Patent ID: 12251200

DETAILED DESCRIPTION OF THE INVENTION

As described above, the necklace according to the invention provides a simple, easy-to-wear sensor that measures all vital signs (HR/PR, SpO2, RR, TEMP, and SBP/DBP), hemodynamic parameters (thoracic fluid levels, CO, SV), and motion-related parameters (posture, degree of motion, activity level, and falls). Perhaps the most complex measurement made by the necklace is that for blood pressure, i.e. SBP and DBP. These parameters are determined from PTT separating heartbeat-induced pulses in the ECG and TBI waveforms, coupled with a PP determined from SV determined from the TBI waveform. Using these measurement systems, the necklace's measurement of SBP and DBP is both continuous and cuffless.

Also innovative is the necklace's measurement of SpO2. Here, an optical sensor featuring red and infrared light-emitting diodes (LEDs) clips on to the patient's ear to measure PPG waveforms. These signals pass through a flexible cable to circuitry within the necklace that processes them to determine SpO2.

All analog and digital electronics associated with these measurements are integrated into the strands of the necklace. This means a single component, shaped like a piece of conventional jewelry as opposed to a bulky medical device, measures a robust set of parameters that can characterize a patient using both one-time and continuous measurements. Measurements can take place over just a few minutes or several hours, and are made in medical facilities and at home. The necklace includes a simple LED in its amulet to indicate high-level conditions (e.g., red/yellow/green illuminations depending on the patient's health, as determined from the vital signs and hemodynamic parameters). Also in the amulet is a battery that is easily replaced for long-term, continuous measurements. The necklace includes a wireless transmitter (operating Bluetooth and/or 802.11a/b/g/n) that sends data to, e.g., a conventional mobile device (e.g. cellular telephone, tablet computer, desktop/laptop computer, or plug-in hub).

More specifically,FIG.1shows the necklace30that, during use, is comfortably worn around the patient's neck like a conventional necklace. In this design, the necklace's cable includes all circuit elements, which are typically distributed on an alternating combination of rigid, fiberglass circuit boards and flexible Kapton circuit boards. Typically these circuit boards are potted with a protective material, such as silicone rubber, to increase patient comfort and protect the underlying electronics. The battery for this design can be integrated directly into the cable, or connect to the cable with a conventional connector, such as a stereo-jack connector, micro-USB connector, or magnetic interface.

The necklace30is designed for patients suffering from CHF and other cardiac diseases, such as cardiac arrhythmias, as well as patients with implanted devices such as pacemakers and ICDs. Using the magnetically connected electrodes described in more detail below, it makes impedance measurements to determine CO, SV, and fluid levels, and ECG measurements to determine a time-dependent ECG waveform and HR. Additionally it measures RR, TEMP, SpO2, PR, location, and motion-related properties such as posture, activity level, falls, and degree of motion. The sensor's form factor is designed for both one-time measurements, which take just a few minutes, and continuous measurements, which can take several days. Necklaces are likely familiar to a patient10wearing this system, and this in turn may improve their compliance in making measurements as directed by their physician. Ultimately compliance in using the necklace may improve the patient's physiological condition. Moreover, it is designed to make measurements near the center of the chest, which is relatively insensitive to motion compared to distal extremities, like the arms or hands. The necklace's form factor also ensures relatively consistent electrode placement for the impedance and ECG measurements; this is important for one-time measurements made on a daily basis, as it minimizes day-to-day errors associated with electrode placement. Finally, the necklace's form factor distributes electronics around the patient's neck, thereby minimizing bulk and clutter associated with these components and making it more comfortable to the patient.

In one embodiment the necklace30features a pair of electrode holders34A,34B, located on opposing sides, that each include magnets as described in more detail with respect toFIG.9. The electrode holders34A,34B each receive a separate 3-part magnetically connected electrode patch35,37. During use, the electrode patches35,37connect to their respective electrode holders34A,34B through the magnetic interface, and then stick to the patient's chest when the necklace30is draped around their neck. An adhesive backing supports each conductive electrode within the electrode patch35,37. The electrodes feature a sticky, conductive gel that contacts the patient's skin. The conductive gel contacts a metal pad that is coated on one side with a thin layer of Ag/AgCl, and connects to a magnet through a via. As described in more detail with respect toFIG.3, the outer electrodes in each electrode patch are used for the impedance measurement (they conduct signals V+/−, I+/−), while the inner electrodes are used for the ECG measurement (they conduct signals ECG+/−). Proper spacing of the electrodes ensures both impedance and ECG waveforms having high signal-to-noise ratios; this in turn leads to measurements that are relatively easy to analyze, and thus have optimum accuracy.FIG.9shows preferred dimensions for these components.

A flexible, flat cable38featuring a collection of conductive members transmits signals from the electrode patches35,37to an electronics module36, which, during use, is preferably worn near the back of the neck. Typically the cable38includes alternating regions of rigid fiberglass circuit boards75A-D and flexible Kapton flex circuits77A-F to house other electronic components (used, e.g., for other measurement circuits) and conduct electrical signals. The electronic module36may snap into a soft covering to increase comfort. The electronics module36features a first electrical circuit for making an impedance-based measurement of TBI waveforms that yield CO, SV, RR, and fluid levels, and a second electrical circuit for making differential voltage measurements of ECG waveforms that yield HR and arrhythmia information. The first electrical circuit, which is relatively complex, is shown schematically inFIG.14; the second electrical circuit is well known in this particular art, and is thus not described in detail here.

FIG.10shows a more detailed view of the electronics module36. During a measurement, the second electrical circuit64measures an analog ECG waveform that is received by an internal analog-to-digital converter within a microprocessor62. The microprocessor analyzes this signal to simply determine that the electrode patches are properly adhered to the patient, and that the system is operating satisfactorily. Once this state is achieved, the first61and second64electrical circuits generate time-dependent analog waveforms that a high-resolution analog-to-digital converter62within the electronics module36receives and then sequentially digitizes to generate time-dependent digital waveforms. Analog waveforms can be switched over to this component, for example, using a field effect transistor (FET)63. Typically these waveforms are digitized with 16-bit resolution over a range of about −5V to 5V. The microprocessor62receives the digital waveforms and processes them with computational algorithms, written in embedded computer code (such as C or Java), to generate values of CO, SV, fluid level, and HR. An example of an algorithm is described with reference toFIG.15. Additionally, the electronics module36features a 3-axis accelerometer65and temperature sensor67to measure, respectively, three time-dependent motion waveforms (along x, y, and z-axes) and TEMP values. The microprocessor62analyzes the time-dependent motion waveforms to determine motion-related properties such as posture, activity level, falls, and degree of motion. Temperature values indicate the patient's skin temperature, and can be used to estimate their core temperature (a parameter familiar to physicians), as well as ancillary conditions, such as perfusion, ambient temperature, and skin impedance. Motion-related parameters are determined using techniques known in the art. Temperature values are preferably reported in digital form that the microprocessor receives through a standard serial interface, such as I2C, SPI, or UART.

Both numerical and waveform data processed with the microprocessor are ported to a wireless transmitter66, such as a transmitter based on protocols like Bluetooth or 802.11a/b/g/n. From there, the transmitter sends data to an external receiver, such as a conventional cellular telephone, tablet, wireless hub (such as Qualcomm's 2Net system), or personal computer, as is shown inFIG.12. Devices like these can serve as a ‘hub’ to forward data to an Internet-connected remote server located, e.g., in a hospital, medical clinic, nursing facility, or eldercare facility.

Referring back toFIG.1, and in more detail inFIG.11, a battery module32featuring a rechargeable Li:ion battery48connects at two points to the cable38using a pair of connectors79A,79B. During use, the connectors79A,79B plug into a pair of mated connectors44A,44B that securely hold the terminal ends of the cable38so that the necklace30can be comfortably and securely draped around the patient's neck. Importantly, when both connectors79A,79B are plugged into the battery module32, the circuit within the necklace30is completed, and the battery module32supplies power to the electronics module36to drive the above-mentioned measurements. The connectors79A,79B terminating the cable can also be disconnected from the connectors44A,44B on the battery module32so that this component can be replaced without removing the necklace30from the patient's neck. Replacing the battery module32in this manner means the necklace30can be worn for extended periods of time without having to remove it from the patient. In general, the connectors79A,79B can take a variety of forms: they can be flat, multi-pin connectors, such as those shown inFIG.1, or stereo-jack type connectors, such as those shown inFIG.11, that quickly plug into a female adaptor. Both sets of connectors79A,79B,44A,44B may also include a magnetic coupling so that they easily snap together, thereby making the sensor easy to apply. Typically an LED27on the battery module indicates that this is the case, and that the system is operational. When the battery within battery module32is nearly drained, the LED27indicates this particular state (e.g., by changing color, or blinking periodically). This prompts a user to unplug the battery module32from the two connectors, plug it into a recharge circuit (not shown in the figure), and replace it with a fresh battery module as described above. Also contained within the battery module is a flash memory card23for storing numerical and waveform data, and a micro-USB port25that connects to the flash memory card23for transferring data to a remote computer24. Typically the micro-USB port25is also used for recharging the battery when the sensor is removed from the patient. In embodiments, these components can also be moved to the electronics module36.

As is clear fromFIG.1, the neck-worn cable38serves four distinct purposes: 1) it transfers power from the battery module32to the electronics module36; 2) it ports signals from the electrode patches35,37to the impedance and ECG circuits; 3) it ensures consistent electrode placement for the impedance and ECG measurements to reduce measurement errors; and 4) it distributes the various electronics components and thus allows the necklace to be comfortably worn around the patient's neck. Typically each arm of the cable38will have six wires: two for the impedance electrodes, one for the ECG electrode, and three to pass signals from the electronics module to electrical components within the battery module. These wires can be included as discrete elements, a flex circuit, or, as described above, a flexible cable.

Non-flexible circuit board75B includes a standard pulse oximetry circuit, such as the one described in the following patent application, the contents of which are incorporated herein by reference: BODY-WORN PULSE OXIMETER, U.S.S.N. 20100324389, filed Sep. 14, 2009. The circuit drives red and infrared LEDs in an alternating, pulsatile manner, and additionally controls a light-sensitive diode. During a measurement, the light-sensitive diode receives radiation from the LED that either transmits through or reflects off of tissue. Signals from the light-sensitive diode pass through amplifier and filter circuitry to yield PPG waveforms emanating from the red and infrared radiation. These waveforms are then digitized with an analog-to-digital converter, and then processed to extract fiducial points as described in the above-referenced patent application. The fiducial points are then processed with an algorithm that operates Eq. 3, below, to determine a SpO2 value.

R=red(AC)/red(DC)infrared(AC)/infrared(DC)(3)
In Eq. 3, the red(AC) and red (DC) represent, respectively, parameters extracted from the AC and DC components of the PPG waveform measured with the red LED. A similar case holds for the infrared(AC) and infrared(DC) values. The term ‘AC’ signals, as used herein, refers to a portion of a PPG waveform that varies relatively rapidly with time, e.g. the portion of the signal originating by pulsations in the patient's blood. ‘DC’ signals, in contrast, are portions of the PPG that are relatively invariant with time, e.g. the portion of the signal originating from scattering off of components such as bone, skin, and non-pulsating components of the patient's blood.

More specifically, AC signals are measured from a heartbeat-induced pulse present in both waveforms. The pulse represents a pressure wave, launched by the heart, which propagates through the patient's vasculature and causes a time-dependent increase in volume in both arteries and capillaries. When the pressure pulse reaches vasculature irradiated by the oximeter's optical system, a temporary volumetric increase results in a relatively large optical absorption according to the Beer-Lambert Law. DC signals originate from radiation scattering from static components such as bone, skin, and relatively non-pulsatile components of both arterial and venous blood. Typically only about 0.5-1% of the total signal measured by the photodetector originates from the AC signal, with the remainder originating from the DC signal. Separation of AC and DC signals is typically done with both analog and digital filtering techniques that are well-known in the art.

The R value in Eq. 3, which is sometimes called a ‘ratio of ratios’ (RoR), represents a ratio of Hb to HbO2. It equates an actual SpO2 value, which ranges from 0-100% O2, to an empirical relationship that resembles a non-linear equation. Above about 70% O2 this equation typically yields values that are accurate to a few percent. Measurements below this value, while not necessarily accurate, still indicate a hypoxic patient in need of medical attention. Additional details for this calculation are described in the above-referenced patent application.

As shown inFIG.1, the pulse oximetry circuit within the circuit board75B connects through a cable51terminated with an optical sensor50. In typical embodiments, the optical sensor is in the form of a simple clip wherein the red54and infrared55LEDs are disposed on one arm of the clip, and the light-sensitive diode56is disposed on the opposing arm. The clip typically includes a spring-loaded mechanism so that it can easily connect to the patient's ear53, as shown inFIG.4. Most preferably, the optical sensor50operates in a transmission mode, meaning the LEDs54,55and light-sensitive diode56are positioned as described above. Radiation from the diodes passes through tissue in the earlobe, and then arrives at the light-sensitive diode56, where this component and the pulse oximetry circuit process it to form the requisite PPG waveforms needed for Eq. 3. Alternatively, the optical sensor50can operate in a reflection mode, meaning the LEDs and light-sensitive diode are disposed on the same side of the sensor50, and radiation emitted from the LEDs reflects off a surface of the earlobe before arriving at the light-sensitive diode. In this case, the radiation interacts with a thin layer of tissue, where it is modulated accordingly to form the PPG waveforms.

FIG.5shows a conventional PPG waveform measured with the above-described optical sensor. It features a sequence of heartbeat-induced pulses, with the time duration separating the pulses being inversely related to PR. The heartbeat-induced pulses represent blood pulsing in an underlying artery that absorbs (or reflects) incident radiation from the red and infrared LEDs. The PPG waveform also includes a slowly varying baseline that is due to underlying optical absorption by the blood. PPG waveform emanating from both waveforms look similar, with that from infrared radiation typically having a relatively high signal-to-noise ratio.

FIG.2shows the above-described necklace30worn around the neck of a patient10. As described above, it includes an electronics module36worn on the back of the patient's neck, a battery module32in the front, and electrode holders34A,34B that connect to the magnetically active electrode patches35,37and secure the cable38around the patient's neck that make impedance and ECG measurements.

As shown in the figure, the necklace30drapes around the patient's neck so that non-flexible circuit boards75B,75C are disposed on opposing sides. Within the circuit board75B is the above-described pulse oximetry circuit. The cable51plugs into a connector on the circuit board75B so that it can be easily detached. With this configuration, the optical sensor50can comfortably connect to the patient's earlobe to measure SpO2 values in an effective manner that minimizes cable clutter, and frees the patient's hands and fingers (where pulse oximetry values are normally made) for other purposes. An added benefit of the configuration shown inFIG.2is the reduction of motion artifacts, which can distort PPG waveforms, thus resulting in erroneous SpO2 values. During everyday activities, the head and neck typically move less than the hands and fingers. This means that a sensor configuration like that shown inFIG.2is less susceptible to motion-related artifacts than one where the optical sensor is worn on the patient's finger. Ultimately this improves the accuracy of SpO2 values measured from the patient.

FIG.3indicates in more detail how the above-described electrode measures TBI waveforms and CO/SV values from a patient. As described above, 3-part electrode patches35,37within the neck-worn sensor attach to the patient's chest. Ideally, each patch35,37attaches just below the collarbone near the patient's left and right arms. During a measurement, the impedance circuit injects a high-frequency, low-amperage current (I) through outer electrodes31C,41C,41A,31A. Typically the modulation frequency is about 70 kHz, and the current is about 4 mA. The current injected by electrodes31A,31C is out of phase by 180 degrees from that injected by electrodes41A,41C. It encounters static (i.e. time-independent) resistance from components such as bone, skin, and other tissue in the patient's chest. Additionally, blood and fluids in the chest conduct the current to some extent. Blood ejected from the left ventricle of the heart into the aorta, along with fluids accumulating in the chest, both provide a dynamic (i.e. time-dependent) resistance. The aorta is the largest artery passing blood out of the heart, and thus it has a dominant impact on the dynamic resistance; other vessels, such as the superior vena cava, will contribute in a minimal way to the dynamic resistance.

Inner electrodes31B,41B measure a time-dependent voltage (V) that varies with resistance (R) encountered by the injected current (I). This relationship is based on Ohm's Law, shown below in Eq. 4:
V=I×R(4)
During a measurement, the time-dependent voltage is filtered by the impedance circuit, and ultimately measured with an analog-to-digital converter within the electronics module. This voltage is then processed to calculate SV with an equation such as that shown below in Eq. 5, which is the Sramek-Bernstein equation, or a mathematical variation thereof. Historically, parameters extracted from TBI signals are fed into the equation, shown below, which is based on a volumetric expansion model taken from the aortic artery:

S⁢V=δ⁢L34.2⁢5⁢(dZ⁡(t)/dt)maxZ0⁢LVET(5)

In Eq. 5, Z(t) represents the TBI waveform, δ represents compensation for body mass index, Zo is the base impedance, L is estimated from the distance separating the current-injecting and voltage-measuring electrodes on the thoracic cavity, and LVET is the left ventricular ejection time, which is the time separating the opening and closing of the aortic valve, and can be determined from the TBI waveform, or from the HR using an equation called ‘Weissler's Regression’, shown below in Eq. 6, that estimates LVET from HR:
LVET=−0.0017×HR+0.413  (6)
Weissler's Regression allows LVET, to be estimated from HR determined from the ECG waveform. This equation and several mathematical derivatives, along with the parameters shown in Eq. 5, are described in detail in the following reference, the contents of which are incorporated herein by reference: ‘Impedance Cardiography, Pulsatile blood flow and the biophysical and electrodynamic basis for the stroke volume equations’, Bernstein, Journal of Electrical Bioimpedance, Vol. 1, p. 2-17, 2010. Both the Sramek-Bernstein Equation and an earlier derivative of this, called the Kubicek Equation, feature a ‘static component’, Z0, and a ‘dynamic component’, ΔZ(t), which relates to LVET and a (dZ/dt)max/Zovalue, calculated from the derivative of the raw TBI signal, □Z(t). These equations assume that (dZ(t)/dt)max/Zorepresents a radial velocity (with units of Ω/s) of blood due to volume expansion of the aorta.

In Eq. 5 above, the parameter Z0will vary with fluid levels. Typically a high resistance (e.g. one above about 30Ω) indicates a dry, dehydrated state. Here, the lack of conducting thoracic fluids increases resistivity in the patient's chest. Conversely, a low resistance (e.g. one below about 19Ω) indicates the patient has more thoracic fluids, and is possibly overhydrated. In this case the abundance of conducting thoracic fluids decreases resistivity in the patient's chest. The TBI circuit and specific electrodes used for a measurement may affect these values. Thus, the values can be more refined by conducting a clinical study with a large number of subjects, preferably those in various states of CHF, and then empirically determining ‘high’ and ‘low’ resistance values.

FIG.6shows derivatized TBI and ECG waveforms measured with the necklace ofFIG.1plotted over a short (about 5 seconds) time window (top), and TBI waveforms plotted over a longer window (bottom, 60 seconds). Referring first to the top portion of the figure, individual heartbeats produce time-dependent pulses in both the ECG and TBI waveforms. The TBI waveform shown in the figure is the first mathematical derivative of a raw TBI waveform. As is clear from the data, pulses in the ECG waveform precede those in the TBI waveform. The ECG pulses, each featuring a sharp, rapidly rising QRS complex, indicate initial electrical activity in contractions in the patient's heart, and, informally, the beginning of the cardiac cycle. The QRS complex is the peak of the ECG waveform. TBI pulses follow the QRS complex by about 100 ms, and indicate blood flow through arteries in the patient's thoracic cavity. These signals are dominated by contributions from the aorta, which is the largest artery in this region of the body. During a heartbeat, blood flows from the patient's left ventricle into the aorta. The volume of blood is the SV. Blood flow enlarges this vessel, which is typically very flexible, and also temporarily aligns blood cells (called erythrocytes) from their normally random orientation. Both of these mechanisms—enlargement of the aorta and temporary alignment of the erythrocytes—improve electrical conduction near the aorta, thus decreasing the electrical impedance as measured with TBI. The waveform shown in the upper portion ofFIG.6is a first derivative of the raw TBI waveform, meaning its peak represents the point of maximum impedance change.

A variety of time-dependent parameters can be extracted from the ECG and TBI waveforms. For example, as shown in the upper portion of the figure, it is well known that HR can be determined from the time separating neighboring ECG QRS complexes. Likewise, LVET can be measured directly from the TBI pulse. LVET is measured from the onset of the derivatized pulse to the first positive going zero crossing. Also measured from the derivatized TBI pulse is (dZ/dt)max, a parameter that is used to calculate SV, as shown in Eq. 5 and described in more detail in the reference described above.

The time difference between the ECG QRS complex and the peak of the derivatized TBI waveform represents a PTT. This value can be calculated from other fiducial points, particularly on the TBI waveform (such as the base or midway point of the heartbeat-induced pulse). But typically the peak of the derivatized waveform is used, as it is relatively easy to develop a software beat-picking algorithm that finds this fiducial point.

PTT correlates inversely to SBP and DBP, as shown below in Eqs. 7-8, where mSBPand mDBPare patient-specific slopes for, respectively, SBP and DBP, and SBPcaland DBPcalare values, respectively, of SBP and DBP measured during a calibration measurement. Without the calibration PTT only indicates relative changes in SBP and DBP. A calibration can be provided with conventional means, such as an oscillometric blood pressure cuff or in-dwelling arterial line. The calibration yields both the patient's immediate value of SBP and DBP. Multiple values of PTT and blood pressure can be collected and analyzed to determine patient-specific slopes mSBPand mDBP, which relate changes in PTT with changes in SBP and DBP. The patient-specific slopes can also be determined using pre-determined values from a clinical study, and then combining these measurements with biometric parameters (e.g. age, gender, height, weight) collected during the clinical study.

SBP=mSBPPTT+SBPcal(7)DBP=mDBPPTT+DBPcal(8)

In embodiments, waveforms like those shown in the upper portion ofFIG.6are processed to determine PTT, which is then used to determine either SBP or DBP according to Eqs. 7 or 8. Typically PTT and SBP correlate better than PTT and DBP, and thus this parameter is first determined. Then PP is estimated from SV, calculation of which is described below. Most preferably, instant values of PP and SV are determined, respectively, from the blood pressure calibration and from the TBI waveform.

PP can be estimated from either the absolute value of SV, SV modified by another property (e.g. LVET), or the change in SV. In the first method, a simple linear model is used to process SV (or, alternatively, SV×LVET) and convert it into PP. The model uses the instant values of PP and SV, determined as described above from a calibration measurement, along with a slope that relates PP and SV (or SV×LVET). The slope can be estimated from a universal model that, in turn, is determined using a population study. Alternatively, a slope tailored to the individual patient is used. Such a slope can be selected, for example, using biometric parameters describing the patient, as described above. Here, PP/SV slopes corresponding to such biometric parameters are determined from a large population study, and then stored in computer memory on the necklace. When a necklace is assigned to a patient, their biometric data is entered into the system, e.g. using a mobile telephone that transmits the data to a microprocessor in the necklace via Bluetooth. Then an algorithm on the necklace processes the data and selects a patient-specific slope. Calculation of PP from SV is described in the following reference, the contents of which are incorporated herein by reference: ‘Pressure-Flow Studies in Man. An Evaluation of the Duration of the Phases of Systole’, Harley et al.,Journal of Clinical Investigation, Vol. 48, p. 895-905, 1969. As described in this reference, the relationship between PP and SV for a given patient typically has a correlation coefficient (r) that is greater than 0.9, which indicates excellent agreement between these two properties. Similarly, in the above-mentioned reference, SV is shown to correlate with the product of PP and LVET, with most patients showing an r value of greater than 0.93, and the pooled correlation value (i.e. that for all subjects) being 0.77. This last result indicates that a single linear relationship between PP, SV, and LVET may hold for all patients.

More preferably, PP is determined from SV using relative changes in these values. Typically the relationship between the change in SV and change in PP is relatively constant across all subjects. Thus, similar to the case for PP, SV, and LVET, a single, linear relationship can be used to relate changes in SV and changes in PP. Such a relationship is described in the following reference, the contents of which are incorporated herein by reference: ‘Pulse pressure variation and stroke volume variation during increased intra-abdominal pressure: an experimental study’,Didier et al.,Critical Care, Vol. 15:R33, p. 1-9, 2011. Here, the relationship between PP variation and SV variation for 67 subjects displayed a linear correlation of r=0.93, and extremely high value for pooled results that indicates a single, linear relationship may hold for all patients.

From such a relationship, PP is determined from the TBI-based SV measurement, and SBP is determined from PTT. DBP is then calculated from SBP and PP.

The necklace determines RR from both the TBI waveform, and from a motion waveform generated by the accelerometer (called the ACC waveform), which is typically located in analog circuitry within the necklace, as described above. The bottom portion ofFIG.6indicates how the TBI waveform yields RR. In this case, the patient's respiratory effort moves air in and out of the lungs, thus changing the impedance in the thoracic cavity. This time-dependent change maps onto the TBI waveform, typically in the form of oscillations or pulses that occur at a much lower frequency than the heartbeat-induced cardiac pulses shown in the upper part ofFIG.5. Simple signal processing (e.g. filtering, beat-picking) of the low-frequency, breathing-induced pulses in the waveform yields RR.

Likewise, the ACC waveform will reflect breathing-induced movements in the patient's chest. This results in pulses within the waveform that have a similar morphology to those shown in the lower portion ofFIG.6for the TBI waveform. Such pulses can be processed as described above to estimate RR. RR determined from the ACC waveform can be used by itself, or processed collectively with RR determined from the TBI waveform (e.g., using adaptive filtering) to improve accuracy. Such an approach is described in the following patent application, the contents of which are incorporated herein by reference: BODY-WORN MONITOR FOR MEASURING RESPIRATION RATE, U.S.S.N. 20110066062, Filed Sep. 14, 2009.

As shown in the lower portion ofFIG.6, the baseline of the TBI waveform, called Zo, can be easily determined. Zo is used to determine SV, as described above in Eq. 5.

FIGS.7and8show how a process called lower body negative pressure (LBNP) affects baseline impedance Zo and SV. LBNP serves as a surrogate for hemorrhage, a process that typically results in dramatic changes in SV.FIG.7shows pooled results from 39 subjects undergoing a gradual increase in LBNP from 0 mmHg (i.e. no change from ambient) to a vacuum of 60 mmHg (corresponding to a loss of blood of about 2 L). The data shown in this figure are averaged over all 39 subjects, and impedance waveforms similar to those described above were measured from the thoracic cavity and analyzed to determine SV and thoracic fluid level. As shown in the top portion of the figure, the change in baseline impedance correlates in a linear manner with the LBNP level, with the agreement between these parameters (Pearson's correlation coefficient r2=0.9998) being extremely high. Here, vacuum applied during LBNP gradually removes conductive fluids from the thoracic cavity, thus decreasing conductivity and increasing baseline impedance. Similarly, the relationship between LBNP level and SV shown in the bottom half of the plot is also linear, with the slope going in the opposite direction as that for the impedance/LBNP correlation. In this case increasing LBNP removes blood from the patient's thoracic cavity, thus reducing their effective blood volume (called ‘pre-load’) and essentially simulating hemorrhage. During hemorrhage, the body is trained to reduce blood flow by decreasing the amount of blood pumped by the heart (the SV) to preserve perfusion of the internal organs. Thus, it is expected that increasing LBNP will systematically decrease SV, which is exactly what is shown in the lower half ofFIG.7. The correlation for this relationship is also quite high, with r2=0.99531.

In conclusion, the results shown inFIG.7indicate that two parameters that change with the onset of CHF—thoracic fluid level and SV—can be accurately measured with an impedance-based technique, such as that deployed with the sensor described herein.

The data shown inFIG.7are averaged over all 39 subjects, while the individual correlation coefficient for each subject for the above-described measurements are shown inFIG.8. As is clear from these data, 36 out of 39 subjects show a correlation between LBNP level (representing a proxy for fluid level, as described above) and baseline impedance characterized by r>0.98, which is extremely high. Similarly, 36 out of 39 subjects show a correlation between LBNP level and SV characterized by r>0.9. Both of these plots indicate that the parameters measured by impedance measurements show promise for being an accurate physiological monitor.

FIG.12depicts how the necklace30shown inFIG.1is designed to facilitate remote monitoring of a patient10. As shown in the top portion of the figure, after the necklace30measures the patient, it automatically transmits data through its internal Bluetooth wireless transmitter to the patient's cellular telephone20. In this case, the cellular telephone20preferably runs a downloadable software application that accesses the phone's internal Bluetooth drivers, and includes a simple patient-oriented application that renders data on the phone's screen. From there, using its internal modem, the cellular telephone20transmits data to an IP address associated with a computer server22. The computer server22, in turn, renders a web-based system that displays data for clinicians at a hospital, medical clinic, nursing facility, or eldercare facility. The web-based system may show ECG and TBI waveforms, trended numerical data, the patient's medical history, along with their demographic information. A clinician viewing the web-based system may, for example, analyze the data and then call the patient10and have them adjust their medications or diet. Alternatively, as shown in the lower half of the figure, the necklace30can automatically transmit data through Bluetooth to a personal computer24, which then uses a wired or wireless Internet connection to transmit data to the computer server22. Here, the personal computer24runs a custom software program to download data from the sensor22, display it for the patient in an easy-to-understand format, and then forward it to the computer server for a relatively complex analysis as described above. In yet another embodiment, the necklace30is directly plugged into the personal computer24through a USB connection, and data are downloaded using a wired connection and forwarded to the computer server22as described above.

FIG.13shows examples of user interfaces90,91,92that integrate with the above-mentioned systems and run on the cellular telephone20, shown in this case as an iPhone. The user interfaces show information such as patient demographics (interface90), patient-oriented messages (interface91), and numerical vital signs and time-dependent waveforms (interface92). The interfaces shown in the figures are designed for the patient. More screens, of course, can be added, and similar interfaces (preferably with more technical detail) can be designed for the actual clinician. The interfaces can also be used to render operational reports, which are then sent off to a clinician for review.

FIG.14shows an analog circuit100that performs the impedance measurement according to the invention. The figure shows just one embodiment of the circuit100; similar electrical results can be achieved using a design and collection of electrical components that differ from those shown in the figure.

The circuit100features a first magnetically connected electrode115A that injects a high-frequency, low-amperage current (I1) into the patient's brachium. This serves as the current source. Typically a current pump102provides the modulated current, with the modulation frequency typically being between 50-100 KHz, and the current magnitude being between 0.1 and 10 mA. Preferably the current pump102supplies current with a magnitude of 4 mA that is modulated at 70 kHz through the first electrode115A. A second magnetically connected electrode117A injects an identical current (I2) that is out of phase from I1by 180°.

Another pair of magnetically connected electrodes115B,117B measure the time-dependent voltage encountered by the propagating current. These electrodes are indicated in the figure as V+ and V−. As described above, using Ohm's law, the measured voltage divided by the magnitude of the injected current yields a time-dependent resistance to ac (i.e. impedance) that relates to blood flow in the aortic artery. As shown by the waveform128in the figure, the time-dependent resistance features a slowly varying dc offset, characterized by Zo, that indicates the baseline impedance encountered by the injected current; for TBI this will depend, for example, on the amount of thoracic fluids, along with the fat, bone, muscle, and blood volume in the chest of a given patient. Zo, which typically has a value between about 10 and 150Ω, is also influenced by low-frequency, time-dependent processes such as respiration. Such processes affect the inherent capacitance near the chest region that TBI measures, and are manifested in the waveform by low-frequency undulations, such as those shown in the waveform128. A relatively small (typically 0.1-0.5Ω) AC component, ΔZ(t), lies on top of Zo and is attributed to changes in resistance caused by the heartbeat-induced blood that propagates in the brachial artery, as described in detail above. □Z(t) is processed with a high-pass filter to form a TBI signal that features a collection of individual pulses130that are ultimately processed to determine SV and CO.

Voltage signals measured by the first electrode115B (V+) and the second electrode117B (V−) feed into a differential amplifier107to form a single, differential voltage signal which is modulated according to the modulation frequency (e.g. 70 kHz) of the current pump102. From there, the signal flows to a demodulator106, which also receives a carrier frequency from the current pump102to selectively extract signal components that only correspond to the TBI measurement. The collective function of the differential amplifier107and demodulator106can be accomplished with many different circuits aimed at extracting weak signals, like the TBI signal, from noise. For example, these components can be combined to form a ‘lock-in amplifier’ that selectively amplifies signal components occurring at a well-defined carrier frequency. Or the signal and carrier frequencies can be deconvoluted in much the same way as that used in conventional AM radio using a circuit that features one or more diodes. The phase of the demodulated signal may also be adjusted with a phase-adjusting component108during the amplification process. In one embodiment, the ADS1298 family of chipsets marketed by Texas Instruments may be used for this application. This chipset features fully integrated analog front ends for both ECG and impedance pneumography. The latter measurement is performed with components for digital differential amplification, demodulation, and phase adjustment, such as those used for the TBI measurement, that are integrated directly into the chipset.

Once the TBI signal is extracted, it flows to a series of analog filters110,112,114within the circuit100that remove extraneous noise from the Zo and ΔZ(t) signals. The first low-pass filter110(30 Hz) removes any high-frequency noise components (e.g. power line components at 60 Hz) that may corrupt the signal. Part of this signal that passes through this filter110, which represents Zo, is ported directly to a channel in an analog-to-digital converter120. The remaining part of the signal feeds into a high-pass filter112(0.1 Hz) that passes high-frequency signal components responsible for the shape of individual TBI pulses130. This signal then passes through a final low-pass filter114(10 Hz) to further remove any high-frequency noise. Finally, the filtered signal passes through a programmable gain amplifier (PGA)116, which, using a 1.65V reference, amplifies the resultant signal with a computer-controlled gain. The amplified signal represents ΔZ(t), and is ported to a separate channel of the analog-to-digital converter120, where it is digitized alongside of Zo. The analog-to-digital converter and PGA are integrated directly into the ADS1298 chipset described above. The chipset can simultaneously digitize waveforms such as Zo and ΔZ(t) with 24-bit resolution and sampling rates (e.g. 500 Hz) that are suitable for physiological waveforms. Thus, in theory, this one chipset can perform the function of the differential amplifier107, demodulator108, PGA116, and analog-to-digital converter120. Reliance of just a single chipset to perform these multiple functions ultimately reduces both size and power consumption of the TBI circuit100.

Digitized Zo and Z(t) waveforms are received by a microprocessor124through a conventional digital interface, such as a SPI or I2C interface. Algorithms for converting the waveforms into actual measurements of SV and CO are performed by the microprocessor124. The microprocessor124also receives digital motion-related waveforms from an on-board accelerometer122, and processes these to determine parameters such as the degree/magnitude of motion, frequency of motion, posture, and activity level.

FIG.15shows a flow chart of an algorithm133A that functions using compiled computer code that operates, e.g., on the microprocessor124shown inFIG.6. The algorithm133A is used to measure TBI waveforms in the presence of motion. The compiled computer code is loaded in memory associated with the microprocessor, and is run each time a TBI measurement is converted into a numerical value for CO and SV. The microprocessor typically runs an embedded real-time operating system. The compiled computer code is typically written in a language such as C, C++, Java, or assembly language. Each step135-150in the algorithm133A is typically carried out by a function or calculation included in the compiled computer code.

Algorithms similar to that shown inFIG.15can be used to calculate other physiological parameters in the presence of motion, such as SpO2, RR, HR, and PR.

In other embodiments, algorithms can process other waveforms, such as the PPG and ECG waveforms, to extract parameters such as RR. Here, the low-frequency envelope of the waveform indicates RR. In other embodiments, a reflective pulse oximetry system can measure SpO2 without requiring an ear-worn optical sensor, such as that shown inFIG.4. In this case the sensor uses reflective-mode optical configurations to measure both the red and infrared PPG waveforms. In still other embodiments, the electronics within the necklace, as shown inFIGS.10and11, are moved within the necklace's geometry. For example, they can be moved from the back portion of the necklace to a side portion proximal to the front of the patient's neck.

Still other embodiments are within the scope of the following claims.