Biometric signal management devices and methods

Abstract

A tissue sensor with a housing, a power source, a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength, a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength, a light detector positioned within the interior of the housing and optically isolated from the first light emitter. The sensor is operable to obtain a timed sampling of light absorption from a tissue with both an increased or decreased absorption of light centered at 685 nm + / - 10 nm relative to an increased or decreased absorption of light centered at 850 nm + / - 10 nm.

Description

BIOMETRIC SIGNAL MANAGEMENT DEVICES AND METHODS CROSS-REFERENCE

[0001] This application claims the benefit of U.S. Provisional Application No. 63 / 578,029 filed August 22, 2023, entitled BIOMETRIC SIGNAL MANAGEMENT DEVICES AND METHODS which application is incorporated herein in its entirety by reference.BACKGROUND

[0002] Field: Disclosed are sensor devices, systems, and methods for monitoring and managing biometric information of the human body.

[0003] Background: Historically, clinical methods of monitoring oxygen, for example, are limited to indirectly measuring the oxygen saturation of blood hemoglobin by pulse oximetry (SpCh), which indicates the supply of oxygen in the blood, with the assumption that if the blood level of oxygen is within the 'normal range,’ vital organ tissue will be safely and effectively supplied with oxygen to meet the cellular need for oxygen. Recently discovered sensor systems for detecting novel biometric information are an improvement over historical methods and have been described in, for example, US 10,638,960 to Hatch for Optical Physiologic Sensor Devices, US Patent 11,278,221 to Hatch for Optical Physiologic Sensor Devices, US Patent 11,426,093 to Hatch for Energy Conversion Monitoring Devices, Systems and Methods, and US Patent 12,004,856 to Hatch for Optical Physiologic Sensor Devices and Methods.

[0004] What is needed are devices and methods for optimizing acquired signals obtained from optical physiologic sensors, such as those described in the ‘960, ‘221, ‘093 and ‘856 patents and management of data produced by those sensors.SUMMARY

[0005] Described are devices and methods for optimizing acquisition of signals obtained by optical physiologic sensor devices. Also described are methods and systems for management of the produced data, and performing a physiologically relevant correlation of the information with existing biometrics. Absorption of light by tissue (such as skin tissue) and blood in the red and infrared portions of the light spectrum provides biometric signal variations that can be used both separately and together. As described in the 960, ‘221 '093 and ‘856 patents, timed sampling of spectral absorption by the skin, optimally centered at 685 nm (Red) and at 850 nm (Infrared), indicates variations in the rate of cellular energy7conversion. Cellular hypoxic and cellular hyperoxic stress are both now detectable as respective changes in absorption at 685 nm and at 850 nm.

[0006] Heartbeat-induced light signal pulsations at 660 nm and 940 nm are produced by tiny changes in the quantity of blood in the fingertip with each heart contraction cycle. The peak and trough values of these photopl ethy smogram (PPG) signals are commonly used in pulse oximetry' to select only the arterial blood signal to measure arterial blood oxygen saturation (SpCh) and to measure heart rate.

[0007] Elsewhere in the body, timed sampling (e.g., 1 Hz) of light intensity also produces useful information. Differential variation in absorption of 685 nm, vs., 850 nm light by the skin has also been found to indicate physiologic stresses. Progressively increasing absorption of 685 nm light, possibly combined with decreased absorption of 850 nm light, appears to indicate interference with skin cellular oxygen delivery’, such as from vasoconstriction in the skin in response to hypoxic stress from exertion, low breathing gas oxygen partial pressure (ppCh). vasoconstriction at the onset of hemodynamic or septic shock, or breathing air contaminated with carbon monoxide (CO). Conversely, the onset of progressively increasing absorption of 850 nm light, combined with decreased absorption of 685 nm light, appears to indicate potentially harmful cellular hyperoxia.

[0008] Tandem variations in timed sample values provide indications of breathing rate and of the effort of each breath. Normal, unrestricted breathing produces tandem variations of 1-2% full scale (FS) in signal intensity, which are more prominent at 850 nm, while obstructed or restricted breathing produces amplitude cycles of as much as 15% FS. Periodic breathing, due to unstable breathing regulation, consists of 20 to 50 second-duration cycles of variation in breathing rate and breath effort / volume during sleep. Periodic breathing becomes abnormal when it results in corresponding cycles of cellular hypoxic stress, as can be detected in the skin, as described herein. Abnormal periodic breathing has been associated with sudden infant death syndrome (SIDS) and sudden unexplained infant death (SUID) and may be lethal to vulnerable infants if it devolves into prolonged apnea, respiratory arrest, and death. Abnormal periodic breathing can be detected in adults during sleep as cycles of increased absorption by the skin at 685 nm without corresponding changes in absorption of 850 nm light. Body’ orientation relative to gravity and acceleration of the body, both of which affect the quantity of blood in the skin at the sensor site, also produce identifiable tandem signal variation signatures. Rostral shift (towards the head) of blood volume after reclining progressively increases light absorption by the skin of the upper arm over the first hour or two of sleep. Conversely, acceleration of the body along the head-to-trunk axis (+GZ), such as occurs with pilots of high-performance aircraft during extreme flight maneuvers, reduces orhalts blood flow to the brain, potentially risking G-induced loss of consciousness (G-LOC) of the pilot. Excess Gzalso decreases the quantity of blood in the skin of the pilot’s forehead or cheek, which is detectable as a large decrease in the absorption of light. Combined with these potentially useful signal responses that are described in greater detail herein are more random signal disturbances caused by variation in sensor contact pressure against the skin. Methods of isolating the physiologically useful biometric information from the mixed signals have been explored by experimental observations to provide a scientific and engineering design basis for exemplary applications.

[0009] Both the foregoing general description and the following detailed description are exemplary and explanatory only and are not restrictive of the disclosed embodiments, as claimed.INCORPORATION BY REFERENCE

[0010] All publications, patents, and patent applications mentioned in this specification are herein incorporated by reference to the same extent as if each individual publication, patent, or patent application was specifically and individually indicated to be incorporated by reference.

[0011] US 10,638,960 for “Optical Physiologic Sensor Methods” to Hatch issued May 5, 2020;

[0012] US 11,278,221 for “Optical Physiologic Sensor Devices” to Hatch issued March 22, 2022;

[0013] US 11,426,093 for “Energy Conversion Monitoring Devices, Systems and Method” to Hatch issued August 30, 2022;

[0014] US 12,004,856 for “Optical Physiologic Sensor Devices and Methods” to Hatch issued June 11, 2024;

[0015] Kalogeris, et. al, “Ischemia / Reperfusion,” Comprehensive Physiology 7: 113-170, 2017, PMID: 28135002;

[0016] Ramanathan, et. al., “Cardiorespiratory' Events Recorded on Home Monitors,” JAMA, 2001285; 2199-2207;

[0017] Reck-Peterson, et. al, “The Cytoplasmic Dynein Transport Machinery and its Many Cargoes, ” Nat Rev Mol Cell Biol. 2018 June; 19(6): 382-398. doi: 10.1038 / s41580- 018-0004-3;

[0018] Redmond, “Transthoracic Impedance Measurements in Patient Monitoring,” Analog Devices, 2013; and

[0019] Rozova, et. al., “Structural and dynamic changes in mitochondria of rat myocardium under acute hypoxic hypoxia: role of mitochondrial ATP-dependent potassium channel,” Biochemistry (Moscow) 2015, Vol. 89, No. 8, pp. 994-1000, PMID: 26547067.BRIEF DESCRIPTION OF THE DRAWINGS

[0020] The novel features of the invention are set forth with particularity in the appended claims. A better understanding of the features and advantages of the present invention will be obtained by reference to the following detailed description that sets forth illustrative embodiments, in which the principles of the invention are utilized, and the accompanying drawings of which:

[0021] FIGS. 1A-D illustrate an exemplar sensor device;

[0022] FIG. 2 is an exemplar schematic architecture for the sensor monitoring device system;

[0023] FIG. 3 is a graph that illustrates net spectral intensity during ischemia (extreme cellular hypoxia);

[0024] FIG. 4 is a graph that illustrates net spectral intensity during reperfusion and cellular hyperoxia;

[0025] FIGS. 5A-5B are graphs that illustrate Cellular Energy Index and raw data during ischemia and reperfusion;

[0026] FIGS. 6A-6C are graphs that illustrate Cellular Energy Index and raw data during two airline flights;

[0027] FIG. 7 is a graph that illustrates rostral shift and body position-induced variations in raw signals;

[0028] FIGS. 8A-8B are graphs that illustrate comparison of thoracic electrical impedance and optical monitoring of breathing;

[0029] FIGS. 9A-9B are graphs that illustrate Cellular Energy Index (CEi) trend and raw data during severe obstructive and central apnea;

[0030] FIGS. 10A-10B are graphs that illustrate CEi and raw data from periodic breathing during sleep:

[0031] FIGS. 11A-11D are graphs that illustrate CEi trend, Cabin Barometric Altitude, raw data, and G-meter data during aerobatic flight maneuvers;

[0032] FIGS. 12A-12C are graphs that illustrate theoretical CEi and raw data trends during CO toxicity;

[0033] FIGS. 13A-13C are graphs that illustrate theoretical CEi trend during current initial breathing therapy for newborn infants;

[0034] FIGS. 14A-14C are graphs that illustrate theoretical CEi trend during proposed monitor-guided initial breathing therapy for newborn infants; and

[0035] FIG. 15 is a table summary of sensor data changes and physiologic correlations.DETAILED DESCRIPTION

[0036] An exemplar sensor 100 suitable for detecting and monitoring the disclosed biometrical signals is illustrated in FIG. 1A. An exterior surface 108 of a sensor aperture component 107 comes in contact with the patient's skin. A light emitter aperture opening 110 and a light detector aperture opening 112 are provided which can be filled with clear plastic potting to present a flat surface to a tissue surface such as patient’s skin, when applied. A plastic resin encapsulation 114 housing hermetically encloses and mechanically protects the internal components of the sensor 100 device.

[0037] FIG. IB shows internal components of the sensor 100 shown in FIG. 1A. As shown, the sensor aperture component 107 has a length, height and width and is adhered to an end of a flexible circuit 116. A component volume space for surface mount technology (SMT) components is illustrated which is mounted on a main printed circuit board (PCB) portion of the main circuit 118. The flexible circuit 116 is configurable to fold over the SMT components. Additional electronic components, such as a BluetoothLEKtransceiver 120, and, optionally, a vibrator motor (not shown) for haptic stimulation, can also be mounted to the main circuit 118.

[0038] FIG. 1C shows the sensor aperture component 107 removed from over the light emitting diodes (LEDs) 124 and 126 and detector 128, such as a photodiode detector. An interior surface 109 of the sensor aperture component 107 is shown. The sensor aperture component 107 has at least one internal wall 122 that, when secured to the flexible circuit 116, provides an optical barrier between a first chamber 140 and a second chamber 142. The first chamber 140 (light emitter cavity) and the second chamber 142 (detector cavity) define three-dimensional (3D) recesses on the interior surface 109 of the sensor aperture component 107. Use of metal for the sensor aperture component 107 can be used in one configuration. However, as will be appreciated by those skilled in the art, a plastic material that is sufficiently opaque to both red and infrared light could also be used to form the sensoraperture component 107, including the internal wall light barrier. Additionally, the sensor aperture component 107 can be attachable, e.g. with a suitable adhesive, such as epoxy.

[0039] A red LED 124, an infrared LED 126, and a detector 128, such as a silicon photodiode, are mountable on the flexible circuit 116. LED power control and signal amplifier components are mounted on the underside of the sensor portion of the flexible circuit 116. The flexed portion 130 of the flexible circuit 116 carries circuit traces from the sensor to the main circuit 132.

[0040] As will be appreciated by those skilled in the art, the actual layout can take a variety of forms without departing from the scope of the disclosure. In some configurations, the lead pattern of the microcontroller can have a pattern that radiates in a plurality of directions. Thus, if the LEDs and sensors are placed in the middle, the microcontroller leads would need to be routed around the LEDs and sensors. By placing the optical components on a flexible circuit, as illustrated, a simplified layout is achieved. In some configurations, the Bluetooth transceiver may have an associated microcontroller that is capable of performing the optical component control, which would allow a sensor positioned in the center.

[0041] FIG. ID illustrates the underside of the main circuit 132, where a battery 134 and battery recharging coil 136 are located with the flexed portion 130 and the optical component portion of the flexible circuit 116 curving around an end of the main circuit 132.

[0042] FIG. 2 is a schematic block diagram of the sensor 200 on a test subject's skin, including the Cellular Energy Monitor (CE monitor) sensor system 202 and a computer- implemented software application (“app”) 206 operating on a smartphone 204. Sensor control commands and data transfer 208 between the CE monitor sensor system 202 and the smartphone 204 are managed via RF transceiver 210 in the CE monitor sensor system 202 and the BluetoothLE® RF transceiver and application program 206 operating in either an iPhone® or Android® OS smartphone, or equivalent. The light emission electro-optics components of the sensor system are enclosed within the visible and infrared light-opaque metal or plastic sensor aperture component 107 and housed within a first chamber 140 in FIG. 1C housing a red LED 124 and an infrared LED 126. The red LED 124 and infrared LED 126 can further be encased together in clear plastic; thus, providing a path for light emission from the LEDs through a light emitter aperture opening 110 in FIG.1A.

[0043] The sensor aperture component 107 has a second chamber 142, which houses a detector 128, such as a silicon photodiode, and can also be filled with optically clear plastic. The housing of the sensor aperture component 107 has an internal wall 122 positionedbetween the first chamber 140 and the second chamber 142. which divides the two chambers and blocks internal transmission of the light from the red LEDs 124 and infrared LEDs 126 to the detector 128. Thus, any light detected by the detector 128 occurs after the light has been transmitted and diffused through the skin.

[0044] The CE monitor sensor system 202, when in use, is placed on the skin 224 of the patient. Emitted light 226 and 228 passes through the epidermis and dermis 230 of the skin, but does not functionally penetrate the superficial facia or underlying muscle 234.

[0045] Top level operational commands are conveyed from the CE monitor sensor system 202 via data transfer 208, such as an RF transmission achieved with a BluetoothLE® transmission, from the application program 206, such as a smartphone app, to initiate sensor function. A sensor microcontroller 240, establishes an RF linkage procedure with the application program 206, then performs a sensor initialization procedure to establish the optimum power levels delivered to the LEDs. Upon completion of the initialization process, the sensor microcontroller 240 performs timed sampling in data acquisition cycles at time intervals set by the application program 206, which is 1 Hz for the 685 nm and 850 nm LEDs for Cellular Energy Index (CEi) monitoring. During each CEi data acquisition cycle, an ambient light (i.e. not illuminated by either LED), a red LED-illuminated, and an infrared LED-illuminated detected light measurement are made and stored in the sensor memory 214. Current data is then communicated as raw data via the data transfer 208, such as an RF transceiver linkage, to the application program 206 for recording, display, and analysis.

[0046] An LED power amplifier 220 receives a digital to analog (D / A) control voltage from the sensor microcontroller 240. The LED power amplifier 220 then sends the corresponding current level to the red LED 124 and / or the infrared LED 126. The detector 128 is in communication with a signal amplifier 242 which is in communication with the sensor microcontroller 240. The sensor microcontroller 240 is also in communication with a sensor memory' 214 and an RF transceiver 210.

[0047] In another configuration, the CE monitor sensor 200 communicates directly or indirectly with an airway therapy device, such as an APAP machine or a CPAP machine. Data from the CE monitor sensor 200, provided directly or indirectly to the airway therapy device can be processed by the airway therapy device to result in a change of the administered airway therapy (e.g., air pressure). The direct submission of data can be wired or wireless. The indirect submission of data can be via a secondary device, such as a smartphone, or via a remote central station that processes the data received from the CEmonitor sensor prior to delivering the processed data or machine operation commands to the airway therapy device.

[0048] Testing of the sensor

[0049] While not used in pulse oximetry, the average (i.e., non-pulsatile, or DC) variations in spectral absorption of light by skin have also been found to respond to physiologic stress in robust and potentially useful ways. Identifying the information obtainable from these DC signal variations involved exploration of the wavelengths of light used and of how know n physiologic stresses and pathologic states affect the detected signals.

[0050] Following the initial observation of unexpectedly large, differential changes in DC absorption of light by skin at 660 nm and at 810 nm in response to briefly breathing nitrogen, a test fixture was created to help identify the optimum infrared wavelength region. The test fixture had four independent sensor channels, all of which used 660 nm LED light (Red). Four different wavelengths of Infrared LED light, 790 nm, 810 nm, 835 nm, and 850 nm, were each paired with a 660 nm LED. The skin absorption signal responses at these wavelengths were compared while the subject breathed nitrogen, air, or oxygen. The most responsive infrared region was found to be centered at about 850 nm.

[0051] An armband-wearable sensor such as described in FIGS. 1 and 2 and further described in the ‘960, ‘221, ‘093 and ‘856 patents was built to record spectral absorption responses during normal life activities. The armband sensor detected absorption responses in upper arm skin at 660 nm (Red) and 850 nm (Infrared) data at 1 Hz sampling rate and digitized at 12-bit resolution. Several adult athletes performing interval exercise for up to 1 hour were recorded. Normal sleep breathing and sleep disordered breathing (SDB) were recorded during sleep at home and during diagnostic sleep lab testing sessions with adult subjects, encountering obstructive and central apnea, periodic breathing, and primary snoring. All diagnostic features of SDB were found to be detected by the wearable sensor as well as, or better than, the pulse oximeter and breathing sensors of the sleep lab instruments. Multiple commercial airline flights by adults were recorded. Unpressurized aircraft flights, including flights at up to 15,000 feet elevation, were recorded with the sensor on the test subject’s upper arm. During flights over 10,000 feet the pilot wore an oxygen mask. Aerobatic aircraft flights and aircraft maneuvers have also been recorded with the sensor worn on the forehead of a pilot while flying to over 13,000 feet, without supplemental oxygen, while encountering an augmented force of gravity (“+GZ”) of up to +5 Gzfor 3 seconds, and during sustained +3 Gzbanked turns for up to 20 seconds.

[0052] Further exploration was performed using broadband light between 600 nm and 1 100 nm projected via optical fiber into the skin, with spectral detection of laterally diffused light using a fiberoptic spectrometer (Ocean Optics, Flame-S) while the subject breathed nitrogen, air, or oxygen. Recordings were then made during exercise and during up to 6 minutes of ischemia (halted blood flow to the forearm), followed by abrupt reperfusion of the forearm skin. The most active regions of variation in spectral absorption by skin tissue in response to these changes in cellular oxygen supply were found to be centered at 685 nm + / - 10 nm (Red) and at 850 nm + / - 10 nm (Infrared).

[0053] Physiologic and biochemical correlations with absorption of 685 nm light by the skin

[0054] Absorption of 685 nm light by the skin, during a timed sampling of detected light intensity at 1 Hz, appears to be robustly sensitive to the activity rate-status of oxygen supply-dependent cellular energy conversion chemistry, i.e., rate of ATP production. Progressively increasing absorption of 685 nm light combined with decreased absorption of 850 nm light is associated with decreased oxygen delivery- to the skin.

[0055] Specifically, skin absorbs more 685 nm light when the test subject’s skin receives less than the current cell-adapted level of oxygen supply, potentially over 10% FS lower than normal detected 685 nm intensity values. Conversely, the skin absorbs less 685 nm light when the subject receives more than the current cell-adapted level of oxygen intake (e.g., breathing up to 100% oxygen), potentially resulting in more than 10% FS higher than normal detected 685 nm intensity values. The heartbeat cycle produces tandem pulsatile variations (e.g., PPG) in detected 685 nm and 850 nm light after diffusion through skin. Breathing also produces tandem cyclic variations in detected signal intensity at these wavelengths. Sensor motion and variation of pressure of the sensor against the skin produces immediate, similarly large tandem variations in absorbance of 685 nm and 850 nm light, apparently due to pressure-induced changes in the amount of light-absorbing blood in the skin. When the sensor is worn on the pilot’s forehead during episodes of elevated Gz, the resulting decreased amount of blood in the facial skin greatly increases detected signal intensity at both wavelengths.

[0056] Observation of 685 nm light absorption during ischemia / reperfusion is especially remarkable. In an exemplary demonstration, an upper arm pneumatic cuff is quickly inflated to 300 mmHg pressure, halting blood flow (i.e., inducing ischemia) to and from the subject's forearm up to 6 minutes. During this ischemia period, there is large-scale,progressively increasing absorption of 685 nm light by the forearm skin that continues to increase (i.e., decreasing detected intensity of 685 nm light) during the ischemic period. Physiologically and spectroscopically, this large, continuously increasing absorption cannot be explained as due to a change in the quantity of blood in the skin or to progressive desaturation of the tiny quantity of blood in skin capillaries during the arrested blood flow. Further, blood oxygen saturation does not change in arteries or veins during arrested blood flow due to the thick walls of these vessels. The continuing increase in 685 nm light absorption appears to be a function of decreasing rate of cellular energy7conversion, i.e., decreasing rate of production of ATP, as the dissolved oxygen within cells and surrounding interstitial fluid is consumed by cellular energy conversion chemistry. During the first 10 - 25 seconds following abrupt return to normal blood flow in the forearm following rapid cuff deflation, the absorption of 685 nm light in forearm skin rapidly decreases to up to 30% less than prior to the ischemic period. This phenomenon also cannot be due to a change in the quantity of blood in the skin or to a change in the oxygen saturation of arterial blood: neither of which occur during and following the ischemic period. It is apparent from the signal intensity rebound during reperfusion that skin cellular adaptation (e.g., mitochondrial movement toward the cell membrane) in response to extreme hypoxic stress had occurred during the ischemic period. The degree of adaptation becomes evident upon abrupt return to normal blood oxygen supply, resulting in a spectral absorption profile that is similar to the profile produced by the subject breathing oxygen.

[0057] Physiologic correlations with absorption of 850 nm light by the skin

[0058] Absorption of 850 nm light by the skin has been found to be unaffected by decreased oxygen supply . However, a threshold to progressively increasing absorption at 850 nm does occur as the oxygen supply to the skin is increased above the acclimated range of oxygen supply, such as when the subject breathes oxygen. The increased absorption of 850 nm light during and following excessive oxygen intake appears to be the result of relatively permanent biochemical reactions in skin tissue and requires up to several hours to return to baseline level following return to normal oxygen intake (e.g., breathing air) following a period of excess oxygen (e.g., 100% oxygen) intake. This threshold to increasing absorption at 850 nm may mark the upper limit of excessive cellular oxygen exposure, and potentially the onset of cellular injury. Monitoring for and detecting this biometric phenomenon, in conjunction with decreased absorption at 685 nm, may have a wide variety of applications. The 850 nm absorption signal threshold may also correlate with the onset of increasedleukoc le / endothelial adhesion in capillaries and venules, which may result in microvascular occlusion by leukocytes and prolonged hypoxic stress to the served tissue. ‘Hypoxia-like’ tissue necrosis pathology following rapid increase in oxygen supply has been found to start with microvascular occlusion by leukocytes. Associated pathologies include ischemia / reperfusion injury' (IRI) during reperfusion therapy for ischemic heart attack and ischemic stroke, during reperfusion of transplant organs, and during resuscitation following cardiac and / or respiratory arrest. Vital organ ‘hypoxia-like’ injuries also become evident shortly after birth of premature newborn infants who are initially supplied with air or higher oxygen content breathing gas, i.e., too-rapidly increasing the supply of oxygen compared with the oxygen supply during fetal life before birth.

[0059] Light at 850 nm is up to about 2% FS more absorbed by blood in the skin than is 685 nm light. As a result, heart rate pulses, breathing cycles, sensor motion, and episodes of elevated +GZproduce more pronounced detected intensity changes at 850 nm. There is a potentially useful signal variation at 850 nm with breathing where breath-related cycles of decreased light absorption by the skin are more pronounced when there is increased breathing effort to overcome upper airway restriction or obstruction during sleep. Normal sleep breathing produces 1 - 5% FS variations in detected intensity with each breath at 850 nm. When there are no breaths, such as during central apnea, there are no changes in signal intensity correlating to breathing effort.

[0060] While a sensor can be placed on different areas of the body, placement of the sensor on the upper arm using an armband has been observed to optimize detection of vasoconstriction in the skin as the first physiologic response to mild cellular hypoxic stress, such as during exercise and during primary snoring during sleep. Primary snoring during sleep is identified during sleep apnea testing as audible snoring that does not result in at least 3-4% decrease in SpCh by pulse oximetry. Vasoconstriction in the skin increases detected 850 nm intensity because vasoconstriction reduces the amount of light- absorbing blood in the skin. Vasoconstriction in the skin also reduces oxygen supply to skin cells, resulting in cellular hypoxic stress and increased absorption of 685 nm light. Thus, it appears that vasoconstriction in the skin is a normal, initial adaptive reflex response that helps to conserve the blood oxygen supply for vital organs during hypoxic stress. If the hypoxic stress becomes greater than can be compensated by vasoconstriction in the skin, such as with airway obstruction during sleep, pulse oximetry will begin to detect decreased SpCh.

[0061] Physiologic correlations with differential absorption of 685 nm and 850 nm light

[0062] Within the normal-to-deficient oxygen supply range, absorption of 850 nm light by the skin is unchanged: except by sensor motion, the heartbeat cycle, breathing effort, body orientation, vs., gravity, and level of Gz. Since these signal changes occur largely in tandem at 685 nm and at 850 nm, subtraction of the 850 nm light signal value from the detected 685 nm signal value cancels most of the tandem variation and enables clearer detection of the differential absorption due to variation in cellular oxygen supply. This calculation produces an analog numerical index of oxygen supply at the skin cellular level, referred to herein as Cellular Energy Index (CEi). At 12-bit A / D signal resolution, the CEi numeric range centers at zero, goes negative with cellular hypoxia and positive with cellular hyperoxia. The normal range of oxygen supply to the skin correlates with CEi between + / - 75. Extreme cellular hypoxia produces CEi values as low as -500, with extreme hyperoxia up to +500, as described in the ‘960, ‘221, ‘093, and ‘856 patents.

[0063] The CE monitor offers the new abil i ty to detect the threshold from cellular normoxia into cellular hyperoxia as the cellular oxygen supply is increased. Pulse oximetry and blood gas measurements cannot detect this cellular chemistry threshold because it is not a blood-based phenomenon. As the cellular oxygen supply is gradually increased, the cellular absorption at 685 nm gradually decreases, producing increasing intensity of the detected 685 nm signal. The cellular hyperoxia threshold is detected when absorption of 850 nm light begins to increase, producing decreasing intensity of the detected 850 nm signal. Regulating a gradual rate of increase in cellular oxygen supply to avoid the onset of decrease in 850 nm signal intensity offers a potentially breakthrough method of preventing microvascular occlusion by leukocytes and the resulting tissue necrotic injuries.

[0064] Correlation with existing biometric signals

[0065] Pulse oximetry and blood gas measurements mainly provide clinically useful insight into the gas exchange function of the lungs relative to changes in total body consumption of oxygen and excretion of carbon dioxide (CO2) by body tissues. The PPG signals used in pulse oximetry to produce blood oxygen saturation (SpO2) data have highest amplitude on the fingertip, due to the high density of pulsating arterioles in this location. Lower amplitude heartbeat-induced light absorption pulsations can also be found on other body surfaces, such as the chest, upper arm, and wrist.

[0066] Newborn infants whose ductus arteriosus has not closed may have a higher SpO2on their head, right chest, and right arm (pre-ductal), and lower SpO2over theremainder of their body (post-ductal). Pulse oximeter sensor motion against the skin distorts the PPG signals, resulting in errors in calculation of SpCty by the typical Logic AC ratio method. These calculation distortions due to sensor '‘motion artifact” can result in many false alarms and missed real alarms. Current methods of decreasing false alarms and decreasing missed real alarms introduce a delay in the displayed output data and in the alarm response. Finally, because blood oxygen measurements indicate only the status of the blood, the adequacy of vital organ cellular oxygen supply must be assumed; leaving a large gap in the information needed for accurate identification of root causes of harm and for optimal guidance of therapy.

[0067] Alternative clinical methods of monitoring breathing include: (1) thoracic electrical impedance that is typically combined with ECG heart rate monitoring, (2) chest and abdomen circumference-sensing respiratory impedance plethysmograph (RIP) belts, (3) nasal airflow sensors, (4) esophageal manometry7, and (5) air hose air pressure and flow in APAP, CPAP, anesthesia, and medical ventilator systems. In clinical sleep apnea studies by polysomnography (PSG), RIP belts are used to detect obstructed breaths that are not detectable by variation in thoracic electrical impedance. Inhalation attempts with an obstructed upper airway result in RIP belt detection of decreased chest circumference with increased circumference of the abdomen; the algebraic sum of the two sensor values being diminished. However, changes in body dimension with breathing cannot indicate breathing effort. PSG studies may also include nasal airflow7detection to provide signal variations analogous to the volume of air exchanged with each breath. As an objective indicator of breathing effort, esophageal manometry7measures the trans-thoracic pressure differential cycle, but is too uncomfortable and invasive for routine sleep studies or clinical monitoring. Modem anesthesia breathing systems, medical ventilators, and APAP and CPAP machines continuously monitor the pressure and air flow delivered to the patient, but this method also has difficulty discerning obstructed breaths from central apnea and cannot indicate breathing effort.

[0068] Turning now to FIG. 3, a graph of net spectral intensity7during ischemia 300 (extreme cellular hypoxia) is illustrated. The Y-axis value scale in the graphs in FIG. 3 and FIG. 4 are derived from the spectrometer’s 16-bit data, as described below7. Timed (1 Hz) data sampling show s the net change in spectral intensity after about 3 minutes of ischemia of forearm skin, which was induced by rapidly inflating an upper arm pneumatic cuff to 300 mmHg. This data was obtained from broadband light (QTH lamp) betw een 600 and 1000 nm(X-axis) that was projected into the forearm skin via an optical fiber. A second optical fiber located 8 mm from the projection fiber collected and returned diffused light from the skin to a spectrometer (Flame-S, Ocean Optics). The light source shutter and spectrometer were controlled, and data was recorded, by a custom LabVIEW (National Instruments) application program. Each spectral sample numeric array was processed, element by element, by subtracting from the sample array (1) a reference spectrum numeric array obtained prior to the ischemia and (2) an unilluminated spectral numeric array. If there is no change in light absorption by the skin within this spectral band, the resulting net spectrum graph would be a straight line along zero. The resulting net value change spectrum array after 3 minutes of ischemia is graphed 300. By reference, pulse oximeter center wavelengths, 660 nm 302 and 940 nm 304, are identified. Two wavelength regions, centered at 685 nm + / - 10 nm 306 and 850 nm + / - 10 nm 308, were selected based on this experimental data as being most differentially active through the ischemia / reperfusion cycle.

[0069] The devices in FIGS. 1-2 are operable to detect cellular hypoxia in skin or other tissue by detecting increased absorption of light centered at 685 nm + / - 10 nm relative to minimally changed absorption of light centered at 850 nm + / - 10 nm. Detecting cellular hypoxia in skin or other tissue allows detection of cellular hypoxia where the decrease in oxygen supply to the skin or other tissue is caused by a decreased oxygen available in breathing gas, decreased blood perfusion, or carbon monoxide (CO) bound to blood hemoglobin.

[0070] FIG. 4 is a graph illustrating net spectral intensity during reperfusion 400. The net value change spectrum graphed in FIG. 4 was obtained about 22 seconds following rapid release of the air pressure in the upper arm cuff. This data w as selected because it was the most changed from the data portrayed in FIG. 3. It is apparent that the pulse oximeter center wavelengths 402 and 404, which are widely accepted as optimum for measuring oxygen saturation of blood hemoglobin, were also differentially active in this experiment, but were less active than the selected regions 406 and 408. In context with the specific physiologic stress involved, this data demonstrates that the two selected spectral regions 406 and 408 are apparently associated with oxygen supply-related changes in molecular resonance at the cellular level, such as mitochondrial enzymes (e.g., rate of ATP Synthase activity ) involved in the production of ATP, rather than blood hemoglobin. The spectrum response pattern during reperfusion 400 is not a simple return back to baseline, and can be interpreted to indicate that a natural, cellular-level adaptive change (e.g., mitochondrialmovement toward their cell membrane) had occurred during the ischemic portion of the test. The increased absorption at 850 nm is especially interesting because of its similarity to the persistent spectral response to induced cellular hyperoxia induced by breathing oxygen. These new biometric insights support development of CE monitor-guided reperfusion and resuscitation therapy; potentially leading to improved outcomes.

[0071] The devices in FIGS. 1-2 are operable to detect cellular hyperoxia in skin or other tissue by detecting increased light centered at 685 nm + / - 10 nm in conjunction with detecting decreased light centered at 850 nm + / - 10 nm. As will be appreciated by those skilled in the art, the detection of cellular hyperoxia can be achieved where the increase in oxygen supply to the skin or other tissue is due to increased oxygen available in reperfusion blood. Tracking the raw and differential detected values due to absorption of 685 nm, vs., 850 nm light by the vital organ tissue can be used to guide a gradual increase in cellular oxygen supply during reperfusion therapy to prevent triggering leukocyte-endothelial adhesion and microvascular occlusion that would result in necrotic tissue injury to previously ischemic organ tissues. Similarly, the beginning oxygen level and the rate of rise in oxygen level of breathing gas during resuscitation therapy can be guided by tracking the differential absorption of 685 nm, vs., 850 nm light by the patient’s skin (i.e., a sensitive analog tissue) to prevent triggering leukocyte-endothelial adhesion and microvascular occlusion that would result in necrotic tissue injury to vital organ tissues. This is especially relevant during initial breathing support of premature newborn infants to prevent triggering leukocyte-endothelial adhesion and microvascular occlusion that likely results in necrotic injury to the infant’s eyes, brain, gut, and ductus arteriosus.

[0072] FIGS. 5A-5B is a graph illustrating Cellular Energy Index and raw data during ischemia and reperfusion 500. The 685 nm 506 and 850 nm 508 data were selected for graphic timeline displays FIG. 5A and FIG. 5B. The spectral data graphs in FIG. 3 and FIG. 4 were obtained at sample times 300’ and 400’ respectively. The Cellular Energy Index trend FIG. 5A was computed from the raw data FIG. 5B, sample by sample, by subtracting each 850 nm intensity value from its corresponding 685 nm intensity value. There is a continuous trend of change in the 685 nm data, resulting in the computed dow nward CEi data trend 504 during the induced ischemia period 506. Other similar experiments have recorded the 685 nm value continuing to decrease during up to 6 minutes of induced ischemia. These data cannot be logically explained as due to blood oxygen desaturation because: (1) the magnitude of change in signal intensity is far greater in than could be caused by desaturationof the tiny quantity of capillary’ blood within the light path, and (2) the 685 nm intensity continues to change. Capillary blood comprises less than 2% of the tissue volume in view, making it physically impossible for blood to be the source of the over 20% FS signal intensity change. If blood oxygen depletion in capillaries was the source of the signal change, the change in signal intensity' would also very likely be fully complete within a few seconds; not still incomplete after several minutes. Blood oxygen tension does not change in arteries or veins because these vessel walls are too thick to allow oxygen diffusion to surrounding tissue.

[0073] Graph FIG. 5B depicts the raw data from which the Cellular Energy Index trend portrayed in the graph shown in FIG. 5A was computed. Up to the point of inflating the upper arm cuff 508, the 685 nm 506’ and 850 nm 508’ data values were near equal, making their subtraction product near zero. Immediately upon inflating the cuff (at 508), the values of the 685 nm data 506’ begin a sharp, continuous decline 510, while the 850 nm data values 508’ remain relatively constant. Within seconds of deflating the cuff (at 512), there is a rapid increase in 685 nm data values 514 that overshoots the pre-ischemia 685 nm data value range by over 20% FS. The 850 nm data 508’ decreases briefly as the cuff deflates and blood flow resumes, then gradually returns toward the pre-ischemia value range. This divergence of data values upon reperfusion also cannot be explained as a change in blood oxygen, which, if due only to blood, would directly return to the baseline because the subject's SpC>2 did not change during the localized ischemia period. This combination of spectral responses can, however, be logically explained as due to cellular adaptation to extreme hypoxic stress during the ischemia phase, as evidenced by' a ‘hyperoxic’ cellular spectral signature upon reperfusion. This non-pulsatile, or DC spectral absorption response is excluded in the pulse oximetry calculation of SpCh. It also does not occur in laboratory spectroscopy of blood hemoglobin during rapid changes in blood oxygen content. However, it may offer a useful biometric indicator of the root cause of TRI and of vital organ injuries in premature newborn infants.

[0074] FIGS. 6A-6C are graphs illustrating Cellular Energy Index and Raw Data measurements taken during two airline flights. The mild hypoxic stress, cellular adaptation, and cellular hyperoxia that were recorded are insensible and are apparently harmlessly experienced by hundreds of thousands of airline passengers and crews daily. This flight originated in San Jose, California and terminated in Newark, New Jersey, with a change of aircraft in Chicago, Illinois. The pressurized cabin barometric altitude shown in FIG. 6C that was recorded on the flight to Chicago, was about 7200 feet above mean sea level (AMSL) (at 602), while the shorter second flight to Newark was about 6200 feet AMSL (at 604). Pulseoximetry would have registered about 3% decrease in SpCh during the cruise phase of each of these flights, due to the decreased partial pressure of oxygen (ppCh) of the cabin barometric altitude, with direct return to the subject’s baseline SpCh upon landing. The Cellular Energy Index in FIG. 6A shows a brief period of cellular hypoxia (at 606) during the climb to cruising altitude, followed by cellular adaptation back to baseline (at 608) during the remainder of the flight. That cellular adaptation to hypoxic stress occurs during airline flight, as it apparently does during ischemia in FIGS. 3-5, is revealed by the cellular hyperoxia response 610 during approach and following landing in Chicago. This latter response did not resolve by the time of the second flight 612, which, likewise induced relative hypoxic stress 614, followed by even more pronounced cellular hyperoxia 616 upon landing. Due to difficulties with data transmission with this prototype sensor, there is some loss of data (at 618 and 618’) during the subject’s somewhat strenuous activities following each flight.

[0075] Cellular hyperoxia in the skin has also been documented in the ‘960, ‘221 ‘093, and ‘856 patents such as from breathing oxygen, to produce an increased data value trend at 685 nm and decreased data value trend at 850 nm. Similar cellular hyperoxia phenomena are recorded just prior to and following landing on both flights 610, 614. Basic research on the mechanisms of cellular adaptation to changing oxygen tension reports that cellular adaptation to decreasing oxygen supply (e.g., mitochondrial movement within their cell toward the cell membrane) is a measurably faster process than cellular adaptation (e.g., mitochondrial movement toward the center of their cell) to rising oxygen supply. This, and the previous recordings of spectrometry data in FIGS. 3-5, also provide insight into cellular dynamics that appear to be relevant to several currently incompletely resolved major medical problems.

[0076] The raw data from which CEi is calculated is portrayed in FIG. 6B, which indicates hypoxic stress by decreased intensity' of 685 nm (at 620), vs., 850 nm (at 622) data during the two assents (at 624, and 626). FIGS. 3-6 provide new sensor-derived evidence of cellular adaptation to decreasing and to increasing oxygen supply, ranging from extreme hypoxic stress due to ischemia to harmless, insensible mild hypoxic stress experienced during airline travel. Pulse oximetry Logio ratio calculations to derive SpO? exclude these non- pulsatile, DC signal changes that are apparently due to cellular adaptation and that, as shown in FIGS. 3 and 4, also occur to lesser extent at 660 nm and 940 nm. Selection of 685 nm and 850 nm as the sensor center wavelengths, based on the spectrometer studies shown in FIGS. 3-5, increases the detected magnitude and resolution of these responses.

[0077] FIG. 7 is a graph illustrating rostral shift and body position-induced variations in raw signals 700. The data portrayed in the raw data graph were recorded during sleep with a prototype armband-wearable sensor that employs 660 nm and 850 nm LEDs as the light sources. The subject was subsequently diagnosed as having severe sleep disordered breathing (SDB), mostly involving abnormal periodic breathing. The rostral shift 702 that typically occurs over the first 1 to 2 hours following laying down is a gradual, gravity-driven redistribution of the total blood volume throughout the length of the subject’s body. Rostral shift 702 was also seen in FIG. 6 as the subject was seated and waited for an hour on the aircraft before the flight to Chicago took off. During upright hours during the day, most of the total blood volume is located within the lower trunk and lower extremities. This phenomenon has been a consistent finding in hundreds of CE monitor sleep recording sessions, and is a normal physiological phenomenon with little or no consequence during normal health. Persons suffering from heart pump failure, however, suffer from increased blood volume in their lungs upon fully reclining, resulting in increased breathing distress. Astronauts living in microgravity also suffer from the continuous increase in blood volume in their head due to lack of gravity to keep the blood mostly in the lower portion of the body during active living.

[0078] Also depicted in FIG. 7 are value changes 704 resulting from changes in body orientation relative to gravity . When the sensor is worn on the sleeping subject's arm positioned above the heart, there is relatively less blood in the skin beneath the sensor than when the sensor is on the arm beneath the heart. Thus, ‘roll-overs’ 704 are apparent in the raw data as abrupt tandem data value offsets, typically followed by brief periods of slower change 706 and 706’. This is presented as objective evidence of even subtle blood distribution phenomena, such as is also involved in elevating an injured extremity higher than the heart to reduce swelling.

[0079] FIGS. 8A-8B illustrate a conceptual comparison of thoracic electrical impedance and optical monitoring of breathing 800. FIG. 8A and FIG. SB represent theoretical comparative depictions of how breathing can be monitored. Commonly applied in intensive care monitoring of breathing, thoracic electrical impedance monitoring FIG. 8A is ty pically combined with multi-lead electrocardiogram (ECG) monitoring of heart rate and cardiac conduction. Variations of thoracic electrical impedance 802 occur as the lungs are inflated and deflated, changing the electrical resistance, or impedance, of the tissue betw een skin contact electrodes placed on the upper right and left anterior chest. The impedance of the skin contact electrodes ranges between 50 and 700 Ohms, and the adult body tissueimpedance ranges between 100 Ohms and 500 Ohms. Variations in impedance due to changes in the inflation of the lungs range between 0.2 and 5 Ohms. This method of monitoring breathing is widely accepted as effective within the near-normal range 804 of lung inflation. Limitations include decreased ability to detect breathing when little or no lung inflation change occurs 806, such as during airway restriction, severe lung disease, and obstructive apnea.

[0080] Conversely, increased breathing effort can be robustly monitored optically using 850 nm LED light diffused through the skin of the upper arm FIG. SB. Venous blood return to the heart from the skin of the upper arm is either sped up, or slowed down, by the difference between atmospheric pressure and pressure within the chest. Contraction of the diaphragm and intercostal muscles results in decreased pressure within the chest relative to atmospheric pressure. This pressure differential drives air flow through the person’s airway into the lungs, and also speeds blood return to the heart. The amplitude of variation of the trans-thoracic differential pressure is a function of breathing effort. Variation of blood present in the skin of the upper arm, therefore, can be detected optically as variations in light absorption by the skin at 850 nm (at 806’). Normal, unrestricted breathing with healthy lungs during sleep requires very little breathing effort, produces minimal trans-thoracic pressure variations, and results in minimal variation in the optical breathing-induced signal 808. However, large trans-thoracic differential pressure cycles that occur with increased breathing effort due to lung disease or restricted or obstructed airways produce pronounced variations in the optical signal 810.

[0081] FIGS. 9A-9B are graphs that illustrate CEi trend and raw data during severe obstructive and central apnea 900. FIGS. 9A-9B show a selected 8-minute portion of a sleep recording of CEi trend 900-A and the raw data FIG. 9B from which the CEi trend is derived. This recording w as made with a prototype wearable armband sensor that used 660 nm and 850 nm LED light and sampled data at 1 Hz. The total recording indicates that the subject has severe obstructive and central sleep apnea and abnormal periodic breathing during a maj only of the sleep period. The 8-minute portion of the total raw data FIG. 9B shows a brief period of regular breathing 902, but most of the recording alternates between obstructed and restricted breath efforts 904 and prolonged periods of central apnea 906. Throughout this recording period, the subject was having hypoxic stress, as indicated by the relatively lower intensity of the 660 nm light, vs., the 850 nm light. The relative intensity of the two signals throughout the recording was normalized with a signal offset elsewhere in the recordingwhen the subject was awake and breathing normally after having been reclined and relaxed for over an hour.

[0082] FIGS. 10A-10B are graphs illustrating CEi and raw data from abnormal periodic breathing during sleep 1000. Repeating cycles of abnormal periodic breathing are depicted in this 8-minute segment of the total sleep recording using a wearable armband sensor using 660 nm and 850 nm LED light. FIG. 10A is the Cellular Energy Index (CEi) derived from the sensor raw data FIG. 10B. Periodic breathing is a common defect of breathing control during sleep and normally and harmlessly occurs during up to 20% of infant sleep. Periodic breathing becomes abnormal when it results in hypoxic stress. Research has documented that abnormal feedback from the person's blood carbon dioxide sensing cells results in cycles of less than adequate breathing rate and depth, resulting in varying degrees of hypoxic stress 1002, alternating with excessive breathing rate and depth, resulting in varying degrees of recovery 1004. In this adult subject, the periodic breathing cycles vary7between about 25 and 35 seconds in duration. The most apparent feature of this recording is the variation in absorption at 685 nm and the resulting decrease in calculated CEi. The Raw Data breathing rate and effort signal, which is most prominent in the 850 nm data 1006, shows only minimal variation because the subject has normal, easily inflated lungs and minimal airway restriction. The abnormal cycles of skin cellular hypoxic stress response are apparent only in the 1 Hz-sampled 660 nm data trend 1008. Adult sleep recordings comparing the response of pulse oximetry, vs., CEi during abnormal periodic breathing and primary7snoring have consistently demonstrated much earlier and more robust CEi response.

[0083] In an example application, monitoring infants for abnormal periodic breathing during sleep is of great interest as a potential means of preventing death from sudden unexplained infant death (SUID). Published reports of SUID research monitoring sleep of infants at home have documented abnormal periodic breathing in vulnerable infants. One ‘extreme event’ record was included in the published 2001 JAMA report of the 1994- 1998 Collaborative Home Infant Monitor Evaluation (CHIME) study and showed periodic breathing with increasing breath depth over three periodic breathing cycles. This sequence was followed by 10 seconds of primary apnea, followed by 11 obstructed breath efforts over 15 seconds, then secondary apnea for 10 seconds. Death will occur unless effective intervention is provided before or immediately after the onset of secondary apnea. This infant was apparently rescued when awakened by the monitor's audible alarm that sounded 20 seconds after the onset of apnea, and that also likely brought caregiver assistance. Decrease inSpO2to below the normal range did not occur until 15 seconds after the onset of apnea. The infant’s SpO2had decreased with each cycle of the preceding periodic breathing, but was back up to 99% (apparent delayed SpO2data display of about 10 seconds, vs., onset of apnea) at the onset of the alarmed apnea period. The CHIME monitor only recorded 75 seconds of data preceding the ‘16 seconds of apnea’ recording trigger; potentially not recording valuable insights, such as extended occurrence and severity progression of abnormal periodic breathing prior to the alarmed event. The lives of several study infants were apparently saved by the loud alarm of the CHIME monitor that sounded after 20 seconds of apnea. However, if the abnormal breathing sequence leading up to the life-threatening period of apnea had been recognized earlier, a less disruptive intervention, possibly even without waking the baby, may have been just as effective.

[0084] During adult sleep, pulse oximetry cannot detect the skin cellular hypoxia that is detectable by CE monitoring and that precedes the onset of decreased SpO2during abnormal periodic breathing. Also in adults, CEi detects hypoxic stress during primary snoring, and detects hypoxic distress prior to decrease in SpO2due to obstructive and central apnea. Therefore, recognition of episodes of cellular hypoxia in the skin due to abnormal periodic breathing during sleep may be an effective identifier of SUID-risk infants. Continuous monitoring of identified infants with a wearable CE monitor could enable effective countermeasures to prevent death during sleep until the infants have matured past their vulnerable age and abnormal periodic breathing no longer occurs. Gentle stimulation, such as vibration from a miniature buzzer motor housed within the sensor, could be triggered by recognition of cycles of decreased CEi. If the abnormal periodic breathing does not resolve in response to this stimulation, an audible alarm can be generated to wake the baby and bring caregiver assistance. The majority of these vibration responses would likely halt the abnormal periodic breathing long before it becomes life-threatening, and would be silent and not wake the baby or family members. Analysis of the recorded data would help to track the maturation of the infant’s breathing control and help to objectively qualify the infant for discontinuation of monitoring. This monitoring method would also alarm for accidental suffocation and for other causes of potentially lethal physiologic stress to the infant.

[0085] FIGS. 11A-11D illustrate CEi trend, Cabin Barometric Altitude, Raw Data, and G-Meter data during aerobatic flight maneuvers 1100. Data recorded during a proof-of- concept test flight is presented in FIGS. 11A-11D. The pilot placed a wearable CE monitor on his forehead with a headband to record CEi data FIG. 11A as he flew a single-engine,high performance aerobatic aircraft to a maximum altitude of 13000 feet AMSL. The unpressurized cabin barometric altitude 1104 in FIG. 11B was recorded by the avionics system. The CEi response to altitude decreased to nearly -100 1102 at 13000 feet AMSL 1104. Normally, the pilot would wear an oxygen mask and breathe supplemental oxygen when flying over 10000 feet 1106. During this flight, multiple brief and prolonged 3- to 4-Gzmaneuvers were performed and recorded 1108 by the avionics G- Meter FIG. 11D. The CE monitor raw data FIG. 11C shows both the altitude- 1110 and G-force-induced 1112 signal changes. The tandem optical signal responses appear to be driven by G-force-induced decrease in light-absorbing blood in the skin of the pilot’s forehead; resulting in spikes of high signal intensity. From this preliminary data, it is also evident that even moderate +GZexposures 1114 at high altitude may have more physiologic impact 1116 than similar exposures at lower altitude. Excessive Gz-induced physiologic stress and, ultimately monitor- detected impending G-LOC, could be used to automatically trigger autopilot-controlled return of the aircraft to straight and level flight at a safe altitude relative to terrain, and initiate radio transmission to alert local air traffic and air traffic control of the emergency. The pilot would be able to abort the response and take back control of the aircraft when able to do so.

[0086] Flying high-performance aircraft also exposes the pilot to high +GZforces that impede blood flow to the pilot’s brain; potentially resulting in G-induced loss of consciousness (G-LOC). Prolonged relatively lower levels of +GZ, and -Gzimmediately prior to +GZ, are known to be as physiologically stressful as briefer periods of higher +GZ. An automated excess-G force / duration biometric monitor response, therefore, could automatically trigger an autopilot response to place the aircraft in straight and level flight at a safe altitude for surrounding terrain. Automated radio communication could also warn nearby aircraft and initiate potentially life-saving remote piloting rescue operations. These new biometric indicators of physiologic stress have been detected during altitude-induced hypoxia and flight Gztests prior to pilot awareness, and prior to the pilot’s performance being measurably compromised.

[0087] FIGS. 12A-12C illustrates theoretical CEi and raw data trends during carbon monoxide (CO) toxicity 1200. An incompletely resolved problem currently exists relative to General Aviation pilot exposure to CO. Existing CO detectors are available and used in vulnerable aircraft. However, prolonged exposure to low levels of CO may have cumulative adverse physiologic and cognitive effects on the pilot that are not factored into the alarm response of the CO monitors. Pulse oximeters operable to detect carboxy -hemoglobin due tohigh levels of CO exposure, such as those used for screening fire fighters and people rescued from a fire, cannot indicate abnormality prior to the onset of adverse physiologic and cognitive effects. Pilots of piston engine-driven aircraft typically have cabin air heating via a heat exchanger that encloses a portion of the engine exhaust pipe. A crack or other defect of the exhaust manifold, firewall, or air heating system could allow CO contamination of the cabin air. Frequent inspection of the engine exhaust system and in-cabin CO detectors are currently the only means of mitigating this risk.

[0088] FIG. 12A depicts the onset of a theoretical continuous breathing gas exposure of about 40 ppM CO 1202. While this level of exposure is considered at the low end of potential toxicity’, CO is known to be cumulatively bound to both blood hemoglobin and to heme-containing molecules of vital organ tissues; and does not leave the body for up to several days. Extended exposure to even low levels of CO, therefore, presents a cumulative risk to the victim that is currently only measurable by repeated carboxy-hemoglobin blood tests and cognitive function evaluations; neither of which are available to pilots actively flying an aircraft. Routine pulse oximetry does not detect carboxy-hemoglobin and cannot indicate the cellular hypoxia that results from inability' of oxygen to leave the blood to supply cellular needs of vital organs.

[0089] FIG. 12B and FIG. 12C theoretically depict the CEi response FIG. 12B and the raw data FIG. 12C from which it is derived. Natural reflex vasoconstriction in the skin induced by even mild hypoxic stress during exertion results in CEi detection of skin cellular hypoxic stress in the skin despite SpCh remaining within the normal range. The onset of progressively worsening skin cellular hypoxia in a pilot flying at safe-breathing altitude, or with unchanged supplemental oxygen in the breathing gas, would be an indicator of CO toxicity. A feature of the cumulative adverse response to even low levels of CO in the breathing gas would be progressively increasing absorption at 685 nm 1204 and the resulting progressive decrease of the CEi trend 1206. While this is not a measure of cognitive function, this spectral absorption phenomenon in the skin is likely to become evident prior to awareness of the pilot and, hopefully, prior to the onset of decreased cognitive function. This signal response can be used to trigger a warning to the pilot. If the pilot does not respond to the alarm and / or if the CEi value continues to decrease, the monitor system could disconnect autopilot, if set, trigger an emergency announcement by radio, and enable automated and remote pilot rescue.

[0090] Finally, any medical episode suffered by the pilot while the aircraft is flying on autopilot risks the pilot not being able to turn off the autopilot: resulting in the aircraft flying on autopilot until its fuel is consumed, then crashing. Even if the aircraft is suitably equipped, this scenario would prevent the aircraft from being remotely piloted to a safe landing. CEi data could also be used as routine feedback to the breathing gas oxygen blending mechanism to maintain a safe breathing gas ppCh. Coupled with this response could be automated radio transmissions to alert local air traffic and air traffic control of the emergency. If needed, controlled flight toward a nearby airport could then be remotely directed, with remote-controlled or auto-landing of the aircraft performed if the pilot continues to be incapacitated. Other major medical issues, such as heart attack or stroke, are ty pically associated with hypoxic stress and could, by this monitor means, also trigger lifesaving rescue responses.

[0091] FIGS. 13A-C illustrates theoretical CEi trend during current initial breathing therapy for newborn infants 1300. Placental transfer of oxygen and the routing of blood flow within the placenta-fetus major blood vessel system results in the average pCh of blood flowing to the fetus’ head and vital organs being limited to no more than about 35 mmHg 1304. This level of cellular oxygen supply is normal and necessary for optimum fetal development and growth and does not produce hypoxic stress or injury. When babies are bom, they need sufficient time 1306 for their blood vessels and vital organ tissues to adapt to the much higher oxygen supply 1308 that normally comes with breathing air. Raising the oxygen supply too rapidly during the first minutes to hours following birth 1306 risks triggering white blood cell blockage of tiny blood vessels in vital organs, such as the eyes, brain, intestines, and ductus arteriosus. Because these injuries do not always occur with premature infants, there appears to be a safe upper limit 1310, or threshold of cellular oxygen supply below which these problems do not occur. Theoretically, if the cellular hyperoxia threshold 1310 is not exceeded during the infant’s first few hours following birth 1306, the vital organ injuries may be prevented. Arterial blood gas (SaCh) and monitoring of SpCh by pulse oximetry cannot detect the infant’s cellular hyperoxia threshold to enable prevention of these injuries during transition care. During labor 1312, the fetus encounters some degree of hypoxic stress 1314 with each uterine contraction. As with adults, this hypoxic stress likely lowers the cellular hyperoxia threshold 1310 in fetuses. If a premature infant's oxygen supply rises too rapidly immediately following birth 1306, the infant’s cellular hyperoxia threshold 1310 will be exceeded.

[0092] FIG. 13A theoretically depicts the rapid increase of blood pCh 1314 that occurs during the current breathing therapy for full term newborn infants. Exceeding the premature newborn infant’s cellular hyperoxia threshold 1310 risks triggering leukocyte adhesion and obstruction of still-developing capillaries and venules, resulting in tissue necrosis in the most vulnerable vital organs: the eyes, the brain, the intestines, and the ductus arteriosus. All premature newborn infants are potentially vulnerable to these problems until natural cellular adaptation raises the cellular hyperoxia threshold 1310 above the blood pCb level 1314.

[0093] FIG. 13B depicts the theoretical CEi value trend during current newborn care. If a sensor were placed on the fetus, the CEi trend would likely indicate episodes of cellular hypoxia 1316 during labor contractions, indicated by the CEi value decreasing below zero 1318. Upon beginning to breathe air, or higher oxygen content gas, the infant’s blood pO2 1316 rapidly rises through the infant’s CEi hyperoxia threshold 1320 The infant’s blood vessels and vital organ tissues will eventually adapt to the increased oxygen intake, raising the cellular hyperoxia threshold to higher than the current pCh 1314. The therapy goal is to safely and effectively regulate the supply of oxygen during transition 1306 to prevent injury.

[0094] FIG. 13C theoretically depicts the sensor raw data corresponding to FIG. 13B CEi trend data. The 685 nm intensity 1322 rises and stabilizes, but the 850 nm intensity 1324 dips rapidly at the onset of breathing and only slowly returns to baseline. It is this immediate, progressive, and persistent drop in 850 nm intensity 1324 that can serve as an analog indicator of the onset of cellular hyperoxia in blood vessels and vital organs.

[0095] FIGS. 14A-C illustrates theoretical CEi trend during proposed monitor- guided initial breathing therapy for newborn infants 1400. FIG. 14A depicts a proposed Cellular Energy Monitor (CE monitorj-guided process of safely and effectively increasing the oxygen supply to newborn infants. During labor and delivery 1412 there may be some variation in the fetus’ CEi 1416 that, as with exertion of adults, results in skin cellular adaptation to harmless hypoxic stress and resulting in lowering of the cellular hyperoxia threshold 1410. The vulnerable newborn transition period 1406 likely varies in duration between infants, as does the rate of rise in the cellular hyperoxia threshold 1410. As depicted in graph FIG. 14B, keeping below an empirically determined maximum CEi value 1420 and, as show n in graph FIG. 14C avoiding a decrease in 850 nm value 1424 will avoid exceeding the infant’s cellular hyperoxia threshold 1420. This process offers the potential benefit of preventing microvascular occlusion by leukocytes in vital organs that are still undergoingvascular development, and preventing the resulting tissue necrosis. At no time during this proposed process will the infant knowingly receive lower oxygen supply than was received as a thriving fetus before birth; thus, no harmful hypoxic stress.

[0096] FIG. 15 is a table summarizing sensor data changes and physiologic correlations. The table divides the non-pulsatile (DC) biometric spectral responses as either differential or tandem. Differential changes of detected intensity of 685 nm light, vs., 850 nm light are then listed, followed by the physiological correlations relative to the patterns of value change. The category that is roughly analogous to pulse oximetry' is the 685 value response to cellular hypoxic stress. However, pulse oximetry cannot detect cellular hypoxia, the adverse cellular effect of carbon monoxide (CO) toxicity, or the vasoconstriction and resulting skin cellular hypoxia associated with the onset of hemodynamic and septic shock. The devices in FIGS. 1-2 are operable to detect the onset of hemodynamic and septic shock in the skin by the combination if increased absorption of light centered at 685 nm + / - 10 nm relative to decreased absorption of light centered at 850 nm + / - 10 nm. Detecting the onset of hemodynamic and septic shock in the skin is achieved from analysis of the combination of light absorption that is results when reflex vasoconstriction in the skin is detected resulting in skin cellular hypoxia. Detecting the onset of hemodynamic and septic shock in the skin is may be independent of the blood oxygen saturation detected by pulse oximetry or blood gas measurement.

[0097] The differential pattern induced by increasing oxygen supply is divided into safe rate of increase and potentially harmful cellular hyperoxia from too-rapid increase. Safe increase in oxygen supply is accompanied by increasing 685 nm value while the 850 nm value remains stable. Excessively rapid increase in oxygen supply produces a progressively- decreasing 850 nm value trend that persists past the discontinuation of exposure to excess cellular oxygen supply; indicative of oxygen-related cellular chemical injury. Theoretically, avoiding this decrease in 850 nm value will avoid triggering the increased adhesion of white blood cells that results in microvascular occlusion and necrosis of the served tissue. Potential benefits include reduced or eliminated tissue loss from ischemic heart attack and ischemic stroke, and improved organ function following organ transplant. Better outcomes from resuscitation for respiratory and cardiac arrest are also likely. A reduction or elimination of retinal damage, brain hemorrhage and brain tissue injury-, intestinal necrosis, and failure of the ductus arteriosus to close that currently occurs in extremely premature newborn infants could result.

[0098] Tandem variations of sensor data value commonly occur with sensor motion and changes of pressure of the sensor against the skin, which changes the amount of blood in the light path in the skin. When the sensor is mounted on the forehead or cheek, positive Gzinduces a large, tandem increase in data value due to decreased blood in the sensor light path. Negative Gzinduces a large tandem decrease in data value due to increased blood in the light path. Recordings during sleep show abrupt tandem increases and decreases when the sensor is worn on the upper arm and the subject rolls over. This, as with Gzeffects, is due to gravity - (1.0 G) driven redistribution of blood volume resulting in changes in the amount of blood within the sensor light path. When the sensor is above the heart, the tandem data value response is higher value, because there is less blood in the skin to absorb the light.Conversely, when the sensor is below the heart, the tandem response is relatively lower than when the sensor is above the heart. A progressive, tandem decrease in detected values occurs upon reclining to sleep; indicative of rostral shift blood volume redistribution throughout the body. When the monitored subject has remained in the same orientation to gravity to stabilize distribution of blood, changes in total blood volume, such as during artificial kidney dialysis, would likely also be evident in tandem data variations. If hypovolemia occurs during dialysis, natural physiologic reflex responses will occur to compensate, likely resulting in detectable cellular hypoxia in the skin, by CEi, as a result of vasoconstriction that, as with the onset of hemodynamic or septic shock, would likely precede detectable changes in arterial blood pressure or SpCh.

[0099] Tandem cyclic variations in data value occur relative to heart beat cycles and to breathing. The heart beat-induced pulse cycles, or PPG, is ty pically detected by signal sampling at 30-40 Hz and is currently used to monitor heart rate. Differences in intensity value at 660 nm and at 940 nm at pulse peaks and troughs (AC) are used in pulse oximetry to calculate the percent blood hemoglobin oxygen saturation (SpCh). Cardiac output and systemic arterial blood pressure also roughly correlate with the amplitude of the PPG.

[0100] Breathing also induces tandem cyclic data variation. Vary ing trans-thoracic pressure generated by breathing muscles results in air flow into and out of the lungs. This pressure differential also cyclically affects the rate of venous blood return to the heart, and is the basis of pulsus paradoxus, where the pulse amplitude increases during the heart cycles immediately following inhalation-induced increased venous return, and decreases following decreased venous return during exhalation.

[0101] Increased breathing effort produces greater amplitude of the breathing- induced data cycles. Normal breathing during sleep, with healthy lungs and no airway restriction of air flow, produces minimal amplitude of data variation with each breath. Restriction or obstruction of the airway produces correspondingly large amplitude data variations; as does severe lung or airway disease. This ‘optical’ method of monitoring breathing overcomes deficiencies of familiar thoracic ‘electrical impedance' monitoring, which cannot detect restricted or obstructed breaths. This is the reason why clinical sleep studies use RIP belts placed around the chest and abdomen to measure variation in the circumference of the body to detect obstructed breaths. Correlation between the ‘optical’ and ‘electrical impedance’ methods of monitoring breathing also offers potential benefit. The ratio of the amplitude values of these two biometrics could be used as a numeric indicator of breathing effort, and as a continuous numeric scale indicator of the severity of breathing distress.

[0102] OTHER CLINICAL APPLICATIONS

[0103] The disclosed sensor technologies are configurable to extend biometric detection in useful ways. The increased biometric detection can potentially result in improved therapy processes. Automated capture of relevant biometric data reduces the need for uninterrupted clinician vigilance, and improves availability of intervention.

[0104] As will be appreciated by those skilled in the art, several areas of clinical medicine are enhanced by the integrated addition of the new information sources described herein and obtainable via suitable sensors. Clinical areas include, but are not limited to:• Lung disease in children and adults• Prevention of IRI in adult medicine and surgery

[0105] Lung disease in children and adults

[0106] Lung disease can be conceptually divided into two main categories of pathology': (1) increased demand for mechanical effort, and (2) impaired gas diffusion. Inflation pressure assistance to breathing can be lifesaving when the patient’s lungs are difficult to inflate and keep inflated. Patients also need mechanical assistance if they have impaired breathing control, or weakness or paralysis of the breathing mechanism. Supply of oxygen to the blood is impaired by thickening of the alveolar wall and / or fluid in the alveolar spaces. Removal of carbon dioxide from the blood is impaired when the volume of gas exchanged with each breath is limited.

[0107] Acute care monitoring, such as during an acute asthma episode or COPD exacerbation, can be performed with an armband-wearable sensor that detects CEi as an indicator of hypoxic stress. The sensor data is transmitted by RF to a smartphone app for graphic display and recording. Breathing effort is also indicated by this sensor by the amplitude of variation in the 850 nm signal with each breath and can be used to continuously evaluate severity of breathing distress and progress of therapy. Use of the armband sensor format for diagnostic monitoring during sleep has demonstrated detection of hypoxic stress during abnormal periodic breathing, cellular hypoxic stress during primary snoring, and both central and obstructive apnea, all of which are only weakly detectable, or not detectable, by pulse oximetry. At-home optimization of CPAP and APAP therapy has also been demonstrated with this format.

[0108] Monitoring of breathing assistance needs to continuously manage the oxygen content of the breathing gas relative to the patient’s need for oxygen. Maintaining the proper level of breathing effort assures adequate gas exchange. Prevention of later difficulties with weaning from support begins as soon as ventilator assistance is started. Modem ventilators and breathing monitors adequately measure airway pressure and the volume of gas exchanged with each breath. However, accurate information about the patient’s breathing effort and breathing rate control is not currently available to clinicians. As a result, pressure and breathing cycle rate settings are typically made to ‘err on the side of caution’ and provide more mechanical and rate support than needed. This approach relieves the patient of too much of the work of breathing, which results in deconditioning of the patient’s breathing muscles and deterioration of the patient’s natural breathing regulation. This approach also often results in avoidable difficulty weaning from support. The missing information is the relationship between the breathing rate and breath volume, and the effort of breathing relative to the patient’s normal, sustainable ability.

[0109] Monitoring of breathing effort using the absorption of 850 nm light by the skin (optical) combined with thoracic electrical impedance monitoring (impedance) would improve clinical monitoring. The optical breathing variation is apparently due to the rate of venous blood return to the heart from the skin beneath the sensor. Increased effort of breathing, such as with airway obstruction or lung disease, results in more rapid venous blood flow back to the chest during inhalation, less blood in the skin beneath the sensor, and greater amplitude of the optical breathing signal. Conversely, the impedance signal amplitude indicates the relative inflation of the lungs with each breath. Breathing rate can be measuredbest with the highest-amplitude signal. Near normal breathing will produce a relatively higher impedance amplitude signal. Severe lung disease or air flow restriction will produce a relatively higher amplitude optical signal.

[0110] Both optical and impedance breath cycle data can be tracked simultaneously to indicate the "breathing effort, vs. tidal volume effectiveness’ of breathing. A calculated ratio of the detected amplitude of these two signals would provide a patient- and situationspecific numeric scale that would extend from extreme breathing distress (high optical amplitude: low impedance amplitude) to normal breathing (low optical amplitude: high impedance amplitude). As lung disease resolves, there will be decreasing amplitude of the optical signal and increasing amplitude of the impedance signal. Ventilator therapy that keeps patient breathing effort higher than ‘healthy -normal,’ and closely tracks the stability of the patient’s breathing control system, could provide needed support without deconditioning breathing muscles or confusing the patient’s breathing control system.

[0111] Prevention of IRI in medicine and surgery

[0112] Current reperfusion and resuscitation methods target rapid restoration to ‘normal’ blood oxygen supply following severe hypoxic / ischemic stress, such as drowning, cardiac or respirator}' arrest, infarction, etc. Traditional belief holds that the long-term tissue impairment, or death, resulting from these events is due only to the duration and severity of the hypoxic / ischemic period. However, microvascular occlusion by leukocytes is well documented as an initial response to abrupt restoration of ‘normal’ cellular oxygen supply following severe cellular hypoxia / ischemia episodes. The resulting prolonged obstruction to blood flow at the microvascular level becomes the predominant root cause of the necrotic tissue injury. A potentially better therapy approach to returning to normal cellular oxygen supply would be to advance the oxygen supply gradually enough to avoid triggering leukocyte adherence to endothelium and resulting microvascular occlusion. However, the natural cellular adaptation that occurs during hypoxia cannot be detected by blood gas measurements or pulse oximeter monitoring.

[0113] CEi monitoring of the patient’s skin, or of target tissues, could avoid cellular hyperoxia as breathing gas oxygen, or the blood oxygen level perfusing the target organ, is gradually increased. Monitoring also identifies when the starting point of cellular oxygen supply will be sufficiently less than ‘normal’ cellular oxygen supply to not immediately exceed the hypoxia-adapted cellular hyperoxia threshold. Thus, these monitoring techniques allow target areas of organs, such as ischemic heart myocardium, or transplant organs whilebeing reperfused, to be monitored optically either on their surface or via needles carrying optical fibers into the tissue.

[0114] The patient’s low-oxygen venous blood, for example, could be used to initially perfuse previously ischemic tissue or a transplant organ. By this means, distortion of electrolyte and / or nutrient levels could be normalized and stabilized. With CEi trend, plus 850 nm light intensity trend, monitoring of the tissue, the oxygen level in the perfusion blood could be gradually increased by blending in an increasing fraction of the patient’s arterial blood. The rate of increasing the perfusion blood oxygen level would be such that there is a positive CEi trend, with no increase in absorbance of 850 nm light. As the transplant organ warms and / or the target ischemic tissue returns to active metabolism, oxygen consumption will naturally increase. Keeping up with this increased oxygen consumption will be accurately guided by the CEi trend. When only arterial blood is perfusing the transplant organ or target tissue without adverse 850 nm response, the reperfusion therapy can be considered complete. The potential benefit of this approach would be a dramatic decrease in necrotic injury and dysfunction of the reperfused tissue compared with the outcome of the current abmpt ‘normal oxygen’ arterial reperfusion approach.

[0115] OTHER APPLICATIONS

[0116] Monitoring of human performance and environmental challenges has, traditionally, borrowed from medical monitoring technology. However, there are several applications where monitoring with existing medical devices is not possible, is inconvenient, or does not adequately detect the environmental threats. Examples include but are not limited to: firefighters and divers.

[0117] Firefighters wear protective breathing gear while they work in potentially toxic smoke-contaminated atmospheric conditions. The most common toxin in smoke is CO, which interferes with release of oxygen from the blood into the tissue, resulting in cellular hypoxia despite normal or even elevated SpO2. It is well known that CO toxicity always begins to compromise mental and physical function prior to the awareness of the victim. Detecting CO toxicity before the victim becomes aware would enable earlier extraction from exposure and earlier and likely more effective treatment. Immediate, objective detection of CO toxicity is, therefore, potentially life critical. The CE monitor sensor response to CO toxicity is, theoretically, a rapid and continuing decline in CEi value as oxygen delivery to the firefighter’s skin, as an analog of vital organs, is compromised.

[0118] Divers breathe compressed gas blends that, at underwater breathing pressure, may result in excessive cellular delivery of oxygen. Review and research literature lists the typical first symptom of oxygen toxicity as a seizure, which is potentially fatal underwater. Automated adjustment of breathing gas oxygen tension based on a relevant biometric indicator, such as CEi, of the cellular oxygen supply status of the diver could improve diving safety.

[0119] The sensors and biometric signal management can include integration with one or more computing devices. Suitable computing devices can include, without limitation, a mobile user device such as a mobile phone, a smartphone and a cellular phone, a personal digital assistant (PDA), such as an iPhone®, a tablet, a laptop and the like. In at least some configurations, a user can execute a browser application over a network, such as the Internet, to view and interact with digital content, such as screen displays. A display includes, for example, an interface that allows a visual presentation of data from a computing device. Access could be over or partially over other forms of computing and / or communications networks. A user may access a web browser, e.g., to provide access to applications and data and other content located on a website or a webpage of a website.

[0120] A suitable computing device may include a processor to perform logic and other computing operations, e.g., a stand-alone computer processing unit (CPU), or hardwired logic as in a microcontroller, or a combination of both, and may execute instructions according to its operating system and the instructions to perform the steps of the method, or elements of the process. The user’s computing device may be part of a network of computing devices and the methods of the disclosed subject matter may be performed by different computing devices associated with the network, perhaps in different physical locations, cooperating or otherwise interacting to perform a disclosed method. For example, a user’s portable computing device may run an app alone or in conjunction with a remote computing device, such as a server on the Internet. For purposes of the present application, the term “computing device” includes any and all of the above discussed logic circuitry, communications devices and digital processing capabilities or combinations of these.

[0121] Certain embodiments of the disclosed subject matter may be described for illustrative purposes as steps of a method that may be executed on a computing device executing software, and illustrated, by way of example only, as a block diagram of a process flow. Such may also be considered as a software flow chart. Such block diagrams and like operational illustrations of a method performed or the operation of a computing device andany combination of blocks in a block diagram, can illustrate, as examples, software program code / instructions that can be provided to the computing device or at least abbreviated statements of the functionalities and operations performed by the computing device in executing the instructions. Some possible alternate implementations may involve the function, functionalities and operations noted in the blocks of a block diagram occurring out of the order noted in the block diagram, including occurring simultaneously or nearly so, or in another order or not occurring at all. Aspects of the disclosed subject matter may be implemented in parallel or seriatim in hardware, firmware, software, or any combination(s) of these, co-located or remotely located, at least in part, from each other, e.g., in arrays or networks of computing devices, over interconnected networks, including the Internet, and the like.

[0122] The instructions may be stored on a suitable “machine readable medium” within a computing device or in communication with or otherwise accessible to the computing device. As used in the present application a machine-readable medium is a tangible storage device and the instructions are stored in a non-transitory way. At the same time, during operation, the instructions may at sometimes be transitory, e g., in transit from a remote storage device to a computing device over a communication link. However, when the machine readable medium is tangible and non-transitory, the instructions will be stored, for at least some period of time, in a memory storage device, such as a random access memory (RAM), read only memory (ROM), a magnetic or optical disc storage device, or the like, arrays and / or combinations of which may form a local cache memory, e.g., residing on a processor integrated circuit, a local main memory7, e.g., housed within an enclosure for a processor of a computing device, a local electronic or disc hard drive, a remote storage location connected to a local server or a remote server access over a network, or the like. When so stored, the software will constitute a “machine readable medium,” that is both tangible and stores the instructions in a non-transitory7form. At a minimum, therefore, the machine readable medium storing instructions for execution on an associated computing device will be “tangible” and “non-transitory” at the time of execution of instructions by a processor of a computing device and when the instructions are being stored for subsequent access by a computing device.

[0123] Additionally, a communication system of the disclosure comprises: a sensor as disclosed; a server computer system; a measurement module on the server computer system for permitting the transmission of a measurement from a detection device over anetwork. Communications capabilities also include the capability to communicate and display relevant performance information to the user, and support both ANT+ and Bluetooth Smart wireless communications. A storing module on the server computer system for storing the measurement in a detection device server database can also be provided. In some system configurations, the detection device is connectable to the server computer system over at least one of a mobile phone network and an Internet network, and a browser on the measurement recipient electronic device is used to retrieve an interface on the server computer system. In still other configurations, the system further comprising: an interface on the server computer system, the interface being retrievable by an application on the mobile device. Additionally, the server computer system can be configured such that it is connectable over a cellular phone network to receive a response from the measurement recipient mobile device.

[0124] The present teachings may also extend to one or more of the following numbered clauses:1. Devices and methods for the detection of cellular hypoxia in skin or other tissue by increased absorption of light centered at 685 nm + / - 10 nm relative to minimally changed absorption of light centered at 850 nm + / - 10 nm. la. The detection of cellular hypoxia where the decrease in oxygen supply to the skin or other tissue is due to decreased oxygen available in breathing gas. lb. The detection of cellular hypoxia where the decrease in oxygen supply to the skin or other tissue is due to decreased blood perfusion. lc. The detection of cellular hypoxia where the decrease in oxygen supply to the skin or other tissue is due to carbon monoxide (CO) bound to blood hemoglobin.2. Devices and methods for the detection of the onset of hemody namic and septic shock in the skin by the combination of increased absorption of light centered at 685 nm + / - 10 nm, relative to decreased absorption of light at 850 nm + / - 10 nm.2a. The combination of light absorption that is due to the combination of reflex vasoconstriction in the skin resulting in skin cellular hypoxia.2b. The combination of light absorption that may be independent of the level of blood oxygen saturation by pulse oximetry or blood gas measurement.3. Devices and methods for the detection of the onset of cellular hyperoxia in the skin or other tissue by decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm.3a. The detection of cellular hyperoxia where the increase in oxygen supply to the skin or other tissue is due to increased oxygen available in breathing gas.3b. The detection of cellular hyperoxia where the increase in oxygen supply to the skin or other tissue is due to rapidly rising oxygen available in breathing gas or perfusion blood.3c. The clinical application of the detection of cellular hyperoxia used to regulate the rate of rise of oxygen available in breathing gas or perfusion blood.3d. Regulation of the rate of rise in oxygen available during reperfusion therapy in 3c to prevent triggering leukocyte-endothelial adhesion to prevent microvascular occlusion that would result in necrotic tissue injury to reperfused ischemic organ tissues.3e. The regulation of the rate of rise in oxygen available in breathing gas during resuscitation therapy in 3c to prevent triggering leukocyte-endothelial adhesion to prevent microvascular occlusion that would result in necrotic tissue injury to vital organ tissues.3f. The regulation of the rate of rise in oxygen available in breathing gas during transition therapy of newborn infants in 3c to prevent triggering leukocyte-endothelial adhesion to prevent microvascular occlusion that would result in necrotic tissue injury' to vital organs. Devices and methods for the detection of local physiologic effects of positive and negative Gzby tandem changes in absorption of light.4a. The detection of the physiologic effect of positive Gzby tandem decrease in absorption of light by the skin of the subject’s skin, such as on the forehead or cheek. 4b. The detection of the physiologic effect of negative Gzby tandem increase in absorption of light by the skin of the subject’s skin, such as on the forehead or cheek.4c. The detection of the physiologic effect of positive and negative Gzthat is used to initiate automated or remote controlled rescue of aircraft pilots suffering from G- induced loss of consciousness. Devices and methods for the detection of adverse changes in blood volume and blood distribution by tandem variation in absorption of light by the skin.5 a. The detection of adverse changes in blood volume where the blood volume change is due to blood loss.5b. The detection of adverse changes in blood volume where the blood volume change is due to artificial kidney dialysis. Devices and methods for the detection of trend changes in breathing distress as indicated in the amplitude of tandem cyclic variation in absorption of light by the skin.6a. The detection of breathing distress where increased distress is indicated by increased amplitude of breathing-induced cycles of absorption of light by the skin.6b. The detection of breathing distress where the variations in absorption of light are compared with variations in thoracic electrical impedance, such as by a ratio calculation of the respective amplitude values. Devices and methods for the detection of breathing effort before, during, and following breathing assistance as indicated in the amplitude of tandem cyclic variation in absorption of light by the skin.7a. The detection of breathing effort where regulation of ventilation assistance parameters is guided by maintaining increased breathing effort of the subject relative to the subject's normal, healthy level of breathing effort. Devices and methods for the detection of abnormal periodic breathing in infants by cyclic occurrence of cellular hypoxia in the skin.8a. The detection of periodic breathing where breathing rate and depth vary in repeating cycles as detected by cyclic variation in absorption of light.8b. The detection of abnormal periodic breathing where the cyclic variation in breathing rate and depth is accompanied by periods of cellular hypoxia in the skin.8c. The detection of abnormal periodic breathing in infants where the response maybe gentle vibratory stimulation to arouse the infant from sleep to restore normal breathing control to prevent infant death from SUID.8d. The vibratory stimulation in 8c where an audible alarm is generated on a reporting device to bring caregiver assistance if the infant does not regain normal breathing control in response to the vibratory stimulation. A sensor comprising: a housing; a power source; a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength; a second light emitter positioned within the interior of the housing configured to emit a second light at a second w avelength different than the first w avelength; a light detector positioned within the interior of the housing and optically isolated from the first light emitter and the second light emitter wherein the light detector is operable toobtain a timed sampling of light absorption from a tissue; and an analyzer operable to analyze the obtained timed sampling of light centered at 685 nm + / - 10 nm and centered at 850 nm + / - 10 nm.9a. the analyzer determines whether the obtained timed sampling of light has an increased absorption of light centered at 685 nm + / - 10 nm relative to a minimally changed absorption of light centered at 850 nm + / - 10 nm.9b. the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.9c. a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.9d. the analyzer determines whether obtained timed sampling of light has both an increased absorption of light centered at 685 nm + / - 10 nm relative to decreased absorption of light centered at 850 nm + / - 10 nm.9e. the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.9f. a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.9g. the analyzer determines whether obtained timed sampling of light has increased absorption of light centered at 685 nm + / - 10 nm relative to decreased absorption of light centered at 850 nm + / - 10 nm.9h. the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.9i. a plurality' of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue. A sensor for detecting positive and negative Gzeffects on a person comprising: a housing; a power source; a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength; a second light emitterpositioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength; and a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue with a decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm.10a. the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.10b. a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue. ethod of using a sensor comprising the steps of: applying the sensor having a housing, a power source, a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength, a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength, a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue with a decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm; powering the sensor; administering the first light at the first wavelength to the tissue; administering the first light at the first wavelength to the tissue; and obtaining the timed sampling of light absorption from the tissue.1 1 a. Additional steps comprise at least one of : determining whether the tissue has an increased absorption of light centered at 685 nm; determining whether the tissue has a decreased absorption of light centered at 685 nm; determining whether the tissue has a minimally changed absorption of light centered at 850 nm; and determining whether the tissue has a decreased absorption of light centered at 850 nm.1 lb. Additional steps comprise at least one of : identifying cellular hypoxia in tissue by increased absorption of light centered at 685 nm + / - 10 nm relative to minimally changed absorption of light centered at 850 nm + / - 10 nm; identifying hemodynamic or septic shock in tissue by a combination of increased absorption oflight centered at 685 nm + / - 10 nm, relative to a decreased absorption of light at 850 nm + / - 10 nm; identifying cellular hyperoxia in tissue by a decreased absorption of light centered at 685 nm + / - 10 nm relative to an increased absorption of light centered at 850 nm + / - 10 nm; and identifying a local physiologic effect of one or more of positive and negative Gz by tandem changes in absorption of light.12. A system comprising: a sensor having a housing, a power source, a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength, a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength, a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue with both a decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm; and at least one secondary device selected from a second sensor, and a remote computing device.

[0125] While preferred embodiments of the present invention have been shown and described herein, it will be obvious to those skilled in the art that such embodiments are provided by way of example only. Numerous variations, changes, and substitutions will now occur to those skilled in the art without departing from the invention. For example, the use of comprise, or variants such as comprises or comprising, includes a stated integer or group of integers but not the exclusion of any other integer or group of integers. It should be understood that various alternatives to the embodiments of the invention described herein may be employed in practicing the invention. It is intended that any claims presented define the scope of the invention and that methods and structures within the scope of these claims and their equivalents be covered thereby.

Claims

CLAIMSWHAT IS CLAIMED:

1. A sensor comprising: a housing; a power source; a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength; a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength; a light detector positioned within the interior of the housing and optically isolated from the first light emitter and the second light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue; and an analyzer operable to analyze the obtained timed sampling of light centered at 685 nm + / - 10 nm and centered at 850 nm + / - 10 nm.

2. The sensor of claim 1 wherein the analyzer determines whether the obtained timed sampling of light has an increased absorption of light centered at 685 nm + / - 10 nm relative to a minimally changed absorption of light centered at 850 nm + / - 10 nm.

3. The sensor of claim 2 wherein the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.

4. The sensor of claim 2 wherein a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.

5. The sensor of claim 1 wherein the analyzer determines whether obtained timed sampling of light has both an increased absorption of light centered at 685 nm + / - 10 nm relative to decreased absorption of light centered at 850 nm + / - 10 nm.

6. The sensor of claim 5 wherein the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.

7. The sensor of claim 5 wherein a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.

8. The sensor of claim 1 wherein the analyzer determines whether obtained timed sampling of light has increased absorption of light centered at 685 nm + / - 10 nm relative to decreased absorption of light centered at 850 nm + / - 10 nm.

9. The sensor of claim 8 wherein the sensor uses a plurality of wavelengths of light selected by in vivo spectrometry.

10. The sensor of claim 8 wherein a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.

11. A sensor for detecting positive and negative Gzeffects on a person comprising: a housing; a pow er source; a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength; a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength; and a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue with a decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm.

12. The sensor of claim 11 wherein the sensor uses a plurality of w avelengths of light selected by in vivo spectrometry.

13. The sensor of claim 11 wherein a plurality of wavelengths of light are selected to maximize a respective variation in detected cellular light absorbance relative to at least one of a known cellular biochemical phenomenon and a known physiologic phenomenon affecting a monitored tissue.

14. A method of using a sensor comprising the steps of: applying the sensor having a housing, a power source, a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength, a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength, a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the lightdetector is operable to obtain a timed sampling of light absorption from a tissue with a decreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm; powering the sensor; administering the first light at the first wavelength to the tissue; administering the first light at the first wavelength to the tissue; and obtaining the timed sampling of light absorption from the tissue.

15. The method of claim 14 further comprising at least one of determining whether the tissue has an increased absorption of light centered at 685 nm; determining whether the tissue has a decreased absorption of light centered at 685 nm; determining whether the tissue has a minimally changed absorption of light centered at 850 nm; and determining whether the tissue has a decreased absorption of light centered at 850 nm.

16. The method of claim 14 further comprising at least one of identifying cellular hypoxia in tissue by increased absorption of light centered at 685 nm + / - 10 nm relative to minimally changed absorption of light centered at 850 nm + / - 10 nm; identifying hemodynamic or septic shock in tissue by a combination of increased absorption of light centered at 685 nm + / - 10 nm, relative to a decreased absorption of light at 850 nm + / - lO nm; identifying cellular hyperoxia in tissue by a decreased absorption of light centered at 685 nm + / - 10 nm relative to an increased absorption of light centered at 850 nm + / - 10 nm; and identifying a local physiologic effect of one or more of positive and negative Gz by tandem changes in absorption of light.

17. A system comprising: a sensor having a housing, a power source, a first light emitter positioned within an interior of the housing configured to emit a first light at a first wavelength, a second light emitter positioned within the interior of the housing configured to emit a second light at a second wavelength different than the first wavelength, a light detector positioned within the interior of the housing and optically isolated from the first light emitter wherein the light detector is operable to obtain a timed sampling of light absorption from a tissue with adecreased absorption of light centered at 685 nm + / - 10 nm relative to increased absorption of light centered at 850 nm + / - 10 nm; and at least one secondary device selected from a second sensor, and a remote computing device.

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