AU784815B2 - Device and method for creating aerosols for drug delivery - Google Patents

Description

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AUSTRALIA

Patents Act 1990 UNIVERSIDAD DE SEVILLA COMPLETE

SPECIFICATION

STANDARD

PATENT

Invention Title: Device and method for creating aerosols for drug delivery The following statement is a full description of this invention including the best method of performing it known to us:cqiu4 uIC 10.40 rAA 01 Z U0iU OCUU r l11f% S tU. cc W&JUOV DEVICE AND METHOD FOR CREATING AEROSOLS FOR DRUG DELWVERY FIELD OF THE INVENTION This invention relates generally to the field of aerosols and more particularly to devices and methods for creating aerosols of pharmaceutical formulations for delivery to a human patient, preferably by inhalation.

BACKGROUND OF THE INVENTION Aerosolizing formulations for inhalation has been considered as a convenient alternative to injection. for decades. This alternative to injections is particularly interesting for drugs which cannot be delivered orally, e.g. insulin. Although most compounds will effectively move from the lungs into the circulatory system there is considerable unpredictability in how much aerosolized formulation reaches the areas of the, lungs where the material can move into the, circulatory system. This results in inefficiency and unpredictability of dosing. A number of devices have been proposed for improving the efficiency of aerosol delivery, monitoring patients and teaching patients to correctly use delivery devices.

There are several different types of devices which use generally different mechanisms and methodologies to produce aerosols for inhalation. The most commonly used device is a metered dose inhaler (MDJ) which comprises a drug formulation container with the formulation including a low boiling point propellant. The formulation is held in the container under pressure and a metered dose of formulation is released as an aerosol when the valve on the container is opened. The low boiling point propellant quickly evaporates or "flashes'0 when the formulation is exposed to atmospheric pressure outside the containe. The particles of formulation containing the drug without the propellant are inhaled into the patient's lungs and thereaftr migrate into the patient's circulatory system.

There are a number of different types of MD1 devices. Devices of this type are disclosed in U.S. Patents 5,404,871 issued April 11, 1995 and 5,364,838 issued November 1994.

Another type of device is the dry powder inhaler (DPI) device. As indicated by the name such devices use formulations of dry powder which powder is blown into an 4£'uiu% V4 10;Z0j rA 01 4 VO.LU osuu r.n. nitn Cc t.U. L' 2 aerosolized cloud via a burst of gas. Typical DPI devices are Shown in U.S. Patents 5,775,320 issued July 7. 1998 and 5,740,794 issued April 21, 1998.

yet another type of aerosol delivery device forces a formuhlation through a porous membrane. Formulation mnoing through the pores breaks up to form small particles which are inhaled by the patient. Devices of this type are shown in U.S.

'Patents 5,554,646 issued August 13, 1995 and 5,522,385 issued June 4, 1996.

Each of these devices has some advantages and disadvantages. The object of each is substantially the sane to repeatedly produce a fine mist aerosol wherein the particles are substantially uniform in size and within a size range of 1 micron to microns. A patient can be accurately dlosed if the device can repeatedly start with a given amount of formula and produce a known amount of aerosol with particles having sizes within a known -range. The present invention endeavours to provide a device and method for obtaining accurate repeatable dosing of a patient with an aerosol.

Any discussion of documents, acts, materials, devices, articles or the like which has been included in the present specification is solely for the purpose of providing a context for the present invention. It is not to be taken as an admission that any or all of the-se matters form part of the prior art. base or were common general knowledge in the field relevant to the present invention as it existed in Australia before the priority date of each claim of this application.

SUMM4ARY OF TBf INENION In a first aspect, the present invention is a method of producing coated particles, comprising the steps of;, forcing a liquid formulation comprising a first liquid through a channel of a first feeding source in a man-ner which causes a stream of the first liquid to be expelled from a first exit opening at a first velocity; forcing a liquid comprising a second liquid through a second channel concentrically positioned around the first channel in a manner which causes a stream of the second liquid to be expelled from a second exit opening at a velocity which is substantially the same as the first velocity whereby the stream of second liquid is concentrically positioned around the stream of the first liquid; forcing a gats through a pressure chamber indamanner which causes the gas to exit the pressure chamber from an exit orifice positioned downstream of the concentrically positioned streams of the first and second liquids; wherein the density of the liquid formulation comprising the first liquid is substantially the same as the density of the liquid comprising the second liquid, and L4I/YV% V4 1.4oI~ rtA 01 z vo±Uv 0£UU r fl1sX% I AJ M I VVYI 2A the gas focuses the concentrically positioned streams to a stable unified jet which flows out of the chamber exit orifice and breaks up into coated partidles of the first liquid coated with the second liquid.

In a second aspect. the present invention is a method of delivering en aerosol to a patient, comprising! forcing a pharmaceutically active liquid through a channel of a feeding source in a manner which causes the liquid to be expelled from an exit opening; forcing a gas through a pressure chamber in a manner which causes the gas to exit the pressure chamber from an eit orifice in front of a flow path of the liquid expelled from the exit opening of the feeding source: wherein a stable liquid-gas interface is maintained and the liqid forms a stable capillary jet focused on the exit orifice of the pressure chamber by the gas.

In ak third aspect, the present invention is a method of delivering aerosolized particles of a pharmaceutically active drug to a patient, comprising: feeding liquid formulation comprised of a pharmaceutically active drug through a liquid feeding source to an outlet; feeding gas through an orifice positioned in front of the outlet in a direction aligned with a direction of flow out of the outlet: wherein the feeding of liquid enad feeding of gas are each at a rate relative to each other so as to maintain a stable capillary microjet of liquid which exits the orifice and formas aerosolized particles having a size in the range of 0.1 Micron to microns.

In a fourth aspect the present invention is mondisperse aerosol of hollow droplets comprised of a carrier and a pharmaceutically active drug wherein the hollow droplets have air in their center and have an actual physical size diameter in a range of from 200 microns to 1 micron.

Aerosolized particles within a desired range 1 micron to 5 microns) are produced from a liquid formulation comprised of a pharmaceutical active drug end a carrier. The particles produced all have substantially the same particle diameter ±03/ to ±3096, e.g. all particles in the aerosol have a diameter of 2 microns ±t3%6 to ±10%6. The formulation is provided in any desired manner forced. through a channel of aL feeding needle and expelled out of an exit opening of the needle).

Simultaneously. gas contained in a pressure chamber (which surrounds at least the area where the formulation is provided, surrounds the exit opening of the needle) is forced out of an opening positioned in front of the formullation. e.g., directly in front of the flow path of the formulation being expelled from the feeding 24/04 '02 15:26 FAX 61 2 9810 8200 F.B. RICE CO. igi UU0 2B needle. Various parameters are adjusted to obtain a super critical flow of liquid characterized by a stable liquid-gas interface and a stable capillary jet of the liquid which forms particles on exiting the opening of the pressurized chamber which particles will all (90% or more) have substantially the same diameter, a monodisperse aerosol.

A preferred feature of the invention is to provide a device for aerosolized delivery of a pharmaceutically active drug formulation or a diagnostic formulation.

Another preferred feature is to provide a method of creating an aerosol of consistent particle size to 30% or preferably ;t3 to 10% difference in diameter) which aerosol is inhalable by a patient for aerosolized delivery of drugs or diagnostics.

Z4/U4 UZ 10:ZU kAA 01 Z 61U OZUU r.b. Mi%'L 1% tAI. U U to A feature of the invention is that the diameter of the opening from which liquid is expelled, the diameter of the opening from which gas is expelled and the distance between these two openings is adjustable and is adjuisted to obtain a stable liquid-gas interface wb~ich results in a stable capillary microjet being formed by the liquid expelled which microjet is focused on an exit opening by the flow of sunrrounding gas.

Another feature of the invention is that the viscosities and velocities of the fluids can be chosen with consideration ro other adjusted parameters to obtain a capillary supercritical flow of liquid.

Aiiothe.- feature of the invention is that the liquid can be a single liquid, rwo or more (miscible or immiscible) liquids mixed, a solution or a suspension.

An advantage of the invention is that the gas flowing with the p~articles prevents the particles frora agglomerating thereby maintaining a monodisperse aerosol.

An advaniage of the invention is that it consistently produces aerosols having particles with a desired particle diameter e.g. I to 5 microns.

Ii0 An advantage of the invention is tha the device of the invention is ener.gy efficient in terms of the energy used to creatm small particles for inhalation.

Another advantage is that the structure of the device and its use are simple.

Another advantage is chat clogging of the exit opening of the pressure chamber is substantially eliminated because liquid is kept out of contact with the surface of the exit opeaing by a surrounding focused funnel of gas which flows out of the pressure chamber exit opening.

Yet another advantage is that particles produced are substantially smaller in size than would be exDected based on the diameter of the exit opening of the pressure chamber due to focusing the flow of the liquid with the flow of surrounding aas.

An aspect of the invention is a hand-held. set f-contained portab le drugr delivery device- which consistently produces small aerosolized particles which are re-!ative!v uniform in size.

9G/UI uc riAA Di Z uO±u OZUU r.n. %AI. gj LU Another aspect of the invention is a device and mnethod which products multiple streams of aerosol thereby quickly aerosolizing a large dose of formulation for inhalation by a patient.

These and other aspects, features and advantages will become apparent to those skilled in the art upon reading this disclosure in combination with the figures provided.

Throughout this specification the word "comprise", or variations such as "comprises" or "comprising", will be understood to imply the inclusion of a stated element, integer or step, or group of elements, integers or steps, but not the exclusion of any other element, integer or step, or group of elements, integers or steps.

L'AlWeV'* JU ri A v0± L voiLU o£UU r.D. A 0& %IV. 4A-J U JL BRIEF DESCRIJ'TTON OF TH-E DRAWINGS Figure 1 is a schematic view showing the basic cozpouenrs of one exnbodinaezt of the inivention with a cylidrical feeding needle as a source of formulation.

Figure 2 is a schematic view of another embodiment of the invention with two 6- lccntric tubes as a source of formulation.

Figure 3 is a schematic view of Yet another embodiment showing a wedgeshaped planar source of formulat~ion', Figure 3a iImusares a cross-sectional side view of the plana feeding source and the interaction of the fluids. Figure 3b show a frontal view of the openings in the pressure chamber, with the multiple openings through which the /0atozate exits the devicc. Figure 3c illustrates; the channls that are optionaly formed within the piana feeding member. The channeLs are aligied with the openings in the pressure chamrber.

Fimure 4 is a schematic view of a stable capillary microjet being formed and flowing through an e=it opeang to thereafter form a monodisperse aerosol- Figure, 5 is a gph of data whr35 measured valves of d, er O/Q. ame plotted.

Finure 6 is a schemiatic view of the critical area of a device of the type shown in Figure I showing zas sizrounded by liquid excpelled into a liquid to form bubbles.

Finure 7 is a scbetiaric view as in Figure 6 but with the bubbles flowing into a 1gas- Figure 8 is aschemnac vew as iaFigure6 but with two immiscible Iiauinds flowing into a zas.

f'i/U1 V% V4o1-.Z rftA 01 4 VOIU OCLJU flsf -U.j V I'd DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS Before the present aerosol device and method are described, it is to be understood that this invention is not limited to the particular components and steps described, as such may, of course, vary. It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments only, and is not intended to be lIIng, since the scope of the present invention will be limited only by the appended claims.

It must be noted that as used herein and in the appended claims, the singular fonns "and," and "the" include plural referents unless the conteic clearly dictates otherwise.

Thus, for example, reference to "a particle" includes a plurality of particles and reference to "a fluid includes reference to a mixture of fluids, and equivalents thereof known to those skilled in the art, and so forth.

Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. Although any methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present invention, the preferred methods and materials are now described- All publications mentioned herein are incorporated herein by reference to disclose and describe the methods and/or materials in connection with which the publications are cited.

The publications disc*ussed herein are provided solely for their disclosure prior to the fiing date of the present application. Nothing herein is to be construed as an admission tba± the present invention is not entitled to antedate such publication by virtue of prior invention- Further, the dates of publication provided may be different from the actual publication dates which may need to be independently confirmed.

DEFINITIONS

The term "formulation" is used to describe any single liquid, mixture, solution, suspension or the like which is a flowable liqufid at room temperature and comprises either a pharmaceutically active compound, a diagnostic compound or a compound which becomes an active compound after entering a patient, e.g. prodrugs or DNA encoding proteins. The formulation is preferably fur-ther comprised of a carrier and more preferably is a l iquid that has physical properties viscosity) such that when the formulation is moved through a device of the invention the formulation is aerosolized into particles 1 i v io;: r"A 0 z vo.u o/uu r.D. A 09 Iej UJ to 10 microns in diameter) which are inhaled into the lungs of a patient and preferably reach the circulatory system. The carrier may be any pharmaceutically acceptable material and is preferably a flowable liquid which is compatible with the active agent.

Formulations are preferably solutions, aqueous solutions, ethanolic solutions, aqueous/ethanolic solutions, saline solutions, colloidal suspensions and microcrystalline suspensions. Formulations can be solutions or suspensions of drugs (including systemically active drugs, e.g. insulin, insulin analogs including monomeric insulin such as insulin lispro) in the above mentioned liquids.

The term "carrier' shall mean a substantially inactive (biologically) component of a formulation such as a pharmaceutically acceptable excipient material which an active ingredient such as a drug, a diagnostic agent, prodrug or a gene vector is mixed with, suspended or dissolved in. The carrier is preferably a flowable liquid. Useful carriers do not adversely interact with the drug, diagnostic or gene vector and have properties which allow for the formation of aerosolized particles preferably particles having a diameter in the range of 0.1 to 10.0 microns (more preferably 1 to 5 microns) when a formulation comprising the carrier and active ingredient is aerosolized out of a device of the invention.

Carriers include water, ethanol, saline solutions and mixtures thereof with pure water being preferred. Other carriers can be used provided that they can be formulated to create a suitable aerosol and do not adversely effect the active component or human lung tissue.

The term "active compound" means a compound which is either a pharmaceutically active drug, becomes or produces a therapeutic compound in the body or is a detectably labeled compound. In turn, a "phamnaceutically active drug" or therapeutic compound shall be interpreted to mean any pharmaceutically effective compound used in the treatment of disease. A "detectably labeled compound" includes compounds which are radioactively labeled see U.S. Patent 5,829,436 issued November 3, 1998 Re: detectable labels.

Pharmaceutical formulations for use in the invention may be comprised of any therapeutic agent that can be atomized for patient delivery, e.g. for pulmonary delivery using an inhalation delivery system. Examples of such pharmaceutical compositions may be found in the Physicians Desk Reference (1998), and Remington: The Science and Practice of Pharmacy 19 Edition (1995), both of which are incorporated herein by reference. These formulations may be in any form capable of atomization, such as a solution, a suspension, an -6- U I u'I 24/04 '02 15:27 FAX 61 2 9810_8200 F.B. RICE COi.

emulsion, a slurry, etc., provided the dynamics of the form render it capable of forming a capillary microjet upon exposure to a second fluid.

The terms "electronic particle sensing device" and "electronic detection system" refer to a device or other means for measuring and counting particles present in an aerosol mist.

The term "reporting means" is used herein to indicate a means by which information obtained from a monitoring device such as an electronic particle sensing device or a device which monitors and records information relating to the patient's respiratory movement, may be provided to the user.

The term 'aerosol bolus" is used herein to describe a volume of air greater than ml and less than 4 liters which has suspended within it particles of a formulation wherein the particles have a diameter in the range of 0. 1 to 10 microns, preferably 1 to microns, and preferably the totial volume of formulation is from 5 ,csI to 10,000 pl. About A41 of particles having a diameter of about 1 to 3 microns are present in a volume of about 50 ml to 2 liters, preferably 100 nml to 1,000 ml.

The term 'aerosol" meaw any part of an aerosol bolus as described above.

The terms air', "particle free air", 'aerosol free air," and the like, are used interchangeably herein to describe a volume of air which is substantially free of other material and, in particular, free of particles intentionally added such as particles of formulation which create the aerosol. The term mumn that the air does not include particles of formulation which have been intentionally added but is not intended to imply that the normal surrounding air has been filtered or treated to remove all particles although filtering can take place. Air is the preferred gas to use with drug delivery it being noted that other non-toxic gases, CO,~ cani be used.

The te r 'easuring'" describes an event whereby the total lung capacity, (2) inspiratory flow rate or in~spiratory volume of the patient is measured and/or calculated. The information is preferably used in order to deterine an optimal point in the inspiratory cycle at which to release an aerosol and/or a particle free volume of air.

An actual measurement of both rate and volume may be made or the rate can be directly measured and the volume calculated based on the measured rate. The total lung capacity can be measured or calculated based on the patient's height, sex and age. It is also preferable to continue measuring inspiratory flow during and after aerosol delivery and to -7- II A uio1 24/04 '02 15:28 FAX 61 2 9810 8200 F.B. RICE CU.

record inspiratory flow rate and volume before, during and after the release of formulation. Such reading makes it possible to determine if aerosolized formulation was properly delivered to the patient.

The term "monitoring" event shall comprise determining the character of an aerosol such as the size, number and speed of the particles in the aerosol, further, monitoring may comprise measuring lung functions such as inspiratory flow, inspiratory flow rate, and/or inspiratory volume so that a patient's lung function as defined herein, can be evaluated during before and/or after delivery thereby making it possible to evaluate the effect of formulation such as respiratory drug delivery on the patient's lung function.

The term "inspiratory flow rate" shall mean a value of air flow rate determined, calculated or measured based on the speed of the air passing a given point in a measuring device assuming atmospheric pressure 5% and a temperature in the range of about to The term 'inspiratory flow' shall be interpreted to mean a value of air flow calculated based on the speed of the air passing a given point along with the volume of the air that has passed that point with the volume calculation being based on integration of the flow rate data and assuming atmospheric pressure, 5% and temperature in the range of about 10"C to about 400C.

The term "inspiratory volume' shall mean a determined, measured or calculated volume of air passing a given point into the lungs of a patient assuming atmospheric pressure 5% and a temperature in the range of 10C to The terms "lung function" and "pulmonary function" are used interchangeably and shall be interpreted to mean physically measurable operations of a lung including but not limited to inspiratory and expiratory flow rates as well as lung volume, i.e., total lung capacity. Methods of quantitatively determining pulmonary function are used to measure lung function. Quantitative determination of pulmonary function may be important when delivering any formulation and in particular respiratory drugs in order to direct the aerosol to a specific area of the lung and to determine effectiveness. Methods of measuring pulmonary function most commonly employed in clinical practice involve timed measurement of inspiratory and expiratory maneuvers to measure specific parameters. For example, forced vital capacity (FVC) measures the total volume in liters exhaled by a patient forcefully from a deep initial inspiration- This parameter, when evaluated in -8conjunction with the forced expired volume in one second (FEV), allows bronchoconstriction to be quantitatively evaluated. A problem with forced vital capacity detrmination is that the forced vital capacity marnuver forced exhalation from maximum inspiration to maximum expiration) is largely technique dependent. In other words, a given patient may produce different FVC values during a sequence of consecutive FVC maneuvers. The FEF 25-75 or forced expiratory flow determined over the midportion of a forced exhalation maneuver tends to be less technique dependent than the FVC. Similarly, the FEV, tends to be less technique dependent than FVC. In addition to measuring volumes of exhaled air as indices of pulmonary function, the flow in liters per minute measured over differing portions of the expiratory cycle can be useful in determining the status of a patient's pulmonary function. In particular, the peak expiratory flow, taken as the highest air flow rate in liters per minute during a forced maximal exhalation, is well correlated with overall pulmonary function in a patient with asthma and other respiratory diseases. The present invention carries out treatment by administering formulation in a delivery event and monitoring lung function in a monitoring event. A series of such events may be carried out and repeated over time to determine if lung (see U.S. Patent 5,404,871) function is improved.

Each of the parameters discussed above is measured during quantitative spirometry. A patient's individual performance can be compared against his personal best data, individual indices can be compared with each other for an individual patient (e.g.

FFVl divided by FVC, producing a dimensionless index useful in assessing the severity of acute asthma symptoms), or each of these indices can be compared against an expected value. Expected values for indices derived from qaantitative spirometry are calculated as a function of the patient's sex, height, weight and age. For instance, standards exist for the calculation of expected indices and these are frequently reported along with the actual parameters derived for an individual patient during a monitoring event such as a quantitative spirometry test.

The terms "particles", "aerosolized particles" and "aerosolized particles of formulation" are used interchangeably herein and shall mean particles of formulation comprised of any pharmaceutically active ingredient, a compound which becomes such in the body a prodrug or DNA encoding a protein) or a diagnostic compound. Any of these is preferably with a carrier, a pharmaceutically active respiratory drug and -9w" OCUJ~ UV UV

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L-W. W--J V.L I carrier). Particles of a liquid formulation are formed upon forcing the formulation from, anywhere it is provided the feteding needle) and then out of the pressure chamber exit orifice. The particles have a size which is sufficiently small such that when the' particles are formed they remain suspended in the air for a sufficient amount of time such that the patient can inhale the particles into the patient's lungs. The particles have a size in the range of 0. 1 micron to about 10 microns, preferably 1 to 5 microns. Particle diameter is an aerodynamic diameter.

In referring to particle diameter, the diameter measurement given is the ALaerodynamic diameter' measurement which is the diameter of a sphere of unit density that has the same terminal sedimentation velocity in air under normal atmosplieaic conditions as the particle in question. T1his is pointe out in that it is difficult to accurately maeasure the diameter of small particles using current technology and the shape may be continually changing. Thus, the diamter of one particle of material of a given density will be sad to have the same diameter as another particle of the same or a different material if the two particles have the same terminal sedimentation velocity in air under the same conditions.

DEVICE IN GENERAL The present disclosure focuses on the use of an aerosol creation device to produce aerosolized liquid formulations for delivery of drugs to patients preferably by inhalation into the lungs.

The basic: technology of the invention comprises a means for supplying a first fluid;. and a pressure chamber supplied with a second fluid. The first fluid is generally a liquid and preferably aqueous. The second fluid is generally a gas and preferably. air.

However, the first fluid may be a gas and second fluid a liquid or both fluids may be liquid provided the first and second fluid are sufficiently different fronm each other (immiscible) so as to allow for the formation of a stable niicrojet of the first fluid moving from the Supply means to an exit port of the pressure chamber. Notwithstanding these different combinations of gas-liquid, liquid-gas, and liquid-liquid the invention is generally described with a Liquid formulation being expelled from the supply means and forming a stable microjet due to interaction with surrounding air flow focusing the microjet to flow Out Of an exit of the pressure chamber.

t'i/UVt U. I 0 FAA OJi. 0±U 01UU r.D. fIl., M uJV.Lo Formation of the microjet and its acceleration and ultimate particle formation are based on the abrupt pressure drop associated with the steep acceleration experienced by the first fluid a liquid) on passing through an exit orifice of the pressure chamber which holds the second fluid. On leaving the chamber the flow undergoes a certain pressure difference between the first fluid a liquid) and the second fluid a gas), which in turn produces a highly curved zone on the first fluid liquid) surface near the exit port of the pressure chamber and in the formation of a cuspidal point from which a steady microjet flows provided the amount of the first fluid the liquid) withdrawn through the exit port of the pressure chamber is replenished. Thus, in the same way that a glass lens or a lens of the eye focuses light to a given point, the flow of the gas surrounds and focuses the liquid into a stable microjet. The focusing effect of the surrounding flow ofgas creates a stream of liquid which is substantially smaller in diameter than the diameter of the exit orifice of the pressure chamber. This allows liquid to flow out of the pressure chamber orifice without touching the orifice, providing advantages including clogging of the exit orifice is virtually eliminated, contamination of flow due to contact with substances (e.g.

bacteria) on the orifice opening is virtually eliminated, and the diameter of the stream and the resulting particles are smaller than the diameter of the exit orifice of the chamber.

This is particularly desirable because it is difficult to precisely engineer holes which are very small in diameter. Further, in the absence of the focusing effect (and formation a stable microjet) flow of liquid out of an opening will result in particles which have about twice the diameter of the exit opening. An additional advantage is that the particles are not prone to agglomeration following exit from the chamber owing to the accelerating effect of the surrounding gas stream, which has a velocity larger than the liquid velocity.

Specific embodiments of aerosol creation devices are now described.

EMBODIMENT OF FIGURE 1 A first embodiment of the invention where the supply means is a cylindrical feeding needle supplying liquid into a pressurized chamber of gas is described below with reference to Figure 1.

The components of the embodiment of Figure 1 are as follows: 1. Feeding needle also referred to generally as a fluid source and a tube.

2. End of the feeding needle used to insert the liquid to be atomized.

-11 z4/va vz ia:zv rAA vi z voiv ozuu r.b. Azt't W.jU WJ VJLV 3. Pressure chamber.

4. Orifice used as gas inlet.

End of the feeding needle used to evacuate the liquid to be atomized.

6. Orifice through which withdrawal takes place.

7. Atoinizate (spray) also referred to as aerosol.

D, diameter of the feeding needle; Do diameter of the orifice through -which the microjet is passed; ef= axial length of the orifice through which withdrawal takes place; H =distance from the feeding needle to the microjet outlet P, =pressure inside the chamber, P,z atmospheric pressure.

An aerosolized drug delivery device of the invention may be any size but is preferably designed to be a hand-held, portable, self-contained device weighing less than 1 kilogram. Although the device can be configured in a variety of designs, the different designs will all include the essential components shown in Figure 1 or components which perform an equivalent function and obtain the desired results. Specifically, a drug delivery device of the invention will be comprised of at least one source of formulation a feeding needle with an opening 2) into which a liquid flowable formulation can be fed and an eWt opening 5 from which the formulation can be expelled. Ile feeding needle 1, or at least its exit opeining 5, is encompassed by a pressure chamber 3. The chamber 3 has inlet opening 4 which is used to feed gas into the chamber 3 and an exit opening 6 through which gas from the pressure chamber and liquid formulation from the feeding needle 3 are expelled creating an aerosol.

IIn Figure 1, the feeding needle and pressaire chamber are configured to obtain a desired result of producing an aerosol wherein the particles are small and uniform in size.

Preferably the particles have a size which is iii a range of0. 1 to 10 microns, more preferably 1 to 5 microns. Particles of less than 1 microu in diameter can be produced via the present invention. However, particles below I micron may be too small for inhalation as the particles may not settle in the lung during a normal breath hold and as such would be exhaled. The particles of any given aerosol all have about the same diameter with a relative standard deviation of 10 to 30% or more preferably 3 to 20%. Stating that particles of the aerosol have a particle diameter in a range of 1 to 5 microns does not mean that different particles will have differnt diameters and that some will have a diameter of 1 EmiCron While -12- 4/Ua UL 10:ZU FAA 01 Z u IiU o0UU r.D. U.

others of 5 microns. The particles in a given aerosol will all (preferably about 90% or more) have the same diameter to For example, the particles of a given aerosol will have a diameter of 2 microns to Such a monodisperse aerosol is created using the components and configuration as described above. However, other components and configurations will occur to those skilled in the art The object of each design will be to supply formulation so that it creates a stable capillary microjet which is accelerated and stabilized by tangential viscous stress exerted by the gas on the liquid surface. The stable microjet created by the gas leaves the area of the pressurized gas leaves the pressure chamber and exits the pressure chamber orifice) and splits into particles which have the desired size and uniformity.

The aerosol created is a monodisperse aerosol meaning that the size of the particles produced are relatively uniform in size. The relative standard deviation in particle size is in the range of from about 10% to about 30%, preferably 3% to 10% and most preferably 3.% or less. The size of aerosolized particles useful for inhalation is a diameter in the range of from about 0.1 micron to about 10 microns, more preferably about 1 micron to about 3 microns.

For purposes of simplicity the remainder of the detailed description of the operation of the device of Figure 1 will refer to the first fluid as liquid and the second fluid as gas. The parameter window used the set of special values for the liquid properties, flow-rate used, feeding needle diameter, orifice diameter, pressure ratio, etc.) should be large enough to be compatible with virtually any liquid (dynamic viscosities in the range from 10" to 1 kg in this way, the capillary microjet that emerges from the end of the feeding needle is absolutely stable and perturbations produced by breakage of the jet cannot travel upstream.

Downstream, the microjet splits into evenly shaped drops simply by effect of capillary instability (see, for example, Rayleigh, "On the instability ofjets", Proc. London Math. Soc., 4-13, 1878), similar in a manner to a laminar capillary jet falling from a half-open tap.

When the stationary, steady interface is created, the capillary jet that emerges from the end of the drop at the outlet of the feeding point is concentrically withdrawn into the nozzle. After the jet emerges from the drop, the liquid is accelerated by tangential sweeping forces exerted by the gas stream flowing on its surface, which gradually decreases the jet cross-section. Stated differently the gas flow acts as a lens and focuses and stabilizes the microjet as it moves toward and into the exit orifice of the pressure chamber.

JIU U -13- £'/iU4 UC9 10;V 1.AA 01 X VOIU OCU r.fD. A1 1 AU. W1G YL The forces exerted by the second fluid a gas) flow on the firs: fluid a liquid) surface should be steady enough to prevent irregular surface oscillations. Therefore, any turbulence in the gas motion should be avoided; even if the gas velocity is high, the characteristic size of the orifice should ensure that the gas motion is laminar (similar to the boundary layers formed on the jet and on the inner surface of the nozzle or hole).

STABLE CAPILLARY MICROJET Figure 4 illustrates the interaction of a liquid and a gas to form atomizate using the method of the invention. The feeding needle 60 has a circular exit opening 61 with an internal radius I which feeds a liquid 62 out of the end, forming a drop with a radius in the range of R to plus the thickness of the wall of the needle. Thereafter, the drop narrows in circumference to a much smaller circumference as is shown in the expanded view of the tube feeding needle) 5 as shown in Figures 1 and 4. The exiting liquid flow comprises an infinite amount of liquid streamlines 63 that after interaction of the liquid with the surrounding gas to form a stable cusp at the interface 64 of the two fluids. The surrounding gas also forms an infinite number of gas streamlines 65, which interact with the solid surfaces and the exiting liquid to create the effect of a virtual focusing funnel 66. The exiting liquid is focused by the focusing funnel 66 resulting in a stable capillary microjet 67, which remains stable until it exits the opening 68 of the pressure chamber 69. After exiting the pressme chamber, the microjet begins to break-up, forming monodispersed particles The gas flow, which affects the liquid withdrawal and its subsequent acceleration after thejet is formed, should be very rapid hut also uniform in order to avoid perturbing the fragile capillary interface (the surface of the drop that emerges from the jet).

As illustrated in Figure 4, the exit opening 61 of the capillary tube 60 is positioned close to an exit opening 68 in a planar surface of a pressure chamber 69. The exit opening 68 has a minimum diameter D 0 and is in aplanar member with a thickmess e. The diameter D 0 is referred to as a minimum diameter because the opening may have a conical configuration with the narrower end of the cone positioned closer to the source of liquid flow. Thus, the exit opening may be a funnel-shaped nozzle although other opening configurations are also possible, e.g. an hour glass configuration. Gas in the pressure chamber continuously flows out of the exit opening. The flow of the gas causes the liquid drop expelled from the tube to -14- &It/ Yq Yf. ±O.Ou FiA vi. 4 DUY O~UU F.P. flA W %IV. iLFdecrease in circumference as the liquid moves away from the end of the tube in a direction toward the exit opening of the pressure chamber.

In actual use, it can be understood that the opening shape which provokes maximum gas acceleration (and consequently the most stable cusp and microjet with a given set of parameters) is a conically shaped opening in the pressure chamber. The conical opening is positioned with its narrower end toward the source of liquid flow.

The distance between the end 61. of the tube 60 and the beginning of the exit opening 68 is H. At this point it is noted that RI, Do. H and e are all preferably on the order of hundreds of microns. For example, R 4 00wm, Do l50/m, H Imm, e =-300/1m.

However, each could be 1/100 to lOx these sizes.

The end of the liquid stream develops a cusp-like shape at a critical distance from the exit opening 68 in the pressure chamber 69 when the applied pressure drop AP, through the exit opening 68 overcomes the liquid-gas surface tension stresses y/R" appearing at the point of maximum curvature e.g. from the exit opening.

A steady state is then established if the liquid flow rate Q ejected from the drop cusp is steadily supplied from the capillary tube. This is the stable capillary cusp which is an essential characteristic of the invention neaeded to form the stable microjet More particularly, a steady, thin liquid jet with a typical diameter di is smoothly emitted from the stable cusplike drop shape and this thin liquid jet extends over a distance in the range of microns to millimeters. The length of the stable microjet will vary from very short 1 micron) to very long 50 mm) with the length depending on the flow-rate of the liquid and the Reynolds number of the gas stream flowing out of the exit opening of the pressure chamber.

The liquid jet is the stable capillary microjet obtaiaed when supercritical flow is reached. This jet demonstrates a robust behavior provided that the pressure drop AP, applied to the gas -is sufficiently large compared to the maximum surface tension stress (on the order of that act at the liquid-gas interface. The jet has a slightly parabolic axial velocity profile which is, in large part, responsible for the stability of the microjet. The stable microjet is formed without the need for other forces, i.e. without adding force such as electrical forces on a charged fluid. However, for some applications it is preferable to add charge to particles, e.g.

to cause the particles to adhere to a given surface. The shaping of liquid exiting the capillary tube by the gas flow forming a focusing funnel creates a cusp-like meniscus resulting in the stable microjet This is a fundamental characteristic of the invention.

I&'II U't VI .L.O rAM VIU 4I& VQU.LIJ Qf6V. VI~.U r.o. nl%,z M W--J V,6 0 The stable capillary microjet is maintained stably for a significant distance in the direction of flow away froDm the exit from the tube. The liquid is, at this point, undergoing "supereritical flow." The microjet eventually destabilizes due to the effect of surface tension forces. Destabilization results from small natural perturbations moving dowstream, with the fastest growing perturbations being those which govern the break up of the rnicrojet, eventually creating a uniform sized monodisperse aerosol 70 as shown in Figure 4.

The microjet, even as it iniially destabilizes, passes out of the exit orifice of the pressure chamber without touching the peripheral surface of the exit opening. This provides an important advantage of the invention which is that the exit opeging 68 (which could be referred to as a nozzle) will not clog from residue anid/or deposits of the liquid. Clogging is a major problem with very small nozzles and is generally dealt with by cleaning or replacing the nozzle. When fluid contacts the sm-faces of a nozzle opening some fluid will remain in contact with the nozzle when the flow of fluid is shut off. The liquid remaining on the nozzle surface evaporaTes leaving a residue. After many uses over time the residue builds up and clogging takes place. The present invention substantially reduces or eliminates this clogging problem.

MATHEMATICS OF A STABLE MICROJET Cylindrical coordinates (rz) are chosen for analyzing the shape of a stable microjet, i.e. a liquid jet undergoing "supercritical flow." The cusp-like meniscus formed by the liquid coming out of the tube is pulled toward the exit of the pressure chamber by a pressure gradient created by the flow of gas.

The cusp-like meniscus formed at the tube's mouth is pulled towards the hole by the pressure gradient created by the gas stream- From the cusp of this mneniscus, a steady liquid thread with the shape of radius r is withdfrawn through the hole by the action of both the suction effect due to AP., and the tangential viscous siresses -r exerted by the gas on the jefs surface in the axial direction. The averaged momentum equation for this configurtion may be written: d 2 fP 2 T -16- 24/04 '02 15:30 FAX 61 2 9810 8200 F.B. KiL. U W- v where Q is the liquid flow rate upon exiting the feeding tube, P, is the liquid pressure, and p, is the liquid density, assuming that the viscous extensional term is negligible compared to the kinetic energy term, as will be subsequently justified. In addition, liquid evaporation effects are neglected. The liquid pressure P, is given by the capillary equation.

PI Pg y (2) where y is the liquid-gas surface tension. As shown in the Examples, the pressure drop AP. is sufficiently large as compared to the surface tension stress to justify neglecting the latter in the analysis. This scenario holds for the whole range of flow rates in which the microjet is absolutely stable. In fact, it will be shown that, for a given pressure drop AP, the minimum liquid flow rate that can be sprayed in steady jet conditions is achieved when the surface tension stress y/ is of the order of the kinetic energy of the liquid pIQ'f(2HV), since the isurface tension acts like a "resistance" to the motion (it appears as a negative term in the right-hand side term of Eq. Thus, min (3) P,

I

For sufficiently large flow rates Q compared to Qj the simplified averaged momentum equation in the axial direction can be expressed as d(Pl.QZ) d,.g Z7.

I z ff iQ 24 where one can identify the two driving forces for the liquid flow on the right-hand side. This equation can be integrated provided the following simplification is made: if one uses a thin plate with thickness L of the order or smaller than the hole's diameter D (which minimizes downstream perturbations in the gas flow), the pressure gradient up to the hole exit is on the average much larger than the viscous shear term 2cr owning to the surface stress. On the other hand, the axial viscous term is of the order O[Q/D 2 dj], since the hole diameter D is -17- Z4/U4 UZ 1i:31 FAA .iUl 2 91U 5ZUU k. KIs. -u UU. ugUi: O actually the characteristic distance associated with the gas flow at the hole's entrance in both the radial and axial directions. This term is very small compared to the pressure gradient in real situations, provided that AP, A/D 2 p, (which holds, eg., for liquids with viscosities as large as 100 cpoises, using hole diameters and pressure drops as small asD 10 pm and AP 100 mbar). The neglect of all viscous terms in Eq. is then justified. Notice that in this limit on the liquid flow is quasi.isentropic in the average (the liquid almost follows Bernoulli equation) as opposed to most nmicromaetric extensional flows. Thus, integrating from the stagnation regions of both fluids up to the exit, one obtains a simple and universal expression for thejet diameter at the hole exit: d. 8p V4 P 1 Q% s J nQ (5) which for a given pressure drop AP is independent of geometrical parameters (hole and tube diameters, tube-hole distance, etc.), liquid and gas viscosities, and liquid-gas surface tension.

This diameter remains almost constant up to the breakup point since the gas pressure after the exit remains constant MONODISPERSE

PARTICLES

Above the stable microjet undergoing "supercritical flow" is described and it can be seen how this aspect of the invention can be made use of in a variety of industrial applications particularly where the flow of liquid through small holes creates a clogging problem. An equally important aspect of the invention is obtained after the microjet leaves the pressure chamber.

When the microjet exits the pressure chamber the liquid pressure P becomes (like the gas pressure P) almost constant in the axial direction, and the jet diameter remains almost constant up to the point where it breaks up by capillary instability. Defining a Weber number We (pv2d)/Y 2 AP./y (where v, is the gas velocity measured at the orifice), below a certain experimental value We,- 40 the breakup mode is axisymmetric and the resulting droplet stream is characterized by its monodispersity provided that the fluctuations of the gas flow do not contribute to droplet coalescence (these fluctuations occur when the gas stream reaches a folly developed turbulent profile around the liquid jet breakup region). Above this -18- 24/04 '02 15:31 FAX 61 2 9810 8200 F.B, xlt.; kou.

We, value, sinuous nonaxisymmetric disturbances, coupled to the axisymmetric ones, become apparent. For larger We numbers, the nonlinear growth rate of the sinuous disturbances seems to overcome that of the axisymmetic disturbances. The resulting spray shows significant polydispersity in this case. Thus, it can be seen that by controlling parameters to keep the resulting Weber number to 40 or less, allows the particles formed to be all substantially the same size. The size variation is about to *30% and move preferably to These particles can have a desired size e.g. 0.1 microns to 50 microns.

The shed vorticity influences the breakup of the jet and thus the formation of the particles. Upstream from the hole exit, in the accelerating region, the gas stream is laminar.

Typical values of the Reynolds number range from 500 to 6000 if a velocity of the order of the speed of sound is taken as characteristic of the velocity of the gas. Downstream from the hole exit, the cylindrical mixing layer between the gas stream and the stagnant gas becomes unstable by the classical Kelvin-Helmholtz instability. The growth rate of the thickness of this layer depends on the Reynolds number of the flow and ring vortices are formed at a frequency of the order of vJD, where D is the hole diameter. Typical values of vg and D as those found in our experimental technique lead to frequencies or the order of MNIz which are comparable to the frequency of drop production (of order of tj 1 Given the liquid flow rate and the hole diameter, a resonance frequency which depends on the gas velocity (or pressure difference driving the gas stream) can be adjusted (tuned) in such a way that vortices act as a forcing system to excite perturbations of a determined wavelength on the jet surface. Experimental results obtained clearly illustrates the different degree of coupling between the two gas-liquid coaxial jets. In one set of experimental results the particle sizes are shown to have a particle size of about 5.7 microns with a standard deviation of 12%. This results when the velocity of the gas has been properly tuned to minimize the dispersion in the size of droplets resulting from the jet breakup. In this case, the flow rate of the liquid jet and its diameter are 0.08ul s' and 3 gm, respectively. Data have been collected using a MASTERSIZER from MALVERN Instruments. As the degree of coupling decreases, perturbations at thejet surface of different wavelengths become excited and, as it can be observed from the size distributions, the dispersion of the spray increases.

It is highly desirable in a number of different industrial applications to have particles which are uniform in size or to create aerosols of liquid particles which are uniform in size. For example, particles of a liquid formation containing a pharmaceutically active drug -19ciiu'a uc 10;4. tAA 01 Z uOIU OVu r.n. ai c IJ fji£ I eduld be created and designed to have a diameter of about 2 microns These particles could be inhaled into the lungs of a patient for intrapulmonary drug delivery. Moreover, particle size can be adjusted to target a particular area of the respiratory tract.

Upstreamn of the orifice exit the gas flow should be laminar in order to avoid a turbulent regime turbulent fluctuations in the gas flow which have a high frequency and would pertuarb the liquid-gas interface. The Reynolds numbers reached at the orifice are

V

91 where is the kinematic viscosity of the gas. Even though this number is quite high, there are large pressure gradients downstream (a highly convergent geometry), so that a turbulent regime is very unlikely to develop.

The essential difference from existing pneumatic atomizers (which possess large Weber numbers) and the present invention is that the aim of the present invention is not to rupture tbe liquid-gas interface but the opposite, i.e. to increase dhe stability of the interface until a capillary jet is obtained- Thej et, which will be very thin provided the pressure drop resulting from withdrawal is high enough, splits into drops the sizes of which are much more uniform than those resulting from disorderly breakage of the liquid-gas interface in existing pneumatic atomizers.

The proposed atomization system obviously requires delivery of the liquid to be atomized and the gas to be used in the resulting spray. Both should be fed at a rate ensuring that the system lies within the stable parameter window. Multiplexing is effective when the flow-rates needed exceed those on an individual cell. More specifically, a plurality of feeding sources or feedng needles may be used to increase the rate at -which aerosols are created. The flow-rates used should also ensure the mnas ratio between the flows is compatible with the specifications of each applieatioa- The gas and liquid can be dispensed by any type of continuous delivery system (e.g.

a compressor or a pressurized tank the former and a volumetric pump or a pressurized bottle the latter). If multiplexing is needed, the liquid flow-rate should be as uniform as possible among cells; this may entail propulsion through several capillary needles, porous media or any other medium capable of distributing a uniform flow among different feeding points.

24/04'02 15:32 FAX 61 2 9810 8200 KI,, C Lu.

Each individual atomization device should consist of a feeding point (a capillary needle, a point with an open microchannel, a microprotuberance on a continuous edge, etc.) 0,002-2 mm (but, preferentially 0.01-0.4 mm) in diameter, where the drop emerging from the microjet can be anchored, and a small orifice 0.002-2 mm (preferentially 0.01-0.25 mm) in diameter facing the drop and separated 0.01-2 mm (preferentially 0.2-0.5 mm) from the feeding point The orifice communicates the withdrawal gas around the drop, at an increased pressure, with the zone where the atomizate is produced, at a decreased pressure. The atomizer can be made from a variety of materials (metal, polymers, ceramics, glass).

Figure 1 depicts a tested prototype where the liquid to be atomized is inserted through one end of the system 2 and the propelling gas in introduced via the special inlet 4 in the pressure chamber 3. The prototype was tested at gas feeding rates from 100 to 2000 mBar above the atmospheric pressure P. at which the atomized liquid was discharged. The whole enclosure around the feeding needle I was at a pressure PO P, The liquid feeding pressure, PI, should always be slightly higher than the gas propelling pressure, P. Depending on the pressure drop in the needle and the liquid feeding system, the pressure difference (PI PO 0) and the flow-rate of the liquid to be atomized, Q, are linearly related provided the flow is laminar which is indeed the case with this prototype The critical dimensions are the distance from the needle to the plate the needle diameter the diameter of the orifice through which the microjet 6 is discharged (do) and the axial length, e, of the orifice the thickness of the plate where the orifice is made). In this prototype, Hwas varied from 0.3 to 0.7 mm on constancy of the distances (D 0 0.45 mm, d, 0.2 mm) and e 0.5 mm. The quality of the resulting spray 7 did not vary appreciably with changes in) hprovided the operating regime stationary drop and microjet) was maintained. However, the system stability suffered at the longer H distances (about 0.7 mm). The other atomizer dimensions had no effect on the spray or the prototype functioning provided the zone around the needle (its diameter) was large enough relative to the feeding needle.

WEBER NUMBER Adjusting parameters to obtain a stable capillary microjer and control its breakup into monodisperse particle is governed by the Weber number and the liquid-to-gas velocity ratio or a which equal VIII. The Weber number or "Wen is defined by the following equation: -21laq/u'i U4 AO- rAA 01 9 VOIV OLIU V

U£

r.D. 06 VV. LIJ-J V L 0 p 2 d

Y

wherein p. is the density of the gas, d is the diameter of the stable microjet, Y is the liquidgas surface tension, and I'2 is the velocity of the gas squared.

When carrying out the irvention the Parameters should be adjusted so that the Weber nmb11er is greater than 1 in order to produce a stable cap illary microjeL However, to obtain a particle dispersion which is uionodizerse =11h Particle has; the same size to 30 the parameters should be adjusted so that the Weber number is less than about The monodisperse aerosol is obtained with a Weber number in a range of abou t I to about (1 We g OUNESORGE

NUMER

A measure of the relative importanc Of viscosity on the jet breakup can be estimated frona the Olinesorge uumber define as the ratio between two characteristic times; the viscous tizue x, aiad the breaking time The breaking time t, is give. by [see Rayleigh (1878)] -b (2) Perturbations on the jet surface are propagated inside by viscous diffusion in times t, of the order of v- p~d 2 pl (3) where /p 1 is the viscosity of the liquid. Then, the Ohnesorgc number, Oh, results -22- Z41U4 Ui ia:jz ?AA 01 Z V5IU OZUU t.D. AJKIL 4x %ii. WJ YOU If this ratio is much smaller than unity viscosity plays no essential role in the phenomenon under considera ion. Since the maximum value of the Oltuesorge number in actual experiments conducted is as low as 3.7x 10 2 viscosity plays no essential role during the process of jet breakup.

EMBODIMENT OF FIGURE 2 A variety of configurations of components and types of fluids will become apparent to those skilled in the art upon reading this disclosure. These configurations and fluids are encompassed by the present invention provided they can produce a stable capillary microjet of a first fluid from a source to an exit port of a pressure chamber containing a second fluid. The stable microjet is formed by the first fluid flowing from the feeding source to the exit port of the pressure chamber being accelerated and stabilized by tangential viscous stress exerted by the second fluid in the pressure chamber on the surface of the frst fluid forming the microjet.

The second fluid forms a focusing funnel when a variety of parameters are correctly tuned or adjusted. For example, the speed, pressure, viscosity and miscibility of the first and second fluids are chosen to obtain the desired results of a stable microjet of the first fluid focused into the center of a funnel formed with the second fluid. These results are also obtained by adjusting or tuning physical parameters of the device, including the size of the opening from which the first fluid flows, the size of the opening from which both fluids exi, and the distance between these two openings.

The embodiment of Figure 1 can, itself, be arranged in a variety of configurations.

Further, as indicated above, the embodiment may include a plurality of feeding needles. A plurality of feeding needles may be configured concentrically in a single construct, as shown in Figure 2.

The components of the embodiment of Figure 2 are as follows: 21. Feeding needle tube or source of fluid.

22. End of the feeding needle used to insert the liquids to be atomized.

23. Pressure chamber.

24. Orifice used as gas inlet 25. End of the feeding needle used to evacuate the liquid to be atomized.

26. Orifice through which withdrawal takes place.

27. Atonizate (spray) or aerosol.

-23 Lqlk V4 LU;00 rIL& UL OIV 04VU r.D. ^Lvru m 4VJ v o 28. First liquid to be atomized (inner core of particle).

29. Second liquid to be atomized (outer coating of particle).

Gas for creation of microjet.

31. Internal tube of feeding needle.

32. External tube of feeding needle.

D diameter of the feeding needle; d= diameter of the orifice through which the microjet is passed; e axial length of the orifice through which withdrawal takes place; H distance from the feeding needle to the microjet outlet; y=surface tension; Po= pressure inside the chamber; P, atmospheric pressure.

The embodiment of Figure 2 is preferably used when attempting to form a spherical particle of one substance coated by another substance. The device of Figure 2 is comprised of the same basic component as per the device of Figure 1 and further includes a second feeding source 32 which is positioned concentrically around the first cylindrical feeding source 31.

The second feeding-source may be surrounded by one or more additional feeding sources with each concentrically positioned around the preceding source. The outer coating may be used for a variety of purposes, including: coating particles to prevent small particles from sticking together, to obtain a sustained release effect of the active compound a pharmaceutically active drug) inside, and/or to mask flavors; and to protect the stability of another compound a pharmaceutically active drug) contained therein.

The process is based on the microsuction which the liquid-gas or liquid-liquid interphase undergoes (if both are immiscible), when said interphase approaches a point beginning from which one of the fluids is suctioned off while the combined suction of the two fluids is produced. The interaction causes the fluid physically surrounded by the other to form a capillary microjet which finally breaks into spherical drops. If instead of two fluids (gas-liquid), three or more are used that flow in a concentric manner by injection using concentric tubes, a capillary jet composed of two or more layers of different fluids is formed which, when it breaks, gives rise to the formation of spheres composed of several approximately concentric spherical layers of different fluids. The size ofthe outer sphere (its thickness) and the size of the inner sphere (its volume) can be precisely adjusted. This can allow the manufacture of coated particles for a variety of end uses. For example the thickness -24- ILfU'i uz ia;oj, Di z voiv ozu Ivv r.D. A111Z IV, W--J V 0 of 'the coating can be varied in different muanufacturing events to obtain coated particles which have gradually decreasing thicknesses to obtain a controlled release effect of the contents, e.g.

a pharmaceutically active drug. The coating could merely prevent the particles from degrading reacting or sticking together.

The method is based on the breaking of a capillary microjet composed of a nucleus of one liquid or gas and surrounded by another or other liquids and gases which are in a concentric manner injected by a special injection head, in such a way that they formu a stable capillary ruicrojet and that they do not mix by diffusion during the time between when the microjet is formed and when it is broken. When the capillary microjet is broken into spherical drops under the proper operating conditions, which will be described in detail below, these drops exhibit a spherical nucleus, the size and eccentricity of which can be controlled.

In the case of spheres containing two materials, the injection head 25 consists of two concentric tubes with an external diameter on the order of one millimeter Through the internal tube 311is injected the material that will constitute the nucleus of the microsphere, while between the internal tube 31 and the external tube 32 the coating is injected. The fluid of the external tabe 32 joins with the fluid of tube 31 as the fluids exit the feeding needle, and the fluids (normally liquids) thus injected are accelerated by a stream. of gas that passes through a small orifice 24 facing the end of the injection tubes. When the drop in pressure across the orifice 24 is sufficient, the liquids form a completely stationary capillary niicrojet, if the quantities of liquids that are injected are stationary. This microjet does not touch the walls of the orifice, but panses through it wrapped in the stream of gas or funnel formed by gas from the tube 32. Because the funnel of gas focuses the liquid, the size of the eit orifice 26 does not dictate the size of the particles formed.

When the parameters are correctly adjusted, the movement of the liquid is uniform at the exit of the orifice 26 and the viscosity forces are sufficiently small so as not to alter either the flow or the properties of the liquids; for example, if there are biochemical molecular specimens having a certain complexity and fragility, the viscous forces that would appear in association with the flow through a micro-orifice might degrade these substances.

Figure 2 shows a simplified diagram of the feeding needle 21, which is comprised of the concentric tubes 30, 31 through the internal and external flows of the fluids 28, 29 that are going to compose the microspheres comprised of two immiscible fluids. The difference in pressures PO P. PJ, through the orifice 26 establishes a flow of gas present in the zqiuq uz ia;JJ rAA oi z voU ovur r.D. AiZ M %U. I JV U-J chamber 23 and which is going to surround the microjet at its exit The same pressure gradient that moves the gas is the one that moves the microjet in an axial direction through the hole 26, provided that the difference in pressures Po P, is sufficiently great in comparison with the forces of surface tension, which create an adverse gradient in the direction of the movement There are two limitations for the minimum sizes of the inside and outside jets that are dependent on the surface tensions yl of the outside liquid 29 with the gas 30 and 2 of the outside liquid 29 with the inside liquid 28, and on the difference in pressures' P P0 P, through the orifice 26. In the first place, the jump in pressures AP must be sufficiently great so that the adverse effects of the surface tension are minimized. This, however, is attained for very modest pressure increases: for example, for a 10 micron jet of a liquid having a surface tension of 0.05 N/m (tap water), the necessary minimum jump in pressure is in the order of 0.05 0.00001 m API 50 mBar. But, in addition, the breakage of the microjet must be regular and axilsymmetric, so that the drops will have a uniform size, while the extra pressure AP cannot be greater than a certain value that is dependent on the surface tension of the outside liquid with the gas yl and on the outside diameter of the microjet. It has been experimentally shown that this difference in pressures cannot be greater than 20 times the surface tension yl divided by the outside radius of the microjet.

Therefore, given some inside and outside diameters of the microjet, there is a range of operating pressures between a minimum and a maximum; nonetheless, experimentally the best results are obtained for pressures in the order of two to three times the minimum.

The viscosity values of the liquids must be such that the liquid with the greater viscosity verifies, for a diameter d of the jet predicted for this liquid and a difference through the orifice AP, the inequality:

Q

With this, the pressure gradients can overcome the extensional forces of viscous resistance exerted by the liquid when it is suctioned toward the orifice.

Moreover, the liquids must have very similar densities in order to achieve the concentricity of the nucleus of the microsphere, since the relation of velocities between the liquids moves according to the square root of the densities vl/v2 (p2/pl) and both jets, the -26- £'Ait,'a u o raA 01 z vo1Uv OSuU r.n. nI In %IV. LI&J I JO't iinide jet and the oursidejer, must assume the most symmetrical configuration possible which does not occur if the liquids have different velocities (Figure Nonetheless, it has been experimentally demonstrated that, on account of the surface tension y~2 between the two liquids, the nucleus tends to migrate toward the center of the microsphere, within prescribed parameters.

When two liquids and gas are used on the outside, the distance between the planes of the mouths of the concentric tubes can vary, without the characteristics of the jet being sub stantilly altered, provided That the int~nal tube 31 is not introduced into the exte rnal one 32 more than one diameter of the external tube 32 and provided that the internal tube 31 does not project wore than two diameters from the external tube 32. The best resulrs are obtained when the inaternal tube 31 projects from the external one 32 a distance substantially the same as the diameter of the internal tube 3 1. This same criterion is valid if more than two tubes are used, with the tube that is surrounded (inner tube) projecting beyond the tube that surrounds (outer tube) by a distance substantially the same as the diameter of the first tube.

The distance between the plane of the internal tube 31 (the one that will normally project more) and the plane of the orifice may vary between zero and three outside diameters of the external tube 32, depending on the surface tensions between the liquids and with the gas, and on their viscosity values- Typically, the optimal distance is found experimentally for each particular configuration and each set of liquids used.

The proposed atomizing system obviously requires fluids that are going to be used in the resulting spray to have certain flow parameters. Accordingly, flows for this use mnust be: -Flows that are: suitable so that the system falls within the parametric window of stability. Multiplexing several sets of concentric tubes) may be used, if the flows required are greater than those of an individual cell.

-Flows that are suitable so that the mass relation of the fluids fails within the specifications of each application. Of course, a greater flow of gas may be supplied externally by any means in specific applications, since this does not interfere with the functioning of the atomizer.

Therefore, any means for continuous supply of gas (compressors, pressur deposits, etc.) and of liquid (volumetric pumps, pressure bottles) may be used, If multplexing is desired, the flow of liquid must be as homogeneous as possible between the -27 £I/1t44 uc ±a~oi 01 4 volu ohUU r.n. %1,Z CC I4J 1)00 vaious cells, which may require impulse trough multiple capillary needles, porous media, or any other medium capable of distributing a homogeneous flow among different feeding points.

Each atomizing device will consist of concentric tubes 3 1 32 with a diameter ranging between 0.05 and 2 mm, preferably between 0. 1 and 0.4 mmn, on which the drop from -which the microjet emanates can be anchored, and a small orifice (between 0.00 1 and 2 mm. in diameter, preferably between 0. 1 and 0.25 mm), facing the drop and separated from the point of feeding by a distance between 0.00! and 2 mm, preferably between 0-2 and mm. The orifice puts the suction gas that surrounds the drop, ax higher pressure, in touch wvith the area in which the atomizing is to be attained, at lower pressure.

EMBODIENT OF FIGURE 3 The embodiments of Figures I and 2 are similar in a number of ways. Both have a feeding piece which is preferably in the form of a feeding needle with a circular exit opening. Further, both have an exit port in the pressue chamber which is positioned directly in front of the flow path of fluid out of the feeding source. Precisely maintaining the alignment of the flow path of the feeding source with the exit port of the pressure chamber can present an engineering challenge particularly when the device includes a number of feeding needles. The embodiment of Figure 3 is designed to simplify the manner in which components are aligned. The embodiment of Figure 3 uses a planar feeding piece, which by virtue of the withdrawal effect produced by the pressure difference across a smallI opening through which fluid is passed permits multiple microjets to be expelled through multiple exit ports of a pressure chamber thereby obtaining multiple aerosol streams.

Although a single planar feeding member is shown in Figur 3 it, of course, is possible to produce a device with a plurality of planar feeding members where each planar feeding member feeds fluid to a linear arry of outlet orifices in the surrounding pressure chamber.

In addition, the feeding member need not be strictly planar, and may be a curved feeding device comprised of two surfaces that maintain approximately the same spatial distance between the two pieces of the feeding source. Such arved devices may have any level of curvature, e.g. circular, semicircular, elliptical, herni-elliptical etc.

The components of the embodiment of Figure 3 are as follows: Feeding piece.

-28- ,G Ivi raA ux voxv otVV r.V. nD v M Iv. 4J UtuU liquid second

H=

chamber, flow rate 42. End of the feeding piece used to insert the fluid to be atomized.

43. Pressure chamber.

44. Orifice used as gas inlet End of the feeding needle used to evacuate the liquid to be atomized.

46. Orifices through which withdrawal takes place.

47. Atomizate (spray) or aerosol.

48. first fluid containing material to be atomized.

49. second fluid for creation ofmicojet wall of the propulsion chamber facing the edge of the feeding piece.

51. channels for guidance of fluid through feeding piece.

d- diameter of the microjet formed; p= liquid density of first fluid pu= density of second fluid velocity of the first liquid v,-velocity of the liquid e axial length of the orifice through which withdrawal takes place; distance from the feeding needle to the microjet outlet; P pressure inside the Ap,= change in pressure of the gas; P, atmospheric pressure; Q=volumetric The proposed dispersing device consists of a feeding piece 41 which creates a planar feeding channel through which a where a first fluid 48 flows, The flow is preferably directed through one or more channels of uniform bores that are constructed on the planar surface of the feeding piece 41. A pressure chamber 43 that holds the propelling flow of a second liquid 49, houses the feeding piece 41 and is under a pressure above maintained outside the chamber wall 50. One or more orifices, openings or slots (outlets) 46 made in the wall 52 of the propulsion chamber face the edge of the feeding piece. Preferably, each bore or channel of the feeding piece 41 has its flow path substantially aligned with an outlet 46.

Formation of the microjet and its acceleration are based on the abrupt pressure drop resulting from the steep acceleration undergone by the second fluid 49 on passing through the orifice 46, similarly to the procedure described above for embodiments of Figures 1 and 2 when the second fluid 49 is a gas.

When the second fluid 49 is a gas and the first fluid 48 is a liquid, the microthread formed is quite long and the liquid velocity is much smaller than the gas -29- 24/0.4 '02 15:34 FAX 61 2 9810 8200 F.B. RICE CO.

velocity. In fact, the low viscosity of the gas allows the liquid to flow at a much lower velocity; as a result, the microjet is actually produced and accelerated by stress forces normal to the liquid surface, i.e. pressure forces. Hence, one effective approximation to the phenomenon is to assume that the pressure difference established will result in the same kinetic energy per unit volume for both fluids (liquid and gas), provided gas compressibility effects are neglected. The diameter d of the microjet formed from a liquid density p that passes at a volumetric flow-rate Q through an orifice across which a pressure difference AP. exists will be given by d. 8p, o, See GaUn-Calvo, Physical Review Letters, 80:285-288 (1998).

The relation between the diameter of the microjet, d, and that of the resulting drops, J, depends on the ratio between viscous forccs and surface tension forces on the liquid on the one hand, and between dynamic forces and surface tension forces on the gas on the other on the Ohnesorge and Weber numbers, respectively) (Hinds (Aerosol Technology, John Sons, 1982), Lefevre (Atomization and Sprays, Hemisphere Pub.

Corp., 1989) and Bayvel Orzechowski (Liquid Atomzation, Taylor Francis, 1993)).

At moderate to low gas velocities and low viscosities the relation is roughly identical with that for capillarity instability developed by Rayleigh: d 1.89d) Because the liquid microjet is very long, at high liquid flow-rates the theoretical rupture point lies in the turbulent zone created by the gas jet, so turbulent fluctuations in the gas desmbilize or rupture the liquid microjet in a more or less uneven manner. As a result, the benefits of drop size uniformity are lost On the other hand, when the second fluid 49 is a liquid and the first fluid 48 is a gas, the facts that the liquid is much more viscous and that the gas is much less dense provide that the gas microthread formed is much shorter; however, because its rupture zone is almost invariably located in a laminar flowing stream, dispersion in the size of the tJ U ot Z4/U4 'UZ 1b:3b A Ul *Z V1U 8ZUU KHit, U. W. uVo microbubbles formed is almost always small. At a volumetric gas flow-rate Q, and a liquid overpressure AP, the diameter of the gas muicrojet is given by: J T2AP, 9 where p, is the gas density. The low liquid velocity and the absence of relative velocities between the liquid and gas lead to the Rayleigh relation between the diameters of the microthread and those of the bubbles d 1.89d).

If both fluids 48, 49 are liquid and scarcely viscous, then their relative velocities will be given by their densities ratio: VA

LP)J

The diameter of a microjet of the first liquid at a volumetric flow-rate of Q and an overpressure of AP, will be given by d 8 P A- 1%Q

TT

2 PB

A

B At viscosities such that the velocities of both fluids 48, 49 will rapidly equilibrate in the microjet, the diameter of the microjet of the first liquid will be given by d 8p 8 I rT 2 Apa I The proposed atomization system obviously requires delivery of the fluids 48,49 to be used in the dispersion process at appropriate flow-rates. Thus: Both flow-rates should be adjusted for the system so that they lie within the stable parameter window.

-31- L4144 UZ 10; )a rAA 01 Z Volu 0xUV r.D. AIIZ M 'MJ V -j 0P The mass ratio between the flow-s should be compatible with the specifications of each application. Obviously, the gas flow-rate can be increased by using an external means in special applications burning, drug inhaation) since this need not interfere with the atomizer operation.

Therefore, the gas and liquid can be dispensed by any type Of continuous delivery systein a compressor or a pressurized tenk the former-and a volumetric pump or a pressurized bottle the latter).

The Atomizer can be made from a variety of materials (metal, plastic, ceramics, glass).

BUBBLES IN~TO LIQUID OR GAS Figures 6 and 7 are useful in showing how bubbles may be formed in either a liquid (Figure 6) or a gas (Figure In Figure 6 a tubular feeding source 71 is continually supplied with a flow of gas which forms a stable cusp 72 which is surrounded by the flow of liquid 73 in the pressure chamber 74 which is continually supplied with a flow of liquid 73. The liquid 73 flows out of the chamber 74 into a liquid 75 which may be the same as or different from the liquid 73.

The cusp 72 of 8as narrows to a capillary supercritical flow 76 and then enter the exit opening 77 of the chamber 74. At a point 78 in the exit opening 77 the supercritical flow 76 begins to destabilize but remains as a critical capillary flow until leaving the exit opening 77. Upon leaving the exit opening 77 the gas steamn breaks apart and forms bubbles 79 each of which are substaatially identical to the others in shape and size. The uniformity of bubbles is such that one bubble differs from another (in termis of measured physical diameter) in an amount in a range of standard deviation of to *30% with a preferred deviation being less than Thus, the uniformity in size of the bubbles is greater than the uniformity of the particles formed as described above in connection with Figure 1 when liquid particles are formed.* -32 10-.40o rAA 01 9 VOIV O9UU jU'&L r.D. nllz M %IV. WJ VV Gas in the bubbles 79 will diffuse into tbe liquid 75. Smaller bubbles provide for greater surface area contact with the liquid 75. Smaller bubbles provide for greater surface area contact with the liquid 75 thereby allowing for a faster rate of diffusion then would occur if the same volume of gas were present in a smaller number of bubbles. For example, ten bubbles each containing I cubic mam of gas would diffluse gas into the liquid much more rapidly than one bubble containing 10 cubic min of gas. Further, smaller bubbles rise to the liquid surface more slowly than larger bubbles- A slower rate of ascent in the liquid means that the gas bubbles are in contact with the liquid for a longer period of time thereby increasing the amount of diffusion of gas into the liquid. Thus, smaller bubbles could allow a greater amount of a pharmaceutically active drug to. diffuse into a carrier. Becamse the bubbles are so uniform in size the amount of gas diffusing into the liquid can be uniformly calculated which is i'nporant to obtain a uniform concentration of drug in a carier, Figure 7 shows the same components as shown in Figure 6 except that the liquid is replaced with a gas 80. When the stream of bubbles 79 disassociate the liquid 73 forms an outer spherical cover thereby providing hollow droplets 81 'which will float in the gas 80. The hollow droplets 81 have a large physical or actual diameter relative to their aerodynamic diameter. Hollow droplets fall in air at a much slower rate compared to liquid droplets of the same diameter. Because the hollow droplets 81 do not settle or fall quicly in air they can be inhaled into the lungs. Eventually the hollow droplets 8 1 will burst and form many smaller particles which can be drawn even deeper into the lunogs. Thus, it is understood that the aerodynamic diameter of the hollow droplets is very small compared to their actual physical diameter. The creation of hollow droplets 81 which burst and form very small particles is applicable to obtain deeper delivery of drug into the lungs.

EMULSIONS

Figure 8issimilartoFigs6 2m7- However, rathertbanagas72asnFigure 6 the feeding source 71 provides a steam of liquid 82 which may be miscible but is preferably immiscible in the liquid 73. Further the liquid 73 may be the same as or different from the liquid 75 but is preferably immiscible in the liquid 75. The creation of emulsions using such a configuration of liquids has applicrability in a variety of fields particularly because the liquid particles formed can have a si=e in the range of from about 1 to about 200 microns with a standard deviation in size of one particle to another being as little as 0-01%.

-33- Z4/q4 VZ 10:JU rAA 01 Z V01V OZVV r.D. ni%,n 1,v.

igj V '4 1 The size deviation Of One particle to another can vary up to about 30% and is preferably less than and more preferably less than =41%.

The system operates to expel the liquid 82 out of the exit orifice 77 to form spheres 83 of liquid 82. Each sphere 83 has an actual physical diameter which deviates from other spheres 83 by a standard deviation of ±0.01% to preferably 10% or less and more preferably I1% or less. The size of the spheres 83 and flow rate of liquid 82 is controlled so that each sphere 83 contain a single particle (etg. a single cell) to be examined.

The stream of spheres 83 is caused to flow past a sensor and/or energy source of any desired type thereby allowing for a deter-milnation of precisely how much drug a patient is inhaling.

EMULSIONS

In Figure 8 the liquid 75 can be a gas air) inside a patient's respiratory system. The liquid 82 could be water or a drug which is surrounded by a second liquid 73 which is a water or surfactant. The system then forms particles 83 which have a center of one liquid and an outer coating of a different liquid.

DRUG DELIVERY

DEVICES

The various embodiment& of components for creating aerosols described above can be used in devices for the delivery of an aerosol to at patient. The device is preferably a hand-held, self-contained device which a patient can easily carry about and use for the administration of drugs.

In one embodiment the pressurized canister of a device as disclosed in US. Patent 5,364,838 or 5,404,871 is replaced with an aerosol generating device of the type shown here in any of Figures 1, 2 or 3. The device preferably includes a means for measuring a patients respiratory flow rate and respiratory volume. The aerosol may be generated when the patient manually actuates release of aerosol. However, the device preferably operates as the device in the 5,364,.838 patent and is actuated automatically in response to a measured inspiratory flow rate and inspiratory volume. This makes it possible to repeatedly deliver aerosol to a patient at the same point in the respiratory cycle thereby improving repeatability of dosing.

In another embodiment the disposable containers and porous membranes of the device disclosed ii, U.S. Patent 5,544,646 is replaced with an aerosol generating device of -34- 44/uvl V9 .O;jD 01 4 VOIu Ouu14

V'*

r.D. Al%,Z ft %IV. W--J V 4 4.

tbd type shown here in Figures 1, 2 or 3. In such an embodiment the aerosol generating device is likely to include a Plurality of sources of liquid drug formulationa a plurality of feeding needles) so that a suffici ent amount of formulation can be aerosolized ini a sufficiently short period of time That the patient can, in a single inhalvition, inhale all of the aerosol needed for a dose of the drug being delivered. An aerosol generating device of the type shown in Figure 3 is generally easier to manufacture and as such is preferred in this embodiment.

In the two embodiments described above it is preferable to create an aerosol with particles having a particle diameter in the range of about I to about 3 microns. This allows the particles to penetrate into the smallest channels of the lungs. However, such particles are not so small that they will not settle on the surface of these channels wben inhald -failure to settle would cause the particles to be exhaled as with particles of smoke.

In yet another embodiment the source of aerosol of a device as disclosed in U.S.

Patent 4,484,577 is replaced with an aerosol generating device of the type described here and shown in Figures 1, 2 or 3. The device of U.S. Patent 4,484,577 includes a large container into which the aerosol is dispersed. The aerosol then "hangs" in the air in the large container until it is inhialed by the patient Using an aerosol generation device of the invention the aerosol particles can be made very small, e.g. about 1 micron or les. Particles of this size wvill be suspended in air in the container until inhaed. The patient can then breath normally in and out of the container until all or substantially Wll particles in the container have deposited themselves on the channels of the lung.

EXAMIPLES

The following examiples are put forth so as to provide those of ordinary skil in the arn with a complete disclosure and description of how to make and use the present invention, and are not intended to limit the scope of what the inventors regard as their invention nor are they intended to represent that the experiments below are all or the only experiments performed. Efforts have been made to ensure accuracy with respect to numbers used (e.g.

amounts, temnper-ature, etc.) but some experimental errurs and deviations should be accounted for. Unless indicated otherwise, parts are parts by weight, molecular weight is weight average molecular weight, temnperature is in degrees Centigrade, and pressure is at or near atmospheric.

zqiUki v4 ia;ov vi z voiu ozuu MUi.

r.D. Al%,r, M W--J V 'V The Properties of sixteen different liquids are provided in Table I TABlLE 1: Liquids used and some of their physical properties at 24.5 C kg/rn 3 Iz: cpoise, y: N/ni). Also given, the symbols used in the plots.

Liquid p IAy Symbol lieptane 684 0.38 0.021 0 Tap Water 1000 1.00 0.056 0 Water glycerol 90/10 V/v 1026 1.39 0.069 Water+ glycerol8 0/20ov/ 1052 1.98 0-068

V

Isopropyl alcohol 755.5 2.18 0,021 Water glycerol 70/3 0 v/v 1078 2.76 0.067 0 Water glycerol 60/40 v/v 1104 4.37 0.067 Water +glycerol 50/50ovv 1030 6.17 0.066 0 I -Octanol 827 7.47 0.024 Water glycerol 40/60 v/v 1156 12.3 0.065A Water glycerol 35/65 v/v 1167 15.9 0.064

V

Water glycerol 30/70 v/v 1182 24.3 0.064

X

Water +glycerol 25/75 v/v 1195 38.7 0.063 Propylene glycol 1026 41.8 0.036 The liquids of Table 1 were forced through a feeding needle of the type shown in Figure 1I The end 5 of the feeding needle had an internal radius The exit orifice 6 had a diamleter D and the wall of the pressur~e chamber 3 had a thickness of L. Three differenti devices were tested havin~g the following dimensions: (D 0. 15, 0.2, and 0.3 mm; L 0. 1, 0.2 and 0.3 5 mum; R& 0.2, 0.4, and 0.6 rm, respectively), and several distnces H fromn the tueot oteoiiernigfoH mm toH-=1.5rmmnhave been used. Thejet diamleter was measured at tihe hole exit and was plotted as a function of the pressure difference AP, and flow rate Q respectively. Although this technique allows for jet diameters even below one micron, larger flow rates and diameters have been used in this stu~dy to diminish the measurin~g errors.

-36z/i U4 u a;oo rAA o01 z VOU OAUU r.D. ti r C U. JU44 In order to collapse all of the data, we define a reference flow rate Q. and diameter t( based on the minimal values, from expressions and that can be attained in stable regime for a given AP:

Q

Y

J d. d-L PIAP AP 6 where y and p, are the liquid-gas surface tension and the liquid density, respectively. These definitions provide the advantage of a nondimensional expression for as d/d r 8 I)y(QQ 0 (7) which allows for a check for the validity of neglecting the surface tension term in Q/Q. should be large).

Notice that if the measured dj follows expression the surface tension cancels out in Also notice that d/d a We/2.

350 measured values of d/d versus Q/Q, are plotted in Fig. 5. A continuous line represents the theoretical prediction independent of liquid viscosity and surface tension.

The use of different hole and tube diameters as well as tube-hole distances does not have any appreciable influence on d. The collapse of the experimental data and the agreement with the simple theoretical model is excellent. Finally, the experimental values of Q are at least four 6ms large than Q 0 (being in most cases several hundreds times larger), which justifies the neglect of the surface tension term in Eq. While the present invention has been described with reference to the specific embodiments thereof, it should be understood by those skilled in the art that various changes may be made and equivalents may be substituted without departing from the true spirit and scope of the invention. In addition, many modifications may be made to adapt a particular situation, material, composition of matter, process, process step or steps, to the objective, spirit and scope of the present invention. All such modifications are intended to be within the scope of the claims appended hereto.

-37-

Claims (12)

1. A method of producing coated particles, comprising the steps of: forcing a liquid formulation comprising a first liquid through a channel of a first feeding source in a manner which causes a stream of the first liquid to be expelled from a first exit opening at a first velocity; forcing a liquid comprising a second liquid through a second channel concentrically positioned around the first channel in a manner which causes a stream of the second liquid to be expelled from a second exit opening at a velocity which is substantially the same as the first velocity whereby the stream of second liquid is concentrically positioned around the stream of the first liquid; forcing a gas through a pressure chamber in a manner which causes the gas to exit the pressure chamber from an exit orifice positioned downstream of the concentrically positioned streams of the first and second liquids; wherein the density of the liquid formulation comprising the first liquid is substantially the same as the density of the liquid comprising the second liquid, and the gas focuses the concentrically positioned streams to a stable unified jet which flows out of the chamber exit orifice and breaks up into coated particles of the first liquid coated with the second liquid; 20 wherein the stable unified jet comprises a diameter d i at a given point A in the stream characterized by the formula: *Q 8p, 25 wherein di is the diameter of the stable unified jet, indicates approximately equally to where an acceptable margin of error is 10%, p, is the average density of the liquid of the unified jet and AP, is change in gas pressure of gas surrounding the stream at the point A and Q is the total flow rate of the stable unified jet; and wherein dj is a diameter in a range of 1 micron to 1 mm. S"

2. The method of claim 1, wherein the stable unified jet has a length in a range of from 1 micron to 50 mm.

3. The method of claim 1, wherein the stable unified jet is maintained, at least in part, by tangential viscous stresses exerted by the gas on a surface of the jet in an axial direction of the jet.

4. The method of claim 1, wherein the stable unified jet is further characterized by a slightly parabolic axial velocity profile. The method of claim 1, wherein the particles of the first liquid coated with the second liquid are characterized by having the same diameter with a deviation in diameter from one particle to another in a range of from 3% to

6. The method of claim 5, wherein the deviation in diameter from one particle to another is in a range of from 3% to

7. The method of claim 1, wherein a given coated particle has a diameter in a range of o.1 micron to 1oo microns and other particles produced have the same diameter as the given particle with a deviation of 3% to

8. The method of claim 1, wherein AP=P-P 1 the difference in pressure through the chamber exit orifice, is equal to or less than twenty times the surface tension of the S."second liquid with the gas, divided by the radius of the stable unified jet. S2 9. A method of delivering an aerosol to a patient, comprising: 25 forcing a pharmaceutically active liquid through a channel of a feeding source in a manner which causes the liquid to be expelled from an exit opening; forcing a gas through a pressure chamber in a manner which causes the gas to •e exit the pressure chamber from an exit orifice in front of a flow path of the liquid expelled from the exit opening of the feeding source; 30 wherein a stable liquid-gas interface is maintained and the liquid forms a stable capillary jet focused on the exit orifice of the pressure chamber by the gas. The method of claim 9, wherein the liquid has a viscosity in a range of from 0.3 x 10-3 to 5 x 10 2 kg/m/sec and the gas is air; wherein the liquid is forced through the channel at a rate in a range of 0.01l nl/sec to loo WL/sec and further wherein the air is forced through the opening of the pressure chamber at a rate in the range of from 50 m/sec to 2000 m/sec; wherein the feeding source is a cylindrical channel and the liquid is expelled from an exit opening having a diameter in the range of from 0.002 to 2mm and wherein the opening in the pressure chamber has a diameter in the range of 0.002 to 2mm and is positioned directly in front of a flow path of the exit opening of the channel; wherein the exit opening in the pressure chamber has a diameter in the range of from 0.ol0mm to o.4mm or more preferably 0.005 mm to 0.25 mm; and wherein the exit opening of the feeding source is separated by a distance of from o.ol01 to 2mm (or more preferably 0.002 to 1 mm) from the exit opening in the pressure chamber.

11. A method of delivering aerosolized particles of a pharmaceutically active drug to a patient, comprising: feeding liquid formulation comprised of a pharmaceutically active drug through a liquid feeding source to an outlet; feeding gas through an orifice positioned in front of the outlet in a direction S- 20 aligned with a direction of flow out of the outlet; •t wherein the feeding of liquid and feeding of gas are each at a rate relative to each other so as to maintain a stable capillary microjet of liquid which exits the orifice and forms aerosolized particles having a size in the range of 0.1 micron to 1O microns. S 25 12. The method of claim 11, wherein gas is forced into an area around the feeding source outlet at a pressure in the range of 10 to 50000, mBar above atmospheric pressure and further wherein the liquid has a viscosity in the range of from 10-4 to 1 kg/m/sec; i wherein gas from the pressure chamber surrounds liquid exiting the feeding 30 source outlet which liquid is drawn into the orifice concentrically being focused by the gas flow out of the outlet, and further wherein the aerosolized particles formed are eeoc *uniform in size to the extent of having a relative size standard deviation of 1% to

13. A monodisperse aerosol of particles of a pharmaceutically active drug, the particles characterized by having the same diameter with a deviation in diameter from one particle to another in a range of to 41

14. The monodisperse aerosol of claim 13, wherein the particles are comprised of a liquid formulation of the drug and a carrier and wherein the particles have a size in a range of from 1 micron to 5 microns and the particles have a standard size deviation of less than The monodisperse aerosol of claim 14, wherein the particles are comprised of a dry powder of drug and wherein the particles have a size in a range of from 1 micron to 5 microns and the particles have a standard size deviation of less than

16. A monodisperse aerosol of hollow droplets comprised of a carrier and a pharmaceutically active drug, wherein the hollow droplets have air in their center and have an actual physical size diameter in a range of from 200 microns to 1 micron.

17. The monodisperse aerosol of claim 16, wherein the diameter is in a range of from 50 microns to 1 micron and the hollow droplets have a standard size deviation in diameter from one particle to another in a range of from +0.01% to 20 18. The monodisperse aerosol of claim 17, wherein the diameter is in a range of from 10 microns to 1 micron and each hollow droplet is substantially identical to other hollow droplets in size with a deviation of less than 1%. Dated this 1st day of March 2006 Universidad De Sevilla Patent Attorneys for the Applicant: F B RICE CO 000 F B RICE CO 0 *00° *o