Molecularly responsive hydrogel-based biosensor

Abstract

The present invention relates to sensor systems and in particular to sensor systems for sensing an analyte of interest comprising a sensor, a hydrogel composition contacting the metal-coated surface of the sensor and processor circuitry. The invention further relates to a method for sensing an analyte of interest, to use of a sensor system for sensing an analyte of interest as well as to a hydrogel composition for use in diagnostic sensing of an analyte of interest. Embodiments of the invention have been particularly developed as implantable biosensors allowing for molecular in vivo monitoring of analytes and will be described hereinafter with reference to this application.

Description

[0001] ULLRICH & NAUMANN

[0002] GEWERBLICHER RECHTSSCHUTZ • I NTELLECTUAL PROPERTY

[0003] 8272 / P / 1001 -EP Heidelberg, July 14, 2025 / , %

[0004] European Patent Application in the name of

[0005] BioMed X Germany GmbH Im Neuenheimer Feld 518 69120 Heidelberg regarding

[0006] “Molecularly responsive hydrogel-based biosensor"

[0007] Schneidmuehlstrasse 21 Office Mannheim: 69115 Heidelberg, Germany Franz-Volhard-Strasse 3 Telephone +49 6221 60430 68167 Mannheim, Germany Facsimile +49 6221 604360 Telephone +49 621 4892590

[0008] E-mail office@un-ip.com www.un-ip.com MOLECULARLY RESPONSIVE HYDROGEL-BASED BIOSENSOR

[0009] FIELD OF THE INVENTION

[0010] The present invention relates to sensor systems and in particular to sensor systems for sensing an analyte of interest comprising a sensor having a metal-coated surface, a hydrogel composition contacting the metal-coated surface of the sensor and processor circuitry. The invention further relates to a method for sensing an analyte of interest, to use of a sensor system for sensing an analyte of interest as well as to a hydrogel composition for use in diagnostic sensing of an analyte of interest. Embodiments of the invention have been particularly developed as implantable biosensors allowing for molecular in vivo monitoring of analytes and will be described hereinafter with reference to this application. However, it will be appreciated that the invention is not limited to this particular field of use.

[0011] BACKGROUND OF THE INVENTION

[0012] Any discussion of the background art throughout the specification should in no way be considered as an admission that such art is widely known or forms part of common general knowledge in the field.

[0013] Continuous, real-time, and quantitative monitoring of small molecules (such as drugs, metabolites, hormones, etc.) and of other biomarkers in vivo remains a significant challenge, despite conventional techniques such as immunoassays and liquid chromatography with mass spectrometry having currently been established as gold standard techniques for therapeutic drug monitoring (TDM). These methods lack realtime continuous monitoring and are limited by their inadequacy to characterize drugs that exhibit a complex relationship with serum or blood components. As such, advancements in techniques capable of measuring in vivo concentrations and dynamics of molecules with high spatio-temporal resolution, while individuals perform their daily activities, have the potential to revolutionize modem healthcare as they offer crucial insights into physiology, pharmacokinetics (PK), pharmacodynamics (PD), and drug toxicity.

[0014] Known technologies attempting to address this need include electrochemical aptamer-based (EAB) sensors. These sensors achieve specificity through aptamers that change their conformation upon analyte binding. However, such systems rely on metal electrode sensors that are subject to fouling when placed in a complex biological environment such as blood and / or tissues. This affects sensor functionality, accuracy, and causes drift, hence limiting the quantitative determination of concentrations.

[0015] Alternatively, various optical sensors can be applied for / 7? vivo sensing, including fiber optic surface plasmon resonance (FO-SPR) sensors. While the probes of such sensors are implantable much like electrodesthey equally suffer from biofouling when placed in complex biological environments. Furthermore, optical sensors are not self- contained and typically require complex detection devices. Some may have the propensity to disintegrate within the body making them unsuitable for continuous in vivo monitoring. As such, numerous sensor systems for sensing an analyte of interest are known.

[0016] The incorporation of hydrogels into such systems, in particular in the context of biosensing, is also known. Hydrogels are three-dimensional networks of hydrophilic polymers capable of holding large amounts of aqueous solutions. Due to their unique properties, hydrogels can be integrated into various types of sensors for analyte biosensing applications, including electrical, mechanical and optical sensors. Electrical biosensors, for example, rely on changes in the electrical properties of a hydrogel composition, such as conductivity or impedance, induced by analyte binding, whereas mechanical biosensors detect changes in the mechanical properties of a hydrogel, such as a change in stiffness or mass, resulting from analyte binding or hydrogel swelling. In optical biosensors, hydrogels are used as matrices for immobilizing biomolecules, such as enzymes or antibodies, and for transducing analyte binding events into measurable optical signals, such as changes in absorbance or fluorescence.

[0017] Known hydrogel compositions exhibit unique swelling properties in response to changes in environmental conditions, such as pH, temperature, or the presence of specific ions or molecules: properties, which have been utilised for signal transduction in biosensor systems. For instance, changes in the volume or mechanical properties of the hydrogel matrix upon analyte binding can be translated into measurable changes in mechanical properties of the hydrogel, enabling the quantitative detection of analytes leading to the measured change in mechanical property. Hydrogels can be prepared such that they selectively bind to specific analytes, for example to specific proteins, enzymes, antibodies or nucleic acids. This selective binding can allow for the detection of target analytes in complex aqueous solutions, including complex biological samples, such that hydrogels can be applied across a wide range of fields, including environmental monitoring, food safety, biotechnology and healthcare.

[0018] For example, hydrogel-based biosensors can detect environmental contaminants, such as heavy metals, pesticides or toxins, in water, soil or air samples with high sensitivity and selectivity. In the food industry, hydrogel-based biosensors can be employed to detect foodborne pathogens, allergens, or adulterants in food products. Not dissimilarly, in healthcare, hydrogel-based biosensors can be used to detect disease biomarkers and biotechnology applications of hydrogel-based biosensors include various bioprocessing applications, such as monitoring cell culture processes, enzymatic reactions, which may be used in real-time monitoring in the context of process optimization, quality control and / or process scale-up.

[0019] However, currently known hydrogel-based biosensors face several limitations and challenges, which have limited their widespread adoption and commercialization even non-critical applications. One significant challenge is the optimization of hydrogel properties, such as porosity, swelling kinetics and mechanical stability, to ensure optimal performance and longevity of biosensors.

[0020] In particular in the context of healthcare, the tuning of hydrogel-based sensor systems to achieve the necessary sensitivity and selectivity (i.e. accuracy), while ensuring biocompatibility and stability (i.e. durability) have been identified as key challenges in hydrogel-based biosensor design generally.

[0021] Specifically, the development of hydrogel-based biosensor systems for continuous in vivo detection of biomarkers of disease or for continuous monitoring of therapeutic agents in a target organ or tissue still requires addressing numerous issues in concert such as biocompatibility, stability and anti-fouling properties. It is accepted that non-specific interactions between hydrogel sensors and biomolecule analytes present in biological samples will lead to false-positive or falsenegative results, compromising the accuracy and reliability of biosensing assays. Importantly, it is also known that physical properties of known hydrogel biosensors, such as overall probe size as well as probe rheology (i.e. stiffness, surface roughness, viscoelasticity of the hydrogel composition, etc.) can (a) impact the integrity of the sensor probe itself, e.g. through delamination either due to shear forces experienced by the probe during its deposition of the probe or while implanted due to micro motion within the tissue, and (b) can lead to the formation of a fibrous capsule around the sensor or sensor probe in a foreign body response - all of which compromise or limit optimal performance of the sensor system.

[0022] In light of the above, it will be appreciated that there is a need in the art for improved hydrogel-based biosensor systems for sensing an analyte of interest, in particular for continuous sensing of such an analyte in vivo.

[0023] It is an object of the present invention to overcome or ameliorate at least one of the disadvantages of the prior art, or to provide a useful alternative. In particular, it is an object of the present invention to provide improved hydrogel-based biosensor systems for sensing an analyte of interest.

[0024] SUMMARY OF THE INVENTION

[0025] As indicated above, the present invention aims at providing hydrogel-based biosensor systems for sensing an analyte of interest with high specificity and sensitivity. The sensor systems of the present invention are particularly useful for continuous biosensing in vivo sensing, as they utilize a label-free transduction method, enable continuous data collection with minimal delays, and remain biocompatible for extended periods of time when placed inside the body.

[0026] Accordingly, in a first aspect, the present invention relates to a sensor system for sensing an analyte of interest, comprising:

[0027] - a sensor to provide sensor data, wherein the sensor comprises a metal- coated surface;

[0028] - a hydrogel composition contacting the metal-coated surface of the sensor, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest; and

[0029] - processor circuitry to determine information on a presence of the analyte of interest based on the sensor data, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel, and wherein the sensor data is indicative of the alteration of the spatial dimensions of the 3-D network structure.

[0030] As outlined above, each allosteric molecule undergoes a conformational change upon binding to the analyte of interest and assembly of the respective analyte binding complex. It is this conformational change of each individual allosteric molecule, which ultimately alters the spatial dimensions of the 3-D network structure of the hydrogel, thereby contributing to measurable effects, such as a surface plasmon resonance wavelength shift in case a surface plasmon resonance sensor is used, or a shift in resonance frequency or energy dissipation in case a quartz crystal-based sensor is used. As such, the specific hydrogel composition of the sensor system of the present invention is molecularly responsive to the analyte of interest and can act as a transducer, converting analyte binding events into measurable effects, such as surface plasmon resonance wavelength shifts.

[0031] Importantly, through assembly of respective analyte binding complexes, the effect measured (e.g., the surface plasmon resonance wavelength shifts or the change in resonance frequency and the change in energy dissipation) is indicative of the number of analyte molecules present in the aqueous environment surrounding the sensor probe. Furthermore, by way of its 3-D network structure, the hydrogel composition of the present invention is significantly more sensitive than a 2-D mono-layer surface plasmon resonance probe as the 3-D network structure, and in particular its alteration of its spatial dimensions in response to analyte binding, allows for the detection of significantly more analyte binding events than a 2-D mono-layer, including binding events distant from the surface. The sensor systems of the present invention may further be designed such that binding of the allosteric molecules or allosteric fragments thereof to the analyte of interest within the hydrogel composition is reversible, at least at body temperature (i.e. at about 37°C), such that the analyte binding complex disassembles upon separation of the analyte of interest from the allosteric molecules or allosteric fragments thereof and the conformational change is reversed. This reversibility of analyte binding makes the sensor system of the present invention suitable for continuous, real-time, in situ sensing, addressing a critical challenge when sensing analytes such as drug concentrations within complex biological solutions or tissues such as the brain.

[0032] As such, if the alteration of the spatial dimensions of the 3-D network structure of the hydrogel composition upon assembly of the analyte binding complex is a contraction, release of the analyte, i.e. disassembly of the analyte binding complex, leads to an expansion or relaxation of the hydrogel composition.

[0033] In a second aspect, the present invention relates to a method for sensing an analyte of interest, the method comprising the following steps:

[0034] (a) providing a sensor, such as a surface plasmon resonance sensor or quartz-based sensor, comprising a sensor probe with a metal-coated surface and a hydrogel composition contacting the metal-coated surface, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel;

[0035] (b) providing, by the sensor, sensor data being indicative of a of the alteration of the spatial dimensions of the 3-D network structure, such as a surface plasmon resonance wavelength shift, shift in resonance frequency of a quartz crystal or shift in dissipation of a quartz crystal; and

[0036] (c) determining information on a presence of the analyte of interest based on the sensor data. As indicated above, either, the method is an ex vivo method for sensing the analyte of interest in an aqueous sample, preferably the aqueous sample is a previously- obtained bodily fluid sample or a tissue sample or step (c) providing the sensor data by the sensor, comprises inserting a probe of the sensor into a bodily fluid or tissue for sensing the analyte of interest in situ.

[0037] In a third aspect, the present invention relates to use of a sensor system according to the first aspect for sensing an analyte of interest in an aqueous sample ex vivo.

[0038] Further, in a fourth aspect, the present invention relates to a hydrogel composition for use in diagnostic sensing of an analyte of interest, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest.

[0039] FIGURES

[0040] Embodiments of the invention will now be described, by way of example only, with reference to the accompanying drawings in which:

[0041] Fig. 1 shows a sensor probe according to one embodiment of the present disclosure.

[0042] Fig. 2 illustrates the components used in exemplary hydrogel compositions of the present disclosure. (A) schematically shows a full allosteric aptamer, the separated allosteric split aptamer fragments of the full aptamer of (A) as well as the re-assembled split aptamer pair substantially recreating the full aptamer of (A); (B) illustrates how an 8-arm functionalized hydrophilic star polymer together with a correspondingly functionalized cross linker can form the 3-D network structure of the hydrogel composition; and (C) illustrates the wavelength shift measurable with a sensor system in accordance with the present disclosure upon analyte binding and assembly of the analyte binding complex. In (C), the sensor data was obtained using a Multi-Parametric SPR (MP-SPR) sensor. However, the proposed concept can be implemented using different types of SPR sensors, such as planar SPR sensor, MP-SPR sensors or Fiber-Optic SPR (FO-SPR) sensors, with FO-SPR sensors being particularly suitable for the purpose of implantation.

[0043] Fig. 3 schematically illustrates the alteration of the spatial dimensions of the 3-D network structure of a hydrogel composition of a sensor system of the present disclosure (connected to the metal surface of the SPR) in response to analyte binding. (A) without analyte binding = expanded or relaxed conformation; (B) with analyte binding = collapsed or contracted conformation.

[0044] Fig. 4 schematically illustrates the basic concept of hydrogel-based sensing using multi-parametric SPR. (A) schematically illustrates the reflectivity of the metal film being a function of the angle 0 of the incident light, with some of the changes in reflectivity being based on the waveguide (i.e. , the thickness and refractive index of the waveguide), and the major change in reflectivity being based on the surface plasmon. (B) schematically illustrates the surface plasmon angle shifting (by an angular shift A0, e.g., influenced by the kinetics of the analyte being bound by the aptamers included in the hydrogel structure.

[0045] Fig. 5 shows a difference in the waveguide pattern of a Multi-Parametric SPR sensor with an empty hydrogel (A) and a hydrogel with an aptamer (B) for different angles of incidence, with Vancomycin 1 mM as analyte of interest, at 37°.

[0046] Fig. 6 shows a difference in the surface plasmon angle shift of a Multi-Parametric SPR sensor with an empty hydrogel (A) and a hydrogel with an aptamer (B), with Vancomycin 1 mM as analyte of interest, at 37°.

[0047] Fig. 7 shows suitability of the proposed concept for detection of analytes of interest, such as Vancomycin. (A) shows the shift in the surface plasmon angle 0 caused by different concentrations of Vancomycin, followed by a wash. (B) shows the sensor specificity, by illustrating the shift in the surface plasmon angle 0 caused by different analytes (Glucose, Vancomycin, Ampicillin, Hygromycin), followed by a wash.

[0048] Fig. 8 (A) shows an aptamer-hydrogel coated gold fiber-optic SPR probe. (B) shows a shift in the surface plasmon wavelength caused by the analyte of interest, Vancomycin (1 mM).

[0049] Fig. 9 illustrates the experimental framework for investigating an aptamer- crosslinked hydrogel composition on a thiolated gold surface of the SPR sensor of the present invention. (A) shows the 3-D network structure of the hydrogel composition as being composed of star-shaped, 8-arm PEG- norbornene and dithiol cross linker (the norbornene to thiol ratio ([ene] / [SH]) of 3 or 4 leaves free norbornenes reactive groups available for covalent bonding with free thiols on the gold surface of the sensor. (B) Thiol-modified aptamers are shown for incorporation into the 3-D network structure with the thiol-based cross linker by varying the [PEG-dithiol] / [aptamer] ratio from 2 to 12. MM (monofunctionalized split aptamer pairs), MD (one monofunctionalized and one difunctionalized split aptamer pair), DD (both split aptamer pairs dithiolated), and F (full aptamer sequence before splitting, with both ends thiolated). (C) Gold surface pre-functionalization is shown as being performed by incubating the surface in DTT, followed by exposure to DDT to promote a higher proportion of upright-oriented free thiols. Subsequently, the hydrogel precursor (pre-gel) solution was cast on the prefunctionalized gold surface and exposed to UV radiation to initiate the crosslinking of the 3-D network structure of the hydrogel.

[0050] Fig. 10 illustrates the two primary surface sensing techniques employed for comprehensive characterization. (A) Principle: For optical analysis, multi- parametric surface plasmon resonance (MP-SPR) was utilized to determine the refractive index and optical thickness of the hydrogel film by examining the waveguide pattern, while the surface plasmon resonance range was used to investigate angle shifts and study kinetics. To complement the optical measurements, quartz crystal microbalance with dissipation monitoring (QCM-D) was employed to assess the wet acoustic mass (via changes in resonance frequency) and energy dissipation, enabling the characterization of the viscoelastic properties of the hydrogel film during analyte-aptamer interactions; (B) Kinetics: Representative time dependent change in surface plasmon resonance angle change in MP-SPR and frequency and energy dissipation in QCM-D, induced by exposure to 1 mM Vancomycin; (C) Properties: Effect of 1 mM Vancomycin binding on aptagel, physical properties as depicted by resonance angular shifts in waveguide mode in MP- SPR indicating thickness and density change, supported by viscoelastic changes observed in QCM-D.

[0051] Fig 11 illustrates the effect of aptamer crosslinking configuration on biomolecular sensing observed by multi-parametric surface plasmon resonance (MP- SPR). (A) Split aptamer pair sequence of SEQ ID Nos:2 and 3, with the full original aptamer sequence consisting of SEQ ID NO:1.; (B) Schematic representation of variations in aptamer crosslinking configurations; (C) Formulations of aptamer-functionalized hydrogels; (D) Surface plasmon resonance (SPR) sensograms for different crosslinking configurations;

[0052] (E) Repeated injection of 1 mM vancomycin comparing the reversibility and reproducibility of DD1 -3, while the response of F1 -3 diminishes over time. Sensograms were normalized between the first data point before injection (representing the baseline) and the highest data point of the dataset;

[0053] (F) Representative case of fitting the experimental data using analysis software to determine the refractive index and optical thickness of the hydrogel. The inset in the figure provides a magnified view of the waveguide mode from the full spectrum; (G) The changes in SPR minimum angle (A0min, empty bars), thickness (Ac / %, striped bars), and refractive index (An) were calculated to compare the effect of vancomycin on different crosslinking configurations.

[0054] Fig. 12 Acoustic characterization of aptamer-functionalized hydrogels was performed to investigate the influence of aptamer crosslinking configuration on the viscoelastic properties (A) The maximum shifts in QCM-D resonance frequency (Af, empty bars) and energy dissipation (AD, striped bars) are presented upon exposure of the hydrogel layer on the gold electrode surface to 1 mM vancomycin (fifth overtone). (B) Time-independent f- D plots were constructed to correlate changes in resonance frequency and dissipation, which can provide insights into the viscoelastic contributions of the adsorbed species. (C) A schematic illustration depicts the proposed mechanism of dithiol split aptamer pairs upon analyte binding, leading to the observed optical and acoustic changes in the hydrogel layer.

[0055] Fig. 13 Characterization of split aptamer-functionalized hydrogel compositions and their response to vancomycin using MP-SPR. (A) Hydrogel formulations were prepared by varying the crosslinking density, represented by the molar ratio of norbornene groups to thiol groups ([ene] / [SH]), and the concentration of split aptamer pairs specific for vancomycin ranging from 0 to 3 mM. DDO-4 denotes blank hydrogels without aptamers, and PA3-4 refers to hydrogels incorporating a non-specific poly(A) sequence (27 nucleotides) as negative controls. (B) MP-SPR measurements were performed to evaluate the hydrogel response upon exposure to vancomycin. The changes in SPR minimum angle (A0min), shown as empty bars, and thickness (Ac / ), shown as striped bars, were calculated to quantify and compare the effect of vancomycin among the various hydrogel compositions.

[0056] Fig. 14 Sensor performance was evaluated through a dose-response study in 200- fold diluted rat plasma, and specificity testing was conducted using negative analytes. (A) Sensograms illustrating the real-time response obtained upon injecting various concentrations of vancomycin ranging from 10 pM to 1 mM. The inset provides a magnified view of the red dotted box from the main figure. (B) Angular reflectivity spectra for a hydrogel film measured in the presence of different vancomycin concentrations, highlighting the greater shift observed in the SPR mode (71.5 to 73 degrees) compared to the waveguide mode (62.5 to 63 degrees, observed from the first peak). (C) Specificity assessment using other commonly employed antibiotics at their maximum concentrations of 1 mM and 10 mM glucose, revealing negligible responses. The calculated limit of detection (LOD) was determined to be approximately 20 pM.

[0057] Fig. 15 Sensor sensitivity was evaluated in relation to characteristics of the 3-D network structure of the hydrogel composition. (A) Dithiol cross linkers of different molecular weights ranging from 0.3 to 10 kDa (0.6, kDa, 1 .0 kDa and 3.4 kDa) were used when preparing an aptamer-based hydrogel composition for use in the senor system of the present invention. (B) Limit of detection (LOD) determination. LOD improvement (i.e. sensitivity) was inversely correlated with the molecular weight of the cross linkers. While cross linkers with a molecular weight of 3.4 kDa showed a beneficial detection limit for the analyte of interest (here: vancomycin), sensitivity was further increased when cross linkers of lower molecular weights were used. LOD for 3.4 kDa cross linker: 16 nM, LOD for 1.0 kDa cross linker: 12 nM LOD for 0.6 kDa cross linker: 7.2 nM.

[0058] Fig. 16 Sensor sensitivity was further evaluated in relation to characteristics of the metal surface coating of the fiber sensor probe. (A) Schematic representation of cross section through a sensor probe for use in the sensor system of the present invention. From centre to perimeter, the shown layers represent: glass core, intermediate silver (Ag) layer, outer gold (Au) layer. (B) Bulk sensitivity of probe) was determined for a sensor probe with a single layer gold (Au) coating in comparison with bilayer (Ag / Au) coated sensor probes using sucrose concentration gradient. Incorporating an intermediate Ag layer increased sensitivity (Ag / Au 60:40). Increasing the ratio of Ag to Au in the bilayer further increased bulk sensitivity (Ag / Au 80:20). (C) Wavelength shifts upon analyte binding (here: vancomycin 1 , 11 , 111 and 55.5 pM) were determined for a sensor probe with a single layer gold (Au) coating in comparison with bilayer (Ag / Au) coated sensor probes. Including an intermediate Ag layer increased sensitivity (Ag / Au 60:40). Increasing the ratio of Ag to Au in the bilayer further increased bulk sensitivity (Ag / Au 80:20).

[0059] Fig. 17 Versatility of aptamer-based hydrogel compositions (aptagels) used in the sensor system of the present invention. (A) Schematic representation of split aptamer and incorporation into aptamer-based hydrogel composition, and subsequent collapse of hydrogel upon binding of analyte, (here: Corticosterone; 0.35 kDa; hydrophobic). (B) Aptamer-based hydrogel composition increases sensitivity in comparison to monolayer assay of splitaptamer directly on gold SPR surface compared to split aptamer in hydrogel. (C) A representative sensorgram of aptamer-based hydrogel sensor depicting fully reversible detection of an analyte of interest in micromolar range (here: Corticosterone; 0.35 kDa; hydrophobic). (D) A representative sensorgram of aptamer-based hydrogel sensor depicting fully reversible detection of an analyte of interest in micromolar range (here: dopamine; 0.15 kDa; hydrophilic). (E) A representative sensorgram of aptamer-based hydrogel sensor depicting fully reversible detection of an analyte of interest in micromolar range (here: vancomycin; 1.4 kDa; hydrophilic). (F) A representative sensorgram of aptamer-based hydrogel sensor depicting fully reversible detection of an analyte of interest in nanomolar range (here: TNF alpha; 17 kDA monomer; 51 kDa homotrimer).

[0060] Fig. 18 Schematic of an implantable fiber optic surface plasmon resonance (FO- SPR) aptagel sensor probe of the sensor system of the present invention designed for in vivo pharmacological monitoring in the brain of rats. (A) FO- SPR aptagel sensor system consisting of a light source and spectrophotometer detector connected to a gold coated optical fiber probe through a Y-splitter. Incident white light is reflected through the optical fiber glass core generating surface plasmon resonance at the gold and dielectric medium interface containing the aptagel. Once implanted into the brain tissue, the hydrogel composition acts as a protective and anti-fouling material against biological interferents such as proteins and cells, while transducing the optical changes caused by the binding of analyte to the allosteric molecules forming an integral part of the 3-D network structure (here: aptamers covalently crosslinked into the hydrogel matrix). Drug analyte binding to split-aptamers in the aptagel causes a collapse / contraction of the hydrogel composition (i.e. deswelling, reducing thickness and increasing density), which raises the refractive index, producing a measurable redshift in the plasmon resonance wavelength. The fast dissociation of the aptamer and drug interaction allows for reversible molecular interaction and hence continuous drug monitoring. (B) The expected analyte dose-dependent increase in SPR wavelength. (C) A schematic sensorgram depicting reversible dose-dependent response to drug analyte.

[0061] Fig. 19 The working principle of the FO-SPR aptagel sensor system according to the invention and performance parameters. (A) Schematic of the in vitro sensor system for an FO-SPR aptagel sensor. Fiber probes are inserted in a flowcell made of polydimethylsiloxane (PDMS) allowing for up to 3 mm contact with liquid medium. A temperature controller permits sensor measurements to be performed at the physiologically relevant temperature of 37°C to mimic the in vivo condition. A steady exchange of fluid samples with varying drug analyte concentrations is introduced through a fluidic inlet and outlet with upwards stream to ensure efficient fluid replacement. (B) Reflection spectra of bare gold and hydrogel coated FO-SPR probes. Hydrogel thin film coating on FO-SPR probe causes shift in the resonant wavelength and decrease in reflection due to the dampening of surface plasmons. (C) Theoretical SPR shifts expected from hydrogel coated FO-SPR probes dependent on refractive index changes compared to that of experimentally measured shifts for bare gold and hydrogel coated FO-SPR probes measured using increasing concentrations of sucrose. (D) Performance parameters of bare gold and hydrogel coated FO-SPR probes. The decrease in reflection causes an increase in the full-width-half-maximum, as broadening of the SPR dip is observed, thereby decreasing the overall figure of merit.

[0062] Fig. 20 The real-time, reversible monitoring of vancomycin using bare gold and aptagel coated FO-SPR probes. (A) Direct comparison of vancomycin monitoring using the split-aptamer P27 functionalized on bare gold using monolayer (black) architecture versus covalently crosslinked into hydrogel, aptagel (grey). Both sensors allow reversible and continuous monitoring of vancomycin, with the major drawback of monolayer architecture being steady supply of free split 2 aptamer segment in solution. Whereas, aptagel allows direct detection of vancomycin in the self-contained assay components in the hydrogel architecture. (B) Dose-dependent response of FO-aptagel sensors to vancomycin in assay buffer (PBS; squares) and artificial cerebrospinal fluid (aCSF; diamonds). The LOD was calculated using linear regression fitting for both PBS and artificial cerebrospinal fluid (aCSF) conditions. Solid lines represent the fitted linear regression, and shaded regions indicate standard deviations based on three independently prepared aptagel-integrated FO- SPR sensors. (C) Comparison of FO-SPR response for vancomycin detection using the monolayer-based split aptamer assay and the aptagel-integrated setup. Dose-response curves depict the SPR shift (AA) as a function of vancomycin concentration for two sensing strategies: (i) the monolayer-based approach using the P27 split aptamer (with both Split 1 and Split 2 present throughout sensing and washing steps)in S2 buffer, which mimics the aptagel environment, and (ii) the aptagel-integrated system of the present invention. The aptagel-integrated system demonstrates enhanced sensitivity, in particular at higher concentrations. Error bars represent the standard deviation from triplicate measurements.

[0063] Fig. 21 Continuous drug monitoring sensor performance in complex biological fluid. (A) Sensorgram depicting continuous monitoring of vancomycin with increasing and decreasing concentrations (0.78-100 pM) directly in undiluted horse serum. Sensor response increases to saturation at the highest concentration and recovers fully to baseline level when analyte is depleted indicating full reversibility. Fast dissociation allows for sensor recovery within a physiologically relevant timeframe of drug clearance, expected in the pharmacological monitoring application. (B) The calibration curve of the sensor established in undiluted serum for detection of vancomycin.

[0064] Fig. 22 A combined dose-response obtained for vancomycin sensing on aptagel FO- SPR probes employing different buffer conditions PBS, BP 1 :5, and HS. The LODs calculated for different buffer conditions are 2 pm (PBS), 1 .8 pm (BP 1 :5) and 2.5 pm (HS). The complex biological fluids do not have a significant effect on the LODs, except a small increase in the case of undiluted Horse Serum (HS). The signal gets elevated by 1 .5-2 times in the lower linear ranges for HS and BP 1 :5 cases compared to the standard PBS assay buffer. Fig. 23 Continuous pharmacological monitoring of vancomycin in the brain of anesthesized rats. Following surgical implantation of FO-SPR aptagel probes into the prefrontal cortex (PFC) of anesthesized rats, the sensor was allowed to stabilize for at least 60 minutes. Vancomycin drug in saline was administered through the tail vein at 75 mg / kg or 150 mg / kg doses. Within 5- 20 minutes after administration, the sensor response was observed and reached plateau after around 1 hour and fully cleared within 1 -2 hours. Sensorgrams in panels (A) and (B) show sensor response at the two different administered doses (75 and 150 mg / kg) depicting clear dose-dependence. (C) One rat received both different doses after clearance of the first administered dose of 75 mg / kg and both profiles show saturation and full clearance.

[0065] DETAILED DESCRIPTION OF THE INVENTION

[0066] In order to provide a clear and consistent understanding of the specification and claims, and the scope to be given such terms, the following definitions are provided.

[0067] Definitions

[0068] While the term "allosteric" is commonly used in the context of molecules with enzymatic or receptor-like activities, in the context of the present disclosure, the term “allosteric molecule or allosteric fragment thereof” refers to any molecule or fragment thereof that undergoes a conformational change upon specific binding to a ligand or binding partner, regardless of its specific function. As such, the term expressly includes molecules or fragments thereof that have no defined activity or function but undergo conformational changes upon binding a ligand or binding partner, i.e. it includes molecules and fragments that have allosteric properties.

[0069] In the context of the present disclosure, the term “conformational change” means a structural alteration in a molecule, typically a biomolecule like a protein or nucleic acid, resulting in a different spatial arrangement of its atoms or components. Conformational changes of molecules are discussed herein as the consequence of an allosteric molecule or allosteric fragment thereof binding to another molecule, typically the analyte of interest of the sensor systems and methods and uses for sensing an analyte of interest disclosed herein. Here the term “biocompatible” or “biocompatibility” of a hydrogel-based sensor system - and in particular of its sensor probe - means that the also the physical properties of known hydrogel biosensors, such as overall probe size as well as probe rheology (i.e. stiffness, surface roughness, viscoelasticity of the hydrogel composition, etc.) are substantially compatible with the environment in which the sensing is to occur such that a foreign body response is avoided altogether or at least avoided for the time period required for the sensor system to acquire the necessary sensor data.

[0070] The term “sensor data being indicative of a surface plasmon resonance wavelength shift” means that the sensor data can be used to derive whether a surface plasmon resonance wavelength shift has occurred and to derive an amplitude of the surface plasmon resonance wavelength shift, as compared to other surface plasmon resonance wavelength shifts measured using the same surface plasmon resonance sensor to. The sensor data being indicative of a surface plasmon resonance wavelength shift does not necessarily include a numeric representation of a surface plasmon resonance wavelength or surface plasmon resonance wavelength shift. It may include a measurement being indicative of (i.e., representing) the surface plasmon resonance wavelength or surface plasmon resonance wavelength shift, such as an emission wavelength or angle of incidence having triggered surface plasmon resonance, or a shift between emission wavelengths or angles of incidence having triggered surface plasmon resonance during at least two different measurements.

[0071] Similarly, “sensor data being indicative of a shift in resonance frequency of a quartz crystal” or “sensor data being indicative of a shift in energy dissipation of a quartz crystal” means that the sensor data can be used to derive the shift in resonance frequency or shift in energy dissipation. For example, such sensor data can be provided by a QCM-D (Quartz Crystal Microbalance and Dissipation) sensor data. The sensor data being indicative of the shift in resonance frequency or energy dissipation does not necessarily include a numeric representation of the resonance frequency, energy dissipation, or shift thereof. It may include a measurement being indicative of (i.e., representing) the resonance frequency, energy dissipation, or shift thereof. In particular, to represent a shift, the sensor data may comprise multiple measurements. In the context of the present disclosure, the term “3-D network structure” refers to the network of 3-dimensionally interlinked molecules providing for the structural integrity of the hydrogel composition. This 3-D network structure comprises hydrophilic polymers crosslinked to each other via cross linkers - predominantly via well-known and established functionalized polymer cross linkers (such as by, e.g., thiolated polyethylene glycol) (PEG) cross linkers) but also - to a varying degree - via functionalized allosteric molecules or allosteric fragments thereof (such as, e.g. by di-thiolated DNA aptamers). Accordingly, the cross-linkers, be they well-known and established functionalized polymer cross linkers (such as by, e.g., thiolated poly(ethylene glycol) (PEG) cross linkers) or functionalized allosteric molecules (such as, e.g. by di-thiolated DNA aptamers), form an integral part of the 3-D network structure. Accordingly, a conformational change of an allosteric molecule or allosteric fragment thereof upon specific binding to a ligand or binding partner directly affects the spatial dimensions of the 3-D network structure.

[0072] The term “sensor data being indicative of an alteration of the spatial dimensions of the 3-D network structure” means that the sensor data can be used to detect a change (i.e., alteration) in the spatial dimensions of the 3-D network structure, i.e., a contraction or expansion of the 3-D network structure. For example, sensor data being indicative of a surface plasmon resonance wavelength shift is sensor data being indicative of the alteration of the spatial dimensions of the 3-D network structure, as the alteration of the 3-D network structure contributes to the surface plasmon resonance wavelength shift. Similarly, sensor data being indicative of a shift in resonance frequency or energy dissipation of a quartz crystal is sensor data being indicative of the alteration of the spatial dimensions of the 3-D network structure, as the alteration of the 3-D network structure contributes to the shift in resonance frequency and energy dissipation.

[0073] In addition to the above definitions, and unless the context clearly requires otherwise, throughout the description and the claims, the words “comprise”, “comprising”, and the like are to be construed in an inclusive sense as opposed to an exclusive or exhaustive sense; that is to say, in the sense of “including, but not limited to”. Further, reference throughout this specification to “one embodiment”, “some embodiments” or “an embodiment” means that a particular feature, structure or characteristic described in connection with the embodiment is included in at least one embodiment of the present invention. Thus, appearances of the phrases “in one embodiment”, “in some embodiments” or “in an embodiment” in various places throughout this specification are not necessarily all referring to the same embodiment, but may. Furthermore, the particular features, structures or characteristics may be combined in any suitable manner, as would be apparent to one of ordinary skill in the art from this disclosure, in one or more embodiments.

[0074] As used herein, unless otherwise specified the use of the ordinal adjectives "first", "second", "third", etc., to describe a common object, merely indicate that different instances of like objects are being referred to, and are not intended to imply that the objects so described must be in a given sequence, either temporally, spatially, in ranking, or in any other manner.

[0075] As used herein, the term “exemplary” is used in the sense of providing examples, as opposed to indicating quality. That is, an “exemplary embodiment” is an embodiment provided as an example, as opposed to necessarily being an embodiment of exemplary quality.

[0076] Embodiments of the invention

[0077] As mentioned above, there is a need in the art for improved hydrogel-based biosensor systems for sensing an analyte of interest, in particular for continuous sensing of such an analyte in vivo.

[0078] This need exists because, for example, current in vivo biosensors lack biocompatibility: While Electrochemical Aptamer-based Bio-sensors (EAB-sensors) are most-advanced in this context, they are currently functional for only up to 5-6 hours and upcoming technologies do not allow for reliable and continuous in vivo sensing of a specific analyte of interest. In vivo Fast-Scan Cyclic Voltammetry (FSCV) and Field-Effect Transistor (FET) based sensors, which may use elastomers and materials to increase biocompatibility of implantable probes. However, they can resolve individual neurotransmitter but cannot be applied to detect molecules that do not have electroactivity. The present disclosure addresses the need in the art by providing a sensor system for sensing an analyte of interest, comprising:

[0079] - a sensor, such as a surface plasmon resonance sensor or a quartz crystalbased sensor, to provide sensor data, wherein the surface plasmon resonance sensor comprises a metal-coated surface;

[0080] - a hydrogel composition contacting the metal-coated surface of the sensor, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest; and

[0081] - processor circuitry to determine information on a presence of the analyte of interest based on the sensor data, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel, and wherein the sensor data is indicative of the alteration of the spatial dimensions of the 3-D network structure, e.g., being indicative of the surface plasmon resonance wavelength shift being contributed to by the alteration of the spatial dimensions of the 3-D network structure.

[0082] As mentioned above, each allosteric molecule or allosteric fragment thereof undergoes a conformational change upon binding to the analyte of interest and assembly of the respective analyte binding complex. It is this conformational change of each individual allosteric molecule, which ultimately alters the spatial dimensions of the 3-D network structure of the hydrogel, thereby contributing to effects such as the surface plasmon resonance wavelength shift or a shift in resonance frequency or energy dissipation of a quartz crystal.

[0083] As is well known in the art, a large variety of diverse allosteric molecules exist. For example, allosteric molecules include, without limitation:

[0084] - enzymes, which undergo conformational changes upon binding of substrates or other molecules, typically at sites distinct from their active sites. The conformational changes regularly either enhance or inhibit the enzyme's catalytic activity; - receptors, in particular cell surface receptors, such as G protein-coupled receptors (GPCRs) and ion channels, wherein ligand binding to these receptors induces conformational changes that can alter their affinity for other ligands or activate downstream signaling pathways.

[0085] - structural proteins, which are involved in cellular scaffolding or structural support can also exhibit allosteric behavior. For example, the hemoglobin protein undergoes allosteric changes in response to binding oxygen, facilitating the release or uptake of oxygen in different tissues.

[0086] - certain DNA or RNA nucleic acid sequences, like aptamers, can exhibit allosteric behavior. For instance, riboswitches are RNA molecules that undergo conformational changes upon binding specific metabolites, thereby regulating gene expression.

[0087] - Antibodies can undergo conformational changes induced through the binding of antigens, wherein such changes typically affect the antibody’s binding affinity or trigger downstream immune responses.

[0088] As such, the specific hydrogel composition of the sensor system of the present invention is molecularly responsive to the analyte of interest and can act as a transducer, converting analyte binding events into measurable surface plasmon resonance wavelength shifts or shifts in resonance frequency or energy dissipation of a quartz crystal.

[0089] As indicated, the sensor system comprises a sensor, such as a surface plasmon resonance sensor or a quartz crystal-based sensor. In general, surface plasmon resonance sensors are sensors that are based on measuring an excitation of surface plasmons, i.e., coherent delocalized electron oscillations that occur at an interface between a thin metal layer and a dielectric, such as air or water. These electron oscillations are excited by a P-polarized light beam that is directed at the metal through glass. Light travelling through the glass excites an evanescent wave at the interface between the glass and the metal when the light undergoes total internal reflection at the boundary, due to the glass and the metal having different refractive indices. The evanescent wave extends beyond the glass into the metal and dielectric and excites the surface plasmons under appropriate conditions. If the light is incident to the metal at a specific angle or at a specific wavelength, a resonance effect occurs. In particular, when the energy and the momentum of the incident light and surface plasmon wave match, a resonance occurs which results in a noticeable dip in the intensity of the reflected light. Whether a resonance effect can be observed depends on the angle and wavelength of the light beam, and on the dielectric functions of the metal and dielectric. To determine the surface plasmon wavelength, or at least observe and measure a change in the surface plasmon wavelength, both the angle of incidence and the emission wavelength can be varied to determine the resonance conditions. For example, in an example implementation, the angle of incidence may be varied to determine the angle of incidence at which the resonance effect occurs. In another implementation, the emission wavelength may be varied to determine, for a constant angle of incidence, the emission wavelength, at which the resonance occurs. For example, the emission wavelength, at which the resonance occurs, may be determined by identifying the emission wavelength at which a dip in reflected light intensity can be observed. Similarly, the angle of incidence, at which the resonance occurs, may be determined by identifying the angle of incidence, at which a dip in reflected light intensity can be observed.

[0090] The surface plasmon resonance effect depends, apart from the emission wavelength and angle of incidence of the light beam, on the dielectric functions of the metal and of the dielectric, with the dielectric function of the dielectric being closely related to the refractive index of the dielectric medium. Thus, when the refractive index of the dielectric medium changes, the resonance effect can be observed at a different surface plasmon resonance wavelength, resulting in a wavelength shift. This wavelength shift is denoted surface plasmon resonance wavelength shift. In the present disclosure, the surface plasmon resonance sensor is used to provide sensor data that is indicative of the surface plasmon resonance wavelength shift, i.e. , that includes a measurement or measurements representing the amplitude of the surface plasmon resonance wavelength shift. However, the sensor data does not necessarily include the surface plasmon resonance wavelength shift in terms of wavelength or frequency of the surface plasmons. For example, the sensor data may comprise a measurement being indicative of (e.g., representing) the surface plasmon resonance wavelength and surface plasmon resonance wavelength shift, such as emission wavelengths and corresponding shifts in emission wavelengths or angles of incidence and corresponding shifts in angles of incidence. In surface plasmon resonance (SPR)-based biosensors, the change in the refractive index causing the surface plasmon resonance wavelength shift is usually caused by biological activity, such as an analyte binding to a receptor that is immobilized at the metal surface. When the analyte binds to the receptor, the refractive index of the surrounding medium (the dielectric) changes, affecting the wavelength at which resonance occurs. In other words, due to the biological activity, a shift in the surface plasmon resonance wavelength occurs, which is represented by the sensor data.

[0091] In one or more of the above-mentioned embodiments, at least some of the hydrophilic polymers of the hydrogel composition connect the 3-D network structure to the metal- coated surface of the sensor. This direct connection transmits a conformational change of the allosteric molecules or allosteric fragments thereof to the metal-coated surface.

[0092] Importantly, through assembly of respective analyte binding complexes, the wavelength shift measured is indicative of the number of analyte molecules present in the aqueous environment surrounding the sensor probe. Furthermore, by way of its 3-D network structure, the hydrogel composition of the present invention is significantly more sensitive than a 2-D mono-layer surface plasmon resonance probe as the 3-D network structure, and in particular its alteration of its spatial dimensions in response to analyte binding, allows for the detection of significantly more analyte binding events than a 2-D mono-layer, including binding events distant from the surface.

[0093] The same effect can also be used in other types of sensors having a metal surface, e.g., quartz crystal-based sensors, such as QCM-D sensors. A QCM-D sensor is a sensor that is based on a Quartz Crystal Microbalance with Dissipation monitoring. It is a type of sensor that is used to monitor changes in mass on a quartz crystal sensor while also measuring the energy dissipation associated with those changes. In the present context, QCM-D sensors can be used to determine a shift in resonance frequency and energy dissipation of a quartz crystal sandwiched between two metal electrodes, such as gold electrodes, with the shift being caused by the contraction and expansion of the 3-D network structure of the hydrogel upon assembly and disassembly of the analyte-binding complex, respectively. As will be understood and is described in further detail below, the mechanical properties of the 3-dimensional (3-D) network structure of the hydrogel can be adjusted for optimal analyte sensing through the choice of (a) the hydrophilic polymers, (b) the allosteric molecules (c) the cross linkers used to connect the majority of the attachment points of the hydrophilic polymers as well as (d) the relative proportions of these components.

[0094] In any one of the before-mentioned embodiments, the hydrophilic polymers may be poly(ethylene glycol) (PEG), dextran, poly(N-isopropylacrylamide) (PNIPAAM), polyvinyl alcohol (PVA), hydroxyethyl methacrylate (HEMA), poly(acrylic acid) (PAA), poly(ethylene oxide) (PEG), alginate and / or hyaluronic acid (HA) polymers.

[0095] The hydrophilic polymers may be multivalent polymers, preferably comprising between 3 and 10 attachment points. The attachment points of the hydrophilic polymers allow for attachment to allosteric molecules or allosteric fragments thereof to the polymers on the one hand but also for attachment of the polymers to each other as well as to the metal coated surface on the other. As such, the number of attachment points of the hydrophilic polymers allows for the tuning of the 3-D network structure of the hydrogel composition with respect to its binding capacity for an analyte of interest via attachment of the allosteric molecule or fragment thereof, its viscoelasticity via the degree of crosslinking between the polymers as well as its transmission properties via connection to the metal-coated surface.

[0096] The length of the cross linker used in formation of the 3-D network structure of the hydrogel composition also influences hydrogel properties. For example, the use of shorter cross linkers leads to a more compact 3-D network structure, which in turn has been associated with increased sensitivity (see Fig. 15). In particular, cross linkers of molecular weights between 0.3 and 10 kDa are useful in formation of suitable 3-D network structures.

[0097] The use of hydrogel compositions having a compact 3-D network structure (improves the hydrogel’s sensitivity, in particular when the analyte of interest is a smaller molecule itself. Without wanting to be bound by theory, this effect may be the consequence of limited diffusion of larger highly abundant proteins (such as, for example, albumins or gamma globulins) into the hydrogel, which possibly interfere with analyte binding by the allosteric molecules. Further, a compact 3-D network structure is beneficial for the sensor system’s overall integrity and longevity as the smaller pores of such a network structure also protect the allosteric molecules from degradation due to limiting the diffusion of large degrading enzymes, such as proteases and / or nucleases. Even further, hydrogel compositions comprising compact 3-D network structure limit the diffusion large aggregating proteins involved in biofouling such as, e.g., albumins, immunoglobulins, mucins, lipoproteins or fibrinogen.

[0098] Advantageously, the hydrophilic polymers comprise multivalent poly(ethylene glycol) (PEG) polymers, preferably the multivalent PEG polymers comprise between 3 and 10 attachment points such as 3, 4, 5, 6, 7, 8, 9, or 10 attachment points.

[0099] In one embodiment, a multivalent PEG polymer with 8 attachment points is utilized. The 8-arm star-shaped PEG-norbornene architecture is particularly useful for tuning the hydrogel composition of the present sensor system. Star polymers, characterized by a minimum of three macromolecular chains radiating from a central core, have garnered significant attention due to their unique topological structures and desirable physical / chemical attributes. For instance, the star-shaped architecture provides a higher number of reactive norbornene end-groups per molecule, facilitating enhanced efficiency and control over click chemistry crosslinking. This topology also allows for the incorporation of various functionalities and improved responsiveness to stimuli, stemming from the high density of functional groups compared to linear polymers having only one or maximally two reactive end groups. Viscosity in polymer solutions arises largely from chain entanglements and hydrodynamic interactions compared to linear polymers and star polymers with between 3 and 10 arms / attachment points exhibit reduced solution viscosity in comparison to linear counterparts of equivalent molecular weight, due shorter individual arms leading to fewer arm entanglements. Solution viscosity is highest for linear polymers and reduces - up to a certain point - with the degree of valency. This is because, at equal molecular weight, the hydrodynamic radius of the polymer decreases with increasing arm number (at same MW), so that the polymer occupies a smaller effective volume (i.e. has an increased “compactness”), which leads to fewer entanglements and, consequently, to a lower intrinsic viscosity. As such, at equal molecular weight, a 3-armed polymer has a lower relative viscosity than a linear polymer, at equal molecular weight, a 4-armed has a lower relative viscosity than a 3-armed polymer, at equal molecular weight, a 5-armed has a lower relative viscosity than a 4-armed polymer and so forth. The reduction is most pronounced from linear to 3-arm and gradually flattens as the arm number increases beyond 5 to 6. However, while an 8-arm polymer displays the lowest relative viscosity (due to maximal compactness), this beneficial effect is necessarily present in multivalent polymers, such as PEG polymers comprising between 3 and 10 attachment points such as 3, 4, 5, 6, 7, 8, 9, or 10 attachment points, which is advantageous for producing thin, homogenous hydrogel films as it enhances processability and coatability.

[0100] As already mentioned, in one or more of the above embodiments, the allosteric molecules or allosteric fragments thereof may be connected with at least some of the hydrophilic polymers of the hydrogel composition. Preferably, each allosteric molecule or allosteric fragment thereof links two hydrophilic polymers such as to form an integral part of the 3-D network structure.

[0101] Allosteric molecules or allosteric fragments thereof are known to undergo a conformational change upon specific binding to binding partner and some aptamers are known to undergo conformational change when binding their respective target molecule such that these aptamers are allosteric molecules in the sense of the present application.

[0102] Accordingly, in one or more of the above embodiments, the allosteric molecules or allosteric fragments thereof are aptamers or fragments thereof, which assemble into an aptamer-analyte binding complex upon binding to the analyte of interest. An aptamer is a single-stranded DNA or RNA molecule that can bind to specific target molecules with high specificity. Preferably, the aptamers or fragments thereof are DNA aptamers or fragments thereof.

[0103] When the allosteric molecules are aptamers, each aptamer comprises between 20 and 90 nucleotides and, when the allosteric molecule fragments are aptamer fragments, each aptamer fragment comprises between 10 and 40 nucleotides. As is well-known, aptamers can fold into complex three-dimensional structures due to intramolecular interactions such as hydrogen bonding, stacking interactions, and loop formations. Aptamers can be selected to bind a wide range of target analytes, including small molecules, proteins, peptides, and even whole cells, typically through Systematic Evolution of Ligands by Exponential enrichment (SELEX). SELEX is an iterative in vitro selection process that allows for the isolation of aptamers with high specificity for a target molecule from a diverse pool of random sequences. SELEX begins with the generation of a random library of unique single-stranded DNA (ssDNA) or RNA sequences of between 20-100 nucleotides in length. Once generated, such a library typically comprises between 1013to 1015different sequences, which are then incubated with the target molecule (i.e. the analyte of interest) under defined conditions that promote binding. After incubation, unbound sequences are separated from the target-bound sequences using available separation techniques well-known in the art, for example size-exclusion chromatography, filtration or magnetic bead-based separation. The analyte-bound sequences of the library are then amplified using a polymerase chain reactions (PCR) for DNA sequences or reverse transcriptase PCR (rtPCR) for RNA sequences to increase the abundance of the analyte-binding sequences in the remaining pool of sequences. Usually, the sequences recovered from the pool are subjected to additional rounds of selection (typically 5-15 rounds), with each round aiming at enriching the pool for sequences with higher affinity and specificity for the analyte. After several rounds of amplification and enrichment, individual aptamer sequences are cloned and sequenced to identify aptamer candidates. These candidates are then synthesized and experimentally validated for their ability to bind to the analyte of interest with high specificity.

[0104] While SELEX is the most common method for aptamer selection, there are also in silico approaches using computational methods for designing and / or predicting aptamers with desired binding properties. However, these methods generally rely on structural and sequence data from previously validated aptamers and may, therefore, not be as versatile or effective as experimental SELEX for discovering truly novel analyte-specific aptamers. In the context of the present disclosure an aptamer-based hydrogel composition (i.e. a hydrogel composition as described above in which the allosteric molecule or allosteric fragment thereof forming an integral part of the 3-D network structure of the hydrogel composition is a DNA or RNA aptamer) is also referred to as “aptagel”.

[0105] Some classes of molecules that aptamers have been demonstrated to bind to with high specificity include, without limitation:

[0106] Small molecules - such as small molecule drugs, metabolites, toxins and organic molecules;

[0107] Proteins - such as enzymes, antibodies, antigens, receptors and cytokines.

[0108] Peptides - such as specific peptides or peptide motifs, allowing for aptamer application in peptide purification or interference with protein-protein interactions.

[0109] Nucleic acids - such as specific DNA or RNA sequences, enabling aptamer application in sequence-specific detection or inhibition of nucleic acid function. Metal ions -useful in aptamer application in metal ion detection or chelation therapy. Viruses - including viral particles or viral proteins, offering potential for aptamers to be utilized in antiviral therapy or viral detection.

[0110] Bacteria or other microorganisms - such as bacterial strains, pathogens or microbial components, facilitating aptamer application in bacterial detection or antimicrobial therapy.

[0111] Cells - such as specific cell surface markers or receptors, enabling aptamer use in targeted drug delivery, cell sorting or diagnostic imaging.

[0112] Lipids - allowing for potential aptamer application in lipid detection or lipid-mediated diseases.

[0113] Organic compounds - such as environmental pollutants, food additives or carbohydrates.

[0114] In this context, it is noteworthy that fragments of an aptamer may hybridize to substantially re-assemble into a near complete aptamer. A single aptamer that has been broken up into two or more fragments, which - when brought into close proximity or assembled together - regains its ability to bind to the aptamers target molecule with high specificity is commonly known as a split aptamer. Depending on the actual splitting site within the original aptamer, split aptamer fragments are typically between 10 and 40 nucleotides in length. While in some instances, each fragment of an allosteric split aptamer alone does not possess the full binding capability of the original aptamer, re-assembly of the split aptamer fragments leads to a split aptamer that can again regained its allosteric properties as well as its ability bind to the target molecule with high specificity and affinity.

[0115] Accordingly, it will be appreciated that the allosteric molecules or allosteric fragments thereof, such as the above-described aptamers or fragments thereof, employed in the embodiments of the present invention can be selected to bind a wide range of targets, including small molecules, proteins, peptides and even whole cells.

[0116] Typically, the analyte of interest is

[0117] - an organic molecule, a pharmaceutically active substance, a metabolite, a toxin or a hormone; or

[0118] - a protein or peptide such as an enzyme, an antibody, a protein antigen, a peptide antibiotic, a proteinaceous viral particle, a viral protein, a protein receptor or a cytokine or

[0119] - a nucleic acid molecule; or

[0120] - a metal ion; or

[0121] - a lipid; or

[0122] - a carbohydrate.

[0123] In the hydrogel compositions of the embodiments of the sensor system described in the present disclosure, the conformational changes of the allosteric molecules or fragments thereof upon binding of the analyte of interest is harnessed and ultimately converted into an optically measurable signal. In some embodiments this capability of the sensor system of the present disclosure is mediated through direct connections between the allosteric molecules or allosteric fragments thereof and the hydrophilic polymers within the hydrogel composition. As such, the 3-D network structure of the hydrogel composition regularly comprises several reaction products of a thiol-ene reaction between one functional moiety of one allosteric molecule or allosteric fragment thereof and a correspondingly reactive group of one attachment point of one of the hydrophilic polymers. Typically, the 3-D network structure of the hydrogel composition comprises several reaction products of thiol-ene reactions between two or more functional moieties of one allosteric molecule or allosteric fragment thereof and correspondingly reactive groups of one respective attachment point of two or more separate hydrophilic polymers. As such, it will be understood that the allosteric molecules or allosteric fragments thereof used in the preparation in the hydrogel composition of the sensor system of the present disclosure originally comprised at least one functional moiety for participation in a thiol-ene reaction, wherein, preferably, this moiety is a thiol moiety.

[0124] The functional moiety allows for the allosteric molecule or allosteric fragment thereof to be connected with at least some of the hydrophilic polymers of the hydrogel composition, preferably via a thiol-ene reaction with a correspondingly reactive group of one respective attachment point of the hydrophilic polymer.

[0125] In some instances, the allosteric molecules or allosteric fragments thereof used in its preparation are mono-thiol-functionalized or di-thiol-functionalized split DNA aptamers or combinations thereof, such as mono-thiol-functionalized split aptamer fragment pairs, a combination of mono-thiol-functionalized and di-thiol-functionalized split aptamer pairs, di-thiol-functionalized split aptamer pairs or the complete original aptamer sequence with functional thiol modifications at both termini.

[0126] In embodiments, where the hydrophilic polymers comprise reactive ene-groups as their attachment points, cross linkers and / or allosteric molecules comprising correspondingly reactive thiol-groups can be used to enable formation of the of the 3- dimensional (3-D) network structure.

[0127] In such embodiments, the properties of the 3-dimensional (3-D) network structure of the hydrogel can be tuned by adjusting the molar [ene] / [SH] ratio, while maintaining a constant [ene] concentration (i.e. a constant concentration of the hydrophilic polymer providing the reactive ene-groups, for example a concentration of 10 mM). At [ene] / [SH] ratios below 1 , the presence of free thiol groups in the hydrogel further allowed for direct covalent bonding of the 3-D network structure of the hydrogel composition to the metal surface of the sensor. However, this approach requires excessively high concentrations of thiol functionalized allosteric molecules or thiol functionalized allosteric fragments thereof, leading to increased potential waste and reduced reproducibility due to the propensity of functionalized allosteric molecules or functionalized allosteric fragments thereof to form mono-tethered attachments to the hydrogel. Consequently, [ene] / [SH] ratios exceeding 1 are considered optimal, as they provided free ene groups capable of forming covalent bonds with a thiolated metal surface of the sensor. Further, while ratios below 2.5 may yield insufficient attachment strength on the metal surface, ratios above 4.5 may exhibit unacceptably low mechanical properties due to decreased crosslinking density. Therefore, [ene] / [SH] ratios of 3 and 4 are particularly suitable to ensure robust attachment of the 3-D network structure to the metal surface of the sensor as well as appropriate gel mechanical characteristics of the hydrogel composition useful in continuous sensing applications.

[0128] Typically though, the allosteric molecules or allosteric fragments thereof used in the preparation in the hydrogel composition of the sensor system of the present disclosure originally comprised two or more such functional moieties for participation in thiol-ene reactions, which during preparation of the hydrogel composition each react with correspondingly reactive groups of one respective attachment point of two or more separate hydrophilic polymers, such that the allosteric molecules or allosteric fragments thereof link two or more hydrophilic polymers within the 3-D network structure. In such embodiments, the connected allosteric molecule or allosteric fragment thereof forms an integral part of the 3-D network structure of the hydrogel composition.

[0129] In a particular embodiment the sensor system of the first aspect of the present invention comprises

[0130] - a fiber optic surface plasmon resonance sensor (FO-SPR) sensor with a bilayer metal-coated surface comprising an inner silver (Ag) layer and an outer gold (Au) layer to provide sensor data;

[0131] - a hydrogel composition contacting the outer gold (Au) layer of the bi-layered metal-coated surface of the sensor, wherein the hydrogel composition comprises a 3- dimensional (3-D) network structure including (a) multivalent poly(ethylene glycol) (PEG) polymers having between 3 and 10 attachment points as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof are connected with at least some of the multivalent hydrophilic polymers such that each allosteric molecule links two multivalent poly(ethylene glycol) (PEG) polymers and, thereby, forms an integral part of the 3-D network structure, wherein at least some of the multivalent hydrophilic polymers are interlinked by poly(ethylene glycol) (PEG) cross linkers of a molecular weight ranging between 0.3 and 10 kDa, and wherein the allosteric molecules or fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest; and

[0132] - processor circuitry to determine information on a presence of the analyte of interest based on the sensor data, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel, and wherein the sensor data is indicative of the alteration of the spatial dimensions of the 3-D network structure.

[0133] In such embodiments:

[0134] • the ratio of silver to gold in the bilayer metal coated surface may be 60 / 40, preferably 80 / 20,

[0135] • the multivalent poly(ethylene glycol) (PEG) polymer comprises eight (8) attachment points, preferably the multivalent poly(ethylene glycol) (PEG) polymer is star-shaped 8-arm PEG norbornene, and / or

[0136] • the cross linkers connecting at least some of the multivalent hydrophilic polymers are PEG) are cross linkers of a molecular weight ranging from about 0.3 kDa to 0.8 kDa, preferably derived from dithiol PEG crosslinkers having a molecular weight of about 0.6 kDa.

[0137] In some embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID:1 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGTGGGTCGGGTC G) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID:3 (TTGTTCGCCTATTGTGGGTCGGGTCG).

[0138] In some embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID:4 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGTGGGTCGGGTC GTTTTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID: 5

[0139] (TTGTTCGCCTATTGTGGGTCGGGTCGTTTTT).

[0140] In some embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID:6 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGTGGGTCGGGTC GTTTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID:7 (TTGTTCGCCTATTGTGGGTCGGGTCGTTTT).

[0141] In some embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID:8 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGTGGGTCGGGTC GTTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID:9 (TTGTTCGCCTATTGTGGGTCGGGTCGTTT).

[0142] In some alternative embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 10 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGCGGGTCGGGTC G) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID: 11 (TTGTTCGCCTATTGCGGGTCGGGTCG).

[0143] In some alternative embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 12 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGCGGGTCGGGTC GATTTTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID: 13 (TTGTTCGCCTATTGCGGGTCGGGTCGATTTTT).

[0144] In some alternative embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 14 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGCGGGTCGGGTC GATTTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID: 15 (TTGTTCGCCTATTGCGGGTCGGGTCGATTTT).

[0145] In some alternative embodiments of the sensor system of the present disclosure, analyte of interest is vancomycin and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 16 (CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGCGGGTCGGGTC GATTT) or the split aptamer fragment pair comprised within the hydrogel composition has the sequences of SEQ ID:2 (CGACCGAGGGTACCGCAATAGTACTTA) and SEQ ID: 17 (TTGTTCGCCTATTGCGGGTCGGGTCGATTT).

[0146] In some embodiments of the sensor system of the present disclosure, analyte of interest is dopamine and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 18 (CGACGCCAGTTTGAAGGTTCGTTCGCAGGTGTGGAGTGACGTCG).

[0147] In some embodiments of the sensor system of the present disclosure, analyte of interest is TNF alpha and the allosteric DNA aptamer comprised within the hydrogel composition has the sequence of SEQ ID: 19 (GAGTGACGCAGCATGCTTAAGGGGGGGGCGGGTTAAGGGAGTGGGGAGGCA GCTGGTGTGGACACGG).

[0148] In all of the above embodiments, the assembly of the analyte binding complex by the allosteric molecules or fragments an upon their binding to the analyte of interest - regardless of where within the hydrogel composition - affects the 3-D network structure such that the processor circuitry of the sensor system can determine information on the presence of the analyte of interest based on sensor data being indicative of the of the alteration of the spatial dimensions of the 3-D network structure, e.g., based on sensor data being indicative of the surface plasmon resonance wavelength shift or sensor data being indicative of a shift in resonance frequency or energy dissipation of a quartz crystal.

[0149] Importantly, the sensor systems of the present disclosure may further be designed such that binding of the allosteric molecules or allosteric fragments thereof to the analyte of interest within the hydrogel composition is reversible, at least at body temperature (i.e. at about 37°C), such that the analyte binding complex disassembles upon separation of the analyte of interest from the allosteric molecules or allosteric fragments thereof and the conformational change is reversed. This reversibility is particularly advantageous as chemical sensor regeneration is not required. This is in contrast to known sensors utilizing aptamer functionalized hydrogel compositions, which require separate sensor regeneration in non-physiological solutions (such as e.g. in strong sodium hydroxide solutions) to reverse analyte binding. The reversibility of analyte binding, therefore, makes the sensor system of the present invention suitable for continuous, real-time, in situ sensing of an analyte of interest, addressing a critical challenge when sensing analytes such as drug concentrations, especially within complex tissue environments such as the brain.

[0150] It will be understood that, to ensure reversibility of the binding of the analyte of interest by the allosteric molecules used in the systems and methods of present invention, the exemplary characteristics listed in Table 1 below need to be considered and adjusted such that suitable reversibility is achieved:

[0151] Table 1

[0152] Table 1 (continued)

[0153] As such, if the alteration of the spatial dimensions of the 3-D network structure of the hydrogel composition upon assembly of the analyte binding complex is a collapse or contraction, release of the analyte, i.e. disassembly of the analyte binding complex, leads to an expansion or relaxation of the hydrogel composition.

[0154] As such, the specific hydrogel composition of the sensor system of the present disclosure is molecularly responsive to the analyte of interest and acts as a biosensor transducer, preferably as molecularly responsive biosensor transducer.

[0155] As already indicated above, the relied upon assembly and disassembly of the analytebinding complex is based on highly specific but low-affinity interactions with fast dissociation kinetics between the allosteric molecules or allosteric fragments thereof. In light of these features, the sensor probes of the present system regenerate passively through release of the analyte throughout and, as a result, do not need to undergo a specific probe regeneration process. This makes the sensor systems of the present disclosure particularly useful for continuous biosensing and, ultimately, for implantation.

[0156] In the case where DNA aptamers are used, and active regeneration is desired, the sensor probe may be regenerated by melting the aptamer DNA at high temperature, high salt concentration or pH. Of course, active regeneration cannot be performed on an implanted sensor probe in situ, which, of course, is one of the hurdles known sensors probes face when considered for implantation.

[0157] Accordingly, in embodiments of the present disclosure, the sensor probe is to provide sensor data acquired in situ, preferably the sensor probe is a sensor probe to be inserted into a bodily fluid, tissue or organ. The particularly small dimensions of the sensor probe are highly advantageous for insertion into a tissue or organ as the trauma caused thereby will be minimal. Typically, the sensor probe is to be implanted. Given that all components of the hydrogel are covalently attached within the hydrogel composition and / or the metal-coated surface of the sensor system and that all components of the system may be contained within the hydrogel composition as well as that the sensor probe is self-regenerating, makes the here-described sensor system particularly suitable for implantation. In addition, as the metal surface of the sensor probe itself is coated with the hydrogel composition, preferably entirely, it may also be contained within the hydrogel composition. Thereby, the metal surface is protected from non-specific protein adsorption though the hydrogel composition. As such, in the present sensor system, the hydrogel composition advantageously can prevent or significantly reduce fouling of the sensor probe (i.e. exert and anti-fouling effect), which extends the sensor system's durability and, thereby, longevity.

[0158] Typically, the surface plasmon resonance wavelength, e.g., the measurement being indicative of the surface plasmon resonance wavelength, or the resonance frequency and / or energy dissipation of the quartz crystal, may be determined at two different points in time. If the respective surface plasmon resonance wavelengths change between the first and second point in time, a surface plasmon resonance wavelength shift occurs. Similarly, if the resonance frequency and / or energy dissipation of the quartz crystal change between the first and second point in time, a shift in resonance frequency and / or energy dissipation occurs. As both measurements are included in the sensor data, the sensor data is thus also indicative of the respective shift between the first and second point in time. In many cases, the surface plasmon resonance wavelength shift or the shift in resonance frequency / energy dissipation may be determined relative to a reference measurement. For example, the reference measurement being indicative of the surface plasmon resonance wavelength, resonance frequency or energy dissipation can be measured at a reference time (e.g., towards the beginning of a continuous or quasi-continuous measurement), and any further measurement may represent a shift relative to the reference measurement. The surface plasmon resonance wavelength shift is the difference between two surface plasmon resonance wavelengths, e.g., between a reference surface plasmon resonance wavelength represented by the reference measurement and a surface plasmon resonance wavelength represented by a subsequent measurement, or two measurements at two different points in time. As pointed out above, this shift does not necessarily need to be expressed in terms of a wavelength or frequency. In many implementations, it can be expressed in terms of the shift in angle of incidence, in terms of a shift in the emission wavelength of the light source, in terms of a wavelength or frequency at which the dip in light intensity of the reflected light occurs, or in terms of a dimensionless number representing a change in the emitted light or reflected light between the two measurements. The emission wavelength of the incident light and the angle of incidence are directly related to the surface plasmon resonance wavelength, e.g., proportional to it. The exact relationship is governed by material properties and the dielectric environment. Similarly, the shift in resonance frequency or energy dissipation is the difference between two resonance frequencies or energy dissipation levels, e.g., between two different measurements, which may include the reference measurement.

[0159] While the present disclosure refers to a surface plasmon resonance wavelength and a surface plasmon resonance wavelength shift, the change in the refractive index can be observed by varying the angle of incidence, by varying the emission wavelength, or by observing the dip in reflected light intensity when the light source is a broadband light source. In other words, sensor data that is based on varying the angle of incidence, based on varying the emission wavelength, and / or representing a dip in reflected light intensity in response to a broadband light source may be indicative of the surface plasmon resonance wavelength shift. As a result, the sensor data may be indicative of an angle of incidence and / or an emission wavelength at which surface plasmon resonance occurs, with the angle of incidence (and changes thereof) and emission wavelength also being indicative of the surface plasmon resonance wavelength and surface plasmon resonance wavelength shift. Both are indicative of the same effect, i.e., the refractive index of the dielectric.

[0160] In some examples of the present disclosure, not only the shift in surface plasmon resonance wavelength, but also the effect that the alteration of the 3-D network structure of the hydrogel has on the waveguide may be used to detect the analyte of interest. For example, as shown in Fig. 10, column “MP-SPR”, properties, Fig. 11 F and Fig. 14 B, depending on the concentration of the analyte of interest, the refractive index changes, leading to changes in the reflectivity that affect the waveguide. Thus, the angle of incidence, at which the waveguide pattern occurs, shifts depending on the concentration of the analyte of interest. While this shift is often smaller than the shift exhibited by the surface plasmon resonance wavelength, it is also influenced by changes in the 3-D network structure that occur further than 200 nm away from the metal-coated surface (e.g., up to 10 pm), in contrast to the surface plasmon resonance wavelength shift. Thus, the presence and / or concentration of the analyte of interest may be determined based on a shift of the angle of incidence (or range of angles of incidence), at which the waveguide pattern occurs.

[0161] In various examples of the present disclosure, the measurement being indicative of the surface plasmon resonance wavelength may be measured using an optical sensor. In general, as shown in Fig. 1 , the surface plasmon resonance sensor 1 comprises at least four components: a light source 2a for emitting the incident light 3, an optical sensor 2b for sensing the reflected light 4, a glass component (such as a glass fiber core 5), and the metal-coated surface (such as a gold film 6) that is arranged (e.g., deposited, sputtered) on the glass component 5. In such a sensor, the light source 2a is used to emit a (P-polarized) beam of light 3 through the glass component 5 towards the metal-coated surface 6. The glass component 5, which is often a prism or, in case of a fiber-optic SPR sensor, an optical fiber 5, is used to increase a wavenumber and momentum of the beam of light 3, in order to excite the evanescent wave at the boundary between the glass component 5 and the metal coating 6 that is being used to cause the surface plasmons. Reflections 4 of the light beam are sensed by the optical sensor 2b, which converts the sensed reflections 4 into the sensor signal. In the example shown in Fig. 1 , processor circuitry 8 is arranged separate from the surface plasmon resonance sensor 1. The surface plasmon resonance sensor 1 may comprise communication circuitry 2d, which is coupled with the light source 2a and the optical sensor 2b, to communicate with the processor circuitry 8, e.g., wirelessly or via a cable 7. Particularly, but not exclusively, if wireless communication is used, the surface plasmon resonance sensor 1 may further include a battery 2e. Alternatively, if communication is conducted via cable 7, the cable may be used to power the surface plasmon resonance sensor 1. In some examples, the processor circuitry 8 may be included in the surface plasmon resonance sensor 1 . In this case, the communication circuitry 2d may be used to provide the output of the processor circuitry 8 to a separate device and / or to control the processor circuitry. Alternatively, the surface plasmon resonance sensor 1 may operate autonomously, with sensor data being collected by reading out a memory of the processor circuitry 8. In this case, the communication circuitry 2d is optional. In the example of Fig. 1 , the light source 2a, the optical sensor 2b, an optical coupler (for coupling the light 3 emitted by the light source 2a into the optical fiber 5 and for coupling the reflected light 4 into the optical sensor 2b), the optional communication circuitry 2d and the optional battery 2e are arranged in a housing together with a fiber coupling for the optical fiber 5.

[0162] In the proposed concept, the measurement representing the surface plasmon resonance wavelength shift may be determined and used to determine presence of the analyte of interest. In general, surface plasmon resonance has the effect that a larger amount of the energy contained in the light beam at the surface plasmon resonance wavelength is transferred into the surface plasmons, which decreases the intensity of the reflections 4. Thus, compared to other emission wavelengths, the reflections 4 are less intense when resonance occurs (i.e. , a dip in intensity compared to other emission wavelengths), which can be observed by the optical sensor 2b. The emission wavelength, at which the dip in intensity compared to other, adjacent emission wavelengths occurs, is thus indicative of the surface plasmon resonance wavelength (but not the same). Thus, to determine a measurement representing the surface plasmon resonance wavelength, the emission wavelength at which the dip occurs may be determined.

[0163] The emission wavelength, at which the dip in intensity occurs, can be determined using different measurement techniques. In some examples, the surface plasmon resonance sensor 1 may comprise a broadband light source as light source 2a (i.e., a light source emitting light 3 over a range of at least 100 nm (or at least 200 nm, or at least 300 nm) simultaneously, e.g., between an emission wavelength at least 350 nm (or at least 360 nm, or at least 400 nm, or at least 500 nm) as lower bound and an emission wavelength at most 850 nm (or at most 800 nm, or at most 750 nm, or at most 700 nm) as upper bound of the range. For example, the broadband light source may be a light source for emitting white light. For example, the broadband light source may be light-emitting diode (LED)-based light source.

[0164] In the case of a broadband light source, the optical sensor 2b may be an optical sensor 2b that is capable of differentiating between light having different wavelengths, such as an RGB (Red Green Blue) sensor or a hyperspectral sensor. For example, the optical sensor 2b may comprise a plurality of color bandpass filters, such as a plurality of color filters in a Bayer configuration, to sense light in different wavelength bands using different pixels of the optical sensor 2b. In the present case, the wavelength range of interest, i.e., the emission wavelength range, in which the surface plasmon resonance is likely to occur, is usually known a priori. Therefore, the color filters being used can be selected based on the wavelength range of interest.

[0165] In more general terms, the sensing spectrum of the optical sensor 2b and the emission spectrum of the light source 2a may be matched to each other and to the wavelength range of interest of the reflected light beam 4. For example, the broadband light source may have an emission spectrum (with a bandwidth of at least 100 nm, as it is a broadband light source) that intersects with the sensing spectrum of the optical sensor 2b (which may also be at least 100 nm, to cover a range of wavelengths). In other words, the surface plasmon resonance sensor 1 may comprise a light source 2a having an emission spectrum across a first bandwidth of at least 100 nm, and a sensor 2b for detecting an intensity of reflected light 4 having a sensing spectrum across a second bandwidth of at least 100 nm, with the first and the second bandwidths overlapping. Both the emission spectrum and the sensing spectrum may include the expected wavelength range of interest. The emission wavelength sensed by the optical sensor 2b, at which the intensity of the reflection 4 dips (e.g., at which the intensity is at least 5% lower (or at least 10% lower, or at least 20% lower) than the intensity of reflections 4 at adjacent wavelengths), may by indicative of the surface plasmon resonance wavelength. The dip in the intensity of the reflection 4, in terms of emission wavelength, may be represented by the sensor data.

[0166] Alternatively, the light source 2a may be a tunable light source for emitting light 3 having a (single) wavelength at a time, e.g., an LED-based light source or a Laserbased light source. For example, the tunable light source may be controlled by the processor circuitry 8, with the processor circuitry 8 being configured to control the wavelength of the emitted light 3 and the activation and deactivation of the light source 2a. The tunable light source may be configured to sequentially output light at different emission wavelengths of the wavelength range of interest. In other words, the surface plasmon resonance sensor 1 may comprise a tunable light source for sequentially emitting light 3 at different wavelengths of a wavelength range of interest. For example, the tunable light source may be configured to sweep the wavelength range of interest, by sequentially outputting light 3 at different emission wavelengths of the wavelength range of interest. For example, the different wavelengths being emitted by the tunable light source may differ by at least one tenth of a nanometer, or at least two tenths of a nanometer, or at least five tenths of a nanometer. They may cover the entire wavelength range of interest. For example, the wavelength range of interest may have a bandwidth of at most 20 nm (or at most 10 nm, or at most 8 nm, or at most 5 nm).

[0167] In the case of the tunable light source, as light of only one wavelength is emitted at a time, the optical sensor 2b does not have to discriminate between light having different emission wavelengths. Thus, the optical sensor 2b may be provided without a color filter. The optical sensor 2b may be configured to sense an intensity of the reflected light 4, with the reflected light 4 having the wavelength currently being emitted by the tunable light source. Thus, the surface plasmon resonance sensor 1 may comprise a sensor 2b for detecting a change in intensity of reflected light 4 when a probe (e.g., the glass fiber 5) of the surface plasmon resonance sensor 1 is illuminated by the tunable light source at the different wavelengths of the wavelength range of interest. The wavelength emitted by the tunable light source, at which the intensity of the reflections 4 sensed by the optical sensor 2b dips (e.g., at which the intensity at least 5% lower (or at least 10% lower, or at least 20% lower) than the intensity of reflections 4 at adjacent wavelengths), may be indicative of the surface plasmon resonance wavelength. The optical sensor 2b is thus used to detect a change in intensity of reflected light 4 when the tunable light source emits light 3 at the different wavelengths of the wavelength range of interest. Again, the dip in the intensity of the reflection 4, and the time at which the dip in intensity has occurred, may be represented by the sensor data. Knowing the wavelength being emitted by the tunable light source at the time at which the dip in intensity has occurred, the measurement being indicative of the surface plasmon resonance wavelength can be determined.

[0168] As a further option, the angle of incidence may be varied. Again, a tunable light source or a single-wavelength light source (such as a Laser or a single-wavelength LED light source) may be used as light source 2a. The processor circuitry 8 may be configured to control the angle, at which the light 3 is emitted towards the coated surface 6 (and, in case of fiber-optic surface plasmon reference, reflected by the walls of the glass fiber 5), to vary the angle of incidence of the emitted light 3. Thus, the surface plasmon resonance sensor 1 may comprise a light source 2a for emitting light 3 at a wavelength and means for varying an angle of incidence of the emitted light 3. For example, the processor circuitry 8 may be configured to control a reflective element, such as a movable mirror, or a light-guiding element, such as a movable prism, to vary the angle of incidence of the emitted light 3 (with the movable mirror and movable prisms being means for varying the angle of incidence of the emitted light 3). Alternatively, or additionally, the light source 2a may be capable of varying the angle of incidence, with the processor circuitry 8 controlling the light source 2a to vary the angle of incidence. In this case, the light source 2a may include the means for varying the angle of incidence. The processor circuitry 8 may be configured to vary the angle of incidence across an angle of incidence range of interest, e.g., by sweeping the angle of incidence across the angle of incidence range of interest. This may be done by varying the angle of incidence by one twentieth of a degree (or one tenth of a degree, or one fifth of a degree, or half of a degree, or a degree) between measurements.

[0169] Similar to the approach used with a tunable light source being used to vary the emission wavelength, the optical sensor does not have to discriminate between light having different emission wavelengths. Thus, the optical sensor may be provided without a color filter. The optical sensor may be configured to sense an intensity of the reflected light 4, with the reflected light 4 being based on the angle of incidence being used to illuminate the metal coating 6. Thus, the surface plasmon resonance sensor 1 may comprise a sensor 2b for detecting a change in intensity of reflected light 4 when a probe of the surface plasmon resonance sensor 1 is illuminated by the light source 2a at different angles of incidence. The angle of incidence, at which the intensity of the reflections 4 sensed by the optical sensor dips (e.g., at which the intensity at least 5% lower (or at least 10% lower, or at least 20% lower) than the intensity of reflections 4 at adjacent angles of incidence), may be indicative of the surface plasmon resonance wavelength. The optical sensor is thus used to detect a change in intensity of reflected light 4 when the angle of incidence is varied across the angle of incidence range of interest. Again, the dip in the intensity of the reflection

[0170] 4, and the time at which the dip in intensity has occurred, may be represented by the sensor data. Knowing the angle of incidence used at the time at which the dip in intensity has occurred, the measurement being indicative of the surface plasmon resonance wavelength can be determined.

[0171] In summary, both the emission wavelength at which resonance occurs and the angle of incidence at which resonance occurs may be measures being indicative of the surface plasmon resonance wavelength. A shift (between two measurements) in the emission wavelengths at which resonance occurs, and a shift (between two measurements) in the angle of incidence at which resonance occurs, may be measures being indicative of the surface plasmon resonance wavelength shift. In the following, if the surface plasmon resonance wavelength shift is being compared to a threshold, it may be one of the measures being indicative of the surface plasmon resonance wavelength shift, i.e. , the shift in emission wavelength or shift in angle of incidence, that is being compared to the threshold.

[0172] The surface plasmon resonance sensor 1 comprises the glass component 5, on which the metal coating 6 is deposited. To make the surface plasmon resonance sensor 1 implantable, a compact construction of the surface plasmon resonance sensor 1 is desired. This can be achieved by using an optical glass fiber 5 as glass component

[0173] 5. In other words, the surface plasmon resonance sensor 1 may be a fiber-optic surface plasmon resonance sensor. The metal coating 6 can be deposited onto the optical glass fiber 5, with the metal-coated surface 6 being a surface of the optical glass fiber 5 of the surface plasmon resonance sensor 1 . For improved implantability and sensitivity, the glass fiber 5 may have a diameter of at most 500 pm (or at most 400 pm, or at most 300 pm), e.g., 300 or 200 pm. Moreover, to further improve implantability in tight spaces, the glass fiber 5 (e.g., an exposed portion extending from a casing of the surface plasmon resonance sensor) may have an exposed length of at most 10 mm (or at most 8 mm, or at most 5 mm, or at most 3 mm), e.g., 3 mm.

[0174] The metal coating 6 being used may be adapted to the use with the Hydrogel. In general, the metal-coated surface 6 comprises at least one of gold (Au), silver (Ag), copper (Cu), aluminum (Al), chromium (Cr), platinum (Pt), iridium (Ir), a titanium oxide (TiO2), a tantalum oxide, (Ta2Os), zinc oxide (ZnO), or silicon dioxide (SiC ). Preferably, the metal-coated surface may be a gold-coated surface, or may comprise multiple layers (i.e., form a multilayer structure) having a gold surface as exposed surface. For example, the metal multilayer structure may comprise two or more of gold, silver, copper and aluminum (e.g., with gold as exposed surface). For example, an Au-AI-Au multilayer or an Ag-Au bilayer may be used as metal coating 6. In some examples, the metal coating may be or comprise a metal oxide coating such as a Ta2Os or a TiO2 coating, e.g., a Ta2Os-coated surface or a TiC -coated surface.

[0175] In one preferred embodiment, the metal coating comprises a bi-metallic, dual-layer silver-gold (Ag-Au) coating, whereas a first layer on the glass component 5 is a silver (Ag) layer, which is again coated with a second layer, namely a gold (Au) layer. In this embodiment, bulk sensitivity of a fiber optic surface plasma resonance sensor (FO- SPR) is increased by 1 .2 to 1 .6 fold compared to a corresponding sensor coated with a single layer gold (Au) coating (see Fig. 16). While the silver layer produces a stronger plasmon resonance wavelength shift, the outer gold layer protects the silver layer from rapid oxidation. To further harness the enhanced silver-mediated plasmon resonance wavelength shift, the ratio of the dual-layer components may be altered. A 60:40 Ag / Au coating increases bulk sensitivity of the sensor by about 1.3 fold, while a 80:20 Ag / Au coating increases bulk sensitivity of the sensor by about 1 .5 fold. In a sensor system of the present invention, a 60:40 Ag / Au coating increases sensitivity of the sensor by about 2-3 fold, while a 80:20 Ag / Au coating increases sensitivity of the sensor by about 4-5 fold.

[0176] For example, the metal coating 6 may have a thickness of at least 10 nm (or at lest 20 nm, or at least 30 nm, or at least 40 nm) and / or at most 80 nm (or at most 70 nm, or at most 60 nm, or at most 50 nm). Such a metal coating may also be used when using a quartz crystal-based sensor. In general, a crystal-based sensor comprises the quartz crystal and two metal electrodes, between which the quartz crystal is sandwiched. One of the electrodes may be the metal surface of the sensor. An alternating current is applied to the quartz crystal via its electrodes, causing the quartz crystal to oscillate at its resonance frequency. This frequency, as well as an energy dissipation of the quartz crystal, is then measured to determine the sensor signal. Thus, the sensor may comprise a quartz crystal, two electrodes, an alternating current source and measurement circuitry for determining the resonance frequency and energy dissipation of the quartz crystal (not shown).

[0177] The processor circuitry 8 uses the sensor data to determine information on a presence of the analyte of interest. For example, the information on the presence of the analyte of interest may indicate whether the analyte is present at the sensor, e.g., bound by the aptamer or aptamer fragments of the hydrogel composition. For example, if the sensor data indicates a surface plasmon resonance wavelength shift relative to the reference measurement (e.g., a shift in emission wavelength or angle of incidence at which resonance occurs), the refractive index of the dielectric medium may have changed, indicating the analyte of interest being bound by the hydrogel. For example, if the measurement(s) representing the surface plasmon resonance wavelength shift (e.g., the shift in emission wavelength or angle of incidence at which resonance occurs), is greater than a threshold, e.g., greater than two tenths of a nanometer (or greater than five tenths of a nanometer) or greater than one tenth of a degree (or greater than one fifth of a degree, or greater than half a degree), presence of the analyte of interest may be determined. In other words, the processor circuitry 8 may be configured to determine presence of the analyte of interest if the surface plasmon resonance wavelength shift, e.g., as represented by the shift in emission wavelength or angle of incidence at which resonance occurs, is greater than the threshold. Accordingly, the processor circuitry 8 may be configured to determine absence or uncertainty about the presence of the analyte of interest if the surface plasmon resonance wavelength shift, e.g., as represented by the shift in emission wavelength or angle of incidence at which resonance occurs, is at most the threshold. Similar determinations may be performed by the processor circuitry 8 based on the shift in resonance frequency and / or energy dissipation of the quartz crystal of the quartz crystal-based sensor. The amplitude of the surface plasmon resonance wavelength shift, e.g., as represented by the shift in emission wavelength or angle of incidence at which resonance occurs, correlates positively with the amount of analyte of interest being bound by the hydrogel. In other words, the more of the analyte of interest is bound by the hydrogel, the greater the surface plasmon resonance wavelength shift (e.g., the shift in emission wavelength or angle of incidence). Up to a certain concentration of analyte (i.e., as long as there is ample opportunity for the analyte to bind to the hydrogel), the amplitude of the surface plasmon resonance wavelength shift (e.g., the shift in emission wavelength or angle of incidence) can thus be used to estimate the concentration of the analyte in the medium. For example, the processor circuitry 8 may be configured to determine information on the estimated concentration of the analyte of interest based on the sensor data, e.g., based on the amplitude of the surface plasmon resonance wavelength shift (e.g., the shift in emission wavelength or angle of incidence) indicated by the sensor data. For example, the information on the estimated concentration may indicate, below a threshold for the wavelength shift, the estimated concentration of the analyte of interest. Above the threshold, the information on the estimated concentration may indicate that the estimated concentration surpasses a measurement range of the surface plasmon resonance sensor. For example, the threshold may be based on the capability of the hydrogel of binding additional analyte of interest. Up to the threshold, the estimation of the concentration may be based on information about a conversion ratio between known concentrations of the analyte of interest and corresponding amplitudes of the surface plasmon resonance wavelength shift (e.g., corresponding amplitudes of the shift in emission wavelength or angle of incidence). Similar determinations may be performed by the processor circuitry 8 based on the shift in resonance frequency and / or energy dissipation of the quartz crystal of the quartz crystal-based sensor, which also correlated with the amount of analyte of interest being bound by the hydrogel.

[0178] In various examples, the goal may be to perform measurements in vitro or in vivo over longer periods of time, e.g. over multiple hours, multiple days, or even multiple weeks, such as over a period of 1 to 60 days, preferably 1 to 30 days. To evaluate the surface plasmon resonance wavelength and corresponding surface plasmon resonance wavelength shift (e.g., emission wavelength or angle of incidence and shift thereof), or shift in resonance frequency and / or energy dissipation of the quartz crystal, relative to the reference measurement over time, the measurement and determination of the information on the presence (and / or concentration) of the analyte of interest may be performed quasi-continuously (e.g., according to a sample interval of the surface plasmon reference sensor, e.g., of the optical sensor, or according to a time interval required for sweeping the wavelength range of interest or for sweeping the angle of incidence range of interest) or according to a pre-defined schedule. In other words, the processor circuitry 8 may determine the information on the presence of the analyte of interest quasi-continuously and / or according to a pre-defined schedule. For example, to achieve a tradeoff between a frequency of the determination of the information on the presence and / or concentration of the analyte of interest and an energy use of the sensor system, the information on the presence and / or concentration of the analyte of interest may be determined with a frequency of at most every ten seconds (or at most every minute, or at most every five minutes, or at most every 15 minutes, or at most every hour). In other words, the pre-defined schedule may define the frequency of determining the information on the presence and / or concentration of the analyte of interest. Measurements may be performed according to the pre-defined schedule, e.g., at most every ten seconds (or at most every minute, or at most every five minutes, or at most every 15 minutes, or at most every hour).

[0179] In the present sensor system, measuring the measurement being indicative of the surface plasmon resonance wavelength may be performed by the surface plasmon resonance sensor to generate the sensor data, while the evaluation of the sensor data is performed by the processor circuitry 8. Similarly, measuring the measurement being indicative of the resonance frequency and / or energy dissipation of the quartz crystal may be performed by the quartz crystal-based sensor, while the evaluation of the sensor data is performed by the processor circuitry 8. Thus, the processor circuitry 8 is coupled with the sensor, such as the surface plasmon resonance sensor 1 (e.g., with the optical sensor 2b and / or with the light source 2a, e.g., via the communication circuitry 2d) or the quartz crystal-based sensor.

[0180] The sensor may be part of a sensor probe (shown as components 1 -6, 9 in Fig. 1 ), with the sensor probe comprising the sensor 1 with the metal-coated surface 6 and the hydrogel composition 9. In some examples, the sensor probe may further comprise the processor circuitry 8. In other words, the processor circuitry 8 may be part of the sensor, such as the surface plasmon resonance sensor 1 , or be co-located with the sensor 1 inside the probe. Alternatively, the processor circuitry 8 may be separate from the sensor probe, e.g., part of a computer or mobile device being communicatively coupled to the sensor probe. In both cases, the sensor probe may further comprise communication circuitry 2d for providing the respective information, e.g., for providing the information on the presence or concentration of the analyte of interest to a separate device. For example, in case the processor circuitry 8 is colocated with the sensor 1 inside the probe, the communication circuitry 2d may be used to provide, by the processor circuitry 8, the information on a presence of the analyte of interest to a separate device. In this case, the communication circuitry 2d may be coupled locally inside the sensor probe with the processor circuitry 8. In case the processor circuitry 8 is separate from the sensor probe, the communication circuitry 2d may be used to provide the sensor data to the processor circuitry 8 being separate from the sensor probe. In this case, the communication circuitry may be coupled locally inside the sensor probe with the sensor 1 , e.g., with the light source 2a and with the optical sensor 2b.

[0181] For example, the optical sensor 2b may be one of a CCD (Charge-Coupled Device) optical sensor and a CMOS optical sensor. CCD sensors and CMOS sensors both comprise an array of pixel sensors. In CCD sensors, the pixels convert photons into electrical charges, which are accumulated in so-called bucket. At read-out time, the accumulated electrical charges are converted, via a voltage, into a digital representation of the amount electrical charges, representing the intensity of light hitting the pixel. In CMOS sensors, each pixel contains a photodetector and an active amplifier. When light strikes the photodetector, it generates an electrical charge that is amplified and then converted into a digital value, creating a digital pixel in the resulting image.

[0182] The processor circuitry 8 may correspond to or comprise one or more of processor, a microcontroller, logic circuitry, a Digital Signal Processor, and a Field-Programmable Gate Array (FPGA) etc. In general, the functionality of the processor circuitry 8 may be provided by the processor circuitry 8 executing machine-readable instructions. In other words, the functionality of the processor circuitry 8 may be defined by software being executed by the processor circuitry. For example, the sensor system, e.g., the processor circuitry 8, may comprise a memory comprising the machine-readable instructions. For example, the memory may be further used to store information on a reference surface plasmon resonance wavelength (e.g., information on a reference emission wavelength or information on a reference angle of incidence) and / or information on a conversion ratio between the measurement representing the surface plasmon resonance wavelength shift and a corresponding concentration of the analyte. For example, the memory may be further used to store and / or the information on the presence and / or concentration of the analyte of interest sensor data permanently or temporarily.

[0183] The communication circuitry 2d may comprise circuitry for communicating with a remote entity. For example, the communication circuitry 2d may comprise transmitter circuitry or transceiver circuitry for transmitting information to the remote entity. For example, the communication circuitry may be configured to communicate wirelessly (e.g., include a wireless transmitter or wireless transmitter) or via one or more wires (e.g., include a transmitter for wired communication.

[0184] As will be understood from the above, in the second aspect, the present disclosure also relates to a method for sensing an analyte of interest, the method comprising the following steps:

[0185] (a) providing a sensor, such as a surface plasmon resonance sensor or a quartz crystal-based sensor, comprising a sensor probe with a metal-coated surface and a hydrogel composition contacting the metal-coated surface, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel;

[0186] (b) providing, by the sensor, sensor data being indicative of the alteration of the spatial dimensions of the 3-D network structure, e.g., of a surface plasmon resonance wavelength shift or shift in resonance frequency or energy dissipation being contributed to by the alteration of the spatial dimensions; and

[0187] (c) determining information on a presence of the analyte of interest based on the sensor data. The above method, generally employing the sensor system of the first aspect, may be used in a laboratory setting such that the method may be described as an ex vivo method for sensing the analyte of interest in an aqueous sample. This is particularly useful for sensing the analyte of interest in an aqueous sample that is a bodily fluid sample or a tissue sample previously-obtained from a patient. In instances where such a patient is undergoing treatment with a pharmaceutically active agent, the method may for example be advantageously employed as a non-invasive companion diagnostic method to determine relevant information on the presence of the relevant pharmaceutical agent or its metabolites (i.e. here the analyte of interest) based on the sensor data in such bodily fluids or tissues. Such information may, for example, allow a clinician to quickly and reliably assess whether sufficient pharmaceutically active agent reaches the target tissue or organ for the respective treatment to be effective. Of course, such information may ultimately provide the clinician with an opportunity to consider varying the relevant dosage or dosage regime of the active agent to increase chances of effective and ultimately successful treatment.

[0188] Alternatively, step (c) providing the sensor data by the sensor, may comprise inserting a probe of the sensor into a bodily fluid or tissue for sensing the analyte of interest in situ. Preferably, the probe of the sensor is implanted for continuous sensing the analyte of interest in situ.

[0189] As will be appreciated from the above, such continuous sensing of the analyte of interest may provide much more detailed information regarding the efficacy of a particular dosage regime of a pharmaceutically active agent in the target tissue or organ as the sensing method can provide information on the presence of the pharmaceutically active agent over a defined time period, for example over the entire time period spanning between administration of a first dose of the pharmaceutically active agent and a subsequent dose. This information may allow resolving effects on drug concentrations in the target tissue or organ, for example caused due to parallel intake of other drugs or even food as well as of circadian changes in metabolism etc.

[0190] In particularly useful embodiments, the method of the second aspect includes continuous sensing over a period of 1 to 60 days, preferably 1 to 30 days. In the third aspect, the present disclosure relates to use of a sensor system according to the first aspect for sensing an analyte of interest in an aqueous sample ex vivo. Again, the aqueous sample is typically a previously-obtained bodily fluid sample or a previously-obtained tissue sample.

[0191] The fourth aspect of the present disclosure concerns a hydrogel composition for use in diagnostic sensing of an analyte of interest, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or allosteric fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest.

[0192] In such a hydrogel composition, typically, at least some of the hydrophilic polymers of the 3-D network structure are connectable of to a metal-coated surface of a sensor, such as a surface plasmon resonance sensor or quartz crystal-based sensor.

[0193] The invention is further described by the following non-limiting Examples.

[0194] 1. Development and establishment of sensor system

[0195] Materials and Methods

[0196] Reagents

[0197] 8-arm PEG norbornene (MW: 10 kDa) was purchased from Creative PEGWorks. DNA aptamers of SEQ ID NO:1 (Full Aptamer), SEQ ID NO:2 (Split Aptamer Fragment 1 ) and SEQ ID NO:3 5’ (Split Aptamer Fragment 2) were obtained from Integrated DNA Technologies. The aptamers were modified with 5' thiol modifier C6 S-S and / or 3' thiol modifier C3 S-S for subsequent conjugation to the hydrogel matrix. All buffer reagents were purchased from Thermo Fisher Scientific as Dnase / Rnase-free stock solutions, including UltraPure™ distilled water, phosphate-buffered saline (PBS, 10X, pH 7.4), and magnesium chloride (1 M). HPLC grade ethanol and 1 ,4-dithiothreitol 99% (DTT) were acquired from Carl Roth GmbH. 1 -Dodecanthiol (DDT), PEG-dithiol (MW: 3.4 kDa), lithium phenyl(2,4,6-trimethylbenzoyl)phosphinate (LAP), Tris(2- carboxyethyl)phosphine hydrochloride (TCEP), vancomycin hydrochloride, and lyophilized rat plasma were obtained from Sigma-Aldrich. Hydrogel film preparation

[0198] The custom-made SPR gold sensor chip (dimensions: 20 x 12 x 0.5 mm) was purchased from LET Optomechanika Praha. The gold surface was functionalized by incubating the chip in a 50 mM solution of DTT in ethanol for 5 h, followed by a brief 5-min incubation in a 50 mM solution of DDT, also in ethanol. The surface was then thoroughly washed with water and ethanol, dried, and stored in a vacuum desiccator before further use. The hydrogel precursor solution was prepared by first mixing the respective amounts of DNA aptamers, PEG-dithiol, and TCEP (final concentration of 80 mM) in Dnase-free water. The mixture was incubated at 37°C for 2 h to reduce disulfide bonds to free thiol groups. Subsequently, 8-arm PEG norbornene and LAP (final concentration of 1 mM) were added to complete the pre-gel solution. A volume of 2 pL of this solution was drop-cast onto the treated gold surface, and a hydrophobic-treated glass slide was gently pressed onto the deposited solution. Finally, the sandwiched assembly was exposed to 365 nm UV light for 2 min to polymerize the pre-gel solution and form a thin hydrogel film.

[0199] Multi-parametric surface plasmon resonance (MP-SPR)

[0200] The MP-SPR measurement was conducted using a dual-wavelength MP-SPR Navi 400 Kontio (BioNavis) instrument. The SPR was monitored at wavelengths of 670 and 785 nm, across an angular range of 40 to 78 degrees, with the angle of the SPR minimum used to generate the sensogram. All data presented herein were reported with a wavelength of 670 nm. The experiment was initiated by measuring the baseline with phosphate-buffered saline (PBS) buffer containing 2 mM MgCL, maintained at a temperature of 37°C. A consistent flow rate of 50 pL / min was used throughout all the measurements, controlled by a peristaltic pump (Ismatec). The SPR gold sensor chip with the hydrogel film was inserted into the instrument, and the sample chamber was filled with the running buffer (PBS containing 2 mM MgCL). The system was allowed to equilibrate for at least 30 min to establish a stable baseline before the start of the experiment. Following the baseline stabilization, the analyte solutions were injected into the flow cell at predetermined concentrations, followed by a dissociation phase where the running buffer was injected to remove any unbound analytes. The collected data were fitted with Winspall 3.01 software to determine the hydrogel’s refractive index (n) and thickness (d). Quartz crystal microbalance-dissipation (QCM-D)

[0201] QCM-D measurements were conducted using a Q-Sense E4 instrument (Biolin Scientific) to monitor the kinetics of aptamer-functionalized hydrogel and analyte interactions. The changes in resonance frequency (Af) and energy dissipation (AD) of a quartz crystal were measured as a function of time. The quartz crystal was excited to generate the thickness shear mode at its fundamental resonance frequency of 5 MHz and at odd overtones (n: 1 , 3, 5, 7, 9, and 11 ). All measurements were performed at 37°C with a flow rate of 50 pL / min, controlled by a peristaltic pump (Ismatec). The gold-coated quartz crystal sensor (QSX 301 , Biolin Scientific) was functionalized as described previously, and all data presented herein were collected at the fifth overtone. The system was allowed to equilibrate in the running buffer (PBS containing 2 mM MgCl2) for at least 30 min to establish a stable baseline. The collected data were analyzed using the Qtools software (Biolin Scientific).

[0202] Experimental design

[0203] The current section is divided into three parts: (1 ) the selection of the hydrogel composition of the sensor system, (2) varieties of modification of the hydrogel 3-D network structure with allosteric molecules and / or allosteric fragments thereof (here specifically with aptamers) depending on the molecule and / or fragment configurations, and (3) functionalization of the gold surface of the surface plasmon resonance (SPR) sensor (see Fig. 9).

[0204] (1 ) The hydrogel composition was carefully designed to enable efficient incorporation of the aptamers and robust attachment of the hydrogel films to the gold surface. PEG-norbornene was selected as the base polymer due to its highly efficient thiol-ene click chemistry capabilities. The ‘ene’ (norbornene) group of PEG- norbornene can rapidly and reliably undergo a thiol-ene coupling reaction with thiol groups, either present in the crosslinking agents or on the gold surface, upon the addition of a photoinitiator (lithium-phenyl-2,4,6-trimethyl, LAP) followed by UV (365 nm) light activation (see Fig. 9(A)).

[0205] The 8-arm star-shaped PEG-norbornene architecture was specifically chosen over a linear polymer design due to its several advantages. Star polymers, characterized by a minimum of three macromolecular chains radiating from a central core, have garnered significant attention due to their unique topological structures and desirable physical / chemical attributes. For instance, the star-shaped architecture provides a higher number of reactive norbornene end-groups per molecule, facilitating enhanced efficiency and control over click chemistry crosslinking. This topology also allows for the incorporation of various functionalities and improved responsiveness to stimuli, stemming from the high density of functional groups. Moreover, star polymers exhibit reduced solution viscosity in comparison to linear counterparts of equivalent molecular weight, due to fewer arm entanglements, which is advantageous for producing thin, homogenous hydrogel films as it enhances processability and coatability.

[0206] The properties of the hydrogel were tuned by adjusting the molar ratio between PEG- norbornene (10 kDa) and dithiol cross linkers, denoted as the [ene] / [SH] ratio, while maintaining a constant PEG-norbornene concentration of 10 mM. At [ene] / [SH] ratios below 1 , the presence of free thiol groups in the hydrogel allowed direct covalent bonding to the gold surface. However, this approach required excessively high aptamer concentrations, leading to increased potential waste and reduced reproducibility due to the propensity of difunctional aptamers to form mono-tethered attachments to the hydrogel.

[0207] Consequently, [ene] / [SH] ratios exceeding 1 were determined to be optimal, as they provided free norbornene groups capable of forming covalent bonds with the thiolated gold surface. Initial investigations revealed that ratios below 2.5 yielded insufficient attachment strength on the gold surface, while ratios above 4.5 exhibited unacceptably low mechanical properties due to decreased crosslinking density. Therefore, [ene] / [SH] ratios of 3 and 4 were chosen to ensure robust gold attachment and appropriate gel mechanical characteristics for continuous flow applications.

[0208] (2) Aptamers were incorporated into the hydrogel system via thiol modifications, while the overall hydrogel structure was maintained by PEG-dithiol (3.4 kDa) cross linkers (see Fig. 9(B)). The proportion of PEG-dithiol to thiolated aptamers ranged from approximately 2 to 12, with aptamer concentrations varying between 1 and 3 mM, the upper limit being constrained by the available stock concentrations. Various aptamer configurations were investigated, encompassing monofunctionalized split aptamer pairs (MM), a combination of monofunctionalized and difunctionalized split aptamer pairs (MD), difunctionalized split aptamer pairs (DD), and the full original aptamer sequence with thiol modifications at both termini (F).

[0209] In previous work, the inventors reported a series of vancomycin-binding split aptamer pairs that can be utilized in a fully reversible monolayer-based assay format using surface plasmon resonance (SPR) detection. Building on this prior research, the P27 split aptamer pair was selected for the current study. This split aptamer design exhibited a high degree of reproducibility and an appreciable dissociation constant (KD) with the target analyte, allowing for a reversible and rapid response upon exposure to vancomycin at the physiologically relevant temperature of 37°C. The selfhybridization properties of the P27 split aptamer pair were also favorable, as they associate at room temperature but dissociate at 37°C, ensuring the required proximity and interaction with the analyte during the hydrogel polymerization process conducted at 25°C. This reversibility is particularly advantageous as it allows the sensor system of the present invention to be used in continuous, real-time, in situ sensing of an analyte of interest.

[0210] (3) The gold surface was functionalized through a series of thiolated molecule incubations (see Fig. 9(C)). First, the gold chip was incubated in a 50 mM solution of dithiothreitol (DTT) in ethanol for 5 h. This was followed by a brief 5-min incubation in a 50 mM solution of dodecanethiol (DDT), also in ethanol. This protocol was adapted from a previous study, which verified that the subsequent DDT incubation increases the amount of upright DTT configuration, thereby maximizing the number of free thiol groups available for reaction with the norbornene groups. Next, the hydrogel precursor (pre-gel) solution was drop-cast onto the treated gold surface. A glass slide was then gently pressed onto the deposited pre-gel solution, sandwiching it between the bottom gold surface and the top glass substrate. Finally, the sandwiched assembly was exposed to UV light for 2 min, initiating the polymerization and forming a thin hydrogel film that was covalently bonded to the gold surface.

[0211] Results

[0212] Characterization of aptamer-functionalized hydrogels depending on aptamer modifications

[0213] After forming the thin hydrogel layer onto gold sensor chips, two different surface sensing techniques were used to interrogate the mechanism and functionality of aptamer-functionalized hydrogels upon analyte introduction (sse Fig. 10). For optical characterization, multi-parametric surface plasmon resonance (MP-SPR), which utilizes a dual-wavelength approach scanning across a wide range of angles (0), was used to determine kinetics from the SPR response and the refractive index and thickness of the hydrogel film by analyzing the hydrogel optical waveguide pattern. This technique is well-suited for analyzing thin hydrogel films up to 10 pm in thickness, as the waveguide mode provides additional insights into the structure and dynamic conformation of the film.

[0214] To further complement the optical analysis, quartz crystal microbalance with dissipation monitoring (QCM-D) was employed. The QCM-D technique relies on the oscillation of a quartz crystal, which is sensitive to changes in water-coupled mass on the surface, as indicated by changes in the resonance frequency (A ) and the energy dissipation (AD). A decrease in A indicates an increase in mass due to bound molecules, while an increase in AD suggests greater energy loss caused by the viscoelastic nature of the hydrogel and coupled analyte. This enabled the comprehensive characterization of the viscoelastic properties of the hydrogel film during the aptamer-analyte interactions.

[0215] The split aptamer pair sequence used in this study, named P27, consisted of DNA segments Aptamer 1 (SEQI ID NO:2) and Aptamer 2 (SEQ ID NO:3), whereas the full original aptamer sequence consisted of Aptamer 1 directly connected to Aptamer 2 (SEQ ID NO:1 ) (see Fig. 11 (A)). The binding site for the target analyte, vancomycin, is estimated to be located in the middle of the aptamer sequence, ensuring that the binding pocket would not be disturbed when the aptamers were split. As briefly mentioned above, four different aptamer configurations were investigated, as schematically represented in Fig. 11 (B). These configurations involved different arrangements and tethering of the split aptamer pairs or the original aptamer.

[0216] For the initial hydrogel investigation, the [ene] / [SH] ratio was fixed at 3, corresponding to a concentration of 10 mM for the 8-arm PEG-norbornene and 12.3 mM for the PEG- dithiol cross linkers (see Fig. 11 (C)). The aptamer concentration was then maintained at a constant 1 mM across all hydrogel samples. The difference in crosslinking density between the fully dithiolated (DD1 -3) and monothiolated (MM1 -3) split aptamer designs was less than 8% (13.3 and 12.3 mM of dithiol linkers, respectively). By maintaining the [ene] / [SH] and aptamer concentrations at a constant value, the mechanical properties of the hydrogel matrix remained relatively consistent across the different aptamer configurations, this experimental design enabled the direct evaluation of how the varying aptamer immobilization strategies impacted the binding response and sensitivity towards the vancomycin target, without confounding factors related to substantial changes in the underlying hydrogel network structure and mechanics.

[0217] After preparing thin hydrogel films coated on MP-SPR gold sensor chips, a continuous flow of 50 pL / min was maintained for monitoring. The SPR was monitored across an angular range of 40 to 78 degrees, with the angle of the SPR minimum (6min) used to generate the sensogram. The extent of 0minshift (A0min), obtained by subtracting the baseline value, was presented to compare the response of aptamer-functionalized hydrogels to the analyte (see Fig. 11 (D)). After stabilization, 1 mM vancomycin was injected into the measurement chamber (point (A) at 3 min), then the association kinetics or binding response between aptamers and vancomycin over time were observed from the sensogram. The monofunctionalized MM1 -3, with split aptamer pairs attached to the hydrogel matrix via a single thiol group at one end, exhibited the smallest binding response compared to other aptamer configurations. The hybrid MD1 -3, incorporating a combination of monofunctionalized and difunctionalized split aptamer pairs, showed an improved binding response compared to the MM1 -3 but lower than the fully difunctionalized (DD1 -3) and the original (F1 -3) aptamer designs. These results suggest that difunctionalized aptamer configurations offer the best performance in terms of binding affinity and sensitivity towards vancomycin. Importantly, after buffer rinsing at 11 min (point (B)), all sensogram curves displayed fast dissociation kinetics, indicating a reversible interaction between the aptamers and vancomycin.

[0218] To assess the reversibility and reproducibility of the best-performing DD1 -3 and F1 -3 aptamer configurations, repetitive vancomycin injection experiments were conducted. Fig. 11 (E) reveals that F1 -3 exhibited partial irreversible binding and low reproducibility, retaining only 90% of the initial signal after the third binding cycle. Conversely, DD1 -3 demonstrated fully reversible binding and high reproducibility over repeated vancomycin injections. Notably, the DD1 -3 exhibited excellent temporal resolution, with both association and dissociation kinetics occurring within a rapid 3- min timeframe, highlighting the system’s potential for real-time monitoring applications. These findings suggest that the engineered split aptamer pairs with both terminal ends securely tethered to the 8-arm PEG-norbornene hydrogel represent the most suitable design for achieving the desired reversibility and consistency in target analyte binding performance.

[0219] For a more thorough examination of the mechanism, the hydrogel’s refractive index (n) and thickness (d) were determined by analyzing the experimental data of waveguide mode with optical models (see Fig. 11 (F)). The changes in the hydrogel’s properties induced by the presence of vancomycin were summarized in Fig. 11 (G). The shifts in SPR minimum angle (A0min) and refractive index (An) were determined by direct subtraction of the values before and after vancomycin introduction, while the change in thickness (Ac / ) was calculated as a percentage change using the formula

[0220] Among all the hydrogels investigated, an intriguing observation was made: upon vancomycin interaction with the aptamers within the hydrogel matrix, the 0minand An increased, while the Ac / decreased. This finding suggests that the aptamer-analyte binding induces a physical change or actuation by altering the hydrogel’s internal structure. The binding of vancomycin to the aptamers may cause them to adopt a more compact conformation, leading to a localized contraction of the hydrogel network. Consequently, this contraction results in a decrease in the overall thickness of the hydrogel film. The extent of this actuation is likely dependent on the specific aptamer configuration, with the magnitude of the response varying accordingly. For instance, the MM1 -3 configuration exhibited minimal changes in both thickness and refractive index upon vancomycin binding, while the MD1 -3 displayed a slightly enhanced response compared to MM1 -3. Remarkably, the DD1 -3 configuration showed the most significant changes in SPR minimum angle and thickness. These findings emphasize the crucial role of the aptamer immobilization configuration in controlling the hydrogel’s responsiveness to the target analyte, with the difunctionalized split aptamer pairs (DD1 -3) demonstrating the most promising performance in terms of hydrogel actuation, sensitivity, and reversibility.

[0221] To further investigate the mechanistic details of the aptamer-functionalized hydrogel, QCM-D monitoring was employed to study hydrodynamically coupled mass changes and energy dissipation caused by the viscoelastic properties of the adsorbed material. The shifts in frequency (A ) and dissipation (AD) signals are displayed in Fig. 12(A). The acoustic mass shift observed from the A generally showed a similar trend to the MP-SPR results, particularly for the 0minand An. However, the MM1 -3 exhibited a slightly higher coupled mass shift compared to MD1 -3, which could be further elucidated by the AD signal. The larger dissipation shift of MM1 -3 compared to MD1 - 3 indicates that the amount of freely interacting vancomycin within the gel matrix was higher for the MM1 -3 configuration. This is likely due to the greater flexibility and mobility of the MM1 -3 aptamer segments, which are attached to the hydrogel matrix at only a single point. The freely moving, unattached end of the MM1 -3 split aptamers may not be optimally positioned or oriented for efficient target binding, reducing the binding affinity and sensitivity towards vancomycin. In contrast, the DD1 -3 and F1 -3 showed similarly high amounts of vancomycin association, as evidenced by their comparable A . However, the AD of DD1 -3 was significantly smaller than that of F1 -3, indicating that the dithiolated split aptamer configuration induced relatively rigid, elastic binding interactions with vancomycin, in comparison to the more viscoelastic binding behavior observed for the F1 -3.

[0222] To analyze the dependence of analyte-mediated hydrogel structural transformations on the aptamer configuration, time-independent frequency-dissipation (f-D) curves, which describe changes in the properties of the adsorbed layer, are presented in Fig. 12(B). Specifically, f-D curves provide insights into the relative conformational changes occurring throughout the association and dissociation phases involving vancomycin binding. The area enclosed within the f-D curve serves as a qualitative indicator of the extent of change in the adsorbed film properties, with a larger enclosed area corresponding to more extensive conformational changes in the film characteristics due to vancomycin-aptamer interactions.

[0223] For example, the F1 -3 was most affected by structural transformations upon vancomycin binding, presumably due to the long, flexible aptamer strand undergoing conformational folding in the presence of the target analyte. This is also indicated by the steeper slope of the F1 -3 compared to DD1 -3, suggesting a more dissipative, viscoelastic binding behavior for the full aptamer. When interacting with a similar amount of vancomycin, the F1 -3 induces higher energy losses compared to DD1 -3, with these effects being particularly pronounced during the dissociation phase rather than the association phase. Similarly, the MD1 -3 exhibited a more rigid adsorbed layer and reduced conformational freedom compared to the MM1 -3. As the MM1 -3 split aptamer pairs are tethered to the hydrogel matrix at only a single end, this configuration showed the largest dissipation shift when exposed to a comparable amount of vancomycin.

[0224] Given the different sensing penetration depths of the techniques employed, a suggested mechanism observed from MP-SPR (sensing penetration of SPR and waveguide mode ranging 200 nm and 10 pm, respectively) and QCM-D (sensing depth of 250 nm) is schematically described in Fig. 12(C). The comparison between the MM1 -3 and DD1 -3 designs highlights how the aptamer immobilization strategy can profoundly impact the binding performance and sensitivity towards the analyte. The secure dual-point attachment of the DD1 -3, with the split aptamer pairs tethered at both ends to the hydrogel matrix, appears to be a crucial factor contributing to its superior binding response compared to the more loosely tethered MM1 -3 configuration. In contrast, the single-point attachment of the MM1 -3 split aptamers allows greater flexibility and mobility, resulting in a more dissipative, viscoelastic binding behavior and reduced binding affinity towards vancomycin. The comparative analysis of these aptamer configurations, facilitated by the complementary sensing depths of SPR, waveguide mode, and QCM-D techniques, provides valuable insights into the binding mechanisms and the critical role of aptamer immobilization in dictating the target recognition and binding performance within the hydrogel matrix.

[0225] Optimization of split aptamer-functionalized hydrogel.

[0226] As the dithiol modification of both split aptamer pairs (DD configuration) has been identified as the most efficient design, the optimized hydrogel composition was developed with two key variations (see Fig. 13(A)). The first variation was the total aptamer concentration among the dithiol cross linkers, which ranged from 1 to 3 mM. The second variation involved the crosslinking density, controlled by the ratio between norbornene and thiol ([ene] / [SH]) at values of 3 and 4, as detailed in the previous section. To validate the specificity of the aptamer-analyte interaction, negative controls were included in the study. These controls consisted of a hydrogel without aptamers (DDO-4) and a hydrogel functionalized with a nonspecific poly A sequence (PAO-4). The inclusion of these negative controls allows for the assessment of the aptamer’s specific binding to the target analyte and the elimination of potential falsepositive responses.

[0227] The MP-SPR results for the various hydrogel compositions, screened with 1 mM vancomycin, are presented in Fig. 13(B). Generally, as the aptamer concentration increases, the shifts in all parameters increased (e.g., DD1 < DD2 < DD3). This trend can be attributed to the increased aptamer density within the hydrogel matrix, facilitating more efficient interactions with the diffusing vancomycin molecules. Another general trend observed was that the extent of shift differences, depending on the aptamer concentration, was more pronounced when the crosslinking density was lower (e.g., DDX-3 vs. DDX-4). This suggests that a lower degree of crosslinking (or a lower amount of neutral PEG-dithiol) provides more flexibility for split aptamers to come together upon vancomycin’s presence. Although DD1 -4 performed less effectively than DD1 -3, the 2 mM samples (DD2-3 and DD2-4) exhibited similar performance. However, at the 3 mM concentration, DD3-4 significantly outperformed DD3-3, establishing it as the top-performing configuration. The negative controls, DDO-4 and PAO-4, showed negligible responses, confirming the specificity of the aptamer-functionalized hydrogel system for vancomycin detection.

[0228] Sensor performance in rat plasma.

[0229] The optimized DD3-4 configuration was employed for a dose-response analysis to evaluate sensor performance. Vancomycin concentrations ranging from 10 pM to 1 mM were continuously injected into the measurement chamber, with rinsing steps performed after each injection to assess reversibility (see Fig. 14(A)). The angular reflectivity spectra for a hydrogel film, measured at different vancomycin concentrations, are presented in Fig. 14(B). The spectra reveal a larger shift in the SPR mode (71.5 to 73 deg) compared to the waveguide mode (62.5 to 63 deg, observed from the first peak). This observation contrasts with previous studies reporting improved resolution and lower detection limits for hydrogel optical waveguide spectroscopy (HOWS) compared to conventional SPR sensors. In those studies, the hydrogel film was first attached to the surface and then modified with receptors via amine coupling chemistry, potentially resulting in a less homogeneous distribution and reduced coupling efficiency within the SPR sensing depth region. In contrast, the inventors’ approach involved mixing the pre-gel solution with aptamers, followed by rapid polymerization, ensuring uniform aptamer distribution throughout the hydrogel matrix. Moreover, the homogeneous pore structure and density of the hydrogel likely facilitate rapid analyte diffusion to the surface. Consequently, the SPR mode was selected for final sensor performance assessment to obtain limit of detection and dissociation constant values.

[0230] From the SPR minimum angle shifts obtained in the dose-response study, the limit of detection (LOD) was calculated to be approximately 20 pM in diluted rat plasma (see Fig. 1 (C)), while a lower LOD of 12 pM was achieved in PBS buffer (see highlighted portion of Fig. 14(A). To assess the sensor’s specificity, other commonly employed antibiotics were) tested at their maximum concentrations of 1 mM, revealing negligible responses. Additionally, 10 mM glucose was evaluated to further support the sensor’s specificity. Furthermore, the relationship between refractive index and thickness as a function of vancomycin concentration was analyzed, also revealing linear relationships (graph on top right portion of Fig. 14(C)). These findings demonstrate the hydrogel platform’s high linearity, which can be easily translated into sensor readouts, facilitating the development of reliable and accurate sensing systems.

[0231] Next, the association rate constant (kon), dissociation rate constant (koff), and dissociation constant (Kd) were calculated from the association and dissociation curves depending on the vancomycin concentration. Compared to measurements using a monolayer setup, the kon of the hydrogel system was 30 times lower (200 vs. 7,000 M’1min’1), while the koff was 3 times lower (0.57 vs. 1.44 min’1). These results suggest that the association process in the hydrogel system was particularly slower than in the monolayer setup, likely due to the densely located aptamers and diffusion limitations, whereas the clearance rate was much faster than expected. The Kd value for the hydrogel system was found to be 15 times higher (2.9 vs. 0.2 mM) compared to the monolayer setup, indicating a significantly higher sensing capacity due to the enlarged binding capacity of the hydrogel platform.

[0232] It is important to note that this study utilized only one type of split aptamer pair derived from the original sequence. Improved versions of aptamer pairs with higher affinity and lower LOD values in monolayer setups have been developed, which can be further incorporated into this hydrogel system to enhance its sensitivity. The inventors hypothesize that there is an optimal range of Kd values that can be tailored to match the target concentration range for detection, enabling the development of highly sensitive and specific biosensors based on this aptamer-functionalized hydrogel platform.

[0233] Conclusions

[0234] In this study, the inventors have successfully demonstrated the superiority of difunctional split aptamer pairs over other configurations for the development of aptamer-functionalized hydrogel biosensors. Through comprehensive MP-SPR and QCM-D analyses, the inventors have validated the enhanced performance and mechanistic advantages of the difunctionalized split aptamer design. The split aptamer pair-functionalized hydrogel system exhibits an excellent linear relationship between the SPR angle shift / refractive index change and analyte concentration over a wide range of 10 to 1000 pM. In conclusion, the combination of a broad dynamic range, precise quantification capabilities, simplicity in data analysis, enhanced sensitivity, and versatility in analyte detection make this difunctional split aptamer pair- functionalized hydrogel system a promising platform for in vivo continuous biomolecular monitoring. The insights gained from this study pave the way for the development of advanced aptamer-based biosensors with improved performance and expanded applicability. Further optimization of aptamer sequences and hydrogel compositions based on the principles established in this work could lead to even greater enhancements in sensitivity, specificity, and overall sensor performance.

[0235] Experimental Section

[0236] Reagents. 8-arm PEG norbornene (MW: 10 kDa) was purchased from Creative PEGWorks. DNA aptamers with the sequences 5'- CGACCGAGGGTACCGCAATAGTACTTA-3' (Aptamer 1 ), 5'-

[0237] TTGTTCGCCTATTGTGGGTCGGGTCG-3' (Aptamer 2), 5'-

[0238] CGACCGAGGGTACCGCAATAGTACTTATTGTTCGCCTATTGTGGGTCGGGTCG -3’ (full aptamer) were obtained from Integrated DNA Technologies. The aptamers were modified with 5' thiol modifier C6 S-S and / or 3' thiol modifier C3 S-S for subsequent conjugation to the hydrogel matrix. All buffer reagents were purchased from Thermo Fisher Scientific as Dnase / Rnase-free stock solutions, including UltraPure™ distilled water, phosphate-buffered saline (PBS, 10X, pH 7.4), and magnesium chloride (1 M). HPLC grade ethanol and 1 ,4-dithiothreitol 99% (DTT) were acquired from Carl Roth GmbH. 1 -Dodecanthiol (DDT), PEG-dithiol (MW: 3.4 kDa), lithium phenyl(2,4,6-trimethylbenzoyl)phosphinate (LAP), Tris(2- carboxyethyl)phosphine hydrochloride (TCEP), vancomycin hydrochloride, and lyophilized rat plasma were obtained from Sigma-Aldrich. Hydrogel film preparation.

[0239] The custom-made SPR gold sensor chip (dimensions: 20 x 12 x 0.5 mm) was purchased from LET Optomechanika Praha. The gold surface was functionalized by incubating the chip in a 50 mM solution of DTT in ethanol for 5 h, followed by a brief 5-min incubation in a 50 mM solution of DDT, also in ethanol. The surface was then thoroughly washed with water and ethanol, dried, and stored in a vacuum desiccator before further use. The hydrogel precursor solution was prepared by first mixing the respective amounts of DNA aptamers, PEG-dithiol, and TCEP (final concentration of 80 mM) in DNase-free water. The mixture was incubated at 37°C for 2 h to reduce disulfide bonds to free thiol groups. Subsequently, 8-arm PEG norbornene and LAP (final concentration of 1 mM) were added to complete the pre-gel solution. A volume of 2 pL of this solution was drop-cast onto the treated gold surface, and a hydrophobic-treated glass slide was gently pressed onto the deposited solution. Finally, the sandwiched assembly was exposed to 365 nm UV light for 2 min to polymerize the pre-gel solution and form a thin hydrogel film.

[0240] Multi-parametric surface plasmon resonance (MP-SPR). The MP-SPR measurement was conducted using a dual-wavelength MP-SPR Navi 400 Kontio (BioNavis) instrument. The SPR was monitored at wavelengths of 670 and 785 nm, across an angular range of 40 to 78 degrees, with the angle of the SPR minimum used to generate the sensogram. All data presented herein were reported with a wavelength of 670 nm. The experiment was initiated by measuring the baseline with phosphate-buffered saline (PBS) buffer containing 2 mM MgCL, maintained at a temperature of 37°C. A consistent flow rate of 50 pL / min was used throughout all the measurements, controlled by a peristaltic pump (Ismatec). The SPR gold sensor chip with the hydrogel film was inserted into the instrument, and the sample chamber was filled with the running buffer (PBS containing 2 mM MgCL). The system was allowed to equilibrate for at least 30 min to establish a stable baseline before the start of the experiment. Following the baseline stabilization, the analyte solutions were injected into the flow cell at predetermined concentrations, followed by a dissociation phase where the running buffer was injected to remove any unbound analytes. The collected data were fitted with Winspall 3.01 software to determine the hydrogel's refractive index (n) and thickness (d). Quartz crystal microbalance-dissipation (QCM-D). QCM-D measurements were conducted using a Q-Sense E4 instrument (Biolin Scientific) to monitor the kinetics of aptamer-functionalized hydrogel and analyte interactions. The changes in resonance frequency (Af) and energy dissipation (AD) of a quartz crystal were measured as a function of time. The quartz crystal was excited to generate the thickness shear mode at its fundamental resonance frequency of 5 MHz and at odd overtones (n: 1 , 3, 5, 7, 9, and 11 ). All measurements were performed at 37°C with a flow rate of 50 pL / min, controlled by a peristaltic pump (Ismatec). The gold-coated quartz crystal sensor (QSX 301 , Biolin Scientific) was functionalized as described previously, and all data presented herein were collected at the fifth overtone. The system was allowed to equilibrate in the running buffer (PBS containing 2 mM MgCl2) for at least 30 min to establish a stable baseline. The collected data were analyzed using the QTools software (Biolin Scientific).

[0241] 2. Increasing sensitivity by tuning characteristics of hydrogel composition The hydrogel composition of the sensors system tested in this Example was prepared as already described in detail in example section 1. above. Briefly, aptagel was prepared using varying weights of dithiol PEG cross linkers, while keeping concentrations of the cross linkers, aptamers and 8-arm PEG polymers constant. Sensor responses were tested on MP-SPR using varying concentrations of vancomycin, and bovine serum albumin, globulin and dextrans at fixed concentrations, in PBS buffer (see Figs. 15(A) and (B)).

[0242] 3. Increasing sensitivity by tuning characteristics of metal-coated surface FO-SPR probe fabrication and readout was performed as described in further detail in below section 5, with only one additional step of silver layer deposition having been added before applying the final gold coating. Bulk sensitivity test using sucrose was also identical to described steps. Vancomycin sensing was performed at room temperature using aptagel describe in section 5, containing 10 kDa 8-arm PEG, 3.4 kDa dithiol PEG cross linker and dithiolated vancomycin split aptamer pair P27 (see Figs. 16(A) to (C)).

[0243] 4. Illustrating versatility of sensor system

[0244] Aptagel versatility tests were performed at 37°C using MP-SPR sensor systems comprising various aptamer sequences in hydrogel compositions prepared using 10 kDa 8-arm PEG, 3.4 kDa dithiol PEG cross linker as described above in example section 1. Dopamine hydrochloride was purchased from Fisher Scientific, corticosterone and vancomycin acquired from Sigma Aldrich and human Tumor necrosis factor alpha was from Sino Biological. All assays were performed in assay buffer consisting of PBS and 2 mM MgCL. To prevent dopamine from forming polydopamine complex, 1 % ascorbic acid was added. Due to limited solubility of corticosterone in aqueous solution, 1 % ethanol was added. Aptamer sequences were synthesized by IDT DNA.

[0245] 5. In vivo sensing Materials and Methods

[0246] Reagents

[0247] The employed ssDNA aptamer sequences were ordered from Integrated DNA Technologies. Vancomycin hydrochloride, MES (2-Morpholinoethanesulfonic acid) monohydrate and ethanolamine were purchased from Sigma-Aldrich. UltraPure™ distilled water, 10x concentrated phosphate-buffered saline (PBS) pH 7.4 (RNase- free), magnesium chloride, calcium chloride, sodium chloride, EDC (1 -ethyl-3-(3- dimethylaminopropyl)carbodiimide hydrochloride) and NHS (N-hydroxysuccinimide), and neutravidin biotin-binding protein were procured from Thermo Fischer Scientific. HPLC-grade ethanol was acquired from Carl Roth GmbH + Co. KG. Functionalized thiols (HS-(CH2)m-EG4-OH and HS-(CH2)m-EG6-OCH2-COOH) were purchased from ProChimia Surfaces, and thiolated PEG (CH3O-PEG-SH o-methoxy-co-mercapto PEG) was obtained from Rapp Polymere GmbH. 8-arm PEG norbornene (10 kDa) was purchased from Creative PEGWorks. P27 ssDNA aptamer sequences were manufactured by Integrated DNA Technologies with 5' thiol modifier C6 S-S and / or 3' thiol modifier C3 S-S for hydrogel matrix conjugation. 1 ,4-dithiothreitol (DTT) 99% were acquired from Carl Roth GmbH. PEG-dithiol (MW: 3.4 kDa), lithium phenyl(2,4,6-trimethylbenzoyl) phosphinate (LAP), 1 -Dodecanthiol (DDT), tris(2- carboxyethyl)phosphine hydrochloride (TCEP were obtained from Sigma-Aldrich. Drug monitoring on bare / aptagel FO-SPR probes were performed in assay buffer (PBS with 2 mM MgCI2 at pH 7.4), aCSF (artificial cerebrospinal fluid) (147 mM NaCI, 1.2 mM MgCI2, 1 mM CaCI2, 3.5 mM KCI, 2.5 mM NaHCO3, 1 mM NaH2PO4), horse serum (from Gibco (cat. no. 16050130)), and human plasma (from teams TMI and PTA at BioMed X Institute GmbH) and characterization of fiber

[0248] Plasmonic optical fiber fabrication

[0249] Fiber optic Surface Plasmon Resonance (FO-SPR) probes were constructed using multimode, step-indexed optical fibers with 300 pm diameter (d) pure silica core with a refractive index (nc) of 1.458 and numerical aperture of 0.39, allowing incident angles between 74.5 and 90 o. The fibers were sectioned into 6.5 cm pieces using a ruby-bladed fiber optic scriber to achieve flat probe end faces. Fiber sensing region (I) was designed as ~3 mm section on one end of the fiber for biosensing.

[0250] To fabricate plasmonic fiber probes, the protective buffer was removed using a specific stripping tool (T16S31 , Thorlabs inc.), and the TECS cladding by dissolving in acetone. A fiber inspection tool (FS201 -SMA adaptor) was used to verify the flatness of probe ends. The exposed core of fiber tips was cleaned with isopropanol and ethanol, followed by drying with N2 gas. Subsequently, fiber tips were uniformly coated with 45 nm of gold (Au) using an EM ACE600 sputter coater (Leica Microsystem, EMBL Imaging Centre) at a 0.11 nm / s sputter rate on a rotating stage. To confirm the thickness of the gold layer, fibers were freshly cut with a ruby blade, and the resulting cross-sections were examined under SEM, and finally, captured images were analyzed using Imaged.

[0251] FO-SPR sensor setup

[0252] The reflective optical fiber sensor setup as illustrated in the schematic Fig. 18(A) comprised of a broadband white light source (HL-2000-LL, Ocean Insight, Orlando, FL) and a high-resolution linear silicon array CCD spectrometer (FLAME-T-XR1 -ES, Ocean Insight) with SMA 905 connection outputs to a 1x2 (50:50) Y-multimode fiber optic coupler (300 pm core, Thorlabs Inc.). The coupler was linked to the light source on one end and the FO-SPR probe on the other using an SMA 905 multimode connector (340 pm Bore, SS ferrule, Thorlabs Inc.) via an SMA-SMA mating sleeve, ensuring optimal optical transmission. The sensing end of the fiber was secured with a strain-free bare fiber terminator (Newport, USA) to create a robust FO sensing probe. Incident light in the visible range (400 to 1100 nm) is reflected due to the resonance condition matching with the surface plasmon polariton waves (SPPWs) at the metal-dielectric interface at the 45 nm gold coated fiber tip. The reflected light is collected by the detector through the Y-coupler output.

[0253] A flow cell with a ~3 mm chamber depth was constructed using PDMS. The inlet and outlet tubes (Tygon, Ismatec with 0.25 mm inner diameter) were connected to a peristaltic pump (Ismatec, Switzerland) and the probe inserted flow cell was placed on a heater and enclosed by a custom-designed thermal insulation Teflon block to maintain a constant temperature. All biosensing measurements were conducted at 37°C.

[0254] First, the reflected light spectrum from the probe in air was normalized against the backlight reflection from the Y-coupler without the probe connected. Normalized reflectivity spectra R(A) were processed using a custom -written LabView software. The plasmon resonant wavelength (Aspr) was tracked over time t to obtain real-time sensor readout.

[0255] Fiber probe evaluation

[0256] The fabricated FO-SPR probes were evaluated for the following parameters prior to application in biosensing to ensure reproducibility. For each probe, the sensitivity (S) was determined by recording changes in Aspr (nm) when immersed in sucrose solutions with concentrations ranging from 0-10 vol%. The reflection spectra obtained in DI water (Rl nw = 1 .33) serves as the standard for analyzing the plasmonic sensor performance parameters, e.g., SPR percent reflection R(A), full width half maximum (FWHM), figure-of-merit (FOM), and detection accuracy (DA).

[0257] Biosensinq on bare FO-SPR probe

[0258] Surface functionalization of bare FO-SPR probe

[0259] Bare FO-SPR probes were fabricated for the detection of vancomycin binding using the split-aptamer assay previously described by Santa et al. (2025). We employed the split-aptamer pair designated P27, as it exhibited excellent reversibility in drug-binding interactions while maintaining an appropriate detection limit of 9 pM on the planar SPR sensor.

[0260] The FO-SPR probes were initially cleaned with 99% ethanol and dried using N2 gas. The fibers were then submerged in a 1 mM ethanolic solution of carboxyl and hydroxyl thiols (1 :9) under nitrogen for 12 hours to create self-assembled monolayers, followed by ethanol rinsing and dried with N2. A flow cell with a 0.8 mm aperture was positioned on a hotplate and connected to a microfluidic pump, at a flow rate of 50 pL / min. After calibrating the spectrometer in air, the probe was inserted into the flow cell containing DW using a controlled X-Y-Z optomechanical stand. The change in refractive index due to the presence of water produced SPR reflection spectra with a resonance dip (Aspr), which was tracked within a narrow wavelength range of about ± 10 nm throughout the sensing measurement. The sensor was subsequently treated with 100 mM thiolated PEG for 10 minutes to passivate the gold surface. After the PEG passivation the probe was functionalized with neutravidin (NA) protein for 20 minutes through activated ester groups by 10 minutes incubation with aqueous solution of EDC (75 mg / mL) and NHS (25 mg / mL), followed by neutralizing unreacted NHS ester groups with 1 M ethanolamine at pH 8. The biotinylated segment of the P27 aptamer split pair was immobilized onto neutravidin for 20 minutes, briefly washed, and lastly, the remaining biotin binding sites were blocked by a 10-minute incubation with biotin (100 mM). The functionalized split-aptamer sensor surface was equilibrated with assay buffer to stabilize the system and established baseline at steady flow and temperature of 37°C. All buffers and solutions were kept on a heating block before introduction to ensure sensor stability.

[0261] Vancomycin monitoring on monolayer FO-SPR surface

[0262] Vancomycin detection measurements were carried out similarly to previously reported by Santa et al. (2025) with free segment of split aptamer pair (S2) kept at constant concentration of 2 pM while introducing varying concentrations of vancomycin (ranging from 0.48 pM to 125 pM) in assay buffer. Sensor responses were recorded by establishing baseline and washing carried out in two different approaches. The first approach involved establishing the baseline and performing washing with assay buffer without S2 present, whereas the second approach kept S2 consistently in assay buffer during baseline and washes. Sensor signal peaks at each concentration of vancomycin were extracted, and dose-response curves were generated. Only at high vancomycin concentrations (64 and 125 pM), bulk response was observed without S2 present and thereby subtracted from specific signals in the dose-response curves. The sensor limit of detection (LOD) was determined using linear regression fitting to the dose-response curves. The intercept concentration with the 3 times standard deviation (o) on that linear regression was considered as LOD. Vancomycin monitoring using aptagel FO-SPR sensor

[0263] Aptagel functionalization on FO-SPR probe

[0264] Here the development and analysis of an implantable plasmonic fiber sensor system in accordance with the present invention - designed for extended continuous monitoring both in vitro and in vivo - is again described. Specifically, the molecularly- responsive hydrogel composition of the sensor system used here, includes the vancomycin binding aptamer P27 split pair as the allosteric molecule fragments forming an integral part of the 3-D network structure. The hydrogel composition (aptagel) is further composed of 8-arm poly(ethylene glycol)-norbornene (PEG-NB) star-shaped macromers crosslinked through dithiolated PEG andm as already mentioned the aptamer fragments. A wide-angle MP-SPR (Bionavis) device was used to examine optical properties (refractive index n and thickness d) of the aptagel, as well as binding kinetics to evaluate the optimal configurations for direct vancomycin sensing on SPR. The optimized aptagel formulation comprises a 3 mM aptamer concentration and a norbornene to thiol ratio ([ene] / [SH]) of 4. A negative control aptagel was tested to verify sensor functionality and specificity, containing a seguence of poly A (3 mM) instead of the analyte-binding aptamer seguences.

[0265] Before implementing the aptagel coating, the FO-SPR probes were incubated in a 50 mM solution of dithiothreitol (DTT) in ethanol for 5h, followed by a short 5 min incubation in a 50 mM solution of DDT in ethanol. Thereafter, the probes were washed with ethanol, water and then dried using N2 gas. Before the gel coating, the prepared pre-gel solutions of aptagel (AH) and negative control (NC- hydrogel with polyA seguence) were incubated at 37°C for 2 h. For gel coating, the fiber probes were submerged in AH or NC pre-gel solutions for 10 sec and then exposed to 365 nm UV light for half a minute of cross linking. To ensure uniform gel formation on the fiber probes, the gel coating procedure was repeated three times for each fiber.

[0266] Aptagel presence on the FO-SPR probes was verified under SEM (ZEISS, EMBL) by examining a freshly cut fiber probe tip coated with aptagel. Both the secondary electrons (SE) and the energy-selective backscatter electrons (ESB) were imaged to visualize the fiber core, gold, aptagel, their interfaces, and the respective thicknesses. Acguired cross-sectional electron micrographs were analyzed using Imaged. The swollen gel thicknesses were then calculated. Vancomycin monitoring on aptaqel FO-SPR

[0267] Vancomycin monitoring in complex biological fluids

[0268] All sensing measurements were conducted under flow conditions maintained at 37°C. Aptagel FO-SPR probes were first inserted into the flow cell filled with assay buffer and allowed to stabilize under 37°C temperature-controlled continuous flow conditions until the hydrogel was fully swollen. Once the Asprsensorgram baseline stabilized, analyte solutions were introduced into the flow cell to observe analyte association (binding) and washed with buffer for analyte dissociation.

[0269] Sensors were additionally validated for vancomycin monitoring in complex matrices such as artificial cerebrospinal fluid (aCSF), human blood plasma (20 %), and horse blood serum (100 %). The linear regression equation was used to fit the lower ranges of dose-response curves and to quantify LOD. All obtained sensorgrams were analyzed using Origin 2024 v 10.1 .

[0270] Vancomycin monitoring in vivo

[0271] Stereotactic surgery and vancomycin pharmacokinetics measurements in live rat brains

[0272] Animals

[0273] Male Wistar rats (Janvier, France) weighing 320-360 g were used for the in vivo procedures. The animals were housed in groups three per cage under standard laboratory conditions on a 12h light / dark cycle (lights on at 06:00) in temperature (21 - 22°C) and humidity (55-65%) controlled rooms with free access to food and water. All in vivo procedures were approved by the appropriate institutional governmental agency (Regierungspraesidium Tuebingen, Germany) and performed in an AAALAC (Association for Assessment and Accreditations of Laboratory Animal Care International-accredited facility in accordance with the European Convention for Animal Care and Use of Laboratory Animals.

[0274] In vivo procedures

[0275] In vivo assessment of the FO hydrogel sensors was performed in anesthetized male Wistar rats. On the day of the experiment, the rats were anesthetized with isoflurane and mounted in a stereotaxic frame (Kopf® Model 1900). Anesthesia was maintained during the entire experiment by using 0.2-2% isoflurane. The skull was exposed, and a small hole was drilled to enable the insertion of the FO hydrogel sensor into the brain tissue. The sensor was placed in the cortex just above the left striatum (stereotaxic coordinates: AP +0.5, ML - 3.0, DV - 3.0 mm from bregma) according to the brain atlas of Paxinos and Watson (1998).

[0276] Commencing ~2 hours after the sensor insertion (an equilibration period), the rats were dosed with vancomycin (75 mg / kg or 150 mg / kg; i.v.; 2 ml / kg) or vehicle (0.9% NaCI; i.v.; 2 ml / kg) and the recording continued for a further 2 hours until the remainder of the experiment. Upon completion of the experiments, the rats were injected with sodium pentobarbital (100 mg / kg, i.p.), and the brains were removed for the histological verification of the sensor localization.

[0277] The recorded data were corrected for baseline drift using the Fityk software. The SPR shift (nm) versus time for the in vivo data was employed in the below equation to generate the pharmacokinetic plots.

[0278] C = m- ASPR shift+ b (1 )

[0279] Where C is the vancomycin concentration, m, and b are the fitted parameters obtained from the linear equation fitting to ASPR shift (nm) vs time data (lower linear ranges) collected with the horse serum.

[0280] Result and discussion

[0281] FO-SPR sensor working principle and design

[0282] A polychromatic light spectrum from the source propagating through the fiber reaches the silica core-metal interface via a Y-splitter. At the resonance angle (0min), resonant coupling of the evanescent field of the reflected light and the surface plasmon polaritons (SPP) on the metal-medium interface occurs, generating surface plasmon resonance modes. Then, a detector monitors the resulting reflected light output, which exhibits minima at the resonance wavelength (Aspr) as shown in Fig. 18. Initially, we evaluated vancomycin sensing capabilities using bare FO-SPR with P27 splitaptamer assay strategy described in Santa et al. (2025).

[0283] An implantable plasmonic fiber sensor design and working principle for extended continuous monitoring of vancomycin is shown in the schematic (Fig. 18). The sensor features an innovative molecularly-responsive “aptagel”27matrix coated on the FO- SPR (Fig. 18(A)), composed of a hydrogel (star-shaped 8-arm poly(ethylene glycol)- norbornene (PEG-NB) macromers crosslinked through dithiolated PEG) integrated with vancomycin-binding aptamer P27 split-aptamer pair. The incorporation of P27 split-aptamers within the hydrogel matrix enables tunable mechanical transitions through the formation of reversible ternary complexes between split-aptamer pairs and vancomycin analyte. These transitions from a swollen unbound state to a collapsed analyte-bound state induce density changes within the aptagel matrix, resulting in refractive index modulation ( AH) in the optical regime. This change produces a measurable shift (AA) and in the resonance wavelength (Aspr) observed in the reflection spectra (Fig. 18(B)). The AA Dnm > is directly proportional to the An near the gold surface, with analyte binding causing SPR spectrum toward higher wavelengths (redshift). This dynamic system allows for continuous, reversible analyte detection under physiological conditions. The design of aptagel FO-SPR provides a self-contained, miniaturized implantable sensor probe with enhanced AA compared to the bare FO-SPR sensors, suggesting a higher optical transduction efficiency. This improvement stems from the uniform response of the 3-D aptagel matrix to analytes through matrix contraction (Ac / ).

[0284] Fig. 18(C) shows the reversible and continuous vancomycin detection by the aptagel FO-SPR probes in accordance with he present invention. The aptagel FO-SPR probe design is one of the crucial factors for making highly sensitive implantable sensors. Brain implantation brings complexity for the sensor, which can be easily fouled by diverse cell types (neuron, glial cells, etc.), extracellular matrix components, and biological molecules that can significantly diminish senor performance and longevity as has been seen for many EAB sensors. Due to the specific hydrogel composition used in combination with the metal-coated senor probe, the sensor system of the present invention offers dual advantages: enhanced optical transduction and reduced biofouling (as illustrated in Fig. 18(A)), providing a miniaturized plasmonic sensing platform for real-time in vivo monitoring.

[0285] Aptagel FO-SPR sensor characterization

[0286] For a minimally invasive implantation and monitoring different brain regions in rodents, the probe diameters (d) in the range of 100 pm or less with sensing length ( / ) of 3-5 mm are applicable. Currently, the size of electrochemical based sensors employed for real-time monitoring of biomarkers has decreased to 10-50 pm using carbon fibers and metal electrodes. Achieving such small micron sizes while maintaining high sensitivity is challenging for plasmonic fiber optic probes. Unlike the electrochemical sensors, the performance paraments of FO-SPR are strongly dependent on the size (d and / ) of the probes39. This can be explained as the sensitivity (S) is directly proportional to the total number of reflections (Nref) performed by the light at the fiber core and metal interface, generating SPR modes. Nref can be expressed as: where the reflections are performed by a ray incident at an angle 0, measured from the normal to the core-metal interface.

[0287] A higher / enhances the S by increasing the interaction length with the surrounding medium. Conversely, while smaller diameters (d) theoretically increase Nref for a given length, they also introduce higher scattering or absorption losses at the fiber walls, ultimately decreasing S. It is desirable to achieve an optimal sensing area by considering all the aspects of geometric restrictions while maintaining higher sensitivity. All previously published FO-SPR sensors have d and / values equal to or higher than 400 pm and 5 mm, respectively. In contrast, in the sensor system of the present invention, the sensing length can be (and was) reduced to ~3 mm, and the diameter to 300 pm, leading to a 2.2-fold reduction in sensing area compared to other reported multimode FO-SPR sensors. This reduced sensing area was specifically designed for in vivo monitoring in the brain of rodents. A schematic representation of the FO-SPR sensor system developed for real-time and continuous in vivo monitoring of biomarker analytes is shown in Fig. 19(A). The FO-SPR probe with a 3 mm sensing length is positioned in a PDMS fluidic chamber with a continuous flow and temperature-controlled system to replicate physiological conditions.

[0288] Unlike angular-modulated SPR systems, our bare / aptagel FO-SPR probes operating in wavelength-modulated SPR detection principles are limited with detection sensitivity. An aptagel with a higher refractive index (nAH ~ 1.348) than the air on the FO-SPR probes can significantly alter the plasmonic resonance conditions by increasing the effective refractive index of the medium leading to enhanced dimpling of surface plasmons. This can lead to change in SPR reflection spectra and plasmonic parameters contributing to the S of the probes. Experimentally obtained SPR reflection spectrum for both type of probes (bare / aptagel FO-SPR) is shown in Fig 19(B). To match the plasmonic conditions, both probes were air referenced, and the wavelength spectra were collected in water medium. The resonance wavelength (Aspr) for the bare probe is around 630 nm, shifting to higher wavelengths for the aptagel incorporated plasmonic probes owing to the higher effective refractive index of the aptagel (JIAH > nw). The observed shift (AA) is approximately 30 nm. Upon aptagel incorporation, a lowering in SPR reflection magnitude, and additionally spectral broadening in the spectrum are observed. The above is supported by theoretical simulations. Here we have employed COMSOL Wave Optics module considering fiber core and a gold layer of 45 nm on top in a Kretschmann configuration to simulate the electric field distribution around the metal-surrounding medium interface, and the resonance angle for generated SP mode. Both angular and wavelength interrogation methods were analytically evaluated. The incident angles were varied while fixing the input wavelength at 650 nm initially. The percentage reflected light obtained by varying the angle of incidences for the surrounding mediums (air (na), water (nw), and aptagel (HAH)) was determined. The change in resonance angles (0min) owing to the refractive indices of the above mediums are found to be 46° for air, 82° for water, and 86° for a gel with n =1.35. The employed optical fiber probes with NA of 0.39 allow the light with the angles varying in between 74.5° to 90°. The theoretical values suggest that an aptagel with DAH surrounded on a FO-SPR can only have a limited range of allowed angles, along with a lower magnitude in SPR reflection, and spectral broadening (Supplementary Fig. 4A).

[0289] Bulk sensitivity (S) values evaluated from the experiment are shown in Fig. 2C. The AA change upon An fitted with linear regression equation to calculate the sensitivity values of the bare / aptagel FO-SPR probes. The theoretical sensitivity value is around 3627 nrn / RIU, the yielded sensitivities for bare and aptagel integrated probes are found to be around 2324 nrn / RIU, and 2128 nrn / RIU, respectively. It is worth mentioning that, the aptagel resulted spectral broadening led to a slight reduction (~ 10%) in sensitivity, but the values are comparable with the FO-SPR sensors with larger sensing areas.

[0290] Knowing the sensitivity (S nrn / RIU) values for the bare FO-SPR, along with the AA collected in the water medium after the aptagel incorporation, we could formulate the refractive index of the aptagel (HAH) by employing the below equations

[0291] S = AA / An (3) nAH = n + nw(4)

[0292] The calculated HAH value, averaged across three different FO-SPR aptagel sensors, is approximately 1.3476 ± 0.002 Rill, which matches well with the values we previously obtained using a standard MP-SPR set up.

[0293] Fig. 19(D) illustrates the key performance parameters of FO-SPR sensors, including percentage reflection (% R(A)), full width half maxima (FWHM), and figure of merit (F0M=S / FWHM), which are crucial for sensor validation and reproducibility. The presence of the aptagel layer impacts the above parameters, making it essential to evaluate their variability before proceeding to analyte sensing. The R% is directly linked to the magnitude of the surface plasmon mode generated at the metalsurrounding medium interface. Higher R% enhances the sensitivity by improving plasmon excitation efficiency. A narrower FWHM contributes to a better signal-to- noise ratio, improving sensitivity and give rise to a higher FOM. As expected, the aptagel incorporation has brought a reduction in R% of 12-15% owing to the dampening of the plasmon modes and increased signal losses as the effective refractive index increases. Consequently, the spectral broadening causes an increase in FWHM of about 10 nm, and a small reduction of ~ 4 RILT1is observed in FOM. While the aptagel coating affects plasmonic parameters through reduced reflection intensity and spectral broadening, these changes remain within acceptable levels, preserving the functionality of sensors for sensitive biomarker detection with the added benefits of self-contained molecular recognition capabilities.

[0294] The cross-sections of the gold-coated fibers and the aptagel coated FO-SPR probes were examined under SEM to confirm the probe architecture outlined above. The gold layer thickness on the FO-SPR determined is approximately 45 nm. The thickness of the dry aptagel was found to vary between 1-3 pm. The aptagel layers on a planar gold surface are known to have a swelling ratio of approximately 9.

[0295] Continuous drug monitoring using monolayer split-aptamer assay and aptagel on FO-SPR

[0296] Bare FO-SPR probes were validated for vancomycin detection using a known splitaptamer pair-based monolayer assay. In this monolayer assay, the analyte binding principle differs from that of the aptagel-integrated FO-SPR probes. Specifically, the monolayer assay involves immobilizing one part of the split aptamer (split 1 ) directly onto the surface of the bare FO-SPR, while the other part (split 2) remains free in the solution. In contrast, in the aptagel-integrated probes, both aptamer splits (split 1 and split 2) are embedded within the gel matrix, making them self-contained and thereby viable as an implantable sensors. Fig. 20(A) shows the sensorgrams obtained using both the monolayer assay and the aptagel-integrated FO-SPR probes. These measurements were conducted across a decreasing concentration range of vancomycin, from 125 pM to 0.48 pM. For the monolayer assay, two distinct experimental approaches were employed as explained before in the materials and methods section. In this context, we compared the sensing responses of the bare plasmonic fiber probes with those of the aptagel-integrated counterparts. In both instances, the baseline was established using the assay buffer (PBS), and specifically for the monolayer assay, split 2 was also consistently present throughout the measurements. In the latter case, the S2 assay buffer deployment was to mimic the similar conditions as it is in the aptagel where both the splits are present within the matrix throughout the sensing. Both the sensors exhibited excellent reversibility and steady responses to the different vancomycin concentrations as seen in Fig. 20(A). A bulk response of 0.10-0.14 nm AA shift was observed only at a high vancomycin concentration (125 pm), with no detectable signal in the lower ranges for both monolayer and aptagel. The bulk change values were subtracted from specific signals in the dose-response curves. The result demonstrates the high specificity and enhanced responsivity of the P27 split-aptamer pair embedded within the gel matrix in accordance with the present invention. In the concentration range > 60 pM, the signal response from the aptagel increases significantly (Fig. 20(C)), highlighting the ability of the aptagel 3-D network to enhance optical transduction upon analyte interaction. This enhancement is particularly pronounced at higher concentration ranges due to the higher density of bioreceptors within the 3-D network, driven by its dynamic and responsive structure. Additionally, the aptagel demonstrated sensing responses across a wide range of analyte concentrations (500 pM to 0.48 pM) exhibiting two distinct SPR shift regimes. This dual-regime behavior of the aptagel sensor was also observed in measurements conducted using angular SPR, which confirms the detection accuracy of the aptagel system with FO-SPR.

[0297] The dose-response curves obtained from aptagel FO-SPR sensors in standard buffer (PBS) and aCSF are outlined in Fig. 20(B). At lower analyte concentrations (< 60 pM), the SPR shifts (in nm) from the aptagel-coated FO-SPR showed a linear relationship with the logarithmic values of the analyte concentration. The sensor limit of detection (LOD) was determined using linear regression fitting to the dose-response curves. The intercept concentration with the 3 times standard deviation (o) on that linear regression was considered as LOD. The calculated LODs for the aptagel-incorporated FO-SPR sensors in detecting vancomycin are 2 pM for PBS and 2.9 pM for aCSF. Subsequently, the LOD for the monolayer assay on a bare FO-SPR were determined to be approximately 1.2 pM for the S2 assay buffer. The aptagel system significantly enhanced the sensing performance by lowering the LOD to 0.25 pM. The clinically relevant concentration range of vancomycin in blood and in the brain of rodents, typically falls between 6-30 pM. As such, the current sensitivity of the aptagel FO-SPR sensor is well suited for in vivo detection of the drug.

[0298] Continuous drug monitoring using aptagel FO-SPR in complex biofluids

[0299] Real-time continuous monitoring of vancomycin under complex biological fluid conditions mimicking the in vivo scenario were accessed to establish the efficiency of our aptagel FO-SPR sensors for in vivo. Diluted 20 % human blood plasma (BP 1 :5) with assay buffer and undiluted horse serum (HS) were availed for the purpose. These biological fluids contain diverse proteins (Albumin ~ 60%, Globulins ~ 30%, Fibrinogen 4%, and other regulatory proteins ~ 1 %) to evaluate the anti-fouling and biocompatible properties of the aptagel. Non-specific interactions in complex biological fluids can significantly interfere with the sensing performance, specifically in the case of plasmonic based sensors. To illustrate the real-time vancomycin sensing performance, a sensogram of a gradually increasing then decreasing vancomycin concentration measurements collected with undiluted horse serum (100%) (HS) is shown in Fig. 21 (A) and (B). The calculated LOD values are 1.8 pM for BP 1 :5 and 2.5 pM for HS, matching closely with the values obtained in the standard assay buffers. These measurements were performed after incubating and stabilizing the aptagel FO-SPR sensors in the presence of these complex fluids for approximately 4-5 hours. The sensors remained functional and allowed for multiple sensing experiments. A combined dose-response for PBS, BP 1 :5, and HS is presented in Fig. 22. Notably, the signal response in the lower concentration range (below 100 pM) increased by 1.5-2 times in complex biological buffer media compared to PBS.

[0300] The specificity of the aptagel fiber probe sensors was also evaluated against several other drug molecules, including a cyclic peptide antibiotic bacitracin (molecular weight similar to vancomycin), rifampicin, which is an antibiotic often used in combination with vancomycin, and ampicillin (a commonly used antibiotic). The concentrations of these molecules were adjusted to match the refractive index of 30 pM vancomycin based on their molecular weights. The sensor responses to these drugs, along with vancomycin, are presented in Supplementary Fig. S11. All responses from the nontarget drugs were below the 3o threshold of the vancomycin calibration curve, confirming the specificity of the aptagel FO-SPR sensor. These results demonstrate the suitability of the label-free all optical aptagel FO-SPR sensor for real-time drug monitoring applications in complex biological environments.

[0301] Continuous drug monitoring using aptagel FO-SPR in vivo

[0302] Here the inventors present the first demonstrations of analyte (here: vancomycin) detection in the brain of a live animal using an FO-SPR sensor system in accordance with the present inventio, i.e. a sensor system comprising a hydrogel composition where the sensing allosteric molecule forms an integral part of the 3-D network structure of the hydrogel composition and where analyte binding is fully reversible in situ. This emphasizes the suitability of the sensor systems of the present invention for long-term and continuous in vivo analyte detection in tissue such as the brain.

[0303] The real-time and continuous pharmacokinetic monitoring of vancomycin was executed by implanting the aptagel FO-SPR sensors of the invention in the left prefrontal cortex (PFC) of five different rats. In four of these rats, an intravenous bolus of 75 mg / kg was administered via the tail vein. The sensor responses are shown in Fig. 23(A) and (C). Vancomycin concentrations were calculated using the undiluted horse serum as a reference, with observed values around 6-7 pM. All sensors demonstrated a response to the drug, characterized by a gradual increase to a peak concentration, followed by a steady state, and eventual clearance from the observation site. Before any drug injection, saline administration was performed through IV bolus, to monitor the vehicle response on our sensor. The sensor showed no measurable response from the vehicle injection compared to the vancomycin. The dose dependency of vancomycin is seen in Fig. 23. A higher dose (150 mg / kg) resulted in a response that was nearly double that of the initial 75 mg / kg dose (see Fig. 23(B)). Additionally, higher doses were associated with prolonged drug clearance times, further highlighting the effectiveness of our sensor system for pharmacokinetic studies. An extended measurement of six hours was performed, allowing for repeated dosing, as demonstrated in Fig. 23(C). For instance, Rat 2 was administered a second dose of 150 mg / kg after the first dose had cleared. The observation period was limited by the anesthesia protocol and the survival of the animal under prolonged anesthesia.

[0304] Reference

[0305] Santa, C.; Park, S.; Gejt, A.; A. Clark, H.; Hengerer, B.; Sergelen, K. Real-Time Monitoring of Vancomycin Using a Split-Aptamer Surface Plasmon Resonance Biosensor. Analyst 2025, 150 (1), 131 141. https: / / doi.org / 10.1039 / D4AN01226G. List of reference signs

[0306] 1 Surface plasmon resonance sensor

[0307] 2 Fiber connector, housing

[0308] 2a Light source

[0309] 2b Optical sensor

[0310] 2c Optical coupler

[0311] 2d Communication circuitry

[0312] 2d Battery

[0313] 3 Incident light

[0314] 4 Reflected light

[0315] 5 Glass core

[0316] 6 Gold film

[0317] 7 Cable

[0318] 8 Processor circuitry

[0319] 9 Hydrogel matrix

[0320] 10 Drug-specific aptamers

[0321] 11 Brain tissue

Claims

1. C l a i m s1 . A sensor system for sensing an analyte of interest, comprising:- a sensor to provide sensor data, wherein the sensor comprises a metal- coated surface;- a hydrogel composition contacting the metal-coated surface of the sensor, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or allosteric fragments thereof, wherein the allosteric molecules or fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest; and- processor circuitry to determine information on a presence of the analyte of interest based on the sensor data, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel, and wherein the sensor data is indicative of the alteration of the spatial dimensions of the 3-D network structure.

2. The sensor system according to claim 1 , wherein the sensor is a surface plasmon resonance sensor and the sensor data is indicative of a surface plasmon resonance shift, with the alteration of the spatial dimensions contributing to the surface plasmon resonance wavelength shift, or wherein the sensor is a sensor comprising a quartz crystal, with the sensor data being indicative of a shift in resonance frequency of the quartz crystal and / or a shift in energy dissipation of the quartz crystal, with the alteration of the spatial dimensions contributing to the shift in resonance frequency and / or energy dissipation.

3. The sensor system according to claim 1 or claim 2, wherein at least some of the hydrophilic polymers of the hydrogel composition connect the 3-D network structure to the metal-coated surface of the sensor.

4. The sensor system according to any one of claims 1 to 3, wherein the hydrophilic polymers are poly(ethylene glycol) (PEG), dextran, poly(N- isopropylacrylamide) (pnipaam), polyvinyl alcohol (PVA), hydroxyethyl methacrylate(HEMA), poly(acrylic acid) (PAA), polyethylene oxide) (PEO), alginate and / or hyaluronic acid (HA) polymers, and / or wherein the hydrophilic polymers are multivalent polymers, preferably comprising between 3 and 10 attachment points; and / or wherein the hydrophilic polymers comprise multivalent poly(ethylene glycol) (PEG) polymers, preferably comprising 8 attachment points, and / or wherein the allosteric molecules are connected with at least some of the hydrophilic polymers, preferably each allosteric molecule links two hydrophilic polymers such as to form an integral part of the 3-D network structure, and / or wherein the allosteric molecules or allosteric fragments thereof are aptamers or fragments thereof, preferably DNA aptamers or fragments thereof, which assemble into an aptamer-analyte binding complex upon binding to the analyte of interest.

5. The sensor system according to any one of claims 1 to 4, wherein the 3-D network structure of the hydrogel composition comprises several reaction products of a thiol-ene reaction between one functional moiety of one allosteric molecule or allosteric fragment thereof and a correspondingly reactive group of one attachment point of one of the hydrophilic polymers, and / or wherein the 3-D network structure of the hydrogel composition comprises several reaction products of thiol-ene reactions between two or more functional moieties of one allosteric molecule or allosteric fragment thereof and correspondingly reactive groups of one respective attachment point of two or more separate hydrophilic polymers, and / or wherein the binding of the allosteric molecules or allosteric fragments thereof to the analyte of interest is reversible such that the analyte binding complex disassembles upon separation of the analyte of interest from the aptamer fragments and the conformational change is reversed, and / or wherein the hydrogel acts as a biosensor transducer, preferably as molecularly responsive biosensor transducer.

6. The sensor system according to any one of claims 1 to 5, wherein the metal- coated surface is a surface of a glass fiber of a surface plasmon resonance sensor, preferably with the glass fiber having an exposed length of at most 10 mm and / or with the glass fiber having a diameter of at most 500 pm; and / orwherein the metal-coated surface comprises at least one of gold, silver, copper, aluminum, chromium, platinum, indium, a titanium oxide, a tantalum oxide, a zinc oxide or a silicon dioxide, preferably wherein the metal-coated surface comprises one of a gold-coated surface and / or a metal multilayer structure comprising two or more of gold, silver, copper and aluminum, a Ta2Os-coated surface and a TiO2 coated surface, and / or wherein(a) the sensor is a surface plasmon resonance sensor comprising a light source having an emission spectrum across a first bandwidth of at least 100 nm, and a sensor for detecting an intensity of reflected light having a sensing spectrum across a second bandwidth of at least 100 nm, wherein the first and the second bandwidths overlap, or(b) the sensor is a surface plasmon resonance sensor comprising a tunable light source for sequentially emitting light at different wavelengths of a wavelength range of interest, and a sensor for detecting a change in intensity of reflected light when a probe of the surface plasmon resonance sensor is illuminated by the tunable light source at the different wavelengths of the wavelength range of interest.

7. The sensor system according to any one of claims 1 to 6, wherein the sensor is a surface plasmon resonance sensor comprising a light source for emitting light at a wavelength, means for varying an angle of incidence of the emitted light, and a sensor for detecting a change in intensity of reflected light when a probe of the surface plasmon resonance sensor is illuminated by the light source at different angles of incidence, and / or wherein the processor circuitry is to determine the information on the presence of the analyte of interest quasi-continuously and / or according to a pre-defined schedule, and / or wherein the processor circuitry is to determine information on an estimated concentration of the analyte of interest based on the sensor data, and / or. wherein the sensor system comprises a sensor probe, the sensor probe comprising the sensor with the metal-coated surface and the hydrogel composition, and wherein, optionally, the sensor probe further comprises the processor circuitry.

8. The sensor system according to claim 7, wherein the sensor probe further comprises communication circuitry for providing the information on a presence of the analyte of interest to a separate device, or wherein the processor circuitry is separate from the sensor probe.

9. The sensor system according to claim 7 or claim 8, wherein the sensor probe is to provide sensor data acquired in situ, preferably the sensor probe is a sensor probe to be inserted into a bodily fluid, tissue or organ, and / or wherein the sensor probe is to be implanted.

10. The sensor system according to any one of claims 1 to 9, wherein the analyte of interest is- an organic molecule, a pharmaceutically active substance, a metabolite, a toxin or a hormone; or- a protein or peptide such as an enzyme, an antibody, a protein antigen, a peptide antibiotic, a proteinaceous viral particle, a viral protein, a protein receptor or a cytokine or- a nucleic acid molecule; or- a metal ion; or- a lipid; or- a carbohydrate, and / or wherein, when the allosteric molecules are aptamers, each aptamer comprises between 20 and 90 nucleotides and, when the allosteric molecule fragments are aptamer fragments, each aptamer fragment comprises between 10 and 40 nucleotides, and / or the hydrogel composition comprisesAptamers of SEQ ID NO:1 , and / orAptamer fragments of SEQ ID NOs: 2 and 3.

11. A method for sensing an analyte of interest in an aqueous sample, the method comprising the following steps:(a) providing a sensor comprising a sensor probe with a metal-coated surface and a hydrogel composition contacting the metal-coated surface, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a)hydrophilic polymers as well as (b) allosteric molecules or fragments thereof, wherein the allosteric molecules or fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest, wherein each allosteric molecule undergoes a conformational change upon assembly of the respective analyte binding complex, which alters the spatial dimensions of the 3-D network structure of the hydrogel;(b) bringing the sensor into contact with the aqueous sample;(c) providing, by the sensor, sensor data being indicative of the alteration of the spatial dimensions of the 3-D network structure; and(c) determining information on a presence of the analyte of interest in the aqueous sample based on the sensor data being indicative of the alteration of the spatial dimensions of the 3-D network structure.

12. The method according to claim 11 , wherein the sensor is a surface plasmon resonance sensor and the sensor data is indicative of a surface plasmon resonance shift, with the alteration of the spatial dimensions contributing to the surface plasmon resonance wavelength shift, or wherein the sensor is a sensor comprising a quartz crystal, with the sensor data being indicative of a shift in resonance frequency of the quartz crystal and / or a shift in energy dissipation of the quartz crystal, with the alteration of the spatial dimensions contributing to the shift in resonance frequency and / or energy dissipation.

13. The method of claim 11 or claim 12, wherein- the method is an ex vivo method for sensing the analyte of interest in an aqueous sample, preferably the aqueous sample is a previously-obtained bodily fluid sample or a tissue sample, or- step (c) providing the sensor data by the sensor, comprises inserting a probe of the sensor into a bodily fluid or tissue for sensing the analyte of interest in situ, and wherein, optionally, the probe of the sensor is implanted for continuous sensing the analyte of interest in situ, preferably the continuous sensing over a period of 1 to 60 days, preferably 1 to 30 days.

14. Use of a sensor system according to any one of claims 1 to 10 for sensing an analyte of interest in an aqueous sample ex vivo, andwherein, optionally, the aqueous sample is a previously-obtained bodily fluid sample or a tissue sample.

15. A hydrogel composition for use in diagnostic sensing of an analyte of interest, wherein the hydrogel composition comprises a 3-dimensional (3-D) network structure including (a) hydrophilic polymers as well as (b) allosteric molecules or fragments thereof, wherein the allosteric molecules or fragments thereof assemble into an analyte binding complex upon binding to the analyte of interest, wherein, optionally, at least some of the hydrophilic polymers of the 3-D network structure are connectable of to a metal-coated surface of a sensor, preferably a metal-coated surface of a surface plasmon resonance sensor or a metal-coated surface of a sensor comprising a quartz crystal.

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