bonding of the enzyme via glutaraldehyde crosslinking and a thorough rinsing process to remove loosely bound molecules, a distinct red shift of 5.4° in the resonance angle was observed. This permanent shift indicates a significant increase in the surface refractive index due to the stable attachment of the protein layer. This high level of surface coverage is attributed to the optimized hydrophilic nature of the ppHA coating, which facilitates maximum enzyme loading and contributes to the high sensitivity of the sensor.

The Au-ppHA biosensor, optimized with a polymerization power of 30 W and a water contact angle of 60°, was tested for its response to different concentrations of creatinine. The concentration was varied from 0.05 mM to 0.6 mM. As shown in Fig. 10 (a), the SPR curve exhibits a significant resonance angle shift as the concentration of creatinine increases. The linear relationship between the resonance angle shift and creatinine concentration is shown in Fig. 10 (b). The sensitivity of the sensor was determined to be 1.1918 °/mM. The sensitivity of the sensor was calculated by dividing the shift in the resonance angle (∆θSPR) by the change in creatinine concentration. As the concentration of creatinine increased, the SPR curve’s angle shifting became less pronounced, indicating a decrease in sensitivity at higher concentrations. This reduced response is likely attributed to the finite loading of immobilized creatininase on the sensor surface, where available enzymatic active sites progressively approach saturation. Once a significant fraction of these binding sites is occupied, the rate of enzyme–substrate interaction decreases, resulting in a diminished incremental SPR response. This observation suggests that enzyme immobilization density is a key factor governing the measurable dynamic range of the biosensor rather than an intrinsic limitation of the Au–ppHA platform.

Fig. 10
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(a) SPR response curve of Au-ppHA-GA-creatininase toward different creatinine concentrations (0.05 mM to 0.6 mM) at a 785 nm wavelength, and (b) the linear relationship between the resonance angle shift and creatinine concentration.

Fig. 11
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(a) SPR response curve of Au toward different creatinine-creatininase concentrations (1–100 mM) at a 785 nm wavelength, and (b) the linear relationship between the resonance angle shift and creatinine concentration.

For comparison, a control experiment was performed using an unfunctionalized bare gold surface (with dissolved creatinine and creatininase enzyme). As shown in Fig. 11, the bare Au sensor exhibited a wider detection range of 1–100 mM but significantly poorer analytical performance, with a sensitivity of only 0.0023°/mM and a minimum detectable concentration of 1 mM. In contrast, the ppHA-functionalized sensor exhibited markedly higher sensitivity with a lower detection limit of 0.05 mM. This substantial performance difference is attributed to improved interfacial sensing conditions, where molecular interactions occur closer to the evanescent field, enabling more efficient translation of binding events into measurable resonance shifts. These findings highlight the surface modification using ppHA to immobilize creatininase and enhance the functionality of creatinine detection from 0.0023°/mM to 1.1918 °/mM.

The performance of the ppHA-based sensor was objectively benchmarked against a previous study by Menon, et al. 50, which utilized the same Kretchmann SPR configuration and 785 nm angular interrogation. Under these identical optical conditions, the bare gold surface in the previous study yielded a sensitivity of 0.006°/mM. In contrast, the introduction of the optimized ppHA functional layer in this work achieved a sensitivity of 1.1918 °/mM, representing a significant enhancement in detection performance. The sensing measurements were performed in multiple trials to ensure experimental reliability, and a representative response is presented here for clarity, with replicate experiments exhibiting consistent response trends.

While the detection range in this work (0.05 to 0.6 mM) is narrower than the previous study (10 to 200 mM), the chosen range (0.05 to 0.6 mM) is highly suitable for blood-based creatinine sensing, as it effectively encompasses the normal and pathological concentration range of creatinine in blood (0.040 to 0.150 mM). Although a direct comparison of linearity (R2) was not possible as it was not reported in the previous study, the high sensitivity and clinical focus of the current ppHA interface represent a clear advancement. The primary driver for this performance gain is the ppHA layer’s ability to support effective plasmon coupling while providing a high-density environment for the immobilization of the creatininase enzyme, a functional capability absent in the bare gold baseline.

To further place this work in a modern context, the platform was compared with a recent SPR creatinine biosensor reported by Almawgani, et al. 39. That study reported a high sensitivity of 250 degree/RIU using a complex silver (Ag)-cerium oxide (CeO2)-graphene nanocomposite at 633 nm. However, it is important to note that while their numerical study explored a broad range (10 to 200 mM), our experimental platform is uniquely tailored for high-sensitivity detection at lower, physiologically relevant concentrations (< 0.6 mM). Furthermore, whereas the nanocomposite structure requires complex multilayer fabrication, our ppHA interface offers a streamlined, single-step plasma deposition process, providing a more practical and reproducible route for clinical diagnostic applications.

Conclusion

This study successfully demonstrated the optimization of surface properties for K-SPR biosensors using plasma polymerized heptylamine (ppHA) coatings. The findings of this work confirm that plasma deposition power is a critical parameter for controlling surface wettability, which in turn significantly impacts biosensor performance. This work found a clear, positive linear correlation (R2 = 0.7584) between plasma power and WCA, indicating that higher power levels result in a more hydrophobic surface. The optimal power range for achieving a hydrophilic surface, essential for efficient biomolecule immobilization, was identified to be between 10 W and 30 W, with a WCA between 40° and 60°. The ppHA coating successfully altered the optical properties of the gold substrate, as evidenced by a red shift of the SPR resonance angle. While the ppHA layer resulted in a slight reduction in sensitivity from ∼69°/RIU (bare gold) to 60°/RIU, this minor trade-off is outweighed by the improved surface functionality for biomolecule attachment. The introduction of amine-rich functional groups via ppHA enhances surface bioactivity and provides a more stable, consistent platform for sensing applications. In conclusion, this research confirms that ppHA coatings effectively optimize both surface wettability and sensitivity. By carefully tuning the plasma deposition parameters, particularly the power, it is possible to fabricate highly efficient SPR biosensors with enhanced biomolecule immobilization. This improved platform enabled a high sensitivity of 1.1918 °/mM, confirming its promising performance for the real-time and label-free detection of creatinine.

While the sensor demonstrated promising sensitivity toward creatinine detection, selectivity against common interfering biomolecules was not evaluated in the present study. Therefore, this work should be regarded as a proof-of-concept platform. Future work will also focus on extending the measurable detection range toward higher creatinine concentrations, such as those found in urine, by optimizing enzyme immobilization density and surface architecture to mitigate early saturation effects. Future studies incorporating interference analysis are necessary to further validate sensor specificity and support its potential for clinical applications. Future work could explore the long-term stability of these coatings in different biological media, incorporate interference studies to evaluate selectivity, and investigate the effect of other deposition parameters, such as gas flow rate and time, on the surface properties and biosensor performance. Additionally, the integration of real-time SPR kinetic analysis and advanced surface characterization, such as XPS or FTIR, will be pursued to further elucidate the molecular orientation and binding density of the immobilized enzymes.