Miniaturized infrared fiber probes

Abstract

An example optical sensor includes a first optical fiber defining an axis configured to transmit light; a second optical fiber including a reflection surface formed on an end face of the second optical fiber, wherein the reflection surface is disposed at an angle relative to the axis; and a sleeve or semipermeable membrane configured to fix the second optical fiber to the axis and to fix the reflection surface at a predetermined separation distance from a distal end of the first optical fiber and thereby define a sensing area between the optical fiber and reflection surface; wherein, the sleeve, first optical fiber, and second optical fiber are configured so that when the first optical fiber is illuminated, light from the first optical fiber traverses the sensing area, reflects from the reflection surface, and is returned through the first optical fiber.

Description

Docket Number: 10046-664W01MINIATURIZED INFRARED FIBER PROBES CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims the benefit of U.S. provisional patent application No.63 / 749,262, filed on January 24, 2025, and titled “OPTICAL SENSORS,” the disclosure of which is expressly incorporated herein by reference in its entirety.STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

[0002] This invention was made with government support under Grant no. R21 AA029770 awarded by the National Institutes of Health. The government has certain rights in the invention.BACKGROUND

[0003] Molecular spectroscopy can be used to measure molecules in a sample.Different wavelengths of light can be used for different types of sensing. For example, infrared spectroscopy can be used for molecular analysis configured for biological systems and measurements. Molecular spectroscopy uses light to interrogate a sample and measures the intensity of light as a function of wavelength after interaction with the sample. From the resulting spectrum, properties such as sample composition and analyte concentration can be estimated. Improving molecular spectroscopy allows for improvements to systems and methods that require accurate chemical and biological measurements.SUMMARY

[0004] In some aspects, implementations of the present disclosure include an optical sensor including: a first optical fiber defining an axis configured to transmit light; a second optical fiber including a reflection surface formed on an end face of the second optical fiber, wherein the reflection surface is disposed at an angle relative to the axis; and a sleeve configured to fix second optical fiber to the axis and to fix the reflection surface at a predetermined separation distance from a distal end of the first optical fiber and thereby define a sensing area between the optical fiber and reflection surface; wherein, the sleeve, first optical fiber, and second optical fiber are configured so that when the first optical fiber isDocket Number: 10046-664W01illuminated, light from the first optical fiber traverses the sensing area, reflects from the reflection surface, and is returned through the first optical fiber.

[0005] In some aspects, implementations of the present disclosure include an optical sensor, wherein the reflection surface includes a metallic coating

[0006] In some aspects, implementations of the present disclosure include an optical sensor, wherein the reflection surface includes a gold coating.

[0007] In some aspects, implementations of the present disclosure include an optical sensor, wherein the sleeve includes a sensing window configured to permit a sample fluid to enter the sensing area.

[0008] In some aspects, implementations of the present disclosure include an optical sensor, wherein the sleeve includes a semipermeable membrane

[0009] In some aspects, implementations of the present disclosure include an optical sensor, wherein the semipermeable membrane includes a cellulose-based material.

[0010] In some aspects, implementations of the present disclosure include an optical sensor, wherein the sleeve includes polyetheretherketone.

[0011] In some aspects, implementations of the present disclosure include an optical sensor, wherein the optical fiber includes silver halide.

[0012] In some aspects, implementations of the present disclosure include an optical sensor, wherein the sensing area is less than about 100 micrometers in length between the reflection surface and the distal end of the first optical fiber.

[0013] In some aspects, implementations of the present disclosure include an optical sensor, wherein the sleeve defines a diameter of less than 2 millimeters.

[0014] In some aspects, implementations of the present disclosure include an optical sensor, wherein the angle relative to the axis is 10 degrees or greater.

[0015] In some aspects, implementations of the present disclosure include an optical sensor including: an optical fiber defining an axis configured to transmit light; a reflection surface; a semipermeable membrane configured to fix the reflection surface in place so that the reflection surface is along the axis of the optical fiber and at a fixed distance from the optical fiber.

[0016] In some aspects, implementations of the present disclosure include an optical sensor, wherein the optical fiber includes an angled surface in contact with the reflection surface, and wherein fixed distance from the optical fiber is defined by a length of the angled surface along the axis.Docket Number: 10046-664W01

[0017] In some aspects, implementations of the present disclosure include an optical sensor, wherein the reflection surface includes an optical fiber having a reflective end face.

[0018] In some aspects, implementations of the present disclosure include an optical sensor, wherein the reflection surface includes a metallic coating.

[0019] In some aspects, implementations of the present disclosure include an optical sensor, further including a semipermeable membrane.

[0020] In some aspects, implementations of the present disclosure include an optical sensor, wherein the semipermeable membrane includes a cellulose membrane.

[0021] In some aspects, implementations of the present disclosure include an optical sensor, wherein the optical fiber includes silver halide.

[0022] In some aspects, implementations of the present disclosure include a system including: an optical sensor, a light source configured to illuminate an optical fiber of the optical sensor; a light sensor configured to measure light from the optical sensor; a controller operatively coupled to the light sensor and light source and configured to determine a concentration of at least one analyte based on a measurement of the light sensor.

[0023] In some aspects, implementations of the present disclosure include a system, wherein the concentration includes a concentration of at least one of: a glucose measurement, an ethanol measurement, and a lactate measurement.

[0024] In some aspects, implementations of the present disclosure include a system, wherein the light source is an infrared light source.

[0025] In some aspects, implementations of the present disclosure include a system, wherein the light source is a mid-infrared light source.

[0026] In some aspects, implementations of the present disclosure include a system, wherein the light source is a quantum cascade laser.

[0027] In some aspects, implementations of the present disclosure include a system, wherein the controller is further configured to estimate an absorbance by eliminating reflections, and wherein the concentration is based at least in part on the absorbance.

[0028] In some aspects, implementations of the present disclosure include a system, wherein the light sensor includes an infrared photodetector.

[0029] It should be understood that the above-described subject matter may also be implemented as a computer-controlled apparatus, a computer process, a computing system, or an article of manufacture, such as a computer-readable storage medium.

[0030] Other systems, methods, features and / or advantages will be or may become apparent to one with skill in the art upon examination of the following drawings and detailedDocket Number: 10046-664W01description. It is intended that all such additional systems, methods, features and / or advantages be included within this description and be protected by the accompanying claims.BRIEF DESCRIPTION OF THE DRAWINGS

[0031] The components in the drawings are not necessarily to scale relative to each other. Like reference numerals designate corresponding parts throughout the several views.

[0032] FIG. 1A illustrates a schematic of an example optical sensor, according to implementations of the present disclosure.

[0033] FIG. 1B illustrates a schematic of an example optical sensor including an angled surface, according to implementations of the present disclosure.

[0034] FIG. 1C illustrates a schematic of an example optical sensor including an angled surface, according to implementations of the present disclosure.

[0035] FIG. 2 illustrates a block diagram of an example optical sensor, according to implementations of the present disclosure.

[0036] FIG. 3 illustrates a system block diagram of a system configured to use the optical sensors described with reference to FIGS. 1 A-2, according to implementations of the present disclosure.

[0037] FIG. 4A illustrates a schematic of an optical sensor fabricated in a study of an example implementation of the present disclosure.

[0038] FIG. 4B illustrates an optical sensor fabricated in a study of an example implementation of the present disclosure.

[0039] FIG. 4C illustrates a schematic of an optical sensor fabricated in a study of an example implementation of the present disclosure.

[0040] FIG. 4D illustrates optical sensors fabricated in a study of an example implementation of the present disclosure.

[0041] FIG. 5 illustrates optical sensors fabricated in a study of an example implementation of the present disclosure.

[0042] FIG. 6 illustrates an example experiment including an optical fiber probe according to an implementation of the present disclosure.

[0043] FIG. 7A illustrates a schematic of an optical fiber probe where the optical interaction pathlength is twice the gap distance.

[0044] FIG. 7B illustrates an example transflection fiber probe according to implementations of the present disclosure.Docket Number: 10046-664W01

[0045] FIG. 8A illustrates Mid-IR absorbance spectra of glucose solutions at different concentrations, according to a study of an example implementation of the present disclosure.

[0046] FIG. 8B illustrates Mid-IR absorbance peak height at 1,035 cm'1at different concentrations, according to a study of an example implementation of the present disclosure.

[0047] FIG. 9 illustrates a comparison of an example implementation of the present disclosure to conventional probes.

[0048] FIG. 10 illustrates an experimental setup for optical fiber probe measurements, according to a study of an example implementation of the present disclosure.

[0049] FIG. 11A illustrates normalized mid-IR reflected signals as a function of the distance between silver halide fiber and a mirror in air, according to a study of an example implementation of the present disclosure.

[0050] FIG. 11 B illustrates normalized mid-IR reflected signals as a function of the distance between silver halide fiber and a mirror in water, according to a study of an example implementation of the present disclosure.

[0051] FIG. 12A illustrates a schematic of transflection optical fiber probe where the optical interaction pathlength is twice the gap d, according to a study of an example implementation of the present disclosure.

[0052] FIG. 12B illustrates an example transflection optical fiber probe, according to a study of an example implementation of the present disclosure.

[0053] FIG. 12C illustrates a magnified view of the sensing area and the optical pathlength according to a study of an example implementation of the present disclosure.

[0054] FIG. 13 A illustrates an example signal from optical fiber probes with fiber mirrors coated with different gold thicknesses, with ambient air in the optical pathlength of approximately 80μm according to a study of an example implementation of the present disclosure.

[0055] FIG. 13B illustrates calculated area under the spectral curve for mirrors with different coating thicknesses, according to a study of an example implementation of the present disclosure.

[0056] FIG. 14A illustrates absorbance of glucose solutions at different concentrations between 0 and 200mmol / L measured using the fiber probe with a pathlength of 76μm, and a reference FTIR spectrum of 100mmol / L glucose using 50μm pathlength, according to a study of an example implementation of the present disclosure.Docket Number: 10046-664W01

[0057] FIG. 14B illustrates absorbance at 1035 cm”1for different glucose concentrations, according to a study of an example implementation of the present disclosure.

[0058] FIG. 15A illustrates a schematic of the fiber probe sensor for transflection IR spectroscopy, according to a study of an example implementation of the present disclosure.

[0059] FIG. 15B illustrates a fiber probe with a fiber angle of 10° according to a study of an example implementation of the present disclosure.

[0060] FIG. 16 illustrates an experimental setup for measurements using optical fiber probes and back reflections, according to a study of an example implementation of the present disclosure.

[0061] FIG. 17A illustrates FTIR absorbance spectra of ethanol in aCSF at different concentrations at an optical pathlength of 50 μm, according to a study of an example implementation of the present disclosure.

[0062] FIG. 17B illustrates calibration curve showing absorbance peak height (at 1046 cm'1) versus concentration, according to a study of an example implementation of the present disclosure.

[0063] FIG. 18A illustrates FTIR spectrum of the rat brain dialysate at an optical pathlength of 50 μm, according to a study of an example implementation of the present disclosure.

[0064] FIG. 18B illustrates calibration curve for peak height (at 1046 cm-1) versus concentration, according to a study of an example implementation of the present disclosure.

[0065] FIG. 19 A illustrates absorbance of ethanol solution at different concentrations as measured using transflection fiber probe with an angle of lOo and an averaged optical pathlength of 68 pm, according to a study of an example implementation of the present disclosure.

[0066] FIG. 19B illustrates Fourier-transformed spectra of different concentrations of ethanol solutions, according to a study of an example implementation of the present disclosure.

[0067] FIG. 19C illustrates peak height (at 1046 cm-1) at different concentrations, according to a study of an example implementation of the present disclosure.

[0068] FIG. 20A illustrates the measured ratio of back reflections signal to total reflected signal for different angles as a function of wavenumbers according to a study of an example implementation of the present disclosure.Docket Number: 10046-664W01

[0069] FIG. 20B illustrates the measured ratio of back reflections signal to total reflected signal at 1046 cm-1for different angles according to a study of an example implementation of the present disclosure.

[0070] FIG. 21 A illustrates a regression curve for obtaining parameters for back reflections removal according to a study of an example implementation of the present disclosure.

[0071] FIG. 21B illustrates corrected calibration curve for the ethanol absorbance peak height versus concentration after back reflection removal according to a study of an example implementation of the present disclosure.

[0072] FIG. 22A illustrates peak height (at 1046 cm-1) at various concentrations for different optical pathlengths fiber probes with 0° angled fiber according to a study of an example implementation of the present disclosure.

[0073] FIG. 22B illustrates noise levels (standard deviation of blank sample from 1020-1070 cm-1) of fiber probes with different optical pathlengths. Values at longer optical pathlengths (>157 pm) reflect only system noise due to complete absorption of the signal according to a study of an example implementation of the present disclosure.

[0074] FIG. 22C illustrates SNR of probes with different optical pathlengths for measuring ethanol, according to a study of an example implementation of the present disclosure.

[0075] FIG. 22D illustrates SNR of fiber probes w i th similar- averaged optical pathlength, fabricated with 0° and 10° angled fiber, according to a study of an example implementation of the present disclosure.

[0076] FIG. 23A illustrates an example configuration of a transflection fiber probe according to an implementation of the present disclosure.

[0077] FIG. 23B illustrates a probe according to an example implementation of the present disclosure.

[0078] FIG. 24 illustrates a schematic of the optical setup for the transflection probe measurements, according to a study of an example implementation of the present disclosure.

[0079] FIG. 25A illustrates a schematic of the experimental setup using ex vivo skin. Both the transflection fiber probe and lab-made microdialysis probe were inserted into the dermis and secured in place, with the setup maintained at 37°C. The photograph on the right shows the fiber probe and microdialysis probe inserted in the human skin sample inside the cell culture insert.Docket Number: 10046-664W01

[0080] FIG. 25B illustrates a schematic of the injection experiment, where ethanol was injected into the skin near the probe.

[0081] FIG. 26A illustrates FTIR absorbance spectra of ethanol in aCSF at various concentrations between 2.5 and 40 niM measured with a pathlength of 50 um, according to a study of an example implementation of the present disclosure.

[0082] FIG. 26B illustrates FTIR absorbance spectra of glucose in aCSF at various concentrations between 2.5 and 40 mM measured with a pathlength of 50 pm, according to a study of an example implementation of the present disclosure.

[0083] FIG. 26C illustrates FTIR absorbance spectra of lactate in aCSF at various concentrations between 2.5 and 40 mM measured with a pathlength of 50 pm, according to a study of an example implementation of the present disclosure.

[0084] FIG. 26D illustrates a calibration curve showing peak heights versus concentration for each compound of FIGS. 26A-26C, according to a study of an example implementation of the present disclosure.

[0085] FIG. 27A illustrates absorbance of ethanol in aCSF measured using the mid-IR transflection fiber probe with an averaged optical pathlength of 63 pm,, according to a study of an example implementation of the present disclosure.

[0086] FIG. 27B illustrates absorbance of glucose in aCSF measured using the mid- IR transflection fiber probe with an averaged optical pathlength of 63 pm,, according to a study of an example implementation of the present disclosure.

[0087] FIG. 27C illustrates absorbance of lactate in aCSF measured using the mid-IR transflection fiber probe with an averaged optical pathlength of 63 pm, according to a study of an example implementation of the present disclosure.

[0088] FIG. 27D illustrates calibration curve showing peak height versus concentration for each compound, according to a study of an example implementation of the present disclosure.

[0089] FIG. 28A illustrates absorbance spectrum of a mixture containing 20 mM ethanol, 20 mM glucose, and 20 mM lactate measured using the fiber probe with an averaged optical pathlength of 63 pm, overlaid with the peak deconvolution fit, according to a study of an example implementation of the present disclosure.

[0090] FIG. 28B illustrates absorbance spectra of three different mixtures with varying concentrations of glucose, lactate and ethanol used to validate the deconvolution method, according to a study of an example implementation of the present disclosure.Docket Number: 10046-664W01

[0091] FIG. 29 illustrates Calibration curves for glucose in aCSF obtained using the probe with and without the semi-permeable membrane, according to a study of an example implementation of the present disclosure.

[0092] FIG. 30A illustrates time-course infrared peak height at 1036 cm-1of 20 mM glucose measured using the fiber probe without a semi-permeable membrane, according to a study of an example implementation of the present disclosure.

[0093] FIG. 30B illustrates time-course infrared peak height at 1036 cm-1of 20 mM glucose measured using the fiber probe with a semi-permeable membrane, according to a study of an example implementation of the present disclosure.

[0094] FIG. 31 A illustrates time-course comparison of ethanol measurements in the skin samples using fiber probe and microdialysis where ethanol was spiked into the medium at t = 0 min to achieve a final concentration of 20 mM, according to a study of an example implementation of the present disclosure.

[0095] FIG. 31B illustrates simultaneous monitoring of ethanol, glucose, and lactate in the skin sample, with the medium spiked at 100 mM for each compound, according to a study of an example implementation of the present disclosure.

[0096] FIG. 31C illustrates monitoring of ethanol dynamics in the skin sample following the injection of 50 μL of 200 mM ethanol at t = 0 min where a rapid increase in ethanol concentration was observed, followed by dilution at t = 30 min with the addition of 20 uL of medium, according to a study of an example implementation of the present disclosure.

[0097] FIG. 32 illustrates silver ion concentration released from the silver halide fiber probe with a semi-permeable membrane immersed in ultrapure water over seven days, according to a study of an example implementation of the present disclosure.

[0098] FIG. 33 illustrates an example computing deviceDETAILED DESCRIPTION

[0099] Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the ait. Methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present disclosure. As used in the specification, and in the appended claims, the singular forms “a,” “an,” “the” include plural referents unless the context clearly dictatesDocket Number: 10046-664W01otherwise. The term “comprising” and variations thereof as used herein is used synonymously with the term “including” and variations thereof and are open, non-limiting terms. The terms “optional” or “optionally” used herein mean that the subsequently described feature, event or circumstance may or may not occur, and that the description includes instances where said feature, event or circumstance occurs and instances where it does not. Ranges may be expressed herein as from "about" one particular value, and / or to "about" another particular value. When such a range is expressed, an aspect includes from the one particular value and / or to the other particular value. Similarly, when values are expressed as approximations, by use of the antecedent "about," it will be understood that the particular value forms another aspect. It will be further understood that the endpoints of each of the ranges are significant both in relation to the other endpoint, and independently of the other endpoint. While implementations will be described for measuring glucose, ethanol, and lactate, it will become evident to those skilled in the art that the implementations are not limited thereto, but are applicable for any type of spectroscopy.

[0100] Implementations of the present disclosure are configured to address challenges in the manufacturing and configuration of optical sensors for spectroscopy. In particular, implementations of the present disclosure enable improved transflection sensors. As used herein, “transflection” refers to double-pass absorption, where the light used for sensing is transmitted and received by the same optical fiber.

[0101] Implementations of the present disclosure enable miniaturized optical sensors by using a first optical fiber to transmit and receive optical signals that are reflected off a second optical fiber. The second optical fiber can be angled to optimize the percentage of the signal recovered that is mirror-reflected (and therefore interacted with the analyte) instead of interface reflected. Implementations of the present disclosure further include sleeves configured to cover the sensing area partially or completely, and thereby restrict the analytes introduced into the sensing area (e.g., by a window or membrane that selectively allows certain analytes to enter the sensing area). The resulting miniaturized sensors can enable continuous in-vivo sensing of analytes in sensitive anatomy (e.g., the brain) to enable research into metabolism, addiction, and other biological processes.

[0102] An example optical sensor according to implementations of the present disclosure is shown in FIG. 1 A. In some implementations, the optical sensor is configured as a transflection probe such that light emitted from the first optical fiber 102 passes through the sensing area 110, reflects from the reflection surface 104, and returns through the sensing area 110 toward the first optical fiber 102 for detection. The example optical sensor includesDocket Number: 10046-664W01a first optical fiber 102 with an axis along which light is transmitted through a sensing area 110 and to a reflection surface 104. In such transflection configurations, an effective optical path length through a sample in the sensing area 110 can be approximately twice the separation distance between a distal end of the first optical fiber 102 and the reflection surface 104. The reflection surface 104 can optionally be formed as a surface of a second optical fiber 103. The first optical fiber 102 can optionally be configured for infrared (e.g. mid- or near- infrared) light, and in some implementations comprises silver halide.Optionally, the first optical fiber 102 and / or second optical fiber 103 can be selected from solid-core fibers, hollow-core fibers, and photonic bandgap fibers and can include protective claddings or jackets to improve durability and / or biocompatibility.

[0103] Optionally, the reflection surface 104 can be a reflective coating 112 on the second optical fiber 103. Alternatively or additionally, the reflection surface 104 can include a dielectric mirror stack, multilayer interference coating, or a microfabricated reflective element coupled to the second optical fiber 103. In some implementations, the reflective fiber includes a metallic coating. A non-limiting example of a metallic coating is a gold coating 112. In some implementations, the reflective coating 112 has a thickness in a range of about lOnm to about 5 micrometers. The coating can, for example, be applied by sputtering, evaporation, electroplating, chemical deposition, or any other suitable method.

[0104] A sleeve 106 can be configured to position reflection surface 104 along the axis of the optical fiber 102 and define the sensing area 110 between the optical fiber and reflection surface. In some implementations, the sleeve 106 can include (or be replaced with) a semipermeable membrane (e.g., cellulose) configured to control what substances enter the sensing area 110. The semipermeable membrane can optionally be configured to permit diffusion of one or more target analytes while substantially limiting passage of cells, proteins, or particulates. Alternatively or additionally, the sleeve 106 can include a plastic structure (e.g., PEEK, or polyetheretherketone).

[0105] Alternatively or additionally, the sleeve 106 can include a sensing window 107 cut into the sleeve 106 to allow fluid to enter the sensing area. It should be understood that some implementations of the present disclosure can include more than one sensing window as well as combinations of sensing window and membranes.

[0106] With reference to FIG. IB, the present disclosure contemplates that the first optical fiber 102 can include an angled surface 102b. The angled surface 102b can be configured to reduce Fresnel back reflections and interferometric artifacts at one or more interfaces while maintaining coupling efficiency into a return path. The plane of the angledDocket Number: 10046-664W01surface 102b can be angled relative to the axis 150 of the first optical fiber 102 and second optical fiber 103 as shown by angle θ in FIG. 1B. For example, the angle can be a 10-degree angle from perpendicular with the axis 150. However, the present disclosure contemplates that the angle θ of the angled surface 102b can be any angle. For example, the angle can be adjusted to optimize the percentage of light received at the first optical fiber that is reflected from the second optical fiber. In some implementations, angle is selected in a range of about 1 degree to about 30 degrees.

[0107] With reference to FIG. 1C, the present disclosure further contemplates forming the angled reflection surface on the second optical fiber 103, so that the reflection surface is an angled reflection surface 103b. The plane of the angled reflection surface can optionally be angled at about 10 degrees from axis 150.

[0108] Some implementations of the present disclosure include optical sensors configured using semipermeable membrane and / or angled surfaces. With reference to FIG. 2, an example optical sensor is shown including the optical fiber 102, reflection surface 104, coating 112, and sensing area 110 described with reference to FIG. 1A. Again, the reflection surface 104 can be formed using a second optical fiber in some implementations, as described with reference to FIGS. 1A-1C. In the implementation of FIG. 2, the reflection surface 104 can be configured with an angled surface 104b that contacts the optical fiber 102 and defines the size of the sensing area by the spacing between the optical fiber 102 and reflection surface 104.

[0109] Optionally, a semipermeable membrane 202 can be used to secure the optical fiber 102 and the reflection surface 104, where the semipermeable membrane allows fluid to enter the sensing area. The semipermeable membrane 202 can optionally be attached to or sealed to the reflection surface 104 and optical fiber 102 by an adhesive 201. Alternatively or additionally, the optical fiber 102 and reflection surface 104 (e.g., another segment of optical fiber) can be secured using a sleeve with a hole configured to allow fluid to enter the sensing area disposed between the optical fiber 102 and the reflection surface 104. The sleeve can optionally be adhered to the optical fiber 102 and / or reflection surface 104 by the adhesive 201. The sleeve can optionally be in addition to, or in place of, the semipermeable membrane 202.

[0110] In some implementations, the membrane can be joined to the optical fiber 102 and / or reflection surface 104 by an adhesive. The adhesive can optionally be configured to seal the sleeve or semipermeable membrane of implementations of the present disclosure to the optical fiber 102 and / or reflection surface 104.Docket Number: 10046-664W01

[0111] With reference to FIG. 3, implementations of the present disclosure further include systems for molecular spectroscopy, wherein the optical sensors disclosed herein can be operably coupled to light sources and light sensors to determine the presence and / or absence of various chemical and biological compounds, for example ethanol, glucose, and lactate. In an example measurement system, an optical sensor 300 according to any implementation of the present disclosure is operatively coupled to a light sensor 302 and a light source 304 configured to illuminate an optical fiber (not shown) of the optical sensor 300 (e.g., using one or more fiber optic couplings).

[0112] The optical sensor 300 can optionally be implanted in a subject. The light sensor 302 can measure the light returned by the optical fiber of the optical sensor 300. The measurement of light returned by the optical fiber of the optical sensor 300 can be an input to a controller 306, which can include any or all of the features of the computing device 3300 shown in FIG. 33.

[0113] The controller 306 can optionally be configured to perform any of the methods described in Examples 1 -4 herein to determine the concentration of an analyte in the optical sensor 300. For example, the light sensor 302 can measure a spectrum of the received light from the light sensor 302 and the controller 306 can determine an absorbance spectrum by comparing the received spectrum to the background spectrum (e.g., a measurement of the optical sensor 300 without a sample in the sensing area). The controller 306 can optionally be further configured to remove back reflections and / or interference artifacts from the received signal by estimating and subtracting back-reflection term and / or by fitting a regression model.

[0114] The controller 306 can further be configured to compare the absorbance of the sample to one or more analyte-specific bands and determine a concentration of an analyte by a calibration. The calibration can optionally be a linear regression for a single-analyte solution. Alternatively or additionally, the calibration can include peak deconvolution or multipeak fitting to separate overlapping spectral contributions for samples with multiple analytes. Non-limiting examples of calibration for different analytes and solutions are provided in the examples hereto. Alternatively or additionally, the calibration can include multivariate or machine-learning models (e.g., partial least squares regression, principal component regression, support vector regression, or neural-network regression).

[0115] Examples:Docket Number: 10046-664W01

[0116] Optical sensors according to implementations of the present disclosure were designed and tested. FIG. 4A illustrates a schematic of an example implementation of the present disclosure, and FIGS. 4B and 4C illustrate views of the example optical sensor including a 500 micrometer diameter silver halide fiber with PEEK sleeve. In FIGS. 4A-4C, two fibers are facing each other, one fiber delivers light from light source to the sensing area and then back to photodetector. The second fiber, placed at a particular distance corresponding to optical pathlength, is coated with gold to enable reflection. In this case the optical pathlength is twice the distance between the two fibers. A PEEK sleeve is used to keep the two fibers in front of each other and at a fixed distance from each other. In this example design, a semiporous membrane was not implemented, the semiporous membrane can optionally be placed around the fibers or around the whole assembly.

[0117] FIG. 4D illustrates an example implementation including two fibers facing each other, with 200 micrometer silver halide fibers coupled by a semipermeable membrane. One can be gold coated. The distance between the two fibers can be adjusted under the microscope and gluing it to the fibers will allow for fixing the pathlength at a particular distance.

[0118] FIG. 5 illustrates two fibers facing each other, one fiber delivers light from light source to the sensing area and then back to photodetector. This fiber has an angle. This angle allows to place one fiber next the opposite fiber and to define the pathlength. The second gold-coated fiber enables light reflection back. A transparent thin membrane that holds the fibers together is the semipermeable cellulose membrane.

[0119] This design overcomes the manufacturing challenge of manufacturing a precise and controlled optical pathlength by allowing placement of the two fibers against each other.

[0120] Example 1:

[0121] Additional example implementations were designed and tested in a study. Mid-infrared (mid-IR) optical fiber sensors offer highly specific and sensitive detection and analysis of various chemical species due to many molecular vibrations and fundamental absorption bands in this range. This example illustrates an example implementation of the present disclosure enabling controlled optical pathlength. The example optical fiber probe was fabricated using a silver halide polycrystalline fiber, a gold-coated short fiber as a mirror, and a connector to align the two parts to face each other. The outer diameter of the connector, 1.59 mm, dictates the overall probe diameter. To demonstrate the sensing performance, aDocket Number: 10046-664W01quantum cascade laser (QCL) was coupled to the optical fiber probe to measure glucose solutions at physiological concentration levels by monitoring the C — 0 stretching vibration at 1,035 cm-1, A detection limit of 8.91 mM for glucose was achieved. The results highlight the potential of the example optical fiber probe for molecular detection and analysis, offering a promising solution for chemical and biomedical applications.

[0122] The mid-infrared (mid-IR) spectral region, spanning wavelengths from 3 to 25 μm, holds significant potential for chemical sensing due to the presence of strong and characteristic absorption bands of molecular vibrations. These absorption features enable the detection and quantification of various analytes with high specificity and sensitivity, making mid-IR sensing technologies crucial in fields such as biomedical diagnostics, environmental monitoring, and industrial process control. Among these, fiber-optic sensors have gained significant attention for their versatility and ability to perform in situ and provide real-time measurements.

[0123] Evanescent wave sensing (EWS) is one of the commonly used types of fiber¬ optic sensors, where light propagates through the core of a fiber and interacts with the surrounding medium through the evanescent field. This approach has been widely employed in mid-IR sensing for detecting gases, liquids, and biomolecules. Other approaches involve refractive index measurements and localized surface plasmon resonance (LSPR). However, miniaturization of EWS sensors is challenging due to limited sensitivity. Moreover, refractive index sensing and LSPR, are inherently non-selective and require (bio)chemical functionalization of the fiber surface to achieve selectivity. This adds complexity to the sensor fabrication and is prone to degradation over time, which can limit the long-term stability and reliability of the sensor.

[0124] This study presents an improved mid-IR transflection optical fiber probe based on silver halide optical fibers. Conventional mid-IR transflection probes are large in size and not suitable for in vivo sensing and / or for applications where probe's size is important. The example implementations described herein enable transflection -based optical fiber probes that can operate in the mid-IR region and are small enough to be suitable for in vivo sensing. In the example probe configuration, the light travels through the solution twice-once during transmission and once after reflection from the fiber-mirror. The probe enables direct interaction with the sample over a defined optical pathlength, which can be precisely specified during the fabrication process. To demonstrate the performance of the exampleDocket Number: 10046-664W01probe, aqueous glucose solutions were used, highlighting its potential as a versatile platform for mid-IR chemical analysis.

[0125] Materials

[0126] The example implementation used polycrystalline mid-infrared fiber, which is made of AgCl — AgBr for optical fiber probe fabrication. The silver halide fiber (refractive index 2.1, NA 0.3 ) with a diameter of 500μm and IR transparency wavelength range from 3 to 17μm. D-glucose was used for preparation of the standard solutions in deionized water having concentrations between 0 and 200mmol / L('mM).

[0127] Experimental setup

[0128] FIG. 6 shows the schematics of the experimental setup for optical fiber probe measurements. A mid-IR quantum cascade laser was used as the light source, providing a spectral range of 800 -- 1,800 cm”1(5,556 − 12,500μm). The collimated laser beam was directed into a beam splitter using mirrors. After the beam splitter, an off-axis parabolic mirror (35-643, Edmund Optics) focused the beam onto the optical fiber probe. An optical fiber clamp was used to secure the optical fiber probe and was fixed on a 3-axis stage to adjust the position for signal optimization. The infrared light propagates through the optical fiber probe and interacts with the sample solution in the sensing area, which is between the optical fiber and the gold-coated fiber that acts as mirror in the probe sensor. The reflected signal is sent back through the fiber and directed to the detector. The zoomed-in figure shows that the optical fiber probe was inserted into a 250μL centrifuge tube, containing the analyte solution, until the sensing area was fully submerged in the liquid when performing measurement for glucose.

[0129] RESULTS

[0130] Assembly of the optical fiber probe

[0131] FIG. 7A shows the schematics of the optical fiber probe design in transflection mode. The probe sensor includes an optical fiber, a short piece of gold-coated silver halide fiber which is used as mirror, and a connector to hold and align the fiber and the mirror facing each other. The gold-coated surface was polished using aluminum oxide lapping film prior to gold sputtering. This polishing process ensured a smooth surface, minimizing signal loss due to light scattering. The connector is made of a 2 cm long hollow polyetheretherketone (PEEK) tube with an inner diameter of 0.51 mm and an outer diameter of 1.59 mm. A 3 mm wide groove in the connector was made using a hand drill, while a needle was inserted into the tubing during the drilling process to prevent the hole fromDocket Number: 10046-664W01collapsing. The mirror was inserted with its coated face oriented toward the cleaved optical fiber. The optical path length - two times of the distance between the fiber and the mirror -was adjusted under a microscope after inserting the fiber and mirror into the PEEK connector. FIG. 7B illustrates the fabricated optical fiber probe.

[0132] Measurement of glucose at different concentrations

[0133] The measurement of glucose solutions at various concentrations demonstrates the performance of the optical fiber probe. An optical fiber probe with a distance of 38μm between two fibers was fabricated. FIG. 8A shows the absorbance spectra of glucose solutions at different concentrations, using water as background. For reference, the spectrum of a 100 niM glucose aqueous solution, measured with a FTIR with a 50μm optical pathlength, was also included. The presence of back reflection in both single beam measurement of water and glucose, used for calculating the absorbance spectrum, introduced extra absorption peaks resulting from the material of the fiber. Each spectrum represents an average of three measurements.

[0134] The peak at 1,035 cm-1, corresponding to C — 0 stretching, was used for concentration quantification. FIG. 8B shows the peak heights at various concentrations. The absorbance at 1,035 cm-1, calculated by taking the logarithm of the single-beam intensity of water divided by the glucose solution, was used as the peak height for concentration determination since the baseline absorbance was close to zero. The trend of peak heights across concentrations was not linear due to back reflection occurring at the air-fiber interface, which affected both the single-beam spectra of the background and the glucose solutions, as described in Equation (1):SW(A) + RW(A)Abs(A,c) = login... ^-— (1)

[0135] where Rwand Rgrepresent the back reflection from the end surfaces when measuring water and glucose solutions, respectively, while Swand Sgrepresent the obtained single-beam signals.

[0136] Based on Beer-Lambert’s law and by assuming back reflection is constant for water and glucose at different concentrations absorbance at a specific wavenumber, equation (1) can be expressed as:Abs(c) = log10C0 / (Sw× 10−C₁·c+ R) (2)Docket Number: 10046-664W01

[0137] where C0is the total signal for water measurement at 1,035 cm-1(including both water and back reflection signals), c is the concentration of glucose solution, C1and R are constants.

[0138] A three-parameter regression model was used to fit the experimental data, as shown in FIG. 8B to derive the calibration curve corresponding to equation (2). This curve enables the calculation of unknown concentration of the glucose solution. The limit of detection is around 8.91 mM, determined as three times the standard deviation of the peak height of water over the slope around the origin, calculated using the fitting curve in FIG. 8B.

[0139] The example optical fiber probe was successfully fabricated in transflection mode, utilizing polished mid-IR silver halide fiber, a lab-made PEEK connector, and a gold- coated polished fiber as a reflective mirror. To validate the probe's performance, glucose solutions with concentrations ranging from 3.13 mM to 200 mM were measured using an optical fiber probe with an optical pathlength of 76μm (corresponding to a distance of 38μm between the two fibers). The calibration curve, affected by back reflections at the fiber-air interface, was modeled using a 3-parameter regression approach, yielding a detection limit of approximately 8.91 mM. The present disclosure contemplates further reducing back reflection to further improve probe performance and to miniaturize the design by developing more compact connectors capable of integrating the fiber and mirror for in-vivo sensing.

[0140] Example 2:

[0141] An additional example implementation of an optical fiber probe configured for transflection mode that can address the challenges of in vivo measurements was constructed and studied.

[0142] The example implementation can include an optical fiber probe in transflection mode to address the challenges of in vivo measurements by combining the delivery and collection of light within a single fiber. In this configuration, light interacts with the analyte in a predefined sensing area and is reflected back to the same fiber, eliminating the need for separate fibers for transmitting / receiving or the use of a bent or a long fiber. This design minimizes tissue damage for in vivo applications. Commercially available mid-IR transflection probes are typically large in size (on the order of centimeters) and thus not suitable for in vivo sensing and / or for applications that require small probe dimensions. The example implementation includes a mid-IR transflection optical fiber probe sensor that is small and sensitive (overall diameter of 1.59 mm). The example implementation is thereforeDocket Number: 10046-664W01suitable for in vivo sensing. In the example probe configuration, the light travels through the sample / solution twice-once during transmission and once after reflection from a gold-coated fiber acting as a mirror. The probe enables direct interaction with the sample over a defined optical pathlength, which can be precisely controlled during the fabrication process. The schematics of the mid-IR transflection optical fiber probe and how it compares to existing probes are provided in FIG. 9. To demonstrate the performance of the example probe, aqueous glucose solutions at physiological concentration levels were measured by monitoring the C — 0 stretching vibration at 1035 cm”1(wavelength of ~ 9.7 g m ), demonstrating its potential for mid-IR chemical analysis. In addition to presenting the probe design and its sensing application, the example implementation and study also identified the source of signal interference and developed an improved regression model for calibration-aspects.

[0143] Materials and Methods

[0144] Materials

[0145] Polycrystalline mid-infrared fiber made of AgCl — AgBr (PIR 500, Art Photonics GmbH) was used for the optical fiber probe fabrication. The silver halide fiber with a diameter of 500g m has an IR transparent wavelength range from 3 to 17 m, numerical aperture (NA) of 0.3, and core refractive index of 2.1. d-glucose ( > 99.5%, G7021 from Sigma Aldrich) was used for the preparation of the standard solutions in deionized water having concentrations between 0 and 200mmol / L.

[0146] Measurement of Gold Coating Thickness

[0147] In the example implementation, the end face of an optical fiber was coated with gold to serve as a reflective mirror, which is described in detail in the later section. Different coating duration time was applied to the fiber end to investigate the optimal thickness of the coating layer. The thickness of the gold layer was determined by placing a glass slide next to the fiber during the sputtering process. After sputtering, a scratch was made on the gold-coated surface of the glass slide using a razor blade, and the depth of the resulting valley was measured using an optical profilometer.

[0148] Fourier-Transjbrm Infrared Spectroscopy

[0149] To compare the spectra measured using the optical fiber probe and a standard instrument, a Fourier-transform infrared (FTIR) spectrometer equipped with a transmission cell with an optical pathlength of 50μm and requiring a minimum sample volume of 10μL was used. Each spectrum was acquired for 300 scans, yielding a total acquisition time of 187 s, with a spectral resolution of 0.482 cm”1.Docket Number: 10046-664W01

[0150] Experimental Setup

[0151] FIG. 10 shows the schematics of the experimental setup for optical fiber probe measurements. A mid-IR QCL was used as the light source, providing a spectral range of 800 to 1800 cm-1(5556 to 12500 nm). The collimated laser beam was directed into a beam splitter using three mirrors. After the beam splitter, an off-axis parabolic mirror focused the beam onto the optical fiber probe. An optical fiber clamp was used to secure the optical fiber probe and was fixed on a three-axis stage to adjust the position for signal optimization. The infrared light propagates through the optical fiber probe and interacts with the sample solution in the sensing area, which is between the optical fiber and the mirror in the probe sensor. The reflected signal is sent back through the fiber and directed to the detector. The zoomed-in figure shows the optical fiber probe inserted into a 250 − μL centrifuge tube, containing the analyte solution, with the sensing area fully submerged in the liquid when performing measurement for glucose.

[0152] A cleaved fiber was used instead of the optical fiber probe when quantifying the reflection signal. A mirror made from a piece of gold-coated silicon wafer, mounted on a translational stage, was positioned directly behind the fiber's end face, allowing the reflection signal to be measured at varying distances from the mirror.

[0153] Results and Discussion

[0154] Estimation of the Optical Pathlength for Optical Fiber Probe

[0155] One of the advantages of the example design is the ability to fabricate the optical fiber probe sensor with a specific optical pathlength. The optical pathlength directly impacts the sensor's performance based on Beer-Lambert’s law. Longer pathlengths increase absorbance, therefore enhancing the sensor's sensitivity at low analyte concentrations. But if too long, light emerging from the fiber diverges as it propagates, consequently, less light is reflected back to the detector. Determining the optimal optical pathlength is particularly important for aqueous samples due to the strong absorption of mid-IR light by water. Optical pathlengths, over ~ 100μm, can result in complete absorption of the light by the sample, leaving no detectable signal reflected back to the detector. To estimate the optimal optical pathlength, a cleaved optical fiber and a mirror made from a gold-coated silicon wafer were used. Infrared light was introduced at one end of the fiber, and the reflected signal from the opposite end was measured at varying distances from the mirror.

[0156] FIG. 11 A shows the reflected signal from the optical fiber probe sensor, showing the summed signal in the 1000 to 1180 cm-1region at various distances from aDocket Number: 10046-664W01mirror in the air. The reflected signal was normalized by subtracting the minimum signal (recorded when the mirror was far away) and dividing by the difference between the maximum signal (recorded when the mirror was almost attached to the fiber) and the minimum signal. The reflected signal decreased almost linearly as the distance from the mirror increased from 0 to 1000μ m. At a distance of 560μ m, the signal dropped to 50% of its maximum value and was further reduced to 10% at a distance of 1023μ m.

[0157] When the optical fiber probe sensor was immersed in water, FIG. 1 IB shows the reflected signal decreased sharply compared to when it was in air. The signal declined significantly when the distance between the fiber end-face and the mirror exceeded 40μ m (corresponding to a total optical pathlength of 80μ m ) and dropped to half its maximum value at a distance of 53μ m. This is explained by the strong IR light absorption by the water.

[0158] This data enables designing pathlengths for both gaseous (in air) and aqueous samples.

[0159] Assembly of the Optical Fiber Probe Sensor

[0160] FIG. 12A shows the schematics of the optical fiber probe design in transflection mode. The probe sensor consists of an optical fiber, a short piece of gold-coated silver halide fiber, which is used as a mirror, and a connector to hold and align the fiber and the gold-coated fiber facing each other. The surface of the shorter fiber was polished using aluminum oxide lapping film prior to gold sputtering. This polishing process ensured a smooth surface, minimizing signal loss due to light scattering. The connector is made of a 1-cm-long hollow polyetheretherketone (PEEK) tube with an inner diameter of 0.51 mm and an outer diameter of 1.59 mm. A 3-mm-wide groove in the connector was made using a hand drill, whereas a needle was inserted into the tubing during the drilling process to prevent the hole from collapsing. The mirror was inserted with its coated face oriented toward the cleaved optical fiber. The optical pathlength — 2 times the distance between the fiber and the mirror-was adjusted under a microscope after inserting the fiber and mirror into the PEEK connector. FIG. 12B shows the fabricated optical fiber probe. FIG. 12C shows a close-up view of the sensing area and the optical pathlength.

[0161] Optimization of the Gold Thickness of the Mirror Coating

[0162] The transflection optical fiber probe sensor relies on light interacting with the analyte in the sensing region having a total optical pathlength of 2d and carrying the spectral information after reflecting off the mirror. Consequently, the quality of the mirrorDocket Number: 10046-664W01significantly affects the sensor’s performance. To optimize the reflected signal from the mirror, the optical fiber probe with various gold coating thicknesses was investigated.

[0163] Short fibers with flat end faces were first polished and then sputtered with gold as fiber mirrors for durations of 0,60,90,180,300, and 600 s, corresponding to coating thicknesses of ~ 0,35,54,122,175, and 267 nm. FIG. 13A presents the reflected spectra measured using the optical fiber probe assembled with fiber mirrors of different gold coating thicknesses. The distance between the two fibers, the cleaved fiber and the fiber mirror, was ~ 40μ m, with the gap filled with ambient air for all measurements. The yellow spectrum at the bottom was measured using an optical fiber probe with a mirror made from a cleaved fiber without a gold coating. As the gold coating thickness increases, the reflection signal improves significantly.

[0164] The obtained spectra consist of multiple peaks originating from back reflections at the fiber-air interface before the light was reflected off the mirror. It is also due to multiple reflections within the small gap between the two fibers. To better understand the periodic modulations observed in the spectral data, a Fast Fourier Transform (FFT) was performed on the spectra in FIG. 13A. The dominant frequency was identified at 0.0218, which corresponds to an optical pathlength of ~ 109μ m. This value is close to the optical pathlength of the probe, ~ 80μ m (twice the air gap between the fiber tip and the mirror), which supports that the modulations result from multiple reflections within the small cavity formed between the fiber end face and the mirror.

[0165] To further evaluate the reflectance in the 1000 to 1180 cm-1region (a range critical for detecting absorption peaks of molecules such as glucose), the area under each spectral curve in FIG. 13A was calculated. This analysis provides a quantitative measure of the mirror's performance. FIG. 13B shows the area under each spectral curve for different coating thicknesses. The reflected signal initially increased significantly with the thickness of the gold coating and then reached a plateau at a thickness of around 170 nm. A previous study showed a maximum reflectance of 90.5% using a gold coating thickness of 153 nm, and a 88.5% reflectance using a gold coating thickness of 235 nm These findings aligned with the measurements herein. To ensure optimal performance, in the subsequent sensors of the present example, the mirror of the fiber sensor was fabricated using fibers coated with gold to a thickness of ~ 267 nm.

[0166] Measurement of Glucose at Different ConcentrationsDocket Number: 10046-664W01

[0167] The measurement of glucose solutions at various concentrations demonstrates the sensing performance of the optical fiber probe. An optical fiber probe with a distance of 38μ m between two fibers (optical pathlength of 76μ m) was fabricated to maximize sensitivity while avoiding total absorption of signal due to water absorption. FIG. 14A shows the absorbance spectra of glucose solutions at different concentrations, using water as a background. Each spectrum represents an average of three measurements. The spectrum of a 100 — mmol / L glucose aqueous solution, measured with an FUR with a 50 − μm optical pathlength, was also included for reference. The discrepancy between the glucose spectra obtained from the fiber probe and the FTIR can be attributed to the presence of interference patterns caused by multiple reflections within the sensing cavity. Furthermore, because the absorbance is calculated using signals that include back reflections, glucose-induced spectral features are partially masked according to Eq. (1). When the magnitude of the back reflection becomes larger, the net absorbance can be significantly reduced, decreasing the sensitivity. These two factors, interference and back reflection, distort the overall spectral shape and are absent in the FTIR reference spectrum.

[0168] To verify the origin of the short-period modulation, the same FFT analysis described in the previous section was applied to the spectrum in FIG. 14A. The dominant frequency component was identified at 0.0216, which corresponds to an optical pathlength of ~ 81μ m. This value is close to the designed optical pathlength of 76μ m for the optical fiber probe, indicating that the short-period modulations observed in FIG. 14A originate from multiple reflections within the small cavity between the fiber and the gold-coated fiber mirror.

[0169] The absorbance at 1035 cm"1, corresponding to C — 0 stretching, was used for concentration quantification. FIG. 14B shows this absorbance at various concentrations. The absorbance at 1035 cm"1at different concentrations was not linear as shown in FIG. 14B primarily due to the presence of back reflection present in both water and glucose spectra. Specifically, the signal received by the detector is a combination of analyte-specific signals reflected from the gold-coated mirror and from back reflections at the fiber-liquid interface. In order to account for the influence of back reflections, the absorbance calculated from the spectra collected with the optical fiber probe can be expressed asSH, (A) + Rw(A)Abs(A C) = log10* 'P rWSg(A, c) + Rg(A, c)Docket Number: 10046-664W01

[0170] where Rwand Rgrepresent the back reflection from the end surfaces when measuring water and glucose solutions, respectively, whereas Swand Sgrepresent the obtained unprocessed spectra.

[0171] Table 1 A Back reflection signals at different concentrations of glucose in aqueous solution.Concentration (mmol / L) Intensity at 1035 cm1(a. u.) 0 1.032 ± 0.03750 1.041 ± 0.023100 1.005 ± 0.041200 1.047 ± 0.028

[0172] All parameters are functions of wavenumber or wavelength 2, whereas Sgand Rgalso vary with glucose concentration c.

[0173] At a specific wavenumber of 1035 cm’1, Sgand R depend only on concentration. Experimental observations show that the back reflection signal varies by less than 5% (at 1035 cm’1) when measuring water and glucose at different concentrations as summarized in Table 1A.

[0174] As a result, it can be assumed that Rw, Re, and R are constants, and equal to each other, in the range of 1000 to 1200 cm”1, and thus— Rg — R- (2)

[0175] Based on Beer's Lambert's law, ideally, the absorbance value at 1035 cm”1is linearly proportional to glucose concentration c. Therefore, the glucose absorbance signal can be expressed asAbs(c)ideal= log10= Q ■ c (3)

[0176] and thereforeSg(c) ~ Swx 10~c^.(4)

[0177] where is a constant. By substituting experimental data and Eq. (4) into Eq. (1), the rearranged equation for absorbance becomesCoAbs(c) = l08<" ^^^

[0178] where Cois the total signal for water measurement at 1035 cm”1(including both water and back reflection signals).Docket Number: 10046-664W01

[0179] A three-parameter regression model was used to fit the experimental data, as shown in FIG. 14B to derive the calibration curve. This curve enables the calculation of the unknown concentration of the glucose solution.

[0180] The limit of detection is around 8.91mmol / L, determined as 3 times the standard deviation of the absorbance of water at 1035 cm-1over the slope around the origin, calculated using the fitting curve. This concentration is in the physiological glucose levels observed after eating or during acute hyperglycemia following traumatic brain injury, typically around lOmmol / L For physiological glucose sensing applications, the limit of detection needs to be further improved, preferably down to 0.5mmol / L. In addition, in complex biological fluids such as blood, the LOD would likely be further compromised due to overlapping infrared absorption from proteins, lipids, and other endogenous compounds. Future improvements to the probe will focus on enhancing detection performance through several approaches: first, optimization of the optical design, which includes minimizing back reflections and optimizing optical alignment; second, incorporation of a semi-permeable membrane, which would exclude large interfering biomolecules, such as proteins, from the sensing region; and finally, applying signal processing techniques, such as spectral smoothing and multivariate calibration (e.g., partial least squares regression, PLSR), may further improve detection limit.

[0181] Discussion

[0182] An optical fiber probe sensor was successfully fabricated according to an implementation of the present disclosure to operate in transflection mode, utilizing a mid-IR silver halide fiber, a lab-made PEEK connector, and an additional gold-coated silver halide fiber to act as a reflective mirror. The overall diameter of the probe is 1.59 mm, which is smaller than conventional transflection probe designs. For optimal performance, a 267 — nm gold coating was applied to the fiber end face to maximize signal reflection back to the detector. For aqueous samples, a distance of less than 50 / / m between the fiber and the fiber mirror is recommended due to the divergence of light exiting the fiber and the strong absorption of water in the mid-infrared region. To validate the probe's performance, glucose solutions with concentrations ranging from 3.13 to 200mmol / L were measured using the probe with an optical pathlength of 76p m (corresponding to a distance of 38 / z m between the fiber end and the fiber mirror). A regression model was developed to address the nonlinearity of the calibration curve, caused by back reflection at the fiber-air interface, yielding a detection limit of ~ 8.91mmol / L. Due to the high selectivity of the mid-IR range,Docket Number: 10046-664W01the probe can also be used to detect other analytes instead or in addition to glucose. The present disclosure contemplates further reducing the overall size of the probe and enhance the probe's performance by optimizing the optical design to reduce back reflections, integrating a semi-permeable membrane to block interfering biomolecules, and applying signal processing techniques to extend the probe's applicability for complex biological fluids.

[0183] Example 3:

[0184] An example implementation of the present disclosure was configured as a real-time in vivo ethanol monitor. Real-time in vivo ethanol monitoring is important for understanding its physiological mechanisms and developing treatments for alcohol-related disorders. Mid-infrared (mid-IR) fiber sensors enable highly specific and sensitive detection of chemical species due to unique molecular vibrational absorption bands. This study presents an example mid-IR transflection optical fiber probe that is small enough to be suitable for in vivo monitoring. Ethanol quantification using mid-IR spectroscopy was demonstrated in synthetic solutions and in a biofluid, rat brain dialysate, spiked with ethanol at physiological concentrations using standard Fourier-transform infrared spectrometer (FTIR). The results showed a linear relationship between absorbance and concentration, with a limit of detection (LoD) of -4 mraol / L in each fluid. The optical probe was fabricated using a 500 pm silver halide polycrystalline fiber aligned with a gold-coated short fiber, acting as a mirror, in a connector to maintain a predetermined distance between two fibers. The sample cavity was formed between the optical fiber and the gold-coated fiber. For performance optimization, optical back reflections were analyzed experimentally and numerically in fibers with end-face surfaces polished at varying angular deviations from the fiber axis.Experimental data were used to estimate and subtract back reflections from the detected signals to improve the signal-to-noise ratio (SNR) resulting in an LoD of 4 mmoi / L for ethanol in synthetic solutions. The effects of optical pathlength and fiber end-face angle on probe’s performance were investigated, showing that a fiber with an end-face whose surface normal deviated by 10° from the fiber axis (averaged optical pathlength of 68 pm) achieved a 32.7% higher SNR compared to fiber with a surface normal aligned with the fiber axis (optical pathlength of 64 pm). This study is important for optimizing the sensor design, enabling advanced biomedical applications such as monitoring alcohol levels in tissues and detecting ethanol in various clinical scenarios.

[0185] Ethanol consumption significantly contributes to numerous fatalities and diseases, and only three pharmacotherapies are approved by the Food and DrugDocket Number: 10046-664WO1Administration (FDA) in the United States. The development of new therapeutics for alcohol use disorder requires continued progress in the elucidation of the underlying neuropharmacological mechanisms in the brain after ethanol consumption. A limitation in the field is the ability to monitor ethanol in tissues (including brain) in real-time so that the experimenter can interpret any findings of neurophysiological or neurochemical changes produced by ethanol. It is important to monitor local ethanol concentrations in the brain tissue before, during, and after ethanol consumption (as well as other tissues). Previous studies have explored various technologies to quantify ethanol levels in the rodent and mice brain, including the use of magnetic resonance spectroscopy to measure ethanol concentrations in rat brains. Proton magnetic resonance spectroscopy, in particular, was used to measure the uptake and clearance of ethanol in the brain following intraperitoneal or intragastric alcohol injection. However, this method requires animals to be under anesthesia, which limits its application in studies such as behavioral research. Additionally, this technique suffers from low SNR ratio due to the weak magnetic moments of protons and the limited sample volume available in in vivo applications. It also requires relatively long acquisition times, particularly for high-resolution spectra.

[0186] Microdialysis is a widely used method for measuring ethanol concentration in the brain. Microdialysis is a sampling technique that uses a semi-permeable membrane to measure the concentration of analytes in extracellular fluid by allowing molecules to diffuse from tissue into the sampling probe. By perfusing the probe, analyte concentration in the tissue can be determined based on the analyte concentration in the collected sample.However, several limitations restrict its broader application. Achieving high recovery of the analyte requires a slow flow rate (typically less than 2 gL / min) due to the diffusion-based nature of this method, which leads to relatively poor temporal resolution (usually around or longer than 10 minutes). Moreover, during sampling, analyte molecules that diffuse into the probe are gradually depleted from the tissue, potentially resulting in misinterpretation of experimental data. Furthermore, for in vivo quantification, the technique requires calibration process performed in vivo to achieve accurate results.

[0187] Mid-infrared (mid-IR) fiber sensors have gained significant attention due to their non-destructive nature and unique ability to operate in a spectral region where many molecular vibrations and fundamental absorption bands occur. This spectral region, typically ranging from 2 to 20 micrometers, provides a powerful means for detecting and analyzing a wide array of chemical species with high selectivity and sensitivity. However, challenges remain in the development and deployment of mid-IR fiber sensors. Most designs rely on9Docket Number: 10046-664W01evanescent wave sensing, which often require either coupling two fibers to a sensing head or using a long fiber segment with an exposed core region for sensing - making probe miniaturization difficult. This limitation is especially challenging for in vivo applications, where minimizing probe size is critical to reduce tissue damage. Furthermore, the optical pathlength of the evanescent field is typically short, which results in lower sensitivity compared to absorption-based measurements. Refractive index sensing based on the evanescent wave field lacks chemical specificity and requires biochemical surface functionalization, adding additional drawbacks and complexities.

[0188] One approach to address these challenges is the use of an optical fiber probe based on transflection mode, where the sample is measured with a predetermined optical pathlength in transmission mode after reflecting from a mirror back to the same optical fiber. This approach enables a more compact probe design, making it suitable for precise in vivo positioning within a tissue. This example discloses a novel sensor probe design in mid-IR transflection mode, featuring an angled fiber end-face to minimize back reflection which enhances its SNR. Tills configuration allows for easy adjustment of optical pathlength to suit various applications. In this work, the performance, in terms of SNR, of the fiber probe with various optical pathlengths and fiber angles was investigated. Additionally, the quantitative determination of ethanol concentration in aqueous solutions was demonstrated. The example implementation described herein represents a significant advancement toward compact and efficient mid-IR fiber sensors for a wide range of applications, including biomedical sensing and chemical analysis.

[0189] Materials and Methods

[0190] Materials

[0191] Uncladded polycrystalline silver halide fibers with a diameter of 500 pm, numerical aperture (NA) of 0.3, and are transparent between 3 and 17 pm wavelength. A hollow polyetheretherketone (PEEK) tube, with an inner diameter of 0.51 mm and an outer diameter of 1.59 mm, was used for fabricating connectors to hold two fibers.

[0192] An artificial cerebrospinal fluid (aCSF) was prepared to simulate the chemical and ionic composition of the interstitial fluid of the brain. The concentration of each component was as follows: NaCl at 149 mmol / L, KC1 at 2.8 mmol / L, CaCl₂·2H₂O at 1.2 mmol / L, MgSO₄·6H₂O at 1.2 mmol / L, ascorbic acid at 0.25 mmol / L, glucose at 5.4 mmol / L.200 proof ethanol (CH₃CH₂OH,111000200) was used for preparation of the standard solutions in aCSF through serial dilution.Docket Number: 10046-664W01

[0193] Rat brain dialysate was collected using a lab-made microdialysis probe from female Long-Evans rats. The microdialysis probe has a 3 mm working distance and a cellulose membrane with 13 kDa molecular weight cutoff. The probe was implanted above the striatal brain region and was continuously perfused with aCSF at a flow rate of 0.3 uL / min. To obtain sufficient sample volume for the FTIR instrument, brain dialysate collected from multiple rats were pooled together

[0194] Fiber processing

[0195] To fabricate angled fibers, custom fiber holders were designed to hold the fibers such that their end-faces were polished with the surface normal deviating from the fiber’s longitudinal axis by 0°, 10°, 15°, and 20°. These fiber holders were constructed from resin using a 3D printer. Next, 2 cm of the fiber end were sputtered with gold using a sputter coater for 120 seconds in order to minimize leakage of radiation through the lateral surface of the optical fiber when submerged in solution, due to increased refractive index in the surrounding medium. Before polishing the fiber, a small section of fiber was cut to remove the gold coated end-surface to allow radiation traveling in and out effectively. Polishing was performed using aluminum oxide lapping film with the fibers securely held in the fiber holder.

[0196] To fabricate the gold-coated fiber ends that function as a mirror for reflecting signal back to the optical fiber probe, a short segment of the silver halide fiber was cut, and one end was polished at 0° angle. The polished end-face was then coated with approximately 100 nm-thick gold layer using a sputter coater (6008, Ted Pella, Inc.).

[0197] Fiber probe sensor assembly

[0198] FIG. 15A shows the schematics of the optical fiber probe design in transflection mode. The fiber probe consists of an optical fiber with both ends polished at 10° angle, a short segment of gold-coated silver halide fiber which acts as mirror, and a connector designed to hold and align the fibers facing each other. The connector is made of a 1 cm long hollow PEEK tube with an inner diameter of 0.51 mm and an outer diameter of 1.59 mm. A 2 mm-wide groove was created in the connector using a hand drill, with a needle inserted into the tubing during the drilling process to prevent the hole from collapsing.

[0199] FIG. 15B shows the photograph of a fabricated fiber probe with the fiber polished at an angle of 10°. The fiber mirror was inserted with its 0° polished face oriented toward the angled optical fiber. The optical path length - the distance between the two fiber ends - was adjusted under a microscope after inserting the fiber and fiber mirror into the PEEK connector.Docket Number: 10046-664W01

[0200] To determine the optical pathlength for fiber probe with an angle, an averaged optical pathlength of the angled fiber probe was defined based on the wedge-shaped geometry formed between the polished fiber end and the mirror. The shortest distance between the fiber tip and the mirror was denoted as I, and the longest distance as L, both measured under a microscope. The optical pathlength varies linearly from a minimum value I to a maximum value L. The average one-way pathlength is therefore (L+ '2, and because the transflection configuration causes the beam to traverse the sample twice, the averaged optical pathlengthis:Lavg= 2 - ™ = L + Z (l)

[0201] In this study, I equals to 0 because the angled fiber was in slight contact with the fiber mirror. Therefore, the averaged optical pathlength is equal to the longest distance L.

[0202] Optical measurement setup

[0203] Fourier-transform infrared spectroscopy (FTIR). A Fourier-transform infrared (FTIR) spectrometer, equipped with a liquid cell operating in transmission mode, was used to demonstrate the ability to quantify ethanol in brain dialysate. The liquid transmission cell has an optical pathlength of 50 pm and calcium fluoride (CaF₂) windows, which corresponds to a minimum sample volume of 10 μL. The number of scans was 300, resulting in spectra collection time of 187 seconds, and the spectral resolution was 0.482 cm-1. The spectrum of aCSF was used as the background for absorbance calculations in the experiments.

[0204] Fiber sensor with quantum cascade laser (QCL)

[0205] FIG. 16 shows the schematics of the experimental setup 1600. A mid-IR quantum cascade laser was used as the radiation source, providing a spectral range of 800- 1800 cm1(5556-12500 nm). The plot of source intensity for different wavenumbers is shown next to the Q(’L block in FIGS. 17A and 17B. The output laser within this range varies between approximately 20 and 25 mW. The collimated laser beam was directed into a beam splitter using three mirrors. After the beam splitter, an off-axis parabolic mirror focused the beam onto the optical fiber probe. An optical fiber clamp was used to secure the optical fiber probe and was fixed on a 3-axis stage to adjust the position for signal optimization. The infrared radiation propagates through the optical fiber probe and interacts with the sample solution in the sensing area, which is between the angled fiber and the mirror in the probe sensor. The reflected signal is sent back through the fiber and directed to the detector. WhileDocket Number: 10046-664W01testing, the optical fiber probe was inserted into a 250 pL centrifuge tube, containing the analyte solution, until the sensing area was fully submerged in the liquid sample.

[0206] FIG. 16 further shows the schematic setup 1650 for testing back reflections using polished fibers with different end-face angles. In order to quantify back reflections from the optical fiber probe, a mirror was positioned directly behind the fiber, with the mirror surface perpendicular' to the fiber. A silicon wafer coated with approximately 150 nm of gold was used as the mirror. The gold coating was deposited through electron beam evaporation to achieve high reflectivity. Spectra were collected for both with and without the mirror to estimate the portion of the signal reflected by the mirror. The mirror-reflected signal represents the portion of the radiation that passes through the optical pathlength in the sensing area which interacts with the analyte and carries the spectroscopic information. By subtracting the mirror-reflected signal from the total reflection, the back reflection can be estimated.

[0207] Results and Discussion

[0208] Measurement of ethanol in aCSF using a FTIR

[0209] To evaluate the performance of measuring ethanol using the developed optical fiber probe, a standard FTIR was used to measure ethanol standards prepared in artificial cerebrospinal fluid (aCSF) to provide a basis for comparison. The use of aCSF simulates the ionic environment of biological fluids, demonstrating the capability of the method to measure ethanol in complex biological matrices. The spectrum of aCSF was used as the background for absorbance calculations.

[0210] FIG. 17A shows the mid-IR absorbance spectra of ethanol solutions at varying concentrations, over the spectral range from 1000 to 1300 cm−1. The C-0 stretching band at 1046 cm−1was selected for ethanol quantification due to its high absorbance, which enhances the sensitivity. In FIG. 17A the overall absorbance in the 1200-900 cm−1region decreases with increasing ethanol concentration. As ethanol is added to the solution, it replaces a portion of the water in the optical path. Since ethanol has lower absorbance than water across tills interval, the net absorbance drops. Peak height at 1046 cm−1was used for quantification of ethanol at different concentrations. The peak height was calculated by subtracting the baseline absorbance, averaged over a selected range (1010 to 1020 cm−1), from the absorbance at 1046 cm−1to account for the slight baseline shifts between measurements. A linear regression model was constructed to correlate ethanol concentration with peak absorbance intensity in FIG. 17B. The coefficient of determination (R2) for the full calibration dataset was 0.999.Docket Number: 10046-664W01

[0211] The LoD of the FTIR spectrometer was calculated as 4 mmol / L for ethanol in aCSF. This value was determined as three times the standard deviation of the peak height of the blank sample (just aCSF) divided by the sensitivity, which is defined as the slope of the calibration curve in FIG. 17B. This LoD demonstrates the ability to detect ethanol at physiological levels using mid-IR spectroscopy. Specifically, ethanol concentrations exceed 20 mmol / L (at a dose of 1 g / kg) following oral alcohol administration and 33 mmol / L (at a dose of 2.4 g / kg) following intraperitoneal alcohol injection, respectively. To further assess the predictive performance of the model, the root-mean squared error of cross-validation (RMSECV) for concentrations below 100 mmol / L was calculated to be 0.00087 absorbance units, corresponding to approximately 1 mmol / L based on the slope of the calibration curve.

[0212] FTIR measurement of ethanol in rat brain dialysate

[0213] In order to determine the feasibility of using a mid-IR spectroscopy for quantifying ethanol in the brain tissue for in vivo applications, a FTIR spectrometer was used to measure ethanol in rat brain dialysate. This evaluation is important for determining the sensitivity and selectivity of the technique in detecting ethanol in a complex biological fluid. Rat brain dialysate represents the in vivo environment, allowing for the identification of any interferences or overlapping absorption bands that could impact sensor performance.

[0214] FIG. 18A shows the FTIR spectrum of rat brain dialysate, using aCSF as the background. The FTIR spectrum of rat brain dialysate reveals an absorbance band near 1360 cm−1, which is primarily attributed to bicarbonate ions, a common component of biological fluids such as cerebrospinal fluid. This assignment is based on a previous study involving spectral analysis of interstitial fluid. Importantly, there is no overlapping absorption bands detected near 1046 cm−1, which is the region used for ethanol quantification. This confirms the desirable mid-IR spectral range for ethanol quantification.

[0215] FIG. 18B shows the absorbance peak height at different ethanol concentration in aCSF, calculated using the same procedure described herein. The coefficient of determination (R2) for the calibration dataset, after excluding the first two points, was 0.994. The RMSESV was 0.00073 absorbance units, corresponding to a concentration of 1 mmol / L. At concentrations below the limit of detection, the ethanol absorbance peak is masked by the baseline noise. This effect was also observed for ethanol in aCSF, however, due to the broader ranges of both axes in FIG. 18B, it is visually not observable.

[0216] The sensitivity, determined as the slope of the fitted calibration line in FIG.18B. For measuring ethanol in brain dialysate, it is comparable to that of ethanol in aCSF. The calculated LoD is 4 mmol / L for ethanol in rat brain dialysate.Docket Number: 10046-664W01

[0217] Fiber sensor measurement of ethanol solutions at different concentrations

[0218] The measurement of ethanol solutions in aCSF at varying concentrations was used to assess the optical fiber probe's performance. A 10° angled fiber probe with an averaged optical pathlength of 68 pm was used. FIG. 19A presents the absorbance spectra of ethanol solutions at different concentrations from 4 to 3428 mmol / L, using aCSF as the background. Each spectrum is an average of three measurements, which results in spectra collection time of 89 seconds. The peaks at 1046 cm−1and 1088 cm−1correspond to C-0 stretching vibrations, with the 1046 cm−1peak utilized for quantification. The peak at 1046 cm-1was selected for its higher absorbance and for direct comparison with the FTIR results. The negative baseline shift is attributed to the lower concentration of water and other absorbing components in the ethanol solutions compared to the aCSF background. The periodic modulation observed in the spectral baseline is attributed to Fabry-Perot interference resulting from internal reflections within the sensing region.

[0219] To confirm the source of the modulation, a Fourier transform was performed on the spectra in FIG. 19 A to estimate the corresponding optical pathlength responsible for the modulation, as shown in FIG. 19B. To avoid interference from ethanol absorbance bands, the spectral range of 1100 to 1300 cm-1was used. A peak at frequency of 0.02 cm was consistently observed across all Fourier-transformed spectra of ethanol at different concentrations, with no measurable shift in this peak was detected as ethanol concentration changes. The expected frequency shift resulting from the refractive index change between water (n=1.23) and pure ethanol (n=1.35) is approximately 0.002 cm at a frequency of 0.02 cm. Because the spectral range of 1000–1300 cm−1provides a frequency resolution of 0.005 cm, which exceeds the expected shift, the frequency shift could not be observed.

[0220] The Fourier transform of the spectrum at 4 mmol / L was then used to calculate the optical pathlength with the refractive index of water (n=1.23). At this low concentration, the refractive index is expected to be close to that of water. The calculated pathlength corresponding to the 0.02 cm frequency was 81 pm. Considering the frequency resolution of 0.005 cm, equivalent to 20 pm in optical pathlength, this value is consistent with the 68 pm average optical pathlength measured by microscope. As a result, the optical pathlength calculated based on the modulation supports the interpretation that the observed modulation results from interference within the wedge-shaped sensing region.

[0221] FIG. 19C shows the calculated peak heights for different ethanol concentrations. Peak heights at 1046 cm−1were determined using the same method as in the previous section for concentration quantification. Compared to the result obtained with FTIR,Docket Number: 10046-664W01the trend of peak height versus concentration is nonlinear, which is attributed to the presence of optical back reflections. 'The estimation and the removal of back reflection effects are discussed in the following section.

[0222] Evaluating and correcting for the back reflections

[0223] Superposition of back reflections and mirror-reflected signal can cause undesired absorption peaks and reduce sensor’s SNR. Back reflections originate from the airfiber interface where the radiation is first introduced into the fiber and the fiber-air interface before the radiation exits the fiber.

[0224] To determine the total amount of back reflections from air-fiber and fiber-air interface in the fabricated probe, spectra were measured from a fiber both with and without a mirror positioned directly behind it. The difference between the two signals corresponds to the mirror reflections, representing the portion of radiation that interacts with the analyte.

[0225] To reduce back reflections, the fiber's end-face was polished at an angle, directing the reflected radiation out of the fiber. Various end-face angles were created through polishing to investigate this. The radiation input end of all fibers was polished at 0°, while the output end, which receives the mirror-reflected radiation, was polished at 0°, 10°, 15°, and 20°. FIG. 20 A shows the percentage of mirror-reflected signal relative to the total reflected signal (comprising both back reflection and mirror reflection). As the end-face angle increased, the proportion of mirror reflection rose, reaching a maximum at 10°, before decreasing with larger angles. FIG. 20B shows the percentage of mirror-reflected signal at 1046 cm−1(corresponding to the C-O stretching used for ethanol quantification in this study) at different angles.

[0226] The back reflection was reduced by 69% with a 10° polishing angle compared to a 0.9° polished fiber, and by 62% compared to a fiber polished at 14.3°. At a polishing angle of 19.7°, back reflection became dominant, accounting for over 99% of the total reflected signal. Based on these results, the 10° fiber holder was selected to polish the optical fiber to minimize back reflection.

[0227] In addition to experimental evaluation, finite-element simulations were performed to assess back reflections for different angles. Both the simulated and experimental results showed a similar trend in back reflections — initially decreasing and then increasing as the polishing angle increased. The discrepancy in the optimal angle, where back reflection is minimized, between the experimental data and the simulation may be attributed to the fiber diameter and number of modes considered in the simulation.Docket Number: 10046-664W01

[0228] The back reflection signal cannot be completely eliminated, which reduces the signal-to-noise ratio (SNR) in absorbance calculations and introduces non-linearity in calibration results. Therefore, it is crucial to estimate and subtract the back reflection signal before data processing. The absorbance, considering back reflection, is given by Equation (2):Sw(λ) + Rw(λ)Abs(λ, c) = log10[Sw(λ) + Rw(λ)] / [Se(λ, c) + Re(λ, c)] (2)Se(λ, c) + Re(λ, c)

[0229] where, Swdenotes the mirror-reflected signal for water, while Rwrepresents the total back reflection, which includes reflections from both the air-fiber interface and the fiber-sample interface when measuring water. Similarly, Seis the measured mirror-reflected signal for ethanol, and Rerepresents the total back reflection under ethanol solution measurements. All parameters are functions of wavenumber or wavelength (A), whereas Seand Realso vary with ethanol concentration c. At a specific wavenumber, Seand Reare only a function of concentration.

[0230] Experimental observations indicate that the change in back reflection signal is less than 2% within the 1000–1200 cm−1range when measuring water and ethanol at different concentrations using fibers polished at 0° and 10°. Additionally, the difference in refractive index between ethanol and water at room temperature is less than 3%, leading to comparable back reflection conditions. Therefore, back reflection can be assumed to be constant for both water and ethanol solutions. Consequently, Rw, Reand R can be considered constant and equal in the 1000–1200 cm−1range, leading to:Rw= Re= R (3)

[0231] According to Beer-Lambert’s law and the FTIR measurements of ethanol solutions at varying concentrations, the absorbance at 1046 cm−1has a linear relationship with ethanol concentration c (R2>0.99) for a constant optical pathlength. As a result, the ethanol absorbance signal can be expressed as:Abs(c) = log10= Ct■ c (4)

[0232] and therefore,Se(c) = Sw× 10−C₁·c(5)

[0233] where Ci is a constant. By incorporating the experimental data and substituting Equation (5) into Equation (2), the equation can be rearranged as:Docket Number: 10046-664W01Abs(c) = log10^>< 10C"Ci.e + R(6)

[0234] where, Co represents the total reflected signal for measuring water at 1046 cm−1, which includes both the water signal and back reflection ( Sw+ R). Swis the unknown water signal, and R is the unknown back reflection signal. To determine these unknowns (Sw, Ci, and R), a three-parameter regression model was applied to fit the experimental data in FIG. 21A which corresponds to the same dataset as FIG. 19C. Using this approach, the back reflection at 1046 cm”1can be estimated and subtracted from the original signal.

[0235] FIG. 21 B shows the absorbance at different concentrations after back reflection removal. Now, the absorbance is linearly proportional to concentration and consistent with FTIR measurements. The sensitivity, represented by the slope of the fitted line, is within 4% of the fitted constant C₁ which predicts the value of the slope, demonstrating that the experimental results are in good agreement with the regression model.

[0236] Additionally, to compare the sensitivity before and after subtracting back reflection, sensitivity was determined by differentiating equations (6) and (4) with respect to concentration, resulting in equations (7) and (8), respectively, as follows:Swx 10-c*cSenh„fnrp— - T - Ci (7)before(SwX 10-G-c + R-)1 v JSen-after. (8)

[0237] The factor / , representing the ratio of sensitivity improvement after subtracting back reflections, is calculated by dividing equation (8) by equation (7), as expressed below:f = Senafter / Senbefore= 1 + R / (Sw× 10−C₁·c) (9)SenbeforeSwX 10 -G-c

[0238] According to equation (9), the approach for calculating and removing back reflections is even more beneficial in cases of strong back reflection, where the mirror- reflected signal is relatively weak compared to the back reflection signal. This can happen when the refractive index of the fiber is much higher than that of the sample in the optical pathlength. While the intensity of mirror-reflected radiation is weak if the distance to the mirror is long, and if the solution that is present in the optical pathlength, is a strong mid-IR absorber such as water. These two factors limit the possible optical pathlength.

[0239] In addition to optical back reflections, the geometry of the sensing region also contributes to the non-linear absorbance response observed before correction. Specifically, the angled fiber end creates a wedge-shaped sensing gap, where the optical pathlength variesDocket Number: 10046-664W01spatially across the fiber’s cross-section. As described by Hirschfeld, this variation leads to photometric non-linearity. To estimate the non-linearity effect caused by the wedge-shaped sensing area, equation (10) was used, as derived by Hirschfeld:(In 10)A,= log- — (10)1 — lu

[0240] where Awis the absorbance corresponding to the averaged optical pathlength w, and Ai is the absorbance corresponding to the maximum optical pathlength I in the wedge- shaped sensing gap. In this study, w is 68 gm and I is 136 pm.

[0241] For the analysis of non-linearity caused by the wedge-shaped sensing gap, one high, one medium, and one low concentration were selected from the dataset. The FTIR data collected at a 50 pm optical pathlength was extrapolated to optical pathlengths of both 68 pm and 136 pm according to Beer-Lambert’s law. The estimated absorbance at 68 pm was used as a baseline for comparison. The estimated absorbance at 136 pm was used to calculate the theoretical absorbance for a wedge-shaped transmission cell with an averaged optical pathlength of 68 pm. Table IB shows the extrapolated FTIR absorbance at a 68 pm optical pathlength as baseline, the calculated absorbance for a wedge-shaped transmission cell with an average optical pathlength of 68 pm, and the experimentally measured absorbance using the fiber probe.

[0242] Table IB. Absorbance measured using normal cell with an optical pathlength of 68 pm, wedge-shaped cell and fiber probe both with an averaged optical pathlength of 68 pm.AbsorbanceConcentration FTIR FTIR with wedge-shaped cell Fiber probe (mmoI / L) (68 (averaged optical pathlength of (averaged optical pathlength of pm) 68 pm) 68 pm)17 0.017 0.017 0.01386 0.083 0.080 0.037429 0.420 0.355 0.099

[0243] The results show no difference between using a wedge-shaped cell and a normal cell at the lowest concentration. For medium and high concentrations, the difference is approximately 3% and 15% respectively. However, the probe measurement has a difference of 24%, 55% and 79% differences with the FTIR results for low, medium and high concentrations respectively. The results in Table IB validate that the wedge-shaped sensing region is not the dominant factor causing the non-linearity and suggest that the back reflection reduces absorbance (or sensitivity) even at low concentrations. Our method forDocket Number: 10046-664W01estimating and subtracting the back reflection signal can help mitigate the absorption non¬ linearity and improve sensitivity.

[0244] Performance of the optical fiber probe sensor and the FTIR

[0245] Before removing the back reflection, the LoD for the transflection probe is 7 mmol / L. After removing the back reflections, the LoD for the same probe and dataset is 4 mmol / L. Demonstrating the improvement of back reflections removal. Additionally, the LoD of fiber probe with an optical pathlength of 68 gm slightly outperforms the FTIR instrument measured in transmission mode with a 50 gm optical pathlength. Furthermore, the fiber probe reduces the sample spectrum acquisition time by 52% compared to a traditional FTIR while maintaining a comparable LoD. This shows that implementations of the present disclosure enable a reliable and sensitive tool for real-time ethanol monitoring.

[0246] The sensitivity of both the fiber probe and the FTIR spectrometer depends on the optical pathlength, with sensitivity being directly proportional to the optical pathlength as described by Beer-Lambert’s law. For an optical pathlength of 68 gm, the estimated sensitivity for FTIR measurement is 0.00096, which is higher than the sensitivity of a 10° angled fiber probe, measured at 0.00041. The reduced sensitivity of the fiber probe can be attributed to factors such as coupling losses, back reflection, and transmission losses in the silver halide optical fiber. Therefore, the optimal optical pathlength for a FTIR spectrometer does not necessarily correspond to the optimal optical pathlength for the fiber probe.

[0247] Effects of optical pathlength and fiber’s angle on probe performance

[0248] To evaluate the performance of the optical fiber probe for different optical pathlengths, transflection probe with optical pathlengths of 64, 92, 157, 227, 317, and 387 gm (equal to twice the distance between the fiber end and the fiber mirror) were used to measure ethanol solutions of different concentrations. To provide proper control of the optical pathlength for the sample, 0° angled fibers were used to fabricate the probes. Ethanol was prepared at concentrations of 200, 150, 100, 50, and 25 mmol / L.. FIG. 22A shows the peak height (at 1046 cm’1) versus concentration for fiber probes with varying optical pathlengths. The analysis followed the same procedures as described herein. Results for optical pathlengths of 227, 317, and 387 μm are not included in FIG. 22 A because the sample absorption was too strong and the signal was completely absorbed, showing no change in absorbance for different concentrations. Probes with longer optical pathlengths demonstrate higher sensitivity, as indicated by the slopes of the calibration lines in FIG. 22A. The probe with an optical pathlength of 157 μm is unable to detect 25 mmol / L, as the signal is near the noise level.Docket Number: 10046-664W01

[0249] FIG. 22B shows the noise levels, defined as the standard deviation of the absorbance values in the range of 1020 to 1070 cm-1(corresponding to the C-0 absorption peak used for quantification), for fiber probes with different optical pathlengths. For pathlengths below approximately 160 pm, the noise level generally increases with optical pathlength due to the greater attenuation of the signal by the absorbing medium. However, for probes with longer optical pathlengths (> 157 pm), the sample absorption becomes so strong that nearly all of the mirror-reflected signal is absorbed, and the detected signal is noise and thus not suitable for quantification. As a result, FIG. 22B only includes data points corresponding to optical pathlengths of 64, 92 and 157 gm.

[0250] FIG. 22C shows the SNR for different optical pathlengths, calculated using the slopes from FIG. 22A and divided by the noise level. Among the tested optical pathlengths, the fiber probe with an optical pathlength of 64 gm has the highest SNR, indicating the best performance.

[0251] To investigate the effect of using angled fiber, the 0° angled fiber probe with a 64 gm optical pathlength was used to compare with a 10° angled fiber probe with a similar averaged optical pathlength (68 gm). As shown in FIG. 22D, the angled fiber probe showed a 32.7% higher SNR. While both probes have similar sensitivity, the 10° angled fiber probe had a lower noise level, which is due to reduced back reflection, resulting in enhanced performance.

[0252] Discussion

[0253] An optical fiber probe sensor designed in transflection mode was fabricated using mid-IR silver halide fiber with a diameter of 500 gm, a gold-coated silver halide fiber acting as a mirror, and a PEEK connector for holding and aligning the two fibers. The probe features an outer diameter of 1.59 mm, which is smaller than conventional mid-IR transflection probes. Desired optical pathlength was achieved by aligning the two fibers under a microscope during the fabrication process. The averaged optical pathlength for the angled fiber probes was determined by averaging the shortest and longest distance between the fiber tip and the fiber mirror.

[0254] The feasibility of quantifying ethanol in biological fluid (rat brain dialysate) using mid-IR spectroscopy was first confirmed using a standard FTIR spectrometer. The optimal polishing angle of 10° was determined by investigating the back reflections at 0°, 10°, 15°, and 20° angles. The back reflections decreased with an increasing end-face angle, reaching a minimum at 10°, before increasing again at higher angles. To demonstrate the probe's performance and its capability to measure concentrations at physiological levels,Docket Number: 10046-664W01ethanol solutions with concentrations ranging from 4 to 1714 mmol / L were tested. To remove the back reflections, calculations based on experimental data were developed to estimate the back reflections signal, which was then subtracted from the signal obtained by the detector. The optical fiber probe achieved a limit of detection of 4 mmol / L for ethanol in aqueous solution, demonstrating its applicability for in vivo measurements for alcohol research. To evaluate the effect of the optical pathlength on probe’s performance, probes with different optical pathlengths were tested, with the 64 pm optical pathlength achieving the highest SNR among all tested optical pathlengths. Additionally, the fiber probe with a 10° angled fiber achieved a 32.7% higher SNR compared to the probe with a 0° fiber of similar optical pathlength, due to its reduced noise level because of the use of angled fiber.

[0255] This study demonstrates the potential of the developed transflection fiber probe for in vivo applications. 'The example method for removing back reflections improved the SNR of the probe and the LoD. The improvement due to the removal of the back reflections is expected to be even greater in scenarios where the back reflections are stronger, i.e. when the refractive index of the fiber is high, and when there are weak mirror-reflected signals - a common case when analyzing aqueous samples due to the strong mid-IR absorption of water. The developed transflection probe for mid-IR spectroscopy offers high sensitivity and selectivity, enabling the developed probe to be utilized in a wide range of molecular monitoring applications. The example implementation described herein can further be miniaturized by utilizing fibers with smaller diameters and creating more compact connectors for accommodating both the fiber and the mirror.

[0256] Example 4:

[0257] Simultaneous, real-time monitoring of multiple biomarkers in tissues is critical for diagnostics, therapy, and metabolic health management. However, few existing platforms have the capability of multi-compound detection in vivo. This study presents a compact transflection optical fiber probe for mid-infrared spectroscopic analysis, enabling label-free, simultaneous monitoring of ethanol, glucose, and lactate. The probe comprises two silver halide fibers — one with an angled sensing tip and one gold-coated acting as a mirror — housed in PEEK tubing and surrounded by a semi-permeable membrane. With an outer diameter of only 1.1 mm and a sample volume of 13 nL in the sensing region, smaller than conventional mid-infrared transflection probes. Using a tunable quantum cascade laser system as the light source, the fiber probe’s performance for quantification of ethanol, glucose, and lactate concentrations is comparable with a benchtop infrared spectroscopy system, achievingDocket Number: 10046-664W01detection limits of 1.0, 0.8, and 1.1 mM, respectively. Peak deconvolution was deployed to resolve overlapping spectral features, enabling quantification of individual compounds in mixtures. Validation was performed in ex vivo human skin against microdialysis - a standard technique for in vivo sampling. Additionally, changes in the concentration of all three compounds in the skin were monitored for 90 minutes with measurements every 90 seconds.

[0258] Real-time monitoring or detection of small-molecule metabolites such as ethanol, glucose, and lactate is crucial for advancing precision medicine, early disease detection, managing metabolic disorders, guiding therapeutic decisions, and improving patient outcomes in critical care settings. These molecules serve as important biomarkers across diverse physiological and pathological conditions; Glucose and lactate levels are used to assess brain metabolic dysfunction following traumatic brain injury (TBI), where elevated lactate and reduced glucose are associated with poor outcomes. Continuous glucose monitoring is fundamental in diabetes management, while elevated lactate serves as a critical prognostic marker in sepsis. Ethanol monitoring is important in contexts such as alcohol intoxication and addiction treatment. Monitoring changes in concentration over time can help with better understanding of the metabolic pathway, which is important for developing new treatments and therapeutics. Ethanol metabolism also influences glucose and lactate dynamics due to hepatic metabolic competition and mitochondrial effects. In scenarios such as ethanol-induced liver injury, simultaneous monitoring of ethanol, glucose, and lactate provides a more complete picture of the metabolic status. Table 1C summarizes clinical relevance and applications of glucose, lactate and ethanol that underscore the importance of developing a sensing platform capable of measuring them simultaneously.

[0259] Table 1C. Physiological levels, clinical relevance, importance of simultaneous monitoring and applications for ethanol, glucose and lactate.TypicalClinical andphysiological Importance of multiCompound physiological Applications concentration analyte monitoringsignificancein tissueRelevant in intoxication, Ethanol alters glucose Addiction Negligible addiction, brain injury; and lactate levels due to treatment, under normal influences glucose and hepatic metabolic forensic conditions; lactate metabolism; competition and toxicology, Ethanolcan reach 100 significant for mitochondrial effects; ICU mM during understanding neural combined monitoring monitoring, intoxication. effects and systemic provides insights into alcohol- > toxicity.gglgtgd brainDocket Number: 10046-664W01during alcohol exposure injury or injury. research. Essential source of Glucose metabolism isDiabetes energy; critical for brain linked to lactatemonitoring, ~3 -9 mM in metabolism; primary production viametabolic blood and in target in diabetes glycolysis; ethanolGlucose studies, interstitial management; changes metabolism competestraumatic fluid. indicate metabolic with glucose utilizationbrain injury dysfunction after and affects bloodtraumatic assessment._ -0.3-1.3 mM Sepsis Indicator of anaerobicunder normal Ethanol metabolism monitoring, metabolism; elevatedconditions; alters NAD⁺ / NADH trauma levels signal tissueelevated (>5 balance, impacting evaluation, Lactate hypoxia, sepsis, ormM) in lactate levels; lactate sports metabolic stress; serveshypoxia, rises when glucose medicine, as a prognostic markersepsis, metabolism is impaired. metabolic in critical care.trauma. research.

[0260] Conventional approaches, such as microdialysis, are widely used for in vivo and ex vivo monitoring of biochemical markers. Microdialysis allows localized sampling of interstitial fluid for subsequent offline analysis, and thus inherently does not provide real¬ time continuous measurement and suffers from low temporal resolution due to slow perfusion rates and long sampling intervals. Additionally, accurate quantification of concentration in tissues requires tedious calibration procedures. Furthermore, this extraction-based technique can deplete molecules locally in the tissue surrounding the microdialysis probe, resulting in changes of the chemical environment and leading to lower concentrations in the collected sample. Electrochemical sensing, particularly through enzyme-based detection, is the primary approach for monitoring metabolites such as glucose and lactate, providing rapid, real-time measurements and integration into wearable and implantable devices. Although effective, electrochemical sensors are limited by fouling and the need for immobilized enzymes with finite lifespans and degradation of the electrode. Although recent studies, such as an implantable near-infrared spectroscopy sensor, have demonstrated simultaneous monitoring of ethanol, glucose and lactate in biological tissue, these systems involve surgical implantation which increases the overall cost and the complexity of the procedure. Moreover, the relatively large device size, requiring an incision of approximately 3.5 cm for implantation, restricts its use to specific tissue sites and limits broader medical applicability.Docket Number: 10046-664W01

[0261] Mid-infrared (IR) spectroscopy enables molecular detection in tissues by directly probing the fundamental vibrational fingerprints of molecular bonds. This approach allows chemical identification without the need for additional targeting markers and provides distinct spectral features that are not captured in the overtone-dominated near-IR region. Each compound shows unique absorbance peaks in the mid-IR region (wavelengths 2.5–20 μm, corresponding to wavenumbers 4000–500 cm⁻¹, which result from characteristic bond vibrations. This enables direct molecular identification and quantification. Several mid-IR sensing modalities have been explored for chemical sensing purpose; Attenuated total reflectance (ATR) spectroscopy, which measures evanescent-wave absorption from a sample in contact with a high-refractive index crystal, has been used for analyzing glucose, and lactate. Due to the limited light-matter interaction, the sensitivity is generally lower than that of other approaches. Evanescent-wave fiber sensors utilize uncladded, bent or tapered mid-IR fibers - typically made from silver halide or chalcogenide materials - where the evanescent field extends beyond the fiber surface to interact with surrounding medium. Bending or tapering of the fiber increases the extent of the evanescent field interaction with the sample, thus enhancing sensitivity. In a previous study, glucose was measured using a triple-coiled silver halide fiber sensor with a bending radius of 2.5 mm. Other studies have demonstrated the use of evanescent- wave sensors for detecting volatile organic compounds in aqueous environments, quantification of keratin and for distinguishing between healthy and cancerous skin tissue. Fiber-optic sensors operating in transmission mode, where light passes directly through the sample between two aligned fibers, have been used to measure various concentration of glucose in different synthetic mixtures. However, the implementation of evanescent-wave and transmission fiber sensors in tissues remains challenging as their size and configuration will cause significant tissue damage (for example, transmission measurements require alignment of two fibers within the tissue). To minimize the invasive impact, alternative design for fiber probes, such as a transflection configuration, is preferable.

[0262] In this study, a transflection fiber probe operating in the mid-IR region with an average pathlength of 63 pm was developed. The mid-IR transflection fiber probe, which uses a quantum cascade laser (QCL) as its light source, has an overall diameter of 1.1 mm. The sensing area of the probe is surrounded with a semi-permeable membrane having pores of 2-5 nm to prevent direct contact of the silver halide fiber with the tissue to enhance biocompatibility. Additionally, the membrane excludes large particles such as cells and proteins from the sensing region, thus preventing biofouling and reducing potential interference in the IR measurement. The probe was first tested by measuring ethanol, glucoseDocket Number: 10046-664W01and lactate in aqueous solutions. The temporal response of the fiber probe was investigated in a glucose solution, comparing probes with and without a membrane. The performance of the probe was validated in a tissue by measuring ethanol concentration over time in ex vivo human skin against a standard technique for sampling small molecules in tissues - microdialysis. Additionally, ethanol, glucose and lactate concentrations were measured in the skin sample simultaneously to demonstrate the multi-compound sensing capability. Peak deconvolution of the measured spectra was performed to resolve overlapping spectral features to quantify the concentrations of each compound. A rapid injection-dilution experiment was performed to show that the probe is capable of capturing concentration changes in tissue with a temporal resolution of 30 seconds. To estimate chemical stability and biocompatibility of the fiber probe, the dissolution of silver halide fiber was measured in ultrapure water over 7 days. The results showed that the maximum silver ion concentration remained below 26 ppb. This design demonstrates a versatile platform for multi-compound sensing in a biologically relevant environment.

[0263] Materials and Methods

[0264] Materials

[0265] Optical fiber. Uncladded polycrystalline silver halide fibers (PIR 500, Art Photonics, Germany) with a diameter of 500 gm, numerical aperture (N. A.) of 0.3, and transparency in the wavelength range of 3 to 17 gm, were used for fabricating fiber probes.

[0266] Reagents

[0267] An artificial cerebrospinal fluid (aCSF) was prepared to simulate the chemical and ionic composition of the interstitial fluid, according to a previous study. The concentration of each component was as follows: NaCl at 149 mM, KCl at 2.8 mM, CaCl₂·2H₂O at 1.2 mM MgSO₄·6H₂O at 1.2 mM (63068, Sigma), ascorbic acid at 0.25 mM, glucose at 5.4 mM. 200 proof ethanol, glucose and lactate were used to prepare standard solutions in aCSF through serial dilution.

[0268] Skin tissue models

[0269] Abdominal skin collected from a 31 -year-old Hispanic female donor after plastic surgery was split into multiple samples used for testing probe performance ex vivo. The skin samples have a diameter of 20 mm and a thickness of approximately 4.6 mm. Each skin biopsy is contained in a cell culture insert designed to fit standard six-well plates. The insert has a porous membrane at its base, allowing nutrient exchange between the culture medium and the skin tissue. A nourishing gel-like matrix of proprietary composition is positioned between the membrane and the skin, providing physical support and keeping theDocket Number: 10046-664W01tissue chemically stable. A silicone ring circles the exposed surface of the skin, preventing the applied liquids from leaking into the culture medium. As a result, the working area of the skin surface is reduced to 1.76 cm2. The skin sample was received one day after production. Upon delivery, the skin samples were refreshed with new medium and kept in culture at least 3 hours prior to experiment. Experiments were performed in accordance with the manufacturer’s instructions and before the expiration date of the tissue. This study is not classified as human subject research, and no Institutional Review Board approval was required.

[0270] Instruments

[0271] Fourier-transform infrared spectroscopy ( FTIR )

[0272] A Fourier-transform infrared (FTIR) spectrometer, equipped with a liquid cell operating in transmission mode, was used to measure solutions of ethanol, glucose and lactate as a reference method. The liquid cell has an optical pathlength of 50 pm and CaF₂ windows, requiring a sample volume of 10 pL. The number of scans was 300, resulting in spectra collection time of 187 seconds, and the spectral resolution was 0.482 cm The spectrum of aCSF was used as the background for absorbance calculations in the experiments.

[0273] Headspace gas chromatography with flame ionization detector (GC-FID)

[0274] A gas chromatograph equipped with a flame ionization detector (FID) and an autosampler was used to analyze the concentration of ethanol in the dialysate samples.Software was used to record and analyze all chromatograms. A capillary column (30 m x 0.53 mm x 1 pm film thickness) with nitrogen as the mobile phase, at a flow rate of 8.5 mL / min, was used. A carboxen / PDMS coated solid phase micro-extraction (SPME) fiber was used for extraction and desorption. The extraction time was set to 3 minutes, followed by a desorption period of 1 minute. The oven and injection temperature were maintained at 65 °C and 220 °C, respectively.

[0275] Fabrication of microdialysis probes

[0276] Microdialysis probes were assembled in a side-by-side configuration, following the procedure described in a prior study. Briefly, two pieces of fused silica tubing used as the inlet and outlet lines and were inserted into a segment of regenerated cellulose membrane with a 13 kDa molecular weight cutoff (MWCO), which corresponds to an effective pore size of 2-5 nm. The active dialysis region was determined by adjusting the relative insertion lengths of the tubing inside the membrane and was set to 2 mm. Epoxy was applied externally to coat the region that is not used for dialysis.

[0277] Fabrication of optical fiber probesDocket Number: 10046-664W01

[0278] FIG. 23A illustrates the schematic design of the mid-infrared transflection fiber probe. The probe consists of two aligned silver halide optical fibers. One fiber is used for light delivery and collection, and the other serves as a reflective mirror. The fiber mirror was fabricated by coating the end-face of a short segment of silver halide fiber with a gold layer of approximately 267 nm thickness, using a sputter coater. Before gold sputtering, the fiber end was polished using aluminum oxide lapping film. The other fiber, used for delivery and collection of the IR light, was polished at a 7° angle on its end face facing the fiber mirror using a custom-made fiber holder. This angled fiber tip helps reduce Fresnel back reflections by redirecting the reflected light away from the fiber core.

[0279] Both fibers were held and aligned face-to-face inside a custom-fabricated connector made from a 6 mm-long hollow polyetheretherketone (PEEK) tube (1532L, IDEX Health and Science) with an inner diameter of 510 pm and an outer diameter of 1588 pm. The outer surface of the PEEK tube was machined down to an outer diameter of approximately 850 pm to fit within a semi -permeable membrane (modified polyethersulfone (mPES), 10 kDa MWCO corresponding to effective pore size of around 3 nm). The alignment of the two fibers was adjusted under a microscope to control the optical pathlength - twice the distance between the two fibers. The average optical pathlength of the fiber probe was 63 pm, which is twice the average distance between the fiber tip and the fiber mirror. This average distance was determined by measuring the shortest and longest distances between the angled fiber tip and the fiber mirror under a microscope and calculating their mean. The space between the fiber tip and the fiber mirror defines the sensing region, which requires a sample volume of 13 nL to fill completely. To enable molecules to reach the sensing region, a 2 mm-wide groove was created in the PEEK tube along the wall using a hand drill. The fiber probe was inserted into a 3 mm-long piece of the semi -permeable membrane enclosing the tube and covering the sensing window and secured using epoxy at both ends. This configuration allows small molecules to diffuse through the membrane into the sensing region while excluding large particles such as proteins. The overall outer diameter of the final probe is 1.1 mm. An illustration of the final assembled probe is shown in FIG. 23B.

[0280] Optical setup

[0281] FIG. 24 shows the schematics of the experimental setup. A mid-infrared quantum cascade laser was used as the light source, delivering a spectral range of 800-1800 cur1(5.56-12.5 pm). A focusing coupler (Art Photonics GmbH), consisting of a gold mirror (39-193, Edmund Optics) and an off-axis parabolic mirror (35-488, Edmund Optics), wasDocket Number: 10046-664W01installed at the laser output, as shown in the schematics. The mirror mounted inside the focusing coupler features three adjustment screws for fine-tuning the x, y, and z positions, enabling signal optimization at the FC -PC connector output.

[0282] The output of the focusing coupler was connected to one input of a fiber splitter, which combined two silver halide fibers (500 pm diameter) for delivering the laser light and transmitting the return signal to the detector, as illustrated in the schematics. The two fibers of the splitter were bundled together into a single fiber cable, which was connected to the transflection probe through an SMA connector. To connect the transflection probe to the fiber splitter, a 3D-printed hollow cylinder was fabricated with an inner diameter matching that of the fiber probe for secure insertion, and an outer diameter designed to tightly fit the SMA connector port. The signal carrying spectral information was directed to a thermoelectrically cooled infrared detector using a dual-mirror fiber coupler. This coupler includes two off-axis parabolic mirrors that collect and focus the light onto the detector, as illustrated in the schematic.

[0283] Ex vivo skin measurements

[0284] To assess the performance of the mid-IR transflection fiber probe, ex vivo experiments were conducted using human abdominal skin samples. For direct comparison of the fiber probe and microdialysis, both the optical fiber probe and a lab-made microdialysis probe were inserted into the dermis of the same skin sample, positioned approximately 2.7 mm and 2.9 mm below the top skin surface, respectively, as shown in FIG. 25A. Insertion of both probes was performed by first making a small incision in the dermis using a scalpel, followed by piercing the tissue with a 27-gauge needle to create a tract for the probes. The fiber and microdialysis probes were then carefully inserted along this path.

[0285] The fiber probe was used to acquire mid-IR spectra in situ, while the microdialysis probe served as a reference method for sampling extracellular molecules. The culture insert, which contains skin tissue, placed in a six-well plate containing medium, was kept inside a humidified incubator at 37°C for the duration of the experiments. The microdialysis system included a syringe pump delivering artificial cerebrospinal fluid (aCSF) at a constant perfusion rate of 0.5 gL / min. Dialysate was collected in vials at 10-minute intervals. A 2 gL aliquot from each dialysate sample was transferred to a glass vial for subsequent analysis using GC-FID.

[0286] To directly compare the measurement results of the fiber probe and the microdialysis probe, both were inserted in the skin sample, and ethanol concentration measurements were initiated simultaneously upon spiking the culture medium with ethanol toDocket Number: 10046-664W01a concentration of 20 mM. During this experiment, spectra were acquired every 1.5 minutes, corresponding to an average of 3 scans. The ethanol measurement results obtained with both techniques are presented and compared herein.

[0287] For simultaneous multi-compound quantification, only the optical fiber was inserted in the skin sample. The medium was spiked with a mixture of 100 mM ethanol, 100 mM glucose, and 100 mM lactate. These high concentrations in the medium were used in order to achieve physiologically relevant concentrations in the dermis of the skin. Spectra were acquired every 1.5 minutes, and the acquired spectra were analyzed using peak deconvolution to resolve and quantify the overlapping absorbance features. The results of simultaneous multi-compound quantification are presented herein.

[0288] To evaluate the fiber probe’s capability for capturing dynamic changes in concentration, an injection experiment was performed with the fiber probe in the tissue, as shown in FIG. 25B. A volume of 50 pl, of 200 mM ethanol was injected into the skin at a site approximately 3 mm from the probe, followed by a dilution step in which 20 pL of culture medium was added directly at the insertion site of the probe to quickly reduce the local ethanol concentration. During this experiment, spectra were acquired every 30 seconds using single-scan acquisition to minimize measurement time, The results of tracking ethanol concentration over time are presented herein.

[0289] Spectral processing and peak deconvolution

[0290] All infrared absorbance spectra were processed and analyzed using MATLAB. The spectra were first smoothed using a moving average method with a window size of 9 points. This window' size was chosen to preserve peak shape and intensity while reducing high-frequency noise. Following smoothing, baseline correction was performed. An asymmetric least squares algorithm was used to estimate and subtract the baseline. This method iteratively minimizes a cost function that balances how closely the baseline follows the data and the smoothness of the estimated baseline. The two parameters, penalization and asymmetry parameters, were tuned empirically to achieve effective baseline correction across the spectral range of 1000–1180 cm⁻¹

[0291] To resolve overlapping spectral features and quantify individual compounds, peak deconvolution was performed using nonlinear least-squares curve fitting with Gaussian functions. Each compound (ethanol, glucose, and lactate) was characterized by a predefined set of characteristic peaks, identified from measured infrared absorbance spectra of individual compounds using the fiber probe and validated with the reference FTIR spectra. The positions of the peaks were fixed based on experimental calibration using fiber probe, whileDocket Number: 10046-664W01the amplitude and width were allowed to vary during fitting. The peak height of each fitted Gaussian function was used for concentration quantification. Calibration curves relating peak height to concentration were generated for each compound.

[0292] To predict the concentration of each compound in the mixture, peak deconvolution was performed using all the characteristic peaks from the absorbance spectra of the pure compounds. The concentrations were then estimated using the corresponding calibration curve. Goodness-of-fit was assessed by calculating the coefficient of determination (R2) between the experimental and fitted spectra.

[0293] Investigation of silver halide fiber dissolution

[0294] To evaluate the potential biocompatibility risks associated with the silver halide fiber probe, a dissolution study to measure the release of silver ions over time was conducted. A fully assembled fiber probe was submerged in 12 mL of ultrapure water and stored at room temperature for 7 days. 1 mL of the solution was withdrawn every day. Prior to analysis, each collected sample was diluted to a final volume of 5 mL with ultrapure water and acidified with nitric acid (HNO3) to a final concentration of 2% (v / v), as required for inductively coupled plasma-mass spectrometry (ICP-MS) measurements.

[0295] The concentration of silver ions in the samples was quantified using inductively coupled plasma mass spectrometry in solution mode. A baseline control sample (ultrapure water before the insertion of probe) was included to confirm background levels.

[0296] Results and Discussion

[0297] FTIR spectroscopy of ethanol, glucose, and lactate in aCSF

[0298] To establish a reference for evaluating the performance of the fiber probe, mid-infrared absorbance spectra of ethanol, glucose, and lactate in artificial cerebrospinal fluid (aCSF) were measured using a benchtop Fourier- transform infrared (FTIR) spectrometer equipped with a 50 um pathlength liquid transmission cell. aCSF was selected as the solvent to simulate the ionic environment of biological fluids.

[0299] FIGS. 26A, 26B and 26C show the absorbance spectra of ethanol, glucose, and lactate, respectively, measured over the spectral range of 1000-1180 cm"1. All compounds were measured at concentrations between 2.5 and 40 mM. 'The spectrum of aCSF was used as the background for absorbance calculation. Ethanol has two prominent C-0 stretching bands at 1046 and 1088 cm⁻¹, consistent with previous literature. Glucose shows multiple peaks, resulting from C-0 stretching vibrations, as expected, the peak positions are centered at 1036, 1066, 1082, 1110, and 1154 cm⁻¹. Lactate presents three distinguishable peaks centeredDocket Number: 10046-664W01around 1042, 1086, and 1124 cm⁻¹. The peak at 1042 cm⁻¹ corresponds to C–C stretching, and the peaks at 1086 and 1124 are corresponding to C–O stretching vibrations.

[0300] For quantification, a single absorbance peak was selected for each compound based on the signal intensity and linear response to concentration. Specifically, 1046 cm⁻¹ was used for ethanol, 1036 cm⁻¹ for glucose, and 1124 cm⁻¹ for lactate as indicated by the red shaded regions. The peak heights were determined by fitting Gaussian functions to the experimental spectra using nonlinear least-squares curve fitting, as described in Section 2.7. The calibration curves shown in FIG. 26D demonstrate strong linearity across the tested concentration range, with correlation coefficients (R2) exceeding 0.99 for all three compounds. The sensitivities, defined as the slope of the calibration curves, for ethanol, glucose and lactate are 2.0 ×10⁻³, 6.8 ×10⁻⁴, 6.3×10⁻⁴ a.u. / mM, respectively. Among the three, glucose has the highest slope in the calibration curve, reflecting its stronger mid-IR absorption due to its higher number of interacting chemical bonds. Lactate showed a slightly steeper slope than ethanol.

[0301] The limits of detection (LoD), calculated as three times the standard deviation of the blank divided by the calibration slope, were 1.8 mM for ethanol, 1.3 mM for glucose, and 1.2 mM for lactate. These results confirm the capability of mid-IR spectroscopy to detect these biologically relevant molecules in aqueous environments at physiologically meaningful concentrations, as summarized in Table 1C.

[0302] Measurement of ethanol, glucose, and lactate solutions using the transflection fiber probe

[0303] To evaluate the performance of the developed mid-infrared transfl ection fiber probe for chemical sensing in aqueous biological environments, ethanol, glucose, and lactate dissolved in artificial cerebrospinal fluid (aCSF) were measured and directly compared with the FTIR results.

[0304] FIGS. 27A, 27B, and 27C each present the absorbance spectra of ethanol, glucose, and lactate, measured using the fiber probe with an average optical pathlength of 63 μm. The same characteristic peaks identified in the FTIR reference measurements were used for quantification: 1046 cm⁻¹ for ethanol (C–O stretch), 1036 cm⁻¹ for glucose (C–O stretch), and 1124 cm⁻¹ for lactate (C–O stretch), as indicated by the red shaded regions. The peak heights were extracted by Gaussian fitting, as in the FTIR measurements. FIG. 27D shows the corresponding calibration curves for all three compounds, each showing linearity with R2values exceeding 0.99. The sensitivities for ethanol, glucose and lactate are 9.4 xlO4, 4.5 xlO’4, 4.3xl0’4a.u. / mM, respectively. Notably, glucose again has the highest sensitivityDocket Number: 10046-664W01among the three compounds. Lactate showed a slightly higher sensitivity than ethanol, with a difference of approximately 5%. Although the fiber probe has a longer pathlength than the FTIR flow cell, it has lower sensitivities for ethanol, glucose and lactate measurements. This may result from the effect of using a wedge-shaped sensing region, which reduces absorbance at higher concentration levels. In addition, the back reflections at the interfaces reduce the sensitivity, according to a previous study. These factors result in lower sensitivity compared to that obtained with FTIR.

[0305] The limits of detection (LoD), determined as three times the standard deviation of the blank divided by the calibration slope, were 1.0 mM for ethanol, 0.8 mM for glucose, and 1.1 mM for lactate. These values are lower than those obtained from the benchtop FTIR setup described herein. Despite lower sensitivity, the fiber probe achieved better limits of detection than the FTIR system, which may be attributed to the high intensity QCL light source, which enables better signal-to-noise ratios.

[0306] One key advantage of the transflection probe is its suitability for in vivo environment. Unlike FTIR transmission cells, which require samples to be contained within a flow cell, the fiber probe interacts directly with the surrounding medium, making it ideal for in situ measurements in tissue.

[0307] These results confirm that the mid-IR transflection fiber probe offers sensitive detection of small biomolecules in aqueous environment, and provides a promising platform for biochemical sensing, especially in settings where traditional FTIR setups are impractical.

[0308] Simultaneous quantification of mixture of ethanol, glucose, and lactate using peak deconvolution

[0309] To demonstrate the capability of simultaneously measuring multiple compounds, mixtures of ethanol, glucose, and lactate in aCSF were measured and analyzed. The same fiber probe with an average optical pathlength of 63 μm was used. Three spectral scans were averaged for each measurement, resulting in a total acquisition time of 1.5 minutes. Peak deconvolution was applied to resolve overlapping absorbance features and quantify each component.

[0310] FIG. 28A shows the mid-IR absorbance spectrum of a mixture containing 20 mM of ethanol, 20 mM of glucose, and 20 mM of lactate. Specific wavenumbers were identified for peak deconvolution based on prior FTIR calibration (Section 3.1). For ethanol, the peaks at 1046 and 1088 cm-1were used; for glucose, 1036, 1066, 1082, 1110, and 1154 cm-1; and for lactate, 1042, 1086, and 1124 cm-1. The fitting achieved an R2of 0.995,Docket Number: 10046-664W01indicating good agreement between the measured spectrum and the sum of individual spectral contributions.

[0311] To validate the peak deconvolution method, three standard mixturescontaining different known concentrations of each compound were tested as summarized in Table 2. FIG. 28B shows the corresponding infrared spectral features of these standard mixtures. The average relative prediction errors for ethanol, glucose and lactate in the three standard mixtures were 10.8, 5.2 and 3.3%, respectively, demonstrating that quantitative multicomponent analysis is feasible with the developed fiber probe. Additionally, theaccuracy of ethanol quantification was slightly more variable, which may be due to peakoverlap with both lactate and glucose. Lactate showed the highest accuracy, likely due to its characteristic peak being separated from those of other compounds, resulting in less spectral interference. Glucose quantification was also reliable, with deviations typically below 1 mM across all conditions.

[0312] Table 2. Prediction of concentration of each compound in the mixture. Values are presented as mean ± SD.Mixture 1 Mixture 2 Mixture 3 Compound Actual Predicted Actual Predicted Actual Picdiclcd IddOSiiObil concentration concentration (mM) (mM) llli^iO||||||||^ Ethanol 20 22.1 ± 1.3 10 9.0 ± 0.6 10 1 1.2 ± 0.5 Glucose 10 9.2 ± 0.3 20 19.3 ± 0.9 10 9.6 ± 0.4 Lactate 11111111111 10 10.3 ± 0.4 11111111111111111^1

[0313] However, due to spectral overlap, the use of peak deconvolution introduces greater uncertainty into the concentration measurements compared to single-component analysis. In single-compound measurements, the signal originates from a single compound, and the quantification relies on direct peak fitting. In contrast, when multiple compounds are present, their spectra overlap in the mid-IR region, leading to shared spectral features and baseline shifts. Deconvolution algorithm resolves these overlapping peaks by fitting multiple parameters, such as peak height and width, for each compound simultaneously. This process increases the degrees of freedom in the fitting model, which in turn propagates fitting errors and increases the variability of predicted concentrations.

[0314] Table 3 shows the relative standard error (RSE) of the peak heights used for quantifying ethanol, glucose, and lactate. The peak heights used to calculate the RSEs wereDocket Number: 10046-664W01measured both individually and within mixtures at concentrations of 10 and 20 mM. For all three compounds, the RSE is consistently higher in the multi-compound measurements compared to single-component analysis. For instance, ethanol shows an RSE increase from 3.4% to 8.9% at 10 mM and from 2.3% to 5.9% at 20 mM when comparing single to multi¬ compound detection. This trend is observed for glucose and lactate as well. The magnitude increase is smaller for lactate, likely due to its more isolated spectral peaks. This increase in RSE directly affects the LoD for each compound in mixtures. The LoD is defined as three times the standard deviation of the measurement divided by the slope of the calibration curve. As the standard deviation increases due to uncertainties in deconvolution, the LoD for multicompound measurements becomes higher than that for single-compound measurements, even if the calibration slope remains similar. As a result, while the fiber probe demonstrates feasibility for simultaneous multi-compound sensing, the achievable detection limits in mixtures are higher than those in single-component analyses (1.0, 0.8, and 1.1 mM for glucose, ethanol and lactate) due to the uncertainty introduced by peak overlap and deconvolution fitting.

[0315] These findings highlight the potential of the fiber probe for simultaneous multi-compound quantification, without the need for prior sample processing or the use of markers to target the molecules. This capability is particularly valuable in biological contexts where monitoring of metabolic markers such as glucose and lactate, along with ethanol, may¬ offer insights into tissue health and metabolic state. However, this capability comes with the trade-off of higher detection limits when resolving complex mixtures through peak deconvolution. Alternative analysis approaches, such as artificial intelligence -based method, could potentially improve sensitivity and robustness in multi-compound measurements.

[0316] Table 3, below, illustrates relative standard error (RSE) of the peak heights used for quantifying ethanol, glucose and lactate. RSE was calculated using peak heights measured as individual components and multi-compound mixtures at 10 and 20 mM concentrations.RSE of ethanol peak RSE of glucose peak RSE of lactate peak height height height Individual In mixiurc Individual In mixture Individual In mixlurcniM 3.4 8.9 3.0 4.0 2.2 3.1 mM 2.3 5.9 2.7 4.6 3.4 3.5Docket Number: 10046-664W01

[0317] Probe’s performance with and without a semi-permeable membrane

[0318] In addition to considerations related to sensor dimensions and ability to quantify concentrations using IR, biofouling - the attachment of cell debris or adsorption of proteins on the fiber - can deteriorate the performance of the sensor. Additionally, previous literature have noted potential biocompatibility concerns for bare silver halide fibers in direct contact with cells. This effect can be mitigated through including a semi -permeable membrane or surface coating. Additionally, a semi-permeable membrane or coating can prevent direct contact of the silver halide fiber to the tissue to prevent potential risk of biocompatibility. A mPES membrane was chosen for this purpose. Its hydrophilic nature helps prevent bubble formation, and its 10 kDa MWCO allows diffusion of small molecules but blocks large particles such as proteins. However, with the membrane, the diffusion dynamics of ethanol, glucose and lactate may change. Two experiments were conducted to investigate the effect of membrane on the sensitivity (slope of the calibration curve) and the response time of the fiber probe.

[0319] Effect of membrane on probe ’s sensitivity

[0320] Glucose calibration curve was measured following the procedure described herein using the same probe with and without a semi -permeable membrane around the sensing region. Measurements were started at least 5 minutes after each sample change to ensure that steady-state concentrations were reached. FIG. 29 shows the calibration curves for glucose in aCSF, measured both with and without tire membrane. The sensitivities of the probe with and without a membrane were 9.38 ×10-4and 9.35 ×10-4, respectively. Difference between the two curves is within 1%. Similar sensitivity with and without membrane is expected, as 5 minutes is sufficient to reach steady-state concentration, as verified in the following section discussing the temporal response of the probe. This also applies to ethanol and lactate, since these smaller molecules diffuse faster than glucose. Consequently, the results demonstrate that the probe performs effectively, despite the presence of the membrane.

[0321] Temporal response with and without a semi-permeable membrane

[0322] Temporal response studies were conducted to examine the effect of a semi-permeable membrane on the fiber probe. Absorbance of 20 mM glucose was monitored over time with a single-scan acquisition time of 25 seconds, both in the presence and absence ofDocket Number: 10046-664W01the membrane. These experiments evaluate the sensor’s capacity to monitor dynamic concentration changes.

[0323] In both conditions, fiber probe with an average optical pathlength of 63 μm was used. The fiber was initially immersed in a vial of aCSF to record baseline signals (t < 0). At t = 0, the vial was changed to another vial containing 20 mM glucose. After a defined period, the vial was changed back to aCSF to observe the signal decay over time. The response curves are shown in FIG. 30A and FIG. 30B for the probe without a membrane and the membrane-protected probe, respectively. The response was quantified by calculating the time constant, defined as the time required for the signal to reach 90% of its maximum (rise).

[0324] For the probe without a membrane, the response was faster, with a time constant of 33.4 ± 1.8 (mean ± SD) seconds. In contrast, the membrane-protected probe showed a slower response, with a time constant of 59.1 ± 3.5 (mean + SD) seconds. The increase in response time reflects the added diffusion resistance introduced by the membrane. The semi-permeable membrane will help with filtering out large macromolecules and particles, potentially reducing biofouling and signal interference, but it slows the transport of molecules to the sensing region. Despite the slower response, the membrane-protected probe still reached 90% of the steady-state signal within one minute, indicating that real-time monitoring remains feasible for many applications.

[0325] The decay curves after switching back to aCSF also support these findings. The decay time constants, defined as the times required for the signal to decrease to 10% of the maximum (when immersed in glucose solution), were 36.1 ± 2.0 (mean + SD) and 61.7 ± 2.9 (mean + SD) seconds for probe with and without a membrane, respectively. The unprotected probe showed a sharp drop in signal, returning near baseline within 1 minute. In comparison, the membrane-equipped probe displayed a more gradual signal decay, consistent with residual glucose diffusing out of the membrane -protected sensing zone.

[0326] Overall, these results demonstrate that while the membrane reduces the rate of mass transfer, and thus slightly compromising the temporal response, the concentration changes of ethanol, glucose and lactate changes in the human body is typically occur slowly enough to be captured with a temporal resolution of 1 minute. The choice of using a membrane should depend on the specific application requirements, whether fast response or selective filtering is prioritized.

[0327] Ex vivo skin measurement using the fiber probe and comparison with microdialysisDocket Number: 10046-664W01

[0328] Side by side comparison with microdialysis. To compare the performance of the fiber probe and microdialysis, concentrations in ex vivo human skin were monitored by both techniques following ethanol spiking of the medium under the skin. The mid-IR fiber probe (with a semi-permeable membrane and an averaged optical pathlength of 63 μm) and the microdialysis probe were inserted into the same skin sample, positioned at depths of 2.7 and 2.9 mm, respectively. Ethanol was added to the culture medium at t=0 to achieve a final concentration of 20 mM. FIG. 31 A shows the ethanol concentrations measures by both techniques over time. The concentrations increased after ethanol addition, with the rate of increasing slowing approximately 30 minutes later. The temporal concentration profiles reflect the gradual penetration of ethanol into the skin. The fiber probe measured a maximum tissue concentration of approximately 8 mM, which is lower than the spiked medium concentration, as expected. This discrepancy is due to diffusion barriers within the skin and underlying nourishing matrix, which slows down the transport process of ethanol to the region where the probes were embedded. Compared to microdialysis, which operates at a low perfusion rate (0.5 μL / min) and requires a 10-minute sampling interval, the fiber probe provides better temporal resolution, allowing faster tracking of concentration changes over time. Additionally, the concentration measured using microdialysis was slightly lower than those obtained with the fiber probe. This may be attributed to the evaporation losses of ethanol during sample collection and handling required for the microdialysis process.Moreover, because the molecules diffused into the microdialysis probe were continuously carried to the sample collection vial at outlet, local depletion of molecules in the surrounding tissue can occur. This is a limitation not encountered with the non-extractive, in situ sensing mechanism of the fiber probe.

[0329] Simultaneous measuremen t of ethanol, glucose and lactate in the skin tissue. To demonstrate the fiber probe’s capability for measuring mixtures in ex vivo human skin, the culture medium beneath the skin sample was spiked with 100 mM ethanol, 100 mM glucose and 100 mM lactate and their concentrations were measured over time. As shown in FIG. 3 IB, all three compounds were simultaneously detected using peak deconvolution. Each compound has a distinct temporal concentration profile, even though their initial spiked concentrations were the same. Ethanol showed the highest concentration, while glucose and lactate showed lower levels. These differences result from their distinct diffusion coefficients of each molecule within skin tissue. Ethanol, due to its small size, diffuses more easily through both the nourishing matrix and the skin tissue above it compared to glucose and lactate. As a result, the steady-state concentration of ethanol is higher than that of glucose andDocket Number: 10046-664W01lactate. Among the three, glucose is the largest molecule, leading to the slowest diffusion rate and the lowest steady-state concentration.

[0330] Temporal response of the probe

[0331] This experiment evaluated the temporal response of the fiber probe in capturing dynamic concentration changes. In order to capture the concentration change, spectra were collected every 30 seconds using single-scan acquisition to minimize measurement time. FIG. 31C shows measurements of ethanol concentration overtime following the injection of 50 μL of 200 mM ethanol into the skin near the probe at t - 0 min.

[0332] A quick increase in measured concentration was observed within the first 10 minutes. At t = 30 min, 20 μL of culture medium was added directly on the skin near the probe to quickly dilute the local ethanol concentration. The corresponding decrease was detected by the fiber probe. The results underscore the capability of the fiber probe to capture concentration variations on sub-minute timescales.

[0333] Overall, these findings highlight two main advantages of the mid-IR fiber probe over microdialysis: (1) in situ measurement without the need for sample extraction or processing, (2) high temporal resolution suitable for capturing dynamic changes. These features make the developed probe a promising tool for in vivo molecular monitoring in tissues.

[0334] Evaluation of silver ion release and biocompatibility considerations

[0335] To assess the potential risk of the release of ions from the silver halide fiber in the probe for biological applications, the dissolution of silver halide in aqueous environments was investigated. The concentration of silver ions was measured as an indicator of the extent of fiber dissolution. This was quantified by immersing a fiber probe with a semi-permeable membrane in ultrapure water for 7 days. Samples were collected daily and analyzed using an ICP-MS.

[0336] As shown in FIG. 32, the measured silver ion concentration remained low throughout the test period, with a maximum concentration below 26 ppb. The control sample, prepared before fiber immersion, showed an undetectable silver ion concentration (0 ppb), confirming minimal background contamination. The highest concentration of silver ions is well below known cytotoxic thresholds and is considered safe for biomedical applications. Toxicity effects are generally reported at levels around 1000 ppb for human fibroblasts. In addition, the amount of dissolved chloride and bromide ions can be estimated based on the measured silver ion concentration, as these halide ions are released in stoichiometric proportions during the dissolution of the silver halide material. Therefore, the concentrationsDocket Number: 10046-664W01of chloride and bromide ions are expected to be lower than the silver ion levels detected in the solution. Notably, the physiological concentrations of chloride and bromide ions in human body are substantially higher than these levels, further indicating that the small quantities potentially released from the fiber probe are unlikely to pose a health risk.

[0337] In the example probe design, a semi-permeable PES membrane encloses the entire sensing region, creating a physical barrier that further prevents direct tissue contact with the silver halide fiber, which can enhance the overall biocompatibility of the device. Based on the low silver ion concentrations measured and the use of the membrane, the results support the potential use of the probe in short-term ex vivo and in vivo applications.

[0338] Discussion

[0339] In this study, a compact mid-infrared (mid-IR) transflection fiber probe capable of label-free quantification of multiple biologically relevant molecules in aqueous environments was developed and demonstrated in ex vivo human skin. The probe design consists of two 500 μm diameter silver halide fibers transparent over the wavelengths from 3 to 17 μm. One fiber end has an angled sensing tip to reduce Fresnel reflections, and the other fiber has a gold-coated end to provide mirror-like reflectivity. The two ends were aligned face-to-face within a miniaturized PEEK tube, forming a sensing region with a sample volume of 13 nL and an averaged optical pathlength of 63 μm. This region was enclosed by a semi -permeable mPES membrane with a 10 kDa molecular weight cut-off and pore sizes of ~3–5 nm. The membrane excludes potential infrared interferants from entering the sensing region and prevents direct contact between silver halide fibers and the tissue, which could enhance biocompatibility and reduce biofouling. The resulting fiber probe has an overall outer diameter of only 1.1 mm.

[0340] The fiber probe’s capability for infrared spectroscopic quantification was evaluated against a conventional laboratory-grade benchtop P IR system. A benchtop FTIR equipped with a 50 μm optical path length transmission cell served as the reference method for quantifying ethanol, glucose, and lactate in artificial cerebrospinal fluid (aCSF). The detection limits for ethanol, glucose, and lactate in aCSF were determined to be 1.8, 1.3, and 1.2 mM, respectively. Fiber probe measurements for detecting these compounds were performed using a quantum cascade laser system as light source. The limits of detection for ethanol, glucose and lactate were 1.0, 0.8, and 1.1 mM, respectively. The limits of detection for the fiber probe are lower compared to those achieved with the conventional FTIR system, which may be attributed to the higher intensity of the QCL light source compared to the thermal emission source used in the FTIR. This demonstrates that the compact fiber probeDocket Number: 10046-664W01can achieve a sensitivity comparable to laboratory-based instruments while offering the advantage of in situ measurements.

[0341] Spectral analysis of mixtures is more complex as infrared peaks of these three compounds overlap. To address spectral overlaps in multi-compound detection, peak deconvolution was used to quantify each compound in mixtures. The average relative errors of ethanol, glucose and lactate were 10.8 ± 1.0%, 5.2 ± 2.5% and 3.0 ± 1.0% (mean + SD), respectively. Lactate showed the highest accuracy, likely due to minimal spectral interference, while the predicted ethanol concentrations have greater variability because the peak used for its quantification overlaps with the absorbance peaks of glucose and lactate, compromising the accuracy of peak deconvolution. While this approach enables simultaneous multi-compound detection without the need of chemical labeling, it introduces greater uncertainty compared to single-component analysis due to fitting errors. Consequently, the relative standard error for each compound is higher in mixture analysis, leading to an increased LoD.

[0342] To avoid direct contact between the silver halide fiber and the tissue and to block potential interferant in biological samples, the sensing region of the probe was enclosed by a semi-permeable membrane. However, incorporating the membrane affects diffusion dynamics. The probe’s temporal response showed a time constant of 59.1 + 3.5 (mean ± SD) seconds with the membrane and 33.4 + 1.8 (mean ± SD) seconds without it when measuring 20 mM glucose in aCSF. These results confirm its ability to track dynamic concentration changes, demonstrating its suitability for applications requiring compound concentration monitoring.

[0343] Ex vivo experiments using human skin models validated the probe’s performance in tissue. The fiber probe achieved simultaneous quantification of ethanol, glucose, and lactate concentrations in human skin samples, showing clear advantages over conventional microdialysis techniques in terms of temporal resolution and multi-compound detection.

[0344] The miniaturized diameter of 1.1 mm and the protective semi -permeable membrane highlight the probe’s potential for in vivo applications where minimally invasive, continuous monitoring is essential. Overall, this mid-IR transflection fiber probe offers a useful platform for biomedical sensing, with applications in metabolic monitoring, disease diagnostics, and therapeutic guidance.

[0345] The present disclosure contemplates further miniaturing both the probe and the overall measurement system, enhancing the probe’s sensitivity and validating its performanceDocket Number: 10046-664W01in in vivo models to advance toward clinical applications. For example, the current large infrared laser system and optical components can be replaced with portable alternatives. Identifying suitable light sources and detectors that balance size, sufficient power, and cost¬ effectiveness will be critical for translating this technology into practical, handheld devices. Additionally, further studies on biocompatibility are important, including long-term evaluations of the effect of silver halide fiber probe on tissue, to ensure safe use in vivo.

[0346] Computing Device

[0347] Referring to Fig. 33, an example computing device 3300 upon which the methods described herein may be implemented is illustrated. It should be understood that the example computing device 3300 is only one example of a suitable computing environment upon which the methods described herein may be implemented. Optionally, the computing device 3300 can be a well-known computing system including, but not limited to, personal computers, servers, handheld or laptop devices, multiprocessor systems, microprocessorbased systems, network personal computers (PCs), minicomputers, mainframe computers, embedded systems, and / or distributed computing environments including a plurality of any of the above systems or devices. Distributed computing environments enable remote computing devices, which are connected to a communication network or other data transmission medium, to perform various tasks. In the distributed computing environment, the program modules, applications, and other data may be stored on local and / or remote computer storage media.

[0348] In its most basic configuration, computing device 3300 typically includes at least one processing unit 3306 and system memory 3304. Depending on the exact configuration and type of computing device, system memory 3304 may be volatile (such as random access memory (RAM)), non-volatile (such as read-only memory (ROM), flash memory, etc.), or some combination of the two. This most basic configuration is illustrated in Fig. 33 by box 3302. The processing unit 3306 may be a standard programmable processor that performs arithmetic and logic operations necessary for operation of the computing device 3300. The computing device 3300 may also include a bus or other communication mechanism for communicating information among various components of the computing device 3300.

[0349] Computing device 3300 may have additional features / functionality. For example, computing device 3300 may include additional storage such as removable storage 3308 and non-removable storage 3310 including, but not limited to, magnetic or optical disks or tapes. Computing device 3300 may also contain network connection(s) 3316 that allow theDocket Number: 10046-664W01device to communicate with other devices. Computing device 3300 may also have input device(s) 3314 such as a keyboard, mouse, touch screen, etc. Output device(s) 3312 such as a display, speakers, printer, etc. may also be included. The additional devices may be connected to the bus in order to facilitate communication of data among the components of the computing device 3300. All these devices are well known in the art and need not be discussed at length here.

[0350] The processing unit 3306 may be configured to execute program code encoded in tangible, computer-readable media. Tangible, computer-readable media refers to any media that is capable of providing data that causes the computing device 3300 (i.e., a machine) to operate in a particular fashion. Various computer-readable media may be utilized to provide instructions to the processing unit 3306 for execution. Example tangible, computer-readable media may include, but is not limited to, volatile media, non-volatile media, removable media and non-removable media implemented in any method or technology for storage of information such as computer readable instructions, data structures, program modules or other data. System memory 3304, removable storage 3308, and non-removable storage 3310 axe all examples of tangible, computer storage media. Example tangible, computer-readable recording media include, but are not limited to, an integrated circuit (e.g., field -programmable gate array or application-specific IC), a hard disk, an optical disk, a magneto-optical disk, a floppy disk, a magnetic tape, a holographic storage medium, a solid-state device, RAM, ROM, electrically erasable program read-only memory (EEPROM), flash memory or other memory technology, CD-ROM, digital versatile disks (DVD) or other optical storage, magnetic cassettes, magnetic tape, magnetic disk storage or other magnetic storage devices.

[0351] In an example implementation, the processing unit 3306 may execute program code stored in the system memory 3304. For example, the bus may carry data to the system memory 3304, from which the processing unit 3306 receives and executes instructions. The data received by the system memory 3304 may optionally be stored on the removable storage 3308 or the non-removable storage 3310 before or after execution by the processing unit 3306.

[0352] It should be understood that the various techniques described herein may be implemented in connection with hardware or software or, where appropriate, with a combination thereof. Thus, the methods and apparatuses of the presently disclosed subject matter, or certain aspects or portions thereof, may take the form of program code (i.e., instructions) embodied in tangible media, such as floppy diskettes, CD-ROMs, hard drives, orDocket Number: 10046-664W01any other machine-readable storage medium wherein, when the program code is loaded into and executed by a machine, such as a computing device, the machine becomes an apparatus for practicing the presently disclosed subject matter. In the case of program code execution on programmable computers, the computing device generally includes a processor, a storage medium readable by the processor (including volatile and non-volatile memory and / or storage elements), at least one input device, and at least one output device. One or more programs may implement or utilize the processes described in connection with the presently disclosed subject matter, e.g., through the use of an application programming interface (API), reusable controls, or the like. Such programs may be implemented in a high level procedural or object-oriented programming language to communicate with a computer system. However, the program(s) can be implemented in assembly or machine language, if desired. In any case, the language may be a compiled or interpreted language and it may be combined with hardware implementations.

[0353] Although the subject matter has been described in language specific to structural features and / or methodological acts, it is to be understood that the subject matter defined in the appended claims is not necessarily limited to the specific features or acts described above. Rather, the specific features and acts described above are disclosed as example forms of implementing the claims.

Claims

Docket Number: 10046-664W01 WHAT IS CLAIMED:

1. An optical sensor comprising:a first optical fiber defining an axis configured to transmit light; a second optical fiber comprising a reflection surface formed on an end face of the second optical fiber, wherein the reflection surface is disposed at an angle relative to the axis; anda sleeve or a semi-permeable membrane configured to fix second optical fiber to the axis and to fix the reflection surface at a predetermined separation distance from a distal end of the first optical fiber and thereby define a sensing area between the optical fiber and reflection surface;wherein, the sleeve or a semi-permeable membrane, first optical fiber, and second optical fiber are configured so that when the first optical fiber is illuminated, light from the first optical fiber traverses the sensing area, reflects from the reflection surface, and is returned through the first optical fiber.

2. The optical sensor of claim 1, wherein the reflection surface comprises a metallic coating3. The optical sensor of claim 2, wherein the reflection surface comprises a gold coating.

4. The optical sensor of claim 1 or claim 2, wherein the sleeve comprises a sensing window configured to permit a sample fluid to enter the sensing area.

5. The optical sensor of any one of claims 1-3, wherein the sleeve comprises a semipermeable membraneDocket Number: 10046-664W016. The optical sensor of claim 5, wherein the semipermeable membrane comprises a cellulose -based material.

7. The optical sensor of any one of claims 1-5, wherein the sleeve comprises polyetheretherketone.

8. The optical sensor of any one of claims 1-7, wherein the optical fiber comprises silver halide.

9. The optical sensor of any one of claims 1-8, wherein the sensing area is less than about 100 micrometers in length between the reflection surface and the distal end of the first optical fiber.

10. The optical sensor of any one of claims 1-9, wherein the sleeve defines a diameter of less than 2 millimeters.

11. The optical sensor of any one of claims 1-10, wherein the angle relative to the axis is 10 degrees or greater.

12. An optical sensor comprising:an optical fiber defining an axis configured to transmit light;a reflection surface;a sleeve or a semi-permeable membrane configured to fix the reflection surface in place so that the reflection surface is along the axis of the optical fiber and at a fixed distance from the optical fiber.Docket Number: 10046-664W01 13. The optical sensor of claim 12, wherein the optical fiber comprises an angled surface in contact with the reflection surface, and wherein fixed distance from the optical fiber is defined by a length of the angled surface along the axis.

14. The optical sensor of claim 12 or 13, wherein the reflection surface comprises an optical fiber having a reflective end face.

15. The optical sensor of any one of claims 12-14, wherein the reflection surface comprises a metallic coating.

16. The optical sensor of any one of claims 12-15, further comprising a semipermeable membrane17. The optical sensor of claim 16, wherein the semipermeable membrane comprises a cellulose membrane.

18. The optical sensor of any one of claims 12-17, wherein the optical fiber comprises silver halide.

19. A system comprising:an optical sensor according to any one of claims 1-18a light source configured to illuminate an optical fiber of the optical sensor;a light sensor configured to measure light from the optical sensor;a controller operatively coupled to the light sensor and light source and configured to determine a concentration of at least one analyte based on a measurement of the light sensor.Docket Number: 10046-664W0120. The system of claim 19, wherein the concentration comprises a concentration of at least one of: a glucose measurement, an ethanol measurement, and a lactate measurement.

21. The system of claim 20, wherein the light source is an infrared light source.

22. The system of claim 21, wherein the light source is a mid-infrared light source.

23. The system of any one of claims 19-22, wherein the light source is a quantum cascade laser.

24. The system of any one of claims 19-23, wherein the controller is further configured to estimate an absorbance by eliminating reflections, and wherein the concentration is based at least in part on the absorbance.

25. The system of any one of claims 19-24, wherein the controller is configured to eliminate reflections according to: >4&s(c) = log10.S\vX101 c+7?26. The system of any one of claims 19-25, wherein the controller is configured to eliminate reflections according to: Abs(, c) = log1027. The system of any one of claims 19-26, wherein the light sensor comprises an infrared photodetector.

Images