WO1997018003A1

WO1997018003A1 - System for detecting target respiratory flow rates

Info

Publication number: WO1997018003A1
Application number: PCT/US1996/018042
Authority: WO
Inventors: Warren J. Warwick; Leland G. Hansen
Original assignee: Regents Of The University Of Minnesota
Priority date: 1995-11-15
Filing date: 1996-11-12
Publication date: 1997-05-22
Also published as: AU7726096A

Abstract

An air flow system (10) for detecting target flow rates as a function of an opening (32, 42) in a flow restrictor (18, 20). A negative pressure sensor (50, 52) is triggered in response to a negative pressure condition in a chamber (23, 25) in the flow restrictor (18, 20). Target flow rates are preselected by using a flow restrictor (18, 20) having a particular size restrictor. Triggering of the negative pressure sensor (50, 52) may be used for controlling various medical treatments. A signal generated when a target flow rate is reached may be used for administering aerosols during inhalation, activation of HFCC therapy during exhalation, or to control various other medical treatments. The air flow system (10) allows for selection of a narrow flow rate range for administering aerosols during inhalation to target particular areas in the respiratory tree as major deposition sites for the aerosols.

Description

SYSTEM FOR DETECTING TARGET RESPIRATORY FLOW RATES

Field of the Invention The present invention relates generally to a flow system that detects target flow rates in a patient's respiratory cycle, and more particularly, to a system in which one or more preselected target flow rates are detected using flow restrictors.

Background of the Invention

Various types of medical treatments that depend on the respiratory cycle of a patient are used for persons suffering from respiratory diseases, or persons who are indisposed or unconscious. These medical treatments include, for example, aerosol generators, chest compressors, respirators, and ventilators. For example, aerosol therapy is widely used due to its recognized clinical advantages over intravenous or oral drug therapy. The advantages include: higher therapeutic effect with a given dose of drug, fewer side effects, and more rapid action ofthe drug. Some drugs designed to treat airway dysfunction and parenchymal disease are more effective when delivered by an aerosol. One requirement of effective aerosol therapy is to deposit the aerosols at the appropriate site of the respiratory tract. The flow rate at which an aerosol is inhaled influences the site of aerosol deposition in the lungs ofthe patient. During rapid inhalation, aerosols are more likely deposited in the oropharynx and large conducting airways, and poor delivery to the peripheral lung. A lower inspiration flow rate, preferably a slow, steady inhalation, results in greater dose of the drug and enhanced aerosol deposition in the lungs. Therefore, aerosol generators should ideally be activated within target minimum and maximum respiratory flow rates to achieve the optimum therapeutic effect. Research detailing the efficacy of aerosol drug delivery as a function of inhalation flow rate is set forth in the following: M. Dolovich, Clinical Aspects of Aerosol

Physics. 36 Respiratory Care 931 (1991); R. Kacmarek & D. Hess. The Interface between Patient and Aerosol Generator. 36 Respiratory Care 952 (1991); S. Newman, D. Pavia, & S. Clarke, How should a pressurized β-adrenegic bronchodilator be inhaled?. 62 Eur. J. Respir. Dis. 3 (1981); and S. Newman, Aerosol Deposition Considerations in Inhalation Therapy. Department of Thoracic Medicine, Royal Free Hospital, London England, all of which are hereby incoφorated by reference.

Typically only ten percent of aerosols reaches the lungs ofthe patient. The other 90 percent is lost to the environment or remains as droplets in the aerosol generating device, adding to the cost ofthe therapy. Concerns have been raised about the health risk to primary care givers exposed to the aerosol lost to the environment, and the cost effectiveness of aerosol delivery systems. Large investments have been made in aerosol drug research but few resources have been allotted to applied research on more effective ways of administering aerosol therapy.

Prior art devices use variable flow rate sensors to trigger a nebulization stream. These devices are expensive, often requiring processing ofthe output from the sensor to an analogue comparator circuit or to digital electronics. Some of these devices attach the variable flow rate sensors directly to the mouthpiece used by the patient, adding to the weight ofthe unit and the possibility of damage due to mishandling. Additionally, the sensitivity of some of these devices is such that an increased respiratory flow rate is required, that in practice cannot always be achieved especially by children and indisposed persons. Chest compressor devices used to loosen and eliminate mucus from the lungs, such as high frequency chest compressors (HFCC), are another example of medical treatments that are dependant on the respiratory cycle of a patient. The degree of airway obstruction due to mucus in the lungs affects the site of aerosol deposition in aerosol therapy. Clearance of mucus from the respiratory tract in healthy individuals is accomplished primarily through air flow and ciliary transport, accompanied with sighing and coughing. Failure of these natural systems result in accumulation of mucus which must be removed to reduce the build-up and the risk of infection. Treatment involves aerosol therapy to obtain bronchial drainage in combination with daily pounding on the chest wall to loosen mucus for expectoration. Some HFCC systems rely on a patient hand held control to activate HFCC during the patient's respiratory cycle. Although activation of HFCC is often found to be most effective during a particular portion ofthe respiratory cycle, it is difficult using present HFCC systems for the patient to identify that particular portion ofthe respiratory cycle in order to activate HFCC precisely.

Summary of the Invention The present invention is directed to an air flow system for detecting one or more target flow rates in a respiratory cycle of a subject.

The target flow rates may be used for controlling various medical treatments. A signal generated when a target flow rate is reached may be used for administering aerosols during inhalation, activation of HFCC therapy during exhalation, or to control various other medical treatments. The air flow system allows for selection of a narrow flow rate range for administering aerosols during inhalation to target particular areas in the respiratory tree as major deposition sites for the aerosols.

The air flow system ofthe present invention is activated by detecting a negative pressure condition in an opening of a flow restrictor. Target flow rates are preselected and thereby quantified by using a restrictor opening of a particular size. When a low or minimum flow rate is required, a small restrictor opening is used while a higher or maximum flow rate requires a larger opening.

The system, in effect, operates as a switch to activate various medical treatments by using predetermined target air flow rates, which are a function ofthe size of the opening of the flow restrictors.

The air flow system includes a patient mouthpiece and at least a first interchangeable flow restrictor. The first flow restrictor has a first chamber fluidly coupled to a first air flow resistance opening and the patient mouthpiece. The first air flow resistance opening has a first cross sectional area less than the cross sectional area ofthe first chamber. A first sampling port is located proximate the first chamber in fluid communication with a first negative pressure sensor so that a first target flow rate generates a negative pressure condition in the first chamber that activates the first negative pressure sensor. The target flow rate is a function of the first cross sectional area.

One or more additional interchangeable air flow restrictors may be releasably attached to the distal end of the first flow restrictor to detect other target flow rates. Alternatively, additional flow restrictors may be inteφosed between the first air flow restrictor and the mouthpiece.

The cross sectional areas are generally in the range of about 6 to 130 mm². The sampling ports are preferably fluidly connected to the sensors by tubes so that the mouthpiece can be remote from the sensors.

When the air flow system is configured to detect target inhalation flow rates, the cross sectional area of the flow restrictor connected to the mouthpiece is preferably less than the cross sectional areas of subsequent flow restrictors. In an embodiment in which two flow restrictors are attached to the mouthpiece to detect target inhalation flow rates, the first target flow rate is generally a minimum target flow rate and the second target flow rate is a maximum target flow rate.

The present invention is also directed to controlling a medical treatment in response to target respiration flow rates. A medical device may be activated or deactivated in response to the activation or deactivation of one or more negative pressure sensors. In one embodiment, the medical device may include an aerosol generator fluidly connected to the mouthpiece that is activated in response to activation of the first negative pressure sensor. The medical treatments may be deactivated in response to activation ofthe second negative pressure sensor. Alternatively, the activation of the second negative pressure sensor may signal an alarm. The medical device may include a high frequency chest compression device activated at the end ofthe inspiration cycle. In an alternate embodiment, a high frequency chest compressor may be used in combination with the aerosol generator. A ventilator may also be activated in response activation or deactivation of one of the negative pressure sensors. Brief Description of the Drawings

The present invention will be described in greater detail below, with reference to the accompanying drawings, in which

Figure 1 is a schematic illustration of an exemplary system for controlling an aerosol generator and/or a high frequency chest compressor as a function of target respiratory flow rates;

Figure 2 A illustrates an exploded, cross-sectional view of two exemplary flow restrictors for detecting target respiratory flow rates;

Figure 2B illustrates an exploded cross-sectional view of three exemplary flow restrictors for detecting target respiratory flow rates;

Figure 3 is a schematic illustration of an exemplary system for controlling a generic medical treatment as a function of target respiratory flow rates;

Figure 4 is a graph of an oscillatory curve representing an exemplary patient's respiratory cycle; and Figure 5 is a graph illustrating respiratory flow rates as a function of cross-sectional areas of openings in flow restrictors.

Detailed Description of the Preferred Embodiments

Figure 1 is a schematic illustration of an exemplary air flow system 10 for controlling an aerosol generator 12 as a function of a minimum and a maximum target respiratory flow rate. The aerosol generator 12 is fluidly connected to a T- tube 14. The T-tube 14 has an opening 16 for a mouth piece at one end and a minimum flow restrictor 20 at the other end. A maximum flow restrictor 18 is fluidly connected to the opposite or distal end ofthe minimum flow restrictor 20.

Turning now to Figure 2, the minimum flow restrictor 20 has a one degree tapered opening 22 sized for frictional engagement with one end of the T- tube 14 (see Figure 1). At base 24 ofthe opening 22 is a negative pressure sensing port 26. A tube 28 is pressure-fit into the port 26. The tube 28 is preferably peφendicular to the direction ofthe air stream "S" and protrudes slightly above the inside surface of the opening 22. The protrusion ofthe tube 28 serves as a stop for the connection with the T-tube 14 (see Figure 1).

The opposite end 30 ofthe minimum flow restrictor 20 has an external one degree taper for frictional engagement with the maximum flow restrictor 18. The minimum flow restrictor 20 has an air flow resistance opening 32 with a cross-sectional area "A_:." The cross sectional area "A₂" of the minimum flow resistance opening 32 is determined by the target minimum flow rate, as will be discussed below.

One end ofthe maximum flow restrictor 18 has an opening 34 with a one degree internal taper sized for frictional engagement with the opposite end 30 of the minimum flow restrictor 20. A negative pressure sensing port 36 is located at base 35 of the opening 34. A tube 38 is pressure-fit into the port 36. The tube 38 is preferably perpendicular to the direction ofthe air stream "S" and protrudes slightly above the inside surface of the opening 34. The protrusion ofthe tube 38 serves as a stop for the connection with the minimum flow restrictor 20. (see Figure 1). The opposite end ofthe maximum flow restrictor 18 has an opposite end 40 with a similar diameter and taper as the opening 22 so that it may optionally be attached to another flow restrictor (see Figure 2B). The maximum flow restrictor 18 has an air flow resistance opening 42 having a cross-sectional area "A,." The cross sectional area "A," ofthe maximum flow resistance opening is determined by the target maximum flow rate, as will be discussed below.

The openings 22 and 34 preferably have the same or similar inside diameter and the opposite ends 30 and 40 preferably have the same outside diameter so that the flow restrictors are interchangeable. In the preferred embodiment, only the air flow resistance openings 32, 42 ofthe flow restrictors

20, 18, respectively, vary in size. It will be understood that a variety of nesting structure would be suitable for use with the interchangeable air flow restrictors of the present invention.

Turning now to Figure 1 , a chamber 23 is formed within the flow restrictor 20 adjacent to the tube 28 and a chamber 25 is also formed in the flow restrictor 18 adj cent to the tube 38 of the flow restrictor 20, so that a venturi effect is created in the chambers 23 and 25. During the inhalation cycle, the air flow rate increases until it is first restricted by the air flow resistance opening 32. When the air flow rate in the air flow resistance opening 32 reaches a particular level, a negative pressure condition is created in the chamber 23 that is transmitted to sensor 50 through tubes 28 and 54. As the flow rate increases, the air flow is next restricted by the air flow resistance opening 42 in the flow restrictor 18. As the restricted air flow moves from the air flow resistance opening 42 to the chamber 25, a second negative pressure condition is created in the chamber 25 that is transmitted to sensor 52 by tubes 38 and 56. The negative pressure sensors 50, 52 each include a relay wired in the open position. A negative pressure sensor/relay combination suitable for use in the present invention is available from World Magnetics of Traverse City, MI under model number PSF100A/5RF100B. It will be understood that in embodiments utilizing more than two air flow restrictors, additional sensors may be added to the air flow system 10 without departing from the scope ofthe present invention. In the embodiment illustrated in Figure 1 , the tubes 54, 56 are constructed from a 2.54 mm ID (0.1 inch) clear, vinyl tubing. Consequently, the present air flow system 10 requires no electrical connections to be attached to the patient's hand-held aerosol generator 12. The sensors 50, 52 are calibrated to trigger when subjected to a negative pressure condition equivalent to 124.5 Pa

(0.5 inches of water). It will be understood that a variety of negative pressure sensors may be used without departing from the scope of the present invention.

If the flow rate is below the target minimum flow rate, a two-way solenoid 60 directs compress air from a compress air source 62 to the atmosphere through a vent 64. When a patient's inhalation flow rate is above a target minimum flow rate, the pressure in the chamber 23 drops. When the sensor 50 is exposed to a negative pressure condition of less than 0.5 inches of water, the relay in the negative pressure sensor 50 closes. The closed relay 50 provides electric power from power supply 58 to activate the two-way solenoid 60. The compressed air is redirected by the activated solenoid 60 through hose 66 to the aerosol generator 12. The relay in the negative pressure sensor 50 will remain closed until the inhalation flow rated drops below the minimum target flow rate. A 1 OOV AC two-way solenoid suitable for use in the present invention is available from Grainger of Arden Hills, MN under model number 74514-01 15. When a patient's inhalation flow rate is above a target maximum flow rate, the pressure in the chamber 25 drops. That pressure drop is communicated to the sensor 52 through the tube 56, causing the relay in the sensor 52 to close and an alarm 70 to be activated. The alarm generally signals the patient to slow the inhalation flow rate. When the air flow rate drops below the selected maximum target flow rate, the signal from the alarm 70 stops. In one embodiment, the alarm 70 is a 9 volt battery wired to a Piezo signal generator.

In an embodiment in which the aerosol generator 12 is an ultrasonic nebulizer, a relay wired to provide electrical power to the ultrasonic nebulizer is substituted for the solenoid 60. In an alternate embodiment, the sensor 52 may be wired to deactivate the solenoid 60 when the target maximum flow rate is achieved. In an alternate embodiment, a high frequency chest compressor 90 may be used in combination with the aerosol generator 12. When a patient's inhalation flow rate is above a target minimum flow rate, the aerosol generator 12 is activated and remains so until the inhalation air flow drops below the minimum target flow rate. The high frequency chest compressor 90 is deactivated by double-throw relay 92 during the inhalation cycle. At the end of the inhalation cycle, when the inhalation flow rate drops below the target minimum flow rate, the double-throw relay 92 activates the high frequency chest compressor 90. A high frequency chest compression device suitable for use in the present invention is disclosed in U.S. Patent Nos. 5,056,505 and 4,838,263, which are hereby incoφorated by reference.

Figure 2B is an exemplary configuration of three interchangeable flow restrictors 70, 72, 74 with air flow resistance openings 76, 78, 80 having progressively smaller cross sectional areas A₃ A₄ A, along airstream "S". The flow restrictors 70, 72, 74 may be "nested" in the order shown in Figure 2B, or alternatively, rearranged in a different order. Openings 79, 81, 82 are sized so that any ofthe flow restrictors 70, 72, 74 may be attached to the T-tubes (see Figures 1 and 3). The tubes 84, 86, 88 are fluidly attached to negative pressure sensors (not shown), as previously discussed herein. The flow restrictors 70, 72, 74 detect maximum, intermediate and minimum target flow rates, respectively. It will be understood that the number of flow restrictors may increase or decrease depending upon the particular application.

Figure 3 is a schematic illustration of an exemplary air flow system 10' for controlling a medical treatment 90 as a function of minimum and maximum target respiratory flow rates. Breathing tube 14' has an opening 16' for a mouth piece at one end and a minimum flow restrictor 20' at the other end. The minimum flow restrictor 20' has an air flow resistance opening 32'. A maximum flow restrictor 18' is fluidly connected to the opposite or distal end ofthe minimum flow restrictor 20'. The maximum flow restrictor 18' has an air flow resistance opening 42'.

The flow restrictors 18', 20' are fluidly coupled to a pair of negative pressure sensors 50', 52' via a pair of tubes 54', 56', as discussed in Figure 1.

When a patient's inhalation flow rate is above a target minimum flow rate, a relay in the sensor 50' close and triggers relay 92' to activate or deactivate medical treatment 90'. Correspondingly, when a patient's inhalation flow rate is above a target maximum flow rate, the relay in sensor 52' closes and activates alarm 70'. The alarm generally signals the patient to slow the inhalation flow rate. When the air flow rate drops below the selected maximum target flow rate, the signal from the alarm 70' stops. In an alternate embodiment, activation ofthe sensor 52' may trigger the relay 92' to activate or deactivate the medical treatment 90'. The flow restrictors 18, 18' and 20, 20' as shown in Figures 1 and 3, are preferably constructed using delrin, although a variety of plastic materials approved for medical use would be suitable for this puφose. For example, disposable, low-cost plastic would be suitable and would eliminate the need for sterilization. The medical treatments include, for example, a high frequency chest compression device, such as discussed above, a ventilator or a respirator. It will be understood that the present air flow system may be used with any medical treatment that is dependent on the respiratory cycle ofthe patient.

Figure 4 depicts an oscillatory curve 100 representing an exemplary respiratory cycle of a patient. At the start 102 of an inhalation cycle 101, the air flow rate is zero. As the patient begins to inhale, the flow rate initially increases and then tapers off to zero 104 at the end of the inhalation cycle. The present air flow system 10. 10' detects target air flow rate 106, 108, 1 10, 1 12 generated during inhalation. When the inhalation air flow exceeds the minimum target flow rate 106, the sensor 50, 50' are activated, and remain activated until the flow rate drops below the minimum target flow rate at 112. Likewise, when the air flow rate exceeds the maximum target flow rate 108, the sensors 52, 52' are activated, and remain activated until the flow rate drops below the maximum target flow rate at 1 10. None of the negative pressure sensors are activated during the exhalation cycle 101. As mentioned above, the cross sectional areas "A" ofthe flow resistance openings 32, 32' and 42, 42' are determined by the target flow rate. When a patient's primary health care giver selects the target flow rate(s) at which activation and/or deactivation of the medical treatment should occur, flow restrictors with the appropriate cross-sectional openings are attached to the patient mouthpiece 14, 14'.

Figure 5 depicts a range of target flow rate values as a function of the cross-sectional area of flow resistance openings for the negative pressure sensors calibrated to trigger when exposed to a negative pressure condition equivalent to 124.5 Pa (0.5 inches of water). Air flow was measured using a calibrated manostat flow meter calibrated at 20 °C +/- 2% for eleven different flow restrictors. Cross-sectional areas that were tested ranged from 7.94 mm² to 103.9 mm² (0.0123 in² to 0.161 1 in² ) corresponding flow rate values ranged from 4 to 60 1/min. The correlation coefficient comparing the cross sectional area of each flow restrictor and flow rate value was 0.99. The regression equation is y = 0.13 + 365.88x where y = flow in L/min and x = cross sectional area of the resistance opening in square inches. The following examples merely illustrate the present invention and is not intended to limit the scope ofthe invention in any way.

EXAMPLE 1 The minimum flow restrictors are made 1.1 in (27.94 mm) in length with an outer diameter of 1.25 in (31.75 mm). The end ofthe minimum flow restrictor that frictionally engages with the end ofthe breathing tube or T-tube is 0.5 in (12.7 mm) long and has a one degree tapered opening that is 0.5 in (12.7 mm) deep, with a 0.85 in (21.59 mm) minimum inner diameter at the base. The negative pressure sensing port is located 0.1 in (2.54 mm) from the base. A one inch long stainless steel hypo tube is pressure-fit into the port peφendicular to the air stream "S". The opposite end of the minimum flow restrictor that frictionally engages with the maximum flow restrictor is 0.5 in (12.7 mm) long and has a one degree external tapered opening, with a maximum diameter of 0.86 in (21.84 mm).

The maximum flow restrictor is also 1.1 in (27.94 mm) in length with an outer diameter of 1.25 in (31.75 mm). The end ofthe maximum flow restrictor that frictionally engages with the opposite end of the minimum flow restrictor is 0.5 in (12.7 mm) long and has a one degree internal tapered opening with a maximum outside diameter of 0.85 in (21.59 mm). The negative pressure sensing port is located 0.1 in (2.54 mm) from the base. A one inch long stainless steel hypo tube is pressure-fit into the port. The tube is preferably peφendicular to the direction ofthe air stream and protrudes slightly above the inside surface ofthe opening to serve as a stop for the connection with the minimum flow restrictor (see Figure 1 ). The opposite end ofthe maximum flow restrictor has an opening with a similar diameter and taper as the opening on the minimum flow restrictor that engages with the breathing tube.

EXAMPLE 2 To determine the efficiency ofthe air flow system 10 at reducing the loss of nebulized aerosol, water vapor samples were captured on fresh CaSO₄ (Dryrite) granules and weighed. Samples were collected from three groups: (1) normal breathing with no aerosol, (2) normal breathing with aerosol without the air flow system 10, and (3) normal breathing with aerosol using the air flow system 10. A 1 in (25.4 mm) inner diameter tube was loaded with 20 grams of CaSO₄ granules. The tube and CaS0₄ were then weighed. A one way valve allowed only exhaled air to pass through the CaS0₄ granules. An individual with average tidal volume of 2.25 liters was instructed to breathe five tidal breaths, with nose plugged, through the water vapor trap. The tube was not removed from the individual's mouth until 5 breaths were complete. The weight of water vapor collected after each five breath cycles was measured. This five breath cycle was repeated five times each for normal breathing with no aerosol, normal breathing with aerosol without the air flow system 10, and normal breathing with aerosol using the air flow system 10.

The efficiency of aerosol delivery to the patient was determined by comparing the amount of aerosol produced over the test period to the amount lost to the environment. The output of the nebulizer water vapor was determined by measuring and weighing the amount of water vapor trapped after operating for ten seconds through the water vapor trap described above. The procedure was repeated six times. The efficiency of controlled and uncontrolled aerosol delivery to patients was estimated by calculating the amount of aerosol produced by each method based on the time the nebulizer was active during the five breath cycle. That amount was then compared to the amount of loss to the environment over that period. The mean weight of water vapor added from five breaths of normal breathing with no aerosol was 0.13 +/- 0.02 gm; for normal breathing with uncontrolled aerosol (i.e., without the air flow system) 0.32 +/- 0.06 gm; and for normal breathing with a controlled aerosol using the present air flow system 0.17 +/- 0.04 gm. The difference between the amount of water vapor added by normal breathing and normal breathing with a controlled aerosol generator was

0.04 gm. The difference between the amount of water vapor added by normal breathing and breathing uncontrolled aerosol was 0.19 gm. Nebulizer water vapor output was measured by using the water vapor trap described above. Six 10 second collections were made with a mean of 0.06 gm +/- 0.002 gm.

The efficiency of deliver}' to the patient was determined by comparing the amount of aerosol produced over the test time period to the amount lost to the environment. The controller delivered aerosol for a mean time of 3.8 seconds +/- 0.18 per breath. Delivery efficiency was calculated by multiplying total time aerosol was delivered during the five breath cycle by 0.006 gm (output per second of the aerosol generator) to determine the total aerosol produced. This amount minus the amount lost to the environment is the amount delivered to the patient. The controlled aerosol generator delivered aerosol to the patient at an efficiency rate of 64%.

The uncontrolled aerosol five breath cycle had a mean total activation time of 43 seconds +/- 2.5 seconds. The mean total activation time multiplied by 0.006 gm (output per second of the aerosol generator) represents the number of grams of aerosol produced. That number minus the quantity lost to the environment is the amount of aerosol delivered to the patient. The uncontrolled aerosol generator delivered aerosol to the patient at an efficiency rate of 27%.

Patients routinely remove the aerosol generator T-tube from their mouths for 20-30% of the time during therapy. During these periods, even more of the aerosol is lost to the environment. An analysis of uncontrolled aerosol delivery factoring in an 1 1 second rest cycle (25% of the total therapy time) yields an aerosol delivery efficiency rate of 22%. The above calculations are summarized in table 1 below.

A. B. C. D. weight of aerosol weight of aerosol weight of aerosol delivery produced lost to the delivered to efficiency (C/A) environment patient (A-B)

Aerosol 0.1 1 gm 0.04 gm 0.07 gm 64% controlled

(5 breath cycle)

Uncontrolled 0.26 gm 0.19 gm 0.07 gm 27% aerosol (5 breath cycle)

Uncontrolled 0.32 gm 0.25 gm 0.07 gm 22% aerosol (5 breath cycle plus 1 1 second rest cycle)

The present invention has now been described with reference to the several embodiments thereof. It will be apparent to those skilled in the art that many changes can be made in the embodiments described without departing from the scope ofthe invention. Thus, the scope ofthe present invention should not be limited to the structures described herein, but only by structures described by the language ofthe claims and the equivalents of those structures. For example, any type of medical treatment that is advantageously synchronized with the respiratory cycle may be used with the present air flow system.

Claims

What Is Claimed Is:

1. An air flow system for detecting at least one target flow rate in a respiratory cycle of a subject, comprising: a mouthpiece; and at least a first interchangeable flow restrictor having a first chamber fluidly coupled to a first air flow resistance opening and the mouthpiece, the first air flow resistance opening having a first cross sectional area less than the cross sectional area of the first chamber, and a first sampling port proximate the first chamber in fluid communication with a first negative pressure sensor so that a first target flow rate generates a negative pressure condition in the first chamber that activates the first negative pressure sensor.

2. The system of claim 1 further including a second interchangeable flow restrictor having a second chamber fluidly coupled to a second air flow resistance opening and the mouthpiece, the second air flow resistance opening having a second cross sectional area less than the cross sectional area ofthe second chamber, and a second sampling port proximate the second chamber in fluid communication with a second negative pressure sensor so that a second target flow rate generates a negative pressure condition in the second chamber that activates the second negative pressure sensor.

3. The system of claim 1 wherein the at least first interchangeable flow restrictor comprises a plurality of interchangeable flow restrictors fluidly coupled to the mouthpiece.

4. The system of claim 1 further including tubes fluidly connecting the first sampling port to the first negative pressure sensor located remotely from the flow restrictor.

5. The system of claim 2 further including first and second tubes fluidly connecting the first and second sampling port to the first and second negative pressure sensor, respectively.

6. The system of claim 1 wherein the first target flow rate is a function ofthe first cross sectional area.

7. The system of claim 1 wherein the first interchangeable flow restrictor is releasably connected to the mouthpiece.

8. The system of claim 2 wherein the second interchangeable flow restrictor is fluidly coupled at a distal end of the first interchangeable flow restrictor.

9. The system of claim 2 wherein the first cross sectional area is less than the second cross sectional area.

10. The system of claim 2 wherein the first target flow rate comprises a minimum target flow rate and the second target flow rate comprises a maximum target flow rate.

1 1. The system of claim 1 further including an aerosol generator fluidly coupled to the mouthpiece that is activated in response to activation ofthe first negative pressure sensor.

12. The system of claim 2 wherein the first and second cross sectional areas are in the range of about 6 to 130 mm².

13. A system for controlling a medical treatment according to the air flow system of claim 1 comprising means for activating a medical device in response to activation of he first negative pressure sensor.

14. The system of claim 13 wherein the device comprises an aerosol generator fluidly connected to the mouthpiece that is activated in response to activation of the first negative pressure sensor.

15. A system for controlling a medical treatment according to the air flow system of claim 1 comprising means for activating a high frequency chest compressor in response to deactivation ofthe first negative pressure sensor at the end ofthe inspiration cycle.

16. A system for controlling a medical treatment according to the air flow system of claim 2 comprising means for activating a device in response to activation of the first negative pressure sensor and means for deactivating the device in response to activation ofthe second negative pressure sensor.

17. A system for controlling a medical treatment according to the air flow system of claim 2 comprising means for activating a device in response to activation of the first negative pressure sensor and means for signaling an alarm in response to activation of the second negative pressure sensor.