Wearable spectrometer for biomolecular investigation in biological tissues

Abstract

A method and device for noninvasively diagnosing the condition of subcutaneous biological tissue using biomolecular Raman spectroscopy. The device is secured to a user's skin so that a light source can emit photons of light through a bottom port that probe physiological biomarkers in molecules of interest within the subcutaneous tissue. The mixed photons of Rayleigh and Raman scattered light return through the port to the internal cavity. The Rayleigh scattered light is filtered to limit it to specific wavelengths, and then an array of photodetectors detects the Raman scattered light by counting photons one by one over a predetermined sampling time. The device and method can be configured in a wearable form, such as a wristband, to monitor various conditions, including blood glucose readings.

Description

[Technical Field]

[0001] Cross-reference of related applications This application claims priority to U.S. Provisional Patent Application No. 63 / 062,478, filed on 7 August 2020, the full disclosure of which is incorporated herein by reference and relied upon.

[0002] The present invention relates to a wearable, compact spectrometer device generally suitable for non-invasive investigation of subcutaneous molecules within biological tissues. [Background technology]

[0003] Diabetes mellitus is an incurable chronic metabolic disease in which the body is unable to produce enough insulin and therefore unable to control the glucose concentration in the blood. Currently, there are more than 100 million people with diabetes worldwide, and the World Health Organization predicts that this number will exceed 400 million in the next decade. Diabetes is diagnosed and managed by measuring blood glucose levels. People with diabetes need to closely monitor and control their glucose levels, measuring them several times a day. The current approach, which has been in use for decades, is to measure glucose using a small amount of blood taken from the finger. This repeated finger-prick blood collection is painful, carries a risk of infection, and is a major cause of treatment non-compliance. Alternative sensor devices that can accurately measure glucose levels without the discomfort of blood sample collection for clinicians and patients are not yet widely implemented due to the complexity and cost of the necessary equipment. An international panel of experts recently concluded that real-time continuous glucose monitoring (CGM) would be a major step forward for diabetes management if the technical challenges could be overcome.

[0004] Effective treatment of diabetes requires frequent (ideally continuous) monitoring of blood glucose levels to help patients maintain their basic health and avoid potentially life-threatening events such as hypoglycemia, stroke, or heart attack. For decades, the most advanced technology in glucose monitoring has been portable electrochemical devices that use physical blood samples obtained through inconvenient and painful fingertip pricks. More recently, so-called minimally invasive technologies have emerged that use remote transmitters and subcutaneous probes / patches (micro-needles, wires inserted under the skin, or other implantable probes) to collect blood glucose data in real time and transmit it to smartphones or other receivers. This approach has limitations, including the quality of the probe / sensor, signal interference from interstitial fluid and tissue structure, and consequently, the need for sophisticated signal analysis to indirectly approximate actual blood glucose levels. Furthermore, the high cost of disposable probes is also a challenge. The next and final step is a truly non-invasive blood glucose meter that enhances patient comfort and compliance, provides accurate, real-time, actionable data, and can be easily integrated into modern wearable smartwatch platforms.

[0005] Currently, many university and industry teams are pursuing the goal of non-invasive glucose meters, publishing research involving lasers, optical fibers, millimeter waves, ultrasound, benchtop confocal Raman spectrometers, and even bio-functional nanoparticles injected into patients. While these studies have shown that transcutaneous spectroscopy can be as effective as finger-prick testing, a fundamental challenge remains: detecting limited signals in a complex and dynamic environment requires large, expensive spectrometers and other equipment, making it unsuitable for low-cost, non-invasive, and wearable applications.

[0006] Patient dissatisfaction is widespread regarding the functional lifespan of disposable sensors (typically 5-14 days), sensor skin penetration, discomfort associated with patch / transmitter attachment and protrusion from the body, and the long-term costs of monitoring. Such dissatisfaction is particularly intense among Type 2 patients who do not understand the cost-benefit value of continuous monitoring.

[0007] More importantly, current solutions still require frequent calibration checks, and while this approach may be less invasive in the sense that it involves minimal physical intrusion during the measurement process, from the end customer / patient's perspective, it is far from minimally invasive because frequent direct measurement of blood glucose levels is still necessary for calibration and safety. At present, there is no technology that monitors blood glucose levels 100% non-invasively and painlessly in a wearable form such as a real-time CGM. Furthermore, current technologies, while implying direct measurement, tend to be indirect measurement methods using artificial intelligence / AI, which usually result in delayed and quantitatively questionable data.

[0008] Since the introduction of the first CGM from MiniMed / Medtronic in 1999, many innovative attempts have been made to develop a truly non-invasive CGM, but to date, no such device is on the market, nor is any nearing commercialization. Due to this drawback of commercially available options, diabetic patients must manage their glucose levels using subcutaneously implanted probes, which must be replaced every few days or weeks, or, in the case of some devices (approved only in the EU), every few months. Even if minimally invasive, current CGM devices still require frequent calibration via painful daily finger punctures in most cases. To date, a truly non-invasive CGM has not received FDA approval, mainly because of its low sensitivity, particularly at the lower end of the diabetes spectrum (<70 mg / dL), and lack of specificity. The latter problem is particularly limited, as many of the methods proposed for non-invasive probes do not directly measure specific correlations with glucose concentrations in the blood or interstitial fluid. In fact, since almost all physiological parameters of the body show some correlation with oral glucose intake, conventional attempts at non-invasive blood glucose probes (e.g., IR transmission / absorption, acoustic coupling, tissue light reflectance, microwave response, etc.) are generally confused by the responses of multiple physiological tissues and cannot provide reliable quantitative measurements of actual blood glucose levels. Therefore, it is crucial to find a way to directly bind the sensor probe to the glucose molecule itself. Unfortunately, there are few non-invasive probe approaches that can do this directly, and even fewer that can be miniaturized sufficiently for wearable devices.

[0009] Raman spectroscopy is an optical technique that can directly detect the intrinsic vibrational fingerprint frequency of glucose molecules. Theoretically, Raman spectroscopy, which offers millimolar sensitivity, can provide accurate measurements over the physiological concentration range of glucose from 4 to 10 mmol / L. For example, see JINGWEI SHAO et al., IN VIVO BLOOD GLUCOSE QUANTIFICATION USING RAMAN SPECTROSCOPY, PLOS ONE 7, E48127 (October 2012). DOI:10.1371 / JOURNAL.PONE.0048127. See Figure 1, reprinted from SOURCEPUBMED. However, there are several challenges to the practical application of Raman spectroscopy in the field of continuous glucose monitoring (CGM). One is the extremely small Raman scattering cross-section (approximately 1 / 10). 7 ) Another factor is the presence of glucose-unrelated scattering, which stems from the complex chemical properties of the blood tissue matrix where the Raman signal is generated. Finally, in terms of commercializing practical / wearable CGM devices, the large volume and geometric shape of conventional Raman spectrometers do not help in miniaturization.

[0010] US20200107756A1 describes a silicon photomultiplier (SiPM) array-based multispectral optical probe for image-guided radiotherapy used for low-light detection of Cherenkov emission (CE) in relation to tumor detection and treatment. However, no system integration of this technology into a wearable, real-time continuous diagnostic spectrometer is known.

[0011] Therefore, in this field, there is a need for non-invasive, wearable, real-time continuous diagnostic spectrometers and methods useful for diagnosing specific target conditions present in subcutaneous biological tissue. There is a need for painless devices and methods for investigating subcutaneous molecules that can continuously generate highly accurate and reliable information. While Raman spectroscopy is promising, there is no known method for performing Raman spectroscopy in a small, wearable form. [Overview of the project]

[0012] According to a first aspect of the present invention, a non-invasive, wearable diagnostic spectrometer device includes a housing. The housing has side walls surrounding an internal cavity. A cover is positioned above the side walls and the internal cavity. A base is positioned below the side walls, at least partially surrounding the internal cavity. The base has a port inside configured to allow the passage of photons. Lashing (fixing elements) extend from the housing and are configured to fix the housing to the user's skin in the area of ​​interest. A light source is configured to emit photons of light through the port to probe (examine) physiological biomarkers of target molecules in subcutaneous tissue, and to return photons, a mixture of Rayleigh-scattered and Raman-scattered light, to the internal cavity through the port. An array of multiple photodetectors is positioned within the housing. Each photodetector has a separate channel configured to detect Raman-scattered photons entering the cavity through the port. Each photodetector is associated with at least one optical filter. At least one optical filter is operably positioned between the associated photodetector and the port. Each optical filter restricts the passage of light reaching its associated photodetector to a specific wavelength and excludes Rayleigh scattered light. A data acquisition electronic module (data acquisition electronic module) is operably associated with the array of photodetectors. The data acquisition electronic module is configured to count single photons of Raman scattered light reaching each photodetector over a predetermined sampling time.

[0013] The present invention also considers a method for non-invasively diagnosing the condition of subcutaneous biological tissue using Raman spectroscopy. This method includes a series of steps, including the placement of multiple photodetectors within an internal cavity. The internal cavity is fixed directly to the user's skin. Light is emitted directly onto the user's skin from a light source within the internal cavity through a port. The light is used to investigate at least one subcutaneous molecule beneath the user's skin. This investigation step generates optical photons composed of Rayleigh scattered and Raman scattered light, which are re-incidentated into the internal cavity through the port. The multiple photodetectors are shielded within the internal cavity from the light emitted from the light source, but not from the Raman scattered light re-incidentated into the internal cavity through the port. Rayleigh scattered light is filtered out from the photons re-incidentated into the internal cavity through the port by means of a narrowband optical filter, etc. The Raman scattered light reaching each photodetector is then limited to specific wavelengths associated with the Raman active spectral line. Finally, the single photons received by each photodetector are counted over a predetermined sampling time to measure the integrated intensity of the selected Raman active line.

[0014] In both apparatus and method forms, the present invention enables a truly non-invasive, painless, and wearable real-time Raman spectroscopy diagnostic tool that dramatically improves users' lives and enhances health outcomes. The multispectral probe architecture, in a wearable form factor, readily adapts to a variety of conditions, such as reading blood glucose and other target health parameters. The present invention utilizes weak Raman scattered light transmitted through tissue (after investigation with a self-integrated light source) to enable spectral measurement of relevant molecular biomarkers. The inelastic scattered light characteristic of Raman spectroscopy is weak, leaving only a few measurable photons that convey information about the target, and only a subset of these photons within a specific frequency range is useful for molecular measurement. Using Raman spectroscopy, the present invention is useful for targeting a small fraction of scattered photons that convey the fingerprint or concentration of the target molecule. Because it does not require a physically large and expensive spectrometer, the present invention enables a small, wearable, low-cost, and low-power apparatus and method for directly and non-invasively measuring target molecular biomarkers.

Brief Description of the Drawings

[0015] These and other features and advantages of the present invention will be more readily understood when considered in connection with the following detailed description and the accompanying drawings.

[0016] [Figure 1] A graph reproduced from Jingwei Shao et al., In Vivo Blood Glucose Quantification Using Raman Spectroscopy, showing the vibration fingerprint frequencies unique to glucose molecules.

[0017] [Figure 2A] An example of a wearable device according to the present invention in a wristband form factor.

[0018] [Figure 2B] An example of a wearable device according to the present invention in a fingertip clamp form factor.

[0019] [Figure 2C] An example of a wearable device according to the present invention in an adhesive patch form factor.

[0020] [Figure 2D] An example of a wearable device according to the present invention in an earlobe clip form factor.

[0021] [Figure 3] A simplified partial side view of a wearable device according to the present invention in a wristband form factor similar to that shown in FIG. 2A, configured for wired and wireless communication with a remote computing and / or treatment device.

[0022] [Figure 4] A bottom view of the wearable device generally along line 4-4 of FIG. 3.

[0023] [Figure 5] This is a cross-sectional view of a wearable device roughly along line 5-5 in Figure 4.

[0024] [Figure 6] This figure is similar to Figure 5, showing a wearable device fixed to the user's skin in a working position within the target region, with the target molecule located in the subcutaneous tissue.

[0025] [Figure 7] This diagram is similar to Figure 6, where photons emitted through the port investigate the target molecule, and the photons, a mixture of Rayleigh and Raman scattered light, return through the port to the internal cavity. [Modes for carrying out the invention]

[0026] This invention describes a non-invasive, wearable, real-time continuous diagnostic spectrometer and method of use that can diagnose and monitor the state of a target in subcutaneous biological tissue. The invention investigates subcutaneous molecules in a target area using a painless method that can continuously generate highly accurate and reliable information. Many deep tissue measurement applications of the invention include, but are not limited to, continuous glucose monitoring (CGM), toxin screening, cancer cell detection, tumor pH, oxygenation, radiation dose during cancer radiotherapy, and proteins (troponin) associated with myocardial infarction risk. In other words, the wearable device of the invention is widely understood as a spectroscopic sensor platform for a wide range of biomolecular investigations. The wearable device can be used to investigate physiological biomarkers of the target molecule regardless of its location, such as in blood, skin, or tissue lipid membranes.

[0027] As a wearable device, the present invention can manifest in various forms. Several examples are shown in Figures 2A to 2D. In Figure 2A, the wearable device 10 is shown in the form of an electronic bracelet with an appearance similar to a smartwatch. In Figure 2B, the wearable device, generally referred to as 10', is shown in the form of a fingertip sensor. In Figure 2C, the wearable device 10'' is shown as a skin patch. Also in Figure 2D, the wearable device 10'''' is in the form of an earlobe clip. Those skilled in the art will understand other forms of wearable medical sensors that can be modified or adapted for use in connection with the present invention, such as configurations that are worn on or around the head, neck, shoulders, arms, hands, chest, waist, legs, ankles, feet, etc.

[0028] For convenience, the wearable device 10 will be described in an exemplary wrist-mounted configuration as shown in Figure 2A. However, it should be understood that the present invention is not limited to the wrist-mounted configuration. Any of the configurations mentioned, and configurations not mentioned but readily understood by those skilled in the art, are considered to be suitable alternative configurations for carrying out the present invention.

[0029] Referring to the examples in Figures 3 to 5, the wearable device 10 includes one (or more) housings that contain highly sensitive electronic and optical hardware. The housing has side walls 12 surrounding an internal cavity. In the illustrated examples, the side walls 12 are substantially circular or cylindrical. However, in other possible embodiments, the side walls 12 may have a different geometric shape or be asymmetrical. A cover 14 is positioned over the side walls 12 and the internal cavity. The cover 14 is generally the uppermost or outermost visible feature of the wearable device 10. In some embodiments, a display screen or graphic user interface (GUI) 16 may be mounted outside the cover 14. For example, a simple display screen 16 can display information in the form of text, images, videos, graphs, etc. It can also be configured as a GUI 16 and can receive input from the user for setting up the wearable device, changing attributes, sending and receiving data, etc. Although not shown, the wearable device 10 may include a speaker to transmit audible messages or tones / beeps that communicate relevant information or alerts to the user. Similarly, additional user interface elements such as buttons, LEDs, dials, and / or touch-sensitive elements may be located on the housing or any other function of the wearable device 10.

[0030] As schematically shown in Figure 3, the wearable device 10 can be configured to communicate with a remote computing device 18 via wired and / or wireless connections. The wearable device 10 may also be equipped with appropriate data transmission / reception capabilities that enable connection to the Internet, the World Wide Web, or other desired networks. Naturally, precautions will be taken to ensure the secure transmission of data to and from the wearable device 10 so that only the user and other authorized individuals can remotely access the wearable device 10. The ability to communicate with the remote computing device 18 provides the user, the user's caregiver, and other authorized individuals with the ability to monitor diagnostic information generated by the wearable device 10 and manage the operation of the wearable device 10. Furthermore, notifications can be sent and received between the user and / or the user's caregiver and / or authorized healthcare professionals via the remote computing device 18. Such notifications may originate from measurements taken by the wearable device 10 and / or alarms triggered when certain conditions are met. Furthermore, it can send and receive data with the controller 19 of remote treatment devices such as insulin and other drug pumps, defibrillators (wearable or implantable), and brain stimulators.

[0031] The base 20 is positioned beneath the side wall 12 and at least partially closes the internal cavity. In many cases, including those shown in Figures 3-7, the base 20 is designed to press directly against the user's skin 22 directly over the area of ​​interest where a diagnostic examination of subcutaneous biological tissue is to be conducted. In the case of the wrist wearable device 10 as shown in Figures 3-7, the subcutaneous biological tissue to be examined must be located very close to the wrist. Furthermore, the base 20 can be curved for comfort, depending on the intended use. In the example in Figure 2B, the base (not easily visible) is molded to hold a human finger.

[0032] The base 20 has a port 24 formed inside, as shown in Figures 4 and 5. The port 24 is configured to allow the passage of optical photons. That is, optical photons can be transmitted through the port 24, as suggested in Figure 7. Such transmittance of optical photons can take various forms. In some hypothetical examples, the port 24 is a film or window-like element that can freely pass desired optical photons and optionally filter out unwanted light (e.g., Rayleigh scattered light) or other electromagnetic signals. However, in the illustrated example, the port 24 has an uncovered opening, thus allowing direct movement of optical photons between the internal cavity and the user's skin 22 beneath it.

[0033] The port 24 may have any suitable shape. In the embodiment shown in Figure 4, the port 24 is substantially circular, but other shapes are certainly possible. The port 24 may also consist of multiple separate, smaller ports rather than one large opening. The accompanying figure shows an elastomer gasket 26 surrounding the port 24. The main function of the gasket 26 may be to provide a complete light-shielding seal between the base 20 and the user's skin 22 directly beneath the internal cavity, or simply to improve comfort. Alternatively, the base 20 may be made flexible to conform to the contour of the user's skin 22, thus achieving the desired light-shielding and / or comfort objectives.

[0034] Several forms of lashing 28 are configured to secure the housing to the user's skin 22 over a target area. Of course, lashing 28 can take many different forms to suit the desired function and / or style. In the examples in Figures 2A and 3-7, lashing 28 includes a type of wristband used to secure watches and bracelets. The example in Figure 2C shows lashing in the form of adhesive. In Figures 2B and 2C, lashing is equipped with a spring clamp. In fact, there are many options for lashing, largely based on the intended application. Just looking at the current fastening elements used for clothing worn on or around the head, neck, shoulders, arms, hands, chest, waist, legs, ankles, and feet can provide inspiration for lashing configurations suitable for use with wearable devices 10. It should also be noted that lashing 28 can also include additional elements or sensors for user interfaces, communication, power, measurement, light sources, etc.

[0035] The components contained within the internal cavity of the housing are described here primarily with reference to Figures 4 and 5. A power supply 30 is located within or on the housing, within or on the lashing 28, or in several suitable locations operably associated with the wearable device 10. While the power supply 30 is shown located within the internal cavity, it may be located in any suitable location operably associated with the wearable device 10. The power supply 30 may comprise any suitable electrical energy storage and supply device, including but not limited to batteries, betavoltaic power supplies, supercapacitors, fuel cells, radio power supplies, solar cells, energy harvesters, etc. The power supply 30 is operably connected to the circuit board 32, or otherwise integrated into the operating system that powers and controls the electronics (electronic circuits) of the wearable device 10.

[0036] The light source 34 is configured to emit light through the port 24. The light source 34 is preferably located in an internal cavity, although embodiments are conceivable in which only the light generated by the light source 34 travels through the internal cavity. A light source 34 useful for the present invention must be able to generate optical photons of relevant characteristics when interacting with subcutaneous biological tissue. That is, when probed, the light source 34 can activate multiple responses from the tissue, including but not limited to Raman scattering, infrared emission, fluorescence, and phosphorescence. Generally, a light source suitable for use in connection with a wearable device generates light in a spectral band of about 200 nm to 1500 nm. In one embodiment, the light source 34 is configured to generate monochromatic light. In another assumed embodiment, the light source 34 is configured to generate broadband light. Thus, a light source 34 suitable for the wearable device 10 can be selected from the group consisting of light-emitting diodes, diode lasers, quantum cascade lasers, continuous lasers, plasma sources, hollow cathode sources, and xenon lamps. Other types of suitable light sources 34 are also possible, and within the scope of the present invention, in particular, light sources capable of generating light in a spectral band of about 200 nm to 1500 nm are also possible. Another type of suitable light source 34 may be a monochromatic light source that excites a reference mode of vibration of the target molecule, for example, a monochromatic light source having a specific frequency in the near-ultraviolet-visible-infrared (UV-VIS-IR) spectral band (200 nm to 1500 nm) selected to resonate or not resonate with the selected excitation mode of the target.

[0037] In some applications, it may be desirable to configure the light source 34 to be modulable with a duty cycle capable of counting optical photons. Optical photons can be counted over the duty cycle when the light source is turned on, or alternatively, when the light source is turned off (to enable background measurement), optical photons can be counted over the duty cycle. The measurement signal is determined as the difference of optical photons counted when the light source is on and off. In particular, the light source 34 of the present invention can be distinguished from those described in US20200107756A1, in which light is generated internally by a radiotherapy beam within the tissue being probed.

[0038] Although the light source 34 is depicted as a single photoproduct in the diagram, it will be understood that the light source 34 may instead comprise multiple separate light sources 34 that sequentially or simultaneously excite distinct quantized excitation modes of interest, including but not limited to electronic, vibrational, and resonant vivielectric or non-resonant vivielectric modes. In particular, measurements performed in resonant modes (i.e., when the excitation source is tuned to a specific frequency response of the sample system) may be useful for increasing the consideration of the signal-to-noise ratio.

[0039] The wearable device 10 further includes an array of multiple photodetectors 36 arranged within the housing. Each photodetector 36 has a separate channel (λ) configured to detect photons entering the internal cavity through the port 24. n The array comprises at least two photodetectors 36 (e.g., λ1 and λ2). Preferably, at least one photodetector is used to generate a reference signal, and the multiple photodetectors 36 are used, if necessary, to create separate, i.e., distinct channels for detecting multiple spectral lines to capture changes in tissue optical parameters.

[0040] In the illustrated example, six photodetectors 36 (λ1-λ6) are strategically arranged in a circular or annular pattern within an internal cavity around the light source 34. Other arrangements of the photodetectors 36 are certainly possible and may be deemed beneficial by the designer. The photodetectors 36 may be any suitable type selected from the group consisting of silicon photomultiplier tubes (SiPMs), photodiodes, avalanche photodiodes, Schottky photodiodes, photomultiplier tubes (PMTs), micro-PMTs, CCDs, CMOS sensors, InGaAs sensors, avalanche photodiode imaging arrays, Fabry-Perot etalons, and prisms.

[0041] At least one optical filter is associated with each photodetector 36. The optical filter restricts the light reaching the associated photodetector 36 to a specific wavelength (lambda) and filters out Rayleigh scattered light. While it is conceivable that the optical filter could be directly integrated with its associated photodetector 36, in the illustrated example, the optical filter is shown as separate from the photodetector 36. Similarly, the optical filter could be a single element, but in Figures 5 to 7 it is shown as a first filter 38 and a second filter 40. The first filter 38 comprises a narrowband filter at a specific lambda, while filter 40 comprises a spike filter that filters out Rayleigh scattered light. Thus, in the embodiment shown in the figures, each channel (λ) n Two filters 38 and 40 are used for the light. The first filter 38 allows only a specific selected wavelength (λ) to pass through. The second filter 40 blocks the remaining light source frequencies. Generally, the order in which the light passes through filters 38 and 40 is not important, and the placement relative to the relevant photodetector 36 is left to the designer's choice. Also, as mentioned above, the second filter 40 for filtering out Rayleigh scattered light can be considered a common filter that is useful for all photodetectors 36, such as at port 24 or some other convenient common location.

[0042] Optical filters suitable for use with the wearable device 10 can be selected from the group including, in essence, bandpass, multibandpass, notch, edgepass, spike, Rayleigh scattering rejection, diffraction grating, Fabry-Perot interferometer, MEM-based interferometry, dye, and nanophotonic types, and combinations thereof. Optionally, a light-shielding coating (or epoxy, or other suitable coating or surface treatment) may be applied to the side edges of one or both filters 38, 40 according to known techniques. Also, as described above, one or both optical filters may be integrated with their respective photodetectors 36 or separated from the photodetectors 36.

[0043] Preferably, the light shield 42 is positioned within an internal cavity between the light source 34 and the multiple photodetectors 36. The purpose of the light shield 42 is to prevent, or at least reduce, direct light from the light source 34 or light reflected from the surface of the skin 22 from reaching any of the photodetectors 36. Naturally, the light shield 42 can take many different forms. In the illustrated example, a light shield 42 surrounding the light source 34 is shown. However, alternative forms include one or more light shields surrounding the photodetectors 36 or otherwise partitioning the light source 34. Given the potential for design variations, the light shield 42 shown in Figure 5 is a substantially cylindrical tubular structure extending substantially from the underside of the cover 14 toward a terminal adjacent to the port 24. The ends of the light shield 42 are preferably positioned close to the surface of the skin 22 to maximize the light-shielding function. Optionally, the distal ends of the light shield 42 can be made conformable or extendable to provide a more perfect light-shielding seal to the skin 22. Figures 5 and 6 show dashed lines illustrating a simple accordion-like member 44 that is biased downward and, when the wearable device 10 is pushed into place, contacts and compresses the skin 22. Alternatively, a flexible lip seal or stretchable member can be attached to the end of the light shield 42. Many alternative configurations are possible if a more perfect light-shielding seal of the light shield to the skin 22 is desired. Furthermore, if a window glass or film covers the port 24, the accordion member 44 or extendable end can be appropriately adjusted to allow passage without sacrificing functionality.

[0044] The aforementioned circuit board 32 preferably includes a suitable data acquisition electronic module capable of counting single photons over a predetermined sampling time. The sampling time may vary depending on the application. In some cases, the allocated sampling time may be in the range of 0 to 1000 ms. In other applications, an allocated sampling time in the range of 1 to 10 seconds may be sufficient. Furthermore, the data acquisition electronic module preferably includes a scalar (counter) or other function that digitally records the intensity of a specific Raman ray of interest resulting from the excitation of a reference mode of a particular quantized molecular vibration. The reference mode of the quantized vibration may include an electronic mode, an ultrasonic mode, an acoustic mode, and / or a vibratory mode.

[0045] Now that we have described the basic physical components of the wearable device 10 above, the operation of the system can be understood in conjunction with Figure 7. Generally, as background, when light is scattered from a molecule or crystal, most photons are scattered elastically (i.e., their frequency does not change). This is called Rayleigh scattering. However, at optical frequencies that are different from and usually lower than the frequency of the incident photon, only a small portion of the light (10 7 Approximately one of the photons is scattered. This inelastic process is known as Raman scattering and can be caused by molecular vibration, rotation, or excitation of electron energy. Vibrational excitations appear in the spectrum of scattered light as weak Raman-shifted sidebands adjacent to the Rayleigh peak. Quantum considerations lead to upshifted (higher wavenumber) and downshifted (lower wavenumber) sidebands. The downshifted peak, called the Stokes-Raman spectrum, is generally stronger than the upshifted peak, so measurements can be limited to the downshifted case. The Raman shift Δν from a particular vibrational mode is given (in wavenumber) by the following equation. TIFF0007859682000001.tif23120

[0046] Here, the subscripts i and s refer to the incident photon and the scattered photon, respectively. This equation is essentially a description of energy conservation and is closely related to the present invention.

[0047] In Figure 7, an exemplary wearable device 10 is fixed in place against the user's skin 22 over the area of ​​interest. Ports 24 of the wearable device 10 are positioned to allow the device 10 to investigate subcutaneous molecules 46 in a non-invasive, real-time, and continuous manner to diagnose the condition of the area of ​​interest. As mentioned above, many different conditions of the area of ​​interest are conceivable. One primary example can be captured as blood glucose monitoring, but this is by no means the only possible application of the wearable device 10. Nevertheless, blood glucose monitoring serves as a good example to illustrate the operating characteristics of the present invention. Blood glucose monitoring may be abbreviated as CGM in relation to continuous glucose monitoring.

[0048] To obtain a signal specific to the characteristics of the target subcutaneous molecule (e.g., glucose concentration in the blood), the Raman detector (λs) is tuned to satisfy the above equation for a specific vibrational frequency of the target molecule (e.g., glucose molecule). This tuning step can be achieved in various ways. In one example, a first narrowband bandpass optical filter 38 (spectral width approximately 10 nm) is modified to a high-gain filter (approximately 10 nm). 7 It can be directly placed on the SiPM photodetector chip 36. For this purpose, an interference filter using an etalon (narrowband filter) selected to transmit (pass through) only light of this wavelength is useful. By using the second optical notch filter 40 in combination with the first narrowband optical filter 38, much stronger Rayleigh scattered light can be excluded from the Raman spectrum, ensuring that the detected signal is from Raman scattered light only.

[0049] This sophisticated multispectral detection approach allows for the direct and accurate reading of the properties of a target subcutaneous molecule, such as blood glucose concentration, by measuring the integrated intensity of selected Raman active lines relative to nearby Raman peaks associated with tissue matrices such as water or hemoglobin, which can serve as a reference (control). This technique simultaneously overcomes two barriers to miniaturization that have hindered the application of Raman spectroscopy to real-time glucose (or other target molecule) monitoring: (1) Each of the spectral bandpass filters 38 is precisely tuned to a specific Raman peak, thus eliminating the need for a large scanning spectrometer; and (2) The high-gain SiPM photodetector 36 is sensitive to weak Raman signals and is extremely compact (<2 mm). 2 The design incorporates, if necessary, a multi-channel array that allows for several separate channels for detecting a reference signal and multiple Raman rays to capture changes in the optical parameters of the tissue (e.g., skin color, skin irritation, etc.). Possible arrays include 2x2 and the 6-channel array proposed in Figure 4. Of course, other array configurations are also possible and may be preferred over the aforementioned 2x2 and 6-channel arrays depending on the application.

[0050] The present invention may be understood to highlight two aspects of Raman spectroscopy that are particularly relevant to exemplary continuous geometry (CGM) applications.

[0051] First, this measurement is related to the complex biological environment in which it is performed, where glucose molecules are immersed in blood / interstitial fluid and have numerous vibrational modes, primarily related to hemoglobin and water, within the frequency band of interest. However, it should be noted that, by quantum selection rules, Raman-active processes are limited to those in which the polarizability of the molecule changes due to the symmetry of a particular vibrational mode. This significantly reduces the complexity of the Raman spectrum to a few strong lines and well-separated lines. For example, the most interesting Raman line here is around 1125 cm⁻¹, associated with the "respiratory mode" of the glucose ring. -1 This is the Raman line in Figure 1. Figure 1 is approximately 572 cm.-1 、 796 cm -1 、 1060 cm -1 and 1360 cm -1 show other Raman lines of the target. The strong Raman line at 1549 cm -1 of hemoglobin (not visible in Figure 1) is also about 436 cm -1 、 456 cm -1 、 527 cm -1 、 855 cm -1 、 912 cm -1 、 1060 cm -1 、 1366 cm -1 and 1456 cm -1 are interesting as possible criteria for calibrating blood glucose concentration, along with the Raman lines at these wavelengths. The idea of this ratiometric approach is to use a reference to remove any fluctuations in glucose Raman intensity caused by changes other than the glucose concentration in the blood. For example, when the wearable device 10 moves relative to the measurement area of the subject, the overall intensity may vary, but the differential signal relative to the reference cannot vary. Also, effective rushing 28 helps reduce or eliminate intensity fluctuations caused by relative movement between the wearable device 10 and the subcutaneous molecule 46 being investigated. By coordinating sampling with pulsed measurements (i.e., sampling during the pulse and sampling between pulses), the blood glucose measurement value can also be explained from the interstitial fluid glucose measurement.

[0052] A second important physiological aspect is that the subcutaneous tissue and dermis through which the incident light must pass before reaching the molecule 46 are opaque and have a rather short absorption length (a few mm). However, by a judicious choice of the light source 34 (e.g., a laser or LED, excitation wavelength, λ i etc.), the absorption of the incident light and scattered light can be kept to a fairly minimum. One effective approach is to use red light or near-infrared (NIR) light (λ iThe method involves using a wavelength of ≥660 nm. A further advantage is the availability of commercially available SiPM-type photodetectors 36 optimized for red light (R-series). Thus, the inherent flexibility and selectivity of Raman spectroscopy can be effectively utilized by the wearable device 10 to advance the design of non-invasive, wearable, real-time continuous diagnostic spectrometers and methods for CGM and many other applications. Therefore, the wearable device 10 of the present invention can be configured as a continuous glucose monitoring (CGM) to dramatically improve the lives of diabetic patients and enhance their health outcomes.

[0053] The high sensitivity and high gain of the described photodetector 36 technology are one of the key enablers of this innovation. The wearable device 10 utilizes weak inelastic scattered light transmitted through tissue to enable spectral measurement of relevant molecular biomarkers. Inelastic scattered light is strongly absorbed and scattered by human tissue, leaving only a small number of measurable photons that carry information about the target, and only a subset within a specific wavelength is useful for molecular measurement. Similarly, the energy-efficient light source 34 generates a very limited number of spectroscopically unique signals, thus requiring a similarly sensitive and sophisticated technical approach. In CGM applications, the wearable device 10 employs Raman spectroscopy, targeting a small fraction of scattered photons that convey the "fingerprint" and concentration of the target molecule 46 in blood or interstitial fluid. The wearable device 10 and its method of use uniquely overcome concerns regarding the intensity of weak signals from the target, which are largely mitigated by the high sensitivity of the SiPM detector. In the applications described herein (biomolecules probed by Raman scattering), only specific components of the Raman scattered light, i.e., components with relevant wavelengths necessary to measure the property of interest (e.g., blood glucose), are targeted. This eliminates the need for physically large and expensive high-performance spectrometers, enabling compact, wearable, and low-cost devices.

[0054] The principles of this invention enable a truly non-invasive Raman probe in a wearable form factor. A unique new architecture allows the wearable device 10 to tightly integrate a highly sensitive multi-channel photodetector 36 incorporating a high-gain photodetector 36 coupled with a narrowband first optical filter 38 to enable measurement of specific Raman peaks of target molecules 46. Furthermore, miniaturization of the digital electronics (electronic circuits) required for data acquisition, signal processing, and real-time analysis has improved the signal-to-noise ratio in other Raman scattering applications, but significantly improved in the photon counting mode. Acquisition of multispectral data is crucial for quantitative analysis using multivariate calibration models, which have been shown to help achieve high quantitative accuracy based on measurements of multiple Raman lines, not just a single peak. The wearable device 10 brings a powerful and established Raman spectroscopy technique to practical application as a wearable biosensor technology. This invention serves not only as a truly non-invasive real-time CGM but also as a platform for measuring many other important physiological biomarkers.

[0055] In a specific example of CGM, the wearable device 10 demonstrates a non-invasive approach to measuring blood glucose concentration based on Raman spectroscopy. In one embodiment, monochromatic light (e.g., from a laser diode 34) is directed into the subcutaneous tissue. The generated light is scattered from the interstitial fluid and blood vessels. Some of this light also interacts with glucose molecules 46 in this matrix, exciting vibrational modes specific to glucose. The frequencies of these characteristic modes (e.g., stretching and bending of bonds) are imprinted on the Raman scattered light and, after leaving the body, are analyzed by an array of small narrowband spectrometers 36. In this way, Raman spectroscopy "fingerprints" molecular bonds in chemistry and biology, providing a reliable quantitative probe for blood glucose concentration.

[0056] The above invention is described in accordance with relevant legal standards, and therefore the description is illustrative and not limiting in nature. Variations and modifications to the disclosed embodiments will be obvious to those skilled in the art and fall within the scope of the invention.

Claims

1. A non-invasive, wearable diagnostic spectrometer device adapted for direct contact with the user's skin, A housing having side walls surrounding an internal cavity, a cover positioned above the side walls and the internal cavity, and a base positioned below the side walls and at least partially enclosing the internal cavity, wherein the base has a port inside configured to allow photons of light to pass through, A lashing extending from the housing and configured to make the diagnostic spectrometer wearable by directly fixing the internal cavity of the housing to the user's skin in the area of ​​interest, A light source configured to emit photons of light through a port that will probe a physiologically targeted molecular biomarker in subcutaneous tissue, and to return photons containing a mixture of Rayleigh scattered and Raman scattered light through the port to the internal cavity, wherein the light source includes a plurality of separate light-emitting diodes, and the plurality of separate light-emitting diodes are configured to generate monochromatic light having a frequency in a spectral band of 200 nm to 1500 nm that sequentially or simultaneously excites a separate quantized excitation mode of target, An array of photodetectors disposed within the internal cavity of the housing, wherein each photodetector comprises a separate channel (λn) configured to detect Raman scattered photons entering the internal cavity through the port, and the photodetectors are silicon photomultiplier tubes (SiPMs), At least one optical filter associated with each of the SiPM photodetectors in the internal cavity, wherein the optical filter is operably positioned between the associated SiPM photodetector and the port, and each optical filter restricts the passage of light reaching the associated SiPM photodetector to a specific wavelength and eliminates Rayleigh scattered light, A data acquisition electronic module fixed to the housing and operably associated with the array of SiPM photodetectors, the data acquisition electronic module configured to count single photons of Raman scattered light reaching each of the SiPM photodetectors over a predetermined sampling time, A device equipped with the following features.

2. The apparatus according to claim 1, wherein the array of SiPM photodetectors comprises at least one SiPM photodetector that generates a reference signal and a plurality of SiPM photodetectors that detect a plurality of separate Raman rays.

3. The apparatus according to claim 1, further comprising a light-shielding body disposed in the internal cavity between the light source and the plurality of SiPM photodetectors.

4. The apparatus according to claim 3, wherein the light shield extends substantially from the light source toward a terminal adjacent to the port.

5. The apparatus according to claim 3, wherein the light-shielding body is substantially tubular and surrounds the light source.

6. The apparatus according to claim 1, wherein the predetermined sampling time is 1000 ms or less.

7. The apparatus according to claim 1, wherein the predetermined sampling time is in the range of 1 to 10 seconds.

8. The apparatus according to claim 1, wherein the data acquisition electronic module includes a counting device for digitally measuring the integrated intensity of Raman scattered light resulting from the excitation of a specific quantized vibration reference mode, the quantized vibration reference mode includes at least one of an electronic mode, an optical vibration mode, an acoustic vibration mode, an ultrasonic mode, and a vibratory mode.

9. The apparatus according to claim 1, wherein the optical filter is essentially selected from the group consisting of bandpass, multibandpass, notch, and edgepass filters.

10. The apparatus according to claim 1, wherein each optical filter comprises a first filter configured to remove Rayleigh scattered light and a second filter configured to restrict the passage of light reaching the associated SiPM photodetector to a specific wavelength.

11. The apparatus according to claim 1, wherein the light source is configured to generate light capable of activating a fluorescence or phosphorescence response from a biological tissue to be probed.

12. The apparatus according to claim 1, wherein the light source is configured to generate tunable monochromatic light that can perform measurements in resonant mode.

13. A non-invasive and wearable diagnostic spectrometer device according to claims 1 to 12, When the diagnostic spectrometer is fixed to the user's skin, the diagnostic spectrometer will The steps include: emitting light directly onto the user's skin from a light source within the internal cavity via a port; A step of investigating at least one subcutaneous molecule beneath the user's skin using the light, the investigating step of generating optical photons of Rayleigh scattered light mixed with Raman scattered light that are re-incidentated into the internal cavity through the port, The steps include: shielding the plurality of SiPM photodetectors in the internal cavity from light emitted by the light source, but not shielding them from the Raman scattered light that re-incidentates into the internal cavity through the port; The steps include removing Rayleigh scattered light from the photons that re-incidentate the internal cavity through the port, The steps include limiting the Raman scattered light reaching each SiPM photodetector to a specific wavelength related to the Raman active line, The steps include counting single photons received by each SiPM photodetector over a predetermined sampling time and measuring the integrated intensity of the selected Raman active line, A device configured to perform the following actions.

14. The subcutaneous molecules to be investigated under the user's skin are essentially selected from the group consisting of hemoglobin and glucose, and the selected Raman active lines are essentially about 436 cm －１ , about 456 cm －１ , about 527 cm －１ , about 572 cm －１ , about 796 cm －１ , about 855 cm －１ , about 912 cm －１ , about 1060 cm －１ , about 1125 cm －１ , about 1360 cm －１ , about 1366 cm －１ , about 1456 cm －１ , and about 1549 cm －１ The apparatus according to claim 13, selected from the group consisting of.

15. The apparatus according to claim 13, wherein the step of emitting light from a light source includes generating light having a frequency in a spectral band of 200 nm to 1500 nm.

16. The apparatus according to claim 13, wherein the step of emitting light from a light source essentially includes simultaneously exciting separate quantized excitation modes selected from the group consisting of electronic, vibrational, and resonant vibratory and non-resonant vibratory.

17. The apparatus according to claim 13, further comprising the step of transmitting data notified by the measured integrated intensity of the selected Raman active line to a remote computing device via a secure communication connection.

18. The apparatus according to claim 13, further comprising the step of transmitting data notified by the measured integrated intensity of the selected Raman active ray to a remote controller of the treatment device.

19. The apparatus according to claim 13, further comprising the step of generating an alarm signal in response to the measured integrated intensity of the selected Raman active line.

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