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An investigation into a piezoelectrically actuated nebulizer with μEDM-made micronozzle array
Experimental Thermal and Fluid Science 31 (2007) 1147–1156 www.elsevier.com/locate/etfs
An investigation into a piezoelectrically actuated nebulizer with lEDM-made micronozzle array Y.R. Jeng
a,* ,
C.C. Su
a,b
, G.H. Feng a, Y.Y. Peng
c
a
Department of Mechanical Engineering, National Chung Cheng University, Ming-Hsiung, Chia-Yi 621, Taiwan Department of Mechanical Engineering, WuFeng Institute of Technology, Ming-Hsiung, Chia-Yi 621, Taiwan c Industrial Technology Research Institute, Mechanical Industry Research Laboratories, Hsin-Chu, 310, Taiwan b
Received 5 June 2006; received in revised form 11 December 2006; accepted 18 December 2006
Abstract This study uses micromachining techniques to fabricate a piezoelectrically actuated nebulizer with an easily assembled structure. The experimental measurement and numerical simulation have been performed to gain the insight of atomized droplet behavior and fluidic mechanics for the developed devices. It is expected that a better understanding of these physical insights can lead to various potential applications. Based on the results obtained from a series of investigations, this study has found the optimal operating conditions for the proposed nebulizer. Various test fluids with different viscosities and surface tensions are considered, including water, dilute alcohol, oil, and a honey-like fluid. A spray stability analysis is performed in which the operating voltage and frequency of the piezoelectric actuator in the nebulizer are varied and the corresponding changes in the ejected particle size, flow rate and ejection distance examined. Air trapped within the nebulizer reservoir adversely affects the injection performance. Hence, the effect of the operating parameters on the volume of air ingested into the reservoir during the nebulizer operation is investigated and discussed. Ó 2006 Elsevier Inc. All rights reserved. Keywords: Micronozzle; Piezoelectrically actuated nebulizer; lEDM
1. Introduction Rapid advances in micro-electro-mechanical-systems (MEMS) technologies have facilitated the development of microfluidic systems for a diverse range of applications in many scientific and engineering fields [1–4]. In piezoelectric nebulizer, the atomization and flow rates of the ejected droplets are governed by the interactions between the viscosity and surface tension of the ejected liquid and the pressure developed within the reservoir. In operation, a voltage is applied to the piezoelectric actuator. The corresponding
* Corresponding author. Present address: Department of Mechanical Engineering, National Chung Cheng University, 160, San-Hsing, MingShiung, Chia-Yi 621, Taiwan. Tel./fax: +886 5 2428189. E-mail address: [email protected] (Y.R. Jeng).
0894-1777/$ - see front matter Ó 2006 Elsevier Inc. All rights reserved. doi:10.1016/j.expthermflusci.2006.12.001
deflection of the piezoelectric actuator produces a volume change in the reservoir and this causes an acoustic pressure wave to propagate towards the nozzles, forcing droplets to be expelled through the orifice plate. When the direction of the applied voltage is reversed, the piezoelectric actuator deflects in the opposite direction. However, the negative pressure created by the backward deflection of the piezoelectric actuator may draw some external air into the reservoir. In the present study, experimental and numerical investigations are performed to investigate the ingestion of air bubbles into the reservoir and to examine the size of the ejected droplets together with their flow rate and ejection distance under different operating conditions. The present results are intended as a reference for the development of microparticle sprays in various applications, including medical liquid sprays, methanol injectors in direct methanol fuel cell (DMFCs), and odor generators, etc.
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Nomenclature f g L t T p u v V Vpk–pk X w q qL
liquid volume fraction gravity acceleration (m/s2) length (m) time (s) thickness (m) fluid pressure (N/m2) velocity in the x-direction velocity in the y-direction voltage (V) voltage for peak to peak (V) displacement (m) velocity in the z-direction mixing density (kg/m3) density of the liquid fraction
The main function of an atomization nozzle is to break a bulk volume of liquid into many small droplets for increasing the surface to volume ratio of the liquid. The changes which occur in the spray effect of an atomization nozzle under different applied voltages and frequencies depend on the viscosity and surface tension of the ejected fluid. In the previous study, Pimbley [5] applied linear one-dimensional analysis and the boundary value perturbation method to investigate droplet formation. Fromm [6] introduced a two-dimensional flow line governing equation to investigate the effect of the fluid behavior on the droplet structure and velocity under different Reynolds numbers and driving pressures. The results were then used to estimate the corresponding variations in the ejected droplet stream. Asai et al. [7] conducted numerical simulations and experimental studies to investigate variations in the droplet stream length and developed a one-dimensional numerical model to simulate the droplet formation. Later, Asai [8] applied the finite difference method to solve the three-dimensional Navier– Stokes equation, and predicted the path of ink droplets in a side-direction hot bubble type inkjet printer. The numerical results showed that the ink droplet volume and velocity increased as the pressure applied to the bubble increased, but reduced as the viscosity and surface tension of the ink increased. Liou et al. [9] applied the finite volume method to discretize the flow field governing equation taking the surface tension effect into account and to predict the free interface of the liquid and gas. Moreover, Yeh [10] used the fractional volume of fluid (VOF) and finite element method (FEM) approaches to investigate the development of the free interface of the liquid and gas during fluid ejection in a piezoelectric type inkjet. Performing experimental observations of the variation in size of droplets ejected through the orifices of a nebulizer is challenging. Hence, this study employs computational fluid dynamics software (CFDRC) to assist the droplet stream analysis and to investigate the ingress of air into the reservoir body during the injection process.
qG l Fr ~ V /L /G / rij sij m k dij
density of the gas fraction dynamic viscosity (kg/m s) surface tension force (N/m) velocity vector (m/s) the value of the quantity for liquid the value of the quantity for gas volume averaged quantity stress tensor shear stress tensor kinematic viscosity second coefficient of viscosity (depend on density of fluid) kronecker delta, when i = j, dij = 1; i 5 j, dij = 0
2. Piezoelectrically actuated nebulizer structure with micronozzles The fabricated nebulizer consists of a piezoelectric actuator, an injection orifice-array plate, a front cover plate, and the reservoir, as shown in Fig. 1. In this study, the nebulizer is made of polyoxymethylene (POM), and incorporates a body with a inner volume of approximately 4 cm3. The function of the piezoelectric actuator, which is located between the injection orifice plate and the reservoir, is to transform the electric energy supplied by an external electrical pulse into mechanical energy. When a voltage is applied to the piezoelectric actuator, it deflects and generates the pressure on the liquid within the reservoir. A deflection of the piezoelectric actuator in the forward direction causes the liquid in the reservoir to be squeezed through the array of orifices in the injection orifice plate, as shown in Fig. 2. The present experiments were conducted using a bimorph ceramic piezoelectric actuator fabricated from lead zirconium titanate (PZT). The bimorph structure was fabricated using a steel film of thickness 100 lm pressed on either side with a PZT ceramic piezoelectric actuator of thickness 200 lm. Therefore, the total thickness of the bimorph piezoelectric structure was 0.5 mm. The length and width of the piezoelectric structure were 20 mm and 10 mm, respectively. In this bimorph structure, the external piezoelectric films have one polarity, while the steel film has the reverse polarity. Hence, when an electrical voltage is applied to the bimorph structure, the piezoelectric actuator on one side of the structure expands, while that on the other side contracts, causing a deflection of the bimorph structure. The injection orifice plate of our present study is designed to the form of conical orifices-array. Fabrication of the injection orifice plate commenced by using a Nd-YAG laser to drill circular straight holes of diameter 50 lm through a stainless steel plate of thickness 50 lm. A micro electrostatic discharge machining (lEDM)
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liquid replenishment tube
injection orifice-array plate
reservoir body
sealing ring
cover plate
piezoelectric actuator
Fig. 1. Schematic illustration of piezoelectric nebulizer.
Fig. 2. Schematic illustration of the working principle of the designed nebulizer.
technique was then used to machine the circular holes into a conical profile, with an entrance diameter of 100 lm and
an exit diameter of 50 lm. As shown in Fig. 3, the injection orifice plate was fabricated with two parallel rows of orifices, with each row containing 25 orifices. 3. Atomized particle size measurement
Fig. 3. Schematic illustration and SEM image of injection orifice plate.
The current particle size measurements were carried out by using an INSITEC particle size analyzer. The analysis system comprised an optical arrangement, a detector module, and data collection and analysis software installed on a PC. The laser light was generated by a luminous diode and passed though a beam expander from which it emerged as a parallel light beam with a diameter of 9 mm. The light beam was passed through the droplets in the atomization flow field causing a diffraction effect to take place. The diffracted light was collected by a Fourier Lens and passed to a photo detector consisting of 31 ring-type light intensity sensors. The sensors detected changes in the diffracted light pattern caused by variations in the spray ejected from the injection orifice plate. The detected light signal was transformed into an electrical signal and, following signal amplification, further transformed into a digital signal by an analog-digital converter (A/D converter). Finally, the
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resulting signal was transported to a PC for particle sizing analysis.
4.2. Boundary condition setting The boundary conditions in the procedure of calculation include two parts: One is liquid in the reservoir, and the other is liquid left the reservoir to form the droplets.
4. Theoretical analysis 4.1. Governing equation The continuity equation is given by: oq ~Þ ¼ 0 þ r ðqV ot
ð1Þ
The momentum equation is given by: o ~u ¼ op þ o 2l ou þ o l ov þ ou ðquÞ þ r qV ot ox ox ox oy ox oy o ou ow þ l þ þ F rx þ qgx oz oz ox o op o ov ou o ov ðqvÞ þ r qV~v ¼ þ l þ þ 2l ot oy ox ox oy oy oy o ov ow l þ þ F ry þ qgy þ oz oz oy o op o ou ow o ov ow ðqwÞ þ r qV~w ¼ þ l þ þ l þ ot oz ox oz ox oy oz oy o ow 2l þ F rz þ qgz þ oz oz
ð2Þ
where u, v and w, represent the velocity in the x- y- and zdirections, p, q and l represent the pressure, density, and viscosity coefficients, respectively, Fr is the surface tension force acting at the liquid/gas interface and qg is the gravity term. The governing equations given above yield the velocity and pressure distribution of the system. The volume of fluid (VOF) method is then applied to obtain the distribution of the flow field volume fraction in order to establish the path of the ejected droplets. For given a flow field and an initial distribution of f on a grid, the manner in which the volume fraction distribution f evolves is determined by solving the passive transport equation: of þ r V~f ¼ 0 ot
ð3Þ
where f is the liquid volume fraction, t is time, $ is the standard spatial ‘‘grad’’ perator, and V~ is the velocity vector. When two different fluids exist simultaneously in the computational grid, they can be regarded as a single fluid for the purposes of establishing the physical property of the flow field in the grid. The physical property of the flow field is determined via a process of linear interpolation based on the liquid state volume fractions of the two fluids in the grid, i.e. /¼
f qL /L þ ð1 f ÞqG /G q
ð4Þ
where /L and /G indicate the value of the liquid and gas of the physical property, respectively, and q is the mixing density, given by q = fqL + (1 f)qG.
(a) The free surface boundary condition. The boundary condition in the free surface when the droplets ejected for the nebulizer are as follows: Because the stress at the free surface of two fluids must be equilibrated, the magnitude of the stress tensor for moving fluid can be expressed as rij ¼ P dij þ sij oui ouj ouk þ And sij ¼ m þk dij oxj oxi ouk
ð5Þ ð6Þ
We assume the fluid is an incompressible flow, and then k = 0. Consequently, the stress tensor ri,j can be obtained as oui ouj þ rij ¼ P dij þ m ð7Þ oxj oui On the free surface boundary, the stress in the normal and tangential direction must be in equilibrium, these conditions can be expressed as ðrij nÞ n ¼ 0 ðrij nÞ m1 ¼ 0
ð8Þ ð9Þ
ðrij nÞ m2 ¼ 0
ð10Þ
(b) The solid wall boundary condition. The boundary condition on the solid wall of the reservoir is that the velocity vanishes at the solid boundary, i.e., vi = 0, on the solid wall. The computation is based on a SIMPLEC algorithm for the solutions of velocity and pressure fields. It is coupled with volume of fluid (VOF) and piecewise linear interface calculation (PLIC) techniques for the transport of volume fractions of liquid in the cells, and construction of the interface. For the treatment of surface tension effects, a continuum surface force (CSF) model is employed. 5. Results and discussions 5.1. Quantitative analysis of ejected droplet size distribution Fig. 4 presents distribution histograms of the dimensions of sprayed particles of the dilute alcohol. The Sauter mean diameter (SMD) is one of the key parameters to estimate the atomization quality, which is often of use in applications where the active surface or surface area is important. It gives the mean value in terms of volume/surface ratio. The ejected particle distribution is consistent with the Rosin and Rammler distribution equation [11], i.e.
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Fig. 4. Distribution of the sprayed diluted alcohol particles.
in which pcumD,d is the cumulative percentage of volume for a particle size; d75v and d50v are the diameters for a droplet cumulative percentage volume of 75% and 50%, respectively; and q is the spread factor, which is usually between 1.5 and 4 [11] and that a distribution is more uniform with the increasing q and spread out with the decreasing q. The current experimental results show that the value of q ranges from 2.55 to 3.84 and is related to the driving frequency of the atomizer. The spray ratio (SR), d90v/d10v, is another important parameter which is often used to determine the merit of particle size distribution region, d90v and d10v are the diameters for a droplet cumulative percentage volume of 90% and 10%, respectively. Basically, the smaller the SR value is, the more uniform the distribution of the ejected droplets size will be, which means that the design and fabrication of the nebulizer should be close to the lower SR to acquire high atomization effectiveness. According to the literature relating to the commercial nebulizer, SR generally lies about in the value of 10.65 [12], the present experimental results revealed the value of SR of our nebulizer is 3.04.
At this particular frequency, the displacement of the piezoelectric actuator, and hence the volume change within the liquid reservoir, is maximized and therefore the flow rate of the ejected droplets is enhanced. Prior to the experimental spray tests, the vibration characteristics of five different piezoelectric actuators were examined to investigate this resonant frequency phenomenon. Fig. 5 illustrates the relationship between the flow rate and the operating frequency when using dilute alcohol as the testing fluid. In this case, the bimorph structure was driven by a sine wave with an operating voltage of 20 Vpk–pk, i.e. ±10 V, and the frequency was increased incrementally from 4 to 60 kHz. It can be seen that the flow rate is maximized in two particular frequency ranges, i.e., one lower frequency range and one higher frequency range. These 1.8
Flow Rate (c.c./min)
q d pcumD;d ¼ 1 exp ð ln 2Þ d 50v ln 2 q¼ lnðd 75v Þ lnðd 50v Þ
1.2
0.6
5.2. Effect of operating frequency on flow rate
0
A previous study [2] has reported that an optimum operating frequency exists for piezoelectric type micronozzles.
Fig. 5. Variation of flow rate with driving frequency.
2
18
34 Frequency (KHz)
50
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Fig. 6. SEM images of spray distributions with and without resonant frequency driving conditions.
is approximately 29.6% of the total liquid volume of the reservoir. Finally, at a driving frequency of 50 kHz, the injection flow per minute is approximately 31% of the total liquid volume of the reservoir. Although a reasonable injection performance is obtained at a driving frequency of 50 kHz, it has been reported that a higher operating frequency prompts the accumulation of gas bubbles within the reservoir, thereby adversely affects the injection performance [13]. Therefore, high values of the operating frequency are not recommended.
5.3. Effect of fluid viscosity on spray flow The results above have shown that the optimum injection flow rate is obtained at an operating frequency of approximately 18 kHz. Accordingly, applying this operating frequency and a driving voltage of 40 Vpk–pk (the maximum voltage the piezoelectric actuator is able to safely withstand), this study investigated the influence of the fluid viscosity on the flow rate using a dilute alcohol with viscosity of 0.29 cP and two other fluids with viscosity of 5.9 cP and 9.8 cP. The corresponding experimental results are presented in Fig. 8. It can be seen that the flow rate reduces with an increasing viscosity. This result is reasonable since a higher viscosity increases the resistance to motion of the liquid and hence reduces its momentum [14].
42
1.5
Experiment
40
Simulation
Injection Mass (g)
Percentage of reservoir volume (%)
frequencies correspond to the resonant frequency ranges of the bimorph structure and indicate appropriate operating conditions for the current micronozzle. The greatest flow rate is obtained in the lower frequency range, which extends from approximately 8 to 20 kHz. The average flow rate in this range is found to be 1.6 cm3/min, while the maximum flow rate, i.e. 1.8 cm3/min, is obtained at a frequency of approximately 18 kHz. In the higher frequency range, the maximum flow rate occurs at a frequency of approximately 50 kHz. According to a previous study [3], the injection volume is maximized when the piezoelectric actuator vibrates at its resonant frequency. This finding is confirmed by the present results shown in Fig. 6. Although, the flow rate seems not sharp like many other nebulizer systems, this may be because the air bubbles ingestion occurs in the reservoir, as well as the piezoelectric plate is immersed within the liquid so that the resistive force of vibration becomes larger. Fig. 7 plots the experimental and simulation results obtained for the injection performance at operating frequencies of 20 kHz, 40 kHz, and 50 kHz, respectively, using dilute alcohol as the testing liquid. Both sets of results indicate that the injection performance is improved at driving frequencies of 20 kHz and 50 kHz. Specifically, at a driving frequency of 20 kHz, the simulated injection flow per minute is approximately 34% of the total liquid volume of the reservoir. At a driving frequency of 40 kHz, the injection flow per minute
38 36 34 32
Liquid No 1 5.9 cP
1.2
Liquid No 2 9.8 cP
0.9 0.6 0.3
30 28 15
Dilute alcohol 0.29 cP
20
25
30
35
40
45
50
55
Frequency (KHz) Fig. 7. Experimental and numerical results for injection performance under near-resonant frequency conditions.
0
15
30
45
60
Time (sec) Fig. 8. Variation of flow rate with viscosity for different fluids with injection time for fluids of different viscosities.
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cal results are in a good agreement with the experimental data.
5.4. Effect of fluid viscosity on particle size In the current particle size analysis experiments, the operating voltage was fixed at 40 Vpk–pk, and the operating frequency was specified as 12 kHz, 14 kHz, 16 kHz, 18 kHz, and 20 kHz, respectively. Fig. 9 presents the particle size results obtained from three kinds of liquid with different viscosity, i.e., water, dilute alcohol, and Stoddard solvent. The results indicate that as the viscosity of the injected liquid increases, the average particle sizes also increase. This finding is consistent with what is reported in the literature [15]. In a previous study, their experimental results indicated the Sauter mean diameter (SMD) increases with the increasing viscosity of the ejection fluid [16]. For a given test fluid, Fig. 9 shows that the average particle size gradually decreases as the operating frequency is increased. The numerical results presented in Fig. 10 reveal that at an operating frequency of 20 kHz, the particle sizes of the Stoddard solvent (the highest viscosity of our tested fluid) are larger than those of the dilute alcohol sample (the lowest viscosity test fluid). Hence, the numeri-
5.5. Effect of operating voltage on particle size Using water as the test fluid, the operating frequency was fixed at 14 kHz and 16 kHz, respectively, and particle size analysis was carried out at different operating voltages in the range 14 Vpk–pk–40 Vpk–pk. When the piezoelectric bimorph structure deflects, the displacement of the free 2 end is described by the equation X ¼ 3ðTL Þ d 31 V , where X is the displacement of the free end, L and T are the length and thickness of the piezoelectric structure, respectively, d31 is the piezoelectric constant, and V is the magnitude of the applied voltage. Clearly, as the voltage is increased, the displacement of the free end also increases, and hence a greater mechanical energy is produced. Therefore, it is reasonable to assume that the size of the atomized particles will reduce. From Fig. 11, it can be seen that the particles do indeed become smaller as the voltage is increased. 5.6. Effect of liquid surface tension
55
Stoddard solvent 84 cP Water 1 cP Dilute alcohol 0.29 cP
SMD Particle Size (μm)
50 45 40 35 30 25 20 10
12
14
16
18
20
22
Frequency (KHz) Fig. 9. Variation of particle size with driving frequency for different fluids.
In investigating the effect of the liquid surface tension on the performance of the present nebulizer, three different liquids at room temperature were considered, namely water, a honey-like fluid, and an olive oil-like fluid, with surface tensions of 0.0725 (N/m), 0.06 (N/m), and 0.031 (N/m), respectively. At an operating frequency of 20 kHz, it was appeared that fluids with a higher surface tension tended to accumulate around the nozzles in the injection orifice plate and were not readily separated. Therefore, as these droplets were ejected, their movement in the forward direction was restrained to a greater extent by the ligament issuing from the orifice, and hence the ejection effect was adversely affected. Conversely, for a fluid with a lower surface tension, e.g., the olive oil-like fluid, the atomization
Fig. 10. Comparison of particle size for fluids of different viscosities at constant driving frequency of 20 kHz.
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SMD Particle Size (μm)
Frequency=14 kHz 35
Frequency=16 kHz
5.7. Air ingestion phenomenon
30
10
consistent with the pressure atomizer using water and kerosene [17] as test fluids in which it was found that the atomized SMD increased as the surface tension of the liquid increased.
20
30
40
Volt
Fig. 11. Variation of particle size with driving voltage for different driving frequencies.
effect was improved and the ejection distance of the ejected droplets increased, as shown in Fig. 12. These results are
As the droplets are squeezed through the injection orifice plate, they initially have a columnar form with a uniform cross-section. However, as they emerge from the orifice, their shape gradually evolves into a teardrop-like form with a spherical front surface and a long thin tail. Subsequently, the droplet breaks away from the nozzle exit as the break-off phenomenon occurs. At the moment that the droplets break free of the orifice, surface tension effects cause the ligament to snap back, causing the ingestion of
Fig. 12. Comparison of ejection distance for fluids with different surface tension effects under driving frequency of 20 kHz.
Fig. 13. Comparison of air ingestion effect for different driving frequencies.
Y.R. Jeng et al. / Experimental Thermal and Fluid Science 31 (2007) 1147–1156
Air ingestion/reservoir volume (%)
air into the reservoir. This ingestion effect is reinforced by the deflection of the piezoelectric actuator in the backward direction, which generates a negative pressure effect [18]. Figs. 10 and 13 indicate that the air ingestion phenomenon
5 4 3 2 1 0 0
10
20
30 40 50 Viscosity (cP)
60
70
1155
is more pronounced at lower viscosity and higher driving frequencies. It shows that the ingested air increases with the reduction of fluid viscosity. This is reasonable since a lower viscosity reduces the break up time of the fluid in the exit of the nozzle, and hence the amount of the ingested air per unit time rises. The relationship between the ingested air and the viscosity of fluid is shown in Fig. 14. The simulation results also reveal that as the operating frequency increases, the amount of the ingested air also goes up. That is the reason that the faster the piezoelectric plate vibrates, the larger the negative pressure the reservoir has. Therefore, the air could be easily ingested into the reservoir, as shown in Fig. 15. Since the accumulation of air within the reservoir has a detrimental effect on the injection performance, and, in severe cases, can block the injection orifices, a lower operating frequency is desirable, as discussed previously.
80
5.8. Orifice number effect
Air ingestion/reservoir volume (%)
Fig. 14. Variation of air ingestion for different viscosities.
Fig. 16 illustrates the ejection distance of the ejected droplets for an orifice plate with two orifices and six orifices, respectively. It is observed that the ejection distance is reduced as the number of orifices is increased since the momentum of the droplets is reduced and they more easily collide and coalesce in the air before reaching the recording medium. In general, increasing the operating frequency tends to reduce the severity of this coalescence phenomenon.
4
2
6. Conclusions 0
0
10
20 30 Frequency (KHz)
40
50
Fig. 15. Variation of air ingestion for different driving frequencies.
The piezoelectrically actuated nebulizer with an easily assembled structure has successfully been fabricated and tested. From the experimental results, we find that the ejected particle from the developed nebulizer and the
Fig. 16. Comparison of ejection distance for orifice plates with different numbers of orifices at driving frequency of 20 kHz.
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corresponding lower value of Spray Ratio (SR) exhibits the characteristic of uniform particle size distribution. This distribution is consistent with the Rosin and Rammler distribution equation. Based on the results obtained from a series of numerical and experimental investigations, this study establishes the optimal operating conditions for the proposed nebulizer. According to the investigations, the diameter of particles can be reduced by increasing the operating frequency, but the greater volume of air will be introduced into the reservoir to degrade the ejection performance. The ejection flow rate can be maximized by setting the frequency of the operating voltage at the resonant frequency of the piezoelectric actuator. As the viscosity of the ejected liquid increases, the flow rate decreases and the average particle size becomes larger for a fixed operating frequency. Increasing the driving voltage will be helpful to the amount of atomization. Moreover, the droplets tend to accumulate around the orifices on orifice plate when the surface tension of the ejected liquid increases. This phenomenon can be lessened by boosting the operating frequency for a given liquid. References [1] W. Peter, Micropumps-past progress and future prospects, Sens. Actuators B 105 (1) (2005) 28–38. [2] L. Cao, S. Mantell, D. Polla, Design and simulation of an implantable medical drug delivery system using microelectromechanical systems technology, Sens. Actuators A 94 (1) (2001) 117–125. [3] P. Dario, N. Croce, M.C. Carrozza, G. Varallo, A fluid handling system for a chemical microanalyzer, J. Micromech. Microeng. 6 (1996) 95–98. [4] J. Kan, Z. Yang, T. Peng, G. Cheng, B. Wu, Design and test of a high-performance piezoelectric micropump for drug delivery, Sens. Actuators A 121 (2005) 156–161.
[5] W.T. Pimbley, Drop formation from a liquid jet: a linear onedimensional analysis considered as a boundary value problem, IBM J. Res. Develop. 20 (1976) 148–156. [6] J.E. Fromm, Numerical calculations of the fluid dynamics of dropon-demand jets, IBM J. Res. Develop. 28 (3) (1984) 322–333. [7] H. Asai, H. Toshitami, E. Ichiro, One-dimensional model of bubble growth and liquid flow in bubble jet printers, J. Japan Soc. Appl. Phys. 26 (10) (1987) 1794–1801. [8] H. Asai, Three-dimensional calculation of bubble growth and drop ejection in a bubble jet printer, ASME, J. Fluids Eng. 114 (4) (1992) 638–641. [9] T.M. Liou, S.W. Chau, S.C. Chen, K.C. Shih, Numerical investigation of droplet behavior in inkjet printing process, in: Proceedings of the 9th National Computational Fluid Dynamics Conference, Taiwan, August, 2002. [10] J.T. Yeh, A VOF-FEM and Coupled Inkjet Simulation, in: Proceedings of ASME Fluids Engineering Division Summer Meeting, FEDSM2001-18182, New Orleans, LA, 2001. [11] D.N. Ralph, Dispersing Powders in Liquids Part 1. Available from:. [12] S.C. Shen, Y.W. Wang, Y.C. Chen, The new micro-atomizer, J. Mechatron. Ind. 271 (2005) 1–6. [13] I. Ederer, P. Raetsch, W. Schullerus, C. Tille, U. Zech, Piezoelectrically driven micropump for on-demand fuel-drop generation in an automobile heater with continuously adjustable power output, Sens. Actuators A 62 (1) (1997) 752–755. [14] I. Ederer, J. Grasegger, C. Tille, Droplet generator with extraordinary high flow rate and wide operating range, in: Proceedings of Transducer’97 conference, IEEE, Chicago, 1997, pp. 809–812. [15] H.C. Simmons, The prediction of sauter mean diameter for gas turbine fuel nozzles of different types, ASME, J Eng Power 102 (1980) 646–652. [16] A.R. Jones, Design optimization of a large pressure-jet atomizer, in: Proceedings of Power Plant 2nd International Conference, Liquid Atomization and Sprays System, Madison WI, 1982, pp. 181–185. [17] H.C. Simmons, C.F. Harding, Some effects on using water as a test fluid in fuel nozzle spray analysis, ASME Paper 80-GT-90, 1980. [18] J.H. Park, K. Yoshida, S. Yokoto, Resonantly driven piezoelectric micropump-fabrication of a micropump having high power density, Mechatronics 9 (7) (1999) 687–702.
An investigation into a piezoelectrically actuated nebulizer with lEDM-made micronozzle array Y.R. Jeng
a,* ,
C.C. Su
a,b
, G.H. Feng a, Y.Y. Peng
c
a
Department of Mechanical Engineering, National Chung Cheng University, Ming-Hsiung, Chia-Yi 621, Taiwan Department of Mechanical Engineering, WuFeng Institute of Technology, Ming-Hsiung, Chia-Yi 621, Taiwan c Industrial Technology Research Institute, Mechanical Industry Research Laboratories, Hsin-Chu, 310, Taiwan b
Received 5 June 2006; received in revised form 11 December 2006; accepted 18 December 2006
Abstract This study uses micromachining techniques to fabricate a piezoelectrically actuated nebulizer with an easily assembled structure. The experimental measurement and numerical simulation have been performed to gain the insight of atomized droplet behavior and fluidic mechanics for the developed devices. It is expected that a better understanding of these physical insights can lead to various potential applications. Based on the results obtained from a series of investigations, this study has found the optimal operating conditions for the proposed nebulizer. Various test fluids with different viscosities and surface tensions are considered, including water, dilute alcohol, oil, and a honey-like fluid. A spray stability analysis is performed in which the operating voltage and frequency of the piezoelectric actuator in the nebulizer are varied and the corresponding changes in the ejected particle size, flow rate and ejection distance examined. Air trapped within the nebulizer reservoir adversely affects the injection performance. Hence, the effect of the operating parameters on the volume of air ingested into the reservoir during the nebulizer operation is investigated and discussed. Ó 2006 Elsevier Inc. All rights reserved. Keywords: Micronozzle; Piezoelectrically actuated nebulizer; lEDM
1. Introduction Rapid advances in micro-electro-mechanical-systems (MEMS) technologies have facilitated the development of microfluidic systems for a diverse range of applications in many scientific and engineering fields [1–4]. In piezoelectric nebulizer, the atomization and flow rates of the ejected droplets are governed by the interactions between the viscosity and surface tension of the ejected liquid and the pressure developed within the reservoir. In operation, a voltage is applied to the piezoelectric actuator. The corresponding
* Corresponding author. Present address: Department of Mechanical Engineering, National Chung Cheng University, 160, San-Hsing, MingShiung, Chia-Yi 621, Taiwan. Tel./fax: +886 5 2428189. E-mail address: [email protected] (Y.R. Jeng).
0894-1777/$ - see front matter Ó 2006 Elsevier Inc. All rights reserved. doi:10.1016/j.expthermflusci.2006.12.001
deflection of the piezoelectric actuator produces a volume change in the reservoir and this causes an acoustic pressure wave to propagate towards the nozzles, forcing droplets to be expelled through the orifice plate. When the direction of the applied voltage is reversed, the piezoelectric actuator deflects in the opposite direction. However, the negative pressure created by the backward deflection of the piezoelectric actuator may draw some external air into the reservoir. In the present study, experimental and numerical investigations are performed to investigate the ingestion of air bubbles into the reservoir and to examine the size of the ejected droplets together with their flow rate and ejection distance under different operating conditions. The present results are intended as a reference for the development of microparticle sprays in various applications, including medical liquid sprays, methanol injectors in direct methanol fuel cell (DMFCs), and odor generators, etc.
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Nomenclature f g L t T p u v V Vpk–pk X w q qL
liquid volume fraction gravity acceleration (m/s2) length (m) time (s) thickness (m) fluid pressure (N/m2) velocity in the x-direction velocity in the y-direction voltage (V) voltage for peak to peak (V) displacement (m) velocity in the z-direction mixing density (kg/m3) density of the liquid fraction
The main function of an atomization nozzle is to break a bulk volume of liquid into many small droplets for increasing the surface to volume ratio of the liquid. The changes which occur in the spray effect of an atomization nozzle under different applied voltages and frequencies depend on the viscosity and surface tension of the ejected fluid. In the previous study, Pimbley [5] applied linear one-dimensional analysis and the boundary value perturbation method to investigate droplet formation. Fromm [6] introduced a two-dimensional flow line governing equation to investigate the effect of the fluid behavior on the droplet structure and velocity under different Reynolds numbers and driving pressures. The results were then used to estimate the corresponding variations in the ejected droplet stream. Asai et al. [7] conducted numerical simulations and experimental studies to investigate variations in the droplet stream length and developed a one-dimensional numerical model to simulate the droplet formation. Later, Asai [8] applied the finite difference method to solve the three-dimensional Navier– Stokes equation, and predicted the path of ink droplets in a side-direction hot bubble type inkjet printer. The numerical results showed that the ink droplet volume and velocity increased as the pressure applied to the bubble increased, but reduced as the viscosity and surface tension of the ink increased. Liou et al. [9] applied the finite volume method to discretize the flow field governing equation taking the surface tension effect into account and to predict the free interface of the liquid and gas. Moreover, Yeh [10] used the fractional volume of fluid (VOF) and finite element method (FEM) approaches to investigate the development of the free interface of the liquid and gas during fluid ejection in a piezoelectric type inkjet. Performing experimental observations of the variation in size of droplets ejected through the orifices of a nebulizer is challenging. Hence, this study employs computational fluid dynamics software (CFDRC) to assist the droplet stream analysis and to investigate the ingress of air into the reservoir body during the injection process.
qG l Fr ~ V /L /G / rij sij m k dij
density of the gas fraction dynamic viscosity (kg/m s) surface tension force (N/m) velocity vector (m/s) the value of the quantity for liquid the value of the quantity for gas volume averaged quantity stress tensor shear stress tensor kinematic viscosity second coefficient of viscosity (depend on density of fluid) kronecker delta, when i = j, dij = 1; i 5 j, dij = 0
2. Piezoelectrically actuated nebulizer structure with micronozzles The fabricated nebulizer consists of a piezoelectric actuator, an injection orifice-array plate, a front cover plate, and the reservoir, as shown in Fig. 1. In this study, the nebulizer is made of polyoxymethylene (POM), and incorporates a body with a inner volume of approximately 4 cm3. The function of the piezoelectric actuator, which is located between the injection orifice plate and the reservoir, is to transform the electric energy supplied by an external electrical pulse into mechanical energy. When a voltage is applied to the piezoelectric actuator, it deflects and generates the pressure on the liquid within the reservoir. A deflection of the piezoelectric actuator in the forward direction causes the liquid in the reservoir to be squeezed through the array of orifices in the injection orifice plate, as shown in Fig. 2. The present experiments were conducted using a bimorph ceramic piezoelectric actuator fabricated from lead zirconium titanate (PZT). The bimorph structure was fabricated using a steel film of thickness 100 lm pressed on either side with a PZT ceramic piezoelectric actuator of thickness 200 lm. Therefore, the total thickness of the bimorph piezoelectric structure was 0.5 mm. The length and width of the piezoelectric structure were 20 mm and 10 mm, respectively. In this bimorph structure, the external piezoelectric films have one polarity, while the steel film has the reverse polarity. Hence, when an electrical voltage is applied to the bimorph structure, the piezoelectric actuator on one side of the structure expands, while that on the other side contracts, causing a deflection of the bimorph structure. The injection orifice plate of our present study is designed to the form of conical orifices-array. Fabrication of the injection orifice plate commenced by using a Nd-YAG laser to drill circular straight holes of diameter 50 lm through a stainless steel plate of thickness 50 lm. A micro electrostatic discharge machining (lEDM)
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liquid replenishment tube
injection orifice-array plate
reservoir body
sealing ring
cover plate
piezoelectric actuator
Fig. 1. Schematic illustration of piezoelectric nebulizer.
Fig. 2. Schematic illustration of the working principle of the designed nebulizer.
technique was then used to machine the circular holes into a conical profile, with an entrance diameter of 100 lm and
an exit diameter of 50 lm. As shown in Fig. 3, the injection orifice plate was fabricated with two parallel rows of orifices, with each row containing 25 orifices. 3. Atomized particle size measurement
Fig. 3. Schematic illustration and SEM image of injection orifice plate.
The current particle size measurements were carried out by using an INSITEC particle size analyzer. The analysis system comprised an optical arrangement, a detector module, and data collection and analysis software installed on a PC. The laser light was generated by a luminous diode and passed though a beam expander from which it emerged as a parallel light beam with a diameter of 9 mm. The light beam was passed through the droplets in the atomization flow field causing a diffraction effect to take place. The diffracted light was collected by a Fourier Lens and passed to a photo detector consisting of 31 ring-type light intensity sensors. The sensors detected changes in the diffracted light pattern caused by variations in the spray ejected from the injection orifice plate. The detected light signal was transformed into an electrical signal and, following signal amplification, further transformed into a digital signal by an analog-digital converter (A/D converter). Finally, the
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resulting signal was transported to a PC for particle sizing analysis.
4.2. Boundary condition setting The boundary conditions in the procedure of calculation include two parts: One is liquid in the reservoir, and the other is liquid left the reservoir to form the droplets.
4. Theoretical analysis 4.1. Governing equation The continuity equation is given by: oq ~Þ ¼ 0 þ r ðqV ot
ð1Þ
The momentum equation is given by: o ~u ¼ op þ o 2l ou þ o l ov þ ou ðquÞ þ r qV ot ox ox ox oy ox oy o ou ow þ l þ þ F rx þ qgx oz oz ox o op o ov ou o ov ðqvÞ þ r qV~v ¼ þ l þ þ 2l ot oy ox ox oy oy oy o ov ow l þ þ F ry þ qgy þ oz oz oy o op o ou ow o ov ow ðqwÞ þ r qV~w ¼ þ l þ þ l þ ot oz ox oz ox oy oz oy o ow 2l þ F rz þ qgz þ oz oz
ð2Þ
where u, v and w, represent the velocity in the x- y- and zdirections, p, q and l represent the pressure, density, and viscosity coefficients, respectively, Fr is the surface tension force acting at the liquid/gas interface and qg is the gravity term. The governing equations given above yield the velocity and pressure distribution of the system. The volume of fluid (VOF) method is then applied to obtain the distribution of the flow field volume fraction in order to establish the path of the ejected droplets. For given a flow field and an initial distribution of f on a grid, the manner in which the volume fraction distribution f evolves is determined by solving the passive transport equation: of þ r V~f ¼ 0 ot
ð3Þ
where f is the liquid volume fraction, t is time, $ is the standard spatial ‘‘grad’’ perator, and V~ is the velocity vector. When two different fluids exist simultaneously in the computational grid, they can be regarded as a single fluid for the purposes of establishing the physical property of the flow field in the grid. The physical property of the flow field is determined via a process of linear interpolation based on the liquid state volume fractions of the two fluids in the grid, i.e. /¼
f qL /L þ ð1 f ÞqG /G q
ð4Þ
where /L and /G indicate the value of the liquid and gas of the physical property, respectively, and q is the mixing density, given by q = fqL + (1 f)qG.
(a) The free surface boundary condition. The boundary condition in the free surface when the droplets ejected for the nebulizer are as follows: Because the stress at the free surface of two fluids must be equilibrated, the magnitude of the stress tensor for moving fluid can be expressed as rij ¼ P dij þ sij oui ouj ouk þ And sij ¼ m þk dij oxj oxi ouk
ð5Þ ð6Þ
We assume the fluid is an incompressible flow, and then k = 0. Consequently, the stress tensor ri,j can be obtained as oui ouj þ rij ¼ P dij þ m ð7Þ oxj oui On the free surface boundary, the stress in the normal and tangential direction must be in equilibrium, these conditions can be expressed as ðrij nÞ n ¼ 0 ðrij nÞ m1 ¼ 0
ð8Þ ð9Þ
ðrij nÞ m2 ¼ 0
ð10Þ
(b) The solid wall boundary condition. The boundary condition on the solid wall of the reservoir is that the velocity vanishes at the solid boundary, i.e., vi = 0, on the solid wall. The computation is based on a SIMPLEC algorithm for the solutions of velocity and pressure fields. It is coupled with volume of fluid (VOF) and piecewise linear interface calculation (PLIC) techniques for the transport of volume fractions of liquid in the cells, and construction of the interface. For the treatment of surface tension effects, a continuum surface force (CSF) model is employed. 5. Results and discussions 5.1. Quantitative analysis of ejected droplet size distribution Fig. 4 presents distribution histograms of the dimensions of sprayed particles of the dilute alcohol. The Sauter mean diameter (SMD) is one of the key parameters to estimate the atomization quality, which is often of use in applications where the active surface or surface area is important. It gives the mean value in terms of volume/surface ratio. The ejected particle distribution is consistent with the Rosin and Rammler distribution equation [11], i.e.
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Fig. 4. Distribution of the sprayed diluted alcohol particles.
in which pcumD,d is the cumulative percentage of volume for a particle size; d75v and d50v are the diameters for a droplet cumulative percentage volume of 75% and 50%, respectively; and q is the spread factor, which is usually between 1.5 and 4 [11] and that a distribution is more uniform with the increasing q and spread out with the decreasing q. The current experimental results show that the value of q ranges from 2.55 to 3.84 and is related to the driving frequency of the atomizer. The spray ratio (SR), d90v/d10v, is another important parameter which is often used to determine the merit of particle size distribution region, d90v and d10v are the diameters for a droplet cumulative percentage volume of 90% and 10%, respectively. Basically, the smaller the SR value is, the more uniform the distribution of the ejected droplets size will be, which means that the design and fabrication of the nebulizer should be close to the lower SR to acquire high atomization effectiveness. According to the literature relating to the commercial nebulizer, SR generally lies about in the value of 10.65 [12], the present experimental results revealed the value of SR of our nebulizer is 3.04.
At this particular frequency, the displacement of the piezoelectric actuator, and hence the volume change within the liquid reservoir, is maximized and therefore the flow rate of the ejected droplets is enhanced. Prior to the experimental spray tests, the vibration characteristics of five different piezoelectric actuators were examined to investigate this resonant frequency phenomenon. Fig. 5 illustrates the relationship between the flow rate and the operating frequency when using dilute alcohol as the testing fluid. In this case, the bimorph structure was driven by a sine wave with an operating voltage of 20 Vpk–pk, i.e. ±10 V, and the frequency was increased incrementally from 4 to 60 kHz. It can be seen that the flow rate is maximized in two particular frequency ranges, i.e., one lower frequency range and one higher frequency range. These 1.8
Flow Rate (c.c./min)
q d pcumD;d ¼ 1 exp ð ln 2Þ d 50v ln 2 q¼ lnðd 75v Þ lnðd 50v Þ
1.2
0.6
5.2. Effect of operating frequency on flow rate
0
A previous study [2] has reported that an optimum operating frequency exists for piezoelectric type micronozzles.
Fig. 5. Variation of flow rate with driving frequency.
2
18
34 Frequency (KHz)
50
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Fig. 6. SEM images of spray distributions with and without resonant frequency driving conditions.
is approximately 29.6% of the total liquid volume of the reservoir. Finally, at a driving frequency of 50 kHz, the injection flow per minute is approximately 31% of the total liquid volume of the reservoir. Although a reasonable injection performance is obtained at a driving frequency of 50 kHz, it has been reported that a higher operating frequency prompts the accumulation of gas bubbles within the reservoir, thereby adversely affects the injection performance [13]. Therefore, high values of the operating frequency are not recommended.
5.3. Effect of fluid viscosity on spray flow The results above have shown that the optimum injection flow rate is obtained at an operating frequency of approximately 18 kHz. Accordingly, applying this operating frequency and a driving voltage of 40 Vpk–pk (the maximum voltage the piezoelectric actuator is able to safely withstand), this study investigated the influence of the fluid viscosity on the flow rate using a dilute alcohol with viscosity of 0.29 cP and two other fluids with viscosity of 5.9 cP and 9.8 cP. The corresponding experimental results are presented in Fig. 8. It can be seen that the flow rate reduces with an increasing viscosity. This result is reasonable since a higher viscosity increases the resistance to motion of the liquid and hence reduces its momentum [14].
42
1.5
Experiment
40
Simulation
Injection Mass (g)
Percentage of reservoir volume (%)
frequencies correspond to the resonant frequency ranges of the bimorph structure and indicate appropriate operating conditions for the current micronozzle. The greatest flow rate is obtained in the lower frequency range, which extends from approximately 8 to 20 kHz. The average flow rate in this range is found to be 1.6 cm3/min, while the maximum flow rate, i.e. 1.8 cm3/min, is obtained at a frequency of approximately 18 kHz. In the higher frequency range, the maximum flow rate occurs at a frequency of approximately 50 kHz. According to a previous study [3], the injection volume is maximized when the piezoelectric actuator vibrates at its resonant frequency. This finding is confirmed by the present results shown in Fig. 6. Although, the flow rate seems not sharp like many other nebulizer systems, this may be because the air bubbles ingestion occurs in the reservoir, as well as the piezoelectric plate is immersed within the liquid so that the resistive force of vibration becomes larger. Fig. 7 plots the experimental and simulation results obtained for the injection performance at operating frequencies of 20 kHz, 40 kHz, and 50 kHz, respectively, using dilute alcohol as the testing liquid. Both sets of results indicate that the injection performance is improved at driving frequencies of 20 kHz and 50 kHz. Specifically, at a driving frequency of 20 kHz, the simulated injection flow per minute is approximately 34% of the total liquid volume of the reservoir. At a driving frequency of 40 kHz, the injection flow per minute
38 36 34 32
Liquid No 1 5.9 cP
1.2
Liquid No 2 9.8 cP
0.9 0.6 0.3
30 28 15
Dilute alcohol 0.29 cP
20
25
30
35
40
45
50
55
Frequency (KHz) Fig. 7. Experimental and numerical results for injection performance under near-resonant frequency conditions.
0
15
30
45
60
Time (sec) Fig. 8. Variation of flow rate with viscosity for different fluids with injection time for fluids of different viscosities.
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cal results are in a good agreement with the experimental data.
5.4. Effect of fluid viscosity on particle size In the current particle size analysis experiments, the operating voltage was fixed at 40 Vpk–pk, and the operating frequency was specified as 12 kHz, 14 kHz, 16 kHz, 18 kHz, and 20 kHz, respectively. Fig. 9 presents the particle size results obtained from three kinds of liquid with different viscosity, i.e., water, dilute alcohol, and Stoddard solvent. The results indicate that as the viscosity of the injected liquid increases, the average particle sizes also increase. This finding is consistent with what is reported in the literature [15]. In a previous study, their experimental results indicated the Sauter mean diameter (SMD) increases with the increasing viscosity of the ejection fluid [16]. For a given test fluid, Fig. 9 shows that the average particle size gradually decreases as the operating frequency is increased. The numerical results presented in Fig. 10 reveal that at an operating frequency of 20 kHz, the particle sizes of the Stoddard solvent (the highest viscosity of our tested fluid) are larger than those of the dilute alcohol sample (the lowest viscosity test fluid). Hence, the numeri-
5.5. Effect of operating voltage on particle size Using water as the test fluid, the operating frequency was fixed at 14 kHz and 16 kHz, respectively, and particle size analysis was carried out at different operating voltages in the range 14 Vpk–pk–40 Vpk–pk. When the piezoelectric bimorph structure deflects, the displacement of the free 2 end is described by the equation X ¼ 3ðTL Þ d 31 V , where X is the displacement of the free end, L and T are the length and thickness of the piezoelectric structure, respectively, d31 is the piezoelectric constant, and V is the magnitude of the applied voltage. Clearly, as the voltage is increased, the displacement of the free end also increases, and hence a greater mechanical energy is produced. Therefore, it is reasonable to assume that the size of the atomized particles will reduce. From Fig. 11, it can be seen that the particles do indeed become smaller as the voltage is increased. 5.6. Effect of liquid surface tension
55
Stoddard solvent 84 cP Water 1 cP Dilute alcohol 0.29 cP
SMD Particle Size (μm)
50 45 40 35 30 25 20 10
12
14
16
18
20
22
Frequency (KHz) Fig. 9. Variation of particle size with driving frequency for different fluids.
In investigating the effect of the liquid surface tension on the performance of the present nebulizer, three different liquids at room temperature were considered, namely water, a honey-like fluid, and an olive oil-like fluid, with surface tensions of 0.0725 (N/m), 0.06 (N/m), and 0.031 (N/m), respectively. At an operating frequency of 20 kHz, it was appeared that fluids with a higher surface tension tended to accumulate around the nozzles in the injection orifice plate and were not readily separated. Therefore, as these droplets were ejected, their movement in the forward direction was restrained to a greater extent by the ligament issuing from the orifice, and hence the ejection effect was adversely affected. Conversely, for a fluid with a lower surface tension, e.g., the olive oil-like fluid, the atomization
Fig. 10. Comparison of particle size for fluids of different viscosities at constant driving frequency of 20 kHz.
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SMD Particle Size (μm)
Frequency=14 kHz 35
Frequency=16 kHz
5.7. Air ingestion phenomenon
30
10
consistent with the pressure atomizer using water and kerosene [17] as test fluids in which it was found that the atomized SMD increased as the surface tension of the liquid increased.
20
30
40
Volt
Fig. 11. Variation of particle size with driving voltage for different driving frequencies.
effect was improved and the ejection distance of the ejected droplets increased, as shown in Fig. 12. These results are
As the droplets are squeezed through the injection orifice plate, they initially have a columnar form with a uniform cross-section. However, as they emerge from the orifice, their shape gradually evolves into a teardrop-like form with a spherical front surface and a long thin tail. Subsequently, the droplet breaks away from the nozzle exit as the break-off phenomenon occurs. At the moment that the droplets break free of the orifice, surface tension effects cause the ligament to snap back, causing the ingestion of
Fig. 12. Comparison of ejection distance for fluids with different surface tension effects under driving frequency of 20 kHz.
Fig. 13. Comparison of air ingestion effect for different driving frequencies.
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Air ingestion/reservoir volume (%)
air into the reservoir. This ingestion effect is reinforced by the deflection of the piezoelectric actuator in the backward direction, which generates a negative pressure effect [18]. Figs. 10 and 13 indicate that the air ingestion phenomenon
5 4 3 2 1 0 0
10
20
30 40 50 Viscosity (cP)
60
70
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is more pronounced at lower viscosity and higher driving frequencies. It shows that the ingested air increases with the reduction of fluid viscosity. This is reasonable since a lower viscosity reduces the break up time of the fluid in the exit of the nozzle, and hence the amount of the ingested air per unit time rises. The relationship between the ingested air and the viscosity of fluid is shown in Fig. 14. The simulation results also reveal that as the operating frequency increases, the amount of the ingested air also goes up. That is the reason that the faster the piezoelectric plate vibrates, the larger the negative pressure the reservoir has. Therefore, the air could be easily ingested into the reservoir, as shown in Fig. 15. Since the accumulation of air within the reservoir has a detrimental effect on the injection performance, and, in severe cases, can block the injection orifices, a lower operating frequency is desirable, as discussed previously.
80
5.8. Orifice number effect
Air ingestion/reservoir volume (%)
Fig. 14. Variation of air ingestion for different viscosities.
Fig. 16 illustrates the ejection distance of the ejected droplets for an orifice plate with two orifices and six orifices, respectively. It is observed that the ejection distance is reduced as the number of orifices is increased since the momentum of the droplets is reduced and they more easily collide and coalesce in the air before reaching the recording medium. In general, increasing the operating frequency tends to reduce the severity of this coalescence phenomenon.
4
2
6. Conclusions 0
0
10
20 30 Frequency (KHz)
40
50
Fig. 15. Variation of air ingestion for different driving frequencies.
The piezoelectrically actuated nebulizer with an easily assembled structure has successfully been fabricated and tested. From the experimental results, we find that the ejected particle from the developed nebulizer and the
Fig. 16. Comparison of ejection distance for orifice plates with different numbers of orifices at driving frequency of 20 kHz.
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corresponding lower value of Spray Ratio (SR) exhibits the characteristic of uniform particle size distribution. This distribution is consistent with the Rosin and Rammler distribution equation. Based on the results obtained from a series of numerical and experimental investigations, this study establishes the optimal operating conditions for the proposed nebulizer. According to the investigations, the diameter of particles can be reduced by increasing the operating frequency, but the greater volume of air will be introduced into the reservoir to degrade the ejection performance. The ejection flow rate can be maximized by setting the frequency of the operating voltage at the resonant frequency of the piezoelectric actuator. As the viscosity of the ejected liquid increases, the flow rate decreases and the average particle size becomes larger for a fixed operating frequency. Increasing the driving voltage will be helpful to the amount of atomization. Moreover, the droplets tend to accumulate around the orifices on orifice plate when the surface tension of the ejected liquid increases. This phenomenon can be lessened by boosting the operating frequency for a given liquid. References [1] W. Peter, Micropumps-past progress and future prospects, Sens. Actuators B 105 (1) (2005) 28–38. [2] L. Cao, S. Mantell, D. Polla, Design and simulation of an implantable medical drug delivery system using microelectromechanical systems technology, Sens. Actuators A 94 (1) (2001) 117–125. [3] P. Dario, N. Croce, M.C. Carrozza, G. Varallo, A fluid handling system for a chemical microanalyzer, J. Micromech. Microeng. 6 (1996) 95–98. [4] J. Kan, Z. Yang, T. Peng, G. Cheng, B. Wu, Design and test of a high-performance piezoelectric micropump for drug delivery, Sens. Actuators A 121 (2005) 156–161.
[5] W.T. Pimbley, Drop formation from a liquid jet: a linear onedimensional analysis considered as a boundary value problem, IBM J. Res. Develop. 20 (1976) 148–156. [6] J.E. Fromm, Numerical calculations of the fluid dynamics of dropon-demand jets, IBM J. Res. Develop. 28 (3) (1984) 322–333. [7] H. Asai, H. Toshitami, E. Ichiro, One-dimensional model of bubble growth and liquid flow in bubble jet printers, J. Japan Soc. Appl. Phys. 26 (10) (1987) 1794–1801. [8] H. Asai, Three-dimensional calculation of bubble growth and drop ejection in a bubble jet printer, ASME, J. Fluids Eng. 114 (4) (1992) 638–641. [9] T.M. Liou, S.W. Chau, S.C. Chen, K.C. Shih, Numerical investigation of droplet behavior in inkjet printing process, in: Proceedings of the 9th National Computational Fluid Dynamics Conference, Taiwan, August, 2002. [10] J.T. Yeh, A VOF-FEM and Coupled Inkjet Simulation, in: Proceedings of ASME Fluids Engineering Division Summer Meeting, FEDSM2001-18182, New Orleans, LA, 2001. [11] D.N. Ralph, Dispersing Powders in Liquids Part 1. Available from: