Patent Application
20200139074
The invention generally relates to devices and systems which both humidify a subject's airways and provide high efficiency delivery of pharmaceutical aerosols.
U.S. Pre-Grant Publication No. 2015-0007817 A1 discloses a prior mixer-heater. A challenge with some embodiments of this prior mixer-heater was the relatively large mixer volume required to accommodate the aerosol before the heating section. The large volume had the advantage of minimizing aerosol loss and was needed for continuous aerosolization. However, the large volume of the previous design could make synchronization of aerosol delivery with inhalation challenging. This is because of a travel time delay between the point of nebulization and the lungs. Only aerosol generated very early in the inhalation cycle could reach the lungs with a mixing volume of ˜500-1000 ml, since this is similar to an adult inhaled volume (˜500 ml) during passive breathing. High efficiency lung aerosol delivery required deep inhalations.
Difficulties arising from large device and system volumes are not the only issues in existing aerosol systems. While existing HFNC systems effectively deliver humidified air, they are very inefficient at delivering pharmaceutical aerosols. For example, Perry et al. reported ex-cannula aerosol dose was <0.4% of the nominal dose at typical adult HFNC flow rates of 20 LPM and above (Perry S A, Kesser K C, Geller D E, Selhorst D M, Rendle J K and Hertzog J H. Influences of Cannula Size and Flow Rate on Aerosol Drug Delivery Through the Vapotherm Humidified High-Flow Nasal Cannula System. Pediatr. Crit. Care Med. 2013; 14:E250-E256).
Positioning a mesh nebulizer upstream of the HFNC humidity unit, Reminiac et al. achieved 2 to 10% of nebulized dose downstream of an in vitro nasal model (Reminiac F, Vecellio L, Heuze-Vourc'h N, Petitcollin A, Respaud R, Cabrera M, Le Pennec D, Diot P and Ehrmann S. Aerosol therapy in adults receiving high flow nasal cannula oxygen therapy. Journal of Aerosol Medicine and Pulmonary Drug Delivery 2016; doi:10.1089/jamp.2015.1219).
A recent in vivo study of aerosol delivered simultaneously with a commercial HFNC system reported lung delivery efficiencies in the range of 1-4% of the nebulized dose (Dugernier J, Hesse M, Jumetz T, Bialais E, Roeseler J, Depoortere V, Michotte J-B, Wittebole X, Ehrmann S and Laterre P-F. Aerosol delivery with two nebulizers through high-flow nasal cannula: A randomized cross-over single-photon emission computed tomography study. Journal of Aerosol Medicine and Pulmonary Drug Delivery 2017; 30:349-358). Over 50% of the nebulized aerosol was lost in the delivery system.
The lung drug delivery efficiencies from nebulizers in infants is unacceptably low (<5% of the nominal dose delivered to the lungs) and therefore there is a need for the development of more efficient delivery systems synchronized to infant breathing, especially for high dose medications such as antibiotics. Patient-related factors combined with aerosol factors have contributed to poor delivery efficiencies. Continuous nebulization throughout the entire respiratory cycle of infants that have very short inspiratory times and small inhalation:exhalation ratios results in 6-9 times more drug lost than deposited in the lung. Furthermore, aerosol particle sizes of 4-6 μm obtained using conventional nebulizers have been associated with high (˜70%) nasal deposition.
It is desired that a combination mixer-heater device be configured to efficiently deliver inhaled medications to the respiratory system (e.g., lungs) while providing a continuous stream of airflow required for high flow nasal cannula (HFNC) ventilation support or other forms of noninvasive ventilation. While existing HFNC systems effectively deliver humidified air, they are very inefficient at delivering pharmaceutical aerosols. By contrast, some exemplary combination mixer-heater flow passages according to embodiments of the invention have approximately 5% or less depositional loss with an emitted aerosol drug dose from the mixer-heater of 80% or higher.
According to an aspect of some exemplary embodiments, a new mixer-heater flow path has a significantly reduced volume for adults (e.g., <150 ml) and for infants (e.g., <40 ml). With this reduced volume, nearly all of the aerosol generated during an inhalation cycle is able to reach the patient without requiring deep inhalation. If aerosolization occurs over the first half of inhalation, then all of the aerosol may be purged from the mixer-heater and reach the patient.
A challenge with reducing a mixer-heater volume (e.g., to 150 ml or below, or 40 ml or below) is that simultaneously maintaining low depositional loss in the unit is difficult. This is because of the momentum of the mesh nebulizer aerosol stream, or any aerosol stream, which pushes the aerosol into any boundary that is sufficiently close. To address this challenge some exemplary embodiments include one or more of the following aspects:
Optimal delivery of the drug aerosol requires synchronizing the mesh nebulizer with inhalation. With high flow nasal cannula therapy (HFNC), or some other forms of noninvasive ventilation such as low flow nasal cannula (LFNC) or continuous positive pressure ventilation, air is delivered constantly. According to an aspect of some exemplary embodiments, feedback control is used to keep the plates at a constant temperature, and this may serve as a standard process.
A primary challenge with a system having a single heating region for all flows (e.g., HFNC gas as well as aerosol) is that a significant amount of energy is required to fully evaporate the aerosol whereas a much smaller amount of energy is used to heat the gas stream when the drug nebulizer is off. For adults, the drug nebulizer may only be on for 0.2-2 s. For infants, the drug nebulizer may only be on for ˜0.1 s or less.
The challenge of a single heating region for all flows of a multi-flow system is illustrated well by a specific numeric example. Consider a HFNC configuration in which a plate temperature in the heating region provides 20.7 Watts (W) to a flow stream to evaporate the aerosol and heat the gas flow from 24° C. to a comfortable 32° C. under adult conditions (30 LPM airflow with liquid mass flow rate of 0.4 ml/min). However, when the drug nebulizer is off, a plate temperature in the heating region is required to provide only 4.6 W to heat a humidified gas stream (air and water vapor but not liquid droplets) from 24 to 32° C. This massive difference in required input power during the period when the (medicament aerosol) nebulizer is on versus when the nebulizer is off is due to the very large heat of vaporization of water. When the drug nebulizer is off, providing 20.7 W of power to a humidified gas stream of air flowing at 30 L/min will heat the gas stream from 24° C. to an uncomfortable and likely unsafe temperature of 60° C. (140° F.).
At the time of this disclosure's filing, it is not possible to control the plate temperature of a single aerosol heater in a way that it can swing from providing 20.7 W to 4.6 W over a fraction of a second (e.g., 0.2-1 s for an adult, ˜0.1 s or less for an infant). One potential solution to this problem is to have separate heaters for the gas stream and aerosol each with separate feedback controllers; however, this solution is overcomplicated, requiring additional space, costs, and maintenance.
According to an aspect of some exemplary embodiments of the invention, a solution to the preceding problem in multiflow systems having a single heating region is to use an alternating nebulizer system in which one nebulizer is used as a humidity source and one nebulizer is used to deliver the medicament aerosol. Both nebulizers deliver approximately the same aqueous liquid flow rate. One of the two nebulizers is actuated at all times in an alternating manner. When the medicament nebulizer is actuated during a period of inhalation, it supplies the drug and necessary humidity and the separate humidity nebulizer is not actuated. During all remaining times, the humidity nebulizer is actuated to humidify the continuously flowing gas stream and the drug nebulizer is off to avoid wasting medication and improve lung delivery efficiency. In this manner, a constant flow of ventilation gas (measured in L/min or LPM) and nebulizer solution (measured in ml/min) moves through the heating section at all times, requiring a constant power input and avoiding temperature swings.
A feedback control may be used to keep the single plate temperature at a constant value (e.g., provides ˜21 W of power for an adult system operating with a nebulization rate of 0.4 ml/min and an airflow rate of 30 LPM) regardless of whether the drug nebulizer is on or off at any given moment. When the drug nebulizer is actuated during a brief inhalation period, sufficient energy is available to fully dry the aerosol and a safe temperature (e.g., of ˜32° C.) is provided to the patient. Similarly during the remainder of the breathing cycle when the humidity nebulizer is actuated (and the drug nebulizer is off), the flow stream is humidified from the evaporating droplets and a safe inhalation temperature of 32° C. is maintained. Providing drug and humidity from separate nebulizers thus produces a simplified system with one heating channel (or pathway).
According to an aspect of some embodiments, a very low volume mixer heater (VLVMH) is provided for exceptionally low inhalation volume drug delivery applications (e.g., with infants) to address both issues of synchronization and timing of aerosol delivery together with reducing the aerosol size to minimize aerosol losses.
According to an aspect of some embodiments, the mixing section of a mixer-heater may be preceded by or else include a flow unifier. A flow unifier may be configured as a perforated plate near an air inlet to help unify incoming airflow. The exit of the mixing region may extend along the top of the device to provide a reduction in depositional loss. The heating section may have an elliptical or rectangular cross-section and end with a streamlined taper leading to outlet tubing. The heating section may align with the gravity vector or be perpendicular to the gravity vector. Whichever the orientation, it may be determined using the major axis of the ellipse or rectangle. In other words, the major axis of the ellipse may be aligned (i.e., parallel) to the gravity vector, or the major axis of the ellipse may be perpendicular (i.e., orthogonal) to the gravity vector.
According to an aspect of some embodiments, mixer-heater volume may be 150 ml or less, in some cases 100 ml or less, in some cases 40 ml or less. The mixer-heater volume may be measured starting at the cross-sectional plane which meets the center of the last nebulizer along the flow path and ending where the mixer-heater meets outlet tubing.
Exemplary embodiments include combination devices, in particular mixer-heaters, and systems which employ such combination devices. As the name implies, a combination device combines a plurality of constituent elements. In the present case, separate nebulizer outputs and therapeutic flows are combined to address problems such as those set out in the Background above.
The streams 125 and 135 may meet the ventilation gas stream 115 in a cross-flow or cross-stream configuration. The streams 125 and 135 may be introduced approximately perpendicular to the primary direction of streams 115.
A ventilation stream 115 is in and of itself common. A variety of patients requiring medical care today are subject to ventilation systems, especially high flow or low flow nasal cannula therapy (HFNC therapy or LFNC therapy, respectively). A ventilation source 110 supplies a continuous gas stream 115 that is ultimately supplied to the patient through a nasal cannula or other inhalation patient interface (e.g., mask, intubation tube, etc.).
Where exemplary embodiments herein notably deviate from existing HFNC and LFNC ventilation systems is the addition of humidity and drug aerosol streams. Exemplary embodiments combine with a continuous ventilation gas flow 115 both a humidity stream 125 and a drug aerosol stream 135. More specifically, a mixer-heater 140 may operate at a constant heating power setting (e.g., constant wattage) while ventilation stream 115 is continuous, humidity stream 125 is intermittent, and drug aerosol stream 135 is intermittent.
In some preferred embodiments, the intermittent streams 125 and 135 are produced at alternating intervals. While humidity nebulizer 120 is actuated (and thus producing an output), the drug nebulizer 140 is not actuated (and thus not producing an output). Conversely, while drug nebulizer 120 is actuated, the humidity nebulizer 140 is not actuated. The alternating actuation of the nebulizers may also be configured such that the drug nebulizer is actuated in synchrony with part or whole of the patient inhalation cycle while the humidity nebulizer is actuated in synchrony with the remaining time of the breathing cycle or the patient exhalation. The following are acceptable ranges of liquid flow rates from each nebulizer: for a typical adult, 0.8 ml/min to 0.1 ml/min; for children: 0.4 ml/min to 0.01 ml/min; for high dose medications or assuring 100% RH: ˜1.2 ml/min to 3 ml/min.
When the drug nebulizer is actuated for the brief periods (e.g., 0.2 s) each inhalation, it provides drug and humidity from the drug formulation, and the humidity nebulizer is off. Drug and humidity nebulizers may deliver the same liquid mass flow rate, e.g., 0.4 ml/min of liquid. In this manner, the same ml/min of liquid (e.g., 0.4 ml/min) is flowing through the heating section at any given time during the course drug delivery to a patient. A constant or substantially constant liquid amount passing the heater enables a constant power input to the heater.
The humidity may be supplied by a first nebulizer 120 and the drug aerosol supplied by a second nebulizer 130. Known nebulizers may be used but frequently produce aerosol droplets which are not sized for efficient administration to the patient, where a target aerosol size is a droplet or particle diameter of approximately 2 μm or below. Straight from the nebulizer, a large percentage of the drug aerosol droplets may deposit on the conduits which conduct the aerosol from the nebulizer to the patient or else deposit at unintended locations of the patient's respiratory tract (e.g., the nose or throat instead of the lung alveoli). An exemplary mixer-heater 140 dries the drug aerosol, thereby reducing droplet size to a predetermined range or below a predetermined threshold. As a result of drying, the size (e.g., mass median aerodynamic diameter, or MMAD) of aerosol droplets is reduced. Reducing the MMAD significantly improves aerosol penetration through the delivery system, patient interface, extrathoracic airways and into the lungs.
Besides drying a drug aerosol, the mixer-heater 140 is importantly configured for supplying a predetermined level of humidity (e.g., expressed as relative humidity, or RH) by combining with the drug aerosol the humidity stream 125 from the nebulizer 120. Predetermined humidity levels may be achieved for conditioning a patient's airways or controlling excipient enhanced growth (EEG) aerosol delivery.
This disclosure distinguishes between continuous streams and intermittent streams. “Continuous” as used herein may mean ongoing without interruption for a predetermined time period (e.g., the full duration, start to finish, of administering nasal cannula therapy to a patient, or of administering/delivering a single dose or round of treatment of a drug to a patient). “Intermittent” may be regarded as the opposite of “continuous” and means one or more intervals exist during the predetermined time period during which the relevant event is not taking place. Generally, “continuous” and “intermittent” are used to describe actuation of a device like a nebulizer or the supply of an airstream or flow. A continuous event may or may not be constant. To illustrate, a continuous air flow over a 10 minute duration may mean that during the 10 minute window the air flow is never zero. However, the airflow may change (e.g., in the first 5 minutes be 20 LPM and in the second 5 minutes be 30 LPM; or cyclically rise and fall in synchronization with a breathing cycle). If a parameter (e.g., flow rate) is constant then the actual numerical value remains the same or substantially the same for the specified duration. An intermittent event may take any of a number of temporal forms, including for example cyclical, sinusoidal, or stochastic. “Alternating” intermittent streams are streams which have a relationship in which, as between two alternating streams, a maximum of one stream is running at any given time. A negligibly small temporal overlap in the alternate intervals may nevertheless occur.
The system 200 may further comprise a single control unit 210 which may be used to actuate the nebulizers 120 and 130 (e.g., at alternating intervals) and control the power setting (e.g., wattage) of the heating section of the mixer heater 140. Depending on the embodiment, nebulizer control 211 and temperature control 212 may be configured as independent controllers instead of aspects of a unitary control unit 210. The system 200 may comprise one or more temperature probes 220 which may be included and arranged with the mixer-heater 140 so the control unit 210 reliability maintains a constant setting in a feedback loop. The one or more probes 220 may be arranged at the heating element and/or at the outlet of the mixer-heater.
Outlet temperature is central to whether or not a mixer-heater is actually suited for its intended use. In embodiments herein which employ a constant power setting for the heating section, the outlet temperature may be of particular consequence. Too low a temperature may result in incomplete drying of an aerosol and discomfort to a patient. Too high a temperature risks not just discomfort but physical harm to a patient (e.g., burns to airways). Outlet temperature may be measured at the outlet of the mixer-heater 140 and/or the final outlet of the delivery system (e.g., the orifices of the nasal cannula 250 from which the combined stream 145 passes into the patient's airways). In an exemplary embodiment, the mixer-heater is configured to maintain an outlet temperature in the range of 27-42° C., or preferably 28 to 37 (body temp) ° C. In cases involving nasal administration the outlet temperature may be maintained by the mixer-heater in the range of 32±2° C. In actual practice, a mixer-heater may be operated over a startup period that precedes use with a patient. The startup period permits the cannula outlet to reach a temperature corresponding with the mixer-heater temperature and is a function of ventilation tubing length. The mixer-heater outlet and nasal cannula outlet may vary a small amount, e.g., 1 or 2° C. due to cooling in the ventilation conduits; this is within acceptable tolerances for outlet temperature targets after steady state operation is reached.
The flow rate of system 200 may conform to flow rates acceptable for adult high flow nasal cannula (HFNC) therapy. The flow rate of flow 145 which passes through the ventilator tubing 260 to cannula 250 and from there to the adult patient may be, for example, 12-45 LPM, specifically 20 or 30 LPM according to conventional rates at the time this disclosure was written. During a single therapy session, the cannula therapy airflow is generally constant and continuous. In general, flow rate of system 200 may be primarily or entirely determined by the ventilation source 110. Though the humidity nebulizer 120 produces a stream 125 and the drug nebulizer a stream 135, the flow rates of the streams 125 and 135 may be negligible compared to the flow rate of the ventilation stream 115 with which they join in the mixing section of the mixer-heater 140. Specifically, with the use of mesh nebulizers, streams 125 and 135 include no additional net airflow, but inject liquid formulation streams as an aerosol into the mixer, which is denoted in a nebulized liquid flow rate of ml/min.
A primary direction of flow in the mixing section of each mixer-heater is perpendicular to the gravity vector, i.e., flow is substantially horizontal. In fluid communication with the mixing sections are nebulizer inlets 307 and 308. Flows admitted to a mixing section from either nebulizer inlet are introduced by a cross-stream configuration which may facilitate mixing. A mixing region's starting boundary may be treated as the most upstream cross-sectional plane of the mixer-heater at which nebulizer output from inlet 307 (e.g., humidified air) is present. Thus inlet sections 303 and 353 are not part of the mixing section 301 or 351, respectively. Inlet sections 303 and 353 may help with ventilation stream unification prior to the mixing section. In some embodiments the inlet sections 303 and 353 may not be present.
A second mixing section boundary that defines where the section ends may be treated as the most downstream cross-sectional plane of the mixer-heater at which no temperature increase in the combination stream has yet to occur. This same boundary may mark the start of the heater section. Downstream of this boundary the temperature of the combination stream rises. The end boundary of the heater section may be defined as the cross-sectional plane of the mixer-heater at which no further temperature increase occurs. Generally this may correspond with the position along the mixer-heater at which the heating element or elements of the heating section ends. The outlet section at the downstream end of the mixer-heater may be configured to reduce the size of the combination stream to that of the interior of a ventilation tube.
The direction of flow in the heating section is perpendicular to the gravity vector. However, the long axis of an elliptical heating section's cross section may be either aligned parallel with or perpendicular to the gravity vector. In
The geometries of the mixing section and heating section may be aligned. As discussed herein, some exemplary embodiments have an elliptical heating section. Correspondingly, the mixing section may also be elliptical. In such case the major axes of the mixing section and heating section may be aligned in parallel. This arrangement may minimize changes in aerosol direction prior to evaporation in the heating section and thereby minimize depositional losses. In alternative arrangements the major axes of the respective sections may not be aligned in parallel (e.g., major axes may be perpendicular with respect to one another). A horizontal orientation such as is shown in
The heating section may comprise one or more heating elements, in particular one or more heating plates. In prototype embodiments discussed in the examples below, two plates were employed. The plates may be arranged parallel to one another. The plates may be heated with one or more heaters, for example, Polyimide Film heaters. An insulative material may be provided to shield the plates from the external environment. The insulative material may simply be the shell of the heating section which defines the general geometry and body shape of the flow conducting structure.
The heating section may comprise one or more temperature probes, e.g., thermocouples, to detect the real time temperature of the heating elements. The probes may in turn be connected to a temperature controller. The temperature controller regulates the heater power to attain a set-point temperature of the heating elements. A majority of the supplied energy goes into evaporating the aerosol due to the high heat of vaporization of water (˜16 W) with much less energy required to heat a ventilation gas airstream (˜5 W). Because either the humidity or drug nebulizer is actuated at all times (but generally there is never a time when both are concurrently actuated), wide temperature swings in the system are avoided as the drug nebulizer cycles on and off, and the system is able attain the thermocouple set-point temperature in a stable manner. The temperature controller may be the same control unit as the nebulizer controller, or the two may be independent control units.
In
A primary characteristic of an exemplary mixer-heater design is a minimized total internal volume, which will improve emptying of the aerosol from a device with limited airflow. Total internal volume of a mixer-heater may be calculated as the sum of its sections, e.g., the combined volumes of mixing section 301, heating section 302, and outlet section 304 for mixer-heater 300 (
For adults, passive inhalation times are typically 1.5 s or greater. An exemplary mixer-heater empties within 20% of this inhalation time, providing an emptying time of 0.3 s or less. To achieve this emptying time at a high-flow nasal cannula (HFNC) flow rate of 20 L/min (LPM) (or 333.3 cm3/s), the total system volume including connective outlet tubing should be 100 ml or less. In reference to
In contrast to the exemplary mixer-heaters 300 and 350 discussed above, a VLVMH 500 may have no internal heater. Instead, the VLVMH 500 may be configured for use in a system that includes a heated air source 501 configured to supply heated air of a temperature of 50-90° C. or 50-70° C., for example 60° C. The VLVMH 500 further comprises a grid inlet 505 configured to unify the flow of air into the mixer-heater. The VLVMH 500 is configured to operate with a drug nebulizer connected to inlet 507 and a humidity nebulizer connected to inlet 508 for humidified high-flow therapy by alternating with the drug delivery during treatment.
A first prototype low-volume mixer-heater 610 is shown in
Additional mixer-heaters considered include two configurations having 16 cm long heating sections. The first is mixer-heater 620 in
The prototype mixer-heaters were produced with a heat resistant material using 3D printing. Aerogen Solo mesh nebulizers were used as the separate humidity and drug sources. Actuations of the nebulizers and heating of the airstream was managed with a control unit. Considering nebulizer actuation, a standard Aeroneb Solo driving signal was alternated between the drug and humidity nebulizers at a set timing interval. To capture a wide range of potential adult breathing conditions, the drug nebulizer was actuated for a period of 1.5 s (approximate inhalation phase) followed by a 6 s pause in which the humidity nebulizer was actuated (approximate exhalation phase). As with conventional HFNC therapy, a constant flowrate of 20 or 30 LPM was passed through the system at all times.
The outer shell of the heating section, which was constructed in 3D printed material, contains parallel aluminum heating plates. The parallel aluminum heating plates were heated with Polyimide Film heaters. The two heating plates were in direct contact with the air with the heaters positioned on the back side of the plates next to the 3D printed material, which forms an insulating layer. Use of the metal plates serves to spread the plate temperature evenly, increasing the surface area for effective heat transfer. Approximately 1 cm from the end of the lower plate, a thermocouple was adhered to the metal. The thermocouple was further connected to a temperature controller. The temperature controller regulated the heater power to attain the set-point temperature at the location of the thermocouple.
In vitro experiments were used to evaluate all three prototypical low-volume mixer-heaters at flow rates of 20 and 30 LPM. These experiments included determination of depositional drug loss within the device and determination of the outlet drug particle size distribution at the exit of the mixer-heater. Among the nebulizers, the drug nebulizer is positioned nearest the heating section to minimize travel distance (device volume) and therefore maximize delivery of the more valuable medication. In all cases of this example, the drug nebulizer was filled with 0.5% w/v solute consisting 50% w/w albuterol sulfate and 50% w/w sodium chloride. The humidity nebulizer was filled with isotonic saline (0.9% w/v NaCl). The system was typically operated with the humidity nebulizer on for 3 minutes to allow for warm-up and stabilization. After the 3 minute warm-up period, the system was operated in alternating mode with the drug nebulizer actuated for 1.5 s increments followed by 6.0 s increments of the humidity nebulizer.
Both the depositional drug loss within the mixer-heater and aerodynamic particle size distribution of drug aerosols at the outlet were determined using the alternating mode at system flow rates of 20 and 30 LPM. The deposition study was performed by connecting the outlet of the mixer-heater to a low-resistance filter (Pulmoguard II™, Queset Medical, North Easton, Mass.) and a vacuum pump. Particle size distribution was measured by replacing the filter with an Andersen Cascade Impactor (ACI) operated at 28.3 LPM flow condition. In both studies, the drug nebulizer was actuated 60 times to ensure reliable dose collection. The apparatus was dissembled after each run and albuterol sulfate was collected by rinsing the flange of the drug nebulizer, mixer-heater, filter or ACI plates with a known amount of deionized water. The drug nebulizer was weighed before and after the experiment to calculate the nominal delivered dose. Samples were analyzed with HPLC using Allure® PFP Propyl column (5 μm, 2.1×150 mm, Restek Corporation, Bellefonte, Pa.) and 70% methanol: 30% 20 mM ammonium formate buffer with pH adjusted to 3.4 (v/v) as mobile phase (flow rate: 0.4 mL/min). Albuterol sulfate was detected using fluorescence detection at 276 nm excitation (ex) and 609 nm emission (em) (2475 FLR Detector, e2996 PDA detector, e2695 Separation Module, Waters, Milford, Mass.). The injection volume was 100 μL and calibration curves were linear in the range of 0.2-10.0 mcg/mL (r2>0.999).
Initial experiments were conducted to determine nebulizer performance. Liquid nebulization rates of three different new Aeroneb Solo nebulizers tested three times each were determined on a gravimetric basis. The nebulizers were filled with 2 ml of a 0.9% w/v NaCl solution and operated for 5 minutes. The mean (standard deviation; SD) liquid nebulization rate was 0.4 (0.02) ml/min. The speed of the aerosol plume exiting the Aeroneb Solo device at a position approximately 2 cm from the mesh (just below the nebulizer outlet flange) was determined using high speed video recordings. The aerosol plume velocity was approximately 3.8 m/s; however, establishing variability was difficult due to inherent transient oscillations. The droplet diameter exiting the Aeroneb Solo device was measured using the ACI operating with near 100% RH air to prevent droplet evaporation. The resulting mean mass median aerodynamic diameter (MMAD) of the initial Aeroneb Solo aerosol with a 0.5% w/v solution of 50% AS and 50% NaCl was 5.3 (0.1) μm with a geometric standard deviation (GSD) of 2.2 (0.4) μm.
Depositional drug loss in each section of the delivery system is reported in Table 1 as a percentage of the nebulized dose of drug. Values are reported as mean (standard deviation) of three or more experiments (n≥3). The outlet filter percentage represents the delivery efficiency out of the mixer-heater device. The mixer-heater 620 reduced depositional loss from ˜11% (i.e., the loss from mixer-heater 610) to approximately 5-6%. The mixer-heater 630 further reduced depositional loss to values below 5%. Considering evaporation of the aerosol, all three prototypes effectively reduced aerosol size to approximately 1.5-1.6 μm, which is likely the fully dried size of the polydisperse liquid aerosol once all of the liquid is evaporated and only dried particles of solute remain. The 16 cm prototypes (mixer-heaters 620 and 630) improved device delivery efficiency to approximately 80%, with the best case of >85% achieved by the vertical 16 cm configuration of mixer-heater 630 with 30 LPM airflow.
Velocity measurements were conducted with a pitot tube pressure measurement device to evaluate the velocity field entering the mixing region. As shown in
Pitot tube measurements were made at six locations that traversed the mixer-heater inlets downstream of the flow unifier just before the first nebulizer inlet. The measurement locations were determined with a 6-point log-tchebycheff method to accommodate the diameter difference between mixer-heaters 610 and 620. Velocities in the primary direction of flow were measured and converted to standard meter/s velocity units. Velocity values along the flow path vertical and horizontal centerlines are plotted in
Computational fluid dynamics (CFD) simulations were preformed to evaluate transport of droplets through the system. Monodisperse 5.3 μm droplets were injected at the nebulizer inlet with an initial downward velocity of 3.8 m/s, based on experimental measurements. The ventilation gas flow rate through the system was 30 LPM. CFD simulations accounted for turbulent flow, heat and mass transfer, turbulent particle dispersion, and evaporation of the droplets including hygroscopic and solute effects. Grid independence of the hexahedral mesh was established and solution convergence was based on reduction of all residuals by at least three orders of magnitude. All equations were discretized to be at least second order accurate.
CFD predictions of droplet trajectories illustrate the downward momentum of the nebulized aerosol combined with the cross-flow of ventilation gas.
Average droplet residence time based on CFD predictions is also reported in
Current HFNC gas delivery systems are clearly inefficient at delivering inhaled pharmaceutical aerosols. The intent of these systems is to provide gas support at airflow rates of approximately 10 LPM and above in a continuous manner that is warmed and humidified. However, the need for the airstream to be fully saturated with water vapor (100% RH) and at 37° C. for nasal inhalation, as provided by current commercial systems, has not been established. Target performance goals for the mixer heater were output temperatures greater than 32° C. up to 37° C. targeting the nasal valve and anterior turbinate region temperature range of 28−32° C. and % RH>30% to avoid surface irritation due to low osmolarity and liquid sputtering observed with saturated RH. A temperature below 38° C. is preferable for comfort and a temperature below 42° C. is preferable for safety.
Temperature and RH were measured at the outlet of the mixer-heater for the alternating operation mode at flow rates of 20 and 30 LPM. Studies were performed with the nominal thermocouple set heating temperatures of 60, 90, 110, 130° C. To capture exiting energy and humidity, a custom shell was prototyped to fit around the temperature and humidity probe tip (M170-HMP75 RH probe, Vaisala, Louisville, Colo.) and the shell positioned the tip parallel with the outlet of the mixer-heater. Temperature and RH measurements were recorded over a 5 minute period after the initial 3 minute startup period and time-averaged values were calculated.
Experimentally measured T and RH values for the mixer-heater 610 with a 60° C. thermocouple temperature are shown in Table 2 for alternating mode delivery (i.e., intermittent delivery). Measured RH values at 20 and 30 LPM were approximately 10% (relative difference) below analytically predicted values, likely because of aerosol depositional loss occurring with the experimental system. However, there was agreement between the measured and CFD predicted values at 30 LPM for both temperature (28.7 vs. 29.0° C.) and RH (40.3 vs. 42.0%). As with the CFD analysis, the desired outlet temperature of 32° C. was not attained in the experiments with a plate thermocouple temperature of 60° C. Further studies were performed at elevated plate temperatures and reveal that the target of >32° C. gas temperatures were achieved for each of the three prototypes (mixer-heaters 610, 620, 630) when operated at plate temperatures between 90-130° C. with flow rates of 20 and 30 LPM. Humidity targets of >30% RH were generally observed, however, the targets was not achieved for the mixer-heater 620 at 30 LPM possibly due to experimental error during humidity reading. In the future, this can be alleviated by small increases in the nebulization rate (liquid flow rate) of the nebulizers above approximately 0.4 ml/min.
Two low volume mixer heater systems were compared for their aerosol delivery performance to an in vitro airway model of a 6-month-old infant via a nasal cannula. A low volume prototype mixer-heater 610 (
An Aerogen® Solo vibrating mesh nebulizer (Aerogen Limited, Galway, Ireland) was used to generate aerosol into the mixer-heaters in an intermittent delivery mode. The aerosol output of the nebulizer was reduced by altering the voltage input (14.1 Vrms) to produce a rate of approximately 0.07 ml/min to ensure adequate drying of the aerosol, in contrast to the output rate of 0.3 ml/min produced by the original Aerogen® controller. The Aerogen Solo nebulizers were adapted by removing inlet collar walls to minimize the volume between the nebulizer mesh and the inlet of the dryer. The Aerogen® Solo nebulizer was filled with 100 μl of 0.5% w/v albuterol sulfate solution and allowed to run to dryness for each experiment.
The in vitro aerosol delivery experiments were performed using a realistic 6-month-old infant nose-mouth-throat model, which was created from a computed tomography scan of a 7.7 kg male infant and constructed using Mimics® (Materialize, Ann Arbor, Mich.) image segmentation software and CFD. The model anatomy includes nostrils, turbinates, nasopharynx, larynx and a portion of the trachea. Aerosol delivered through the infant nasal model was captured on a low resistance respiratory filter positioned at the exit of the trachea and was considered as the delivered in vitro lung dose. Albuterol sulfate deposition in the mixer heater devices, cannula, infant model and filter was recovered by washing and quantified by HPLC. Losses during exhalation and loss of inhaled dose through exhalation (i.e., drug loss) were estimated based on the difference between the nominal dose and the total recovery of albuterol. Studies were performed in both mixer-heaters to optimize the delivery of aerosolized drug to the simulated in vitro lungs to determine when nebulization should begin in the breathing cycle and the duration of nebulization.
A breath simulator was used to produce a realistic breathing profile in the airway model.
These values were a significant improvement over the current standard of care and larger mixer-heaters. Table 4 compares the aerosol size characteristics of the Aerogen Solo nebulizer and the particle size from the reduced output nebulizers following drying in the mixer-heaters.
In order to reduce the total aerosol delivery time while achieving maximum lung delivery efficiency, the duration of nebulization was increased in subsequent studies. From the realistic breathing profile, the time during which the inspiratory flow was greater than the entrained airflow of 4-6 LPM, determined to be 0.3 seconds, was hypothesized to enable maximum drug delivery. Accordingly, nebulization durations of 0.1, 0.2 and 0.3 s were investigated to study their effects on aerosol deposition. These were also controlled by a digital relay connected to the breath simulator. Optimized delivery was achieved using the lowest nebulization duration of 0.1 s for each of the mixer heaters as shown in
While exemplary embodiments of the present invention have been disclosed herein, one skilled in the art will recognize that various changes and modifications may be made without departing from the scope of the invention as defined by the following claims.
This application claims the benefit of U.S. Provisional Patent Application Nos. 62/512,750, filed May 31, 2017, and 62/659,985, filed Apr. 19, 2018. The complete contents of both provisional patent applications are herein incorporated by reference.
This invention was made with government support under Grant No. 2R01HL107333-05A1 awarded by the National Institutes of Health. The government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2018/035456 | 5/31/2018 | WO | 00 |
| Number | Date | Country | |
|---|---|---|---|
| 62512750 | May 2017 | US | |
| 62659985 | Apr 2018 | US |
