Inhalation therapy device

Abstract

This disclosure relates, in particular, to an inhalation therapy device comprising a sensor (47) configured to detect flow parameters of the flow in a channel (7), and a controller (18) configured to determine at least the user's inhalation phase and / or exhalation phase based on the output of the sensor (47), and to repeatedly operate an actuator (6) during the inhalation phase and not operate the actuator (6) during the exhalation phase during a single treatment time, wherein the aerosol output rate of the inhalation therapy device at the first opening (8) during a single treatment time is 0.1% (M / V) of albutyric acid in 0.9% sodium chloride as a liquid. When using the roll test solution, the flow rate is at least 0.48 g / min when the median mass diameter of the aerosol droplet is between 4.5 μm and 5.0 μm, or at least 0.42 g / min when the median mass diameter of the aerosol droplet is between 4.0 μm and 4.5 μm, or at least 0.29 g / min when the median mass diameter of the aerosol droplet is between 3.5 μm and 4.0 μm, or at least 0.20 g / min when the median mass diameter of the aerosol droplet is between 3.0 μm and 3.5 μm, or at least 0.13 g / min when the median mass diameter of the aerosol droplet is between 2.5 μm and 3.0 μm.

Description

[Technical Field]

[0001] This disclosure relates to an inhalation therapy device that generates an aerosol and delivers it to a user. [Background technology]

[0002] Inhalation therapy devices, also known as (inhalation) nebulizers or aerosol delivery devices, are widely used to deliver therapeutically effective amounts of pharmaceuticals, such as drugs and vaccines, in aerosol form to the user through their respiratory system. Inhalation therapy devices can also be used for diagnostic purposes, such as in lung load tests using radioisotopes. Regarding therapy, aerosol inhalation is a preferred source of administration for certain pharmaceuticals that may be intended for the treatment of systemic or respiratory diseases.

[0003] To achieve the intended treatment, aerosol particles must deposit in specific locations within the user's airways, such as the lungs. Different particle sizes tend to deposit in different parts of the respiratory system. In this field, it is generally known that particles with a median mass diameter (MMD) between 1 μm and 5 μm, or between 1 μm and 4 μm, and optionally at least 1 μm, are necessary to enable deposition in the intended location of the respiratory system.

[0004] Aerosols for therapeutic purposes are generated using an inhalation therapy device and delivered to a desired location within the user's body, particularly to their airways. Therefore, fluids to be aerosolized or atomized, especially liquids such as drugs, medications, and vaccines, are supplied to a fluid or liquid reservoir within the inhalation therapy device. To aerosolize or atomize the liquid in the reservoir, the inhalation therapy device may include, for example, a membrane unit.

[0005] Such a membrane unit comprises a membrane having multiple apertures and actuators directly or indirectly coupled to the membrane (via support plates supporting the membrane) for vibrating the membrane.

[0006] A fluid or liquid can come into contact with a first side of the membrane by gravity. When the membrane is vibrated, the liquid or fluid (i.e., drug) passes through the aperture from the first side of the membrane, and an aerosol is generated on the second side of the membrane opposite to the first side.

[0007] The aerosolized or atomized liquid or fluid, i.e., aerosol, within the inhalation therapy device is supplied to the user's respiratory system within the scope of inhalation therapy when the user inhales it through the device. In other words, when the user applies suction force through breathing, particularly inhalation, the aerosol is at least partially transported (by suction) from inside the inhalation therapy device into the patient's airway.

[0008] An example of a conventional inhalation therapy device having such a configuration can be derived, for example, from Patent Document 1. The membrane unit of the disclosed inhalation therapy device comprises a cylindrical liquid reservoir container, one end of which is separated by a disc-shaped membrane. The liquid placed in the reservoir container is in contact with the side of the membrane facing the reservoir.

[0009] Patent Document 2 further discloses a membrane connected (welded) to a substrate (support plate) and a vibration generator, such as a piezoelectric crystal, that surrounds the membrane circumferentially and is connected (bonded) to the substrate, thereby allowing the membrane to be vibrated by the vibration generator and an electric drive circuit. As a result, a liquid applied to a first side of the membrane is carried through the aperture of the vibrating membrane to a second side of the membrane opposite to the first side and discharged into the chamber as an aerosol. The disclosed inhalation therapy device further embodies a mouthpiece or mask connected to the outlet of the chamber and having a discharge valve. As a result, when the user exhales, the discharge flow does not enter the chamber, or enters only to a very limited extent, and is rather discharged to the external environment of the device through the discharge valve.

[0010] The disclosed device continuously drives a piezoelectric crystal throughout the treatment time (a single dosing interval). During this time, substantially all of the liquid in the reservoir is aerosolized by the membrane. During the user's exhalation, the generated aerosol is temporarily stored in the chamber, forming an aerosol cloud (a large clump). This aerosol cloud is inhaled upon inhalation and delivered to the airway for treatment. Thus, it is possible to keep the treatment time relatively short. However, one potential drawback is that as the total power rate of the aerosol generator (membrane, piezoelectric crystal, drive circuit, etc.) increases, the amount of aerosol that accumulates on the walls of the device, particularly in the chamber (rainout), increases, which increases the loss of administerable aerosol / pharmaceutical.

[0011] In other conventional inhalation therapy devices, the piezoelectric crystal is driven only during inhalation and stopped during exhalation. These inhalation therapy devices are often called respiratory trigger devices or respiratory actuation devices. These inhalation therapy devices often have the problem of long treatment times (long dosing intervals), which can be inconvenient for the user and may result in poor user adherence, such as interruption or premature termination of treatment before the intended dose of liquid in the reservoir is atomized and delivered to the user's airway for treatment.

[0012] All of the conventional inhalation therapy devices described above are typically operated by the user simply inhaling through the device (the inhalation flow is generated through the device's pathway), and the device is removed from the user's mouth and / or nose during exhalation, or has an exhalation valve in the mouthpiece or mask to expel the exhaled air. In all cases, the user can exhale without resistance, but because data such as flow parameters in the device's pathway cannot be obtained during exhalation, the entire treatment cannot be adequately monitored. Furthermore, in the case of continuous operation, the generated aerosol is lost without providing any therapeutic effect to the user.

[0013] Therefore, there remains a need for a simple and effective inhalation therapy device that can shorten the total treatment time (with a short single-dose interval).

Prior Art Documents

Patent Documents

[0014]

Patent Document 1

Patent Document 2

Summary of the Invention

Problems to be Solved by the Invention

[0015] Based on the above, the object of the present disclosure is to establish an inhalation therapy device that keeps the total treatment time as short as possible while enhancing or at least maintaining the overall quality of drug delivery.

Means for Solving the Problems

[0016] This object is solved by an inhalation therapy device according to independent claim 1. Different embodiments are derived from the dependent claims and the following description. According to a first aspect, an inhalation therapy device is proposed. This inhalation therapy device includes a housing body and a reservoir for holding liquid, and the reservoir is accommodated within the housing body or formed within the housing body. The reservoir can be an integrated part of the housing body. As an example, the reservoir and the housing body can be integrated as one part (for example, in an injection molding process). In another example, the reservoir and the housing body can be manufactured separately and integrated by being fixed to each other (for example, non-removably fixed). Alternatively, the reservoir can be removably inserted into the housing body and may be an ampoule that is opened. The normal amount of liquid accommodated in the reservoir is from 1 ml to 12 ml, or from 1 ml to 8 ml, although from 1 ml to 6 ml is also possible.

[0017] The inhalation therapy device further comprises a membrane having a plurality of apertures (through holes). The membrane is a circular disk having an effective area at its center, and the apertures are formed in the effective area. The membrane can be manufactured from a metal, such as stainless steel, or other suitable material. The membrane is disposed or housed within the housing body. The liquid in the reservoir can be supplied to the first side of the membrane by gravity. For this purpose, the reservoir may abut the surface of the membrane on the first side and have an outlet opening surrounded by a seal lip surrounding the effective area.

[0018] The inhalation therapy device further comprises an actuator coupled to the membrane for vibrating the membrane. The actuator can be a piezoelectric element, such as a piezoelectric ring. The actuator can be directly coupled or attached to the membrane. That is, it can be coupled or attached to either the first side or the second side opposite the first side of the membrane. Alternatively, the membrane can be coupled or attached to a support plate (such as an annular support plate), and the actuator may be directly coupled or attached to the support plate. When the actuator is activated or operated, the membrane vibrates. When the membrane vibrates, the liquid passes through the apertures, and an aerosol is generated on the second side opposite the first side of the membrane. The membrane and the actuator can constitute components of a membrane unit (nebulizer head).

[0019] The inhalation therapy device further includes or defines a flow path. Specifically, the flow path is located within or defined by the housing body and has a first opening at one end of the flow path to the outside (or environment) of the housing body (or inhalation therapy device), and at least one second opening at the other end to the outside (or environment) of the housing body (or inhalation therapy device). Thus, when a user inhales through the first opening, a flow may be generated in the flow path in a first flow direction (inhalation flow direction) from the second opening to the first opening, entraining and delivering the generated aerosol. In other words, when a user inhales through the first opening, an suction force is generated at the first opening. This creates negative pressure in the flow path, drawing air in from the second opening into the flow path. This generates an airflow that entrains the generated aerosol and delivers it to the user through the first opening. Similarly, when a user exhales through the first opening, a flow may be generated in the flow path in a second flow direction (exhalation flow direction) from the first opening to the second opening. When the user exhales through the first opening, the exhaled air is "blown" from the first opening into the flow path, creating positive pressure within the flow path, causing the exhaled air to flow through the flow path and exit through at least the second opening.

[0020] To deliver the generated aerosol to the user, the membrane or membrane unit may be placed in the flow path between the first and second openings. This configuration is not limited to the in-line placement of the membrane or membrane unit, in which case the membrane unit is placed in the flow path itself, in the airflow between the first and second openings. Thus, the membrane or membrane unit can be placed directly / in-line in the flow path such that it is at least partially surrounded by the flow in the flow path. In other words, an envelope flow or covering flow is generated around the membrane or a portion of its periphery (described later). Alternatively, the membrane or membrane unit can be placed in a separate channel that is in fluid communication with the flow path so that the generated aerosol can be equally released or transported into the flow path. One example of such a configuration is to connect the channel to the flow path in a T-shape.

[0021] The inhalation therapy device further includes a sensor configured to detect or sense the flow or flow parameters within the flow path. In this situation, the sensor itself does not need to be located within the flow path. Rather, it is sufficient that some fluid communication is established between the flow path and the sensor. For this purpose, the sensor or measuring port may be located within the flow path or within the wall of the flow path and may be connected to the sensor via, for example, a tube or hose. In this regard, various flow parameters to be detected can be considered, such as pressure (absolute pressure, static pressure, total pressure, dynamic pressure, or differential pressure), volumetric flow rate, mass flow rate, acoustic noise, vortex detachment, etc.

[0022] Furthermore, the inhalation therapy device is equipped with a controller. This controller may be housed within the main body of the housing. The controller is configured to determine at least the user's inhalation phase and exhalation phase based on the sensor output. In other words, the controller can determine the user's respiratory cycle, including the start of inhalation, the end of inhalation, the start of exhalation, and the end of exhalation, as well as respiratory rate, inhalation / exhalation ratio (the ratio of inhalation time to exhalation time), flow rate (current, maximum, average, minimum), inhalation volume (current, maximum, average, minimum), exhalation volume (if exhaled through the device), minute ventilation volume, differential value of ventilation volume, etc., through the sensor output. For example, the controller may interpret the pressure signal in a different way to calculate, for example, inhalation volume / exhalation volume, and this calculated value may be used for triggering (membrane activation / operation) as described later.

[0023] Furthermore, inhalation therapy devices are designed to administer drugs or medications contained in a liquid in a single session. In other words, in a single treatment (a single dosing interval or a single treatment time), the inhalation therapy device converts substantially all of the liquid in the reservoir into an aerosol and delivers that aerosol to the user. The treatment time may be the time required to deliver a predetermined amount of medication to the user (their lungs). A single treatment begins after the inhalation therapy device is switched on and the actuator is first operated / activated (e.g., when the sensor first detects the inhalation flow) and ends when there is no liquid left in the reservoir or when there is residual liquid in the reservoir that can no longer be aerosolized (e.g., trapped in the reservoir). In other words, the treatment time can be defined as the period from the first operation of the actuator to the last operation, or when the inhalation therapy device indicates that substantially all of the liquid in the reservoir has been atomized or aerosolized (end of administration). Regarding the latter, it should be emphasized that a small amount of liquid may remain in the reservoir, e.g., trapped (reservoir residue). However, the amount of residual liquid after complete atomization should be less than 1.2 ml, less than 0.7 ml, or less than 0.5 ml.

[0024] According to this disclosure, the controller is configured to repeatedly operate and deoperate an actuator during a single treatment (dosing interval or treatment time). In other words, the actuator is repeatedly activated and deactivated. In this context, operating the actuator is understood as controlling the actuator to vibrate the membrane and generate an aerosol, while not operating the actuator is understood as controlling the actuator to not vibrate the membrane and not generate an aerosol. Specifically, during the inhalation period (at least a certain period during the inhalation / inhalation phase), the controller is configured to operate / activate the actuator, thereby vibrating the membrane and generating an aerosol. During the subsequent exhalation period (at least a certain period during the exhalation / exhalation phase), the controller is configured not to operate / activate the actuator, thereby not vibrating the membrane and not generating an aerosol.

[0025] According to the first embodiment, the aerosol output rate (AOR) at the first opening during a single treatment time is in units of g / min. a) The aerosol droplet has a median mass of at least 0.48 g / min, optionally at least 0.64 g / min, and more optionally at least 0.80 g / min, when the median mass of the aerosol droplet is between 4.5 μm and 5.0 μm, or b) The aerosol droplet has a median mass of at least 0.42 g / min, optionally at least 0.54 g / min, and more optionally at least 0.70 g / min, when the median mass of the aerosol droplet is between 4.0 μm and 4.5 μm, or c) The aerosol droplet has a median mass diameter between 3.5 μm and 4.0 μm, and the aerosol droplet has a concentration of at least 0.29 g / min, optionally at least 0.38 g / min, and more optionally at least 0.51 g / min, or d) The aerosol droplet has a median mass diameter between 3.0 μm and 3.5 μm, and the aerosol droplet has a concentration of at least 0.20 g / min, optionally at least 0.29 g / min, and more optionally at least 0.38 g / min, or e) The aerosol droplet has a median mass diameter between 2.5 μm and 3.0 μm, and the aerosol droplet has a concentration of at least 0.13 g / min, optionally at least 0.19 g / min, and more optionally at least 0.26 g / min.

[0026] Aerosol output rate (AOR) is defined as the mass of aerosol released per unit of time by an inhalation therapy device through its first opening. Therefore, AOR is an indicator of the amount of aerosol received per unit of time by the user.

[0027] DIN EN ISO 27427 (February 2020 edition, in particular Annex C1) specifies a method for measuring the aerosol output rate of inhalation therapy devices. For the measurement, a test solution of 0.1% (M / V) albuterol in a 0.9% sodium chloride solution, as mentioned in DIN EN ISO 27427, was used. A silicone tube was used as a removable connector, for example, for connection to the first opening. A PALL filter was used as the collection filter. A Data M AZS-M13-V7 or Copley BRS-300 was used for the measurement. The test conditions were set as follows:

[0028] [Table 1]

[0029] [Table 2]

[0030] The median mass diameter (MMD) is measured by laser diffraction. This method is also described in "Influence of flow pattern, apparatus, and formulation on particle size distribution of atomized aerosols" (Junhua Hu et al., International Journal of Pharmaceutics Volume 560, April 5, 2019, pp. 35-46).

[0031] The mass median aerodynamic diameter is USP <1601> Alternatively, it is measured by the cascade impact method described in ISO 27427. Ideally, MMAD and MMD are identical for spherical particles with a density of 1.0 g / ml, as is the case for most aqueous aerosols produced by a nebulizer. In one embodiment, the geometric standard deviation (GSD) is less than 2.0, preferably between 1.6 and 1.8, and most preferably less than 1.5.

[0032] Since laser diffraction measurements are far easier to perform than impactor experiments, MMD data are presented here. The droplet size distribution by laser diffraction (LD) was determined according to the following test method: For this purpose, a Sympatec GmbH Helos BF or BR laser diffractometer, including lens R3, was used. An inhalation therapy device was placed inside the chamber, and conditioned air was introduced to deliver the aerosol to the laser beam. The equipment used included an automated control unit (ACU), a Mettler Toledo XS603S DR chemical balance, a Walter Roller GmbH SV462 Euroline plus air conditioning unit, and an Eppendorf multipipette.

[0033] The measurement conditions were as follows: Atomization time: 2 minutes, or total treatment time. Optical density and MMD [μm] are measured every 10 seconds. Ambient temperature: 23.0±2.0℃ Ambient humidity: 50.0±5.0% rh Inspiratory flow rate: 20.0 ± 1 L / min Exhaust flow rate: 90±5L / min Filling amount: 4ml The test solution was pipetteed into the reservoir. The inhalation therapy device was weighed after filling with the test solution and connected to the laser diffraction test system. Measurements were taken for 2 minutes or throughout the entire treatment time, with optical density (opt. conc.) and MMD measured every 10 seconds. After the atomization process was complete, the inhalation therapy device was weighed again. The residue in the inhalation therapy device represents the weight of the remaining solution in the reservoir and the condensed water in the flow path.

[0034] The treatment time for a desired dose is determined by the aerosolization rate and the efficiency of the device. The aerosolization rate can be calculated from the weight loss of the nebulizer during continuous atomization over a certain period. The total power factor (TOR) is defined as the weight loss of the test solution during constant atomization divided by the atomization time, and is expressed in g / min. Therefore, TOR serves as an indicator of the aerosol generation rate during constant atomization. In the laser diffraction method described above, the total power factor (TOR) of the membrane unit can be calculated from the weight loss over the atomization time.

[0035] TOR and AOR are related through the efficiency of the nebulizer, i.e., the proportion of aerosolized fluid delivered to the patient. In particular, the aerosol output rate is governed, among other things, by the total output rate (TOR) of the membrane unit and the potential loss of aerosol in the flow path of, for example, an inhalation therapy device. The higher the TOR of the membrane unit, the higher the aerosol output rate. The less loss, the higher the aerosol output rate. Simply put, the aerosol output rate (AOR) is the value obtained by multiplying the TOR by the amount delivered (the delivered amount, expressed as salbutamol μg, is the amount of drug deposited on the inhalation filter, measured by high-performance liquid chromatography). The delivered amount represents the dose inhaled by the user. For example, if a membrane unit continuously generates an aerosol with an MMD of 4.4 μm at 0.8 g / min (TOR), and 45% of it is delivered to the patient, the corresponding AOR is 0.8 g / min (TOR) × 45% = 0.36 g / min. In the above example, if the membrane unit is only partially activated during inhalation (for example, activated by respiration for 40% of the entire respiratory cycle (inhalation phase)), 80% of the aerosolized liquid is delivered to the user. The resulting AOR is 0.8 g / min (TOR) × 0.4 (activation time / total time) × 0.8 = 0.26 g / min.

[0036] All parameters in this disclosure, including those related to AOR, TOR, MMAD, MMD, etc., are defined as values ​​under atmospheric pressure within the reservoir, i.e., when there is no negative pressure within the reservoir.

[0037] In one embodiment, the total power output (TOR) of the membrane unit is: a) At the median mass diameter of aerosol droplets between 4.5 μm and 5.0 μm, the concentration may be at least 1.5 g / min, optionally at least 2.0 g / min, and more optionally at least 2.5 g / min, or b) At the mass median diameter of aerosol droplets between 4.0 μm and 4.5 μm, the concentration may be at least 1.3 g / min, optionally at least 1.7 g / min, and more optionally at least 2.2 g / min, or c) At the mass median diameter of aerosol droplets between 3.5 μm and 4.0 μm, the concentration may be at least 0.9 g / min, optionally at least 1.2 g / min, and more optionally at least 1.6 g / min, or d) At the mass median diameter of aerosol droplets between 3.0 μm and 3.5 μm, the concentration may be at least 0.6 g / min, optionally at least 0.9 g / min, and more optionally at least 1.2 g / min, or e) At the mass median diameter of aerosol droplets between 2.5 μm and 3.0 μm, the density may be at least 0.4 g / min, optionally at least 0.6 g / min, and more optionally at least 0.8 g / min.

[0038] By combining the triggering of actuator operation / activation, the resulting aerosol generation, and the minimum aerosol output rate defined above, it is possible to achieve a relatively short treatment time and reduce losses caused by factors such as aerosol deposition on the inner surface of the flow path or increased droplet size due to droplets and other aggregations within the housing. With a reservoir liquid volume of 0.5 ml or 1 ml, treatment may be possible with 10 breathing cycles (10 exhalation phases and 10 inhalation phases). Therefore, overall, this inhalation therapy device is more convenient to use and has improved aerosol generation quality.

[0039] It should be noted that all of the following embodiments can also be individually implemented in inhalation therapy devices that do not achieve an aerosol output rate in g / min units of aerosol released from the inhalation therapy device at the first opening during a single treatment as defined in claim 1.

[0040] According to the second embodiment, the controller is configured to initiate the operation / activation of the actuator within each respiratory cycle, which includes one inhalation phase and one exhalation phase, specifically at the end of each exhalation phase and before the start of the next inhalation phase (inhalation period). This process is sometimes called a “pre-on” process because aerosol generation has already begun at the end of the exhalation phase. However, in this context, “each” refers only to all complete respiratory cycles, which include one exhalation phase (inhalation period) and one inhalation phase (inhalation period). In other words, the “pre-on” process is activated only when a preceding exhalation is detected. The duration for which the actuator is operated during the exhalation phase can be between 50ms and 500ms, optionally between 100ms and 300ms, and most optionally between 200ms, for a standard sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml. The length of the “pre-on” phase depends on various factors, for example, the volume of the storage chamber if the storage chamber described later is implemented. The larger the volume of the storage chamber, the longer the "pre-on" phase can be set.

[0041] Therefore, by the time the user begins inhaling, a sufficient amount of aerosol is already present and is immediately inhaled and transported into the airways. Thus, the required dose can be delivered to the user very efficiently in a short amount of time.

[0042] According to a third embodiment, the inhalation therapy device further comprises a storage chamber (aerosol storage chamber) formed in the flow path between the membrane or membrane unit and the second opening. This storage chamber is formed by the flow path itself or by expanding the internal dimensions of the flow path between the membrane or membrane unit at the second opening, thereby ensuring a sufficiently sized storage chamber while maintaining a relatively compact housing body.

[0043] This embodiment is particularly advantageous when combined with the second embodiment described above. Specifically, if the actuator is already activated during the discharge phase, especially at the end of the discharge phase, the discharge flow tends to wash the generated aerosol from the aerosol therapy device into the environment, resulting in a loss of aerosol for therapeutic use. The storage chamber has a sufficiently large volume for the flow path between the membrane or membrane unit and the second opening, temporarily storing the aerosol that is transported to the second opening by discharge, thereby preventing the aerosol from being discharged from the aerosol therapy device, such as from the second opening. The aerosol generated during discharge, especially at the end of discharge and temporarily stored in the storage chamber, is again drawn from the storage chamber upon user inhalation and delivered to the user through the first opening. As a result, the loss of aerosol, and therefore, for example, pharmaceuticals, is minimized, while the treatment time is shortened.

[0044] According to the fourth embodiment, the volume of the storage chamber is between 10 ml and 100 ml, optionally between 20 ml and 80 ml, and more optionally between 25 ml and 75 ml. This volume is defined by the plane formed by the membrane and the second opening.

[0045] These volume ranges have been shown to enable storage chambers large enough to achieve the above effects for a variety of users and breathing patterns, while simultaneously allowing for a relatively compact housing and, consequently, an inhalation therapy device.

[0046] According to a fifth embodiment, the sensor is configured to sense flow parameters in the membrane or between the membrane or membrane unit and the second opening. In this configuration, the membrane itself may be used as the sensor, or the sensor or at least a sensor support may be placed between the membrane or membrane unit and the second opening. In embodiments comprising the storage chamber described above, the sensor or sensor support may be located inside the storage chamber or within the wall of the storage chamber.

[0047] As a result, the sensor or sensor support is positioned within the housing body in a part of the flow path that is less likely to become excessively wetted. In particular, this is a location where aerosols are less likely to collide with and accumulate on the flow path walls. This improves the reliability and sensitivity of the sensor.

[0048] According to the sixth embodiment, the membrane is positioned in a flow channel, i.e., in line, thereby forming an envelope flow or covering flow around at least a portion of the membrane. For example, the membrane may have an annular outer ring, a concentric circular disk having the effective region having an aperture, and a plurality of spokes connecting the annular outer ring and the circular disk. Through holes / openings are formed circumferentially between adjacent spokes. These holes / openings may become part of the flow channel.

[0049] According to the sixth embodiment, collision of the generated aerosol with the inner surface of the flow path can be prevented, or at least minimized, during both inhalation and discharge. During inhalation, the airflow generated in the flow path forms an envelope flow or covering flow around the membrane, encompassing the aerosol and discharging it from the first opening. During discharge, the airflow generated in the flow path may also encompass the aerosol generated during discharge (at the end of discharge) when the “pre-on” process is performed, and may flow into an optional storage chamber. In either case, collision of the aerosol with the inner wall of the flow path, including the optional storage chamber, is reliably prevented. Therefore, aerosol deposition on the inner wall of the flow path and the resulting loss can be avoided.

[0050] In particular, towards the end of a flow, at the end or start of an inhalation or discharge, the flow rate decreases, increasing the risk of aerosol collision with the opposing wall. If the central axis of the membrane (perpendicular to the plane of the membrane), for example the central axis of the effective area, is oriented so that it exits the first opening without interfering with the inner surface of the housing body (flow channel), the opposing wall is moved as far away as possible, thereby further reducing aerosol deposition and associated losses within the flow channel.

[0051] According to the seventh aspect, the controller is configured to stop the operation / activation of the actuator during each respiratory cycle, which includes one inhalation phase and one exhalation phase, within the inhalation phase (inhalation period) and before the start of the exhalation phase (exhalation period), particularly at the end of each inhalation phase and before the start of the exhalation phase. This process is sometimes called a “pre-off” process, in that aerosol generation is already stopped before the end of the inhalation phase (inhalation period). However, “each” in this context refers only to all complete respiratory cycles, which include one exhalation phase (exhalation period) and one inhalation phase (inhalation period). In other words, the “pre-off” process is activated only when a preceding inhalation is detected. The period during which the actuator is stopped at the end of the inhalation phase may be 100ms to 1000ms before the start of the exhalation phase, optionally 300ms to 800ms before, and most optionally 600ms before, for a standard sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml.

[0052] The advantage of stopping aerosol generation before inhalation is that all of the generated and inhaled aerosol reaches the desired location in the respiratory system, rather than remaining in the trachea or throat and consequently being exhaled in the next breathing cycle. Thus, the exhalation of aerosols that do not reach the desired location in the subsequent exhalation phase and do not provide therapeutic effect can be reduced. Therefore, it is possible to achieve improved drug use efficiency and higher deposition accuracy within the respiratory system while keeping the treatment time reasonably short. In addition, the flow path, particularly the flow path between the second opening and the membrane, can be kept dry. Since the atomized / aerosolized liquid may contain compounds that serve therapeutic purposes for the user's disease but are harmful to individuals in the environment in which the user inhales, it is known in the art to use exhalation filters to prevent these components from being released into the environment. An example of such an exhalation filter is shown in European Patent No. 1 868 570 B1. With a "pre-off" process, this filter may be unnecessary because there is no or only a small amount of aerosol released from the device during exhalation. Furthermore, the time required for aerosol generation, and therefore for the actuator to be active, is reduced, which is advantageous in terms of power consumption, especially when the device is battery-powered.

[0053] According to the eighth aspect, the inhalation therapy device further comprises a mixing chamber formed in a flow path between the membrane and the first opening. The mixing chamber may be a volume into which an aerosol is generated or into which an aerosol is released from the membrane.

[0054] Therefore, the mixing chamber provides sufficient volume to prevent the aerosol from colliding excessively with itself or the inner surface (walls) of the flow path, thus avoiding losses. According to the ninth embodiment, the volume of the mixing chamber is between 5 ml and 50 ml, optionally between 10 ml and 40 ml, and more optionally between 12.5 ml and 37.5 ml. This volume is defined between the first opening and the plane formed by the membrane.

[0055] This volume is considered to be large enough to avoid excessive collisions between the aerosol and the inner surface (walls) of the flow path, thereby preventing losses, while also not leading to an increase in the size of the device. In embodiments with an optional storage chamber, the volume ratio of the storage chamber to the mixing chamber is at least 1, optionally at least 1.5, and more optionally at least 2, i.e., the volume of the storage chamber is greater than or equal to the volume of the mixing chamber.

[0056] According to the tenth embodiment, the mixing chamber tapers toward a first opening. In other words, the mixing chamber has its maximum cross-sectional area at the membrane and contains the generated aerosol in such a way that collision with the inner surface of the mixing chamber is minimized. The cross-sectional area of ​​the mixing chamber may taper gradually and / or continuously toward the first opening toward the user from the membrane. Alternatively, the mixing chamber may begin with a constant cross-sectional area and then gradually or continuously taper toward the first opening.

[0057] The first opening needs to have a certain (smaller) cross-sectional area so that the user can comfortably use the device and put it in their mouth through the first opening. In this respect, the aerosol is assisted to reach the first opening without excessive collision with the inner surface of the mixing chamber and without loss due to aerosol deposition on the inner surface of the device defining the flow path.

[0058] According to the eleventh embodiment, the flow parameter is pressure, optionally a differential pressure, and more optionally a differential pressure between the pressure in the flow path and the ambient pressure. Using pressure as a flow parameter has been shown to enable cost-effective and reliable detection / sensing of suction and discharge, thereby allowing for precise triggering (starting / stopping) of actuators. In the case of absolute pressure, two absolute pressure sensors can be provided, for example, upstream and downstream of a flow limiter located in the flow path. If the limiter is located in the outer wall of the housing body (e.g., embodied as one or more second openings), one absolute pressure sensor may be provided to measure ambient pressure, while the other absolute pressure sensor may be provided to measure the pressure in the flow path of the device. Differential pressure can also be used in a similar manner. For example, one pressure port of a differential pressure sensor can be placed upstream of the flow limiter and the other pressure port can be placed downstream. If the limiter is located in the outer wall of the housing body (e.g., embodied as one or more second openings), one pressure port of the differential pressure sensor may be provided to measure ambient pressure, while the other pressure port may be provided to measure the pressure in the flow path. However, it has been demonstrated that using a differential pressure sensor that takes into account the pressure within the flow path (optionally between the second opening and the membrane, and more optionally within the storage chamber) and the ambient pressure is the most reliable and structurally simple approach.

[0059] According to the twelfth aspect, the flow path is configured to provide a first flow resistance (for example, a maximum absolute value within the following range: 160 Pa ± 100 Pa; preferably 160 Pa ± 60 Pa; more preferably 160 Pa ± 30 Pa; measured at the peak flow rate of a sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml) when the user inhales, and a second flow resistance lower than the first flow resistance (for example, a maximum absolute value within the following range: between 120 Pa ± 75 Pa, preferably 120 Pa ± 45 Pa, more preferably 120 Pa ± 22.5 Pa; measured at the peak flow rate of a sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml) when the user exhales. This standard breathing pattern has been used for definitional purposes, but is generally applicable to any “symmetrical” breathing pattern, i.e., where the duration of inhalation and exhalation are the same and have the same progression. In other words, the second flow resistance during exhalation (of the user, pump, or breathing simulator) is at least 25%, at least 30%, or at least 40% lower than the first flow resistance during inhalation (of the user, pump, or breathing simulator).

[0060] This allows for rapid exhalation against reduced exhalation resistance. On the other hand, increasing the inhalation resistance compared to the exhalation resistance can extend the inhalation time. In this regard, the configuration of an inhalation flow resistance of 160 Pa ± 30 Pa and the exhalation resistance of up to 80 Pa is a preferred range that is comfortable for the user while still delivering the generated aerosol to the user's airway. This is because inhalation is sufficiently extended by inhaling against this inhalation flow resistance without placing too much burden on the user, and because the exhalation resistance of 80 Pa or less does not create excessive resistance during exhalation, the exhalation through the device does not give the feeling that it deviates significantly from natural breathing behavior. By increasing the inhalation resistance while simultaneously decreasing the exhalation resistance, four main advantages can be obtained: i) The maximum inhalation flow rate is limited by increased respiratory resistance, resulting in increased drug accumulation in the lungs.

[0061] ii) The duration of inhalation is extended, which increases the time window for aerosol generation, resulting in a shorter treatment time. iii) Increased respiratory resistance prolongs inhalation, while decreased respiratory resistance shortens exhalation. This reduces respiratory effort, allowing the patient to maintain respiratory rate and minute ventilation.

[0062] iii) Inhalation pressure signals generally play a more important role in actuator activation / operation. The increase in absolute pressure during inhalation facilitates triggering (actuator activation / operation) and results in improved therapy (reliability, duration).

[0063] Inhalation resistance can be achieved by shaping the geometric form of the second opening as a fixed-shape passive valve. This enables a particularly simple, compact, and efficient inhalation therapy device. If all other cross-sectional areas in the flow path are larger than the second opening, the inhalation resistance is determined solely by the second opening.

[0064] Discharge resistance can be achieved by shaping the geometric form of the second opening as a fixed-shape passive valve. This enables a particularly simple, compact, and efficient inhalation therapy device. For example, the second opening can be configured such that, due to the hysteresis effect, the resistance in one flow direction (in this case, inhalation) is higher than the resistance in the other flow direction (in this case, discharge). This can be observed, for example, in a Tesla valve. In this case, air flows in (during inhalation) and out (during discharge) substantially through the second opening (with some leakage within the device).

[0065] Alternatively, the inhalation therapy device may further include a third opening or a plurality of third openings leading to the outside of the housing body, the third openings being optionally located in the flow path between the first and second openings, optionally in the flow path between the membrane and the second opening, or more optionally in an optionally selected storage chamber. The third openings in the inhalation therapy device may be located upstream of the membrane unit in the flow path when viewed along the first flow direction. Thus, the third openings can be positioned far enough away from the first openings that come into contact with the user's mouth to avoid discomfort from the air discharged from the inhalation therapy device.

[0066] These third openings are equipped with check valves, which are configured to substantially restrict the flow through the third opening in the first flow direction (intake) (except for some leakage through the valve), and to allow the flow through the third opening in the second flow direction (discharge) from the first opening to the second opening, thereby enabling at least partial discharge of air through the third opening. For example, less than half or one-third of the air is discharged through the third opening (half or one-third of the volumetric flow rate).

[0067] In such an inhalation therapy device, during inhalation, the check valve of the third opening blocks the opening in the flow path, allowing the inhalation flow through the second opening. However, during exhalation through the inhalation therapy device, the third opening and its check valve reduce the air discharge resistance, thereby reducing the resistance to air discharge from the inhalation therapy device. Therefore, a portion of the discharged air (see above) can be discharged through the freely opening check valve of the third opening and does not need to pass through the second opening and its high discharge resistance. Thus, the second flow resistance is reduced compared to the first flow resistance, preferably at least 25% lower than the first flow resistance.

[0068] According to the 13th embodiment, the first flow resistance is 1.25 to 6 times, optionally 1.5 to 3 times, and more optionally 1.5 to 2.5 times, the second flow resistance. Such ratios have been shown to be advantageous in limiting inhalation flow, helping to prolong the inhalation phase and shorten the exhalation phase, and increasing the inhalation flow parameter signal, resulting in a higher overall therapeutic success rate without causing discomfort such as shortness of breath to the user.

[0069] According to the 14th embodiment, the aperture penetrates the membrane in an elongation direction from the first side to the second side, and each aperture has a nozzle portion (minimum diameter portion) along its elongation direction on the second side. The length of this nozzle portion is between 5 μm and 16 μm, optionally between 5 μm and 14 μm, and more optionally between 5 μm and 12 μm. In one embodiment, the length of the nozzle portion may be 9 μm. In this regard, the nozzle portion may be funnel-shaped or tapered depending on its manufacturing method, and the minimum cross-sectional area (diameter) is perpendicular to the elongation direction of the aperture on the second side.

[0070] This invention is based on the discovery that the length of the nozzle portion of an aperture formed on a vibrable membrane has a significant effect on the total power factor (TOR), and therefore the aerosol power factor (AOR). In particular, the length of the nozzle portion is directly proportional to the TOR; the shorter the nozzle portion, the higher the TOR, and vice versa. On the other hand, the diameter and length of the upstream portion of the nozzle portion within the aperture have little effect on the TOR if the nozzle portion is sufficiently short and small in diameter compared to the upstream portion of the aperture. In this situation, the diameter of the nozzle portion on the second side of the membrane is preferably between 1 μm and 5 μm, optionally between 1 μm and 4 μm, and more optionally between 2 μm and 3 μm. Meanwhile, the diameter distribution of the nozzle portion also affects the geometric standard deviation (GSD) of the droplet size distribution. A low GSD characterizes a narrow droplet size distribution (droplets of uniform size), which is advantageous for targeting the aerosol to the respiratory system. For particles with a size of less than 5 μm, the GSD is less than 2.0, preferably between 1.6 and 1.8, and most preferably less than 1.5.

[0071] The length of the nozzle portion can be defined as the portion from the second side towards the first side up to the next laser drilling stage (described later). Furthermore, the total length of the aperture in the extension direction, and therefore the film thickness, is preferably at least 40 μm, optionally at least 50 μm, preferably at least 70 μm, optionally at least 80 μm, most preferably at least 90 μm, optionally at least 100 μm, and a maximum of 110 μm.

[0072] According to the 15th embodiment, the aperture is laser-perforated in at least two steps, for example, three steps, with the nozzle portion being formed in one step and the remainder of each aperture being formed in the remaining steps.

[0073] A vibrable film may comprise a single layer of a single material. A single layer of a single material does not have a clear boundary, interface, or constraining layer in its polished cross-section image. In one embodiment, the vibrable film may consist of this single layer.

[0074] Instead of achieving an aperture that completely penetrates a single layer of a single material as a through-hole through two or more laser drilling steps, multiple recesses can also be formed in a single layer, for example, by laser milling or etching. Each recess has an opening on the fluid side (first side) of the vibrable membrane. Each recess has a bottom surface on the side opposite the opening, i.e., facing the aerosol side (second side). Recesses are sometimes called blind holes introduced from one side of the single layer (from the fluid side (first side) of the vibrable membrane). In one embodiment, at least two recesses are provided. In another embodiment, at least three recesses are provided. Furthermore, the number of recesses is not limited, but fewer than 3,000, fewer than 1,000, or fewer than 500 may be preferred. Alternatively, fewer than 50 or fewer than 20 recesses may be provided. To achieve high stability of the vibrable membrane, a relatively large number of small recesses, such as 500 to 3,000 or 500 to 1,000, may be beneficial. The more recesses a vibrating membrane has, the more the remaining reinforcement between the recesses ensures the full thickness of the single material layer, thus increasing stability. Conversely, a smaller number of large recesses, such as 2 to 50 or 2 to 20, can be beneficial in reducing dead areas and increasing the effective area for atomization. In this context, a "dead area" refers to a region of the vibrating membrane that lacks a penetrating aperture and therefore does not participate in the atomization process. In other words, fluid does not pass through the vibrating membrane in these regions and is not atomized on the fluid side. Furthermore, a smaller number of recesses can simplify the manufacturing process. For example, the area of ​​the recesses does not exceed 50% of the total area of ​​the vibrating membrane.

[0075] Furthermore, multiple apertures are formed on the bottom surface of each recess. Each of these multiple apertures may have an inlet opening formed on the bottom surface of the respective recess and an outlet opening on the aerosol side (second side) of the vibrating membrane. In one embodiment, no additional apertures are formed on the vibrating membrane other than those formed on the bottom surface of the recesses. In another embodiment, at least no additional apertures are formed in the single layer between the recesses; they are blind holes. Therefore, when the vibrating membrane vibrates, the fluid on the fluid side of the vibrating membrane passes through the recesses and apertures and is atomized as it exits the aperture through the outlet opening on the aerosol side of the vibrating membrane. In this way, an aerosol is formed on the aerosol side of the vibrating membrane and inhaled by the user. In this respect, since the inlet openings of the apertures can be made relatively small, the aperture density can be increased compared to the density of apertures laser-perforated in three stages, thus enabling an improvement in TOR (Transient Ratio).

[0076] Furthermore, TOR and MMD, and consequently the aerosol output rate, strongly depend on the nozzle section, i.e., the aperture shape (especially its length and diameter). The formation of a recess results in a smaller, relatively uniform, or constant thickness at the bottom of the recess. Therefore, the aperture length can be determined very precisely. Moreover, the formation of a relatively short aperture allows for higher precision in terms of diameter at both the inlet side (corresponding to the fluid side) and the outlet side (corresponding to the aerosol side). As a result, the geometric consistency between apertures can be greatly improved, and GSD is reduced. Consequently, the aerosol (atomized fluid) properties become more uniform even between different aperture plates obtained from the same manufacturing process.

[0077] In one embodiment, the recess is laser-milled and the aperture is laser-perforated. Using a laser milling process to form the recess is also beneficial in terms of maintaining a relatively uniform and constant residual material thickness in the recess bottom region. This allows for more precise formation of the aperture in terms of diameter and length, providing advantages related to the reproducibility of certain aerosol properties. In a specific example, an ultrashort pulse laser is used for laser milling. An ultrashort pulse laser is defined as a laser that uses laser pulses of less than 10 picoseconds. Ultrashort pulse lasers offer high precision and further enhance the advantages related to the reproducibility of certain aerosols. The aperture can also be formed using a short pulse laser (e.g., a nanosecond laser) or an ultrashort pulse laser. A short pulse laser is defined as a laser that uses laser pulses of more than 10 picoseconds but less than 500 nanoseconds.

[0078] Furthermore, it has been found that increasing the number of apertures in the membrane can further improve the TOR (Total Error). This can be achieved by performing laser drilling in two or more stages to create through-holes that completely penetrate the membrane from the first side to the second side, or by first forming a recess and then forming an aperture only on its bottom surface. This can be achieved by increasing the effective perforated surface (effective area) of the membrane while keeping the spacing between apertures constant, or by reducing the spacing between through-holes while maintaining the effective area of ​​the membrane. Moreover, it is possible to combine these measures. From this perspective, it is advantageous for the membrane to have 4,000 to 12,000 apertures, optionally 5,000 to 11,000 apertures, and more optionally 5,000 to 10,000 apertures.

[0079] Furthermore, the aperture density is 140 per mm². 2 ~3800 pieces / mm 2 During this period, optionally 170 pieces / mm 2 ~3800 pieces / mm 2 During this period, more selectively 170 pieces / mm 2 ~2300 pieces / mm 2 It can be between 3 mm. The effective area of ​​the film is 3 mm.2 ~28 mm 2 Between, optionally 7 mm 2 ~20 mm 2 Between, more optionally 10 mm 2 ~15 mm 2 It can be between. The distance between two adjacent apertures, that is, the distance between the central axes of adjacent apertures, can be between 15 μm and 100 μm, optionally between 30 μm and 80 μm, and more optionally between 40 μm and 60 μm.

[0080] Hereinafter, non-limiting examples of the present disclosure will be described with reference to the drawings.

Brief Description of Drawings

[0081] [Figure 1] An inhalation therapy device according to the present disclosure is shown in an isometric view. [Figure 2] A side view of the inhalation therapy device of FIG. 1 placed on a horizontal plane is shown. [Figure 3] A longitudinal sectional view of the inhalation therapy device of FIG. 1 is shown. [Figure 4] An exploded view of the inhalation therapy device of FIG. 1 is shown. [Figure 5] A side view of the inhalation therapy device in the use position is shown. [Figure 6] The control of the inhalation therapy devices of FIGS. 1 to 5 is schematically shown. [Figure 7] A partial cross-section of the active region of a vibratable membrane according to one embodiment is shown. [Figure 8] A partial cross-section of the effective region of a vibratable membrane according to another embodiment is shown. [Figure 9] Curves (pressure over time) of two consecutive breathing cycles (n and n + 1) are shown, and each breathing cycle includes inhalation (n, n + 1) and exhalation (n, n + 1).

Modes for Carrying Out the Invention

[0082] The inhalation therapy device described in this specification will be described with reference to the accompanying drawings. It should be understood that such inhalation therapy devices 1 are often also called "aerosol delivery devices" or "nebulizers" when used for therapeutic purposes, for example, in / through the user's respiratory system.

[0083] Furthermore, the term “user” as used throughout this disclosure may refer to a patient. Where a particular angle is defined in this disclosure, a positive angle is always taken in a counterclockwise direction between each side.

[0084] Figure 1 shows an exemplary isometric view of the inhalation therapy device 1. The inhalation therapy device comprises a housing body 4. The housing body 4 is divisible into two main components 16 and 17, as can be most clearly seen in Figure 4.

[0085] In particular, the housing 4 comprises a controller housing (sometimes called a holding portion or handle portion) 16 and a nebulizer housing 17 (sometimes called an aerosolizing portion). However, the disclosure is not limited to such a “split” configuration. It may also be possible to form the controller housing 16 and the nebulizer housing 17 as a single unit.

[0086] In this embodiment, the controller housing 16 also functions as or constitutes a handle portion 12 during therapy, and the user holds the inhalation therapy device 1 by gripping the handle portion.

[0087] The controller housing 16 further optionally includes a support surface 19 within the handle portion 12 on its underside, which is for placing the inhalation therapy device 1 on a horizontal surface such as a table. The support surface 19 can be defined by a plurality of legs 20 positioned on the bottom / underside of the controller housing 16. In other words, the housing body 4, in particular the controller housing 16, has legs 20, and the support surface 19 is a level surface formed by the respective stand surfaces of the legs 20.

[0088] The nebulizer housing 17 has, or defines, a reservoir 2 for holding the liquid to be aerosolized (see Figure 3). The reservoir 2 has a filling opening 25 for receiving the liquid and a lid 26 for closing the filling opening 25 (see Figure 4). The angle α2 between the support surface 19 and the central axis CA2 of the filling opening 25 is greater than 90°±5° in the side view (longitudinal cross-section), and is an obtuse angle between 100° and 160°, and between 110° and 150°, arbitrarily selected. The angle α3 between the central axis CA2 of the filling opening 25 and the central axis CA1 of the mouthpiece is between 70° and 110°, arbitrarily selected between 80° and 100°, and further arbitrarily selected between 90°±5° in the side view (longitudinal cross-section).

[0089] The nebulizer housing 17 further defines a flow path 7 having a first opening 8 at one end that leads to the outside of the housing body 4, particularly the nebulizer housing 17, and at least one second opening 9 at the other end that leads to the outside of the housing body 4, particularly the nebulizer housing 17, thereby enabling a flow to be generated in the flow path 7 in a first flow direction A from the second opening 9 to the first opening 8 to carry and deliver the generated aerosol when the user inhales through the first opening 8, and enabling a flow to be generated in the flow path 7 in a second flow direction B from the first opening 8 to the second opening 9 when the user exhales through the first opening 8.

[0090] In this regard, the aforementioned flow in the first flow direction A from the second opening 9 to the first opening 8 can also be understood as an "inhalation flow." Since the user's exhalation is also performed through the flow path 7 of the inhalation therapy device 1, the second flow direction B from the first opening 8 to the second opening 9 in the flow path 7 when the user exhales at the first opening 8 is understood as an "exhalation flow."

[0091] In this example, the first opening 8 is part of the mouthpiece 27 and is defined by the outer edge 28 of the mouthpiece 27. The mouthpiece 27 is configured to be inserted into the user's mouth to inhale the generated aerosol.

[0092] The mouthpiece 27, particularly its first opening 8, defines a central axis CA1. The central axis CA1 is not parallel to the support surface 19. The central axis CA1 of the first opening 8 intersects the virtual extension 13 of the support surface 19 on the mouthpiece 27 side relative to the support surface 19, and the angle α1 between the virtual extension of the support surface 19 and the central axis CA1 of the first opening 8 is an acute angle in the side view (longitudinal section view) (Figure 3) that is less than 75°, optionally between 10° and 70°, optionally between 15° and 60°, or optionally between 20° and 50°.

[0093] Therefore, the outer edge 28 of the mouthpiece 27 is oriented toward the support surface 19. However, as is clear from Figures 2 and 3, the outer edge 28 does not interfere with the virtual extension 13 of the support surface 19, and therefore does not come into contact with it when placed on a horizontal surface such as a table. The minimum distance D between the virtual extension 13 of the support surface 19 and the outer edge 28 of the mouthpiece 27 in a direction perpendicular to the virtual extension 13 of the support surface 19 is greater than 3 mm, optionally greater than 4 mm and less than 15 mm, and optionally less than 12 mm.

[0094] Furthermore, a mixing chamber 42 is formed between the first opening 8 and the membrane unit 3, particularly the membrane 5, i.e., downstream of the membrane unit 3 in the first flow direction A. The mixing chamber 42 is a volume defined between the first opening 8 and the membrane unit 3. Referring to Figure 3, the mixing chamber 42 has a first portion 42.1 adjacent to the membrane unit 3 with a constant cross-sectional area. In the direction toward the first opening 8, the cross-sectional area of ​​the mixing chamber 42 decreases sharply, and in the second portion 42.2, the cross-sectional area becomes substantially constant. In the third portion, a continuously tapering portion 42.3, the mixing chamber 42 continuously tapers toward the first opening 8. Downstream of the third portion 42.3, the mixing chamber 42 is a fourth portion 42.4 again with a constant cross-sectional area, ending at the first opening 8. However, it is conceivable, and in some cases even preferable, to omit portion 42.2.

[0095] Furthermore, the nebulizer housing 17 comprises a peripheral wall 10, an upper wall 11 with a filling opening 25, and a bottom wall 51. The transition between the upper wall 11 and the peripheral wall 10 may be chamfered 52. The mouthpiece 27 is connected to the peripheral wall 10.

[0096] The second opening 9 is positioned, for example, adjacent to the upper wall 11 and, in this embodiment, within the chamfered portion 52. Therefore, the second opening 9 is positioned as far away from the handle portion 12 as possible, minimizing the risk of blockage of the second opening 9 by the user's hands or fingers. The second opening 9 functions as the only opening to the flow path 7 through which air can enter when the user inhales. Furthermore, because the second opening 9 provides the smallest cross-sectional area throughout the entire flow path 7, it exclusively dominates the first flow resistance, i.e., the inhalation flow resistance.

[0097] The nebulizer housing 17 may further include a third opening 14 or a plurality of third openings 14, which are sometimes referred to as inhalation openings. The third openings 14 may be provided in the circumferential wall 10. Each of the third openings 14 is covered by a check valve 15, in particular a flap valve (see Figures 1 and 2), which covers the third opening 14 on the outside of the nebulizer housing 17. The check valve 15 is configured to restrict the flow through the third opening 14 in a first flow direction A. That is, during inhalation, the check valve 15 completely prohibits the flow through the third opening 14 in the flow path 7. However, the check valve 15 is also configured to allow the flow through the third opening 14 in a second flow direction B from the first opening 8 to the second opening 9 in order to allow at least partial discharge of air through the third opening 14 during discharge. The third opening 14, which leads to the outside of the housing body 4 / nebulizer housing 17, is located within the flow path 7 between the first opening 8 and the second opening 9, particularly between the membrane unit 3 (described later) and the second opening 9, that is, within the storage chamber 40 (described later) in this embodiment. In other words, the third opening 14 is located upstream of the membrane unit 3 when viewed in the first flow direction A.

[0098] The storage chamber 40 is formed within the flow path 7 between the second opening 9 and the membrane unit 3, particularly the membrane 5, that is, upstream of the membrane unit in the first flow direction A. The storage chamber 40 is the volume defined between the second opening 9, the third opening 14 if present, and the membrane unit 3.

[0099] The membrane unit 3 is housed within the nebulizer housing 17. The membrane unit 3 comprises a membrane 5 and an actuator 6. In the art, such a membrane unit 3 is also called a “head unit” or “head”.

[0100] The membrane 5 has a first side 5.1 (liquid side) and a second side 5.2 (aerosol side). The membrane 5 may be circular and may have a circular effective region 62 in its center. Multiple apertures 110 are provided in the effective region 62. Furthermore, the effective region 62 may rise in a dome shape toward the second side 5.2.

[0101] The aperture 110 can be perforated as shown in Figure 7. The aperture 110 penetrates the membrane 61 from the liquid side 5.1 to the aerosol side 5.2 in the effective region 62. During use, when the membrane 61 is vibrated, the liquid solution 26 moves through the aperture 110 from the liquid side 5.1 to the aerosol side 5.2, thereby generating an aerosol on the aerosol side 5.2, which is then released into the mixing chamber 42. This aerosol is then inhaled by the user through the mouthpiece 27 from the mixing chamber 42.

[0102] Figure 7 shows a cross-sectional view (schematic CT image) of three apertures 110 of such a vibrable membrane 5. The through-holes 110 in this particular embodiment are formed by laser drilling using three stages, each with different process parameters. In the first stage, a portion 114 is formed. In the second stage, a portion 116 is formed, and in the third stage, a nozzle portion 112 is formed. In this embodiment, the average length of the nozzle portion 112 is 29.4 μm, while the average length of the portion 116 in the second stage is 55.7 μm. The average length of the portion 114 in the first stage is 16.3 μm. As a result, the total length of each aperture 110 is the sum of the lengths of the first portion 114, the second portion 116, and the nozzle portion 112, i.e., 101.4 μm in this particular example. Thus, the ratio of the total length of each aperture 110 to the length of the corresponding nozzle portion 112 is approximately 4.3.

[0103] Alternatively, as shown in Figure 8, the aperture 110 can also be formed by laser milling and laser drilling. In this example, the vibrable membrane 5 consists of a single layer 126 of a single material (e.g., stainless steel), similar to the previous example. However, in other embodiments, the vibrable membrane 5 may include additional layers attached to the liquid side 5.1 and / or the aerosol side 5.2. Furthermore, other biocompatible metals may be used instead of stainless steel.

[0104] The exemplary vibrable membrane 5 comprises a plurality of recesses 118. The recesses 118 may be circular or annular. The recess 118 has an opening 120 on the liquid side 5.1 of the vibrable membrane 5. The recess 118 is formed as a blind hole with a bottom surface 122 on the opposite side of the opening 120.

[0105] The recess 118 has a circumferential side wall 124. In the case of a circular recess 118, the circumferential side wall 124 corresponds to the covering of a cylinder. However, in the cross-section of Figure 8, several parts of the circumferential side wall 124 are facing each other.

[0106] The recess 118 can be formed within the single layer 128 using an ultrashort pulse laser. In this case, the recess 118 is preferably formed by laser milling, in which the laser beam and / or the single layer 128 of the vibrating film 5 move relative to each other, thereby gradually removing the material of the single layer 128.

[0107] Minimum dimension D of recess 118 R This can be 300 μm. However, other dimensions are also possible. For example, the minimum dimension D of the opening 120 of the recess 118 facing the liquid side 5.1. R This can be between 40 μm and 500 μm, between 70 μm and 400 μm, or between 90 μm and 300 μm.

[0108] In this situation, the minimum dimension D of the circular central recess 118 R This corresponds to the diameter of the circle. The thickness T of the single layer 128 can be 100 μm. Depth or length L of recess 118 R The thickness T of the single layer 128 may be selected between 80% and 95%. Therefore, the depth L of the recess 10 R This can be between 80 μm and 95 μm.

[0109] However, the thickness T of the single layer 128 can be set within the range of 50 μm to 200 μm. Furthermore, the depth L of the recess 10 R This can be selected so as to be 50% or more of the thickness T of the single layer 128.

[0110] Each recess 118 is formed on the bottom surface 122 and has a plurality of apertures 110 that extend from the bottom surface 122 to the aerosol side 5.2 of the single layer 128. Each recess 118 may have at least 2, at least 20, at least 50, or at least 100 apertures 110 formed on the bottom surface 122. The maximum number of apertures 110 that each recess 118 may have may be 5,000 or 10,000.

[0111] Each aperture 110 has an inlet opening 128 on its bottom surface 122 and an outlet opening 130 on the aerosol side 5.2 of the single layer 128. The aperture 110 may be substantially cylindrical or conical, with the smaller opening located on the aerosol side 9, and thus being the outlet opening 130. However, the shape is not limited in this respect, and other shapes may be similarly conceivable.

[0112] The size of the outlet opening 130 of aperture 110 has a significant impact on MMD. Depending on the fluid, the MMD depends on the diameter D of the outlet opening 130. A It is linearly larger than and therefore directly proportional. Thus, aperture 110 is circular, with diameter D facing the aerosol side 5.2. A The diameter of the outlet opening 130 is between 1 μm and 7 μm, preferably between 1.5 μm and 5 μm, more preferably between 1.5 μm and 4.5 μm, and most preferably between 1.5 μm and 3.5 μm. These values ​​also apply to the embodiment shown in Figure 7.

[0113] Length L of aperture 110 in single layer 128 A The height (also called the height or depth) can be between 3 μm and 50 μm. In another example, the length L of the aperture 118 in monolayer 128. A This can be between 5 μm and 20 μm. In yet another example, the length L of the aperture 110 in monolayer 128. A The aperture can be between 10 μm and 20 μm. Other ranges may be 5 to 15 μm, 5 to 12 μm, or 10 to 15 μm. In this example in Figure 8, the aperture forms the nozzle portion, especially when formed by a single laser milling step. These values ​​also apply to the embodiment in Figure 7.

[0114] Furthermore, TOR and MMD, and consequently the aerosol output rate, strongly depend on the shape of the nozzle portion 112, that is, the shape of the aperture 110 (especially its length and diameter). The aperture 110 may be formed by the same ultrashort pulse laser used to form the recess 118. However, short pulse lasers, such as those used in the prior art, may also be used to form the aperture 110.

[0115] In any case, it is preferable to form the aperture 110 in a single laser drilling step while keeping the laser fluence substantially constant. It is preferable to form the aperture in only one drilling step. This improves the shape accuracy of the aperture, i.e., the nozzle portion 112. When using only one laser drilling step, the diameters of the inlet side (corresponding to the liquid side) and the outlet side (corresponding to the aerosol side) are defined very clearly, and the aforementioned advantages are obtained.

[0116] In another embodiment, the aperture 110 at the bottom of the recess 118 can also be perforated in two laser perforation stages having different fluences, similar to the embodiment using three perforation stages described above. For example, a first laser perforation stage with a first fluence can be used to form a first portion of the aperture at the bottom of the recess, and a second laser perforation stage with a second fluence lower than the first fluence can be used to complete the aperture, i.e., to completely penetrate the bottom. In this case, the nozzle portion is defined only by the second laser perforation stage. Therefore, when multiple laser perforation stages are used to form the aperture, the bottom can remain thicker (the recess is shallower). This makes it possible to increase the mechanical rigidity of the aperture plate.

[0117] The manufacturing method may include, as a first step, forming a recess 118 from the liquid side 5.1 of the single layer 128. As a second subsequent step, an aperture 110 is formed on the bottom surface 122 of the recess 118 from the liquid side 5.1 of the single layer 128.

[0118] Subsequently, the vibrable membrane 5 can be formed into a dome shape. The membrane 5 can be fixed to a support. However, in this embodiment, the membrane 5 has an annular portion around the active region 62. The membrane 5 further comprises an annular ring 29 concentric with the membrane 5, particularly its annular portion and the active region 62. The annular ring 29 is connected to the membrane 5 via a plurality of spokes 32. Furthermore, the membrane 5 has an annular portion around its active region 62. When the actuator 6 is directionally coupled, it is attached to the annular portion around the active region 62 of the membrane 5.

[0119] The actuator 6 is electrically connected to the plug 33 of the membrane unit 3 for electrical connection to the controller 18, which will be described later. As is particularly clear from Figure 3, the membrane unit 3 is positioned within the nebulizer housing 17 to close the supply opening 34 of the reservoir 2, which is located at the bottom of the reservoir 2. More precisely, the membrane 5 closes the supply opening 34 at its first side 5.1. To ensure a seal, a seal lip 35 is provided around the supply opening 34 (see Figure 4). The membrane 5 abuts the seal lip 35 at its first side, thereby enclosing the effective area 62 of the membrane 5 with the seal lip 35. As a result, the liquid filling the reservoir 2 is supplied by gravity to the supply opening and, consequently, to the first side 5.1 of the membrane 5.

[0120] To allow for replacement and / or cleaning of the membrane unit 3, the membrane unit 3 is replaceable. In one embodiment, the nebulizer housing 17 comprises a body 30 and a door 31. The peripheral wall 10 is formed by the body 30 and the door 31. The door 31 is hinged to the body 30. Thus, the door 31 is movable between an open position and a closed position. In the open position (see Figure 4), the membrane unit 3 is inserted into the head mount 36 so that the periphery of the effective area 62 on the first side 5.1 of the membrane 5 abuts against the seal lip 35. To prevent the membrane unit 3 from being installed incorrectly, the head mount 36 is provided with a keyway 37 into which the key element 38 of the membrane unit 3 is inserted during installation. When the door 31 is closed, the membrane unit 3 is pressed toward the body 30, thereby pressing the first side 5.1 of the membrane 5 against the seal lip 35. In the closed position, the door 31 is secured to the body 30 using a latch 5.

[0121] When the membrane unit 3 is placed in the flow path, a cavity is formed between the membrane unit 3 and the second opening 9, which is considered difficult to clean. By configuring the nebulizer housing 17 to include a main body 30 and a hinged door 31, the membrane unit 3 can be removed, and the cavity can be accessed for cleaning. However, it is not essential to provide a latch 18 or to configure the nebulizer housing 17 to include a main body 30 and a hinged door 31. It is also possible to provide a nebulizer housing 17 that can be opened in a different way by configuring the main body 30 and the door 31 separately and / or providing some kind of fastening device (magnet, latch, press-fit mechanism, zipper, push button, etc.) to fasten them together. Furthermore, it is possible to provide a nebulizer housing 17 that cannot be opened by the user, operator, hospital staff, or the user themselves. This can be achieved, for example, by forming a nebulizer housing 17 in which the membrane unit 3 is fixed in place and cannot be replaced. In this case, the entire nebulizer housing 17 is disposable and may be discarded from time to time.

[0122] As is most evident from Figure 3, the angle α4 between the support surface 19 and the membrane 5 exceeds 90°±5° in the side view (longitudinal cross-sectional view), and is arbitrarily between 100° and 160°, and further arbitrarily between 110° and 150°.

[0123] Considering the points mentioned above, the nebulizer housing 17 positions the mouthpiece 27, accommodates the membrane unit 3, and forms a flow path 7 between the first opening 8 and the second opening 9. As is clear from Figure 3, the membrane 5 of the membrane unit 3 is positioned within the flow path 7. Furthermore, during inhalation, fluid enters the housing body 4, particularly the nebulizer housing 17, through the second opening 9 and flows into the storage chamber 40 formed between the second opening 9 and the membrane 5 of the membrane unit 3. In addition, a flow path opening 39 (see Figure 4) is formed in the wall 41 with the head mount 36, thereby allowing air to flow from the storage chamber 40 through the flow path opening 39 and the open space between the spokes 32 of the membrane 5 into the mixing chamber 42. Thus, an envelope flow is formed around at least a portion of the circumference of the membrane 5, more specifically, in this embodiment, around a circular disk having the effective area 62 of the membrane 5. In other words, the membrane unit 3 may be positioned within the flow path 7 so that a covering flow or envelope flow can pass around the membrane unit 3 within the flow path 7 during inhalation and exhalation by the user during breathing via the inhalation therapy device 1. In one embodiment, the membrane unit 3 may be positioned substantially (±10° or ±5°) at the axial center. In particular, it is useful that the normal plane of the membrane is as parallel as possible to the mouthpiece axis CA1). Direct spraying in the direction of the first opening 8 is desirable to minimize aerosol deposition within the device. For this purpose, it is desirable that the membrane 5 be positioned (almost) perpendicular to the mouthpiece axis CA1. Therefore, the normal plane of the membrane (CA3) must not intersect the inner wall of the channel 7.

[0124] The aerosol is generated in the mixing chamber 42 when the user inhales it, or when the actuator 6 operates on the second side 5.2 of the membrane 5 while the user is not breathing through the device. Thus, the envelope flow generated during inhalation encompasses the generated aerosol and is discharged through the first opening 8 (the outlet opening of the mouthpiece 27) and administered / supplied to the user.

[0125] When the user exhales, the exhaled air enters the flow path 7 through the first opening 8 of the mouthpiece 27 and flows into the storage chamber 40 via the mixing chamber 42 and the open space provided around the membrane 5 between the spoke 32 and the flow path opening 39. The exhaled air can then be discharged from the flow path 7 through the second opening 9 and the third opening 14. During exhalation, i.e., in the second flow direction B, the check valve 15 is configured to open and discharge the air passing through the third opening 14 to the outside of the flow path 7 and the nebulizer housing 17. Conversely, during inhalation, i.e., in the first flow direction A, the check valve 15 closes the third opening 14. As a result, when the user inhales, air can enter the flow path 7 almost exclusively through the second opening 9, and when the user exhales, air can exit the flow path 7 almost exclusively through the second opening 9 and the third opening 14.

[0126] To evaluate the quality of inhalation therapy and send signals to control the actuator, it is necessary to observe the inhalation and exhalation process through the inhalation therapy device 1. In this regard, "good therapy" refers to therapy in which a large amount of medication is delivered to the user in a short time. This can be achieved by limiting the inhalation flow rate and extending the inhalation duration while shortening the exhalation duration, without placing an excessive burden on the user's inhalation action. Such a configuration can avoid user discomfort (which may increase the risk of the user interrupting the inhalation therapy).

[0127] To date, preferred embodiments of this disclosure have found solutions that are easy to manufacture and disinfect, and achieve beneficially short total therapeutic times. These solutions allow for extended inhalation time, keep exhalation time as short as possible, and avoid unpleasant resistance during exhalation.

[0128] To achieve an extended suction time, the user must inhale against at least a certain resistance, i.e., a certain flow resistance. With the above configuration, the desired suction flow resistance can be achieved while simultaneously reducing the discharge flow resistance, because the check valve 15 facilitates the discharge of air through the check valve 15 and the third opening 14 of the flow path 7 during discharge.

[0129] In another embodiment, the third opening and the check valve 15 may be omitted. In this embodiment, during intake, air may enter the flow path 7 exclusively or substantially exclusively through the second opening 9, and during discharge, air may be discharged exclusively or substantially exclusively from the flow path 7 through the second opening 9. Thus, in this scenario, the entire intake and discharge flow rate is guided through the second opening 9 of the flow path 7. The increase in intake resistance and / or decrease in discharge resistance in this configuration is achieved using a fixed shape valve that functions as the second opening.

[0130] Next, let's take a closer look at the controller housing 16. The controller housing 16 houses a controller 18 configured to control the actuator 6 of the membrane unit 3.

[0131] The top / upper surface 21 of the handle portion 12 of the controller housing 16 has a rounded or curved shape, which provides good ergonomic operability for the controller housing / handle portion 16. The top / upper surface 21 is located opposite the bottom surface of the controller housing 16, which defines the support surface 19. The top surface 21 defines a ridge line 22 in the side view (see Figure 2). In this context, a ridge line is understood as a line formed along the highest point of the top surface in the side view (longitudinal section view). The angle α6 between the ridge line 22 and the support surface 19 is less than 20° in the side view (longitudinal section view) and is optionally between 5° and 15°.

[0132] At least one control button 23 and / or at least one light-emitting indicator 24 are located on the top surface 21, centered on the ridge line 22. The user can turn the aerosol therapy device ON / OFF using the control button 23. The light-emitting indicator 24 may provide the user with feedback regarding the progress of the treatment.

[0133] In this embodiment, the nebulizer housing 17 can be separated from the controller housing 16. In this configuration, the nebulizer housing 17 has a peripheral rim 43 defining a recess 44. The controller housing 16 has a boss 45, and when the nebulizer housing 16 is assembled to the controller housing 16, the recess 44 and the boss 45 engage.

[0134] Furthermore, the controller housing 16 is equipped with a socket 46. When the nebulizer housing 17 is attached to the controller housing 16, the plug 33 of the membrane unit 3 can be detachably inserted into the socket 46. This electrically connects the membrane unit 3 to the controller 18 via the socket 46.

[0135] As an example, the angle α5 between the support surface 19 and the insertion direction of the plug 33 into the socket 46 is greater than 90°±5°%, preferably between 100° and 160°, and more preferably between 110° and 150°, in the side view (longitudinal cross-sectional view).

[0136] Furthermore, the sensor 47 is housed within the controller housing 16. The sensor 47 is electrically connected to the controller 18. In this example, the sensor 47 is a sensor that detects the difference pressure between the ambient pressure and the pressure in the flow path, here between the second opening 9 and the membrane 5, i.e., the storage chamber 40 in this embodiment. For this purpose, the sensor 47 has a sensor support 48. In this embodiment, the sensor support 48 is located on the bottom wall 51 of the nebulizer housing 17. A tube or hose 49 connects the sensor support 48 and the sensor 47. In this configuration, the ambient pressure is measured from inside the controller housing 16, and since the controller housing 16 is not airtight, its internal pressure is equal to the ambient pressure.

[0137] Furthermore, the controller housing comprises and accommodates one or more batteries 50. One or more batteries may be rechargeable and / or replaceable. One or more batteries are positioned near the support surface 19, and their length extends along the support surface 19. In this context, "along the support surface 19" does not necessarily mean extending parallel to the support surface 19. However, the longitudinal direction of one or more batteries should be within 0° to 20°, optionally between 0° and 15°, with respect to the support surface 19.

[0138] Considering the foregoing, the controller housing 16 may contain the most relevant electronic features necessary to operate the membrane unit 3, preferably all electronic features, and other features of the inhalation therapy device 1, so that the nebulizer housing 17 can be configured as one or more throwaway (disposable) units, which are used for a certain period of time, such as a specific number of therapy sessions, weeks, months, six months, or up to one year. At the same time, the controller housing 16, which has all the relevant electrical and sensing elements to operate the membrane unit 3 and other features of the inhalation therapy device and to monitor the therapy process, can be configured as a long-life unit that does not need to be specially disinfected before each therapy session.

[0139] The illustrated embodiment is an inhalation therapy device 1 in which the user inhales and exhales during therapy. That is, the inhalation therapy device 1 is in contact with the user's mouth, and the user's normal inhalation and exhalation are performed through the inhalation therapy device 1 via the first opening 8. In an alternative embodiment, a mask may also be attached to the first opening 8. In either case, during treatment, the inhalation therapy device 1, and especially its first opening 8, must not be removed from the user's mouth (or face if a mask is used). However, if the patient removes the inhalation therapy device 1, and especially its first opening 8, from the user's mouth (or face if a mask is used) during treatment, the membrane unit 3 is simply deactivated and aerosolization stops. Thus, although unintended, interruption is possible.

[0140] The method for generating aerosols inside the inhalation therapy device 1 and the method for actually performing the therapy are described below. Hereafter, and in the following, it will be understood that Figure 5 shows the orientation of the inhalation therapy device when it is held in the ideal position while the user is performing therapy in a standing or sitting position.

[0141] In this scenario, a reservoir 2 for holding a liquid or fluid (i.e., a drug) is provided in the upper section of the housing body 4, and the central axis CA2 of the filling opening 25 is substantially vertical.

[0142] Furthermore, as can be deduced from Figure 5, the membrane unit 3 is also oriented vertically. Therefore, the central axis CA3 of the membrane 5 is oriented horizontally. Consequently, the liquid in the reservoir 2 is supplied to the membrane 5 by gravity.

[0143] As mentioned above, the first opening 8 is the opening that comes into contact with the airway through the user's mouth in order to enable inhalation and exhalation through the inhalation therapy device during aerosol therapy. When using the device, if the user wishes to begin therapy, the first step is to fully assemble the inhalation therapy device 1, which includes inserting the membrane unit 3 into the nebulizer housing 17, closing the housing 17, and attaching the nebulizer housing 17 to the controller housing 16. Furthermore, the lid 26 must be unscrewed from the reservoir 2, and the liquid to be administered must be poured into the reservoir 2 through the filling opening 25. When the inhalation therapy device 1 is placed with its support surface 19 on a horizontal surface such as a table, the plane of the filling opening 25 is tilted toward the mouthpiece 27 relative to the horizontal plane. However, it is still possible to pour the liquid into the reservoir 2 through the filling opening 25 without the liquid flowing out of the reservoir 2. After that, the lid 26 is screwed back onto the reservoir 2 to seal the liquid. As a final step, the user turns on the power to the inhalation therapy device 1 by pressing the control button 23. At this stage, the user can begin treatment.

[0144] To begin treatment, the user places the mouthpiece 27 in their mouth and starts inhaling and exhaling. In this state, one respiratory cycle consists of only one inhalation phase followed by one exhalation phase.

[0145] Upon inhalation, the user inhales against a first maximum absolute value of flow resistance of approximately 160 Pa ± 100 Pa, optionally 160 Pa ± 60 Pa, and more optionally 160 Pa ± 30 Pa. This resistance value is measured at the peak flow rate of a sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml. In another embodiment, the first maximum absolute value of flow resistance may be between 120 Pa and 500 Pa, optionally between 125 Pa and 400 Pa, more optionally between 130 Pa and 300 Pa, and most optionally between 135 Pa and 180 Pa. In yet another embodiment, the first maximum absolute value of flow resistance during the user's inhalation may be between 60 Pa and 260 Pa, optionally between 100 Pa and 220 Pa, and more optionally between 130 Pa and 190 Pa.

[0146] During the suction process, the suction flow A described above is guided into the flow path 7. In the current embodiment, one pressure port (or connection) of the differential pressure sensor in the controller is fluidically connected to the flow path 7 via the sensor support 48. The other pressure port (or connection) of the differential pressure sensor in the controller is fluidically connected to the environment so that the differential pressure sensor 47 can measure the pressure difference between the environment (ambient pressure) and the inside of the atomizing unit. During suction, the pressure in the flow path becomes negative.

[0147] Based on this output from the sensor 47, the controller 18 determines that the user is inhaling. When it determines that inhalation or exhalation has ended (in the case of active "pre-on"), the controller 18 activates the actuator 6. This causes the membrane 5 to vibrate, generating an aerosol from the liquid in the reservoir 2 at the second side 5.2 of the membrane 5. The generated aerosol is inhaled by the user by the inhalation flow A and transported to a predetermined location in the airway.

[0148] At the end of the initial inhalation phase, the controller 18 determines from the pressure signal in the flow path that the inhalation phase is nearing its end and deactivates the actuator 6 to stop aerosol generation. In certain cases, the controller 18 deactivates the actuator 6 before the end of the inhalation phase, thereby stopping aerosol generation ("pre-off" process). Therefore, at the end of the inhalation phase, the user no longer inhales aerosol. This timing is preferably adapted so that all aerosol reaches the desired location (lungs) and does not remain unused in the trachea or throat (respiratory dead volume) at the start of exhalation, thereby maximizing the dose delivered to the user's lungs. The period during which the actuator 6 is deactivated at the end of the inhalation phase can be between 100 and 1000 milliseconds before the start of the exhalation phase, optionally between 300 and 800 milliseconds, and most optionally between 500 milliseconds, which corresponds to a standard sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml.

[0149] Instead of time, the controller may determine an inhalation target volume (target respiratory dead volume or dead volume) of 150 ml, optionally 130 ml, more optionally 100 ml, or even more optionally 90 ml.

[0150] For smaller users (individuals, patients), such as adolescents or children, smaller target volumes (dead volumes), such as 80 ml, optionally 60 ml, or even more optionally 40 ml, may be used.

[0151] As a further example, as will be discussed later with reference to Figure 9, the controller 18 may determine the “pre-off” timing based on the preceding respiratory cycle of the two respiratory cycles. Figure 9 shows curves (e.g., pressure or flow rate over time) for two consecutive respiratory cycles (n and n+1), each respiratory cycle including inhalation (n, n+1) and exhalation (n, n+1).

[0152] In the first step, the maximum value (of the flow / pressure curve) is determined from the first breathing action (inhalation n), and the total inhaled volume (for example, 400 ml in Figure 9) is calculated by integrating the function. In the second step, a predetermined value is subtracted from this total to determine the "pre-off" value. In this state, the predetermined value may be in the range of 80 ml to 120 ml. In other examples, the predetermined value may be a maximum of 130 ml or 150 ml. The minimum predetermined value may be 60 ml or 70 ml. In the example in Figure 9, the predetermined value is 90 ml.

[0153] These predetermined values ​​may be adjustable via interfaces such as apps on mobile devices and / or wireless data connections to the cloud or databases, and may be available to users, manufacturers, clinical research organizations (CROs), and / or physicians.

[0154] Alternatively, predetermined values ​​may be automatically set from given date of birth and / or lung function data via an interface such as an app on a mobile device and / or a wireless data connection to the cloud or database, and are available to the user, manufacturer, clinical research organization (CRO), and / or physician.

[0155] Subsequently, the “pre-off” value is defined as a value on a function from which the subtraction result is obtained. Therefore, by subtracting the target volume (e.g., dead volume) from the total inhalation volume (e.g., total volume, tidal volume, TV), the “pre-off” value, i.e., the point at which actuator 6 is deactivated during inhalation, thereby stopping aerosol generation, can be determined.

[0156] In the example in Figure 9, the total volume is 400 ml, and the predetermined value is 90 ml. Therefore, the result of the subtraction is 400 ml - 90 ml = 310 ml. The "pre-off" value in the function is determined to be 310 ml (corresponding to the left shaded area in Figure 9).

[0157] In the third step, the controller 18 calculates a percentage of the “pre-off” value relative to predetermined maximum values ​​(e.g., maximum pressure value, “volume gradient,” flow rate over time) in this first breathing action (inhalation n). In this example, the “pre-off” value corresponds to 46% of the maximum flow rate / pressure in this first breathing action (inhalation n).

[0158] This percentage value is used in conjunction with the second maximum value of the second breathing action (inhalation n+1), i.e., in a series of inhalations n+1 (e.g., maximum pressure value, "volume gradient", flow rate over time). In particular, actuator 6 is deactivated when it reaches a percentage value from the second maximum value, e.g., 46%.

[0159] As a result, conclusions about the next measurement are always drawn from the previous measurement (e.g., pressure / volume, or flow rate / volume). A "default" option is also available. For example, a fixed percentage value such as 30% or 40% may be used for the first inhalation immediately after starting inhalation therapy. Alternatively, it is possible to not use a fixed percentage value for the first inhalation immediately after starting inhalation therapy, and instead stop aerosol generation at the end of inhalation by disabling actuator 6 ("pre-off" value = 0%).

[0160] Furthermore, or instead, if an inhalation is deemed invalid (e.g., too short a time, too small a volume), that inhalation is not counted, and the percentage value or default value of the previous "correct" inhalation is used. An inhalation may be considered valid and counted in the calculation of the "pre-off" value if the inhalation time is at least 1 second and / or the inhalation volume is at least 150 ml. Alternatively, an inhalation may be considered valid if the inhalation time is at least 1.4 seconds and / or the inhalation volume is at least 180 ml.

[0161] The pre-off process ensures that all or at least the majority of the aerosol reaches the desired location in the airway, with only a small amount of aerosol being exhaled again during the subsequent exhalation phase. In the subsequent exhalation phase, the user exhales, generating exhalation flow B. This exhalation flow is generated in the flow path against a maximum absolute value of a second flow resistance in the range of 120 Pa ± 75 Pa, preferably 120 Pa ± 45 Pa, more preferably 120 Pa ± 22.5 Pa, which is measured by the peak flow rate of a sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml.

[0162] When exhalation begins, the pressure in the flow path 7 changes to overpressure again, and the controller 18 determines the user's exhalation based on the output of the sensor 47. At the end of the exhalation phase, the overpressure in the flow path 7 decreases and approaches the ambient pressure, which causes the controller 18 to determine that the exhalation phase is nearing its end. In certain cases, the controller 18 activates the actuator 6 before the end of the exhalation phase, thereby initiating aerosol generation (pre-on-process). The duration for which the actuator is operated during the exhalation phase can be between 50ms and 500ms, optionally between 100ms and 300ms, and most optionally between 200ms, in relation to a standard sinusoidal breathing pattern with an inhalation time of 2 seconds, an exhalation time of 2 seconds, and a tidal volume of 500 ml. Alternatively, the duration for which the actuator is operated during the exhalation phase can be up to 500ms, up to 300ms, up to 200ms, or up to 100ms.

[0163] Aerosol generation begins during the discharge phase while the discharge flow B is still present, so a portion of the generated aerosol is transported downstream of the membrane 5 towards the second opening 9 and an optional third opening 14. A storage chamber 40 is provided to prevent the aerosol generated during this period from flowing out of the inhalation therapy device 1, i.e., through the second opening 9 and the optional third opening 14. In other words, the generated aerosol flows into the storage chamber 40 during this period, temporarily storing the aerosol as a large mass until the next continuous inhalation phase begins.

[0164] Subsequently, the user initiates the next breathing cycle in the next inhalation phase. At the start of inhalation, the aerosol temporarily stored in the storage chamber 40 is again entrained by the inhalation flow A, passes through the membrane unit 3, and is inhaled through the first opening 8. When the next inhalation phase begins, the pressure in the storage chamber 40 changes again, which allows the controller 18 to determine the next inhalation phase. Therefore, the controller 18 activates the actuator 6 again and starts aerosol generation. At this stage, the above process is repeated.

[0165] Furthermore, aerosol generation may be continuous between the "pre-on" and "pre-off" (operation phases) of the process. However, the power (power adaptation), frequency (frequency adaptation), etc., supplied to the actuator 6 may change between consecutive operation phases or even within the same operation phase.

[0166] When there is no liquid remaining in reservoir 2, or when the amount of residual liquid in the reservoir reaches the minimum amount required for aerosol generation, administration is terminated and the process (therapy) is stopped. The end of administration can be notified to the user visually using light incorporated in the light indicator 24 or control button 23, and / or audibly by outputting an audio signal. In this regard, the end of administration is detectable as described in European Patent No. 1 558 315 B1 (U.S. Patent No. 7,458,372 B2) or European Patent No. 3 082 918 B1 (U.S. Patent No. 10,744,277 B2).

[0167] Finally, the user can use the control button 23 to switch the inhalation therapy device off again and end the treatment. Alternatively, the device can be set to automatically switch off after administration is complete and a certain period of time has elapsed.

Claims

1. Inhalation therapy device, Housing body (4) and A reservoir (2) for holding liquid inside the housing body (4), A membrane (5) disposed within the housing body (4) and having a plurality of apertures (110), wherein the liquid can be supplied to the first side (5.1) of the membrane (5) by gravity, An actuator (6) coupled to the membrane (5) for vibrating the membrane (5), wherein the liquid passes through the aperture (110) and an aerosol is generated on the second side (5.2) of the membrane (5) opposite the first side (5.1), A flow path (7) defined within the housing body (4), having a first opening (8) at one end leading to the outside of the housing body (4), and having at least one second opening (9) at the other end leading to the outside of the housing body (4), wherein when a user inhales through the first opening (8) to deliver the generated aerosol, a flow can be generated in the flow path (7) in a first flow direction (A) from the second opening (9) to the first opening (8), and / or when a user exhales through the first opening (8), a flow can be generated in the flow path (7) in a second flow direction (B) from the first opening (8) to the second opening (9), A sensor (47) configured to detect flow parameters of the flow within the flow path (7), The system includes a controller (18) which determines at least the user's inhalation phase and / or exhalation phase based on the output of the sensor (47), and which is configured to repeatedly operate the actuator (6) during the inhalation phase and not operate the actuator (6) during the exhalation phase during a single treatment time, The aerosol output rate of the inhalation therapy device at the first opening (8) during the single treatment time is determined when using a test solution containing albuterol at a concentration of 0.1% (M / V) in a 0.9% sodium chloride solution as the liquid. a) The aerosol droplet has a median mass between 4.5 μm and 5.0 μm, and the amount is at least 0.48 g / min, optionally at least 0.64 g / min, and more optionally at least 0.8 g / min, or b) The aerosol droplet has a median mass between 4.0 μm and 4.5 μm, and the amount is at least 0.42 g / min, optionally at least 0.54 g / min, and more optionally at least 0.70 g / min, or c) The mass median diameter of the aerosol droplet is between 3.5 μm and 4.0 μm, and the amount is at least 0.29 g / min, optionally at least 0.38 g / min, and more optionally at least 0.51 g / min, or d) The aerosol droplet has a median mass between 3.0 μm and 3.5 μm, and the saturation is at least 0.20 g / min, optionally at least 0.29 g / min, and more optionally at least 0.38 g / min, or e) The aerosol droplet has a median mass between 2.5 μm and 3.0 μm, and the amount is at least 0.13 g / min, optionally at least 0.19 g / min, and more optionally at least 0.26 g / min. Inhalation therapy device.

2. The inhalation therapy device according to claim 1, wherein the controller (18) is configured to start operating the actuator (6) within the exhalation phase and before the start of the inhalation phase in each respiratory cycle which includes one inhalation phase and one exhalation phase.

3. The inhalation therapy device according to claim 1 or 2, further comprising a storage chamber (40) formed between the membrane (5) and the second opening.

4. The inhalation therapy device according to claim 3, wherein the volume of the storage chamber (40) is between 10 ml and 100 ml, optionally between 20 ml and 80 ml, and more optionally between 25 ml and 75 ml.

5. The inhalation therapy device according to any one of claims 1 to 4, wherein the sensor (47) is configured to sense the flow parameter in the membrane (5) or between the membrane (5) and the second opening.

6. The inhalation therapy device according to any one of claims 1 to 5, wherein the membrane (5) is disposed within the flow path (7) and an envelope flow is formed around at least a portion of the membrane (5).

7. The inhalation therapy device according to any one of claims 1 to 6, wherein the controller (18) is configured to stop operating the actuator (6) within the inhalation phase and before the start of the exhalation phase in each respiratory cycle comprising one inhalation phase and one exhalation phase.

8. The inhalation therapy device according to any one of claims 1 to 7, further comprising a mixing chamber (42) formed between the membrane (5) and the first opening.

9. The inhalation therapy device according to claim 7, wherein the volume of the mixing chamber (42) is between 5 ml and 50 ml, optionally between 10 ml and 40 ml, and more optionally between 12.5 ml and 37.5 ml.

10. The inhalation therapy device according to claim 8 or 9, wherein the mixing chamber (42) tapers toward the first opening.

11. The inhalation therapy device according to any one of claims 1 to 10, wherein the flow parameter is pressure, optionally a differential pressure, and more optionally a differential pressure between the pressure in the flow path (7) and the ambient pressure.

12. The inhalation therapy device according to any one of claims 1 to 11, wherein the flow path (7) is configured to provide a first flow resistance when the user inhales and a second flow resistance lower than the first flow resistance when the user exhales, and optionally the second flow resistance when the user exhales is at least 25%, at least 30%, or at least 40% lower than the first flow resistance when the user inhales.

13. The inhalation therapy device according to claim 1, wherein the first flow resistance is 1.25 to 6 times, optionally 1.25 to 3 times, and more optionally 1.25 to 2.5 times greater than the second flow resistance.

14. The aperture (110) penetrates the membrane (5) in the extension direction from the first side (5.1) to the second side (5.2), and each aperture has a nozzle portion (112) on the second side (5.2) along the extension direction (E) of the aperture, and the length (L) of the nozzle portion (122) Ａ The inhalation therapy device according to any one of claims 1 to 13, wherein the particle size is between 5 μm and 16 μm, optionally between 5 μm and 14 μm, and more optionally between 5 μm and 12 μm.

15. The inhalation therapy device according to claim 14, wherein the aperture (110) is formed by at least two laser perforations, one of which forms the nozzle portion (112) and the other forming the remainder (114, 116) of the aperture (110).

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