Abstract:
An electronic gas flow triggering circuit for use in an aerosol drug dispensing device has a voltage source and a self-nulling circuit to which voltage is supplied from the voltage source. A hot wire anemometer filament forms a component of the bridge circuit which is adapted to maintain a constant resistance of the anemometer filament, the bridge drive voltage being dependent upon the gas flow across the anemometer filament. A comparator compares the bridge output voltage with a reference voltage and provides a triggering signal to operate the device if the bridge drive voltage is greater than the reference voltage.

Description:
BACKGROUND OF THE INVENTION 
     This invention relates to a triggering circuit for use in an aerosol drug dispensing device, and in particular, to a triggering circuit for detecting the operation of the device by a user. 
     Prior trigger circuits do not readily discriminate between a real triggering event, i.e. a patient inhaling on the device, and a false event, i.e. the movement of the device through the air or changes in ambient conditions. If the circuit were to be triggered by a false event, the device would be operated when the user was not ready to receive the drug and therefore that dose of the drug would be wasted. Thus, the present invention aims to provide a control circuit which overcomes the above problem. 
     WO92/07599 discloses a portable inhalation device for the administration of medicament in the form of aerosolised fine particles or droplets of liquid to the respiratory system of a patient. The inhalation device comprises a bridge circuit which creates an electrical output signal dependent upon the resistance in a hot wire anemometer. The resistance of this element varies as the flow across it changes and accordingly, the bridge circuit is not self nulling. 
     SUMMARY OF THE INVENTION 
     According to the present invention, there is provided an electronic gas flow triggering circuit for use in an aerosol drug dispensing device, the circuit including:
         a voltage source;   a self-nulling bridge circuit to which voltage is supplied from the voltage source;   a hot wire anemometer filament forming a component of the bridge circuit which is adapted to maintain a constant resistance of the anemometer filament, the bridge drive voltage being dependent upon the gas flow across the anemometer filament; and   a comparator for comparing the bridge drive voltage with a reference voltage such that a trigger signal is provided if a certain rate of change of gas flow is exceeded.       

     As real and false events can be distinguished based on the time frame within which they occur, the present invention compares the bridge drive voltage with a reference voltage composed of the low frequency signals contained within the bridge drive voltage. When the bridge drive voltage reaches a predetermined level relative to the reference voltage, the gas flow circuit is triggered. The present invention therefore is sensitive to rate of change of gas flow and is only triggered by the higher frequency signals that correspond to a patient inhaling on the device. 
     Accordingly, the present invention provides a self nulling bridge circuit in which the resistance of the hot wire anemometer is maintained at a constant value. By maintaining the resistance of the hot wire anemometer at a constant value, the sensor element can be remotely located away from the instrumentation electronics. The resistance of the sensor element is significantly larger than the resistances of the connecting wires and any electrical connectors employed. A constant sensor element resistance also eliminates errors resulting from potential divider ratio variations that would occur in a variable resistance sensor element system. When compared to the prior art, the present invention provides a simpler trigger mechanism in which there is no significant temperature change and only a single measurement is required. 
     Preferably, the bridge circuit is a Wheatstone bridge. 
     The reference voltage may be derived from the bridge drive voltage, thus preventing erroneous triggering due to changes in ambient temperature. 
     The reference voltage may be derived from the bridge drive voltage and may be filtered to pass only low frequency signals. Thus, the reference voltage does not change appreciably with rapid changes in air flow whereas the unfiltered bridge drive voltage signal (the measurement drive) responds quickly to rapid changes in air flow. The sensitivity of the detection circuit may be set by dividing down the measurement voltage such that the more the voltage is divided, the less sensitive the trigger becomes, and the faster a patient needs to inhale on the device in order to trigger the circuit. 
     Thus, the comparison of filtered and unfiltered voltage signals provide a simple and effective means of sensing the rate of change of air flow rather than just the change in the air flow so that the control system discriminates between real and false triggering events. 
     Air passing over the anemometer filament causes it to cool which, in turn, effects the power delivered to the filament because, as the bridge circuit is operating in a self-nulling mode, it attempts to maintain the filament temperature by means of a closed loop control system. Monitoring the voltage supply to the bridge provides a means for sensing changes in the air flow, i.e. the anemometer and the bridge form an air flow sensor. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       An embodiment of the present invention will now be described with reference to the accompanying drawings, in which: 
         FIG. 1  shows a cross sectional side view of a device for dispensing a pharmacologically active liquid; 
         FIGS. 2 to 6  show successive stages during operation of the device of  FIG. 1 ; 
         FIG. 7  shows a trigger circuit for controlling the dispensing of the pharmacologically active liquid; and 
         FIGS. 8   a  and  8   b  illustrate, in section, preferred forms of the perforate membrane for the device of  FIG. 1 . 
     
    
    
     DESCRIPTION OF THE INVENTION 
     As can be seen from  FIGS. 1 to 6 , a drug dispensing device  25  forming part of a drug delivery system has a hollow casing  1  in which a reservoir body  2  is slidably mounted. A spray head mounting assembly  3  is attached to one end of the casing  1 , the mounting assembly  3  comprising spray head assembly  3   a , clamping means  4 , diaphragm clamp  11   a  and seal  17 . A main return spring  5  is provided in a recess  29  in the casing  1  between a lower face  2   a  of the reservoir body  2  and a notch  29   a  in diaphragm ring  30  to support the reservoir body  2 , biassed towards its topmost position (as shown), within the casing  1 . The reservoir body  2  is movable between a first and second position within the casing  1 . 
     The reservoir body  2  has a reservoir cavity  6  in which a liquid drug  7  is stored and which can be topped-up via a re-fill port  23 . A plunger  8  is provided within the reservoir body  2  and extends out of the upper (as shown in the figures) end of the reservoir body  2 , such that it engages with a cap  9  which covers the top of the device  25 . A plunger return spring  10  is provided which extends from the reservoir body  2  and biases the plunger  8  out of the reservoir cavity  6 . 
     A metering cavity  12  is provided at the lower end of the reservoir cavity  6  and the metering cavity  12  is in liquid communication, via a one-way ball valve  13 , with a dispensing conduit or tube  14  provided within the lower end of the reservoir body  2 . An elastomeric diaphragm  15  is provided around the lower portion of the reservoir body  2  and is joined at its outer edge to the casing  1  by diaphragm clamp  11   a  and a seal  11 , which is an O-ring, to define, together with the reservoir body  2  and the casing  1 , a chamber V 1 . 
     The spray head mounting assembly  3  is provided with an outlet  26  covered by a membrane  16  which, together with the lower portion of the reservoir body  2 , the spray head mounting assembly  3  and the elastomeric diaphragm  15 , defines a chamber V 2 . The membrane  16  is retained by a seal  17  and is provided with a perforate portion  27  overwhich a holding reservoir  18  is provided and which can receive liquid drug  7  through the dispensing conduit or tube  14  in use. The perforate portion  27  of the membrane  16  has perforations  50  (see  FIGS. 8   a  and  8   b ) of a reverse taper and this is described in greater detail with reference to  FIGS. 8   a  and  8   b.    
     The spray head mounting assembly  3  is provided with a damping inlet valve  19  leading into the chamber V 1 , a first one-way exhaust valve  20  from the chamber V 1  and a second one-way exhaust valve  21  from the chamber V 2 . 
       FIG. 1  shows the device  25  at rest prior to metering a predefined dose of drug  7  into the holding reservoir  18 . In this state, both the main return spring  5  and the plunger return spring  10  are extended so that the reservoir body  2  and the plunger  8  have been pushed fully upwards against physical stops (not shown). 
     The elastomeric diaphragm  15  is stretched from its normally flat state and acts as a flexible wall, separating the two chambers V 1  and V 2 . These chambers are initially at ambient pressure and the perforations  50  in the perforate portion  27  of membrane  16  are dry and open. 
     The operation of the device shown in  FIG. 1  will now be described with reference to  FIGS. 2 to 6 . 
     To initiate operation of the device  25 , the cap  9  is pressed downwards  9   a  (see  FIG. 2 ). The reservoir body  2  and the plunger  8  are moved downwards at approximately the same rate, compressing the main return spring  5  until the reservoir body  2  reaches physical stops  28 . At this stage the force exerted on the plunger return spring  10  is less than its preload. The lower end of the dispensing conduit or tube  14  is now positioned close to the inner surface of the membrane  16 . 
     The reservoir body  2  forms an air-tight sliding seal with the casing  1 . Therefore, as the reservoir body  2  descends, air is displaced from the chamber V 1 , through the first one-way exhaust valve  20  and the elastomeric diaphragm  15  returns towards its relaxed, almost flat state. At the same time, air is displaced from the chamber V 2  through the second one-way exhaust valve  21 , through the damping valve  19 , and through the perforate portion  27  of the membrane  16 . 
     As can be seen in  FIG. 3 , once the reservoir body  2  has reached its physical stops  28 , further depression of the cap  9  compresses the plunger return spring  10 . This pushes the plunger  8  into the metering cavity  12  (see  FIG. 1 ). The plunger  8  is provided with a plunger seal  24  which engages with the side wall  12   a  of the metering cavity  12  to form a seal therewith. This traps a metered dose of the liquid drug  7  in the metering cavity  12 . The sealing pressure of the plunger seal  24  against the side-wall  12   a  of the metering cavity  12  is greater than the sealing pressure of the elastomeric ball valve  13  against the ball valve seat  13   a . Therefore, as the plunger  8  is depressed further, the trapped dose of liquid drug  7  is displaced from the metering cavity  12  past  13   b  the one-way elastomeric ball valve  13 , through the dispensing conduit or tube  14  and into the holding reservoir  18 . The holding reservoir  18  has a side-wall  22  which is either coated with or formed of PTFE. In this configuration, all of the perforations  50  in the perforate portion  27  of membrane  16  are now covered with liquid drug  7   a . 
     Once the cap  9  is released, as shown in  FIG. 4 , the plunger return spring  10  starts to expand, pushing the plunger  8  out of the metering cavity  12 , thus creating a partial vacuum in the metering cavity  12 . The elastomeric ball valve  13  remains closed. The plunger return spring  10  is strong enough to overcome this partial vacuum so as to push the plunger  8  clear of the metering cavity  12  until it reaches a physical stop (not shown). This partial vacuum pressure reduces as the plunger  8  is withdrawn from the metering cavity  12  until the pressure drop across the plunger seal  24  exceeds the sealing pressure against the side-wall  12   a  of the metering cavity  12 . A small amount of liquid drug  7  can then flow past the plunger seal  24  into the metering cavity  12  as the plunger  8  is withdrawn. The main return spring  5  is able to extend, pushing  2   a  the reservoir body  2  upwards as shown in  FIG. 5 . This then causes the chambers V 1  and V 2  to expand and the elastomeric diaphragm  15  to stretch. 
     Air cannot enter through the one-way valves  20 ,  21  or through the perforate portion  27  of membrane  16  as the perforations  50  are now covered by liquid drug  7   a . The perforations  50  resist air ingress because of the surface tension of the liquid drug forming menisci across the perforations  50  in the perforate portion  27  and therefore, as the chambers V 1  and V 2  expand, the pressures within those chambers decrease below ambient pressure. The reservoir body  2  continues to move upwards  2   a  until the restoring force of the main return spring  5  is stalled by the opposing forces created, in the main, by the negative pressures in the chambers V 1  and V 2  and by the tension in the elastomeric diaphragm  15 . 
     When the damping valve  19  is fully closed, no air can enter the chamber V 1  and so the main return spring  5  will not extend far before it is stalled by the negative pressures generated in the chambers V 1 , V 2  and by tension in the elastomeric diaphragm  15 . 
     However, when the damping valve  19  is slightly open, it will allow air to enter into, and conversely exhaust from, the chamber V 1  at a rate determined by the restoring force of the main return spring  5 . Therefore, as the main return spring  5  pushes the reservoir body  2  upwards  2   a , the volume of chamber V 1  expands at a controlled rate until the reservoir body  2  reaches a predefined physical stop (not shown) (see  FIG. 6 ). 
     The main return spring  5  also causes the elastomeric diaphragm  15  to stretch at a controlled rate which, in turn, causes the volume of the chamber V 2  to expand at a controlled rate. The pressure within the chamber V 2  behind the spray head assembly  3   a  therefore reduces at a controlled rate. 
     The physical stop (not shown) for the upward movement of the reservoir body  2  limits the extension of the stretched elastomeric diaphragm  15  and defines the final volume change of the chamber V 2  and thus the target negative pressure achieved in V 2 . At this stage, the pressure in the chamber V 1  is at ambient pressure. 
     Therefore, the net result is that the target (operating) negative pressure, which is close to the bubble point pressure for the spray head assembly  3   a  (the bubble point pressure being the pressure differential at which air bubbles enter the chamber V 2  through the perforations  50  in the perforate portion  27 , by overcoming the surface tension of the menisci of liquid drug), can be reached at a controlled rate without air being ingested through the membrane  16 . 
     If the negative pressure in the chamber V 2  is generated too quickly, i.e. by opening the damping valve  19  too far, then the surface tension of the liquid drug covering the perforations in the perforate portion  27  of the membrane  16  will be overcome. Air will then be ingested rapidly through the perforate portion  27 , hence cancelling the negative pressure in the chamber V 2 . In the absence of a negative pressure, the drug dispensing device  25  will not generate a spray. 
     It should be noted that the rate of air admission and exhaustion through damping, valve  19  is much less than the rate of exhaustion through the first one way exhaust valve  20 . 
     An electronic control circuit  30  as shown in  FIG. 7  triggers the delivery of the liquid drug through the perforate portion  27  of the membrane  16  by controlling membrane actuating means (not shown) such as piezoelectrically activated vibrating means. The control system uses a small hot wire anemometer  39  to monitor air flow into the drug delivery system. The anemometer  39  senses when a patient is inhaling and trying to trigger the spray head delivery. 
     The maximum spring load of the plunger return spring  10  preferably should not exceed 20N. This therefore limits the force that a user must exert in order to operate the device. 
     The plunger return spring  10  is preloaded and this preload should be greater than the fully compressed load in the main return spring  5 . This ensures that the reservoir body  2  is fully down against its physical stop  28  before the dose is metered into the holding reservoir  18 . If the dose is metered from the dispensing aperture  14  when the reservoir body  2  is not fully down, all or part of the dose may miss the holding reservoir  18 . The target negative pressure in the chamber V 2  will then not be reached when the cap  9  is released because the perforations in the perforate portion  27  of the membrane  16  will remain dry and open. 
     As the plunger  8  is withdrawn from the reservoir cavity  12  a partial vacuum is created between the closed one-way elastomeric ball-valve  13  and the plunger seal  24 . The spring force range of the plunger return spring  10  must be sufficient to lift the plunger  8  out of the reservoir cavity  12  when the cap  9  when the cap is released, in order to create a pressure difference across the plunger seal  24  (between the liquid drug  7  in reservoir cavity  6  and the elastomeric ball valve  13 ) which exceeds the sealing pressure of the plunger seal  24  against the side-wall  12   a  of the reservoir cavity  12 . Liquid drug  7  is then above to flow past the plunger seal  24  and into the reservoir cavity  12  as the plunger  8  is withdrawn. This has the advantage of minimising foam generation within the reservoir cavity  12  especially with liquid drugs containing surfactants as part of their formulation. 
     The spring force range of the main return spring  5  must be sufficient to lift the reservoir body  2  against the negative pressures generated within the chambers V 1 , V 2 , the stretching of the elastomeric diaphragm  15  and the friction of the sliding seal between the reservoir body  2  and the casing  1 . 
     The maximum tension in the elastomeric diaphragm  15 , which is achieved when the reservoir body  2  is at rest, must be high enough to maintain the target negative pressure within the chamber V 2  without the elastomeric diaphragm  15  deforming under that negative pressure. This tensional force must not be high enough to prevent the main return spring  5  from lifting the reservoir body  2  up against the predetermined physical stop during the generation of the negative pressures. 
     In summary, the extension force of the plunger return spring  10  is greater than the extension force of the main return spring  5  which is greater than the tensional force of the elastomeric diaphragm  15 . 
     From  FIG. 7 , it can be seen that the control circuit employs two operational amplifiers (op-amps)  31 ,  32 , which can supply output voltages over the full power supply range, i.e. between the supply voltage Vcc and ground GND. Op-amp  31  is in closed loop configuration with a Wheatstone bridge to maintain a constant resistance of the anemometer filament  39 . The drive voltage  40  for the Wheatstone bridge is supplied to the op-amp  32  which is configured as a voltage comparator. The comparator is a switching device having a measurement input and a reference input. When the measurement voltage  33  is higher than the reference voltage  34 , the output  35  from the op-amp  32  switches from GND to Vcc. 
     Circuits of this type are well known in anemometry. However the circuit of  FIG. 7  shows a simple implementation of a circuit which can discriminate between high and low rates of change of air flow. The RC time constant used for the reference voltage  34  is set so that it can respond, in a damped manner, to slowly varying conditions such as changes in ambient conditions and to the normal slow breathing of a patient. Therefore, the reference voltage  34  is insensitive to fast rates of change in air flow during the patient inhalation cycle. The measurement voltage  33  at the comparator op-amp  32  is an attenuated signal from the Wheatstone bridge drive voltage  40 , but does not have a long time constant filter and so reacts directly with the patient&#39;s inhalation. 
     If the patient inhales at a sufficiently high volume rate, then the measurement voltage  33  will exceed the reference voltage  34  due to the difference in the time constant of the reference and measurement voltages. In order for this to happen, a certain threshold of rate of change of air flow must be exceeded. When the measurement voltage  33  exceeds the reference voltage input  34 , there will be a change in the voltage state at the output  35 . Since the comparator circuit does not have any feedback, it is operating in a high gain mode and so has only two output states, 0 v and Vcc. Thus, it is possible to couple this circuit directly to TTL logic devices. 
     Two variable resistors  36 ,  37  are provided within the control circuit  30 . The variable resistor  36  in the comparator is used to adjust the level of attenuation applied to the measurement voltage  33 , which sets the sensitivity of the circuit to the rate of change of patient&#39;s inhalation. The second variable resistor  37 , in the Wheatstone bridge, is provided to set the temperature of the anemometer filament  39  and is used to set the minimum rate of change of air flow that the circuit can detect. Thus, these adjustments allow a high degree of flexibility in the installation of the anemometer as a breath trigger device. 
     A field effect transistor  38  is used to alter the reference ratio of the Wheatstone bridge thus enabling or disabling the air flow sensor via “ENABLE”. In this way, the power consumption of the control circuit  30  is reduced, allowing the device to have a useful operating life when powered by batteries. 
       FIG. 8   a  shows cross-sectional detail of a first example perforate membrane  16 , which is operable to vibrate substantially in the direction of arrow  58  and which is suitable for use with the drug dispensing device  25  to produce fine aerosol sprays. In one embodiment, the membrane  16  comprises a circular layer of polymer which contains a plurality of tapered conical perforations  50 . Each perforation  50  has openings  53  in the front exit face and openings  54  in the rear entry face, which perforations are laid out in a square lattice. Such perforations may be introduced into polymer membranes by, for example, laser-drilling with an excimer laser. 
       FIG. 8   b  shows cross-sectional detail of a second example perforate membrane  2016 , which membrane is operable to vibrate substantially and suitable for use with drug dispensing device  25  in the direction of arrow  58 . The membrane is formed as a circular disc of diameter 8 mm from electroformed nickel, and is manufactured, for example, by Stork Veco of Eerbeek, The Netherlands. Its thickness is 70 microns and is formed with a plurality of perforations shown at  2050  which, at ‘front’ face  2051 , are of diameter shown at “a” of 120 microns and at ‘rear’ face  2052  are of diameter shown at “b” of 30 microns. The perforations are laid out in an equilateral triangular lattice of pitch 170 μm. The profile of the perforations varies smoothly between the front and rear face diameters through the membrane thickness with substantially flat ‘land’ regions (shown at “c”) of smallest dimension 50 μm in front face  2051 .