Patent Publication Number: US-10772541-B2

Title: System and method for noninvasive analysis of subcutaneous tissue

Description:
CROSS REFERENCE TO RELATED APPLICATION 
     This application is a continuation-in-part of application Ser. No. 15/627,470, filed Jun. 20, 2017, which is a continuation-in-part of application Ser. No. 14/465,311, filed Aug. 21, 2014, both of which are incorporated by reference herein in its entirety. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates to noninvasive analysis. More particularly, the present invention relates to a system and method for noninvasive analysis of subcutaneous tissue for determination of substance concentration in the blood stream and/or in other subcutaneous tissue. 
     BACKGROUND OF THE INVENTION 
     Various medical conditions are characterized by accumulation of liquids under or behind the skin surface. Such conditions may include otitis media, pressure ulcers, or other types of deep tissue injury under intact skin. 
     Medications or other substances may be introduced or delivered into the bloodstream. For example, a substance may be delivered orally to a patient, or may be injected into tissue or via an intravenous infusion. In some cases, it is important to monitor the concentration or amount of the substance in a patient&#39;s blood or tissue. Monitoring the concentration or the amount may include drawing a blood or tissue sample from the patient. 
     In order to provide respiratory and cardiovascular support for patients undergoing medical procedures, as well as monitoring the vital signs of the patient, anesthesiologists need to properly monitor the anesthetics. The degree of anesthesia reflects the degree of blockade of sensory, reflex, mental, and/or motor functions, which can be achieved by using inhalational agents and/or intravenous anesthetics. Current methods cannot provide precise determination of the degree of anesthesia and/or the concentration of anesthetic agents in the blood stream and/or precisely determine the drug quantity required to provide a sufficient “degree of anesthesia”. 
     Propofol is a short-acting intravenously administered anesthetic that is most commonly used for induction and maintenance of general anesthesia and for procedural sedation. For example, it can be used in a regimen that provides hypnosis, analgesia and muscle relaxation. Propofol is usually administered as a constant intravenous infusion in order to deliver and maintain a specific plasma concentration. It is therefore desirable to evaluate plasma concentrations in real time to accurately maintain anesthetic efficacy. Currently there is no clinically useful method for measuring blood Propofol concentrations on line and in real time. 
     SUMMARY OF THE INVENTION 
     There is thus provided, in accordance with some embodiments of the present invention, a method for noninvasive analysis of tissue, the method including: irradiating a surface of the tissue with short wave infrared (SWIR) radiation in a first spectral band that is strongly absorbed by water, and with SWIR radiation in a second spectral band such that an interaction of the radiation in both spectral bands with a component of the tissue other than water is substantially identical; measuring an intensity of the radiation that emerges from the tissue in each of the spectral bands; calculating a relative absorption by the tissue of radiation in one of spectral bands relative to absorption by the tissue of radiation in the other of the spectral bands; and determining a state of the tissue in accordance with the calculated relative absorption. 
     Furthermore, in accordance with some embodiments of the present invention, the first spectral band is in the wavelength range of 1400 nm to 1500 nm. 
     Furthermore, in accordance with some embodiments of the present invention, the second spectral band is in the wavelength range of 1000 nm to 1350 nm or 1500 nm to 2100 nm. 
     Furthermore, in accordance with some embodiments of the present invention, a gap in wavelength between the first and second spectral bands is less than 200 nm. 
     Furthermore, in accordance with some embodiments of the present invention, measuring the intensity includes measuring the intensity of the radiation that is transmitted across the tissue (e.g., passing through the tissue and not reflected back towards the radiation source). 
     Furthermore, in accordance with some embodiments of the present invention, the tissue includes tissue of a finger or an ear. 
     Furthermore, in accordance with some embodiments of the present invention, measuring the intensity includes measuring the intensity of the radiation that is reflected by the tissue. 
     Furthermore, in accordance with some embodiments of the present invention, measuring the intensity includes measuring the intensity of the radiation that emerges from the tissue at a plurality of lateral distances from a location of the irradiating of the tissue. 
     Furthermore, in accordance with some embodiments of the present invention, the state of the tissue includes a concentration of a substance in blood. 
     Furthermore, in accordance with some embodiments of the present invention, the substance is an introduced substance. 
     There is further provided, in accordance with some embodiments of the present invention, a system for noninvasive analysis of tissue, the system including: at least one source of infrared radiation to irradiate the tissue, the infrared radiation including SWIR radiation in a first spectral band that is strongly absorbed by water, and including radiation in a second spectral band such that an interaction of the radiation in both spectral bands with a component of the tissue other than water is substantially identical; at least one radiation detector to measure an intensity of radiation in each of the two spectral bands that emerges from the tissue (e.g., reflected from the tissue or passing through the tissue, to come out at another side), and a processor that is configured to calculate a relative absorption by the tissue of radiation in one of spectral bands relative to absorption by the tissue of radiation in the other of the spectral bands and determine a state of the tissue in accordance with the calculated relative absorption. 
     Furthermore, in accordance with some embodiments of the present invention, wherein the at least one radiation detector is configured to measure the intensity of the radiation that emerges from a surface of the tissue that is irradiated by the at least one radiation source. 
     Furthermore, in accordance with some embodiments of the present invention, the at least one radiation detector is configured to measure the intensity of the radiation that emerges from the surface of the tissue at a plurality of lateral distances from the at least one radiation source. 
     Furthermore, in accordance with some embodiments of the present invention, the at least one radiation detector includes a plurality of radiation detectors separated by different lateral distances from the at least one radiation source. 
     Furthermore, in accordance with some embodiments of the present invention, the at least one radiation detector is configured to measure the radiation that emerges from a surface of the tissue that is substantially opposite a surface of the tissue that is irradiated by the at least one radiation source. 
     Furthermore, in accordance with some embodiments of the present invention, the system includes a removable cover for placement over an aperture of the at least one radiation source or of the at least one radiation detector. 
     Furthermore, in accordance with some embodiments of the present invention, the at least one radiation source includes two radiation sources, one of the sources being configured to emit radiation in the first spectral band and the other being configured to emit radiation in the second spectral band. 
     Furthermore, in accordance with some embodiments of the present invention, the at least one radiation detector includes two radiation detectors, one of the detectors being configured to measure an intensity of radiation in the first spectral band and the other being configured to measure an intensity of radiation in the second spectral band. 
     Furthermore, in accordance with some embodiments of the present invention, the system includes a dispersive element to separate spectral components of the infrared radiation and a micro-mirror array, the micro-mirror array being operable to direct a selected spectral component of the infrared radiation to the tissue or to the at least one radiation detector. 
     Furthermore, in accordance with some embodiments of the present invention, the first spectral band is in the wavelength range of 1400 nm to 1500 nm. 
     Furthermore, in accordance with some embodiments of the present invention, the second spectral band is in the wavelength range of 1000 nm to 1350 nm or 1500 nm to 2100 nm. 
     There is further provided, in accordance with some embodiments of the present invention, a method for determining a state of tissue, the method including: irradiating a surface of the tissue with SWIR radiation in a first spectral band in the wavelength range 1300 nm to 1430 nm, and with SWIR radiation in a second spectral band such that an interaction of the radiation in both spectral bands with a component of the tissue other than water is substantially identical; measuring an intensity of the radiation that emerges from the tissue in each of the spectral bands; and calculating an absorption by the tissue of radiation in the two spectral bands, the absorption being indicative of the state of the tissue. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       In order for the present invention to be better understood and for its practical applications to be appreciated, the following Figures are provided and referenced hereafter. It should be noted that the Figures are given as examples only and in no way limit the scope of the invention. Like components are denoted by like reference numerals. 
         FIG. 1A  is a schematic drawing of a system for noninvasive analysis of subcutaneous liquids based on reflection of infrared radiation, in accordance with an embodiment of the present invention. 
         FIG. 1B  is a schematic illustration of a plurality of component apertures of an optical head of the system shown in  FIG. 1A . 
         FIG. 2A  is a schematic drawing of a measurement unit of a system for noninvasive analysis of subcutaneous liquids based on transmission of infrared radiation, in accordance with an embodiment of the present invention. 
         FIG. 2B  schematically illustrates attachment of the measurement unit of  FIG. 2A  to a finger. 
         FIG. 2C  schematically illustrates attachment of the measurement unit of  FIG. 2A  to an ear. 
         FIG. 3  shows an example of a graph of spectral reflectance. 
         FIG. 4  shows a graph of an example of a relationship of absorbance to concentration of a substance. 
         FIG. 5  is a flowchart depicting a method for noninvasive analysis of subcutaneous liquids, in accordance with an embodiment of the present invention. 
         FIG. 6A  schematically illustrates a system for measurement of reflection at a distance from a radiation source, in accordance with an embodiment of the present invention. 
         FIG. 6B  schematically illustrates an arrangement of a measurement head for concurrent measurement of radiation that emerges from a surface at different lateral distances from the radiation source, in accordance with an embodiment of the present invention. 
         FIG. 7  schematically illustrates paths of incident radiation to different detector locations, in accordance with an embodiment of the present invention. 
         FIG. 8  is a block diagram showing elements of a measurement unit, in accordance with an embodiment of the present invention. 
         FIG. 9  shows an example graph of detection of injected substance over time, in accordance with an embodiment of the present invention. 
         FIG. 10  shows an example of a transmission measurement unit, in accordance with an embodiment of the present invention. 
         FIG. 11A  shows a top view of a reflectance measurement unit, in accordance with an embodiment of the present invention. 
         FIG. 11B  shows a side view of the reflectance measurement unit, in accordance with an embodiment of the present invention. 
         FIG. 11C  shows a cross-section perspective view of the reflectance measurement unit, in accordance with an embodiment of the present invention. 
         FIG. 12A  shows a portion of flowchart for a method of noninvasive analysis of tissue, in accordance with an embodiment of the present invention. 
         FIG. 12B  shows the second portion of the flowchart of  FIG. 12A , in accordance with an embodiment of the present invention. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     In the following detailed description, numerous specific details are set forth in order to provide a thorough understanding of the invention. However, it will be understood by those of ordinary skill in the art that the invention may be practiced without these specific details. In other instances, well-known methods, procedures, components, modules, units and/or circuits have not been described in detail so as not to obscure the invention. 
     Although embodiments of the invention are not limited in this regard, discussions utilizing terms such as, for example, “processing,” “computing,” “calculating,” “determining,” “establishing”, “analyzing”, “checking”, or the like, may refer to operation(s) and/or process(es) of a computer, a computing platform, a computing system, or other electronic computing device, that manipulates and/or transforms data represented as physical (e.g., electronic) quantities within the computer&#39;s registers and/or memories into other data similarly represented as physical quantities within the computer&#39;s registers and/or memories or other information non-transitory storage medium (e.g., a memory) that may store instructions to perform operations and/or processes. Although embodiments of the invention are not limited in this regard, the terms “plurality” and “a plurality” as used herein may include, for example, “multiple” or “two or more”. The terms “plurality” or “a plurality” may be used throughout the specification to describe two or more components, devices, elements, units, parameters, or the like. Unless explicitly stated, the method embodiments described herein are not constrained to a particular order or sequence. Additionally, some of the described method embodiments or elements thereof can occur or be performed simultaneously, at the same point in time, or concurrently. Unless otherwise indicated, us of the conjunction “or” as used herein is to be understood as inclusive (any or all of the stated options). 
     Some embodiments of the invention may include an article such as a computer or processor readable medium, or a computer or processor non-transitory storage medium, such as for example a memory, a disk drive, or a USB flash memory, encoding, including or storing instructions, e.g., computer-executable instructions, which when executed by a processor or controller, carry out methods disclosed herein. 
     In accordance with an embodiment of the present invention, reflection of radiation or transmission of radiation by tissue in a region of the body is measured. In some embodiments, reflected radiation may refer to radiation that emerges from the tissue via a surface through which the tissue was irradiated or illuminated. Thus, reflected radiation may result from one or more of radiation that is reflected from an interface between dissimilar media and radiation that is scattered in the backward direction. For example, a measurement unit may be placed on or near a surface of the tissue, e.g., a skin surface of the patient&#39;s body with superficial blood vessels, on oral mucosa of inner lip, an eardrum, or another surface. The measurement unit may include one or more radiation sources and one or more radiation detectors. The measurement unit may be configured to measure reflection. For example, the radiation sources may be configured to irradiate or illuminate a tissue surface (e.g., a skin surface of a patient). When irradiating the tissue surface, the radiation penetrates into the tissue. The radiation detectors may be aimed at the irradiated surface so as to detect radiation that is reflected or backscattered by the tissue that is covered by the surface. In other cases, the radiation sources and detectors may be configured to measure transmission of the radiation through the tissue. For example, measurement unit may be configured such that radiation that is emitted by the radiation sources is directed toward the radiation detectors. Thus, when tissue (e.g., a part of the patient&#39;s body such as an ear, finger, toe, fold of skin, or other part of the body) is placed in the optical path from the radiation source to the corresponding radiation detector, transmission of the radiation through the tissue may be measured. 
     The reflection or transmission measurement may be indicative of a state of the tissue. The tissue may include subcutaneous liquids such as blood or other fluids. In some embodiments, the term “subcutaneous”, may refer to a depth within tissue, which may or may not be covered with skin. For example, subcutaneous liquid may be within or behind a membrane, such as within or behind the eardrum, or within lung tissue and/or such as oral mucosa of inner lips or within superficial blood vessel. The state of the tissue may be indicative of a medical condition in the patient. For example, a medical condition may include otitis media, early stages of pressure ulcers, or other types of deep tissue injury under intact skin in which liquids accumulate subcutaneously. A state of the tissue may include a concentration of a substance in the blood or other subcutaneous fluids. The substance (e.g., a medication, contrast agent, food component or supplement, or other administered product) may have been administered to the patient (e.g., orally, or via injection or infusion), may be a product of physiological processes on an administered substance, or may be produced by the patient&#39;s body. For example, the most common pressure ulcers are above the heel bone under the skin. The heel bone is covered with a thin fatty layer which is normally composed of triglycerides. When ischemia starts the triglycerides fatty tissue decomposes into glycerol and free fatty acids. According to some embodiments, a lookup table containing different concentrations of, for example, glycerol and free fatty acids may be created, and may further include data regarding such substances absorption (e.g., in different wavelengths). A processor, may determine state of the tissue above the heel bone (or at any other location) by comparing the optical readings from the examined tissue, to the absorption values in the lookup table, and based on the determined concentration of the different substances, determine the existence and degree or severity of DTI. 
     The reflection or transmission is measured in at least two spectral bands of the visible (VIS) and/or near infrared (NIR) and/or shortwave infrared (SWIR) spectral region of the electromagnetic spectrum. As used herein, the SWIR spectral region is used to include the wavelength range of about (e.g., ±10%) 1000 nm to about 2500 nm. The shorter wavelengths of this spectral region are sometimes referred to as near infrared (NIR). It should be noted that as used herein, the VIS-NIR-SWIR spectral region may include the wavelengths range of ˜350 nanometers (nm) to ˜2500 nm, and MIR may include the wavelengths range of ˜2500 nm to 15000 nm waveband. 
     A first spectral band may be in a portion of the SWIR spectral region where radiation is strongly, or more strongly, absorbed by water (e.g., in the wavelength band from about 1400 nm to about 1500 nm) as compared to adjacent or other bands. As used herein, radiation is considered to by strongly absorbed when absorption (e.g., as characterized by an absorption coefficient) is at least an order of magnitude (approximately 10 times or more) greater than in the comparison band. 
     A second spectral band is in an adjacent SWIR spectral band, e.g., in the wavelength band from about 1000 nm to about 1350 nm, or in the wavelength band from about 1500 nm to about 2100 nm. The second spectral band is sufficiently close to the first spectral band such that an interaction of radiation in both spectral bands (e.g., absorption or scattering) with tissue components other than water is substantially identical (e.g., having the same or very similar behavior interaction with the tissue). In some embodiments, substantially identical may refer to measurements at two wavebands for the change of measured backscattered power over sampling time with the change of 3 dB and/or the mean intensity “I” (or mean square of intensity) of √2&lt;I&gt;. In some embodiments, substantially identical may refer to measurements at two wavebands with a similarity test (e.g., t-test, Kruskal-Wallis test, etc.) with the p-value being above 0.05. In some embodiments, substantially identical may refer to measurements at two wavebands while assuming that for a given waveband all chromophores other than water may absorb and scatter the radiation in the same way, for instance the photon-tissue interaction may be constant for a given waveband. For example, absorption of radiation in the spectral band of about 1000 nm to about 1800 nm by such tissue components other than water such as hemoglobin, melanin, and other chromophores is approximately constant (with absorption coefficients in the range of about 0.1 cm −1  to about 1 cm −1 ). The scattering coefficient is about 5 cm −1  to about 15 cm −1 . Thus, radiation may penetrate as deep as 8 cm to 10 cm into tissue. 
     For example, a gap between the two bands may be no more than 200 nm. In some cases, the gap between the two bands may be no more than 100 nm. 
     For example, two or more of the radiation sources may be configured to emit radiation in different spectral bands of the SWIR spectral region. As another example, two or more of the radiation detectors may be configured to detect different radiation in different spectral bands of the SWIR spectral region and/or the visible spectral region (also referred to as VIS). Both the radiation sources and the radiation detectors may be limited to particular spectral bands. 
     For example, the source or detector may include a wavelength selection arrangement that incorporates a spectrally dispersive element (e.g., a grating, prism, or spectrally selective optical coating), focusing optics (e.g., lenses or mirrors), and a micro-mirror array that contains individually rotatable micro-mirrors (or rotatable in groups). Radiation originating from the radiation source (for irradiating the tissue) or from the tissue (e.g., after reflection or transmission) may be focused onto the dispersive element to spatially separate different spectral components (e.g., wavelength ranges or spectral bands) of the radiation. For example, different wavelengths of the radiation may be directed in different directions. Spectrally separated radiation from the dispersive element may be focused onto the micro-mirror array. For example, each spectral component may be incident on a different micro-mirror of the array. Therefore, each micro-mirror may be selectively rotated to direct a particular spectral component of the radiation either toward or away from a source aperture (to irradiate the tissue surface) or a detector (to be detected as transmitted or reflected radiation). 
     In some cases, either the spectral bands in which radiation is emitted by different radiation sources, or the spectral bands to which different radiation detectors are sensitive, may partially or completely overlap. In such a case, spectral separation may be effected by another of the components (e.g., source, detector, or optics). 
     In biological tissue, the absorption of radiation in a particular spectral band (e.g., in the SWIR range) may be determined the contributions of various substances or chromophores that absorb electromagnetic radiation of different wavelengths. Chromophores are functional groups of molecules that absorb light or electromagnetic radiation to various extents in a spectral band. Each chromophore is characterized by a particular characteristic absorption as a function of wavelength, and which may be used to identify the presence of that molecule. 
     In practice, only a few chromophores contribute to the absorption in the NIR-SWIR range from about 800 nm to about 2400 nm. In this region, major contributions to absorption in the body arise from the presence of (oxy- or deoxy-) hemoglobin, water, and fat and melanin. Other chromophores that may potentially contribute slightly (e.g., about 1%-3%) to the absorption include myoglobin, cytochrome, bilirubin, lipids and other substances. 
     In SWIR radiation with wavelength greater than about 1000 nm, the contribution of water to absorption is greater than that of hemoglobin, melanin and other chromophores. 
     For example, one of the spectral bands may include the SWIR band from about 1400 nm to about 1500 nm, which is strongly absorbed by water. Another of the spectral bands may include the SWIR band from about 1000 nm to about 1350 nm, or the SWIR band from about 1500 nm to about 2100 nm. Radiation in the spectral region from about 1350 nm to 1400 nm, where the absorption by water rapidly changes with wavelength (with a large slope in a curve of absorption versus wavelength), may not be used in some cases 
     In some cases, a rate of change of absorption of radiation as a function of the wavelength may be measured in the spectral region from about 1300 nm to about 1430 nm. In this spectral region, the absorption by water rapidly changes with wavelength. The high rate of change of absorption with wavelength (absolute rate of change for absorption in 1 mm of water being greater than 0.007 nm −1 ) may be exploited to detect the presence of water (or an amount or concentration of water) in the tissue. For example, a slope of a graph of absorption (or reflection or transmission) versus wavelength may be calculated. As an example, it may be noted that in  FIG. 3 , the slopes of normal tissue curve  72  and of disease tissue curve  74  (having different water content) noticeably differ from one another in the spectral range of 1300 nm to 1430 nm. 
     The measurement of the reflection or transmission may be performed in at least two different but adjacent spectral bands of the SWIR region of the electromagnetic spectrum. The results of the reflection or transmission measurement may be analyzed by comparison with previously measured or calibrated results. The comparison may be utilized to determine the state of the subcutaneous tissue. 
     For example, one of the spectral bands may be selected such that the reflection or transmission in that band is relatively unaffected by the presence or absence of the medical condition. The measured reflection or transmission in this spectral band may serve as a reference. For example, the reference measurement may set a baseline value or temporary calibration for measurements in the presence of conditions that are specific to the measurement. The conditions may be related to the patient&#39;s body (e.g., dimensions or other properties of a skin or tissue surface or body part on which the measurements are made). The conditions may be related to drift or transient variation in properties of the measurement unit that is used to perform the reflection or transmission measurements. 
     Measurements may be further calibrated by monitoring a dark signal (e.g., when a radiation detector is shielded from radiation) and an intensity of the radiation source (e.g., by providing a direct channel of radiation from the radiation source to the corresponding radiation detector). 
     Another of the spectral bands may be selected such that a reflection or transmission measurement in that band is sensitive to the state of the tissue. Thus, when adjusted in accordance with the various reference and calibration measurements, the measurement in that other spectral brand may be indicative of the state of the tissue. Thus, the state of the tissue may be detected noninvasively. 
     For example, one of the spectral bands may include wavelengths of SWIR electromagnetic radiation that are strongly (e.g., almost completely) absorbed by water. Another of the spectral bands may include SWIR electromagnetic radiation that is largely transmitted by water. In some cases, one of the spectral bands may include SWIR electromagnetic radiation whose absorption varies rapidly with wavelength (e.g., instead of, or in addition to, the spectral band in which radiation is strongly absorbed). 
       FIG. 1A  is a schematic drawing of a system for noninvasive analysis of subcutaneous liquids based on reflection of infrared radiation, in accordance with an embodiment of the present invention. 
     Liquid analysis system  10  includes reflection measurement unit  12 . Reflection measurement unit  12  includes infrared radiation source  14  and radiation (e.g., infrared) detector  16 . Unit aperture  19  on optical head  13  of reflection measurement unit  12  is configured to be placed on or near a tissue surface  22  of a patient so as to measure infrared radiation that is reflected from tissue that is covered by tissue surface  22 . Liquid analysis system  10  is configured to determine the state of subcutaneous tissue, such as the presence of absence of a subcutaneous medical condition in tissue that is covered by tissue surface  22 , or to measure the presence or absence of an administered substance in blood that flows via the tissue. 
     Reflection measurement unit  12  may be configured to be held and manipulated by a single hand of a user. For example, the user may include a healthcare professional, a caregiver, or the patient. Reflection measurement unit  12  may include an internal power source in the form of a battery or other self-contained power source, or may be connectable to an external power source. Reflection measurement unit  12 , optical head  13 , or both may have a generally or substantially cylindrical form, or may have another geometrical form or shape. Unit aperture  19  may thus have a substantially round or elliptic shape, or another shape. 
     Unit aperture  19  of optical head  13  may include one or more component apertures. 
       FIG. 1B  is a schematic illustration of a plurality of component apertures of an optical head of the system shown in  FIG. 1A . Unit aperture  19  includes a plurality of component apertures  40 . Each component apertures  40  may be configured to enable passage of radiation from a single component infrared source of infrared radiation source  14  or to a component detector of radiation (e.g., infrared) detector  16 . In the configuration shown, unit aperture  19  is round and component apertures  40  are arranged in a hexagonal pattern. Other configurations may include other shapes of unit aperture  19  or other arrangements of component apertures  40 . 
     Infrared radiation source  14  is configured to irradiate tissue surface  22  with SWIR radiation via unit aperture  19  of optical head  13 . Infrared radiation source  14  may include one or more separate component infrared sources. For example, two or more component infrared sources may each produce SWIR radiation in one or more spectral bands. For example, infrared radiation source  14  may include tungsten-halogen or other incandescent lamp, a xenon lamp or other gas emission radiation source, a fluorescent radiation source, an electronic radiation source (e.g., light emitting diode, laser diode, or laser), or other radiation source. 
     Infrared radiation source  14  may include a single wideband infrared source that emits radiation over two or more spectral bands. In some cases, infrared radiation from a single wideband infrared source may be separately channeled via separate spectral band selection devices (e.g., that include filters, prisms, or gratings) to form effective single-band sources. For example, the separate channeling may be performed sequentially to radiate infrared radiation in different spectral bands in quick succession (e.g., less than a millisecond) via a single component aperture  40  of unit aperture  19 . As another example, the radiation from the wideband source may be divided (e.g., using a beam splitter) and concurrently channeled via different band-selection devices to concurrently radiate in different spectral bands via separate component apertures  40  of unit aperture  19 . 
     Optics  18  of optical head  13  may direct infrared radiation from infrared radiation source  14  out unit aperture  19  (or from a component infrared source of infrared radiation source  14  out a component aperture  40  of unit aperture  19 ) to tissue surface  22 . For example, optics  18  may include one or more mirrors, reflectors, light pipes or optical fibers, lenses, filters, gratings, polarizers, beam splitters, prisms, apertures, collimators, shutters, or other components. Similarly, optics  18  may direct radiation from tissue surface  22  (e.g., reflected radiation) via unit aperture  19  to infrared radiation detector  16  (or via a component aperture  40  of unit aperture  19  to a component detector of infrared radiation detector  16 ). One or more components of optics  18  may function both to direct radiation from infrared radiation source  14  to tissue surface  22 , and to direct radiation from tissue surface  22  to infrared radiation detector  16 . Alternatively or in addition, separate components of optics  18  may be provided for either directing radiation from infrared radiation source  14  to tissue surface  22  or for directing radiation from tissue surface  22  to infrared detector  16 . 
     Optics  18  may be or may include a dispersive element (e.g., grating, prism, element with spectrally selective optical layers or coating, or another dispersive element), and a micro-mirror array for directing radiation of one or more selected wavelengths toward unit aperture  19  (e.g., for limiting irradiation of tissue surface  22  to selected wavelengths), or toward infrared detector  16  (e.g., for limiting detection of reflected or transmitted radiation to selected wavelengths). 
     Optics  18  may be configured to direct a portion of radiation that is emitted by infrared radiation source  14  to infrared detector  16 . Thus, an intensity of the radiation that is emitted by infrared radiation source  14  may be monitored. Optics  18  may include a shutter or other component that is configured to block radiation (e.g., that is emitted by infrared radiation source  14 ) from reaching infrared detector  16 . When the radiation is blocked, a baseline measurement may be made (e.g., a dark current or a detection level that is due to stray radiation). 
     Infrared detector  16  may be configured to detect SWIR radiation from tissue surface  22  that enters reflection measurement unit  12  via unit aperture  19 . Infrared detector  16  may include one or more component detectors. For example, radiation that is reflected by tissue surface  22  may be enabled to impinge on a component detector of infrared detector  16  via one of component apertures  40 . 
     Two or more different component radiation detectors may be configured to detect, or be optimized to detect, SWIR radiation in one or more spectral bands. A component detector may include a solid state or other photoelectric transducer or photodetector that is configured or optimized for one or more spectral bands of SWIR radiation. A component detector may include a thermal detector, a photon detector (e.g., including InGaAs), or another type of wideband detector. The temperature of a component detect of infrared detector  16  may be regulated (e.g., via thermoelectric cooling or heating) or may be unregulated. 
     Infrared detector  16  or controller  28  may include an amplifier to amplify a detection signal that is produced by infrared detector  16 . For example, the amplifier may include a trans-impedance amplifier or other amplifier. 
     Infrared detector  16  or controller  28  may include a logarithmic converter that enables direct calculation of an absorbance value of the tissue, or of a quantity that is proportional to an absorbance. The absorbance data may be used to analyze a liquid, such as water or blood, below tissue surface  22  (e.g., detect a pressure ulcer or a concentration of a drug or other substance in the blood). 
     Component apertures  40  may be arranged such that radiation in a particular wavelength band that is emitted by a particular component infrared source of infrared radiation source  14  and that is reflected by tissue surface  22  is likely to impinge on a corresponding (e.g., configured or optimized to detect radiation in that same wavelength band) component detector of infrared detector  16 . For example, the positions of a pair of corresponding component apertures  40  may be arranged such that radiation that irradiates tissue surface  22  via one of the corresponding component apertures  40  may be specularly reflected by tissue surface  22  into the other of the pair of corresponding component apertures  40 . 
     Unit aperture  19  of optical head  13  may be configured to be placed against or near tissue surface  22 . Optical head  13  may be provided with a removable protective cover  20  that may be placed over unit aperture  19  when reflection measurement unit  12  is in use. At least an outer surface of removable protective cover  20  may be constructed of materials that are suitable (e.g., approved by an appropriate organization) for contact with human skin or other tissue surfaces. At least a region of protective cover  20  (e.g., a region that is configured for placement over unit aperture  19 ) is substantially transparent or translucent in the spectral bands in which reflection measurement unit  12  is configured to operate. Suitable materials may include, for example, rigid vinyl, polycarbonate, POLY IR® plastic materials, or other materials such as poly urethane (PU), thermoplastic elastomers (TPE), silicones (LSR) and the like. 
     Removable protective cover  20  may include a structure (e.g., tab, projection, notch, clip, or other structure) that cooperates with corresponding structure on optical head  13  to prevent or inhibit removable protective cover  20  from accidentally or unintentionally falling off of optical head  13 , e.g., during use. 
     Protective cover  20  may be disposable, cleanable, or sterilizable. Removable protective cover  20  may be removed from optical head  13  and replaced (e.g., with a different removable protective cover  20 , or with the same removable protective cover  20  after cleaning and sterilization) between uses of reflection measurement unit  12  on different patients. Use of removable protective cover  20  may enable sanitary use of reflection measurement unit  12  on different patients while not exposing reflection measurement unit  12  from repeated cleaning or sterilization. 
     Liquid analysis system  10  may include a controller  28 . Controller  28  may include a microcontroller unit (MCU), or one or more other types of controller, microprocessor or processor. Controller  28  may include for example two or more intercommunicating devices or units. Controller  28  may be configured to control operation of infrared radiation source  14 , and to receive signals that are indicative of detected radiation from infrared detector  16 . For example, controller  28  may include circuitry that is configured to control operation of infrared radiation source  14  and infrared detector  16 . Controller  28  may be configured to operate in response to operation of user controls  27 . For example, user controls  27  may include one or more user touch-operable controls, such as pushbuttons, dials, switches, levers, touch-sensitive surfaces, or other touch-operable controls. User controls  27  may include other types of controls, such as light-sensitive controls, sound-operated controls, electromagnetically-operable controls, proximity sensors, pressure sensors, or other types of controls. 
     Controller  28  may be configured to dynamically adjust the intensity of radiation that is emitted by infrared radiation source  14 , e.g., in accordance with intensities that are detected by infrared detector  16 . For example, the intensity may be adjusted to accommodate various tissue thicknesses, skin coloration, or other characteristics. The intensities of a component infrared source may be adjusted in accordance with output of another component infrared source. 
     Controller  28  may be configured to digitally filter the signals of infrared detector  16 , e.g., to remove effects of baseline wandering and artifacts caused by patient movement. 
     Controller  28  may include a processor or processing units that may be configured to operate in accordance with programmed instructions. Controller  28  or a processor of controller  28  may communicate with an external device  30  via connection  36 . External processing device  30  may represent a device with processing capability, such as a computer, smartphone, or other device. External processing device  30  may be portable (e.g., a portable computer or smartphone) or may be fixed (e.g., a server). External processing device  30  may include or communicate with an input device  34  (e.g., keyboard, keypad, touch screen, pointing device, or other input device), an output device  32  (e.g., display screen or other output device), or both. Connection  36  may represent a wire or cable connection, a wireless connection (e.g., Bluetooth), a network connection, or another communications connection. 
     External processing device  30  may be utilized to communicate commands or programmed instructions to control operation of controller  28 . For example, external processing device  30  may be operated using input device  34  to download parameters or instructions (e.g., a measurement protocol) to controller  28 . Measured results from operation of reflection measurement unit  12 , or results of analysis of the measured results, preformed by, for example, external processing device&#39;s processor  38 , may be output by output device  32  of external processing device  30  for examination or review by a user of liquid analysis system  10 . 
     External processing device  30  may communicate (e.g., via a network such as the Internet) with one or more other processors, computers, or servers. For example, measure spectral reflection or transmission data may be communicated to a remote server. The remote server may analyze the transmitted data and return a diagnosis or other indication of a state of a medical condition. 
     Controller  28  may communicate with memory  24  (and/or external processing device&#39;s memory  39 ). Memory  24  may include one or more volatile or nonvolatile memory devices. Memory  24  may be incorporated within reflection measurement unit  12 , external processing device  30 , or elsewhere. Memory  24  may be utilized to store, for example, programmed instructions for operation of controller  28 , data or parameters for use by controller  28  during operation, or results of operation of controller  28 . 
     Controller  28  and or processor  38  may communicate with data storage device  26 . Data storage device  26  may include one or more fixed or removable nonvolatile data storage devices. Data storage device  26  may be incorporated within reflection measurement unit  12 , external processing device  30 , or elsewhere. For example, data storage device  26  may include a computer readable medium for storing program instructions for operation of processing unit of controller  28  or of external processing device  30 . It is noted that data storage device  26  may be remote from the processing unit. In such cases data storage device  26  may be a storage device of a remote server storing an installation package or packages that can be downloaded and installed for execution by the processing unit. Data storage device  26  may be utilized to store data or parameters for use by controller  28  during operation or results of operation of controller  28  (e.g., detection of radiation). 
     Data storage device  26  may be used to store data that relates spectral absorption, transmission, or reflection characteristics of tissue surface  22  to one or more medical conditions. The data may be stored in the form of a database. Processor or controller  28 , or another processor or controller may be configured to carry out methods as described herein. 
     In accordance with an embodiment of the present invention, a system for noninvasive analysis of subcutaneous liquids may be based on measured transmission of SWIR radiation. 
       FIG. 2A  is a schematic drawing of a measurement unit of a system for noninvasive analysis of subcutaneous liquids based on transmission of infrared radiation, in accordance with an embodiment of the present invention. 
     Transmission measurement unit  50  may be used in a system for noninvasive analysis of subcutaneous liquids, such as in liquid analysis system  10  (e.g., in place of, or in addition to, reflection measurement unit  12  of  FIG. 1A ). Transmission measurement unit  50  is configured to measure transmission through a body part  52 . For example, body part  52  may represent a part of the body (e.g., ear, finger, fold of skin) through which a measurable fraction of SWIR radiation is transmitted. 
     Transmission measurement unit  50  includes radiation source arm  54  and detection arm  55 . 
     Radiation source arm  54  may include infrared radiation source  14  and source optics  53   a . In some cases, infrared radiation source  14  may be located outside of radiation source arm  54 . In such a case, source optics  53   a  may be configured (e.g., with a mirror, light pipe, or optical fiber) to convey radiation from infrared radiation source  14  to source arm aperture  57   a.    
     As described above, infrared radiation source  14  may include two or more separate component infrared sources. Source optics  53   a  may be configured to convey radiation from the component infrared sources, concurrently or sequentially, to source arm aperture  57   a , or to separate component apertures of source arm aperture  57   a.    
     Source optics  53   a  may be or may include a dispersive element (e.g., grating, prism, element with spectrally selective optical layers or coating, or another dispersive element), focusing optics, and a micro-mirror array for directing radiation of one or more selected wavelengths of radiation from infrared radiation source  14  toward source arm aperture  57   a.    
     Detection arm  55  may include infrared detector  16  and detector optics  53   b . In some cases, infrared detector  16  may be located outside of detection arm  55 . In such a case, detector optics  53   b  may be configured (e.g., with a mirror, light pipe, or optical fiber) to convey radiation from detection arm aperture  57   b  to infrared detector  16 . 
     As described above, infrared detector  16  may include two or more separate component detectors. Detector optics  53   b  may be configured to convey radiation from detection arm aperture  57   b , concurrently or sequentially, to component detectors of infrared detector  16 , or from separate component apertures of detection arm aperture  57   b  to component detectors of infrared radiation detector  16 . 
     Detector optics  53   b  may include a dispersive element (e.g., grating, prism, element with spectrally selective optical layers or coating, or another dispersive element), focusing optics, and a micro-mirror array for directing radiation of one or more selected wavelengths of radiation from detector arm aperture  57   b  toward infrared detector  16 . 
     Radiation source arm  54 , detection arm  55 , or both, may be rotated outward (away from one another) or inward (toward one another). Outward rotation of radiation source arm  54  or detection arm  55  may enable insertion of body part  52  between the arms. Inward rotation of radiation source arm  54  or detection arm  55  may bring source arm aperture  57   a  and detection arm aperture  57   b  into contact with or near to body part  52 . Source arm aperture  57   a  and detection arm aperture  57   b  may each be covered with a removable protective cover  20 . 
     A rotation mechanism  56  may be configured to enable the outward or inward rotation of radiation source arm  54  and detection arm  55 . For example, rotation mechanism  56  may include a hinge, gimbal, bearing, or other mechanism to enable rotation of radiation source arm  54  or detection arm  55 . For example, a separate rotation mechanism  56  for one of radiation source arm  54  and detection arm  55 . Separate rotation mechanisms  56  may be provided for both radiation source arm  54  and detection arm  55  (e.g., as shown schematically in  FIG. 2A ). A single rotation mechanism  56  may be provided (e.g., a single hinge mechanism) between radiation source arm  54  and detection arm  55  (e.g., as shown schematically in  FIGS. 2B and 2C ). Rotation mechanism  56  may include a spring, latch, or other mechanism to hold radiation source arm  54  and detection arm  55  against body part  52  when body part  52  is inserted between radiation source arm  54  and detection arm  55 . Thus, rotation mechanism  56  may attach transmission measurement unit  50  to body part  52 . 
     Rotation mechanism  56  may be configured to enable measurement of a thickness of body part  52 . For example, rotation mechanism may include an encoder or other measuring device for measuring an angle of rotation of rotation mechanism  56 . Alternatively or in addition, rotation mechanism may include an angular scale or mechanical rotation gauge for determining an angle of rotation of rotation mechanism  56 . A measured rotation angle, together with a known distance from (e.g., and axis of rotation of) rotation mechanism  56  from source arm aperture  57   a  or from detection arm aperture  57   b  may be used (e.g., by a processor or controller) to calculate the thickness. 
     When source arm aperture  57   a  and detection arm aperture  57   b  are positioned on or near body part  52 , transmission measurement unit  50  may be operated to measure of transmission of SWIR radiation from infrared radiation source  14  through body part  52  to infrared detector  16 . 
       FIG. 2B  schematically illustrates attachment of the measurement unit of  FIG. 2A  to a finger. 
     For example, transmission measurement unit  50  may be clipped to finger tip  60  to measure transmission of SWIR radiation through finger tip  60 . For example, the transmission measurement may be indicative of a medical condition, such as the concentration of a substance in blood that flows through finger tip  60 . 
       FIG. 2C  schematically illustrates attachment of the measurement unit of  FIG. 2A  to an ear. 
     For example, transmission measurement unit  50  may be clipped to outer ear  62  to measure transmission of SWIR radiation through outer ear  62 . For example, the transmission measurement may be indicative of a medical condition, such as the concentration of a substance in blood that flows through outer ear  62 . 
     In accordance with an embodiment of the present invention, a reflection or transmission measurement may be utilized to characterize tissue in a patient. 
     A spectral reflectance measurement R(λ) may be expressed as 
               R   ⁡     (   λ   )       =         I   ⁡     (   λ   )       -       B   0     ⁡     (   λ   )               I   0     ⁡     (   λ   )       -       B   0     ⁡     (   λ   )                 
where I 0 (λ) is a measured source intensity, I(λ) is a measured reflected intensity, and B 0 (λ) is a baseline measurement (e.g., measured when infrared radiation source  14  is turned off or when infrared detector  16  is covered, e.g., by a shutter). Source intensity I 0 (λ) may be monitored continuously (e.g., by a dedicated detector), or may be measured in the absence of tissue (e.g., a skin surface or body part) in the optical path from infrared radiation source  14  to infrared detector  16 .
 
     In some cases, baseline measurement B 0 (λ) may be ignored when B 0 (λ) is much smaller than I 0 (λ) or I(λ). 
     The relative spectral absorbance A(λ), which may be used to characterize the tissue, may be calculated by: 
               A   ⁡     (   λ   )       =       -     log   ⁢           [       R   ⁡     (   λ   )         α   R       ]       ∼     -       log   ⁢           [       I   ⁡     (   λ   )           α   R     ⁢       I   0     ⁡     (   λ   )           ]       I   ,       I   0     &gt;&gt;     B   0                     
where the dimensionless value A(λ)=α A L is the relative spectral absorbance, α A  is the absorption coefficient, L is the path-length or tissue penetration depth; and α R  is the reflection coefficient. In general, the relative absorbance A(λ) corresponds to spectral extinction, which results from both absorption and scattering.
 
     Reflectance measurements in two or more spectral bands, Δλ i  may be performed on a single region of skin or tissue to yield separate measured values of A(Δλ i ) or R(Δλi). For example, the spectral bands may include two or more of the wavelength ranges ˜1400 nm-1500 nm (strongly absorbed by water), ˜1000 nm-1350 nm (no strong absorption by water), and ˜1500 nm-2100 nm (no strong absorption by water). 
     The differential absorption A Diff  may be calculated from measurements in two wavelength bands, Δλ i  and Δλ j , where i≠j: 
     
       
         
           
             
               A 
               Diff 
             
             = 
             
               
                 
                   A 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
                 - 
                 
                   
                     A 
                     ref 
                   
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         j 
                       
                     
                     ) 
                   
                 
               
               ∼ 
               
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     R 
                     ⁡ 
                     
                       ( 
                       
                         Δλ 
                         i 
                       
                       ) 
                     
                   
                   ] 
                 
                 - 
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     Δ 
                     ⁢ 
                     
                         
                     
                     ⁢ 
                     
                       
                         R 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           λ 
                           j 
                         
                         ) 
                       
                     
                   
                   ] 
                 
               
               ∼ 
               
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     
                       R 
                       ⁡ 
                       
                         ( 
                         
                           Δ 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             λ 
                             i 
                           
                         
                         ) 
                       
                     
                     
                       
                         R 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           Δ 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             λ 
                             j 
                           
                         
                         ) 
                       
                     
                   
                   ] 
                 
                 . 
               
             
           
         
       
     
     A ref  and R ref  refer the absorbance and reflectance in one of the spectral bands that serves as a reference band. For example, radiation in the reference band may be largely absorbed, scattered, or transmitted whether or not a medical condition to be detected is present. For example, the wavelength range of ˜1400 nm-1500 nm (strong water absorption), a portion of this range, or another similarly unaffected spectral range may be selected as the reference band. 
     Absorption, scattering, or transmission of radiation in one of the other spectral bands, referred to as the operating band, may be detectably dependent on the presence of absence of the medical condition. For example, the operating band may include one or both of the spectral ranges ˜1000 nm-1350 nm, ˜1550 nm-2100 nm, one or more portions of one or both spectral ranges, or another suitable spectral range. Here, and throughout the specification, the symbol ˜ indicates an approximation (e.g., ±10%). 
     The differential absorption may be related to the state of the tissue (e.g., presence, absence, degree, or other state of a medical condition). For example, a database of previous measurement results may associate a value of a differential absorption with a state of a medical condition such as inflammation (e.g., otitis media, or other inflammation), tumor (e.g., in the colon, or elsewhere), or other conditions. The differential absorption value may be used to differentiate between conditions (e.g., inflammation and tumor, healthy and diseased tissue), detect or measure liquid within tissue, or other conditions. 
     In some cases, reflection measurements may be made on a region of a tissue surface when the underlying tissue is expected to be healthy (e.g., based on other medical indications), and another where presence of unhealthy tissue is suspected. 
     The differential absorption A Diff  of two measurements with the same setup and in the same wave-band Δλ i ,i=1, 2, 3 yield two different spectral absorbance values, A ref (Δλ) and A SUS (Δλ), which correspond to healthy tissue (reference absorbance, A ref ) and suspicious tissue A SUS , respectively: 
     
       
         
           
             
               A 
               Diff 
             
             = 
             
               
                 
                   
                     A 
                     SUS 
                   
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
                 - 
                 
                   
                     A 
                     ref 
                   
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
               
               ∼ 
               
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     
                       R 
                       SUS 
                     
                     ⁡ 
                     
                       ( 
                       
                         Δ 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         
                           λ 
                           i 
                         
                       
                       ) 
                     
                   
                   ] 
                 
                 - 
                 
                   log 
                   ⁢ 
                   
                       
                   
                   ⁢ 
                   
                     ⌊ 
                     
                       
                         R 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           Δ 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             λ 
                             i 
                           
                         
                         ) 
                       
                     
                     ⌋ 
                   
                 
               
               ∼ 
               
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     
                       
                         R 
                         SUS 
                       
                       ⁡ 
                       
                         ( 
                         
                           Δ 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             λ 
                             i 
                           
                         
                         ) 
                       
                     
                     
                       
                         R 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           Δ 
                           ⁢ 
                           
                               
                           
                           ⁢ 
                           
                             λ 
                             i 
                           
                         
                         ) 
                       
                     
                   
                   ] 
                 
                 . 
               
             
           
         
       
     
     To improve detectability, all three spectral ranges Δλ i i=1, 2, 3 can be used simultaneously. 
     In some cases, chromophore content may be measured quantitatively. In some cases, differentiation is limited to two states, e.g., healthy or diseased (e.g., presence of pressure ulcer indicated by accumulation of subcutaneous liquid in the) tissue. 
     In some cases, a state of a medical condition (e.g., presence of diseased tissue) may be determined by calculating the tissue liquid index (TLI), the sub-dermal fluid index (SDFI), or another quantity. A parameter C, such as TLI, SDFI, or another parameter, can be defined as a normalized difference of the reflectance as measured at two different wavelength bands Δλ i  and Δλ j , where i≠j: 
     
       
         
           
             C 
             = 
             
               
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
                 - 
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         j 
                       
                     
                     ) 
                   
                 
               
               
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
                 + 
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         j 
                       
                     
                     ) 
                   
                 
               
             
           
         
       
     
     In some cases, C may be approximated by 
     
       
         
           
             C 
             = 
             
               
                 
                   
                     R 
                     ⁡ 
                     
                       ( 
                       
                         Δ 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         
                           λ 
                           i 
                         
                       
                       ) 
                     
                   
                   - 
                   
                     R 
                     ⁡ 
                     
                       ( 
                       
                         Δ 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         
                           λ 
                           j 
                         
                       
                       ) 
                     
                   
                 
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         j 
                       
                     
                     ) 
                   
                 
               
               ⁢ 
               
                   
               
               ⁢ 
               or 
             
           
         
       
       
         
           
             C 
             = 
             
               
                 
                   R 
                   ⁡ 
                   
                     ( 
                     
                       Δ 
                       ⁢ 
                       
                           
                       
                       ⁢ 
                       
                         λ 
                         i 
                       
                     
                     ) 
                   
                 
                 
                   
                     R 
                     ⁡ 
                     
                       ( 
                       
                         Δ 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         
                           λ 
                           i 
                         
                       
                       ) 
                     
                   
                   + 
                   
                     R 
                     ⁡ 
                     
                       ( 
                       
                         Δ 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         
                           λ 
                           j 
                         
                       
                       ) 
                     
                   
                 
               
               ⁢ 
               
                   
               
               ⁢ 
               or 
             
           
         
       
       
         
           
             C 
             = 
             
               slope 
               ⁢ 
               
                   
               
               [ 
               
                 R 
                 ⁡ 
                 
                   ( 
                   
                     Δ 
                     ⁢ 
                     
                         
                     
                     ⁢ 
                     λ 
                   
                   ) 
                 
               
               ] 
             
           
         
       
     
       FIG. 3  shows an example of a graph of spectral reflectance. 
     Graph  70  shows measured reflectance in arbitrary units as a function of wavelength. Normal tissue curve  72  may represent spectral reflectance for normal, or healthy, tissue. Diseased tissue curve  74  may represent spectral reflectance for diseased, or unhealthy, tissue. Water curve  76  represents spectral transmittance for water (e.g., for a particular optical path such as 1 mm; transmittance=1−absorbance) in arbitrary units. 
     It may be noted that in the wavelength band of 1400 nm-1500 nm (low water reflectance due to strong absorption of radiation), there is little difference between normal tissue curve  72  and diseased tissue curve  74 . However, in the adjacent bands (e.g., wavelength less than about 1350 nm), the difference is more pronounced. 
     Quantification of chromophores may enable estimation of changes in concentration levels of substances or materials (e.g., drugs, or other substances) that may be administrated by injection, infusion, or otherwise. 
     The spectral transmittance of blood T B  may be expressed as
 
 T   B (λ)= I (λ)/ I   0 (λ)= e   −(α     B     (λ)+α(λ))·L  
 
where α B (λ) represents the absorption coefficient of blood and (e.g., in units of cm −1 ), respectively, at wavelength λ, α(λ) is the absorption coefficient of additional components of the tissue, and L is the length of the absorbing path (e.g., in cm). I(λ) is the detected intensity of transmitted radiation, and I 0 (λ) is the intensity of incident radiation.
 
     Similarly, T % S , the spectral transmittance of a mixture of blood and an introduced substance may be expressed as
 
 T   % S (λ)= I (λ)/ I   0 (λ)= e   −(α     S     (λ)+α     B     (λ)+α(λ))·L  
 
where α S (λ) represents the absorption coefficient of the introduced substance.
 
     The absorption coefficients may relate to the concentrations of blood C B  and of the introduced substance C S :
 
α B (λ)=ε B (λ)· C   B , and
 
α S (λ)=ε S (λ)· C   S .
 
where ε B  and ε S  represent the absorptivity coefficients of blood and of the introduced substance (e.g., in units of l·mol −1 ·cm −1  or 1·g −1 ·cm −1 ; also referred to as the specific absorption coefficient or mass absorption coefficient), respectively.
 
     The relative absorbance A S  (dimensionless) at wavelength λ may be related to the concentration of introduced substance: 
     
       
         
           
             
               
                 A 
                 S 
               
               ⁡ 
               
                 ( 
                 
                   λ 
                   , 
                   
                     C 
                     S 
                   
                 
                 ) 
               
             
             = 
             
               
                 log 
                 ⁢ 
                 
                     
                 
                 ⁢ 
                 
                   ( 
                   
                     
                       T 
                       
                         % 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         S 
                       
                     
                     
                       T 
                       B 
                     
                   
                   ) 
                 
               
               = 
               
                 
                   ( 
                   
                     
                       
                         
                           ɛ 
                           S 
                         
                         ⁡ 
                         
                           ( 
                           λ 
                           ) 
                         
                       
                       · 
                       
                         C 
                         S 
                       
                     
                     + 
                     
                       α 
                       ⁡ 
                       
                         ( 
                         λ 
                         ) 
                       
                     
                   
                   ) 
                 
                 · 
                 
                   L 
                   . 
                 
               
             
           
         
       
     
     The relative absorbance may be expressed as a linear equation:
 
 A   S   =p   1   ·C   S   +p   2  
 
with coefficient p 1  (e.g. in l·g −1 ) and p 2  (dimensionless).
 
     The normalized spectral transmittance S at a wavelength λ n  may be calculated from a measurement k as: 
     
       
         
           
             
               S 
               ⁡ 
               
                 ( 
                 
                   
                     λ 
                     n 
                   
                   , 
                   k 
                 
                 ) 
               
             
             = 
             
               
                 
                   
                     I 
                     meas 
                   
                   ⁡ 
                   
                     ( 
                     
                       
                         λ 
                         n 
                       
                       , 
                       k 
                     
                     ) 
                   
                 
                 - 
                 
                   
                     I 
                     dark 
                   
                   ⁡ 
                   
                     ( 
                     
                       λ 
                       n 
                     
                     ) 
                   
                 
               
               
                 
                   
                     I 
                     ref 
                   
                   ⁡ 
                   
                     ( 
                     
                       
                         λ 
                         n 
                       
                       , 
                       
                         k 
                         0 
                       
                     
                     ) 
                   
                 
                 - 
                 
                   
                     I 
                     dark 
                   
                   ⁡ 
                   
                     ( 
                     
                       λ 
                       n 
                     
                     ) 
                   
                 
               
             
           
         
       
     
     I meas (λ n , k) is a measured radiation intensity at wavelength λ n  for measurement k, I meas (λ n , k) being proportional to T % S . I ref (λ n , k 0 ) is a reference signal for a measurement k 0  made prior to introduction of the substance into the blood, I ref (λ n , k 0 ) being proportional to T B . I dark (λ n ) represents a baseline measurement that is made in the absence of a radiation source, e.g., when the radiation source is switched off. 
     The differential spectral absorbance A Diff  for measurement k at two wavelengths λ 1  and λ 2  may thus be calculated as 
     
       
         
           
             
               
                 A 
                 Diff 
               
               ⁡ 
               
                 ( 
                 
                   
                     λ 
                     1 
                   
                   , 
                   
                     λ 
                     2 
                   
                 
                 ) 
               
             
             = 
             
               
                 
                   A 
                   ⁡ 
                   
                     ( 
                     
                       λ 
                       1 
                     
                     ) 
                   
                 
                 - 
                 
                   A 
                   ⁡ 
                   
                     ( 
                     
                       λ 
                       2 
                     
                     ) 
                   
                 
               
               = 
               
                 
                   
                     log 
                     ⁢ 
                     
                         
                     
                     [ 
                     
                       S 
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             1 
                           
                           , 
                           k 
                         
                         ) 
                       
                     
                     ] 
                   
                   - 
                   
                     log 
                     ⁡ 
                     
                       [ 
                       
                         S 
                         ⁡ 
                         
                           ( 
                           
                             
                               λ 
                               2 
                             
                             , 
                             k 
                           
                           ) 
                         
                       
                       ] 
                     
                   
                 
                 = 
                 
                   log 
                   ⁢ 
                   
                       
                   
                   [ 
                   
                     
                       S 
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             1 
                           
                           , 
                           k 
                         
                         ) 
                       
                     
                     
                       S 
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             2 
                           
                           , 
                           k 
                         
                         ) 
                       
                     
                   
                   ] 
                 
               
             
           
         
       
     
     For example, T B  and T % S  may represent relative spectral transmittances of blood and of a mixture of blood and a introduced substance measured at two wavelengths λ 1  and λ 2 : 
     
       
         
           
             
               
                 T 
                 B 
               
               = 
               
                 
                   
                     T 
                     ⁡ 
                     
                       ( 
                       
                         λ 
                         1 
                       
                       ) 
                     
                   
                   
                     T 
                     ⁡ 
                     
                       ( 
                       
                         λ 
                         2 
                       
                       ) 
                     
                   
                 
                 = 
                 
                   e 
                   
                     
                       - 
                       
                         ( 
                         
                           
                             K 
                             B 
                           
                           + 
                           α 
                         
                         ) 
                       
                     
                     · 
                     L 
                   
                 
               
             
             , 
             
               
 
             
             ⁢ 
             
               
                 T 
                 
                   % 
                   ⁢ 
                   
                       
                   
                   ⁢ 
                   S 
                 
               
               = 
               
                 
                   
                     T 
                     ⁡ 
                     
                       ( 
                       
                         λ 
                         1 
                       
                       ) 
                     
                   
                   
                     T 
                     ⁡ 
                     
                       ( 
                       
                         λ 
                         2 
                       
                       ) 
                     
                   
                 
                 = 
                 
                   
                     e 
                     
                       
                         - 
                         
                           ( 
                           
                             
                               K 
                               S 
                             
                             + 
                             
                               K 
                               B 
                             
                             + 
                             α 
                           
                           ) 
                         
                       
                       · 
                       L 
                     
                   
                   . 
                 
               
             
           
         
       
     
     K B =(α B (λ 2 )−α B (λ 1 )) and K S =(α S (λ 2 )−α S (λ 1 )) represent the differential absorption coefficients of blood and of the introduced substance, respectively. 
     The relative differential absorbance A S  (dimensionless) at two wavelengths λ 1  and λ 2  is related to the concentration of introduced substance as: 
     
       
         
           
             
               
                 A 
                 S 
               
               ⁡ 
               
                 ( 
                 
                   
                     λ 
                     1 
                   
                   , 
                   
                     λ 
                     2 
                   
                   , 
                   
                     C 
                     S 
                   
                 
                 ) 
               
             
             = 
             
               
                 log 
                 ⁢ 
                 
                     
                 
                 ⁢ 
                 
                   ( 
                   
                     
                       T 
                       
                         % 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         S 
                       
                     
                     
                       T 
                       B 
                     
                   
                   ) 
                 
               
               = 
               
                 
                   ( 
                   
                     
                       
                         ( 
                         
                           
                             
                               ɛ 
                               S 
                             
                             ⁡ 
                             
                               ( 
                               
                                 λ 
                                 1 
                               
                               ) 
                             
                           
                           - 
                           
                             
                               ɛ 
                               S 
                             
                             ⁡ 
                             
                               ( 
                               
                                 λ 
                                 2 
                               
                               ) 
                             
                           
                         
                         ) 
                       
                       · 
                       
                         C 
                         S 
                       
                     
                     + 
                     
                       α 
                       ⁡ 
                       
                         ( 
                         λ 
                         ) 
                       
                     
                   
                   ) 
                 
                 · 
                 L 
               
             
           
         
       
     
     The normalized spectral transmittance S at two wavelengths λ 1  and λ 2  may be calculated from measurement k as: 
     
       
         
           
             
               S 
               ⁡ 
               
                 ( 
                 
                   
                     λ 
                     1 
                   
                   , 
                   
                     λ 
                     2 
                   
                   , 
                   k 
                 
                 ) 
               
             
             = 
             
               
                 
                   [ 
                   
                     
                       
                         I 
                         meas 
                       
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             1 
                           
                           , 
                           k 
                         
                         ) 
                       
                     
                     - 
                     
                       
                         I 
                         dark 
                       
                       ⁡ 
                       
                         ( 
                         
                           λ 
                           1 
                         
                         ) 
                       
                     
                   
                   ] 
                 
                 / 
                 
                   [ 
                   
                     
                       
                         I 
                         meas 
                       
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             2 
                           
                           , 
                           k 
                         
                         ) 
                       
                     
                     - 
                     
                       
                         I 
                         dark 
                       
                       ⁡ 
                       
                         ( 
                         
                           λ 
                           2 
                         
                         ) 
                       
                     
                   
                   ] 
                 
               
               
                 
                   [ 
                   
                     
                       
                         I 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             1 
                           
                           , 
                           
                             k 
                             0 
                           
                         
                         ) 
                       
                     
                     - 
                     
                       
                         I 
                         dark 
                       
                       ⁡ 
                       
                         ( 
                         
                           λ 
                           1 
                         
                         ) 
                       
                     
                   
                   ] 
                 
                 / 
                 
                   [ 
                   
                     
                       
                         I 
                         ref 
                       
                       ⁡ 
                       
                         ( 
                         
                           
                             λ 
                             2 
                           
                           , 
                           
                             k 
                             0 
                           
                         
                         ) 
                       
                     
                     - 
                     
                       
                         I 
                         dark 
                       
                       ⁡ 
                       
                         ( 
                         
                           λ 
                           2 
                         
                         ) 
                       
                     
                   
                   ] 
                 
               
             
           
         
       
     
     The differential spectral absorbance A Diff  for measurement k at two wavelengths λ 1  and λ 2 , then is:
 
 A   Diff (λ 1 ,λ 2 )=log [ S (λ 1 ,λ 2   ,k )]
 
     The differential spectral absorbance measurement may eliminate the effects of background materials. For example, if the absorption and scattering by the background materials (e.g., tissue components other than water, or a substance that is introduced into the blood) are substantially constant in both measured spectral bands, than the differential spectral absorbance may be indicative of the water content of the tissue (e.g., as indicative of the presence or absence of a medical condition in which fluids accumulate in the tissue, or indicative of water content of blood). 
     A known relationship between the differential spectral absorbance and a concentration of the substance in the blood may be applied to the measured differential spectral absorbance to determine a concentration of the substance in the blood. For example, the known relationship be applied as a parameterized formula expressing the relationship (e.g., a polynomial or other formula), as a lookup table, or in another manner. 
       FIG. 4  shows a graph of an example of a relationship of absorbance to concentration of a substance. 
     Line  82  of graph  80  shows a relationship between a relative absorbance (dimensionless) and the concentration of a substance (e.g., Propofol) in blood, as plotted on a logarithmic scale (e.g., in units of μg/ml). The relationship may be derived from laboratory measurements  84 , e.g., from transmission measurements on cuvettes containing various concentrations of the substance in blood. A relationship may be derived by application of a fitting technique to fit line  82  to laboratory measurements  84 . 
       FIG. 5  is a flowchart depicting a method for noninvasive analysis of subcutaneous liquids, in accordance with an embodiment of the present invention. 
     It should be understood with respect to any flowchart referenced herein that the division of the illustrated method into discrete operations represented by blocks of the flowchart has been selected for convenience and clarity only. Alternative division of the illustrated method into discrete operations is possible with equivalent results. Such alternative division of the illustrated method into discrete operations should be understood as representing other embodiments of the illustrated method. 
     Similarly, it should be understood that, unless indicated otherwise, the illustrated order of execution of the operations represented by blocks of any flowchart referenced herein has been selected for convenience and clarity only. Operations of the illustrated method may be executed in an alternative order, or concurrently, with equivalent results. Such reordering of operations of the illustrated method should be understood as representing other embodiments of the illustrated method. 
     Operations of subcutaneous liquid analysis method  100  may be executed by a processor of a controller of a device for subcutaneous liquid analysis, or by a processor that is in communication with a controller of a device for subcutaneous liquid analysis. 
     Execution of subcutaneous liquid analysis method  100  may be initiated by a user of a device for subcutaneous liquid analysis. For example, a user may operate a control to initiate execution of subcutaneous liquid analysis method  100 . As another example, execution of subcutaneous liquid analysis method  100  may be initiated automatically when a device for subcutaneous liquid analysis is activated (e.g., turned on), and when it is detected (e.g., by an optical sensor or by a proximity sensor) that one or more apertures of the device are in contact with a tissue surface. 
     The tissue may be irradiated with SWIR radiation in a spectral band that is strongly absorbed by water (block  110 ). For example, a tissue surface may be irradiated with SWIR radiation in the wavelength range from about 1400 nm to about 1500 nm. The radiation may originate from a wideband source (e.g., an incandescent or other thermal source, or from a fluorescent source), or from a narrowband source (e.g., laser diode or light emitting diode). The irradiation may be filtered or otherwise manipulated. For example, radiation that is emitted by a wideband radiation source may be filtered or otherwise manipulated to select only that radiation that is within the water-absorbed spectral band. 
     The tissue may be irradiated with SWIR radiation in a spectral band that is adjacent to the water-absorbed spectral band (block  120 ). For example, a gap between the adjacent spectral band and the water-absorbed spectral band may be no more than 200 nm. In some cases, the gap may be no more than 100 nm. A single wideband source may produce both the radiation in the water-absorbed spectral band and in the adjacent spectral band. In some cases, the radiation in the adjacent spectral band may be isolated from radiation that is emitted by a wideband source prior to irradiation of the tissue. 
     Radiation that emerges from the tissue in each of the spectral bands may be detected (block  130 ). The detector is configured to produce a signal that is indicative of an intensity of the emerging radiation. 
     For example, one or more detectors may be configured to detect radiation that emerges from the tissue surface that is irradiated. For example, optics of the radiation source and the radiation detector may be aimed at a single tissue surface. In this case, the detector is configured to measure backscattered or reflected radiation. As another example, the detector may be configured to detect radiation that emerges from a tissue surface on the opposite side of the tissue from the tissue surface that is irradiated. In this case, the detector is configured to measure radiation that is transmitted by the tissue. 
     In some cases, different detectors may be configured to measure emerging radiation in each of the wavelength bands. For example, each detector may be constructed of a material that produces an electric signal only when irradiated with radiation in one of the spectral bands. As another example, detector optics (e.g., including a filter or grating) may restrict radiation that is outside of that spectral band from irradiating the detector. In some cases, the detector may be configured to detect radiation in both spectral bands. In this case, separate measurement of the emerging radiation in the different spectral bands may be effected by separate (e.g., sequential or alternating) irradiation of the tissue with radiation in each of the spectral bands. In some embodiments, emerging radiation may be measured at different distances from a location of the irradiation. For example, emerging radiation may be measured concurrently by a plurality of detectors that are arranged at different distances from a location on the tissue surface that is irradiated. As another example, emerging radiation may be measured sequentially in time at different distances by one or more detectors whose distance from the location of irradiation may be changed (e.g., automatically or manually). In this case, a detector may be provided with a sensor or mechanism (e.g., encoder or other sensor or mechanism) that is configured to measure a distance between a detector and the radiation source. In some embodiments, signals received from each detector may correspond to the distance of the detector from the light source. In some embodiments, the shape and/or intensity of a signal received by a detector may correspond to a medical condition in the tissue, for example shape of a signal changed in measurements for healthy tissue and for deep tissue injury, so that the shape of the signal received for a healthy tissue changes when measuring an injured tissue of the same subject (e.g., person). The shape of a signal may change because of different types of chromophores and different types of chemical functional groups on a chromophore. For example, a phenol ring on a chromophore will have different absorption graph shape in comparison to a CO chemical bond, CH chemical bond, or CN chemical bond. Absorption graph different shape means different location of the absorption peak or peaks in different wavelengths and at different peak heights (amplitudes). In some embodiments, for a single detector the shape of a signal may refer to the intensity (or amplitude) of absorption lines as a function of the wavelength (e.g., at different wavelengths). In some embodiments, for a plurality of detectors with different distances from the radiation (or illumination) source, the shape of the signal may refer to a slope, such as the intensity (or amplitude) of absorption line at given wavelength for all detectors. The processor may determine the change in shape of a signal by calculating (and/or testing) similarity of measured data sets with the predefined data with known (or calibrated) spectral response (e.g., for healthy tissue) and/or calculate a combined absorption graph (e.g., for a plurality of detectors). The similarity test may be performed by using at least one of the following techniques: correlation coefficient estimation, minimum of signal difference at given wavelength or waveband, t-test, Kruskal-Wallis test, Wilcoxon rank sum test, Kolmogorov-Smirnov test, etc. In some embodiments, a plurality of detectors may be used, each of the plurality of detectors may have a different distance from the radiation source. It should be appreciated that the signal received by each detector may have different amplitudes. According to some embodiments, the processor or controller may combine the readings of the plurality of detectors and calculate the slope of the combined absorption graph (e.g., a graph that represents the result of combining all detected amplitudes). A change in the state of the tissue may be determined, according to some embodiments, when a change in the slope of the combined absorption graph is detected (e.g., by processor  38  and/or controller  28  in  FIG. 1 ). 
     Detection of the emerging radiation may be preceded by, followed by, or may be concurrent with one or more calibration, baseline, or reference measurements. For example, a baseline or dark measurement may be made when the radiation source is not being operated. The baseline measurement may determine a signal that is produced by the detector (e.g., due to detector electronics or to stray radiation impinging on the detector surface) when no radiation of interest is present. A calibration measurement may be made when radiation that is emitted by the source is directly (e.g., not via tissue) conveyed to the detector. Such a calibration measurement, when made concurrently with, or immediately prior to or after, the measurement of emerging radiation may enable compensation for drifting in source intensity or detector sensibility. In some cases, a reference measurement may be made on the radiation that emerges from the tissue under known conditions (e.g., on a skin surface that is known to overlie healthy tissue, or prior to introduction of a substance into the blood). 
     The measurements of emerging radiation may be used to calculate a relative absorption by the tissue (block  140 ). 
     For example, a relative reflectance may be calculated by calculating a ratio of a measured intensity of reflected (e.g., due to backscattering) radiation in one of the spectral bands to the measured intensity of reflected radiation in another spectral band. A relative absorption may be inferred from the relative reflectance, e.g., on the assumption that a characteristic penetration depth of the radiation is the same in both spectral bands. If the characteristic penetration depth of the radiation is known (e.g., from laboratory experiments), a differential absorbance may be calculated. 
     A relative transmission may be calculated by calculating a ratio of a measured intensity of transmitted radiation in one of the spectral bands to the measured intensity of transmitted radiation in another spectral band. Since the path length through the tissue is the same for both spectral bands, a relative absorption may be inferred from the relative transmission. If the thickness of the tissue (L) is known (or may be estimated) and constant, a differential absorbance of the tissue may be calculated. 
     The relative differential absorbance A S  (dimensionless) at two wavelengths λ 1  and λ 2  is related to the concentration of introduced substance as: 
     
       
         
           
             
               
                 A 
                 S 
               
               ⁡ 
               
                 ( 
                 
                   
                     λ 
                     1 
                   
                   , 
                   
                     λ 
                     2 
                   
                   , 
                   
                     C 
                     S 
                   
                 
                 ) 
               
             
             = 
             
               
                 log 
                 ⁢ 
                 
                     
                 
                 ⁢ 
                 
                   ( 
                   
                     
                       T 
                       
                         % 
                         ⁢ 
                         
                             
                         
                         ⁢ 
                         S 
                       
                     
                     
                       T 
                       B 
                     
                   
                   ) 
                 
               
               = 
               
                 
                   ( 
                   
                     
                       
                         ( 
                         
                           
                             
                               ɛ 
                               S 
                             
                             ⁡ 
                             
                               ( 
                               
                                 λ 
                                 1 
                               
                               ) 
                             
                           
                           - 
                           
                             
                               ɛ 
                               S 
                             
                             ⁡ 
                             
                               ( 
                               
                                 λ 
                                 2 
                               
                               ) 
                             
                           
                         
                         ) 
                       
                       · 
                       
                         C 
                         S 
                       
                     
                     + 
                     
                       α 
                       ⁡ 
                       
                         ( 
                         λ 
                         ) 
                       
                     
                   
                   ) 
                 
                 · 
                 L 
               
             
           
         
       
     
     According to some embodiments, Raman spectroscopy as described in further detail below, may be employed to analyze the measurement data, and determine changes in concentration of any chemical compound inside the tissue due to, for example, a deep tissue injury. The calculated relative absorption may be utilized to determine a state of the tissue (block  150 ). For example, a calculated differential absorbance, relative reflection, relative transmission, or other calculated value that is related to relative absorption may be compared to previously measured values. The previously measured values may relate to a particular body part, suspected medical condition, introduced substance, or may otherwise relate to a specific state that is being examined. The comparison may include substitution of the calculated value in a functional relationship, may be used to retrieve an indication of the state of the tissue from a lookup table, or may be otherwise utilized in determining a state of the tissue. 
     It should be noted that deep tissue injuries may include presence of liquids at depths larger than about 5 mm compared to a corresponding region of a healthy tissue (e.g., of a healthy patient). Furthermore, deep tissue injuries may include increasing concentration (over time) of substances that are products of ischemic processes (causing damage to the tissue) such as free fatty acids and/or glycerol which are the breakdown products of the fatty part of the tissue (triglyceride). In the blood stream (unlike other tissues) triglyceride may appear in the form of very low density lipoproteins (VLDL) while in other tissue the structure of triglyceride is maintained. Accumulation of other substances (e.g., protease or myoglobin) may also indicate presence of a deep tissue injury. Therefore, in order to determine a state of deep tissue injury, calibration may be initially performed to identify substances in tissue indicating such an injury. Using the calculated relative absorption a state of deep tissue injury may be therefore determined, as described above. 
     According to some embodiments, a reflectance measurement may include detecting reflected radiation at different lateral distances from a radiation source. In some embodiments, analysis of such reflectance measurements at different distances may indicate a depth within the tissue of a detected feature. 
     According to some embodiments, non-invasive measurement may be carried out for determination of drug (e.g., intravenous delivered substance such as Propofol) or other substance concentration in the blood stream and/or in other subcutaneous tissue (e.g., of a mammal) using optical sensors. 
       FIG. 6A  schematically illustrates a system for measurement of reflection at a distance from a radiation source, in accordance with an embodiment of the present invention. 
     According to some embodiments, reflection measurement system  200  may include a source unit  202  and one or more detection units  204 . Source unit  202  may include at least one infrared (IR) radiation source  14  and source optics  18   a  (e.g., lenses etc.) for directing a beam of radiation, e.g., into a tissue surface  22 . Each detector unit  204  may include at least one IR detector  16  and detector optics  18   b  (e.g., lenses etc.) for directing radiation, for instance radiation emerging from tissue surface  22 , toward infrared detector  16 . Detector unit  204  may be located at a measurement distance  206  from source unit  202 . For example, measurement distance  206  may correspond to a distance between a center of an aperture of source unit  202  to a center of detector unit  204 , and/or another measurement that characterizes a lateral distance between source unit  202  and detector unit  204 . 
     In some embodiments, at least one of detector unit  204  and source unit  202  (or assemblies that include a plurality of detector units  204  and/or source units  202 ) may be moveable relative to one another so as to change measurement distance  206 . In this case, reflectance measurements at different measurement distances  206  may be measured sequentially in time. The distance  206  between detector unit  204  and source unit  202  at the time of a measurement may be determined with a dedicated mechanism and/or sensor. For example, detector unit  204  and source unit  202  may be mounted on a fixture that includes a mechanism for adjusting a distance therebetween. The fixture may include a telescoping rod, a bendable joint, or any other mechanism to adjust a distance between detector unit  204  and source unit  202  in a controllable manner. The fixture may include fixed stops at known distances between detector unit  204  and source unit  202 . Alternatively or in addition, a sensor may be provided to measure a distance between detector unit  204  and source unit  202 . For example, a telescoping rod or bendable joint may be provided with an encoder to measure relative movement between detector unit  204  and source unit  202 . As another example, a rangefinder sensor may directly measure a distance between detector unit  204  and source unit  202 . 
     In some embodiments, at least one source unit  202  and at least one detector unit  204  may be combined in a single measurement unit to concurrently measure radiation that emerges from tissue surface  22  at different lateral distances from source unit  202 . 
       FIG. 6B  schematically illustrates an arrangement of a measurement head for concurrent measurement of radiation that emerges from a surface at different lateral distances from the radiation source, in accordance with an embodiment of the present invention. 
     According to some embodiments, each measurement head  208  may include a single source unit  202  surrounded by a plurality of detector units  204  and separated by two different lateral distances form source unit  202 . In the configuration shown in  FIG. 6B , source unit  202  is surrounded by an arrangement of inner detector units  204   a , each laterally separated from source unit  202  by a distance that is substantially equal to first distance  206   a . Inner detector units  204   a  are surrounded by an arrangement of outer detector units  204   b , each separated from source unit  202  by a lateral distance that is substantially equal to second distance  206   b . In some embodiments, second distance  206   b  may be larger than first distance  206   a.    
     It should be noted that the arrangement shown in  FIG. 6B  has been selected for illustrative purposes only. An actual arrangement may differ from the arrangement shown in  FIG. 6B . For example, an arrangement may include detector units  204  at more than two distances from a source unit  202 . An arrangement may include a different pattern of detector units and/or of source units. An arrangement may include more than one source unit  202 . For example, different source units  202  may produce radiation with different wavelengths. A center unit may include a detector unit  204  (e.g., surrounded by sources  202 ). 
     In some embodiments, measurement of a distance to a detected feature may be advantageous. For example, measurement of a distance may enable differentiation between a superficial pressure ulcer and a deep ulcer. In some embodiments, such measurement of a distance may enable differentiation between a deep pressure injuries and superficial pressure ulcers. In some embodiments, a distance between the light source and the light sensor may be predetermined. In some embodiments, determination of such distance may allow determination of a medical condition in the tissue (e.g., deep tissue injury) due to changes in signal from a detector having a determined distance to the light source. 
     Pressure ulcers are areas of soft tissue breakdown that result from sustained mechanical loading of skin and underlying tissues. They can interfere with quality of life, activities of daily living, and rehabilitation and, in some cases, may be life threatening. Pressure ulcers can develop either superficially or deep within the tissues, depending on the nature of the surface loading and the tissue integrity. The superficial pressure ulcer type forms within the skin, with maceration and detachment of superficial skin layers. When allowed to progress, the damage may result in an easily detectable superficial ulcer. 
     In contrast, deep tissue ulcers arise in muscle layers covering bony prominences and are mainly caused by sustained compression of tissue. Deep tissue ulcers may develop at a faster rate than superficial ulcers, and result in more extensive ulceration with an uncertain prognosis. 
     In addition to absorption of radiation that traverses tissue, NIR and SWIR radiation may be strongly scattered by such tissue. The free scattering length may be in the range of about 0.3 mm to about 1 mm. In some embodiments, the scattering may be strongly forward peaked. Beyond a free transport scattering length (e.g., about 1 mm), directional correlation with the direction of the incident irradiation (the correlation between the actual direction and the direction of incidence) is lost such that radiation transport may be modeled as photon diffusion. In a photon diffusion model, scattering may be isotropic at locations that are more distant from radiation sources and boundaries than several times the free scattering length. Photons may follow complex trajectories that are considerably longer than the geometrical distance between the radiation source and the radiation detector. 
       FIG. 7  schematically illustrates paths of incident radiation to different detector locations, in accordance with an embodiment of the present invention. 
     According to some embodiments, a deep lesion  210  (or other deep tissue injury) may be distant from tissue surface  22 . Incident radiation  214  may enter tissue surface  22  (e.g., from a source unit  202  as shown in  FIG. 6A ). Emerging radiation  216   a ,  216   b , and  216   c  may emerge from tissue surface  22  at different distances from incident radiation  214 , having traversed radiation paths  218   a ,  218   b , and  218   c  respectively within the tissue. It should be noted that as shown in  FIG. 7 , only radiation path  218   c  passes through deep lesion  210 . Thus, only emerging radiation  216   c  may correspond to and be attenuated by water absorption in deep lesion  210 . 
     According to some embodiments, a connection between measured absorption and differences in concentration of subcutaneous liquids may be described with reference to a modified Beer-Lambert law: 
               A   ⁡     (   λ   )       =       -     log   ⁢           [       I   ⁡     (   λ   )       /       I   0     ⁡     (   λ   )         ]       =         (         ɛ   water     ⁢     C   water       +       ∑   i     ⁢       ɛ   i     ·     C   i         +     μ   S       )     ·     &lt;   L   &gt;             
where I 0 (λ) is a measured source intensity, I(λ) is a measured reflected intensity, A(λ) is the measured absorption (attenuation), ε water  is the absorptivity coefficient of water, C water  is the concentration of water, and ε i  and C i  are the absorptivity coefficients and concentrations of different absorbing compounds. The path length &lt;L&gt;˜(F·d) represents the total mean optical path in the tissue (e.g., in units of centimeters), where d is the distance between the point of incident radiation  214  and one of emerging radiation  216   a ,  216   b , or  216   c , and F is a scaling factor (related to the optical path length). The scattering coefficient may be expressed as μ s =μ s0 (1−g), where g is the scattering anisotropy.
 
     For example, it may be assumed that the ratio for a penetration depth below tissue surface  22  may be twice the lateral distance d between the points of incident radiation  214  and the emerging radiation  216   a ,  216   b , or  216   c.    
     Differences in concentration of subcutaneous liquids may be estimated from the slope (ΔA/Δd) of the attenuation with respect to distance d, where A 1 −A 2 ≡ΔA=ε·(ΔC·&lt;L&gt;), with A 1  and A 2  representing differential absorptions A Diff  as described above that are measured with two different lateral distances d between incident radiation  214  and emerging radiation  216   a ,  216   b , or  216   c.    
     In some embodiments, the change in differential measurements may be analyzed to obtain ΔA/ε=ΔC&lt;L&gt;. When the relationship between path length &lt;L&gt; and lateral distance d is known, the measurements may be analyzed to yield a change in concentration of water. 
     According to some embodiments, Raman spectroscopy may be employed to analyze the measurement data. In Raman spectroscopy a complementary scattering mechanism to excite the molecules into the vibrational states may be achieved via the visible excitation wavelengths. Raman scattering may be an inelastic scattering which is usually generated by intensive monochromatic light (e.g., laser) in the visible, near infrared, or near ultraviolet (UV) region. The energy of laser photons may be changed after the excitation laser interacts with the vibrating molecules or the excited electrons in the sample. As a spontaneous effect, photons may transfer the excitation energy to change the molecule from the ground state to a virtual state. The excited molecule may then return to a different rotational or vibrational state after emitting a photon. 
     As may be apparent to one of ordinary skill in the art, the difference between wavelengths of photons in the incident wavelength λ 0  (excitation wavelength) and the scattered light is known as a Raman shift. It is related to characteristic oscillation frequencies of the molecule, and may correspond to the oscillations of a single molecular bond or the larger fragment of a molecular network. For λ 0  far from the molecule absorption band, intensity of the Raman signal may be inversely proportional to λ 0   4  so application of a VIS or UV laser as the excitation source may be more effective than an IR one if the intensity of Raman scattering was considered. However, practical efficiency of Raman scattering versus excitation wavelength may also depend on dimensions of the investigated structures. 
     Moreover, fluorescence induced by the laser beam maybe also taken into account. Fluorescence is the strongest for the excitation wavelength λ 0  range extending from 270 to 700 nm but its influence can be different for various materials. It is particularly strong for organic materials, so the excitation range in VIS and near UV is not appropriate for their sampling. 
     In some embodiments, application of NIR lasers (e.g., in the range of 785-1064 nm) may be effective with selection of appropriate detector types (e.g., InGaAs, MCT, etc.) that can ensure high efficiency of the measurement system in a wide Raman range of 800 cm −1  to 4000 cm −1 . Raman range of systems using such detectors may begin at about 800 cm −1  for excitation wavelength equal to 830 nm. 
     In some embodiments, a measurement system (such as reflection measurement system  200  shown in  FIG. 6A ) to employ Raman spectroscopy may include at least one light source (e.g., such as source unit  202  shown in  FIG. 6A ). The source unit may include a diode laser (e.g., ˜100 mW), light emitting diodes (LEDs) and/or a combination thereof. In some embodiments, Raman measurement system may include at least one light sensor (e.g., such as detection unit  204  shown in  FIG. 6A ). The light sensor may include a thermoelectric cooled charge coupled detector (CCD) and/or a spectrograph. In some embodiments, the Raman measurement system may include optical elements (e.g., such as optics  18   a ,  18   b  shown in  FIG. 6A ) with at least one of beam expanding/focusing lenses, laser-line filter, dichroic mirrors, holographic rejection (notch) filters, low-pas filter, and fibers. In some embodiments, at least two elements of the Raman measurement unit may be embedded in a plastic polymer casing within a wrist band which positions the sensing unit firmly above a wrist blood vessel. The unit may be wirelessly connected to an external processor such as external device  30  shown in  FIG. 1A  (e.g., via Bluetooth protocol). In some embodiments, in order to reduce the influence of background fluorescence signal, the 830 nm excitation wavelength may be used. According to some embodiments, Raman measurement system may indicate a change in concentration of a chemical compound within the tissue, for example due to deep tissue injury after calibration with healthy tissue is carried out. 
     According to some embodiments, a Raman measurement system may include at least one processor (such as controller  28  shown in  FIG. 1A ) and/or communicate with at least one processor (such as external processing device  30  shown in  FIG. 1A ) so as to allow analysis of the measured data according to Raman spectroscopy. In some embodiments, such analysis may include measurements of the distance between the light source and the light sensor. In some embodiments, such analysis may allow determination of liquid accumulation and/or determination of concentration of chemical compounds within the tissue (e.g., myoglobin, triglycerides, proteins, etc) and thereby determine deep tissue injuries if a change in the concentration corresponds to deep tissue injuries. In some embodiments, such analysis may allow determination of at least one substance secreted as a result of a deep tissue injury. 
     In some embodiments, the changes in the spectra (e.g., change in intensity of absorption spectra over time) of blood with Propofol mixture may overlap with the most prominent lines of the Propofol solution, for instance at spectral regions: 875 cm −1 -1050 cm −1 , 1250 cm −1 -1260 cm −1 , and 1450 cm −1 . 
     The Raman intensity may be 10 −6  to 10 −9  times less than that of Rayleigh scattering. Therefore, well-controlled high-power light sources (e.g., ˜100 mW) and sufficient accumulation time (e.g., tens of a second) may be required in order to produce a sufficient number of Raman-scattering photons. In the IR spectroscopy, abundance of water in most biological environments, a very strong IR absorption by water may interrupt the emitted photons of the target. In contrast, a weak Raman scattering by water may have detection of bio-molecular signals in water-abundance environments, such as body fluids, cells and/or other tissues. 
     According to some embodiments, a controlled cyclic change in external conditions applied on a region of the skin may also allow determination of deep tissue injuries. In some embodiments, applying pressure (e.g., measuring with a pressure sensor) in varying magnitude and/or applying heat (e.g., measuring with a temperature sensor) in varying magnitude may cause different absorption and/or scattering of radiation in tissues with deep tissue injuries (compared to absorption and/or scattering in healthy tissue). For example, pressure may be applied in a periodical (e.g., sinusoidal) regime to determine presence of a deep tissue injury. 
     In some embodiments, an at least partially transparent (to visible and/or IR radiation) biodegradable element, for instance an elastomer or other soft plastic, may be placed upon the at least one radiation detector as a cover so as to allow protection of sensors (e.g., from gels applied on the skin). In some embodiments, such biodegradable element may be transparent at least 90%. 
     According to some embodiments, specific regions of a measured MIR may be compared to VIS-NIR-SWIR spectrum for non-invasive monitoring of substance concentration (e.g., Propofol, etc.) in tissue (e.g., in blood). In some embodiments, substances that exhibit unique VIS-NIR-SWIR-MIR signatures during measurements may be monitored. In some embodiments, the spectral signatures may be used as input for multivariate classification in order to automatically quantify and/or identify the response. The multivariate classification methods may include at least one of partial-least square (PLS), principal component regression (PCR), linear discriminant analysis (LDA), k-nearest neighbors (kNN), naive bayesian classifier (NBC), support vector machine (SVM), artificial neural networks (ANN), etc. Spectral attenuation (e.g., in the VIS-NIR-SWIR range) in tissue may occur due to absorption from chromophores (e.g., of fixed or variable concentration) and/or due to scattering from the tissue. In some embodiments, absorption and scattering properties of the tissue may be characterized by absorption (μa) and scattering (μs) coefficients as: μa˜εC and μs=μs0(1−g), where ‘ε’ is the absorptivity, ‘C’ is the concentration of absorbing compounds, and ‘g’ is the scattering anisotropy. The spectral attenuation (e.g., due to absorption) may be represented as:
 
 A (λ)=−log [ I (λ)/ I   0 (λ)]=(Σε· C+μ   s )·&lt; L&gt; 
 
where ‘I(λ)’ is the detected intensity of transmitted and/or scattered radiation, I 0 (λ) is the intensity of incident radiation, &lt;L&gt;˜(F·d) is the total mean optical path in the tissue (in centimeters), ‘d’ is the distance between the light source and the detector, and ‘F’ is the scaling factor as a function of (μs, μa, g). For example, absorption and/or scattering for soft tissues (e.g., for skin) in the range of 800-1300 nm may be μa˜0.1-2 cm −1  and μs˜1-25 cm −1 .
 
     According to some embodiments, some substances (e.g., Propofol) may have unique effect on the absorption properties of the blood, such that these substances may be monitored within the blood stream in particular. In some embodiments, the spectral transmittance of blood ‘T B ’ may be expressed as:
 
 T   B (λ)= I (λ)/ I   0 (λ)= e   −(α     B     (λ)+α(λ))·L  
 
where α B (λ) is the absorption coefficient of blood (e.g., in units of cm −1 ) at wavelength λ, α(λ) is the absorption coefficient of other components of the tissue, and ‘L’ is the length of the absorbing path (e.g., in centimeters). ‘I(λ)’ is the detected intensity of transmitted radiation, and ‘I 0 (λ)’ is the intensity of incident radiation.
 
     In some embodiments, the spectral transmittance of a mixture of blood and an introduced substance (e.g., Propofol) ‘T % S ’ may be expressed as:
 
 T   % S (λ)= I (λ)/ I   0 (λ)= e   −(α     S     (λ)+α     B     (λ)+α(λ))·L  
 
where α S (λ) is the absorption coefficient of the introduced substance.
 
     In some embodiments, the absorption coefficients may relate to the concentrations of blood ‘C B ’ and the introduced substance ‘C S ’:α B (λ)=ε B (λ)·C B , and α S (λ)=ε S (λ)·C S . where ‘ε B ’ and ‘ε S ’ represent the absorptivity coefficients of blood and of the introduced substance (e.g., in units of liter·mol −1 ·cm −1  or liter·g −1 ·cm −1 , also referred to as the specific absorption coefficient or mass absorption coefficient), respectively. 
     In some embodiments, the relative absorbance ‘A S ’ at wavelength ‘λ’ may be related to the concentration of introduced substance as a linear equation: 
     
       
         
           
             
               
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     In some embodiments, the normalized spectral transmittance ‘S’ at a wavelength ‘λ n ’ may be calculated from a measurement ‘k’ as: 
               S   ⁡     (       λ   n     ,   k     )       =           I   meas     ⁡     (       λ   n     ,   k     )       -       I   dark     ⁡     (     λ   n     )               I   ref     ⁡     (       λ   n     ,     k   0       )       -       I   dark     ⁡     (     λ   n     )                 
where I meas (λ n , k) is a measured radiation intensity at wavelength λ n  for measurement ‘k’, I meas (λ n , k) being proportional to ‘T % S ’ as described above. I ref (λ n , k 0 ) is a reference signal for a measurement ‘k 0 ’ made prior to introduction of the substance into the blood, I ref (λ n , k 0 ) being proportional to ‘T B ’ as described above. I dark (λ n ) represents a baseline measurement that may be made in the absence of a radiation source (e.g., when the radiation source is switched off).
 
     In some embodiment, the change concentration (e.g., of Propofol) due to transmission and/or reflection may be estimated from the differences in attenuation between two blood states: 
               Δ   ⁢           ⁢     A   ⁡     (   t   )         =           A     B   -   P       ⁡     (   t   )       -       A   B     ⁡     (     t   0     )         =       log   (       I   B       I     B   -   P         )     =     &lt;   L   &gt;         ·     ɛ   P     ·   Δ     ⁢           ⁢     C   P       +   G                 
where A B (t0) is the spectral attenuation without the substance (e.g., Propofol), A B-P (t) is the spectral attenuation with blood-substance mixture, &lt;L&gt; is the mean path-length, ‘ε P ’ is the absorptivity of the substance (e.g., Propofol), ΔC P  is the concentration difference for the substance (e.g., Propofol), and G is a constant dependent on at least one of μs, μa, g, and measurement geometry. In some embodiment, the differences in attenuation between two blood-substance states may be estimated from the slope (ΔA/Δt) of the attenuation with respect to time ‘t’, where:
 
 A   1   −A   2   ≡ΔA =ε·(Δ C·&lt;L &gt;),
 
where ‘A 1 ’ and ‘A 2 ’ represent absorptions that are measured at two different time units with a given lateral distance between incident radiation and emerging radiation.
 
     According to some embodiments, some measurements may be carried out with detection of transmission radiation, for instance radiation passing through the ear (e.g., using the transmission measurement unit  50  as shown in  FIGS. 2A-2C ). In some embodiments, a transmission measurement unit may include at least one of a light source (e.g., in VIS-NIR-SWIR range), optical elements (e.g., collimating lenses, filters, focusing lenses), a detector (e.g., for the VIS-NIR-SWIR range), spectrometer, a motion sensor (e.g., gyro, accelerometer, etc.), and an attachable segment (e.g., a clip). In some embodiments, the attachable segment may be configured to attach at least one light source and corresponding at least one sensor to be aligned accurately against each other from both sides of ear cartilage. 
     In some embodiments, the attachable segment (e.g., ear clip) may include adhereable tips configured to allow stable positioning on the ear cartilage. In some embodiments, at least one of the light source and corresponding sensor may include a (thin) plastic polymer shaped in an ergonomic half disc shape, so as to allow easy positioning of the clip inside the round curvature of the ear cartilage border. 
     In some embodiments, the at least one sensor may be battery operated and/or with a wireless connectivity protocol (e.g., Bluetooth connectivity). The ear clip may contain the sensing unit on one arm and the light source unit on the other arm. The ear clip may include an adjustment knob to regulate pressure on ear cartilage avoiding over pressure which reduce blood flow in the ear cartilage capillaries. In some embodiments, the transmission measurement unit may include a motion sensor with a corresponding algorithm configured to adjust (and or compensate) to sudden movements during measurements. For example, a compensating mechanism may stop the measurements during movement of the subject and/or calculate and/or interpolate missing readings for the measurements. 
     According to some embodiments, some measurements may be carried out with detection of reflected radiation, for instance measuring backscattered radiation. In some embodiments, a reflectance measurement unit may include at least one of a light source (e.g., in VIS-NIR-SWIR range), optical elements (e.g., collimating lenses, filters, focusing lenses), a detector (e.g., for the VIS-NIR-SWIR range), spectrometer, one or more motion sensors (e.g., gyro, accelerometer, etc.), and an attachable segment (e.g., guide wire attachable on a limb of the subject). In some embodiments, the attachable segment may be configured to be attached on a limb with at least one light source and corresponding at least one sensor to be aligned in proximity to topical blood vessels (e.g., carotid arteries, jugular veins, grate saphenous veins, cephalic vein, arcuate arteries, etc.). In some embodiments, the reflectance measurement unit may include a motion sensor with a corresponding algorithm configured to adjust (and or compensate) to sudden movements during measurements. It should be noted that patients administered with Propofol are known to move during anesthesia due to short duty cycles of Propofol, and sudden (e.g., mechanical) movements during measurements may alter the optical absorption and/or reflection signal. In order to make sure that signal change is received from mechanical movements (e.g., and not from a heart attack) a movement sensor may be attached to the limbs or head so as to detect and compensate sudden movements during measurements. For example, a compensating mechanism may stop the measurements during movement of the subject and/or calculate and/or interpolate missing readings for the measurements. 
       FIG. 8  is a block diagram showing elements of a measurement unit, in accordance with an embodiment of the present invention. A measurement unit  300  (e.g., optical transmission and/or reflectance unit) may be used to measure signals of the body in order to determine concentration of predetermined substances (e.g., Propofol). The measurement unit  300  may include a central controller or processor  301  (e.g., such as controller  28  shown in  FIG. 1A ), at least one light source  302 , and at least one light sensor  303 . In some embodiments, measurement unit  300  may further include at least one motion sensor  304  and/or an attachable segment  305  (e.g., a clip) configured to be attach to a portion of the body of the patient. 
     In some embodiments, at least one light source  302  may be coupled to at least one optical element  306  (e.g., collimating lenses, filters, focusing lenses), and at least one light sensor may be coupled to a spectrometer  307 . Processor  301  may communicate (e.g., receive and/or send signals) with coupled elements (such as the light source and/or the light sensor) and analyze the received measurement data. Processor  301  may communicate (e.g., receive and/or send signals) with an external device using a communication module  308  (e.g., via Bluetooth). 
     According to some embodiments, measurements may be carried out with attenuated total reflection (ATR) spectroscopy using optical elements and/or sensors operated in the ATR regime in the range of 4000 cm −1  to 400 cm −1  (or 2.5 um to 25 um). In some embodiments, analysis of such measured data may be carried out with evanescent wave Fourier transform infrared (EW-FTIR) spectroscopy and/or FTIR-ATR spectroscopy. 
     As may be appreciated by someone with ordinary skill in the art, in the MIR range some fundamental molecular vibrations may occur as well as many of the first overtones and combinations thereof. The spectral bands in the MIR range tend to be sharp and may have high absorption. Since the bands are sharp, most small molecules may have distinctive spectral “fingerprints” that can be readily identified in mixtures (e.g., a mixture of Propofol and blood). Moreover, since individual peaks can often be associated with individual functional groups, it may be possible to detect changes in the spectrum of individual objects during the monitoring due to the corresponding specific reaction. 
     Fourier transform infrared (FTIR) spectroscopy is based on the interaction between the radiation and the sample (e.g., Propofol), which absorbs the IR wavelengths causing transitions between vibrational energetic levels, and thus vibrational modes of different chemical bonds and/or molecules may be detected. 
     According to some embodiments, the FTIR-ATR spectroscopy may utilize the total internal reflection phenomenon, occurring when a beam of radiation enters from a denser medium with a higher refractive index, n 2 , into a less-dense medium with a lower refractive index, n 1 . In FTIR-ATR spectroscopy a crystal with a high refractive index and IR transmitting properties may be used as internal reflection element (or an ATR crystal). The ATR element  309  may be placed in contact with the sample (e.g., liquids or a moist epithelial tissue). 
     In some embodiments, the beam of radiation propagating in ATR may undergo total internal reflection at the interface of ATR-sample. Total internal reflection of the light at the interface between the two media of different refractive index (crystal-tissue) may therefore create an “evanescent wave” that penetrates into the medium of lower refractive index (or tissue). The intensity of evanescent waves may decay exponentially with distance from the interface at which they are formed. Such distance may be for example in the 1-25 um range. 
     In some embodiments, the evanescent wave may be attenuated in regions of the spectrum where the sample absorbs energy, and the attenuated energy may be passed back to the optical element. The radiation may then exit the optical element and impinge on a detector through optical waveguide and/or fiber. The detector may record the attenuated radiation, which may be transformed to generate a spectrum (e.g., an absorption spectra). 
     The depth of penetration ‘dp’ into the sample of the evanescent wave may be defined as the distance from the ATR-sample boundary where the intensity of the evanescent wave decays to about 1/e (or 37% of its original value), and may be represented as:
 
 dp =( n   ATR   2  sin 2   θ−n   S   2 ) −1/2   /k,k =2π/λ
 
where ‘λ’ is the wavelength of the radiation, ‘θ’ is the angle of incidence of the wave front (or light beam), and ‘n ATR ’ and ‘n S ’ are the refractive indices of ATR element and sample respectively.
 
     The sample under test may be placed in tight contact with the ATR element along the IR radiation pathway, between the source and the detector side. When the ATR element is in contact with the sample, the evanescent wave may be either partially or totally absorbed at specific absorption lines as determined by the biochemical composition of the sample. The total transmission through the ATR element and the sample may decrease at the absorption lines that correspond to molecular bonds in particular classes of bio-molecules. In some embodiments, a core silver halide fiber may be used as the ATR element, since the polycrystalline silver halide (e.g., AgClxBr1-x) fibers are useful for applications in the mid-IR. These fibers may have a wide transparency range (˜2-20 um wavelength), they are non-toxic, flexible and may be insoluble in water. 
     According to some embodiments, FTIR-ATR unit may include a central controller or processor (e.g., such as processor  301  shown in  FIG. 8 ), at least one IR light source (e.g., such as light source  302  shown in  FIG. 8 ), for instance a broadband IR source for emitting an IR beam into the ATR element  309 . In some embodiments, the FTIR-ATR unit may include at least one light sensor (e.g., such as light sensor  303  shown in  FIG. 8 ), for instance a beam delivery and collection optics (e.g., a guide wire or fiber). 
     In some embodiments, the FTIR-ATR unit may include ATR element  309  at the tip of a guide wire (e.g., embedded in a clip arm, such as attachable segment  305  shown in  FIG. 8 ) thereby positioning the ATR element  309  in an intimate fixed contact with a moist epithelial tissue, for instance of the inner lower or upper lip (where there is no stratum corneum). Such a clip may have for example a curved arced shape with the ATR element  309  embedded on the clip arm in such a way that the curve positions the ATR probe on the inner lower or upper lip. In some embodiments, the FTIR-ATR unit may include an FTIR spectrometer (e.g., such as spectrometer  307  shown in  FIG. 8 ) and/or IR sensors for simultaneously measuring absorbance of specific regions of the IR spectrum. 
     The ATR tip may have different forms: for example loop, flat form, lens-like form, hexagon, triangle, etc. The ends of ATR element may be connected to the IR source and detecting system by using the waveguide for the FTIR-ATR device implementation. In some embodiments, the ATR element may be configured to allow multiple internal reflections (e.g., 3-15 internal reflections) against the measurement surface prior to measurement by the IR sensors. 
     According to some embodiments, an ATR element may be placed within running urine in a urine collecting catheter or in the urine collecting bag in order to preform measurements of the urine. In some embodiments, the ATR element frontal surface may be dipped in the urine stream while the rest of the ATR element may be embedded in the plastic polymer catheter tube or bag. 
     According to some embodiments, in order to receive the benefit from the data of concentration of a particular substance (e.g., Propofol) in the blood stream there may be a need to combine the substance (e.g., Propofol) readings with other parameters of the patient, for instance utilizing an algorithm which combines a set of parameters. According to some embodiments, in order to quantify concentration of anesthetic drugs (e.g., Propofol, etc.) and/or identify and/or classify deep tissue injury (DTI), the algorithm may use the multivariate classification methods, for instance including at least one of partial-least square (PLS), principal component regression (PCR), linear discriminant analysis (LDA), k-nearest neighbors (kNN), naive bayesian classifier (NBC), support vector machine (SVM), artificial neural networks (ANN), etc. In some embodiments, such a process may include training (or calibration) and validation steps. The Propofol concentration (e.g., in g/mL, mol/mL, etc.) may be considered as the dependent variables (‘Y’), and the other variables (e.g., data matrices A, B, C, D, E) may be the independent variables (‘X’). Data matrices A, B, C, D, and E may refer to the matrices of measured values: such as the spectral dataset with Raman data (matrix A), FTIR-ATR data (matrix B), reflectance mode data (matrix C), transmittance mode data (matrix D) and deterministic and/or vital signs parameters (matrix E). The multivariate classification (e.g., with multiple regression) may allow finding a solution that best predicts the ‘Y’ variable as a linear (or non-linear) function of the ‘X’ variables. For example, the general equation for the multiple regression may be for example: 
             Y   =         b   0     +       b   1     ⁢     x   1       +     …   ⁢           ⁢     b   k     ⁢     x   k       +   f     =         b   0     +       ∑     i   =   1     k     ⁢       b   i     ⁢     x   i         +   f     =       b   0     +   f   +   BX               
where ‘Y’ is the response (e.g., Propofol concentration), x k  are the predictors (A, B, C, D, and E), bk are the regression coefficient to be determined, b 0  is the offset and a constant factor, and ‘f’ is the residual. If ‘X’ and ‘Y’ are mean-centered, then b 0 =0. This equation may also be written in matrix form: Y=Xb+f.
 
     The parameter ‘b’ may be estimated for example by a least squares (LS) fit minimizing the sum of squared residuals, where multiple linear regression (MLR) may be used for estimating the regression vector ‘b’. The solution for regression coefficient for the LS may be for example: b=(X T X) −1 X T Y, where ‘T’ refers to the transpose of the matrix. In some embodiments, after training and/or validation the regression coefficients bk may be determined and/or stored. In some embodiments, the processor may use the coefficients bk during the real-time quantification of changes in concentration of anesthetic drugs in blood during the anesthesia. 
     Deterministic parameters may include the patient&#39;s age, gender, weight, bio mass index (BMI), background diseases, and other changing parameters of vital signs such as temperature, breathing rate, blood pressure, PO 2 -pulse oximetry, EEG, bi-spectral index (BIS), signal sensors and patient movements (e.g., measured with a gyro) to create a minimal anesthetic concentration (MAC) model. 
     The recombined set of parameters may continuously or iteratively determine (e.g., in a closed loop controlled process of continuous measurements) the minimal dose of substance (e.g., Propofol) for injection to achieve the desired anesthetic effect. It should be noted that the combination of the standard vital signs and other parameters with substance (e.g., Propofol) concentration data in real time and the algorithm which controls the syringe pump may provide optimal measurements. The movement sensors (e.g., gyro and/or accelerometers) may be combined in a wrist band, a leg band, or a temporal band (along the temporal bone), and wirelessly communicate with the main CPU (e.g., such as such as processor  301  shown in  FIG. 8  or external device  30  shown in  FIG. 1A ), for instance via Bluetooth protocol. 
     According to some embodiments, a measurement system may include a measurement unit (e.g., as shown in  FIG. 8 ) such as a transmission unit, a reflectance unit, FTIR_ATR unit, Raman unit, and urine collection unit. The measurement unit may also receive parameters of the subject body (e.g., vital signs) measured with dedicated sensors. The measurement unit may include at least one optical sensor units positioned at least on one wrist and/or on the mouth lip and/or on the outer ear and/or on the leg and/or in a urine collection catheter or bag. Each sensor may be operated simultaneously and/or simultaneously collect data from different sensors. The collected data may be transmitted to a main CPU (e.g., such as processor  301  shown in  FIG. 8  or external device  30  shown in  FIG. 1A ) which may process the data together with data collected from other vital sign sensors and other preprogrammed data to control continuous injection of the substance (e.g., Propofol) to the patient. 
     In some embodiments, the sensors which are located on the wrist band (e.g., via a sticker) may have a transparent casing polymer to allow measuring blood vessel on the wrist therethrough. The sensing unit may include at least one light source and at least one light sensor (e.g., positioned 2-3 mm next to each other) and positioned perpendicular to a wrist blood vessel. In some embodiments, the light source may detect the position of the blood vessel (using strongest absorption measurement analysis), for instance using at least one of visible and NIR light. 
     In some embodiments, the at least one sensor may be mounted on a transparent plastic polymer in a rotatable dial with an open slit of 4-6 mm wide along its diameter axis allowing a good view of the wrist blood vein from a top view and allowing positioning by rotation of the dial with sensors and light sources in such a way that the light sources and sensors may be aligned above a wrist blood vessel. In some embodiments, the sensors may have an external ring that is adhered to the skin around the blood vessel. In some embodiments, the measurement unit is battery operated and has a wireless communication module. 
     It should be noted that analysis (e.g., with a processor such as processor  301  shown in  FIG. 8 ) of such measurements may allow real-time quantification of changes in concentration of anesthetic drugs in blood during anesthesia. In some embodiments, several multivariate classification methods may be used, including partial-least square (PLS), principal component regression (PCR), linear discriminant analysis (LDA), k-Nearest Neighbors (kNN), Naive Bayesian classifier (NBC), Support Vector Machine (SVM), Artificial Neural Networks (ANN), etc. With a linear approximation, the Propofol concentration (in g/mL, mol/mL, etc.) may be considered as the dependent variables (Y), and the rest of the variables, for instance data matrices of different measurement units are the independent variables (X). The purpose of a multivariate classification (e.g., multiple regressions) may be to find a solution that best predicts the ‘Y’ variable as a linear (or non-linear) function of the ‘X’ variables. 
     For example, the general equation for the multiple regression may be: 
             Y   =         b   0     +       b   1     ⁢     x   1       +     …   ⁢           ⁢     b   k     ⁢     x   k       +   f     =         b   0     +       ∑     i   =   1     k     ⁢       b   i     ⁢     x   i         +   f     =       b   0     +   f   +   BX               
where ‘Y’ is the response (for Propofol concentration), ‘x k ’ are the predictors (e.g., data matrices of different measurement units), ‘b k ’ are the regression coefficient to be determined, b 0 ′ is the offset and a constant factor, and ‘f’ is the residual. If ‘X’ and ‘Y’ are mean-centered, then b 0 =0. Thus, the above equation may be written in matrix form: Y=Xb+f.
 
     According to some embodiments, variety of spectral regions in different measurement units (e.g., transmission and/or reflectance and/or FTIR-ATR) may be examined separately and/or in combination to obtain optimal classification accuracy. 
     According to some embodiments, a measurement unit may further include at least one of a display (e.g., a touchscreen displaying patient data), a power supply unit, a patient data unit (e.g., with oxygen saturation, blood pressure, and heart rate), an anesthetic drug monitor (ADM) clip, and an anesthesia degree analysis unit (e.g., collects in real time the patient data and the ADM data and gives a new combined value (deep of anesthesia value, DAV). In some embodiments, the display may include a Capnometer to display the capnogram, respiratory rate, and end-tidal CO 2  (EtCO 2 ). In some embodiments, the display may include a pulse Oximeter to display the plethysmogram and oxygen saturation. The Capnometer and pulse Oximeter may operate in tandem to detect indicators of over sedation, such as low respiratory rate, apnea, and low oxygen saturation. 
     In some embodiments, the display may include an electrocardiogram to display the electrocardiogram and heart rate. In some embodiments, the display may include anon-invasive blood pressure display with last-measured systolic and diastolic blood pressure. In some embodiments, the display may include an ADM to assess the level of sedation by measuring patient Propofol blood level. In some embodiments, the display may include a deep anesthesia value (DAV) to calculate the level of sedation after collecting all the patient data and Propofol level in the blood. 
       FIG. 9  shows an example graph of detection of injected substance over time, in accordance with an embodiment of the present invention. The graph in  FIG. 9  was obtained by assembling a light sensor on a blood vessel of a wrist of a patient while Propofol was injected intravenously by a target controlled injection (TCI) syringe pump. The ‘X’ axis is a time scale and the ‘Y’ axis is the relative absorption. Both injection timing and continuous measurement with SWIR sensor were carried out on the same time scale. The data log provided by the syringe pump demonstrates time of injection  401  commencement and finish. 
     The graph shown in  FIG. 9  may be obtained with a measurement unit  300  as shown in  FIG. 8 . As may be appreciated by one of ordinary skill in the art, after injection  401  of a drug (e.g., injecting 30 milligrams of a predetermined drug) into the bloodstream, a peak  402  in absorbance (e.g., a peak of about 50%) may be observed in less than a minute. The graph shows a typical response of the system at a given wavelength on Propofol injections. The injections are accompanied by a sharp increase of the signal with subsequent trend. The amplitude of peaks  402  and the form of trend (or rate of change) depend on the concentration of the injected Propofol. 
       FIG. 10  shows an example of a transmission measurement unit  500 , in accordance with an embodiment of the present invention. Transmission measurement unit  500  may include an attachable body  505  (e.g., a clip, such as attachable segment  305  shown in  FIG. 8 ) configured to attach onto an outer ear portion  62  of a patient. In some embodiments, attachable body  505  may have a pressure mechanism configured to control pressure applied by the attachable body  505  (e.g., with elastic material), for example a user may reduce the pressure if the patient experiences pain or discomfort. 
     Transmission measurement unit  500  may further include, for instance embedded within the attachable body  505 , a processor  501 , a light source  502  and a light detector  503  coupled therebetween such that processor  501  may analyze measurement data received at light detector  503  for light transmitted through outer ear portion  62  from the light source  502 . In some embodiments, at least one of light source  502  and light detector  503  may be coupled to optical elements (e.g., lens, such as optical elements  306  shown in  FIG. 8 ). 
     In some embodiments, transmission measurement unit  500  may further include a communication module (e.g., for Bluetooth communication with an external device, such as communication module  308  in  FIG. 8 ) and/or a motion sensor  504  embedded within the attachable body  505 . In some embodiments, the attachable body  505  may include a contact portion  515  (e.g., with elastic material) configured to contact the outer ear portion  62  of the user, where light source  502  may be embedded within the contact portion  515 . 
       FIGS. 11A-11C  show a reflectance measurement unit  600 , in accordance with an embodiment of the present invention.  FIG. 11A  shows a top view,  FIG. 11B  shows a side view and  FIG. 11C  shows a cross-section of a perspective view. Reflectance measurement unit  600  may be removably attached (e.g., with a bio-adhesive strip) onto a body portion of the patient where a blood vessel is visible through the skin such as the arm or the skull for infants to allow optical reflectance measurements of predetermined substances (e.g., Propofol) in the blood stream. 
     Reflectance measurement unit  600  may include a hollow base  601  (e.g., disc shaped) removably attachable to the patient or to patient&#39;s tissue (e.g., with a bio-adhesive strip or other suitable attachment device), and coupled to a clipping element  602  and a transceiver portion  603 . In some embodiments, base  601  may be an adhesive strip. In some embodiments, clipping element  602  may include optically transparent material such that light may pass therethrough, for instance light beams may pass into a hollow center  604  of base  601  to penetrate skin tissue  22  of the patient. In some embodiments, the user may view skin tissue  22  through hollow center  604  and thereby position reflectance measurement unit  600  over a blood vessel. In some embodiments, transceiver portion  603  may include at least one indicator  609  configured to indicate to the user at least one of reflectance measurement unit  600  operation and positioning over a blood vessel (e.g., detecting a change in reading when positioned over a blood vessel). 
     In some embodiments, clipping element  602  may be manually moved (e.g. in at least one direction) to allow the user to view skin tissue underneath (e.g., via hollow center  604 ) and thereby position the reflectance measurement unit  600  over a blood vessel of the patient. In some embodiments, clipping element  602  may be manually rotated  611  around the hollow center  604  and/or clipping element  602  may be manually moved along axis ‘X’ along arms  607  of the clipping element  602 . In some embodiments, arms  607  of the clipping element  602  (e.g., including elastic material) may be gripped by a user to release/attach clipping element  602  to base  601  using abutting elements  608 . For example, the user may push arms  607  towards transceiver portion  603  and thereby release clipping element  602  from base  601  so as to allow relative movement (e.g., manual rotation of clipping element  602  relative to base  601 ). 
     In some embodiments, transceiver portion  603  may be configured to emit light beams towards the skin of the patient with at least one light source  605  (e.g., such as light source  302  in  FIG. 8 ) and detect light reflected back with at least one light sensor  606  (e.g., such as light sensor  303  in  FIG. 8 ). In some embodiments, transceiver portion  603  may further include a processor  610  configured to operate the at least one light source  605  and at least one light sensor  606  and analyze the received data to determine concentration of substances (e.g., Propofol) in the blood stream. It should be noted that while at least one light source  605  and at least one light sensor  606  are schematically shown at a distance from the plane of base  601 , in some embodiments at least one light source  605  and at least one light sensor  606  may be at the plane of base  601  so as to allow contact of skin tissue with base  601 , at least one light source  605  and at least one light sensor  606 . 
     In some embodiments, transceiver portion  603  may further include at least one of motion sensor (e.g., such as motion sensor  304  in  FIG. 8 ) and a communication module (e.g., via Bluetooth, such as communication module  308  in  FIG. 8 ) so as to communicate with an external device. 
     According to some embodiments, transceiver portion  603  may communicate (via the communication module) with the transmission measurement unit  500  (e.g., via Bluetooth) to compare and/or analyze combined measurements, where measurements by transceiver portion  603  and transmission measurement unit  500  may be carried out simultaneously. 
     In some embodiments, transceiver portion  603  may further include a visible light switch (e.g., manually operated by the user) such that during illumination of visible light (e.g., wavelengths of 400-900 nm), signals reflected from the skin may be read with visible light. In case of minimal reflectance detected by the at least one light sensor  606 , for instance due to maximal absorbance, positioning of transceiver portion  603  over a blood vessel may be determined. Once positioning of transceiver portion  603  over a blood vessel is determined, measurements in other wavelengths (e.g., IR, SWIR, etc.) without the visible light may be carried out in order to determine concentration of substances in blood. 
       FIGS. 12A-12B  show a flowchart for a method of noninvasive analysis of tissue, in accordance with an embodiment of the present invention. In some embodiments, a surface of the tissue may be irradiated  701  (e.g., with at least one source of infrared radiation) with infrared radiation in a first spectral band that is strongly absorbed by water, and with infrared radiation in a second spectral band such that an interaction of the radiation in both spectral bands with a component of the tissue other than water may be substantially identical. In some embodiments, an intensity of the radiation may be measured  702  (e.g., with at least one radiation detector) that emerges from the tissue in each of the spectral bands. 
     Change in at least one of shape and intensity of signals received by the at least one radiation detector may be determined  703  (e.g., by a processor). A relative absorption by the tissue of radiation may be calculated  704  (e.g., by the processor) in one of the first and second spectral bands relative to absorption by the tissue of radiation in the other of the first and second spectral bands. Concentration of a predetermined substance may be determined  705  (e.g., by a processor) in accordance with the calculated relative absorption and in accordance with determined change in the received signal. The interaction of the radiation may include at least one of absorption and scattering. Other or different operations may be performed. 
     Unless explicitly stated, the method embodiments described herein are not constrained to a particular order in time or chronological sequence. Additionally, some of the described method elements can be skipped, or they can be repeated, during a sequence of operations of a method. 
     Various embodiments have been presented. Each of these embodiments may of course include features from other embodiments presented, and embodiments not specifically described may include various features described herein.