Patent Publication Number: US-6342040-B1

Title: Patient monitor and method of using same

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This is a Continuation of U.S. Pat. application Ser. No. 09/030,221 filed Feb. 25, 1998, now U.S. Pat. No. 6,017,315. 
    
    
     BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The present invention pertains to a patient monitor for monitoring and/or quantitatively measuring a physiological characteristic of the patient, and, in particular, to an apparatus and method for monitoring and/or quantitatively measuring a physiological characteristic based, at least in part, on a pressure differential between a pressure within a user interface and an ambient atmospheric pressure outside the user interface. 
     2. Description of the Related Art 
     There are many situations in which it is necessary or desirable to measure a physiological characteristic of a patient, such as characteristics associated with respiration. Examples of characteristics associated with respiration include the patient&#39;s flow, inspiratory period, expiratory period, tidal volume, inspiratory volume, expiratory volume, minute ventilation, respiratory rate, ventilatory period, and inspiration to expiration (I to E) ratio. It is also important in many situations to identify still other characteristics associated with respiration, such as identifying the start, end and duration of a patient&#39;s inspiratory phase and expiratory phase, as well as detecting patient snoring. For example, when conducting a sleep study to diagnose sleep disorders or when conducting other pulmonary monitoring functions, it is common to measure the respiratory rate and/or the air flow to and from the patient. Distinguishing between inspiration and expiration is useful, for example, in triggering a pressure support device that provides breathing gas to a patient. 
     There are several known techniques for monitoring patient breathing for these purposes. A first conventional technique involves placing a thermistor or thermocouple in or near the patient&#39;s airway so that the patient&#39;s breath passes over the temperature sensing device. Breathing gas entering the patient has a temperature that is generally lower than the exhaled gas. The thermistor senses this temperature difference and outputs a signal that can be used to distinguish between inspiration and expiration. 
     A primary disadvantage of the thermistor or thermocouple air flow sensing technique is that these devices cannot quantitatively measure the flow and/or volume of breathing gas delivered to and/or exhaled from the patient, because the signal from the sensor is a measure of air temperature, not air flow or pressure. Typically, a thermistor air flow sensor is only used to differentiate between inspiration and expiration. Sensors that detect humidity have similar uses and similar disadvantages. 
     A second conventional technique for measuring the airflow to and from a patient is illustrated in FIG.  1  and involves placing a pneumotach sensor  30  in a breathing circuit  31  between a supply of breathing gas, such as a ventilator or pressure support device, and the patient&#39;s airway. In a conventional pneumotach, the entire flow of breathing gas Q IN  is provided to a patient  32  from a pressure source  34 . Conversely, all of the gas expelled from patient  32 , passes through pneumotach  30  so that during operation, there is a two-way flow of gas through pneumotach  30 . 
     In its simplest form shown in FIG. 1, pneumotach  30  includes a flow element  36  having an orifice  38  of a known size defined therein. Flow element  36  provides a known resistance R to flow through the pneumotach so that a pressure differential ΔP exists across of flow element  36 . More specifically, flow element  36  causes a first pressure P 1  on a first side of the flow element to be different than a second pressure P 2  on a second side of the flow element opposite the first side. Whether P 1  is greater than P 2  or vice versa depends on the direction of flow through the pneumotach. 
     In a first type of conventional pneumotach, a major portion Q 1  of the total flow Q IN  of gas delivered to pneumotach  30  passes through orifice  38 . The pressure differential ΔP created by flow element  36  causes a lesser portion Q 2  of the gas delivered to the pneumotach to be diverted through a bypass channel  40 , which is connected to breathing circuit  31  across flow element  36 . An airflow sensor  42  in bypass channel  40  measures the flow of gas therethrough. Because the area of orifice  38  and the area of bypass channel  40  are known and fixed relative to one another, the amount of gas Q 2  flowing through bypass channel  40  is a known fraction of the total gas flow Q IN  delivered to pneumotach  30 . Airflow sensor  42  quantitatively measures the amount of gas Q 2  passing through bypass channel  40 . Once this quantity is known, the total flow Q IN  of gas passing through pneumotach  30  can be determined. 
     In a second type of conventional pneumotach, a pressure sensor, rather than an airflow sensor, is provided in bypass channel  40 . Gas does not pass through the pressure sensor. Instead, each side of a diaphragm in the pressure sensor communicates with respective pressures P 1  and P 2  on either side of flow element  36 . The pressure sensor measures pressure differential ΔP across flow element  36 . For example, for flow in the direction illustrated in FIG. 1, pressure differential ΔP across flow element  36  is P 1 -P 2 . Once pressure differential ΔP is known, the flow rate Q IN  of gas passing through pneumotach  30  can be determined using the equation, ΔP=RQ 2 , where R is the known resistance of flow element  36 . 
     Another conventional pneumotach  44  is shown in FIG.  2 . Pneumotach  44  improves upon pneumotach  30  in FIG. 1 by providing a first linear flow element  46  in place of flow element  36 . First linear flow element  46  functions in the same manner as flow element  36  by creating a pressure differential in breathing circuit  31 . However, flow element  46  has a plurality of honey-comb like channels that extend in the direction of gas flow to linearize the flow of gas through the pneumotach. The previous flow element  36  in FIG. 1 can create downstream turbulence that hinders the flow of gas through the bypass channel or causes fluctuations in the downstream pressure, thereby degrading the airflow or pressure differential signal output by sensor  42 . Flow element  46  solves this problem by providing a plurality of honeycomb-like channels having longitudinal axis parallel to the axis of the breathing circuit. The honeycomb channels ensure that the flow across the downstream port of the bypass channel is linear, i.e., non-turbulent. 
     To ensure that the flow of gas across the port in bypass channel  40  upstream of flow element  46  is also linear, i.e., non-turbulent, other linear flow elements  48  and  50  are provided in the breathing circuit Flow elements  48  and  50  have the same honeycomb configuration as flow element  46 . Because gas can flow in both directions through pneumotach  44 , flow elements  48  and  50  are respectively located on each side of flow element  46  so that each entry port for bypass channel  40  is downstream of one of these additional flow elements regardless of the direction of flow through the pneumotach. 
     Although a pneumotach improves upon a theremistor in that it quantatively measures the flow and/or volume of gas passing therethrough, it also has significant disadvantages. For example, a pneumotach is relatively complicated and therefore difficult and costly to manufacture. It is also difficult to clean and is relatively large. Because of its size, which is dictated by the need to measure the pressure differential or flow across the flow element in the breathing circuit, it creates a relatively large amount of dead space in the patient breathing circuit, which is not conducive to minimizing rebreathing of CO 2 . Because of its complexity, a pneumotach may leak, and its operating capabilities can suffer as a result of heat and moisture buildup. 
     A third type of conventional airflow meter, illustrated in FIG. 3, is a nasal cannula airflow meter  52 . Nasal cannula airflow meter  52  is similar to a nasal oxygen cannula in that it includes a pair of ports  54  and  56  that insert into nares  58  and  60  of the user. A hollow tubing  62  carries a fraction of the total amount of breathing gas to a sensor, such as an airflow or pressure sensor, If the total area of the user&#39;s nares relative to the total area of the ports  54  and  56  is known, the nasal cannula airflow meter can provide a quantitative measure of the patient airflow. 
     However, bececse the total area of each user&#39;s nares can vary from person to person, a commonly sized nasal cannula airflow meter cannot provide an accurate, quantitative measure of the airflow for all users. If two people have different sized nasal openings, the fraction of the exhaled air that is being delivered to the ports of the nasal cannula cannot be known for both users. For example, a first user may deliver 30% of the exhaled gas to the ports of the nasal cannula, while a second user may deliver only 10% of the exhaled to the same sized nasal cannula This variation in the percentage of gas delivered to the same size cannula is due to the variation in the total cross-sectional area of the nares of both users. For the same size nasal cannula, a user with larger nares will deliver a smaller percentage of the total exhaled gas to the ports of the nasal cannula than a user with smaller nares. Tlus, a conventional nasal cannula cannot accurately measure the airflow for a plurality of users having different sized nares. 
     In addition to detecting and measuring quantities associated with the rate of volume of air being delivered to a patient, there are also many instances where it is important to detect other characteristics associated with respiration, such as snoring. The onset of snoring andlor the intensity of snoring can be used, for example, as a trigger to initiate or control the level of a positive pressure therapy provided the patient. Also, the presence, intensity and/or duration of snoring can be used as a diagnostic tool in determining whether the patient suffers from a sleep and/or breathing disorder. 
     It is known to use a microphone or pressure sensor mounted on the exterior of the patient&#39;s neck to detect sounds or throat vibrations generated by the snore. In many situations, these sensors are mounted on the user as an individual unit and are not connected to other structures worn by the patient. This can result in incorrect or inefficient placement of such sensors. Also, conventional snore sensing devices are quite susceptible to noise. For example, microphones can pick up external sounds not produced by the patient, such as snoring of a person or animal near the patient, and/or sounds not resulting from snoring, such as coughing. Pressure sensors can be adversely effected by body movements, such as normal movements that take place during the night and/or throat vibrations resulting from coughing. 
     SUMMARY OF THE INVENTION 
     Accordingly, it is an object of the present invention to provide a patient monitoring device for monitoring and/or quantitatively measuring a physiological characteristic of the patient, and, in particular, a characteristic associated with respiration, that does not suffer from the disadvantages of convention airflow/volume meters and snore detectors. This object is achieved by providing a user interface having an interior portion that communicates with an airway of a user such that substantially all gas inhaled and exhaled by the user enters the interior portion of the user interface. At least one vent element is associated with the user interface and connects the interior portion of the user interface with the ambient atmosphere outside the user interface. The vent element and the user interface define a flow element across which a pressure differential is created during inspiration and expiration. This pressure differential is a pressure difference between a first pressure within the interior portion of the user interface and the pressure of the ambient atmosphere outside the user interface. A sensor coupled to the interior portion of the user interface measures a fluid characteristic resulting from the pressure differential and outputs a signal indicative of that fluid characteristic. This signal can be used to monitor and/or measure physiological characteristics of the patient. In a preferred embodiment of the present invention, the signal output by the sensor corresponds to a characteristic associated with respiration and a processing unit receives this signal and determines a quantitative value for the characteristic associated with respiration based thereon. 
     It is yet another object of the present invention to provide a patient monitoring method for monitoring and/or quantitatively measuring a physiological characteristic of the patient that does not suffer from the disadvantages of conventional patient monitoring methods. This object is achieved by providing a method that includes the steps of providing a user interface having an interior portion adapted to communicate with an airway of a user such that substantially all gas inhaled and exhaled by the user enters the interior portion of the user interface. The user interface also has at least one vent element associated therewith for communicating the interior portion of the user interface with the ambient atmosphere outside the user interface. The vent element and the user interface define a flow element across which a pressure differential is created during inspiration and expiration. This pressure differential is the pressure difference between a first pressure within the interior portion of the user interface and the pressure of the ambient atmosphere outside the user interface. The next steps in the method of monitoring and/or quantitatively measuring a physiological characteristic of the patient include passing a gas across the flow element during inspiration and expiration, measuring a fluid characteristic resulting from the pressure differential between the pressure within the interior portion of the user interface and ambient atmosphere, and outputting a signal based on the measured fluid characteristic. In a preferred embodiment of the present invention, the method also includes using the output signal. to determine a quantitative value for the physiological characteristic of the patient. 
     It is a further object of the present invention to provide a patient monitoring apparatus and method for detecting an analyzing a patient&#39;s snore. This object is achieved by providing a patient monitoring apparatus that includes a user interface having an interior portion that communicates with the airway of a user, a device for measuring gas flow between the user and the user interface or a pressure within the user interface created by the gas flow, and a processing unit that determines a quantitative volume for an amount of gas displaced during at least a portion the user&#39;s snore based on a signal output by the measuring device. In a futher embodiment of the present invention, the processing unit determines a location of a structure in the user that causes the snore based on this quantitative volume. 
     These and other objects, features and characteristics of the present invention, as well as the methods of operation and functions of the related elements of structure and the combination of parts and economies of manufacture, will become more apparent upon consideration of the following description and the appended claims with reference to the accompanying drawings, all of which form a part of this specification, wherein like reference numerals designate corresponding parts in the various figures. It is to be expressly understood, however, that the drawings are for the purpose of illustration and description only and are not intended as a definition of the limits of the invention. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIGS. 1-2 are schematic diagrams of conventional pneumotach airflow meters; 
     FIG. 3 is a schematic diagram of a conventional nasal cannula airflow meter; 
     FIGS. 4A and 4B schematically illustrate a first embodiment of first portion of an interface meter according to the principles of the present invention; 
     FIG. 5 is a schematic diagram a second portion of the interface meter illustrated in FIG. 4; 
     FIGS. 6,  7  and  8  are charts illustrating the relationship between the output from a sensor coupled to an interface device and the actual flow through the interface; 
     FIG. 9 is a more detailed circuit diagram of the interface meter illustrated in FIG. 5; 
     FIGS. 10A and 10B are waveforms illustrating the flow and volume of patient respiration measured using the interface meter according to the first embodiment of the present invention; 
     FIG. 11A is a waveform illustrating the uncalibrated flow signal output by the sensor portion of the interface meter in the presence of user snoring, and 
     FIG. 11B is a waveform illustrating the calibrated (actual) flow signal output from the interface meter (inhale only) in the presence of snoring; 
     FIG. 12 is a waveform illustrating a flow signal produced by the interface meter of the present invention in the presence of snoring that demonstrates how the present invention is used to analyze patient snoring; 
     FIG. 13 is a schematic diagram of a circuit used to analyze a patient&#39;s snore according to the principles of the present invention; 
     FIG. 14 illustrates various configurations for a first embodiment of the interface meter according to the principles of the present invention; 
     FIG. 15 illustrates a second embodiment of an interface meter according to the principles of the present invention; 
     FIG. 16 illustrates a third embodiment of an interface meter according to the principles of the present invention; 
     FIG. 17 illustrates a fourth embodiment of an interface meter according to the principles of the present invention; 
     FIG. 18 illustrates a fifth embodiment of an interface meter according to the principles of the present invention; and 
     FIG. 19 illustrates a sixth embodiment of an interface meter according to the principles of the present invention. 
    
    
     DETAILED DESCRIPTION OF THE PRESENTLY PREFERRED EMBODIMENTS OF THE INVENTION 
     FIGS. 4A and 41 schematically illustrate a first embodiment of an interface meter  70  according to the principles of the present invention. Meter  70  includes a user interface  72  in which in this embodiment is a mask worn over the nose and/or mouth of the patient (not shown). It should be noted that the terms “user” and “patient” are used synonymously through this document. A wall  73  of user interface  72  defines an interior portion  74  that receives the user&#39;s nose and/or mouth when worn by the patient. As the user breathes into the user interface, gas is transferred between the user and interior portion  74  of user interface  72 . A plurality of holes  76  are provided in wall  73  of user interface  72  to exhaust exhaled gas firom interior portion  74  to the ambient atmosphere outside user interface  72 . See FIG.  4 A. Conversely, gas inhaled by the user enters interior portion  74  of user interface  72  through holes  76  before being inhaled by the user. See FIG.  4 B. 
     A sensor  78  is coupled to a hole  80  in the user interface to measure a fluid characteristic, such as a flow rate or a pressure differential, associated with the transfer of gas between interior portion  74  of user interface  72  and ambient atmosphere. In the embodiment illustrated in FIGS. 4A and 4B, sensor  78  is coupled to user interface  72  such that a portion of the gas entering or exiting interior portion  74  of user interface  72  passes through the sensor. The size and shape of sensor  78 , hole  80  and a tubing  82  connecting sensor  78  to hole  80  are selected so as to minimize the resistance to flow between interior portion  74  and the area outside the mask imposed by sensor  78 , hole  80  and tubing  82 . In the illustrated embodiment, sensor  78  is an air flow meter that measures the rate of flow of gas passing through the meter. 
     Holes  76  provided in user interface  72  function in much the same manner as the flow element in a conventional pneumotach. Namely, holes  76  create a slight resistance to the flow of gas into or out of interior portion  74  of user interface  72  so that during inhalation and exhalation, a pressure differential is created between interior portion  74  of user interface  72  and the ambient pressure outside the mask. This pressure differential causes gas to flow through the circuit defined by tubing  82  and sensor  78  so that the rate of flow of gas through sensor  78  can be qualitatively measured by sensor  78 . 
     For an incompressible fluid or gas, the flow of a fluid into an area must equal the flow of the fluid out of that area (Q IN  =Q OUT ). It should be noted that the terms “fluid” and “gas” are used interchangeably throughout this document. Applying this principle to interface  72 , establishes that the flow of fluid into interior portion  74  from the user during exhalation Q TOT IN  must equal the flow of fluid Q 1 , Q 2 , . . . Qn from the mask out holes  76 , assuming that there are no unaccounted for leaks in the mask or at the user/mask interface. See FIG.  4 A. Similarly, the flow to the user during inhalation Q TOT OUT  must equal the flow into the mask through holes Q 1 , Q 2 , . . . Qn, again, assuming that there are no unintentional leaks in the mask or at the user/mask interface. See FIG.  4 B. Thus, Q TOT =Q 1 +Q 2 +. . . Qn. 
     While the illustrated embodiment describes the mask interface as having a plurality of holes defined directly in the wall of the interface, it is to be understood that the present invention is not limited to this particular configuration for communicating the interior portion of the interface to the ambient atmosphere. On the contrary, the present invention contemplates that any venting structures that communicates the interior portion of the interface to the ambient atmosphere, while creating a sufficient pressure differential can be used. For example, venting can be achieved in a mask that has no exhaust holes by attaching an adapter tube to the inlet/outlet port in the mask. Holes can be provided in the adapter tube that communicate the interior portion of the adapter tube, and hence the interior portion of the mask, to ambient atmosphere. The combined mask and adapter is equivalent to user interface  72  illustrated in FIGS. 4A and 4B. It can also be appreciated that the venting structures need not be provided directly in the mask. Also, the venting mechanism, such as holes  16 , can have any shape, pattern, or number of holes so long as they function for their intended purpose—to communicate the interior of the user interface to ambient atmosphere while creating a sufficient pressure differential to produce a fluid characteristic that can be measured by sensor  78 . Also, the venting mechanism need not be defined by fixed diameter holes. On the contrary, the diameter or degree of opening of the venting structure can vary. 
     In the illustrated embodiment of the present invention, the area of hole  80  is fixed relative to the total area of the remaining holes  76  in user interface  72 , so that the flow of gas Q 5  out of the mask through sensor  78  is a known fraction of the total flow of gas out of interior portion  74  of user interface  72  during expiration. Conversely, the flow of gas Q 5  into the mask through sensor  78  is a known fraction of the total flow of gas into interior portion  74  during inspiration. Sensor  78  measures the flow of gas Q 5  passing therethrough in either direction and outputs a signal  84  indicative of that flow and of the direction of the flow through the sensor. The rate of flow through the sensor is a characteristic of the gas passing through the mask interface and, as noted above, results from the pressure differential created by the flow element, which in this embodiment is defined by providing holes directly in the mask. 
     Because the portion of gas passing through sensor  78  is a known fraction of the total amount of gas passing through holes  76  and  80 , the total flow of gas to and from the interior portion of user interface  72  can be determined from the measured flow through meter  78 . Ideally, the measured flow through sensor  78  is linearly related to the total flow Q TOT  into or out of interior portion  74  of user interface  72 , so that once the flow through sensor  78  is known, the total flow into or out of the mask can be readily determined by applying a multiplying factor to signal  84  output from sensor  78 . This can be accomplished, for example, by amplifying signal  84  by a predetermined amount. 
     It has been determined, however, that the flow measured by sensor  78  is typically not linearly related to the total flow through the user interface. This is so because the relationship between the total flow Q TOT  through interior portion  74  of user interface  72  and the measured flow through sensor  78  is dependent upon a number of factors, such as the number and size of holes  76 , the shape of interface  72 , the distance of the sensor sampling port from the pressure source, the resistance to flow through the sensor and associated components, and the location of hole  80  in the mask to which the sensor is attached. Thus, additional processing typically must be performed on signal  84  before that signal accurately indicates the actual total flow through the user interface. 
     Regardless of whether the relationship between the flow through the sensor and the total flow through the mask is linear or non-linear, as long as the structure of the interface meter does not change, the determination of the total flow Q TOT  into interior portion  74  of user interface  72  using the measured flow through sensor  78  will be substantially the same for all users regardless of the physical characteristics of the patient wearing the interface meter. Thus, once sensor  78  is calibrated for a particular interface  72  with fixed structural features, i.e., once the relationship between the output of sensor  78  to the total flow through the mask interface is established, the same interface meter  70  can be used on a wide variety of patients to measure characteristics associated with respiration quantitatively, such as the flow and/or volume of gas provided to the patient. 
     In a preferred embodiment of the present invention, sensor  78  is a mass airflow sensor, such as the AWM2100V sensor manufactured by Honeywell Inc., which outputs a range of analog voltages corresponding to a predetermined range of airflow rates through the sensor. The output from the AWM2100V is a positive and negative differential signal that corresponds to the rate and direction of flow through the sensor. The AWM2100V sensor is particularly well suited for use in measuring the amount of gas passing through a portion of user interface  72  because the AWM2100V is capable of accurately measuring a very small flow. For example, it has been determined that the pressure drop needed to generate flow across the AWM2100V at full scale is only 0.5 cm H 2 O. Because the flow of breathing gas through sensor  78  can be quite small, the pressure drop across user interface  72  needed to create a flow through the AWM2100V is also quite small. As a result, user interface  72  can have an extremely low resistance so that gas flows relatively easily into and out of interior portion  74 . Decreasing the flow resistance, which is accomplished by reducing the pressure drop across the user interface, i.e., across the flow element defined by the mask and the holes in the mask, is achieved, for example, by providing more holes in the mask andlor increasing the size of the holes so that breathing gas flows more freely between interior portion  74  and the area outside user interface  72 . 
     One advantage achieved by making the mask resistance as low as possible is to provide a good (leak free) mask seal with the patient. The lower the mask resistance, the more likely there will be no leaks at the mask/patient interface. Unintentional leaks in the mask or in the mask/patient interface can be taken into consideration in determining the total flow to and from the patient based on the measured flow through sensor  78 . For example, the leak estimation algorithms taught by U.S. Pat. Nos. 5,148,802; 5,239,995; 5,313,937; 5,433,193; and 5,632,269, the contents of which are incorporated herein by reference, can be used to determine unintentional leaks in the mask or mask/patient interface. If these unintentional leaks are minimized to an insubstantial amount, the use of leak estimation and correction techniques can be avoided. 
     It has been determined that a good seal is achieved as long as the pressure within interior portion  74  of user interface  72  is between −2 cm H 2 O to 2 cm H 2 O. The relatively low flow resistance through the AWM2100V allows the pressure within the mask to be within this range. Thus, the assumption that there are no mask leaks other than through holes  76  and  80  is valid. For example, it has been determined that the pressure in user interface  72 , even with the pressure drop caused by tubing  82  and a bacteria filter (not shown) placed between user interface  72  and sensor  78 , is 1 cm H 2 O at a flow rate of 100 liters per minute (1pm). 
     In any event, even if the pressure drop needed to generate flow across the sensor exceeds 2 cm H 2 O, unintentional leaks at the mask/patient interface can be eliminated by increasing the sealing force applied on the mask to hold the mask on the patient and/or by providing an improved seal between the mask and user, such as an adhesive seal or a larger sealing area. 
     The present invention also contemplates that sensors other than an airflow sensor can be use as sensor  78 . For example, sensor  78  can be a pressure sensor. An example of a suitable pressure sensor in a differential pressure sensor that directly measures the differential between interior portion  74  of user interface  72  and a pressure of the ambient atmosphere outside the user interface. This pressure differential, like the flow of gas through the sensor in the previous embodiment, is due to the restriction in flow between ambient atmosphere and the interior of the mask created by the flow element, which in this embodiment is defined by the holes provided in user interface  72 . Another suitable sensor is an absolute pressure sensor that measures the pressure of interior portion  74  relative to a fixed reference pressure. Any sensor, such as an airflow, pressure or quantitative temperature sensor, that is capable of measuring a fluid characteristic created by the pressure differential caused by the flow element and that is capable of outputting a signal indicative of that characteristic can be used as sensor  78 . 
     If sensor  78  is a pressure sensor, gas does not pass through the sensor. Instead, for a differential pressure sensor, one side of a diaphragm in sensor  78  communicates with interior portion  74  of user interface  72  and the other side of the diaphragm communicates with ambient atmosphere. The pressure sensor measures the pressure differential ΔP between a first pressure within the interior portion of the user interface and the ambient atmospheric pressure outside the user interface. Once this pressure differential is known, the actual flow rate Q IN  of gas passing through the mask interface can be determined, for example, using a look-up table based on the known relationship between pressure and flow, i.e., ΔP=RQ 2 . A similar approach is used of the pressure sensor is an absolute pressure sensor. 
     Regardless of whether sensor  78  is an airflow sensor, a pressure sensor, or any other type of sensor, the signal output by the sensor is typically an analog signal. If sensor  78  is an airflow sensor, signal  84  corresponds to the rate of flow of gas through the sensor and indicates the relative direction of flow. If sensor  78  is a differential pressure sensor, signal  84  corresponds to a pressure differential across the flow element and also indicates a relative direction of flow based on whether the pressure within the interface is greater or less than ambient pressure. It can be appreciated that signal  84  corresponds to a characteristic associated with respiration because signal  84  can be used to quantitatively determine a characteristic associated with respiration, such as patient flow or volume. Even in its raw, uncalibrated form, signal  84  can be used to differentiate between inspiration and expiration and/or to detect snore, and thus, corresponds to the respiratory characteristic of inspiration, expiration and/or snore. The value of analog signal  84  represents a valve of one of these characteristics of respiration. Examples, of characteristics associated with respiration that can be determined using signal  84  are discussed in greater detail below. 
     As shown in FIG. 5, signal  84  from sensor  78  is provided to an amplifier  86  and the output of amplifier  86  is provided to an analog-to-digital (A/D) converter  88 . The digital output  90  of A/D converter  88  is provided to a processor  92  that corrects for the non-linearity in the output of sensor  78  so that the signal  94  output from processor  92  is a digital signal indicative of a quantitative value for a characteristic associated with respiration. 
     For example, in one embodiment of the present invention discussed above, signal  84  from sensor  78  is a signal indicative of the rate of flow of gas through the sensor. However, as discussed above, this signal is typically not linearly related to the total rate of flow of gas Q TOT  to or from the interior portion  74  of the user interface. To correct for this non-linearity, signal  84  is provided to processor  92 . Processor  92  determines the quantitative (actual) value for the total flow of gas Q TOT  entering or exiting the interior portion  74  of user interface  72  based on signal  84 . The details as to how this is accomplished are discussed below. It is to be understood, that, based on signal  84 , processor  92  can determine characteristics associated with respiration other than flow rate. For example, by integrating the corrected flow signal, processor  92  can output a signal representing the total volume of gas V TOT  exiting or entering the interior portion of the user interface. 
     In the illustrated embodiment, a digital-to-analog converter  96  converts signal  94  from processor  92  into an analog signal  98  and provides analog signal  98  to an output and/or storage device  100 . In a preferred embodiment of the present invention, output device  100  is a monitor or an LED display, that converts signal  98  into a human perceivable output indicative of the characteristic associated with respiration, such as rate or volume of flow to and from the user. It is also preferable to store signal  98  in a memory for use in evaluating the respiratory conditions of the patient. Alternatively or in addition to the above embodiment, output  94  from processor  92  can be provided in its digital format to a digital output device  99 , such as a digital display, memory, terminal, and/or communication system. 
     As noted above, because the output from sensor  78  is typically not linearly related to the rate of flow through the mask interface over the entire range of airflows to and from the patient, processor  92  must correct for the non-linearity in the output of sensor  78 . In a preferred embodiment of the present invention, processor  92  calculates the total flow QTro entering or exiting interface  72  based on the output from sensor  78  using a lookup table, which is determined from a preestablished relationship between the output from sensor  78  and the flow through the interface. FIG. 6 is a graphical representation of this relationship. It should be understood that the graph in FIG. 6 is determined for a specific mask interface. The relationships established by the curve in FIG. 6 do not apply to all interfaces. Thus, for each different type of interface to which the processor is to be used, the relationship between the output of the sensor and the actual, quantitative value for the respiratory characteristic of interest must be determined beforehand so that this relationship can then be used to determine the quantitative value for the desired respiratory characteristic. 
     Curve  102  in FIG. 6 illustrates the relationship between the signal output by sensor  78  for a first type of mask interface and the flow through that interface. The vertical axis of the graph in FIG. 6 corresponds to the output of sensor  78 , which is typically in a range of −60 mV to +60 mV for the AWM2100V sensor. The horizontal axis represents the total flow Q TOT  into or out of the interface. The portion of curve  102  to the right of the zero flow mark on the horizontal axis represents flow in a first direction through the sensor, for example expiration, and the portion of curve  102  to the left of the zero flow point represents flow in a second direction, opposite the first direction, for example inspiration. 
     It can be appreciated from FIG. 6 that for the particular sensor and type of interface used to generate curve  102 , the output from sensor  78  is not linearly related to the rate of flow through the mask interface. This is particularly true near the zero flow rate. However, by knowing the relationship between the output of sensor  78  and the total flow, the actual, quantitative flow through the mask can be readily determined. 
     It can be further appreciated that curve  102  will have different shapes depending on the type of sensor and interface being used. However, once the relationship between the sensor output and the flow through the interface is determined, this relationship remains valid independent of the physical characteristics of the patient using the interface meter. Thus, unlike nasal cannulas, the same interface meter can be used to quantitatively determine a characteristic associated with respiration, such as the flow rate, for a wide variety of users. 
     FIG. 7 is similar to FIG. 6 in that it is a graph illustrating the relationship between the signal output by sensor  78  for a particular typ of mask interface and the flow through that interface. However, the vertical axis in FIG. 7 denotes a linearly amplified output of sensor  78 , which corresponds to signal  122  in FIG.  5 . The signal output from sensor  78  illustrated in FIG. 7 has been amplified so that the voltage range of the signal is between −5 V and +5 V. FIG. 7 includes a first curve  104 , illustrated by a solid line, that represents the voltage-total flow relationship for flow through sensor  78  in a first direction (typically during exhalation) and a second curve  106 , illustrated by a dotted line, that represents the voltage-total flow relationship for flow through sensor  78  in a second direction (typically during inhalation) opposite the first direction. In the illustrated embodiment, the output from sensor  78  is positive during expiration and negative during inspiration. It is to be understood, however, that this relationship could be reversed 
     In FIG. 7, curves  104  and  106  representing the flow during expiration and inspiration, respectively, are superimposed on one another to demonstrate that the voltage-flow characteristics are substantially the same regardless of the direction of flow through sensor  78 . Thus, the same relationship between the sensor output and the flow through the interface can be used regardless of the direction of flow through the interface, i.e., during inspiration and expiration, thereby simplifying the determination of the flow through the interface based on the measured output of sensor  78 . It is possible, however, to use separate relationships to determine a quantitative value for a characteristic associated with inspiration and a characteristic associated with expiration. 
     In a preferred embodiment of the present invention, the known relationship between the output of sensor  78  and the flow through the mask interface, as illustrated by the curves in FIGS. 6 and 7 for example, is used to generate a lookup table. This table is used to determine the actual flow through the mask interface from the output of sensor  78 . However, the present invention contemplates that techniques other than a look-up can be used to determine a quantitative measure of a characteristic associated with respiration from the raw signal output from sensor  78 . For example, once the voltage-total flow relationship for the interface meter is established, the flow can be calculated from an equation defining this relationship. For example, curves  104  and  106  in FIG. 7 can be generally defmed by the following third order polynomial equation: 
     
       
         y=−2.208×10 −6 x 3 +5.982×10 4 x 2 −2.731×10 −3 x+9.165×10 −3 , 
       
     
     where y is the linearly amplified output of sensor  78  and x is the flow into or out of the mask interface. Once y is determined by sensor  78 , processor  92  can solve for x to determine the total flow into or out of the interface. As noted above, separate equations or lookup tables can used to determine the flow through the mask during expiration and expiration, thereby improving the accuracy of the output of the interface meter if the relationship between the output of sensor  78  and flow through the interface is not the same for flow in both directions through the sensor. 
     It should be understood that the above equation and the graph illustrated in FIGS. 6 and 7 defining the relationship between the output of sensor  78  and the total flow into or out of the mask interface apply only to a particular type of interface having a predetermined structure. If, for example, more holes are added or the mask shape or size is altered, the relationship between the output of sensor  78  and the total flow into or out of the mask interface may change, requiring recalibration of processor  92  so that a different curve is used to determine the quantitative value for the desired respiratory characteristic based on the output of the sensor. 
     For example, FIG. 8 illustrates three curves  101 ,  103  and  105  defining the relationship between the pressure measured by sensor  78  and the flow through the interface for three masks having different structural characteristics. Curve  101  associated with a first mask is nearly a straight line, meaning that there is nearly a linear relationship between the pressure measured by sensor  78  and the total flow through the interface. FIG. 8 also demonstrates that if sensor  78  is a pressure monitor, the same techniques used to generate the total flow through the mask, i.e., using a look-up table or equation derived from the relationships illustrated in FIG. 8, can be used to determine the quantitative value for characteristic associated with respiration, such as flow through the interface. 
     So long as a batch of mask interfaces are manufactured with the same structural characteristics, the same calibration, i.e., voltage—flow curve, can be applied to all of the mask interfaces in that batch. By providing each processor with the same voltage-total flow relationship there is no need to calibrate each interface meter individually. In short, the interface meters of the present invention can be commonly calibrated so long as they share the same structural characteristics for the interface. The operating characteristics of the interface meter do not vary with the physical characteristics of the user, as is the case with conventional nasal cannula flow meters. 
     It is to be further understood that processor  92  can contain a number of different lookup tables and/or equations associated with a variety of different interface devices so that the same processor can be used in conjunction with a number of different types or configurations of patient interfaces, so long as the proper lookup table or voltage-total flow equation is used with the selected interface. In this embodiment, a selector is provided so that the user can select the type of interface being used with the interface meter. Processor  92  then uses the correct lookup table or equation or other technique for determining the quantitative value for a patient&#39;s physiological characteristic based on the selected interface. For example, a memory portion in processor  92  can contain three lookup tables associated with three different mask sizes. The user selects the mask size being used and inputs this selection to processor  92 . Processor  92  then uses the correct lookup table for the selected mask size to determine the quantitative value for the flow through the mask interface based upon the output from sensor  78 . 
     As discussed above, a primary function of processor  92  in the present invention is to convert the signal from sensor  78  into a signal that accurately represents the flow of breathing gas into or out of the user interface. This is necessary because it is believed to be difficult to situate the various structural elements of the interface and sensor such that the signal output from sensor  78  is linearly related to the flow through the mask to which the sensor is attached. If, however, a suitable configuration can be established, the linearizing function performed by the processor will not be necessary. Instead, processor  92  will merely provide a multiplying function to calculate the total amount of breathing gas passing through the mask interface from the known fraction of breathing gas passing through sensor  78 . Alternatively, processor  92  can be eliminated and the multiplying function can be performed using circuitry, for example, by adjusting the gain in amplifier  86  of FIG.  5 . 
     FIG. 9 is a more detailed diagram of the circuit schematically illustrated in FIG.  5 . Gas passes through sensor  78  in a first direction, as indicated by arrow  107 , during expiration and in a second direction opposite the first direction, as indicated by arrow  108 , during inspiration. Amplifier  110  sets the control for a heater that is used in the Honeywell airflow sensor to measure the flow rate therethrough. Outputs  112  and  114  of sensor  78  are positive and negative differential signals representing the flow measured by sensor  78  and are provided to a pair of amplifiers  1   16  and  118 , respectively. Outputs of amplifiers  116  and  118  are provided to a differential amplifier  120 . Amplifiers  116 ,  118  and  120  define amplifier  86  in FIG.  5  and convert the dual outputs of sensor  78  into a single analog signal  122 . In a preferred embodiment of the present invention, amplifiers  110 ,  116 ,  118  and  120  are provided on a same integrated circuit, such as the LMC660CN Quad OP-AMP manufactured by National Semiconductor. 
     Signal  122  from amplifier  86 , which is referred to as a raw or uncalibrated signal because it typically does not linearly correspond to the respiratory characteristic of interest, is provided to A/D converter  88 , such as an ADC10831 converter manufactured by National Semiconductor. Digital output  90  of A/D converter  88  is provided to processor  92 . In the illustrated embodiment, processor  92  is the PIC16C84 manufactured by Microchip Inc. Processor  92  operates at a clock speed set by oscillator  124  to calculate the flow Q TOT  entering or exiting interface  72 , for exanple, based on the output from sensor  78  as discussed above. It is to be understood, that any combination of the circuit components illustrated in FIG. 9 can be provided on a single chip. For example, A/D converter  88 , processor  92 , and D/A converter can be fabricated on the same chip for ease of manufacturing the interface meter of the present invention. 
     In one embodiment of the present invention, processor  92  uses a lookup table or a voltage-total flow equation established for a particular type of interface  72  to determine the flow Q TOT  entering or exiting the interface based on signal  90  from A/D converter  88 . In the illustrated embodiment, output  94  of processor  92  is a signal indicative of the flow entering or exiting the interface and is provided to D/A converter  96  where it is converted into a pair of analog signals  98 , which are positive and negative signals, respectively, depending on the direction of flow through sensor  78 . In the illustrated embodiment D/A converter  96  is a DAC0854 converter manufactured by National Semiconductor. A first pair of variable resistors  126  set the positive gain for the analog output of D/A converter  96  and a second pair of variable resistors  128  set negative gain. Analog signals  98  are provided to a display  100 , such as an LCD or LED display, where they are converted into an output that is capable of being perceived by humans. 
     In the illustrated embodiment, analog signals  98  are also provided to a pair of output terminals  130  and  132  so that signals  98 , which represent the actual (quantitative) flow of breathing gas passing through the interface, can be provided to external components, such as a display, data storage device, alarm system, printer, additional processing elements, and/or data communication system, such as a modem. It is to be understood, however, that any of these components could be provided within the circuitry illustrated in FIG. 9 on the same card or circuit board. 
     FIG. 10A illustrates a waveform  134  of the flow through sensor  78  during inspiration and expiration in liters per minute (1pm) generated by a computer using signal  98  taken at terminals  130  and  132  in FIG.  9 . FIG. 10A is one example of how the signal produced by processor  92  could be output in human perceivable format. FIG. 10B illustrates a waveform  136  of the tidal volume for the same flow rate of breathing gas passing through sensor  78  in liters, which is also generated by a computer using signal  98  taken at terminals  130  and  132 . Waveform  136  in FIG. 10B can be generated, for example, by integrating the flow signal  134  illustrated in FIG.  10 A. The smoothness of waveforms  134  and  136  illustrated in Figs. 10A and 10B can be improved by increasing the processing speed of processor  92 . This could be accomplished, for example, by increasing the oscillating frequency of oscillator  124  in FIG.  9 . 
     FIG. 11A illustrates a waveform  138  that corresponds to the uncalibrated analog flow signal  122  output from amplifier  86  in FIG. 8 during inhalation and exhalation. The points in FIG. 11A where waveform  138  crosses the X axis correspond to points where the patient&#39;s breathing switches from inspiration to expiration or from expiration to inspiration. Thus, these points can be used as trigger points or reference points for the application of a respiratory therapy, such as an application of positive pressure to the airway or an application of electrical stimulation to the muscles in the patient. 
     Waveform  138  was generated while the user was asleep and snoring. The rapid signal fluctuations  137  at each apex of inhalation in waveform  138  correspond to the rapid flow variations that take place in the user&#39;s respiratory system during snoring. One embodiment of the present invention detects these rapid fluctuations in the raw signal  122  output from the sensor to determine the onset, intensity and duration of snoring. This can be accomplished in a variety of ways, for example, by comparing the rate of change in signal  138  to predetermined thresholds. Because this rapid variation in flow (snore) can be easily detected from waveform  138 , the signal from sensor  78 , even if not corrected by processor  92 , can be used, for example, as a trigger for a therapy intended to relieve such snoring or as a reference point from which therapy is to begin. 
     FIG. 11B illustrates a waveform  140  that corresponds to the signal output from processor  92  based on the signal illustrated in FIG.  11 A. In other words, waveform  140  corresponds to the quantitative signal produced by processor  92  based on the raw signal illustrated in FIG.  11 A. It should be noted that FIG. 11B illustrates only the inspiration portion of the patient&#39;s flow, which is the equivalent of the output at one of terminals  130  and  132  in FIG.  9 . As with wavefonn  138  in FIG. 11A, waveform  140  in FIG. 11B exhibits relatively large and rapid fluctuations  139  during inspiration due to the patient&#39;s snoring. These rapid fluctuations can be detected in a variety of fashions, for example, by using a threshold detector, to signal the onset of snoring. It can be appreciated if the processing speed of processor  92  is increased, the rapid fluctuations in the apex of waveform  140  would be even more well defined. In fact, the sensitivity of the present invention is so great that the gas displaced by each individual snore vibration can be determined. 
     It can be further appreciated that the present invention can determine a wide variety of information based on the output from sensor  78 . For example, as noted above, by integrating the quantitative value for flow, which can be done either by processor  92  or using addition components that are either internal or external to the circuit illustrated in FIGS. 5 and 9, the interface meter of the present invention also calculates the volume V TOT  of breathing gas entering or exiting the interface. Calculating the volume V TOT  can be done in place of or in addition to determining the flow Q TOT  of breathing gas passing through the interface. The present invention contemplates providing additional digital-to-analog converters similar to D/A converter  96 , additional output devices similar to output device  100 , as well as additional output terminals similar to terminals  130  and  132  so that any additional information, such as volume V TOT , can be calculated and provided to the user, a third party, or to a data output and/or storage medium. 
     Knowing the quantitative value for the patient flow makes it possible to determine a number of physical characteristics associated with respiration. This can be done using processor  92  or other circuitry based on the signal output from processor  92  and/or, where possible, the raw signal output from sensor  78 . For example, the present invention contemplates using either the raw output of sensor  78  or the flow signal output from processor  92 , such as that illustrated in FIG. 10A, to determine the patient&#39;s breathing rate, typically in breaths per minute (bpm), minute ventilation, peak expiratory flow, inhalation time, exhalation time, and inhalation to exhalation (I:E) ratio. Also, the present invention contemplates using the volume signal, such as that illustrated in FIG. 10B, to determine the patient&#39;s exhalation volume and inhalation volume. 
     In addition to determining a number of physical characteristics, the patient flow, which is characterized by the raw signal from sensor  78  (FIG. 11A) or the quantitative signal from processor  92  (FIG.  10 A), can be used for a variety of purposes. For example, as noted above, the presence, frequency, duration or intensity of rapid fluctuations indicative of snoring can be used to trigger the application of a therapy, such as an airway pneumatic pressure support, to relieve the snoring. The detection of snoring using the patient&#39;s flow signal (raw or quantitative) can be used to auto-titrate a pressure support device. Auto-titration is accomplished, for example, by increasing the pressure provided by a pressure support device if the presence or intensity of snoring, or more generally, the presence of any event indicative of the onset of an airway obstruction, is detected, and by decreasing the pressure if such events are not detected for a predetermined period of time. This same principle can be employed with other devices, such as an electrical stimulation device, that is used to relieve the obstruction. Auto-titration can also be accomplished based on the rise time of the flow signal. An increase in rise time can indicate an increase in airway resistance, and hence, the onset of an airway obstruction. This increase in rise time can be detected by the present invention and used to increase the pressure support provided to the patient. The opposite process can be carried out if a decrease in the flow signal rise time is detected. 
     It is also possible to determine specific characteristics of a patient&#39;s snore based on the signal output from sensor  78 . For example, the frequency of the snore can be determined from the rapid fluctuations in the flow signal, either from the raw signal output from the sensor or the calibrated, quantitative signal derived from the raw signal. It is known that the frequency of the snore signal can indicate the physical location of the structure or structures in the patient causing the snore. See, for example, the article by S. J. Quinn et al. entitled, “The Differentiation of Snoring Mechanisms Using Sound Analysis,” pages 119-123 of Clinical Otolarnynhology, Vol. 21, 1996. Knowing the location of the tissue that is causing the snore is important in determining how to best treat the snore. 
     As noted above, the sensitivity of the interface meter of the present invention is great enough that it can detect the amount of gas displaced by each individual snore vibration. For example, FIG. 12 illustrates a flow signal  151  generated by the interface meter of the present invention in the presence of patient snore  153 . Snoring  153  appears in flow signal  151  as a series of high frequency oscillations  155  in flow signal  151  that oscillate about a central axis  157 . Each oscillation displaces an amount of gas corresponding to the area  159  defined by axis  157  and the curve defining the oscillation. 
     As noted above, the frequency of a snore can be used to determine the location of the structure or structures in the patient that cause the snore. In a similar manner, the amount of gas displaced by each individual snore vibration can also be used to determine the location of the snore. The amount of gas displaced by each snore vibration is related to the frequency of that snore vibration. For example, the lower the frequency of the snore, the more gas will be displaced by each individual snore vibration. Therefore, by knowing the amount of gas displaced by the individual snore vibrations, the present invention can determine the location of the structure in the patient that is causing the snore. Furthermore, because the present invention accomplishes this function based on the amount of gas displaced by each snore vibration, rather than based on the sound produced by the snore, it is more accurate and less prone to noise than conventional frequency analysis techniques. 
     In addition to determining the volume of gas displaced by each vibration in a patient&#39;s snore, the present invention also quantitatively determines the volume of the patient&#39;s entire snore signal. Quantitatively determining the volume of gas displaced by the patient snore can be accomplished, for example, as shown in FIG.  13 . The output  141  of sensor  78  is provided to a low pass filter (LPF)  142  that removes the relatively high frequency snore from the flow signal so that output  143  of low pass filter  142  corresponds to the patient flow without any snore. The flow signals  141  and  143  output from sensor  78  and LPF  142 , respectively, are provided to a subtractor circuit  144  so that the output  145  thereof is the raw, uncalibrated analog snore flow signal. Snore flow signal  145  is provided to a processor  146 , which uses a look-up table or other technique, to determine the quantitative value of the snore flow  147 . Integrating only the positive portion of the snore flow signal  147  in integrator  148  provides a volume accurate snore signal  149 , which can be used to analyze the patient&#39;s snore. It is to be understood, that only the negative portion of the snore flow signal can be integrated and the same result achieved. 
     It is to be further understood that other techniques for determining a volume accurate snore signal are contemplated by the present invention. For example, the positive portion of analog signal  145  can be integrated and then software can be used to determine the derivative, which is then converted into a quantitative flow signal to determine a quantitative snore flow signal. This quantitative snore flow signal can then be integrated to provide the volume accurate snore signal. Also, the determination of patient flow, either raw or quantitative, can be made using a conventional flow measuring device. 
     The information generated by the interface meter of the present invention can also be used in conjunction with other information about the patient&#39;s physiology to determine other characteristics of the patient. For example, if a capnometer is used to measure the patient&#39;s expired CO 2 , the flow signal and the capnometer information can be used to determine the volume of CO 2  expired by the patient. The volume of CO 2  expelled from a patient during exhalation can be determined from the following equation:            V     CO   2       =         V   MIX          [       PCO   2       P   MIX       ]               t         ,                   
     where V MIX  is the volume of gas expired by the patient, PCO 2  is the pressure of carbon dioxide in the gas expired by the patient, and P MIX  is the pressure of the gas expired by the patient. As discussed above, V MIX  can be quantitatively determined by the present invention. PCO 2  is determined using a capnometer, and P MIX  is determined using a conventional barometer. 
     Similarly, the volume of CO 2  expelled from a patient during exhalation can be determined based on the quantitative flow signal using the following equation:          V     CO   2       =       ∫     t   1       t   2              [       Q   Patient          (       PCO   2       P   MIX       )       ]                        t                         
     where: t 2 −t 1 =inhalation period and Q patient  is the flow of gas from the patient. This same principle can be used to the measure the volume of other elements expelled by the patient, such as nitrogen, O 2 , CO, water vapor and any other trace elements that can be detected. 
     Furthermore, the quantitative flow signal output by the present invention, in combination with other sensing devices, can be used to determine a patient&#39;s effective minute ventilation, effective tidal volume, airway dead space, and alveolar volume using conventional techniques. If the patient&#39;s arterial PCO 2  is also known, further information, such as the physiologic V D /V T , physiologic dead space, and alveolar dead space can also be determined using conventional techniques. 
     While the items discussed above describe physiological parameters that are capable of being measured using the present invention, either alone or in combination with other measuring devices, and processes that can be performed or controlled based on the information produced by the present invention, this list is not intended to be exclusive. On the contrary, the present invention can be used to determine any characteristic about a patient that can be derived from the information output by sensor  78  and/or processor  92 . Also, the present invention can be used in conjunction with any process that is controlled or requires information of the type produced by the present invention, either directly from the signal output by sensor  78  or processor  92 , or indirectly when used combination with other measured physical characteristics. 
     Although the embodiment of the present invention discussed above has been described for use with a mask-like user interface, it is to be understood that a wide variety of user interfaces, which are discussed in greater detail below, can be used in conjunction with the interface meter of the present invention. Also, the mask serving as a user interface in the embodiment illustrated in FIGS. 4A and 4B can have a wide variety of configurations. For example, user interface  74  can be a nasal mask that covers only the user&#39;s nose, a total face mask that encompasses the user&#39;s entire face from chin to forehead, or a helmet type mask that encapsulates the user&#39;s head. It should also be understood that the term “user interface” is not limited to the mask-like structure illustrated in the figures. Quite the contrary, the “user interface” of the present invention can include structures that attach to the mask-like portion. User interface  72  and tube  82  can be made from any suitable material. In addition, a bacteria filter can be provided anywhere along the length of tube  82 . It is preferable to use a bacterial filter and tubing  82  that have a sufficiently low resistance so that a suitable amount of gas flows through sensor  78 . 
     FIG. 14 illustrates an. example of a plurality of interface meters  150 ,  152 , and  154  according to the first embodiment of the present invention. Each interface meter includes a user interface, which in this embodiment is a mask-type interface, a venting element that communicates the interior of the interface to ambient atmosphere and a sensor for measuring a fluid characteristic, such as pressure or flow, resulting from the pressure differential between the interior of the mask and ambient atmosphere created by the venting element. 
     Interface meter  150 , for example, includes a user interface  158  similar to the interface schematically illustrated in FIGS. 4A and 4B. The venting element in interface meter  150  is a plurality of holes  160  provided in user interface  158 . A hollow tube  162  having one end selectively coupled to user interface  158  and a second end selectively coupled to a housing  164  communicates the interior of user interface  158  with a sensor (such as sensor  78  in the previous figures) in housing  164 . Housing  164  also contains the circuitry illustrated in FIGS. 5 and 9 associated with the sensor. In the illustrated embodiment, a bacteria filter  166  is provided between the first and second ends of tube  162 . 
     Housing  164  includes a display  167  that corresponds to output device  100  in FIGS. 5 and 9 and an on/off activating mechanism  168 . Housing  164  also includes a selector  170  so that the user can manually select the type of interface being coupled to housing  164 . As discussed above, this enables the processor to use the appropriate look-up table for determining the flow through the interface. Selector  170  and on/off activating mechanism can be any suitable input device for controlling the circuitry and/or processing elements of the present invention. In the illustrated embodiment, interface meter  150  is AC powered. It is to be understood, however, that any suitable power supply, such as batteries, can be used to energize the interface meter. 
     Interface meter  150  also includes a wireless communication link  169  for communicating with a base unit  17 . Any suitable wireless communication system, such as an rf communication link or a modem and land line telephone, cellular, and/or satellite communication system is contemplated by the present invention. 
     Interface meter  152  is similar to interface meter  150  except that user interface  174  in interface meter  152  does not have holes defined therein. An example of such masks are the nasal mask sold by RESPIRONICS Inc. under the trademark “GOLD SEAL”™ and the full face mask that covers the nose and mouth sold by RESPIRONICS Inc. under the registered trademark “SPECTRUM”®. The venting element that communicates the interior of the interface to ambient atmosphere is an attaching element  176  that selectively couples to a hole defined in user interface  174 . Attaching element  176  includes a plurality of holes  178  that communicate the interior of user interface  174  to ambient atmosphere. A headgear  180  attaches the user interface to the patient. As with interface meter  150 , a hollow tube  162  couples a sensor in housing  164  to the interior of user interface  174 . Interface meter  152  communicates information with base unit  172  via a hard wired link  182 . 
     Interface meter  154  includes a first user interface  184  and a second user interface  186 . Unlike interface meters  150  and  152 , the interior of user interfaces  184  and  186  communicate directly with a sensor  188  and  190 , respectively, that is provided on, in or at the user interface itself, thereby eliminating the hollow tube of the previous embodiments. Sensor  188  in interface meter  154 , like sensor  78  in the previous embodiments, measures a fluid characteristic, such as the flow therethrough or the absolute pressure within the mask or the pressure in the mask relative to ambient atmosphere, and outputs a signal via a wire  192  to a processor within housing  164 . Sensor  190  performs a similar function except that there is a wireless communication  194  between sensor  190  and housing  164 . It is to be further understood that the sensor can be provided within the mask. 
     Base unit  172  processes the information provided by each interface. For example, the signal from each interface meter can be the raw flow signal from the sensor (sensor  78  in FIG. 5) or the quantitative flow signal from the processor (processor  90  in FIG.  5 ). Base unit  172  can use these signals, as discussed above, to determine a variety of respiratory characteristics for each patient. Base station  172  can communicate this information, either wirelessly or via wires, to other information processing devices. The illustrated embodiment of the present invention also contemplates providing information from base station  172  to various-outputlstorage devices, such as a display  196 , a printer  198 , and a storage device  200 . 
     The multiple interface meter system illustrated in FIG. 14 is particularly suited for the hospital or sleep lab environment where multiple patients are monitored by one caregiver. By employing wireless communications between the components of the interface meter, the respiratory characteristics of a patient can be monitored from a remote location, such as the patient&#39;s home or while the patient is in transit to a hospital. 
     While the above embodiment for the interface meter uses a mask-like interface that communicates with the airway of the user, the present invention is not limited to a mask type interface. Quite the contrary, any interface that communicates with the airway of the user is contemplated by the present invention. For example, in a second embodiment of the present invention, as shown in FIG. 15, a pair of nasal prongs  202  replace user interface  72  of FIGS. 4A and 4B. In all other respects, the second embodiment of the present invention and the first embodiment discussed above are the same. 
     Nasal prongs  202  include protruding portions  204  that insert into the nares of the user. The diameters at each proximal end  206  of protruding portions  202  are sized to seal the nares into which the protruding portion are inserted so that gas does not leak around the periphery of proximal end  206  of protruding portion  204 . An opening  208  is defined at a distal end  210  of each protruding portion for communicating an interior portion of the protruding portion with a nasal cavity of the user. At least one vent hole  212  is provided in the proximal end of protruding portions  204 . Vent holes  212  perform the same function as holes  76  in the user interface illustrated in FIGS. 4A and 4B. A sensor (not shown) that performs the same function as sensor  78  in FIGS. 4A and 4B is coupled to the interior portion of both protruding portions  204  via a hollow tube  214  and short, connecting tubes  216 . 
     FIG. 16 illustrates a third embodiment of an interface meter according to the principles of the present invention. The interface meter in this embodiment includes a incubator chamber  220  as the interface that communicates with the airway of the user. Vent elements  224  are provided in the wall of incubator chamber  200  for communicating the interior portion of the chamber with ambient atmosphere, in the same manner as holes  76  in user interface  72  of FIGS. 4A and 4B. A sampling port is provided in the wall of chamber  220  to communicate a sensor  224  with the interior of the chamber via a hollow tube  226 . As with the previous embodiment, tubing  226  can be eliminated and the flow or pressure sensor provided in direct communication with the interior of chamber  220 . Sensor  224  corresponds to the circuitry illustrated in FIG. 5 and 9. 
     Typically, a breathing gas, such as oxygen or an oxygen mixture, is delivered to the incubator chamber via a gas supply  222 . As a result, there is a constant leak from the chamber through vent elements  224 . This leak will offset the raw flow or pressure signal from the sensor, as well as the quantitative flow signal output from the processor, so that the flow or pressure signal and quantitative flow signal no longer varies about a zero flow or zero pressure axis. Instead, these signals will fluctuate about a level that corresponds to the leak from the chamber, which corresponds to the flow of breathing gas to the chamber via gas supply  222 . In the illustrated embodiment, the processor accounts for this offset caused by the supply of breathing gas so that the output from the processor in sensor  224  is a true representation of the patient&#39;s inspiration and expiration. This can be done, for example, by subtracting the leak, once determined, from the quantitative signal output by the processor. Thus, the present invention outputs a quantitative representation of the flow through the chamber even in the presence of a constant supply of gas to the chamber. 
     FIG. 17 illustrates a fourth embodiment of an interface meter  230  according to the principles of the present invention. This embodiment is similar to the embodiments illustrated in FIGS. 4A and 4B except that a breathing gas supply provides a constant supply of breathing gas, such as oxygen or an oxygen mixture, to the interior of mask  232 . This embodiment of the present invention is particularly advantageous in that it permits a wide variety of diagnostic information to be garnered from the patient while the patient is being provided with a breathing gas, which is a common medical procedure. 
     Mask  232  in FIG. 17 includes a first port  234  into which breathing gas from a suitable supply, such as an oxygen tank  233  or oxygen concentrator, is supplied and a second port  236  that communicates a sensor  238  to the interior portion of the mask. It is to be understood that the breathing gas need not be directly provided to the user interface, as shown in FIG.  17 . On the contrary, the breathing gas can be provided to the tube connecting sensor  238  to interface  232 , thereby avoiding the need to provide two ports in the mask. 
     In the illustrated embodiment, sensor  238  corresponds to the circuitry illustrated in FIGS. 5 and 9 of the previous embodiment As with the previous embodiments, a plurality of holes  240  are provided in the mask so that the mask defines a flow element. It is to be understood, however, that any venting system for communicating the interior of the mask with ambient atmosphere, while creating a pressure drop across the flow element, is contemplated by the present invention. 
     As with the third embodiment illustrated in FIG. 16, the constant supply of a breathing gas to mask  232  produces a substantially continuous leak from the mask. This supply of gas will skew the signals output from the sensor or from the processor so that these signal do not fluctuate about zero during the patient&#39;s breathing cycle. Instead, these signals will have a bias that corresponds to the flow of breathing gas into the mask and hence, the leak from the mask. As in the previous embodiment, the present invention compensates for this bias, for example, by subtracting the known leak from the signal output by the sensor or processor. Of course, any other technique for correcting the signals output from the sensor or processor to account for this leak are also contemplated by the present invention. For example, the vertical axis in the waveform diagram for the patient&#39;s quantitative flow can be re-labeled so that the bias level caused by the leak is defines as the effective zero flow axis. The flow signal will fluctuate about this effective zero flow axis if a constant supply of gas is delivered to the mask. 
     FIG. 18 illustrates a fifth embodiment of an interface meter according to the principles of the present invention. This embodiment is similar to the embodiment discussed above with respect to FIG. 17 except that a positive pressure device  244  supplies a breathing gas to an interface  246  via a breathing circuit  248 . In the illustrated embodiment, interface  246  is a mask interface that covers the user&#39;s nose or the user&#39;s nose and mouth. There are no holes in the mask to serve as a venting element. Instead, an adapter device  250  is coupled to the mask. Adapter device  250  attaches an end of breathing circuit  248  to mask  246 . Adapter device  250  also includes at least one hole  252 , which can have a variety of configurations, that communicates the interior portion of mask  246  to the ambient atmosphere. A hollow tube  254  is coupled to a port defined in adapter device  250  to communicate an a sensor  256  with the interior of mask  246 . Sensor  256  performs the same function as the circuit illustrated in FIGS. 5 and 9. It is to be understood, however, that sensor  256  can be coupled to other portions of the mask a breathing circuit. For example, sensor  256  can be coupled directly to a pick-off port defined in mask  246  or can be provided along breathing circuit  248 , so long as sensor  256  measures a fluid characteristic associated with the pressure differential caused by venting the interior of the mask to ambient atmosphere. 
     The present invention also contemplates that holes  252  can be removed from the interface and/or breathing circuit so that there is no venting element between the positive pressure device and the patient. Instead, the gas inlet to the positive pressure device serves as the primary venting element, i.e., gas inlettout, for the patient circuit. During inhalation, the patient&#39;s inhalation and the pressure provided by the positive pressure provide breathing gas to the patient. During exhalation, the force of the patient&#39;s expiration causes gas to be backed up into the positive pressure device and out of the gas inlet provided thereon. 
     As with the third and fourth embodiments illustrated in FIGS. 16 and 17, the constant supply of a breathing gas to mask  246  produces a substantially continuous leak from the mask via holes  252 . As in the previous embodiments, the present invention compensates for the bias caused by this supply of gas, for example, by subtracting the known leak from the signal output by the sensor or processor. If bi-level pressure or variable pressure is provided by positive pressure device  244 , compensations techniques such as those discussed above, can be employed to correct for the bias imposed by the variable pressure. 
     Although FIGS. 17 and 18 illustrate providing a supply of breathing gas to a mask-type patient interface, it is to be understood that a breathing gas, such as oxygen, can be supplied to other types of patient interfaces, in addition to the incubation chamber illustrated in FIG. 15, according to the principles of the present invention. FIG. 19, for example, illustrates a nasal prong patient interface that is similar to that illustrated in FIG. 15 except that nasal prong interface  260  in FIG. 19 includes a supply of oxygen to the patient. In all other respects, the sixth embodiment of the present invention and the embodiment illustrated in FIG. 14 are the same. 
     In the illustrated embodiment, nasal prongs  260  include protruding portions  262  that insert into the nares of the user and opens are provided in each end of the protruding portions. The proximal end of protruding portions  262  include at least one vent hole  264  that perform the same function as vent holes  212  in the nasal cannula illustrated in FIG. 15. A sensor (not shown) that performs the same function as sensor  78  in FIGS. 4A and 4B is coupled to the interior portion of both protruding portions  262  via a first hollow tube  266  and short, connecting tubes  268 . A breathing gas, such as oxygen, is provided to the interior of protruding portions  262  via a second hollow tube  270  and short, connecting tubes  272 . The constant supply of breathing gas to nasal prongs  262  produces a substantially continuous leak from the protruding portions holes  264 . As in the previous embodiments, the present invention compensates for the bias caused by this supply of gas, for example, by subtracting the known leak from the signal output by the sensor or processor. 
     The present invention also contemplates that a breathing gas can be provided to the tubing connecting the nasal prong interface to the sensor. This embodiment is advantageous in that it eliminates that need for two hollow tubes and two connecting tubes to be connected to each protruding portion of the nasal prong interface. 
     Although the invention has been described in detail for the purpose of illustration based on what is currently considered to be the most practical and preferred embodiments, it is to be understood that such detail is solely for that purpose and that the invention is not limited to the disclosed embodiments, but on the contrary, is intended to cover modifications and equivalent arrangements that are within the spirit and scope of the appended claims. For example, while processor  92  and  146  have be described in terms of an integrated circuit that is carries out a predetermined program, it is to be understood that these functions could be accomplished using hardwired circuit elements.