Patent Application
20240399079
Therapeutic aerosol inhalations are deposited in the respiratory tract to treat respiratory diseases and syndromes. More widespread usage of aerosol delivery devices has led to marked improvement in respiratory function and mortality, especially those devices capable of safely delivering small molecules and micrograms of the therapeutics. However, there is an unmet need to more effectively treat those patients with more severe respiratory distress due to multiple etiologies. While respiratory distress impacts people of all ages, the young and elderly are among those most vulnerable to such ailments. In addition to high cost of health care, these demographics of patients are associated with high mortality rates.
Globally, over 120 million episodes of pneumonia in children under the age of 5 are reported annually, 14 million of which progress to severe episodes, and 888,000 result in death. Most of these patients are less than 2 years of age. Other illnesses such as Idiopathic Pulmonary Fibrosis (IPF), TB, COVID-19 and Chronic Obstructive Pulmonary Disease (COPD) also require effective deposition of the minimum concentrations of antivirus vaccines or anti-infectives throughout the lungs for positive outcomes. Moreover, life-saving rescue therapy for hypoxemia patients requires noninvasive treatments with pulmonary surfactant aerosols having life-saving biophysical properties.
Many potential therapeutic treatments with aerosols fail due to inadequate deposition of the active agent into the distal lungs to resolve physiological dysfunction or disease malady. The death rate of 40% of patients suffering from respiratory distress, which includes neonates, infants, children and adults, is unacceptable. Delivery of insufficient masses of agents to the lungs using present inadequate aerosol devices cannot be expected to achieve a therapeutic benefit, and therefore should not be used.
Atomizers generate aqueous aerosols of ˜3 μm diameter at output flows of 0.3 ml/minute using Venturi liquid flow. Compressed central axial air flow creates a negative Venturi liquid flow to the aerosolizing nozzle where the plume is generated by Raleigh breakup phenomena. Larger droplets are recirculated. Output aerosol flow in these systems is typically 0.3 ml/minute with 12 L/minute gas flow.
Mesh nebulizers include a vibration mesh typically having one thousand holes, each 3 micron in size, and vibrate at 128,000 cycles per second. Aerosol generation in these mesh nebulizers is severely limited by the drug liquid viscosity. Generally, viscosities greater than 6 cP markedly reduce aerosol generation or even prevent it altogether. Moreover, the mesh is susceptible to cloggage by liquid suspensions and large molecular entities. The output rates in these systems are limited to less than 1 ml/min. These devices have been incorporated with separate ventilators to provide aerosol therapy in conjunction with respiratory support.
Windtree's capillary aerosol generation system incorporates Sinclair-Lamer aerosol generation. Evaporation of liquids results in aerosol nuclei. Condensation of the vapor on the nuclei results in uniform aerosols. This method can generate surfactant aerosols at 3 l/min through a delivery tube through the nose of a neonate. Losses in the aerosol generation system and tubing, combined with minute volumes of 20-40 ml/breath, results in delivery of only approximately 1% of the initial volume of the surfactant liquid loaded to the drug vial for delivery.
Trudell's Solarys nebulizer uses compressed air pressure. Using this device, aerosolization occurs external to the liquid feed tip and its surrounding compressed gas ports. Generally, larger aerosol diameters and broader sprays are generated with large particle size dispersion. However, high exit gas port pressures and aerosol velocities can cause injury and local impaction in the oral cavity or conducting airways in the lungs.
There is an unmet need to generate aerosols with a range of effective therapeutic moiety agents, from both low to high concentrations of drugs, over a range of low and high molecular weights, structural configurations, as well as low and high viscosities. Aerosol drugs delivered using nebulizers include broncho dilators, antibiotics, mucolytics and future biologics that include pulmonary surfactant, proteins, oligomers antibodies nano suspensions.
Generation of fine atomized aqueous aerosols suitable for penetration through the upper respiratory airways and deposition in the peripheral lungs of neonates, infants and children with low tidal and minute volumes is a challenge due to the high gas volumes and high jet velocities required, as predicted by the Weber number. Thus, pediatric patients typically inhale only a small volume of aerosol. These limitations have resulted in a preference to use mesh nebulizers as a medication delivery system, despite the aforementioned disadvantages and limitations.
It is therefore an object of the invention to provide an aerosol delivery system and method that overcomes the aforementioned shortcomings in the prior art. This object is achieved as by the various embodiments of this disclosure (“AeroPulsR”), which overcomes these challenges by use of its inventive liquid aerosol dispensing system to provide suitable inhalation therapeutics for treatment of a wider range of illnesses and patients.
Efficiently generating and delivering fine aqueous aerosols using pressure driven nozzles, for instance as described in the U.S. Pat. No. 10,661,033, herewith incorporated by reference, during inspiration in children requires that small volumes of the plumes from the nozzle are either delivered directly and/or reduced in velocity using a mini aerosol concentrator as described in this specification.
To achieve this capability, aerosol generation must be able to be started and stopped within milliseconds. The fine particle generation by the aforementioned nozzle utilizes pressurized drug containing liquids along with pressurized gas such as air, oxygen or heliox. Previously, a high-pressure syringe drive was used to deliver the liquid to the nozzle. This syringe was considered to be too specialized and cumbersome to be used by untrained operators, and leakage and misuse routinely resulted in liquid spillage. Additionally, the syringe drive was unable to respond to the rapid changes of liquid flow to the nozzle necessary to meet the rapid temporal precise dosimetry requirements during inspiratory breaths in pediatric patients. The required rapid precise control of aerosol delivery during the inspiratory phase, especially for neonates and infants.
In one general aspect, the aerosol concentrator may include an input tube including: a lumen; an input tube entrance port having an input tube entrance port cross-sectional shape differing from a cross-slit shape; and an input tube exit port having a cross-slit shape, where the input tube entrance port cross-sectional shape transitions between the input tube entrance port and the input tube exit port to the cross-slit-shaped exit port cross-sectional shape. The aerosol concentrator may also include an output tube including: a lumen; an output tube entrance port having a cross-slit-shaped cross-section, said output tube entrance port being aligned with and spaced apart from the input tube exit port by a gap between the input tube exit port and the output tube entrance port; and an output tube exit port where the output tube entrance port cross-sectional shape transitions between the output tube entrance port and the output tube exit port to an output tube exit port cross-sectional shape. The concentrator may furthermore include a housing encompassing a plenum, said plenum encompassing the gap between the input tube and the output tube, where the plenum is connected to an exhaust port through which an exhaust aerosol exits the plenum.
The input tube of the aerosol concentrator may converge from its input tube entrance port to its input tube exit port from an input tube entrance port cross-sectional area that is larger than an input tube exit port cross-sectional area.
The output tube of the aerosol concentrator may diverge from its output tube entrance port to its output tube exit port from an output tube entrance port cross-sectional area that is smaller than an output tube exit port cross-sectional area.
The input tube and the output tube of the aerosol concentrator may have a joint longitudinal axis and have radial sizes transverse to the longitudinal axis that are less than a third of a radial size of the plenum transverse to said longitudinal axis.
The housing of the aerosol concentrator may compromise an input housing part holding the input tube, an output housing part holding the output tube and an exhaust tube, and a seal is provided between the input housing part and the output housing part.
The housing of the aerosol concentrator may share the same joint longitudinal axis with the input tube and the output tube, and the plenum may be encompassed at least for the most part by either one of the input housing part or the output housing part, while the other part of the output or input housing, respectively, is essentially disc-shaped, with the housing part encompassing at least for the most part the plenum also having the exhaust tube which extends essentially radially in a transverse direction with respect to said longitudinal axis.
The aerosol concentrator may further comprise a pressure relief valve connected to the exhaust tube, configured to keep the pressure within and across concentrator and consequently at the output tube exit port essentially constant.
The aerosol concentrator may be configured to concentrate an input aerosol having fine particles of a particle size distribution of 1-6 μm MMAD suspended in gas.
The aerosol concentrator may be configured to concentrate an aerosol using virtual impaction at a small positive pressure within and across concentrator all the way to the output tube exit port such that the concentrated respirable aerosol can be delivered to the patient at the small positive pressure without the use of pumps to remove the exhausted gas.
The present disclosure involves an aerosol generation system together with adjoining components to be used by a clinician for providing aerosol inhalation therapy with concurrent respiratory support. A housing may encompass several system parts including electronic components, regulators, valves, a liquid dispenser system, and an aerosol nozzle chamber. The system is controlled by a console, which includes a user interface, a microcontroller, a hardware controller, and internal hardware. The microcontroller housed in the console is responsible for communicating with the user and initiating preset commands that control the hardware controller. The hardware controller comprises a set of solenoid valves that precisely control the flow of regulated pressurized gas such as air, oxygen and/or heliox. The pressurized gas is regulated by four pressure regulators. Three regulators supply gas that control a variety of system functions, including: (i) the flow of hydraulic driven liquid from a container to a nozzle via a capillary, (ii) the liquid flow within an aerosol space, and (iii) the gas pressure and velocity for shearing the liquid to form an aerosol. The liquid drug dispenser is pressurized and is configured to deliver liquid to the nozzle. A counterflow regulator may be configured to regulate gas counterflow that decelerates the aerosol plume from the nozzle. The aerosol chamber may be configured to process the aerosol, and may further include a detachable cone which converges the generated aerosol downstream towards an aerosol delivery outlet. The delivery outlet may be configured to include a desired level of respiratory support pressure. The processed aerosol comprises fine aqueous particles effective for respiratory treatment.
A control module, also known as electronics control and interface module (ECIM), comprises components configured to control the system. The microcontroller activates a series of fast switching solenoid valves, which in turn control the onset and offset sequence of the gas flows that operate the AeroPulsR system. The internal hardware may comprise four precision regulators (Bellofarm 960-007-009 modified with bleed 803-000-001 to reduce the heliox use). A primary regulator R regulates the input pressure P for the system. The primary regulator R regulates pressurized gas from a high-pressure source and maintains the gas pressure P at a constant level, typically between 40 and 70 psi. Three additional regulators—R1, R2 and R3—regulate gas pressure P1, P2 and P3, respectively. R1/P1 correspond to the liquid drug dispenser, R2/P2 correspond to the aerosolizing nozzle, and R3/P3 correspond to the counterflow tube. Activation of these pressures is controlled by rapid action solenoid valves. The output pressures of each of the precision regulators are preset prior to delivering the aerosol drug/agent treatment to the patient. The gas flow is metered by the regulators and modulated by the solenoid valves to control a plurality of system components and parameters, including: the pressure of the liquid in the vial, the rate of flow of the liquid to the nozzle, the cessation of aerosol generation, the aerosol arresting counter-flow gas flow, and the bidirectional valve that regulates the inspiration of aerosol and expiration of the breath. The control module may have tactile, visual, and/or audible input and output elements for controlling the system and administering the therapy. A green button may be included on the electronics control and interface module (ECIM) for activating the microcontroller that controls the timing and actuation of solenoid valves. These solenoid valves enable precise aerosol generation and delivery as well as the timing of the respiration ventilation. Opening and closing of specific solenoid valve may be achieved by an Arduino microcontroller. Additionally, the control module may include a silver button for switching the device ON or OFF. The ECIM may include a simplistic layout having three LEDs and a built-in speaker, which work together to instruct the user on inhalation and exhalation timing. The prescribed treatment duration for a given patient may be pre-loaded to the ECIM in the form of software code. A therapy-specific software program may be uploaded before the treatment begins, or alternatively a pre-set therapy program may be accessed.
The pneumatic and hydraulic controls of the aerosol generation and delivery system involve a network of regulators and valves. The gas pressure of regulator R (which may be set at 60 psi), sets a maximum pressure for the entire system, followed downstream by a junction of a plurality of parallel conduits feeding the pressure from the pressure regulator R into several pressure regulators R1, R2 and R3, each of which feeding into respective downstream pressure control valves V1, V2 and V3 for fine-tuning the pressure control to fall within narrow margins of tolerance. While the 2-way control valve V1 downstream of the pressure regulator R1 controls the pressure of the liquid to be aerosolized, a 2-way control valve V2 downstream of the pressure regulator R2 controls the pressure of the gas aerosolizing the liquid and consequently the pressure in an aerosolizing space surrounding a liquid exit orifice where the liquid enters the aerosolizing space. Raising the gas pressure in the aerosolizing space by the pressure control valve V2 above the liquid pressure controlled by the pressure control valve V1 results in shutting off the liquid flow through the liquid exit orifice into the aerosolizing space, which prevents liquid aerosolization. Conversely, lowering the liquid pressure at the liquid exit orifice below the gas pressure in the surrounding aerosolizing space likewise shuts off liquid flow into the aerosolizing space. Dropping the liquid pressure can be accomplished by the pressure control valve V1 and/or the pressure regulator R1. For example, setting the pressure at the pressure regulator R1 to zero shuts down the liquid flow. However, during operation, a fine-tuning of the liquid pressure is necessary, which is accomplished by the pressure control valve V1.
During operation, it is more favorable to keep the liquid pressure at a constant value within narrow tolerances, and to modulate the gas pressure in the aerosolizing space via the pressure control valve V2. Such control can accomplish a pulsed aerosol generation, e.g. in pulses a short as 10 ms. Keeping the gas pressure in the aerosolizing space constant but modulating the liquid pressure is also possible, but tends to be a slower. Also, both the liquid pressure and the aerosolizing gas pressure may be modulated simultaneously. From a practical standpoint, keeping the liquid pressure constant within narrow margins and modulating the aerosolizing gas pressure is the preferred solution, both as to the complexity of the control, as well as the response time to a modulating signal controlling the pressure control valve V2.
For controlling the liquid pressure, the 2-way valve V1 is activated so that a gas pressure in a vial holding the liquid to be aerosolized is set, the gas pressure forcing the liquid out of the vial and into the aerosolizing nozzle where it flows through a capillary inside the nozzle that terminates at the aforementioned liquid exit orifice.
Finally, a third pressure can be controlled by the control valve V3 for a gas flow into a counterflow tube ejecting a counterflow gas, e.g. air, for “arresting” an aerosol plume that enters from the aerosolizing nozzle into an aerosol chamber. Controlling this pressure simultaneously controls the volume flow of gas exiting the counterflow tube, which in turn controls the gas flow speed of the gas exiting the counterflow tube. This does not only control the “arresting” action distributing the aerosol plume in the aerosol chamber, but also determines an amount of dilution gas diluting the aerosol in the aerosol chamber, which may be beneficial for avoiding liquid aerosol particle deposition within the aerosol chamber.
Further, a 5-2-way valve V4 is provided proximal to a patient for controlling the aerosol flow. During inspiration, the 5-2-way valve V4 opens an aerosol delivery channel to allow aerosol to flow to the patient. In contrast, during exhalation the 5-2-way valve V4 shuts off or at least reduces the aerosol flow to the patient, while simultaneously opening an aspiration port venting to outside of the system. It is also possible to synchronize the actuation of control valve V4 which is controlled in synchrony with the inhaling/exhaling cycles of the patient with the pressure modulation by the valve V2 controlling the aerosol generation, and optionally along with that to also synchronized the gas flow through the counterflow tube controlled by the pressure control valve V3. The aerosol generation duration can be controlled within wide ranges, ranging from pulses as short as 10 ms to a quasi-continuous aerosol generation over several minutes. During the inhalation cycle, a plurality of pulses can be delivered to the patient, while aerosol generation may be shut off completely during the exhalation cycled. It is also possible to make the aerosol generation continuous over the entire inhalation cycle, and stop the aerosol generation completely during the exhalation phase, so that the pulse length of aerosol generation spans over the entire inhalation phase.
The system may include a liquid dispenser system that enables the drug to be aerosolized to be easily loaded onto the AeroPulsR console. Moreover, the pneumatic liquid dispenser is designed to enable precise control of the aerosol delivery during the inspiratory cycle, especially for treatment of neonates and infants. The dispenser facilitates the delivery of surfactant or other viscous or non-viscous agents using the AeroPulsR system, and allows liquids of known viscosity to be precisely metered to the aerosolizing nozzle.
Compressed gas (air, heliox or oxygen) is used to pressurize the dispenser system for transporting the liquid drug into the nozzle for aerosolization. A dedicated solenoid valve precisely synchronizes the hydraulic liquid flow for aerosolization with each inspiratory breath. No mechanical liquid pumps or electronic flow devices are required. In addition, handling of the liquid dispenser is configured to be easy for clinicians and patients to use.
The liquid dispenser system may comprise a sliding mount, a cap and drug-containing vial. The sliding mount may be fixed to the AeroPulsR housing. The cap may be attached to the top of the vial, which may contain a liquid drug agent. The vial (which may for instance have a volume of 15 ml or 50 ml) may be secured to the cap with a one-handed 180° twist, simultaneously achieving easy vial loading and a leak proof seal. A capillary tube for transporting the liquid drug may extend to the bottom of the vial. The drug agent may be added to the vial as needed without changing the capillary or the nozzle. After the vial is attached to the cap, it may easily be inserted into the sliding mount affixed to the housing. This insertion may be performed with one hand, in a manner similar to filling an atomizer. The sliding mount may include two ports, namely an inlet for pressurized gas and an outlet with a capillary attachment for transporting the liquid agent to the nozzle. The cap may include corresponding gas inlet and liquid outlet ports, which mate with the mount holes via face seals. A first face seal may be included on the vial cap to seal the liquid capillary with the capillary on the cap-mount. A second face seal may be included to seal the gas pressure port. The system may also include a fluidic check valve disposed between the capillary and nozzle barrel, and may be configured to deliver the liquid in the vial to the nozzle and prevent back-flow of liquid into the vial. In a preferred embodiment, the total volume of the liquid inside the capillary and the nozzle barrel is less than 0.5 ml. The capillary feeds the liquid through the capillary to the barrel knob and through the axial channel to the liquid jet orifice through which the liquid enters the aerosolizing space. The flow rate of the viscous liquid in the vial to the nozzle is dependent on four parameters: 1) Pressure difference between the vial and nozzle chamber, 2) capillary diameter, 3) capillary length, and 4) liquid viscosity. Thus, the equation for the liquid flow rate (Q) in a capillary based on the parameters of the pressure difference (ΔP) across of the capillary, liquid viscosity (μ), inside diameter (D), and length (L), and is given by:
A capillary with larger diameter may be selected for liquid agents with higher viscosities, which enables adequate liquid flow generation with a pressure differential of only 0.01 psi. This technique results in lower pressure losses in the system.
Precision regulator R provides the highest pressure to be used by the other regulators R1, R2, and R3. Precision regulator R is set at a constant pressure, for instance between 40 psi and 60 psi. This compressed gas output is connected to the inputs of R1, R2, and R3. A venting 2-way solenoid valve V1 is placed in line between R1 and the vial, and passes the gas input through a cap on the vial to pressurize the liquid in the vial. When this 2-way venting valve V1 is ON, no gas pressure is provided to the vial. When valve V1 is OFF, the regulated output pressure gas from R1 pressurizes the liquid vial and maintains a precise pressure, typically at 50 psi. A capillary tube is included with an inlet located at the bottom of the vial. The liquid from the vial passes through the capillary through an inline check valve to the nozzle.
It is noted that the liquid flow is independent of the vessel volume or its liquid contents, indicating a broader use for this liquid delivery system. The liquid dispenser system precisely meters the delivery of liquid solutions and suspensions of low or high molecular moieties with viscosities including, but not limited to, 1 to 60 cP. Notably, this novel pneumatic system is readily manufacturable, cost effective and the vial is disposable.
The system includes a unique proprietary drip-less, clog-free nozzle that allows low and high viscosity liquids to be converted into fine aqueous aerosols. This nozzle resolving current shortcomings of other nozzles by enabling generation of aqueous viscous liquids having viscosity between 1 and 60 cP. Moreover, with the present system, liquid aerosols are generated 10 times faster (0.4 ml/min up to 4 ml/min) than with current atomizers and mesh nebulizers. The system is tuned to deliver respiratory aerosols with narrow size distributions between 1-6 μm in diameter. The liquid in the vial is maintained at a constant pressure and enters through a liquid microchannel extending axially through the center of the nozzle barrel. The axial liquid microchannel is surrounded by gas microchannels. In a preferred embodiment, four radially positioned air microchannels are included within the nozzle to allow for high pressure gas to be transported into the nozzle. Either air, oxygen or heliox enters the four ports on the nozzle barrel. The input ports for the gas microchannels may be made at an acute angle of 35 degrees to assist in smooth gas inflow and reduce flow resistance. The output of the liquid nozzle and the outputs of the gas microchannels terminate at a small aerosolization space within the nozzle. Typically, the gas pressure is maintained between 30 psi and 70 psi inside the aerosolization space. Aerosolization occurs within the aerosolization space when the gas and liquid converge. The aerosol exits the nozzle through an exit orifice, and is expelled as a columnar plume within a sheath of gas essentially devoid of aerosol. The resulting generated aerosol plume forms a central columnar aerosol plume that is surrounded by a sheath of gas that prevents interaction of the aerosol with the output orifice. This results in a clean, drip and clog free nozzle resistant to exit orifice wear.
The cross section of the nozzle barrel and nozzle aerosolization space is shown U.S. Pat. Nos. 8,820,662 B2; 9,573,147 B2; and 10,661,033 B2, and Japanese Patent Publication JP 6743280 B2, Chinese Patent Publication CN 109562237 B, and Australian Patent Publication AU 2016402362 B2, herewith incorporated by reference. It is advantageous to have a greater number of gas channels, large gas channel diameters, and short channel lengths, as these parameters increase the gas flow rate and reduces the gas flow resistance. Collectively these conditions reduce the loss of gas pressure in the nozzle barrel and thus the aerosolization space experiences higher pressure compared to the previous nozzle, markedly improving the performance of the revised nozzle, thereby enabling the aerosol to be rapidly switched on or off.
The nozzle end contains a minute aerosolizing space where the liquid is aerosolized and transported through an exit orifice of the aerosol chamber. The present disclosure describes a novel approach for controlling the liquid flow to be aerosolized within the aerosolization space. The differential pressure between the gas in the aerosolizing space and the liquid vial (ΔP) controls the liquid flow rate, and thus the aerosolization rate. The pressure in the aerosolizing space is controlled by a dedicated fast switching solenoid valve. When the gas pressure is less than the liquid pressure in the aerosolization space, the liquid is aerosolized. Conversely, when the gas pressure is higher than the liquid pressure in the aerosolizing space, the liquid flow is stopped and no aerosol is generated. More specifically, liquid flows from the liquid vial to the nozzle when the pressure in the aerosolization space is less than the liquid pressure in the vial. When aerosol generation should be stopped, e.g. during a time period between aerosol generation pulses, valve V2 interchanges P2 to pressurized gas P, thereby increasing the pressure inside the aerosolization space.
Liquid flow can be stopped by increasing the pressure of the gas surrounding the liquid jet within the aerosolizing space down-stream from the aerosol exit orifice. As described by Venturi, an increase in gas pressure increases its velocity which decreases the pressure of the liquid jet. The compressed gas within this nozzle chamber causes an extremely rapid acceleration of the liquid jet as the diameter doubles from 0.01 inch to 0.02 inch at the aerosol exit of nozzle, thereby allowing shearing and expansion of the liquid such that aerosol is generated within this space. However, the flow resistance offered by the narrow orifice in the nozzle cap increases the gas pressure within the aerosolization space. The overall result is such that despite the decrease in Venturi pressure, the summation of the pressures on the gas acting on the liquid jet arrests the liquid flow within the liquid jet nozzle.
A one-way check valve may be attached to the capillary in the dispenser cap to prevent back-flow of liquid in the nozzle and to allow liquid flow towards the nozzle cap orifice. The pressure differential between the gas and liquid is precisely controlled within 0.01 psi, and preferably within less than 0.001 psi. When aerosol generation is desired, valve V2 facilitates an interchange of pressure between P2 and P, as discussed. When the gas pressure in the aerosolization space is reduced below the liquid pressure, liquid flows into the aerosolization space and fine aerosols are generated. Precise control of gas flows and liquid flows liquid flow rate through the nozzle to be controlled between 0.5 ml/min and 4.5 ml/min. This is achieved by reducing the pressure within the aerosol space according to the viscosity of the liquid, with differential pressures constraints between 0 and 5 psi. High viscosity liquids require a higher pressure differential and/or a larger internal diameter capillary.
The gas pressure interacts with the liquid from the liquid nozzle to aerosolize the liquid into a fine plume. When the gas pressure regulated by regulator R2 is lower than the liquid pressure, the flow is proportional to the pressure differential between the liquid and the gas. When the gas pressure regulated by system regulator R is larger than 0.5 psi, the liquid flow is arrested and the aerosol generation stops. This sequence is repeatable as needed. The pressure in the vial is vented to the atmosphere via the vent valve V1, and closing vent valve V1 prevents the pressure in the vial from increasing. This stops further liquid entering the nozzle and prevents dripping from the nozzle.
The velocity of the aerosol plume generated by the nozzle is reduced as the aerosol passes through a chamber. The chamber contains an axial counterflow tube through which a precisely regulated gas flow opposes the velocity of the aerosol emanating from the nozzle. The control of this counter-flow gas finely controls the arrest of the nozzle exit plume, as well as stops gas flow during expiration, thereby minimizing the use of heliox during expiration. This virtual baffling phenomenon reduces the aerosol velocity and results in a uniformly distributed aerosol plume, and enables an aerosol output greater than 70% of the aerosol generated by the nozzle. Absence of a counterflow gas results in aerosol losses due to high velocity deposition on the sidewalls of converging cone. In one embodiment, the nozzle and counterflow tube have a common port and receive gas meted from valve V2. In another embodiment, the counterflow tube has an independent gas source, and receives gas at pressure P3 regulated by regulator R3. The gas flow from the counterflow tube is switched on or off independently using a dedicated fast switching solenoid valve. Typically, the onset of the counterflow is activated prior to activating the aerosol generation through the nozzle and deactivated following generation of aerosol from the nozzle More specifically, a 3-2-way solenoid valve V3 interchanges the gas flow on or off 50 ms before and after the aerosol is switched on or off, respectively. This procedure is performed to conserve the amount of gas required for aerosolization and to control the total output flow rate from the system. In one embodiment of the invention, the counterflow tube control and gas supply can be made independent by having a common gas port for both the nozzle and the counterflow tube. This further simplifies the pneumatic liquid aerosol system by eliminating R3 and V3.
The slowed aerosol generated within the chamber may then be focused using a delivery cone, which may be fitted to the chamber using a lip seal or other sealing mechanism that mates with the chamber exterior wall of the chamber. The length of the cone and internal shape are optimized to minimize wall deposition of the aerosol and to provide a soft aerosol exiting the output cone. The cone output port tip is designed to connect with other AeroPulsR related complementary aerosol delivery devices.
The combination of the aerosol chamber and delivery cone forms the shape of a pod, which rests on a pedestal. Aerosol that is deposited on the sidewalls of the chamber and cone may drain through a drain hole into a condensation collection well built into a pedestal. The combination of the chamber and delivery cone is able to house the nozzle, counterflow tube and the aerosol before delivery in a compact manner with minimum losses of aerosol. The chamber may include two ribs that match with two matching notches on the pedestal to allow for easy installation and removal. This allows the user to disassemble the elements within the chamber and cone for cleaning and maintenance.
In another embodiment, the AeroPulsR system may be configured to deliver aerosols on-demand, for which several of the system's capabilities are not required. In this embodiment, the vial may be charged with a prescribed dose volume and mounted onto the console. Similar to generic atomizers, the system may include a wye piece comprising a mouthpiece and exhalation valve along with an exhaust filter. The aerosol generation may for instance be set at a dose rate between 0.1 and 5 ml/minute, i.e. up to 10 times greater than alternative aerosol generation delivery devices. A dose rate of 0.3 ml/min is typical for atomizers used to output aerosol diameters of 3 μm median mass aerodynamic diameter (MMAD), whereas the AeroPulsR system allows 3 ml/min or more to be inhaled through the mouthpiece. Upon activating the AeroPulsR system, the patient breathes until the dose of the liquid in the vial is depleted. This allows the medicine delivery time to be reduced to 1/10th of the delivery time of a generic atomizer.
In a preferred embodiment, the AeroPulsR system may be configured to deliver aerosol with a pressure assist. This enables expansion of the patient's lung and airways volumes to improve aerosol drug penetration and deposition. In one embodiment, a bidirectional valve may be attached directly to a cone. In another embodiment, the bidirectional valve may be connected to the cone via a wye, wherein the wye may include a filter and a pressure release valve. The bidirectional valve assembly comprises an aerosol inlet, an aerosol outlet, and an aspiration port. During inhalation, the bidirectional valve allows the aerosol to enter through the inlet and pass through an aerosol delivery channel port to the patient, while the bidirectional valve concurrently occludes the aspiration port. The patient's inspiratory volume is governed by the gas pressure of the aerosol in the chamber and cone. During exhalation, the bidirectional valve opens the aspiration port, while occluding the aerosol delivery channel port. The opening of the aspiration port enables the patient to exhale through a very low resistance. Simultaneously, to vent the pressure generated inside the chamber-cone, the excess pressure is vented through a pressure relief valve, which may be set or adjustable. In a preferred embodiment, the bidirectional valve assembly comprises a flap valve operated pneumatically through a 5-2-way solenoid valve V4 in the console. Notably, there are no electrically conducting wires near the patient. Using pressure P regulated by the primary regulator R, the 5-2-way valve V4 may switch the direction of pressure, resulting in a 90 degree angular motion of the bidirectional valve. When the aerosol is generated for inspiration, a corresponding control signal is transmitted to the 5-2-way valve V4 to switch the pressure. This procedure results in a 90° rotation of the bidirectional flap valve, thereby opening the aerosol delivery channel port so that aerosol may be delivered to the patient. The rotation of the bidirectional flap valve simultaneously closes the aspiration port. When the 5-2-way valve V4 switches the pressure to an alternate position, the bidirectional flap valve closes the aerosol passage, while simultaneously opening the exhaust port, thereby facilitating exhalation by the patient.
In one embodiment, the nozzle may include a 500 μm diameter exit orifice and a gas flowrate of 30 L/min. Aerosol may be generated and delivered either continuously or in defined boluses during inspiration. This functionality is achieved by the synchronous activation and deactivation of aerosol generation dictated by the nozzle and the bidirectional valve. As discussed, aerosol may be generated or arrested in the nozzle by controlling the pressure in the aerosolization space within the nozzle. The aerosol velocity may be reduced in the chamber and cone by a counterflow tube. Subsequently, the bidirectional valve may delay the aerosol's delivery times, as necessary. To deliver aerosol pulses between 50 ml and 2 L, the liquid aerosol generation pulse duration may be selectable between 100 ms and 4 s, respectively. The final aerosol inspiratory and expiratory breaths by the patient are synchronized by the bidirectional valve.
Tidal volume for inspiration and periodic breath rate are determined according to the size of the patient. To increase the drug concentration while retaining the tidal volume, an aerosol concentrator may be placed in between the cone and the bidirectional valve. The cone, concentrator and bidirectional valves maybe configured to be easily interconnected or disconnected, but with connections sufficiently secure to prevent leakages. The volume of the aerosol is dependent on the volume of the aerosol chamber as well as additional gas from the counterflow tube. These parameters control the dilution of the overall aerosol concentration. An aerosol concentrator may be incorporated to reduce the volume of the gas. The concentrator may comprise accelerating/converging slits and decelerating/diverging slits separated by 1.5 mm gap.
Due to their momentum, the higher inertial aerosol particles (2-5 μm in size) pass though the slits and through the gap into the decelerating slits, whereas lighter particles and dilution gas (<1 μm in size) escape through the sides of the gap and exit through the exhaust port in the concentrator. By having the dilution gas and other non-aerosol particles escape through the exhaust port, a high concentration of aerosols is outputted. The concentrator reduces gas flow to the patient by ⅕th of the initial flow (i.e., 6 l/min) while maintaining the same mass of the aerosolized drug. This procedure may increase the drug concentration up to fivefold.
Pressure may be routed to a liquid vial 23 through a gas supply line 13, along which a liquid pressure regulator 22 and a 3-2-way solenoid valve 24 are disposed. The liquid pressure regulator 22 may be configured to pressurize the liquid vial 23, which supplies liquid to the aerosol nozzle 7 via a liquid supply line 11. A 3-2-way solenoid valve V1 24 may be disposed between the liquid pressure regulator 22 and the liquid vial 23 to control the flow of liquid from the liquid vial 23 to the nozzle liquid input 60 of the nozzle 7. In a preferred embodiment, the 3/2-way solenoid valve 24 is by default in a closed position, and prior to aerosolization may switch to an open position, thereby turning the pressure on.
The system may additionally include a three-way solenoid valve 26 configured to control the flow of liquid within the liquid microchannel 70 inside the nozzle barrel 66. Regulated pressures from the system pressure regulator 21 and a nozzle gas regulator 25 are routed through the three-way solenoid valve 26. The pressures P, P1 and P2 respectively corresponding to the system regulator R 21, liquid pressure regulator R1 22 and nozzle gas regulator R2 25, are adjusted such that P>P1>P2. When the V2 26 is open to pressure P from regulator R 21, the pressurized liquid flow in the aerosol nozzle 7 is arrested. However, when the V2 26 is open to pressure P2 from regulator R2 25, the liquid flows through the aerosol nozzle 7 for aerosolization.
Pressure may be routed to counterflow tube 8 via a counterflow gas tube supply line 14 and expelled out of a counterflow gas tube exit opening 16. A counterflow tube regulator 27 may be included to control the gas pressure in the counterflow tube 8. A second 3-2-way solenoid valve 28 may be included intermediate to the counterflow tube regulator R3 27 and counterflow tube 8. When in its default closed position, the 3-2-way solenoid valve 28 allows for independent control of the counterflow tube 8.
A 5/2-way solenoid valve V4 29 may be included to control the pressure output to the aerosol delivery respiratory support 9, which comprises a bidirectional valve assembly 130. Two pneumatic lines 31 extend between the 5/2-way solenoid valve 29 and a pneumatic actuator 43 located at the bidirectional valve assembly 130. Pressure regulated by the system pressure regulator 21 may be switched between the two pneumatic lines 31 in order to open and close the aerosol delivery channel port 136 of the bidirectional valve assembly 130, thereby controlling whether the aerosol output to the patient is allowed or arrested.
The dispenser mount 42 may include a front recess through which the dispenser cap 41 is inserted, a back wall opposite the front recess, a top surface, and two side walls. The back wall of the dispenser mount 42 may include two threaded holes 46 for mounting the dispenser assembly onto the console 1. The liquid vial 23 is pressurized through a threaded port 47 located between the two threaded holes 46 on the back wall of the dispenser mount 42. The threaded port 47 extends to a perpendicular hollow channel 48 inside the dispenser mount 42 that directs the compressed gas downwards into the vial 23 (
In one configuration (as shown in
Diverging of the input tube is optional but may be preferred for increasing the aerosol velocity through the cross-slit and therefore the momentum of the aerosol particles, resulting in a higher particle concentration rate. Converging of the output tube is likewise optional but preferred for reducing the velocity of the aerosol flow for a smoother for inhalation action and distribution of the aerosol particles. In lieu of converging/diverging profiles within the input/output tubes, a similar effect could also be accomplished by including additional components with similar profiles connected upstream/downstream of the input/output tubes. In this context, converging is to be understood as reducing the total cross-section of the lumen in the input tube along the direction of aerosol flow, while diverging is to be understood as increasing the total cross-section of the lumen in the output tube along with the direction of flow.
In the following, additional embodiments of the invention are described:
Embodiment 1. An aerosol concentrator 105 dividing an input aerosol of an input particle concentration into a concentrated respirable aerosol of an increased particle concentration that is higher than the input particle concentration and an exhaust aerosol of a lower particle concentration that is lower than the input particle concentration, said concentrator 105 comprising:
Embodiment 2. The aerosol concentrator 105 according to embodiment 1, wherein the input tube 109 converges from its input tube entrance port 119 to its input tube exit port 120 from an input tube entrance port cross-sectional area that is larger than an input tube exit port cross-sectional area.
Embodiment 3. The aerosol concentrator 105 according to embodiments 1 or 2, wherein the output tube 110 diverges from its output tube entrance port 121 to its output tube exit port 122 from an output tube entrance port cross-sectional area that is smaller than an output tube exit port cross-sectional area.
Embodiment 4. The aerosol concentrator 105 according to one of embodiments 1-3, wherein the input tube 109 and the output tube 110 have a joint longitudinal axis and have radial sizes transverse to the longitudinal axis that are less than a third of a radial size of the plenum 124 transverse to said longitudinal axis.
Embodiment 5. The aerosol concentrator 105 according to one of the embodiments 1-4, wherein the housing 106, 107 compromises an input housing part 106 holding the input tube 109, and output housing part 107 holding the output tube 110 and an exhaust tube 111, and a seal 114 is provided between the input housing part 106 and the output housing part 107.
Embodiment 6. The aerosol concentrator 105 according to embodiments 4 or 5, wherein the housing 106, 107 shares the same joint longitudinal axis with the input tube 109 and the output tube 110 and the plenum 124 is encompassed at least for the most part by either one of the input housing part 106 or the output housing part 107, while the other part of the output 107 or input housing 106, respectively, is essentially disc-shaped, with the housing part encompassing at least for the most part the plenum 124 also having the exhaust tube 111 which extends essentially radially in a transverse direction with respect to said longitudinal axis.
Embodiment 7. The aerosol concentrator 105 according to embodiments 5 or 6, wherein a pressure relief valve is connected to the exhaust tube 111 keeping the pressure within and across concentrator 105 and consequently at the output tube exit port 122 essentially constant.
Embodiment 8. The aerosol concentrator 105 according to one of embodiments 1-7 wherein the concentrator 105 is designed to concentrate an input aerosol having fine particles of a particle size distribution of 1-6 μm MMAD suspended in gas.
Embodiment 9. The aerosol concentrator 105 according to one of embodiments 1-8 wherein the concentrator 105 is designed to concentrate an aerosol using virtual impaction at a small positive pressure within and across concentrator 105 all the way to the output tube exit port 122 such that the concentrated respirable aerosol can be delivered to the patient at the small positive pressure without the use of pumps to remove the exhausted gas.
The following is a list of reference numerals as shown in the drawings:
The present invention, in part, was supported by the National Institute of Health, Heart, Lung and Blood Institute under grants R44HL142335 and R43HL127834. The US government has certain rights to this invention.
