Patent Application
20080271732
The lung is the essential respiration organ in air-breathing vertebrates, including humans. Its principal function is to transport oxygen from the atmosphere into the bloodstream, and to excrete carbon dioxide from the bloodstream into the atmosphere. This exchange of gases is accomplished by a mosaic of specialized cells that form millions of tiny, thin-walled air sacs called alveoli. Beyond respiratory functions, the lungs also act as an efficient drug delivery mechanism.
In recognition of the potential of pulmonary drug delivery, various efforts have been made toward developing effective pulmonary drug delivery devices. Current pulmonary drug delivery devices include metered dose inhalers (MDIs), dry powder inhalers (DPIs), and nebulizers. MDIs are pressurized hand-held devices that use propellants for delivering liquid medicines to the lungs. DPIs also use propellants, but deliver medicines in powder form. Nebulizers, also called “atomizers,” pump air or oxygen through a liquid medicine to create a vapor that is inhaled by the patient.
Each of the above-described devices suffer from disadvantages that decrease their attractiveness as a mechanism for pulmonary drug delivery. For example, when MDIs are used, medicine may be deposited at different levels of the pulmonary tree, and therefore may be absorbed to different degrees, depending on the timing of the delivery of the medicine in relation to the inhalation cycle. Accordingly, actual deposition of medicine in the lungs during patient use may differ from that measured in a controlled laboratory setting. Furthermore, a portion of the “metered dose” may be lost in the mouthpiece or the oropharynx.
Although DPIs reflect an effort to improve upon MDIs, small volume powder metering is not as precise as the metering of liquids. Therefore, the desired dosage of medicine may not actually be administered when a DPI is used. Furthermore, ambient environmental conditions, especially humidity, can adversely effect the likelihood of the medicine actually reaching the lungs.
Nebulizers may also exhibit unacceptable variability in delivered dosages, especially when they are of the inexpensive, imprecise variety that is common today. Although more expensive nebulizers are capable of delivering more precise dosages, the need for a compressed gas supply that significantly limits portability and the need for frequent cleaning to prevent bacterial colonization renders such nebulizers less desirable. Furthermore, the relatively high cost of such nebulizers also makes their use less attractive.
From the above, it can be appreciated that it would be desirable to have an improved pulmonary drug delivery system or device that avoids one or more of the above-described disadvantages.
Of the various applications in which such an improved pulmonary drug delivery system and device could be used, the delivery of nicotine as a method to achieve smoking cessation is one of the most compelling. The adverse health care consequences of smoking tobacco are enormous and incontrovertible. According the World Health Organization (WHO), tobacco is the second major cause of death in the world, currently accounting for one in ten deaths worldwide (5 million each year), and is the single largest preventable cause of disease and premature death. Of the 1.1 billion smokers in the world today, half will die from tobacco-related illness. For example, it is estimated that smoking will contribute to the death of one third of all Chinese males under 30 years old currently alive. In the United States, the 1999 National Health Interview Survey estimated that 46.5 million adults smoke and that 440,000 die each year from smoking related causes. In men, smoking is estimated to decrease life expectancy by 13.2 years and in women by 14.5 years.
Furthermore, it is now understood that cigarette smoke is not only harmful to the smoker, but also can affect the health of non-smokers when they passively inhale the smoke of other peoples' cigarettes. Such “secondhand smoke” is a risk factor for numerous types of adult ailments including lung cancer, breast cancer, and heart disease. Secondhand smoke exposure also increases the risk of various diseases in children and infants.
Despite recognizing the health risks associated with their habit, smokers continue to smoke. The primary reason for this phenomenon relates to the effect that nicotine has on the central nervous system (CNS). At low serum levels, nicotine provides stimulatory effects, primarily through activation of the locus ceruleus within the cerebral cortex. Such stimulatory effects include increased concentration, decreased anxiety, improved mood, decreased appetite, and improved memory. At high serum levels, nicotine activates the limbic system and produces a sense of euphoria, commonly referred to as a “buzz.” Cigarette smokers are accustomed to achieving both of these effects.
After inhaling cigarette smoke, nicotine is absorbed across the alveolar membrane in the lungs, leading to a rapid rise of serum nicotine levels within a few seconds. Within five minutes of smoking, the average maximum concentration of nicotine in arterial blood rises to 49 nanograms per milliliter (ng/ml), thereby providing the euphoric buzz. As nicotine levels fall, the stimulant effects predominate for the next 1-2 hours. Soon after, however, withdrawal symptoms begin to develop. These symptoms include irritability, anger, impatience, restlessness, difficulty concentrating, increased appetite, anxiety, and depressed mood. Such withdrawal symptoms are normally relieved by smoking the next cigarette, thereby creating a potentially endless cycle.
Over the years, many efforts have been made to develop effective means for assisting smokers in quitting. Currently, there are several Federal Drug Administration (FDA) approved nicotine replacement treatments (NRTs) intended for use in smoking cessation available both over-the-counter and as a prescription. Significantly, none of those NRTs deliver significant amounts of nicotine to the alveolar level of the lungs. Instead, they rely on the absorption of nicotine across the skin or across the nasal, buccal, or oropharyngeal mucosa. As a result, absorption is much slower and much less efficient than that typical of smoking and therefore leads to slower and much lower peak nicotine concentrations compared to that produced by cigarettes. Notably, this is true for existing nicotine inhalers, which are purported to have delivery characteristics most like cigarettes. Studies have confirmed that nicotine absorption resulting from use of such inhalers primarily occurs across the buccal mucosa, not the lungs, and that the arterial nicotine concentration spike that results from cigarette smoking does not occur with such inhalers.
The peak serum levels achieved with the current NRTs may be adequate to ameliorate or prevent withdrawal symptoms. However, they do little to satisfy the acute craving for the “buzz” created by the rapid onset and high peak serum nicotine levels typical of tobacco smoke. This may be the primary reason why so few habitual smokers that have used NRT have achieved long-term success. Instead, such persons typically give in to the persistent cravings, which currently can only be satisfied through smoking.
Given the enormity of the health problems caused by smoking, it is agreed upon by physicians and laypersons alike that the best thing that smokers can do is quit smoking. However, given the limited success that previous cessation solutions have had, it is clear that more effective alternatives are needed. It stands to reason that an alternative capable of providing nicotine to the user in ways analogous to smoking could save numerous lives.
The disclosed pulmonary drug delivery devices can be better understood with reference to the following drawings. The components in the drawings are not necessarily to scale.
As described above, it would be desirable have a pulmonary drug delivery system or device that is effective in enabling absorption of medicines, such as nicotine, via the lungs. Embodiments of a pulmonary drug delivery device are described in the following disclosure.
Disclosed herein are various embodiments of apparatuses and methods for pulmonary drug delivery. It is noted that those embodiments comprise mere implementations of the disclosed inventions and that alternative embodiments are both possible and intended to fall within the scope of the present disclosure.
Referring to the drawings, in which like numerals indicate corresponding parts throughout the several views,
Extending from the top side 18 of the housing 12 is a mouthpiece 34 that is used to deliver medicine to a patient who uses the device 10 (i.e., a “user”). In the embodiment of
Also provided within the interior space 42 is a fan 50 that is used to generate airflow within the drug delivery member 44. As indicated in
Further provided within the interior space 42 is a circuit board 60, which is more clearly shown in the exploded view of
With continued reference to
As is further indicated in
Turning to
The above-mentioned support structure will now be described with reference to
Referring next to
Adjacent the top end of the platform 122 is a seat 132 that is adapted to receive and support a head 136 of an electrical cable 138 that electrically couples the droplet ejection device 98 (
Example configurations for the pulmonary drug delivery device 10 having been described in the foregoing, examples of operation of the device will now be described. As explained above, the device 10 can be activated to deliver medicine to the respiratory system of the user upon detecting user inhalation as indicated by a drop in pressure within the upper tube 48 of the drug delivery member 44. The pressure drop can be detected by the pressure sensor 64 and an appropriate detection signal can then be sent from the sensor to the device microcontroller 62. The microcontroller 62 can then activate the fan 50 to cause it to draw in air from the environment, for example through the inlet 28 provided in the front cover 24, and exhaust the air through the opening 80 of the lower tube 46, as indicated by flow arrow 156 in
Due to the nature of the fan 50, the air is exhausted at a relatively precise angle relative to the lower tube 46. By way of example, the exhaust angle, α, is approximately 10 to 40 degrees relative to a horizontal direction that is parallel to the longitudinal axis of the lower tube 46. As mentioned above, a sharp angle is formed between the lower tube 46 and the upper tube 48. By way of example, that angle is approximately 70 to 120 degrees, for example approximately 90 degrees. Due to that sharp angle, the air exhausted by the fan 50 impinges upon the walls of the upper tube 48 and becomes highly turbulent within a turbulence zone 158 adjacent the intersection between the lower and upper tubes 46, 48 (i.e., at the sharp “bend” of the drug delivery tube). As is schematically indicated by flow arrows 160, the air vigorously circulates with the turbulence zone 158 before being forced up through the upper tube 48, as indicated by flow arrow 162.
Simultaneous to or soon after activation of the fan 50, the microcontroller 62 activates the droplet ejection device 98 to cause droplets of medicine to be ejected from the nozzles 154 of the ejection head 150. In some embodiments, the nozzles 154 are selectively activated to ensure a desired separation in terms of both distance and time. For example, the nozzles 154 can be activated such that only non-adjacent nozzles eject in sequence and a period of at least approximately 150 to 500 microseconds (μs) passes between firing of any two nozzles. Such an activation scheme ensures that the droplets are physically spaced to a degree at which evaporation of a droplet is not significantly influenced by the proximity of one or more other droplets.
Irrespective of the nozzle activation scheme that is implemented, the ejected droplets travel along the pathway 120 of the medicine injection tube 118 in the direction of arrow 164, which forms an angle, β, of approximately 30 to 60 degrees relative to the horizontal direction and which is generally opposite to the direction of the airflow generated by the fan 50. As indicated in
In order to achieve effective systemic absorption, it is normally desirable to deliver a medicine directly to the alveoli located deep within the lung structure where transport to the bloodstream is most quickly and efficiently accomplished. Lung deposition curves, such as those published by the International Commission on Radiological Protection (ICRP), indicate that the locations within the pulmonary tree in which inhaled particles are deposited depends to a substantial degree upon particle size. Specifically, lung deposition curves based on both theoretical modeling and experimental data typically show that particle deposition rates in the alveolar regions of the lung are greatest for particles having a diameter of approximately 1 to 3 μm. In view of this, the device 10 can be configured to deliver droplets having a diameter of approximately 1 to 3 μm from the opening 40 of the mouthpiece 34. In other embodiments, the droplets have even smaller exit diameters, for example approximately 0.1 to 1 μm, to enable hygroscopic growth of the droplets within the respiratory tract.
With further reference to
After the desired quantity of medicine has been injected into the airflow during the current inhalation cycle, ejection of medicine droplets ceases and the fan 50 is powered down. The process can then be repeated for further inhalation cycles of the user until a desired dosage of medicine has been administered. If desired, the entire process can be repeated at a later time, such as later that day or the next day. In some embodiments, appropriate controls can be integrated into the device 10 to limit the frequency with which the medicine can be administered. For example, the microcontroller 62 can be programmed to limit operation of the device 10 once every hour, once every few hours, once each day, and the like.
As mentioned above, it may be desirable to deliver droplets having a diameter of approximately 1 to 3 μm from the opening 40 of the mouthpiece 34. It is possible, however, for the size of the droplets to fall outside of that range in some circumstances. For example, environmental conditions, such as temperature, humidity, and pressure, can cause the ejected droplets to shrink or grow. In some embodiments, measures may be taken, substantially in real time, to control the size of the droplets relative to feedback that is collected by the device 10. Such feedback can comprise, for example, one or more of the current atmospheric temperature, humidity, and pressure, or the size of the droplets that are being delivered. In the former case, the device 10 comprises an open feedback loop and, in the latter case, the device comprises a closed feedback loop. Irrespective of which feedback scheme is used, the actions to be taken can be determined through reference to a look-up table or through application of an appropriate algorithm, either of which can be stored within memory provided on the circuit board 60.
Generally speaking, the size of the droplets can be controlled before droplet ejection, during droplet ejection, and after droplet ejection. Various embodiments for controlling droplet size before, during, and after ejection are described in the following.
Prior to droplet ejection, the temperature of the medicine to be administered can be adjusted to control the size of the droplets that will be ejected. For example, relatively smaller droplets can be generated when the medicine is heated given that elevated temperatures decrease both the viscosity and surface tension of liquids, which translates into smaller droplets being ejected. In some embodiments, liquid temperatures in the range of approximately 45 to 110° C. are effective in reducing droplet diameter, with temperatures of approximately 90 to 99° C. being preferred in some embodiments.
Medication used in the device 10 can be heated using a variety of methods. Generally speaking, any method with which the medication is heated prior to its ejection (i.e., is preheated) can be used.
In a further embodiment, preheating can be achieved using the ejection elements of the droplet ejection head 150. For example, when the ejection elements comprise heater resistors, a relatively low voltage can be applied to the resistors when they are not being used to eject droplets so as to heat the medicine prior to such ejection.
As indicated in
In a variation on the control scheme described above in relation to
Yet another parameter that has a significant effect on the size of the droplets that are ejected is the composition of the medicine. In particular, the nature of the medicine used to form the droplets can have a significant effect on the rate at which the droplets evaporate. The evaporation rate of droplets depends to a significant degree on the properties of the solvent and the solutes present within the solvent. Volatile liquids (i.e., those with relatively high vapor pressures) evaporate more quickly than non-volatile liquids. Various solutes tend to affect the vapor pressure of the droplet surface in particular ways. Saline solutions, which comprise water and sodium chloride, are widely used as carriers for medicinal compounds due to their similarity to and compatibility with human tissues and biological processes. The presence of sodium chloride in such solutions tends to lower vapor pressures.
Evaporation and condensation typically occur simultaneously at the air-liquid interface of liquid droplets. The ratio of evaporation rate to condensation rate is dependent upon the vapor pressure at the droplet surface. As the concentration of sodium chloride in a saline solution increases, the ratio of evaporation to condensation decreases. At low relative humidity and elevated temperatures, saline solutions (e.g., a 0.9% solution) tend to have evaporation rates that are higher than condensation rates with a net result of evaporation and droplet shrinkage. As relative humidity increases, the rate of condensation relative to evaporation becomes larger until the droplet begins to gain mass and increase in size. Increasing the solute concentration in such a case will shift the point at which evaporation and condensation are at equilibrium to a point of lower humidity and higher temperature.
When separate compartments 180, 182 are used as described above, the composition of the medicine that is ejected can be controlled. For example, the first compartment 180 can contain a concentrated medicine while the second compartment 182 can contain an inert liquid, such as saline solution. The amounts of liquid that are provided into the mixing chamber 188 from each compartment 180, 180 can be controlled with the control elements 184, 186 relative to measured feedback as to environmental conditions and/or droplet sizes.
As mentioned above, certain parameters affect droplet size at the time of ejection. One such parameter is the size of the nozzles that are used to eject the droplets, which directly affects the size of the droplets. Such nozzle size variability can be achieved by providing a droplet ejection device that comprises nozzles of various different sizes.
As also mentioned above, the size of the droplets can be controlled after ejection. Therefore, even if the droplets entering the drug delivery member 48 are undesirably small or large, their size can be adjusted to ensure that the droplets enter the user's mouth at the optimal size (e.g., approximately 1 to 3 μm). The exercise of such control may generally be referred to as post-processing of the droplets.
It has been determined that droplet size can be significantly reduced due to evaporation of the ejected droplets during their flight to the user's respiratory tract. Such evaporation may naturally occur as a consequence of the current environmental conditions in which the system is used, such as temperature, humidity, and pressure. As the droplets evaporate, they lose fluid (e.g., water), which results in a corresponding loss of mass and volume and, ultimately, droplet diameter. Discussed in the following are several parameters that affect droplet evaporation rate and which therefore can be used to control droplet size.
One parameter that has a significant impact on droplet evaporation is air temperature. Specifically, the higher the temperature of the air that is being used to deliver the droplets to the respiratory tract, the greater the evaporation rate. Therefore, droplet size can be reduced by heating the air that flows through the system. In some embodiments, the air is heated from an ambient temperature (e.g., room temperature) to a temperature of approximately 20 to 60° C. The extent of droplet evaporation and size reduction obtained is dependent upon the particular air temperature that is reached as well as the duration of time the droplets are present within the heated air (i.e., time of flight to the respiratory tract), with higher temperatures and longer times of flight resulting in greater evaporation. The time of flight corresponds to the distance the droplets must travel to reach the respiratory tract and the speed with which the air is flowing toward the user. Therefore, the temperature to which the air is heated, the position at which the drug delivery unit is located relative to the patient interface, and the speed setting for the air supply blower can each be selected to obtain desired evaporation results.
In a further embodiment, ejected droplets can be heated by exposing the droplets to electromagnetic radiation. Such exposure can result in rapid temperature increase and, consequently, evaporation of the droplet. Generally speaking, substances absorb electromagnetic energy to different degrees depending on the wavelengths of the energy that is applied with the greatest overall absorption for a given liquid being achievable by selecting a wavelength that offers the greatest absorption for that liquid. By calculating the volume of the droplet and then using the heat of vaporization for the liquid of interest, the amount of energy required to evaporate the droplet can be determined
As a consequence of the droplets 202 passing through the electromagnetic energy 200, the droplets are rapidly heated and therefore rapidly evaporated so as to shrink. Such shrinkage is depicted in
In either of the embodiments of
Although a laser diode has been explicitly identified above, it is noted that other electromagnetic energy sources may be used with desirable results. For example, in some embodiments, light emitting diodes (LEDs) can be used in place of laser diodes.
Another parameter that has a significant effect on droplet size is the relative humidity of the air used to carry the droplets to the user. As one would expect, the lower the relative humidity of the air, the greater the droplet evaporation rate and therefore the smaller the diameter of the droplets when they reach the respiratory tract. In some embodiments, the air is dehumidified from an initial relative humidity to a reduced relative humidity. The extent of droplet evaporation and size reduction that can be achieved is dependent upon the particular environmental relative humidity and the duration of time the droplets are present within the airstream (time of flight), which corresponds to both the distance the droplets must travel to reach the respiratory tract and the speed with which the air that carries the droplets is flowing. Therefore, the relative humidity to which the air is reduced, the position at which the drug delivery unit is located relative to the patient interface, and the speed setting for the air supply blower can each be selected to obtain desired evaporation results. Just as dehumidification may be used as a means to decrease the size of the medicine droplets, humidification may be used to increase the size of the medicine droplets.
Another method with which droplet size can be controlled, and more particularly reduced, is to break up relatively large droplets into smaller droplets as they travel along the flow path of the drug delivery device.
In the foregoing, various parameters have been described that affect droplet evaporation and that therefore can be manipulated to control droplet size. Although each parameter is discussed separately, two or more of the parameters can be individually or simultaneously controlled in order to achieve a desired droplet size.
In cases in which the current atmospheric temperature, humidity, and pressure are to be measured (i.e., open-loop feedback), the circuit board 60 can, as described above, include one or more sensors 65 for detecting those conditions (see
The light source 230 and light detector 232 together comprise a droplet size sensing apparatus configured to capture light data regarding the droplets flowing through the upper tube 224. As indicated in
Droplet size can be measured in alternative ways. In one such alternative, ejected droplets are electrically charged and passed through an electric or magnetic field that laterally (relative to the direction of flight) deflects the droplets. When the resultant degree of deflection is determined, the mass, and therefore the size, of the droplets can be inferred.
As shown in
Later along the flow path 245, the droplets 240 pass through an electric or magnetic field 250 generated by a field generator 252. In some embodiments, the field generator 252 comprises one or more permanent magnets or electromagnets. In other embodiments, the field generator 252 comprises opposed plates provided on opposite sides of the flow path 245 (not shown) having a large potential difference. For an electric field, a force is imposed upon the droplets 240 given by the following relation:
F=qE [Equation 1]
where q is the charge on the droplets and E is the strength of the electric field. For a magnetic field, a force is imposed upon the droplets 240 given by the following relation:
F=qvβsinθ [Equation 2]
where q is the charge on the droplet, v is the velocity of the droplet, β is the strength of the magnetic field, and θ is the angle between the direction of travel and the magnetic field. The direction of the force, F, whether due to an electric or magnetic field, is perpendicular to direction of travel of the droplet 240. In
Because the charge on the droplets 240 remains substantially constant as evaporation occurs, the droplets enter the field 250 with nearly identical charges. Therefore, the lateral force imposed on the droplets 240 is substantially constant. However, the masses of the droplets will differ depending upon how much evaporation has occurred. It follows then that the smaller the droplet 240, the greater the force will affect the droplet and the greater the degree of deflection. Therefore, the size of the droplet 240 can be inferred from the amount of deflection of the droplet.
The deflection of the droplets 240 can be determined using conductive pads 254 placed on the wall 247. As indicated in
Notably, if the amplifier circuits are highly sensitive, they further can be used to detect passing droplets. In such a case, it would be possible to measure the speed of the droplet by determining the times at which the droplets pass the various contact pads. It is further noted that the applied charges can, in some embodiments, be used to adjust the size of the particles. Once a droplet is charged, the excess electrons arrange themselves on the surface of the droplet. Having like charges, the electrons naturally repel each other. As the droplet evaporates and the electrons are forced closer together, the repulsive forces increase. If the repulsive forces become great enough, they may break the droplet apart into multiple smaller droplets. Therefore, at a given charge level and percentage evaporation, droplet charge create smaller particles.
Various modifications can be made to the embodiments described in the foregoing. For example, in one alternative embodiment, an extension tube can be connected to the mouthpiece of the device and used to increase the distance between the device housing and the point at which medicine enters the user's mouth. In another alternative embodiment, the cap to the medicine container can include a vent port that equalizes the pressure within the container with that of the surrounding environment. In a further alternative embodiment, a screen can be placed over the passage formed within the container to filter particulate matter that could clog the droplet ejection device and/or to reduce surface tension that could interfere with the flow of medicine to the drug ejection device.
It is further noted that appropriate regulatory measures can be taken to avoid abuse of the device or the medicine(s) that the device is intended to administer. For example, each medicine storage and delivery unit can be sold separately as one-time use component that comprises identification data that can be read by the pulmonary drug delivery device when the unit is installed on the drug delivery member. If the device microcontroller determines from the identification data that the medicine storage and delivery unit does not contain a medicine for which the device has been prescribed, for example by a doctor, operation of the device can be disabled.
Finally, it is noted that absolute spatial terms such as “horizontal” and “vertical” have been used herein relative to the orientations of the device components shown in the drawings. Therefore, it is to be understood that such terms may not strictly apply in cases in which the orientation of the device is changed from that shown in the figures.
This application claims priority to U.S. provisional application Ser. No. 60/915,379 entitled “Controlling the Droplet Size in a Drug Delivery System Using Temperature Modification” and filed May 1, 2007, U.S. provisional application Ser. No. 60/915,390 entitled “Aerosol Generating Device” and filed May 1, 2007, and U.S. provisional application Ser. No. 60/915,408 entitled “Droplet Delivery Methods and Systems” and filed May 1, 2007. This application also comprises a continuation-in-part of U.S. non-provisional application Ser. No. 11/950,180 entitled “Systems, Methods, and Apparatuses for Pulmonary Drug Delivery” and filed on Dec. 4, 2007, and U.S. non-provisional application Ser. No. 11/950,154 entitled “Apparatuses and Methods for Pulmonary Drug Delivery” and filed on Dec. 4, 2007. Each of the foregoing applications is hereby entirely incorporated by reference into the present disclosure.
| Number | Date | Country | |
|---|---|---|---|
| 60915379 | May 2007 | US | |
| 60915390 | May 2007 | US | |
| 60915408 | May 2007 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | 11950180 | Dec 2007 | US |
| Child | 12037513 | US |
