Estimation of Flow and Detection of Breathing in CPAP Treatment

This invention relates to methods and apparatus for the estimation of flow and the detection of breathing (respiration) in continuous positive airway pressure (CPAP) treatment.

The administration of CPAP is common in the treatment of Obstructive Sleep Apnea (OSA) syndrome and Upper Airway Resistance syndrome. The fundamental disclosure of CPAP is made in the specification of PCT AU82/00063, published under WO 82/03548. CPAP treatment effectively acts as a pneumatic splint of a patient's upper airway by the provision of a positive air pressure at approximately 10 cm H.sub.2 O, although pressures in the range of approximately 2-20 cm H.sub.2 O are encountered. More sophisticated forms of CPAP, such as bi-level CPAP and autosetting CPAP, are described in U.S. Pat. No. 5,245,995. Common to all forms of CPAP is a nose, mouth or face mask fitted to a patient having connection via a flexible air delivery tube to an air flow generator.

In more advanced forms of CPAP treatment, the measurement of airflow in the air delivery tube is used to detect the average volume breathed by the patient and to determine whether that person is inhaling (inspiring) or exhaling (expiring). Currently this is done using an in-line sensor to measure flow directly, or by measuring the pressure drop across a restriction in the air delivery tube (or alternatively, the pressure drop along the air delivery tube). These methods require the use of additional transducers and, in some cases, additional wiring or tubing to connect the transducer to the appropriate point in the circuit.

In this specification any reference to a "mask" is to be understood as embracing a nose, mouth or combination nose and mouth (face) mask suitable for the administration of CPAP treatment. Furthermore, a"mask" can include nasal prongs (cannulae) that are inserted into the nares at the entrance to the airway.

It is a preferred object of the present invention to offer an improvement over existing methods at least by removing the need for additional transducers, wires or tubing to measure airflow. It is a further preferred object of the invention to detect respiration, and particularly inspiration and expiration.

Therefore the invention discloses a method for detecting respiration during the administration of continuous positive airway pressure (CPAP) treatment to a patient by CPAP apparatus, said CPAP apparatus including a flow generator for supplying air to a `mask` via a gas delivery tube, said method comprising the steps of:

measuring a parameter indicative of power consumed by said flow generator to derive a power signal;

filtering said power signal to remove non-respiratory components; and

sensing a change in said filtered signal as indicative of respiration.

The invention further discloses a method for detecting inspiration and expiration during the administration of CPAP treatment to a patient by CPAP apparatus, said CPAP apparatus including a flow generator supplying air to a `mask` via a gas delivery tube, said method comprising the steps of:

measuring a parameter of said flow generator to derive an operating signal;

calculating from said operating signal a signal representative of flow due to respiration;

sensing a change in said flow signal;

comparing said change with a first threshold to detect start of inhalation if said first threshold is exceeded; and

comparing said change with a second threshold to detect start of exhalation if said second threshold is exceeded.

The invention yet further discloses apparatus for the detection of respiration during the administration of CPAP treatment to a patient by CPAP apparatus, the CPAP apparatus including a flow generator for supplying air to a `mask` via a gas delivery tube, said apparatus comprising:

means for measuring a parameter indicative of power consumed by said flow generator and deriving a power signal thereof;

means for filtering said power signal to remove non-respiratory components; and

means for sensing a change in said filtered signal as being indicative of respiration.

The invention yet further discloses apparatus for detecting inspiration and expiration during the administration of CPAP treatment to a patient by CPAP apparatus, said CPAP apparatus including a flow generator supplying air to a `mask` via a gas delivery tube, the apparatus comprising:

means for measuring a parameter of said flow generator to derive an operating signal;

means for calculating from said operating signal a signal representative of flow due to respiration;

means for sensing a change in said flow signal;

means for comparing said change with a first threshold to detect start of inhalation if said first threshold is exceeded; and

means for comparing said change with a second threshold to detect start of exhalation if said second threshold is exceeded.

Embodiments of the invention will now be described with reference to the accompanying drawings, in which:

FIGS. 1a and 1b are schematic diagrams of two flow detection systems;

FIG. 2 is a schematic diagram of an inhalation/exhalation detector; and

FIGS. 3a and 3b are graphs of airflow, and the time derivative of airflow.

A CPAP flow generator typically is realised by a small turbine powered by an electric motor. The speed of the motor is controlled by a feedback loop in that either the motor speed or the air pressure in the breathing circuit is measured and an error signal generated to increase or decrease drive to the motor or other regulating device, thus attempting to maintain either a constant motor speed or a constant CPAP treatment delivery pressure. Airflow in the breathing circuit is normally dictated by one or more of:

a) a constant, deliberate leak situated near the mask (or at the end of the nasal prongs) to ensure that patient rebreathing is minimal;

b) the patient's respiration; and

c) any additional unintended leaks.

With changing airflow due to respiration, parameters such as the turbine motor's speed and/or current will alter in a manner defined by the motor or solenoid-operated spill valve control circuitry. The solenoid-operated spill valve has the function of maintaining constant delivery pressure.

Signals can be derived from the motor speed and power measurements, or from the spill valve position and power measurements. These measured signals vary cyclically with the patient's respiration. In general they also bear a non-linear relationship to the actual volumetric flow which can, if required, be linearised using previously determined pressure/flow/speed characteristics of the turbine system to give a volumetric measure of patient respiration (flow).

Without linearisation, the signal can be used to detect the points at which the patient starts to inhale and exhale. These points are crucial to the correct working of bi-level CPAP machines which present to the patient a smaller pressure at exhalation than inhalation, and in machines which seek to automatically vary the CPAP pressure in line with the patient's clinical needs. The cyclical variation of the respiratory component in the signal is extracted by identifying the more slowly changing non-respiratory component and subtracting this from the original signal. Alternatively, the minimum value of the cyclic component can be detected and stored and thereafter subtracted from the original signal.

Embodiments implementing this methodology, are shown in the accompanying drawings. In FIGS. 1a and 1b, a flow generator 10 comprises an electric motor 12 that drives a turbine 14. The turbine 14 provides a supply of air for the administration of CPAP treatment, transported to the patient's mask by an air delivery tube (not shown). The motor 12 receives electrical power from a power supply 16. Also comprising a part of the flow generator 10 is a motor controller 18, which issues a control signal 20 to the power supply 16 to control the motor speed and thus the speed of the turbine 14 and, in turn, the flow rate in the air delivery tube. Furthermore, a speed feedback signal 22 is input to the motor controller 18 to provide a signal upon which speed regulation can be based.

As shown in FIG. 1a, one of the interconnecting wires between the power supply 16 and the motor 12 is provided with a current sensing resistor 24. This resistor therefore detects current demanded by the motor 12, represented as the voltage appearing across the resistor 24. That voltage is sensed and input to a differential amplifier 26, thus producing an output signal representative of motor current (and thus also motor power). This output signal is then provided to a low-pass filter circuit 28 having an upper limiting frequency of, say, 20 Hz. The low-pass filter circuit 28 removes high frequency electrical noise, and also tends to average the signal. The filtered signal is then passed through a high-pass filter 30, typically with 0.5 Hz cut-off, to remove the non-respiratory components.

FIG. 1b shows an alternative arrangement to that shown in FIG. 1a. Optionally, the voltage drop across the motor 12 also can be measured via a differential amplifier 55. The output voltage thereof is then multiplied with the motor current signal previously derived by the differential amplifier 26 by a multiplier 27 to produce a measure of the time instantaneous power consumed by the motor. This time instantaneous power signal then is provided to the low pass filter circuit 28, and processing proceeds as described above. In many implementations, current alone will be a sufficient indicator of motor power, however in other cases the real instantaneous power will advantageously be determined as shown in. FIG. 1b.

The output of the high-pass filter circuit 30, shown as output signal 32, is supplied to an inhalation/exhalation detector (not shown) which functions to determine the start of inhalation and exhalation by locating the sudden changes of signal amplitude and/or polarity following a segment having a low rate of change with time and being of a given minimum duration, typically 0.2 second. FIG. 3a shows a typical flow characteristic, and identifies both the start of inhalation and exhalation events, that are detected by the methodology described. Therefore, the output signal 32 (FLOW) provides an indication of the instances of inhalation and exhalation, and these points are critical to the correct working of bi-level CPAP machines, for example.

Alternatively, the output from the low-pass filter 28 can be input to a negative peak detector (not shown) which triggers a sample and hold circuit (also not shown) to latch the minimum point of the respiratory fluctuation. The detected minimum point of respiratory fluctuation can be updated periodically, typically as an exponentially past-time weighted sum of present and previous minimum measurements.

FIG. 2 shows another method and apparatus for the detection of inhalation and exhalation events. As shown, a band-limited differentiator 42 receives the FLOW signal 32, derived from the high-pass filter circuit 30. The output from the differentiator 42 is supplied to a pair of comparators 44,46. This output signal, as a time differential, is represented in FIG. 3b. Associated with each comparator 44,46 is a threshold generator 48,50. These generators 48,50 respectively have a threshold value for detecting inhalation and exhalation, as is shown in FIG. 3b. Therefore, the comparator 44 will compare the time differentiated flow signal output from the differentiator 42 with the threshold value supplied by the threshold generator 48, and when that value is exceeded (in the negative sense) an inhalation detection signal 52 is output. The other comparator 46 functions in a similar way to output an exhalation detection signal 54 in the event that the time differentiator flow signal exceeds (in a positive sense) the threshold value set by the threshold generator 50.

In another embodiment, the low-pass filtered. signal (the output of the low-pass filter 28) derived from the motor current can be digitised and input to a microcomputer where the subsequent signal processing described above is performed digitally. In this implementation threshold levels and decision times can be varied dynamically to track the changing shape of the patient's respiratory flow.

If the output signal 32 is linearised by a linearisation element 34 to produce a linearised flow signal 36 (FLOW'), then, as previously described, this gives a volumetric measure of patient respiration by comparison with a previously determined pressure/flow/speed characteristics of the turbine system.

Modifications and alterations, as would be apparent to one skilled in the art, can be made without departing from the basic inventive concepts. All such modifications and alterations are to be considered within the scope of the present invention, embodiments of which have been hereinbefore described.