Faraday instability-based micro droplet ejection for inhalation drug delivery

Abstract

We report here the technology and the underlying science of a new device for inhalation (pulmonary) drug delivery which is capable of fulfilling needs unmet by current commercial devices. The core of the new device is a centimeter-size clog-free silicon-based ultrasonic nozzle with multiple Fourier horns in resonance at megahertz (MHz) frequency. The dramatic resonance effect among the multiple horns and high growth rate of the MHz Faraday waves excited on a medicinal liquid layer together facilitate ejection of monodisperse droplets of desirable size range (2–5 µm) at low electrical drive power (<1.0 W). The small nozzle requiring low drive power has enabled realization of a pocket-size (8.6 × 5.6 × 1.5 cm³) ultrasonic nebulizer. A variety of common pulmonary drugs have been nebulized using the pocket-size unit with desirable aerosol sizes and output rate. These results clearly provide proof-of-principle for the new device and confirm its potential for commercialization.

INNOVATION

Drugs designed to treat pulmonary diseases or for systemic absorption through the massive surface area of the lung require optimum particle sizes (1 to 6 µm). Current advanced commercial devices such as Omron, Pari eFlow, and Philips I-neb produce droplets or aerosols by a vibrating mesh. All these devices suffer from broad particle size (poly-disperse) distributions and lack of size control capability, and also are plagued by clogging of mesh orifices used.

Here we report a new delivery device that has demonstrated capability for control of particle size within the optimum size range at low drive power and freedom from clogging. The new device employs a novel silicon-based ultrasonic nozzle with multiple Fourier horns in resonance designed to operate based on the newly discovered phenomena of Faraday waves at the frequency range of 1 to 2.5 MHz. The superior performance and batch fabrication economy of the centimeter-size nozzles have paved the way for commercialization of the new device.

INTRODUCTION

Inhalation is an increasingly important route for non-invasive drug delivery for both systemic and local applications^1–5. Control of particle (aerosols or droplets in air) size and output plays a critical role in the efficient and effective delivery of often expensive medications to the lung. Drugs designed to treat pulmonary diseases or for systemic absorption through the alveolar capillary bed require optimum particle sizes (1 to 6 µm) for effective delivery. Based on deposition of aerosol particles in human respiratory tract following a slow inhalation (250 mL/s air rate) and a 5-second hold^4,5, particles 1 to 4 µm (in aerodynamic diameter) deposit primarily in the alveolar region while particles 3 to 6 µm deposit in the airways. Particles larger than 6 µm deposit primarily in the mouth and throat. It is to be noted that respirable particles in the 1–4 µm range are optimal for alveolar deposition, but contain less drug than particles in the 2–5 µm range by 4 to 8 times per particle. With proper control of particle size, drugs ranging from small molecules such as pentamidine to large molecules such as insulin could be effectively delivered as aerosols via inhalation for systemic absorption or local treatment in airways and lung parenchyma^6–10.

Current commercial devices such as Misty-Neb, AeroEclipse, Omron, Pari eFlow, and Philips I-neb produce droplets or aerosols by compressed air, a vibrating piezoelectric plate together with a metallic mesh, or a vibrating mesh. As the examples (plots a and b) show in Fig. 1A, all these devices suffer from broad particle size (poly-disperse) distributions (with high geometrical standard deviation (GSD) > 1.5), making it difficult to deliver sufficient drug of desirable sizes to targeted sites precisely and rapidly. Furthermore, the eFlow and I-neb which utilize vibrating mesh technology¹¹ and are considered the most advanced commercial devices suffer from clogging of the mesh orifices used¹². Thus, delivery devices with control capability in particle size and less susceptibility to clogging are clearly needed for rapid administration at sufficient quantities of often expensive drugs such as agents to treat pulmonary hypertension and toxic inhalation agent antidotes^13,14. Here we report a new delivery device that utilizes a silicon-based megahertz (MHz) multiple-Fourier horn ultrasonic nozzle (MFHUN) and temporal instability of Faraday waves^15,16 to fulfill the unmet needs. As plots c, d, and e of Fig. 1A show, by varying the drive frequency of the MFHUN from 1.0 to 2.5 MHz the new device was capable of controlling the particle sizes within the optimum size range (1 to 6 µm) with a much narrower size distribution.

Figure 1.

Comparison of (A) particle diameters and size distributions (MMAD/MMD µm, GSD), and (B) Cumulative undersize percentage between the commercial devices and the new device: (a) Philips I-neb⁹, (b) Pari eFlow⁸, (c) 1.0 MHz, (d) 1.5 MHz, and (e) 2.5 MHz. Note that MMAD for (a)–(b) and MMD for (c)–(e) stand for mass median aerodynamic diameter and mass median diameter, respectively; MMD and geometrical standard deviation (GSD) for the new device were obtained using Malvern/Spraytec particle sizer (Model #STP 5311); and GSD of 1.0 represents a single particle size and the closer the GSD to 1.0 the higher the monodispersity.

The much narrower droplet size distribution achieved by the new device (MFHUN), compared to advanced commercial devices, is also illustrated by its steeper cumulative undersize percentage curves as shown in Fig. 1B. Notably, plots c to e show that 100% or nearly 100% of droplets are smaller than 6 µm in diameter for the new device while plots a and b show only 66% and 85%, respectively, for the advanced commercial nebulizers. The much narrower droplet size distribution will have significant economic impact on lung deposition of drugs, especially of expensive drugs. For example, as shown in Fig. 1A the MMD of plot c obtained using the new device at 1 MHz drive frequency is similar to the MMAD of plot a obtained using one of the most advanced commercial nebulizers — the I-neb Adaptive Aerosol Delivery (AAD) System that delivers aerosol only during inspiration (i.e. breath actuation). However, as shown in plot c of Fig. 1B the narrower size distribution with a smaller GSD obtained by the former (1.16 versus 1.70) leads to a much greater respirable (inhaled) fraction (nearly 100% versus 66% with particle diameter <6 µm) than by the latter. Therefore, the inhaled dose delivered to the lung using the new device together with breath actuation should be much higher than the 60 ± 5% achievable using the I-neb AAD System⁹. Note that as presented in the final section of “Methods and Results”, the capability of breath actuation with the pocket-size MFHUN unit has been demonstrated (Fig. 6C).

Figure 6.

Pocket-size silicon-based ultrasonic nebulizer (8.8 cm × 5.9 cm × 1.9 cm): (A) Layout of ultrasonic nozzle and other mechanical components. (B) Layout of battery-powered electronic driver. (C) Externally actuated ON/OFF nebulization using a Harvard ventilator at 55% inspiratory, 750 mL tidal volume, and 10 breaths/min.

Faraday waves, also known as standing capillary waves, were first observed as wavy surface of a water layer resting on a solid surface subjected to perpendicular mechanical vibration, as depicted in Fig. 2A, at a very low drive (or forcing) frequency (5 Hz) by Faraday in 1831¹⁷. They were subsequently analyzed by Rayleigh in 1883¹⁸ and many others¹⁹. Faraday instability, the underlying physical mechanism for Faraday wave formation and amplification, was studied^20–24 extensively based on the above classical planar geometry during the 1990s, and continued^25–29 in the 2000s, but mostly at very low drive frequency ranging from tens to thousands hertz (Hz). At such low drive frequencies, various standing-wave patterns were observed^25,30 when the mechanical vibration amplitude on the solid surface reached the onset threshold for Faraday wave formation. However, no droplet ejection (atomization) was described in the theoretical treatments of these earlier low-drive frequency studies. In the few reports on atomization at low drive frequencies, droplet ejection was found to take place only when the mechanical vibration amplitude on the solid surface was much higher than the onset threshold for Faraday wave formation. In stark contrast, our recent discovery as reported here using a novel device, silicon-based MHz MFHUN (see Fig. 2B), shows that at the much higher drive frequencies of MHz the onset threshold for Faraday wave formation is much lower and the mechanical vibration amplitude required for subsequent droplet ejection is only slightly higher than the onset threshold for Faraday wave formation¹⁵. These findings have also been verified by the rigorous theoretical treatment^31,32 summarized in the first section of “Methods and Results”.

Figure 2.

(A) Classical planar geometry for Faraday waves formation at low drive frequency (5 Hz). (B) 3-D architecture of the MHz multiple-Fourier horn ultrasonic nozzle. (C) Geometry of nozzle end face and planar liquid layer d in depth.

METHODS AND RESULTS

Linear theory on MHz Faraday wave instability for monodisperse micro droplet ejection

Classical model of planar liquid layer geometry

Figure 2C depicts the classical model geometry for theoretical study of Faraday instability reported here — a planar liquid layer of depth (d) resting on a planar solid surface mimicked by the end face of the silicon-based multiple-Fourier horn ultrasonic nozzle (MFHUN) that vibrates longitudinally along the nozzle axis (Z-axis) at the resonance frequency (f). The longitudinal vibration of the nozzle end face exerts a periodic pressure on the planar liquid layer. Faraday waves are formed on the free surface of the liquid layer when the peak vibration displacement of the nozzle end face (h) reaches the critical value or onset threshold (h_cr). The following theoretical treatment for the Faraday wave with amplitude ξ(x, t) that represents the time-dependence displacement of the free liquid surface is based on mass-conservation and linearized Navier-Stokes equations for incompressible Newtonian liquids, such as water, with density ρ, surface tension σ, and kinematic viscosity ν³³.

Due to symmetry in the XY-plane, it is convenient to separate the coordinates into the perpendicular Z-direction and the XY-plane with x designating (x, y). First, we utilize Navier-Stokes equations and incompressibility condition, and apply the boundary conditions on the nozzle end face and the free liquid surface of the planar liquid layer geometry of Fig. 2C. The Navier-Stokes equations together with the Laplace’s equation are solved by means of the Fourier transform in the horizontal coordinates (x) to result in the following equation for the temporal evolution of the amplitude ξ_k of the k^th mode Faraday waves for low-viscosity liquid such as water:

$${\partial }_t^2{\xi }_k(t)+4\nu k^2{\partial }_t{\xi }_k(t)+({\omega }_k^2+k{g}_e(t)){\xi }_k(t)=0$$ (1)

where the wave number (k) equals 2π/λ and ${\omega }_k^2\equiv \sigma k^3/\rho$ and g_e(t) is the external acceleration, equal to h(2πf)²cos(2πft). Finally, Eq. (1) is solved using the established technique for the k^th mode Faraday waves amplitude ξ_k(x, t) in the most unstable region (the first tongue-like region) of the stability chart^15,34 and the result is given by Eq. (2).

Temporal instability of MHz Faraday waves and onset threshold for droplet ejection

The amplitude of the fastest growing k_th mode Faraday wave ξ_k(t) in the most unstable region is obtained as follows:

$${\xi }_k(t)={\xi }_0{e}^{(\pi kfh-\beta )t}\text{sin}(\pi ft-\pi /4),\mathrm{i}.\mathrm{e}.{\xi }_k(t)={\xi }_0{e}^{\pi kf(h-h_{cr})t}\text{sin}(\pi ft-\pi /4),$$ (2)

where ξ₀ designates the initial wave amplitude and h_cr ≡ β/(π kf), in which ≡ α2νk². Equation (2) clearly shows that the single-mode Faraday waves excited carry a frequency equal to one-half of the drive frequency (f), namely, ω_k = πf, and the corresponding wavelength (λ) and onset threshold (h_cr) are given as follows:

$$\lambda =2\pi /k=2\pi {(\rho {\omega }_k^2/\sigma )}^{-1/3}=2\pi {(\rho {\pi }^2{f}^2/\sigma )}^{-1/3}={(8\pi \sigma /\rho )}^{1/3}{f}^{-2/3}.$$ (3)

$$h_{cr}\equiv \beta /(\pi kf)=2\nu k/(\pi f)=2\nu {\rho }^{1/3}{(\pi \sigma )}^{-1/3}{f}^{-1/3}.$$ (4)

Note that the onset threshold (h_cr) is at the minimum at the critical wave number (k_cr), i.e., the k^th mode Faraday wave in the aforementioned stability chart and also referred to in Eq. (1).

Dynamics of droplet ejection

Equation (4) shows the specific dependence of the onset threshold on the drive frequency (f) and the liquid properties (ρ, σ, and ν). It should be emphasized that h_cr decreases with the drive frequency in accordance with f^−1/3, and the wave amplitude ξ_k(t) grows exponentially in time when h > h_cr. It is to be emphasized also that h_cr not only falls within the most unstable region but also decreases from 0.33 to 0.29 and to 0.24 µm, respectively, for water as an example, as f increases from 1.0 to 1.5 and to 2.5 MHz. The corresponding Faraday wavelengths are 12.2, 9.2, and 6.6 µm, and the respective wave numbers k_cr’s are 5204, 6820, and 9498 cm⁻¹. While h_cr decreases with the drive frequency in accordance with f^−1/3, the exponent πk f (h − h_cr)t in the exponential factor of ξ_k(t) increases with the drive frequency in accordance with f^4/3. Thus, the temporal growth of the single-mode Faraday wave amplitude at MHz drive frequency is very rapid once the nozzle end face excitation displacement h exceeds the onset threshold h_cr. Take the Faraday waves at the 1.5 MHz drive frequency with the corresponding wavelength of 9.2 µm in water as an example, when h exceeds the predicted h_cr of 0.29 µm by as little as 0.01 µm the growth rate factor e^{π kf (h−h_cr)t} of the Faraday wave amplitude in Eq. (2) in a time increment of 0.5 ms is ~10⁷ times that with (h − h_cr) as large as 100 µm at 200 Hz²⁵ drive frequency at the same time increment (0.5 ms). The amplitude growth rate factor at 1.5 MHz with (h − h_cr) of 0.01 µm in a time increment of 0.5 ms is still greater (by 8%) than that at 200 Hz with (h − h_cr) of 100 µm in a much longer time increment of 143 ms. When the wave amplitude grows and becomes too great to maintain wave stability, the Faraday waves break up, and thus droplets are ejected from the free surface of the liquid layer^31,32.

In summary, the scientific discovery on dynamics of droplet ejection at MHz drive frequency as described above together with the greatly magnified displacement on the end face of the MHz MFHUN enables generation of micrometer-size monodisperse droplets at low electrical drive power. This discovery also serves to explain the observations on atomization experiments at low drive frequencies described in “Introduction”.

Size of ejected droplets

Now the ejected spherical droplets, with radius a and at the lowest oscillation mode frequency^33,35, are dispersed in air and free from external acceleration (g_e). Thus, by setting the frequency of the droplet’s lowest oscillation mode equal to the nozzle drive frequency, the following theoretical droplet diameter (D_p ≡ 2a) in terms of Faraday wavelength (λ) is obtained:

$$D_p=2{(2/{\pi }^2)}^{1/3}{(\sigma /\rho )}^{1/3}{f}^{-2/3}=0.40\lambda .$$ (5)

Namely, the predicted droplet diameter equals four tenths of the Faraday wavelength involved. Note that this is the first theoretical formula for the droplet diameter in capillary wave-based ultrasonic droplet ejection. Clearly, for a given liquid to be nebulized, the desired size of the droplets can be controlled by the drive frequency (f) of the MFHUN in accordance with f^−2/3; the smaller the droplet size the higher the nozzle drive frequency.

Design, simulation, fabrication, characterization of MFHUNs, and droplet ejection studies

Design and simulation of MFHUNs

The innovation of the new ultrasonic delivery device lies in the multiple Fourier horns in cascade that vibrate in a single longitudinal mode along the nozzle axis at the same and single resonance frequency. As shown in Fig. 2B, the nozzle consists of a drive section and a resonator section. A lead zirconate titanate (PZT) piezoelectric transducer is bonded on the drive section to excite mechanical vibrations along the nozzle axis (Z-axis). The resonator section is made of multiple (3 in the example) Fourier horns in cascade³⁶. The nozzle is designed to vibrate at the resonance frequency of the multiple Fourier horns. The resultant vibration amplitude (displacement) on the nozzle end face (tip of the distal horn) is greatly magnified with a gain of Mⁿ for a n-Fourier horn nozzle with a magnification of M for each Fourier horn¹⁵. In order to produce monodisperse droplets of desirable size range (2–5 µm) at high output rate and low electrical drive power (sub Watt), the optimum magnifications of 1.8 and 1.3 for the first three and the last horn of the 4-Fourier horn nozzles, respectively, were used in the design of optimized MFHUN.

Each Fourier horn is of half acoustic wavelength design^36,37. The longitudinal axis of the nozzle (Z-axis) is in the direction of the primary flat, namely, <110> of the silicon wafer with the highest longitudinal acoustic velocity. A three-dimensional (3-D) finite element method (FEM) simulation is carried out using the commercial ANSYS (ANSYS Inc., Canonsburg, PA) Program first for vibration mode shape analysis and then for electrical impedance analysis³⁶. The mode shape analysis determines the nozzle resonance frequency of pure longitudinal vibration mode; the electrical impedance analysis determines the longitudinal vibration displacement on the nozzle end face and the impedance at the nozzle resonance frequency. The resultant longitudinal vibration displacement on the nozzle end face is then used to determine the threshold voltage required to produce the onset threshold (h_cr) given by Eq. (4) in the first section of “Methods and Results”. The threshold voltage thus obtained together with the resistive part of the impedance is then used to calculate the electrical drive power required for droplet ejection. Both the simulation results with the acoustical and electrical losses taken into account and the experimental results show that the optimum number of Fourier horns (n) in terms of electrical drive power requirement is 3 or 4. Due to the design of the nozzle, the onset threshold (h_cr) cannot be reached at any frequency outside the bandwidth (<10 kHz) of the designed resonance frequency; this design feature is critical to producing monodisperse droplets of the desired diameter at very low drive power.

The greatly enhanced peak vibration displacement (h) on the nozzle end face of the MFHUN facilitates the onset threshold required for initiation of temporal instability of the Faraday waves on the surface of the liquid layer and subsequent ejection of droplets at low electrical drive.

Fabrication and characterization of MFHUNs

The silicon-based MFHUNs of various designs in terms of the drive (resonance) frequency and the number of Fourier horns in cascade were fabricated using micro-electromechanical system (MEMS) technology.³⁸ The drive and resonator sections of the MFHUN were formed in a single-fabrication step using an inductive coupled plasma (ICP) process³⁹. Note that a large number of nozzles with identical or different design specifications can be fabricated in one batch in a common silicon wafer. Figure 3A shows the layout of a large number of nozzles. Aft er bonding of the PZT plate and the connecting wires, the resulting centimeter-size nozzles with optimum designs are shown in Fig. 3B.

Figure 3.

(A) Layout of nozzles with various designs in a common silicon wafer. (B) 1.0, 1.5, 2.0, and 2.5 MHz nozzles fabricated and studied.

The only essential characterization of the fabricated nozzle prior to droplet ejection experiments is to measure the impedance curve using Agilent impedance analyzer model 4294 A¹⁵ from which the drive frequency and the electrical drive power are determined.

Droplet ejection studies

All the droplet ejection (atomization) experiments presented below were conducted using either the established bench-scale setup¹⁶ or the pocket-size nebulizer. Major components of the setup are: (i) a PZT transducer drive system to provide a MHz electrical drive to the ultrasonic nozzle, (ii) a Syringe Pump (kd Scientific Model #101) to provide a controlled flow rate of liquid, (iii) a CCD camera to take pictures or movies of the droplets produced, and (iv) a Malvern/Spraytec System (Model #STP 5311) for analysis of the size distribution of the droplets.

Figures 4A and 4B show photos of a stream of droplets produced from a water layer resting on the nozzle end face of the MHz nozzles using a high-resolution high-speed camera system with back-lighting and a normal speed camera system with front lighting, respectively. The droplet rates for the outputs of 40 and 350 µL/min (set by external feeding of water) are 1.4 × 10⁷ and 4.0 × 10⁸ droplets/sec, respectively. These high droplet output rates are in contrast to the continuous ink-jet printing (<10⁶ droplets/sec (10 µm in diameter)⁴⁰) resulting from the Rayleigh-Plateau instability. The short depth of focus provided by the high-resolution high-speed camera system enabled visualization of individual droplets ejected from the liquid layer and the thin-lighted curve caused by the Faraday waves excited on its surface as shown in Fig. 4A. The wavy liquid surface caused by the Faraday waves acts like optical gratings to scatter the incident light. Note that light scattered by Faraday waves excited on the liquid surface also resulted in intense brightness over the entire spherical surface of a water ball³² (see Fig. 2a of Ref. 32) ejecting micro droplets.

Figure 4.

Stream of droplets ejected from the liquid layer resting on the end face of the nozzle tip: (A) High-speed photograph (10⁴ frame/sec, Vision Research Model Phantom® v7.2) with short depth of focus for droplets stream at output rate of 40 µL/min with 1.0 MHz nozzle. (B) CCD image (20 frame/sec) of droplets stream at output rate of 350 µL/min and electrical drive power of 0.27 W with optimized 2.0 MHz nozzle. Note that the scale bars in (A) and (B) are 25 µm and 105 µm, respectively.

The Malvern/Spraytec System is used to analyze the size distribution of droplets of Fig. 4B. The system is a non-invasive particle sizing instrument based on laser light diffraction. Since the laser beam had a cross sectional area (1 cm in diameter) much larger than the stream (spray) of droplets (~0.07 × 0.1 cm²) produced by the nozzles, the data reported in this study were obtained from the entire cross-section of the droplet stream. The data include mass median diameter (MMD) and geometrical standard deviation (GSD). GSD and geometric mean of a data set with a log-normal distribution are, respectively, equivalent to the standard deviation and arithmetic mean of a data set with normal distribution. GSD is simply calculated by the geometric mean of quotients (ratios) D₈₄/D₅₀ and D₅₀/D₁₆, where D₈₄, D₅₀, and D₁₆ are, respectively, the droplet diameters at 84.1%, 50.0%, and 15.9% of the cumulative undersize percent of the droplet size distribution as shown in Fig. 1B. Note that the GSD was determined by least square fit (R² = 0.9987) of the cumulative undersize percentage curve obtained using the Malvern/Spraytec System. Figure 5 shows clearly the excellent agreement between the measured droplet diameter and the theoretical prediction of Eq. (5) given in the first section of “Methods and Results”.

Figure 5.

Comparison between the measured droplet diameters and the theoretical predictions.

Pocket-size ultrasonic nebulizer and nebulization of common drugs

As presented in the last subsection, the droplet ejection (atomization) experiments using the bench-scale unit with a single nozzle have demonstrated the capability of the silicon-based MHz MFHUNs for production of monodisperse droplets of desirable size (2 to 5 µm) and moderate output (up to 0.2 mL/min) at low electrical drive power (sub Watt). At the typical output rate of 0.15 mL/min with 3.5 µm-diameter droplets, which equals 1.1 × 10⁸ droplets/sec, the required electrical drive power of 0.27 W corresponds to 2.4 nano joule per droplet generated. The centimeter-size nozzles with the low electrical power requirement enabled most recent realization of the first pocket-size ultrasonic nebulizer (8.6 × 5.6 × 1.5 cm³). The nebulizer contains a single ultrasonic nozzle with optimum design as shown in Fig. 3B, IC electronic driver, battery, micro pump, drug reservoir, and liquid feed tube. The layout of the nozzle with other mechanical components and that of the battery-powered electronic driver are shown in Figs. 6A and 6B, respectively. A range of drug substances for asthma, diabetes, pulmonary fibrosis, cyanide poisoning, etc. such as albuterol (salbutamol), Humulin U-100,⁴¹ cobinamide, interferon-γ¹⁰ and budesonide suspension⁴² have been nebulized using the pocket-size nebulizer with desirable aerosol size and output rate as summarized in Table 1. Some drugs such as albuterol and cobinamide, and aqueous glycerol were successfully nebulized using the bench scale and/or the pocket-size units at varying concentrations with the corresponding range of viscosity up to as high as 4.5 cP and flow rate up to 350 µL/min. Furthermore, the pocket-size nebulizer has demonstrated its capability for nebulization of commercial budesonide suspension medicine that none of the current commercial ultrasonic nebulizers is capable of. It is to be noted that the demonstrated moderate output rate of 100−350 µL/min by the pocket-size unit would provide a higher effective dosage over the current commercial nebulizers in light of the higher aerosol monodispersity produced by the former. Clearly, a higher output rate can be accomplished readily by using an array of identical ultrasonic nozzles.

Table 1.

Summary of drugs nebulized.

Medicine	Medicine concentration	Ultrasonic drive freq. (MHz)	Nebulizer unit	Droplet diameter (µm)	Output rate (µL/min)	Disease
Albuterol	25 mg/ml	1.0	Bench-scale	4.5	150	Asthma
Humulin, U100	100 units/ml	1.0	Bench-scale	4.5	100	Diabetes
Cobinamidea	100 mM	1.5	Pocket-size	3.9	130	Cyanide poisoning
Interferon-γb	100 µg/0.5 ml	2.0	Pocket-size	2.9	100	Pulmonary fi brosis
Budesonide suspension	0.5 mg/2.0 ml	2.0	Pocket-size	3.1	350	Asthma

^a,bProvided by Prof. G. Boss at UC San Diego and Prof. G.C. Smaldone, State Univ. of New York, Stony Brook, respectively.

Breath actuation is an important operational requirement of a nebulizer in order to save the often expensive medicine. As shown in Fig. 6C, nebulization of medicine is turned ON/OFF by turning ON/OFF the feeding of the medicine using a piezoelectric micro pump to the nozzle end face in synchronization with the inspiratory cycle of a ventilator (Harvard piston pump). Specifically, a probe senses the UP/DOWN movement of the piston pump and sends a corresponding electrical signal to turn ON/OFF the liquid feed. A pressure sensor, with or without an ON/OFF disk/membrane valve, can be used for breathing without a ventilator. Thus, the pocket-size ultrasonic nebulizer is also capable of fulfilling this operational requirement.

DISCUSSION

Recently, we mimicked the Faraday’s classical planar geometry of Fig. 2A using the end face of silicon-based MHz multiple-Fourier horn ultrasonic nozzle (MFHUN) and carried out theoretical and experimental studies at much higher drive frequencies (0.5 to 2.5 MHz)^15,16,31,32. The resonance effect among the multiple Fourier horns designed for the single vibration mode at the single resonance frequency greatly magnifies the vibration displacement of the nozzle end face (by a factor of 2ⁿ, where n designates the number of Fourier horns each with a magnification of 2) and, hence, readily facilitates formation and subsequent high-rate growth of Faraday waves on the free surface of the liquid layer. The single-mode Faraday wave excitation and amplification leads to generation of monodisperse droplets of desirable size range (~2 to 5 µm) at very low electrical drive power (<1 W)¹⁶. Such Faraday instability-based droplet ejection device is in stark contrast to all other existing ultrasonic devices employing a vibrating piezoelectric plate that simultaneously involve various droplet ejection mechanisms such as cavitation, impinging, and jetting in addition to capillary wave mechanism^25,43,44. These ejection mechanisms require much higher electrical drive power (by 1 to 2 orders of magnitude) and produce broad droplet size distribution. The low electrical drive power requirement for the new ultrasonic device reported here not only facilitates miniaturization, but also minimizes temperature rises that might damage the medications to be aerosolized.

The precise control of particle size and much narrower particle size distribution achieved by the new device will improve targeting of treatment within the respiratory tract and improve delivery efficiency, resulting in better efficacy, fewer side effects, shorter treatment times, and lower medication costs compared with existing nebulizers. Additionally, high output rate of medicinal aerosol is desirable because it will shorten treatment times. The new device with a single nozzle alone has demonstrated an output rate up to 350 µL/min. Clearly, the output rate can be readily increased by using an array of identical nozzles. Furthermore, nozzle arrays with individual nozzles designed at different resonance frequency will provide the unique capability for production and mixing of aerosols of different sizes and/or medicines. Such nozzle arrays will provide a unique platform for basic research. The MEMS-based fabrication technology involved will enable construction of nozzle arrays of various designs.

The micron-size monodisperse droplets produced by the miniaturized low-power ultrasonic nozzles reported here may find a wide range of attractive applications other than inhalation drug delivery. Some examples of such applications include synthesis of nano particles via spray pyrolysis, sample injection in chemical analysis (e.g. mass spectroscopy), high-quality thin-film coating,⁴⁵ 3-D photoresist coating of micro- and nano-structures in fabrication of nano-electronic and -photonic devices, fuel injection in combustion, delivery of lipid-based micro-encapsulated biological entities or gene,^46,47 pharmaceutical preparation such as double emulsion,^48,49 rapid heat removal in laser surgery, and rapid administration of cosmetics.

The work reported here is the result of an interesting cross-fertilization between science and technology. Our original quest for an atomizer capable of producing micron-size monodisperse droplets led to the invention of the silicon-based MHz multiple Fourier horns in cascade. Subsequently, the study of the underlying mechanisms for droplet ejection led all the way back to Faraday’s 1831 classical experiments involving a water layer at the very low drive frequency of 5 Hz and to many subsequent theoretical and experimental studies, but still at low drive frequencies up to ten’s kHz. Our theoretical study on Faraday waves at the much higher drive frequency of MHz uncovered the frequency-sensitive dynamics of droplet ejection. The multiple-Fourier horn ultrasonic nozzle provides experimental verification of the theoretical predictions and facilitates the production of monodisperse droplets of desirable micron-size range at low drive power. Despite more than 180 years of continued interest and studies on Faraday waves it was only through our invention and realization of MHz ultrasonic nozzles with multiple Fourier horns in resonance that the technical potential of Faraday waves at MHz frequency range was discovered and led to a technology with immediate and important medical applications.