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Int. J. Mol. Sci. 2011, 12, 3648-3704; doi:10.3390/ijms12063648
International Journal of
Molecular Sciences
ISSN 1422-0067
Micro Electromechanical Systems (MEMS) Based Microfluidic
Devices for Biomedical Applications
Muhammad Waseem Ashraf *, Shahzadi Tayyaba and Nitin Afzulpurkar
School of Engineering and Technology, Asian Institute of Technology (AIT), Bangkok 12120,
Thailand; E-Mails: [email protected] (S.T.); [email protected] (N.A.)
* Author to whom correspondence should be addressed;
E-Mail: [email protected]; Tel.: +66-8-73516061; Fax: +025245697.
Received: 10 April 2011; in revised form: 3 May 2011 / Accepted: 19 May 2011 /
Published: 7 June 2011
Abstract: Micro Electromechanical Systems (MEMS) based microfluidic devices have
gained popularity in biomedicine field over the last few years. In this paper, a
comprehensive overview of microfluidic devices such as micropumps and microneedles
has been presented for biomedical applications. The aim of this paper is to present the
major features and issues related to micropumps and microneedles, e.g., working
principles, actuation methods, fabrication techniques, construction, performance
parameters, failure analysis, testing, safety issues, applications, commercialization issues
and future prospects. Based on the actuation mechanisms, the micropumps are classified
into two main types, i.e., mechanical and non-mechanical micropumps. Microneedles can
be categorized according to their structure, fabrication process, material, overall shape, tip
shape, size, array density and application. The presented literature review on micropumps
and microneedles will provide comprehensive information for researchers working on
design and development of microfluidic devices for biomedical applications.
Keywords: drug delivery system; microfluidics; micropumps; microneedles
1. Introduction
Microfluidics is a relatively new branch of science and technology which has made extensive
progress in the last few years. Microfluidic systems deal with the fluid flow in diminutive amounts,
Int. J. Mol. Sci. 2011, 12
typically a few microlitres (μL) in a miniaturized system. The main functions performed by these
systems are sample preparation, purification, separation, reaction, transport, immobilization, labeling,
biosensing and detection. Fluid behavior at macro scale is quite different from micro and nano scale.
Factors such as surface tension may become dominant in microfluidic devices. When the size of
biological samples is close to the flow channels or needles through which the samples are transported,
then the sample flow may not be envisaged on the basis of conventional fluidic systems. Considerable
research has been made in recent years in the field of microfluidic components, devices, systems and
fabrication methods. The use of micro and nano electromechanical systems (MEMS and NEMS)
technology has been increasing rapidly to fabricate microfluidic devices for biomedical applications.
Due to MEMS and NEMS technology, the fabrication of miniature size and high performance medical
devices has become practicable to congregate the critical medical requirements like controlled delivery
with negligible side effects, improved bioavailability and therapeutic effectiveness [1,2]. In recent
years, the most important advancement of MEMS and NEMS in biomedicine is microfluidic
transdermal drug delivery (TDD) systems [3]. TDD systems deal with the movement of
pharmaceutical compound through the skin to reach the systemic circulation for subsequent
distribution in the human body [4]. TDD system consists of micropumps, microneedles, reservoir,
micro-flow sensor, blood pressure sensor, and required electronic circuit for necessary operations.
Among them, micropumps and microneedles are the most important components of microfluidic
system particularly for drug delivery applications. Micropumps are used for delivery and treatment
purposes. Microneedles can be used as stand-alone devices and part of complicated microfluidic
system in which microneedles are integrated with other devices in the system. The schematic
illustration of transdermal drug delivery system is shown in Figure 1.
Figure 1. Schematic illustration of transdermal drug delivery (TDD) system.
Int. J. Mol. Sci. 2011, 12
In recent years, a few TDD products have been reported and approved by the US FDA. IONSYS
(Fentanyl ionophoretic), a product by Alza Corporation was approved in 2006 for patient controlled
pain management. Emsam, a product by Bristol-Myers Squibb (Princeton, NJ, USA) was approved in
2006 for major depressive disorder. Fentanyl generic by Watson Pharmaceuticals was approved in
2007 as an analgesic. Neupro, by Schwarz Pharma (Mequon, WI, USA) was approved in 2007 for
Parkinson’s disease. Exelon, by Novartis (East Hannover, NJ, USA) was approved in 2007 for
dementia [5]. Similarly various researchers have presented microfluidic devices for different medical
applications. Particularly micropumps and microneedles have been extensively studied in this decade
for biomedicine. But there is still a need to present the latest updates on the development of micropumps
and microneedles for biomedicine because these devices are still at the research level and have limited
availability for commercial use. Some earlier reviews on various applications of MEMS in the
biomedical field have been reported, such as the therapeutic microsystem, surgical microsystem and drug
therapy. These reviews provide basic information on various devices such as microneedles, micropumps,
micro-reservoirs, etc. [6–9]. Various researchers have reported reviews on design and development of
micropumps only [10–14]. Laser and Santiago [10] presented a comprehensive review on micropumps.
But the review did not cover some actuation methods, e.g., ion conductive polymer film (ICPF),
development of evaporation micropump and advance applications of micropumps in biomedicine.
Woias [11] presented a concise overview of different types of micropumps and their applications.
However, the electrowetting micropump, evaporation micropump and ICPF have not been described in
the review. Tsai and Sue [12] reported introductory overview on the importance of micropumps for
medical applications, but significant details about the applications of various kinds of micropumps for
drug delivery have not been presented. Nisar et al. [13] presented a comprehensive and good review on
various types of micropumps and their applications in biomedical applications. Some key features of
micropumps like actuation techniques, performance parameters, working principles, structure,
fabrication and applications have been reported, but this review has not covered latest developments in
micropumps for biomedical applications. The review does not provide up-to-date information about
bio-MEMS devices as there is an exponential increase in design and development in the bio-medicine
field. Amirouche et al. [14] presented a review on current developments in micropumps. The focus of
this review was on mechanical micropumps and their applications in the biomedical field. However
this review has not covered non-mechanical type of micropumps. Grayson et al. [15] reported a brief
review on various integrated MEMS devices such as biosensors, stents, immunoisolation devices,
reservoirs, microneedles, etc. This review has not described all parameters of MEMS devices like
design, development, actuation methods, fabrication techniques, etc. Karman et al. [16] reported a very
basic and introductory review on drug delivery devices like micropumps, microneedles, microvalves,
microactuators, microreservoirs, etc. This review has not covered important parameters such as
actuation techniques, working principles, performance constraints, design, fabrication and applications
of MEMS devices. Bao-jian et al. [17] presented information on the development and applications of
MEMS based microneedles. This review has not covered some important aspects of design and
development, forces experienced by microneedles, testing, structural/fluidic analyses, etc.
Khanna et al. [18] reported a review on the particular design requirements of microneedles for diabetic
therapy. This review has not covered the key parameters like development, fabrication, failure
analysis, etc. Sachdeva and Banga [19] reported good comprehensive review on microneedles design,
Int. J. Mol. Sci. 2011, 12
development, safety and regulatory issue, therapeutic applications and limitations of microneedles for
commercialization. However, this review has not described the fabrication techniques of microneedles,
failure of microneedles due to various applied forces, structural and fluidic analysis and integration
issues of microneedles with micropumps. All reviews that have been discussed above present the
information about micropumps or microneedles only. Here the authors have presented a review on
micropumps and microneedles that covers most recent advancement of MEMS technology in
biomedicine. This is the first comprehensive and updated review that covers latest information of
microfluidic devices regarding the design, development, actuation methods, performance parameters,
working principles, structure, fabrication techniques, material used for fabrication, safety issue,
challenges, limitations of commercialization and applications. This comprehensive review will be
helpful for researchers who would like to work in the fast growing field of bio-MEMS and bio-NEMS.
2. Micropumps
Pioneering work on micropumps started in the 1970s and developments based on microfabrication
technology was initiated in the 1980s. The MEMS based micropump was developed in 1990s. The
micropump is the main component of drug delivery system that provides the actuation mechanism to
deliver specific volumes of therapeutic agents/drugs from the reservoir. The requirements for drug
delivery include a minimum flow rate in order of 10 µL per minute or more, small size and high
reliability [13]. Normally a micropump consists of the following components: diaphragm membrane,
chamber, actuator, microchannels, microvalves, inlet, outlet, etc. Micropumps can be categorized into
two classes: One type has a mechanical moving part and is known as a mechanical micropump; the
other has no moving part and is known as a non-mechanical micropump.
2.1. Design Specifications and Parameters of Micropumps
Design of micropumps plays an important role for practical applications of devices. To develop a
suitable design of micropumps for real time applications, it is very important to understand terms like
actuator, valves, chamber or reservoir, nozzle diffuser mechanism and pumping parameters properly.
2.1.1. Actuator
The actuator is the necessary and driving part of a micropump that converts energy into motion. It is
used to provide force for fluid flow in micropumps. The actuator takes energy from electricity, heat,
liquid pressure, air pressure and converts it into some kind of motion. In most micropumps reported in
literature, the actuation disk is attached with membrane which is used to push the fluid. Some types of
time diaphragm are fabricated in such a way that it produces energy itself which pushes the fluid. In
peristaltic micropumps more than one actuator is fabricated sequentially.
2.1.2. Valves
In micropumps, valves are used to control the fluid flow by opening, closing and partially hindering
passageways. In microfluidic systems, active and passive valves have been reported. In passive valves
there is no actuation mechanism. The control of fluid flow is dependent on the pressure difference in
Int. J. Mol. Sci. 2011, 12
liquid chamber and the fluid flow is normally in one direction. In active valves, active elements are
present for opening and closing that are operated by an external actuation source. Mostly, separate
components have been reported for active micro-valves for regulating the fluid flow in microfluidic
systems. It is very easy to control the active valves but they are more complicated in integrated
microfluidic system.
2.1.3. Chamber or Reservoir
Chamber design is very critical in microfluidic systems and it can significantly influence the
volume stroke, pressure characteristics and nozzle-diffuser loss coefficients. Most of the micropumps
reported in literature have a single chamber configuration. But in order to improve the performance,
two or three chamber micropumps have also been reported. Micropumps in which pumping chambers
are arranged sequentially or fabricated in such a way that the multiple chambers are in series or in
parallel arrangements, are known as peristaltic micropumps.
2.1.4. Nozzle/Diffuser Element
Nozzle/diffuser element is mostly used in valveless micropumps as a flow rectifier. A schematic
illustration of the nozzle/diffuser action in micropumps is shown in Figure 2. Nozzle/diffuser element
works in such a way that during supply mode more fluid enters in the chamber through an inlet than
fluid that exiting the outlet. While in pump mode the reverse action occurs. Stemme and Stemme [20]
were the first to report valveless miniature micropumps in which they used a nozzle/diffuser element
as flow rectifying element.
Figure 2. Schematic of nozzle/diffuser element.
2.1.5. Pumping Parameters
Various design parameters are important to optimize the performance of micropumps such as
maximum flow rate (Qmax), pump power (Ppump), maximum back pressure (hmax) and pump efficiency
(η). Qmax is highest at zero hmax and Qmax is zero when highest value of hmax. For incompressible flow,
the pump head (h) can be calculated from the steady flow energy equation [21].
Int. J. Mol. Sci. 2011, 12
is pressure,
The pump efficiency
is pressure head,
is velocity head and
is elevation.
in the form of power can be expressed as:
Ideally, losses are zero and both quantities
are identical. Efficiency is governed
by frictional losses, fluid leakage losses and losses due to imperfect pump construction. The total
efficiency can be expressed as [21].
is mechanical efficiency,
is volumetric efficiency and
is hydraulic efficiency.
2.2. Mechanical Micropumps
The mechanical micropumps have moving parts so require a physical actuator for the pumping
process. The most common mechanical micropumps are displacement type micropumps that involve a
pumping chamber which is closed with a flexible diaphragm. The fluid flow is achieved by the
oscillation of a diaphragm. Due to these oscillations, the pressure ( P) is created. This pressure is a
function of stroke volume
inside the chamber produced by the actuator. The actuator has to run
itself with the dead volume
in chamber. Compression ratio is the important parameter for
mechanical diaphragm type micropumps. The compression ratio is defined by the equation (4):
The performance of mechanical micropump is normally limited by its mechanical components. The
piezoelectric, electrostatic, thermopneumatic, electromagnetic, bimetallic, ion conductive polymer
films (ICPF), phase change and shape memory alloy (SMA) are examples of mechanical micropumps.
A detailed description of mechanical micropumps is given below.
2.2.1. Piezoelectric Micropumps
The conversion of mechanical energy to electronic signal (voltage) and vice versa is known as the
piezoelectric effect. The materials which exhibit piezoelectric effect normally have no center of
symmetry in their structure. A stress applied to such materials will alter the separation between the
positive and negative charges that leads to the net polarization at the surface. An electrical field with
voltage potential is created in those materials due to the polarization. This property can be used to form
the actuator, micropump, inkjet printer head, etc. The effectiveness of energy and vice versa can be
expressed by factor :
Int. J. Mol. Sci. 2011, 12
Piezoelectric actuator shows large actuation and fast response time, but the fabrication of such
materials is complicated on a single chip. Piezoelectric micropumps exhibit small stroke volume at
high voltages. A schematic of a piezoelectric micropump is shown in Figure 3.
Figure 3. Piezoelectric micropump.
The first piezoelectric micropump was fabricated using micromachining technology by
Van Lintel et al. [22]. The micropump consisted of a pumping chamber, passive silicon (Si) check
valve, and a thin glass membrane actuated by piezo disk. The maximum flow rate of 8 μL/min and
back pressure of 9.8 kPa were observed at applied 125 V with 1 Hz frequency. Esashi et al. [23]
reported a three layers piezoelectric pump with flow rate of 15 μL/min and back pressure of 6.4 kPa at
applied 90 V with 30 Hz frequency. Olsson et al. [24] reported a two chamber piezoelectric
micropump to improve the performance. Koch et al. [25] presented piezoelectric micropump based on
screen printing of PZT (Lead Zirconate Titanate) on Si membrane. The flow rate of 120 μL/min and
back pressure 2 kPa were observed at applied 600 V with 200 Hz frequency. Schabmueller et al. [26]
fabricated piezoelectric micropump with passive valves. The flow rate of 1500 μL/min and back
pressure of 1 kPa were achieved using ethanol. Feng and Kim [27] reported piezoelectric micropump
that consisted of one way parylene valves. The flow rate of 3.2 μL/min and back pressure of 0.2 kPa
were observed at applied 80 V with lower power consumption of 3mW. Geipel et al. [28] reported a
novel design of micropump with back flow pressure independent flow rate. The back pressure
independency was reported up to 20 kPa at low frequency. Trenkle et al. [29] reported a piezostack
actuated peristaltic micropump. The flow rate of 40 μL/min was obtained at the frequency of 28.6 Hz
using water. The flow rates were observed to be independent of backpressure up to 7 kPa, with a
maximum backpressure of 45 kPa at 140 V. Johari et al. [30] reported the fabrication of a piezoelectric
micropump for drug delivery system using two optical masks. Fluidic characteristics analysis was
performed using CoventorWare simulator. Wang et al. [31] studied the effect of longitudinal flow
asymmetry on pumping capability by using a simple pumping system comprised of a piezoelectric
buzzer imbedded in a channel. Ali et al. [32] studied the dynamic piezoelectric micropump process.
The quantitative measurement of the pressure generated, applied electrical field, frequency and length
of the actuator, were observed. Liu et al. [33] proposed a disposable high performance piezoelectric
micropump with four chambers in serial connection for closed loop insulin therapy system. Outflow
resolution of 6.23 × 10−5 mL/pulse was observed. The maximum backpressure of 22 kPa was reported
at applied voltage of 36 Vpp and 200 Hz frequency.
Int. J. Mol. Sci. 2011, 12
2.2.2. Electrostatic Micropumps
Electrostatic micropumps involve electrostatic forces for actuation mechanism. Electrostatic force
is defined as ―the electrical force of attraction and repulsion induced by an electric field
‖. The like
charges repel each other and unlike charges attract each others. The electrostatic force applied on the
electrostatic plates can be expressed by the equation (6):
Where, is electrostatic attraction force,
is energy stored, is dielectric constant, is area of
electrodes, is electrode spacing and is applied voltage.
Electrostatic actuation is widely used in microfluidic devices. The fabrication of such mechanisms
on electronic chip is very easy, but electrostatic actuator has only a small stroke, typically 10 μm. The
main advantages of electrostatic micropump are low power consumption and fast time response. The
schematic of an electrostatic micropump is shown by Figure 4.
Figure 4. Electrostatic micropump.
The first electrostatic micropump was fabricated by Judy et al. [34] using surface micromachining
technology. It consisted of active check valve, chamber and active outlet valve. Pumping results were
not reported. The first experimental results of electrostatic micropump were reported by
Zengerle et al. [35]. The flow rate of 70 μL/min and back pressure of 2.5 kPa were observed at applied
170 V with frequency 25 Hz. Cabuz et al. [36] presented dual diaphragm electrostatic micropump
using injection molding technique. Micropump was capable of bidirectional operation but only used
for gases. The flow rate of 30 μL/min was observed at applied 160 V with frequency of 30 Hz and
power of 8 mW. Machauf et al. [37] presented membrane based electrostatically actuated micropump
across the working fluid. The concept was based on high and low electric permittivity of working fluid.
This pump was limited only for conducting fluid. The flow rate of 1 μL/min was achieved at 50 V.
Astle et al. [38] proposed a pumping mechanism using electrostatic actuation for gas chromatograph
applications. The flow rate of 3 mL/min and backpressure of 7 kPa were observed at frequency of
14 kHz. Lee et al. [39] fabricated and tested a peristaltic electrostatic gas micropump that employed
fluidic resonance for high flow rate and multi stage peristaltic configuration. The micropump presented
the pressure ranges from 7.3 to 3.3 kPa and flow rates from 0.29 to 0.07 sccm at the duration time
Int. J. Mol. Sci. 2011, 12
ranges from 0.05 and 0.35 cycles for opening of valves. Liu [40] reported the ―pull in phenomena‖ in
electrostatic micropump using reduced order model of membrane. Various parameters like radius,
thickness, initial gap, residual stress on pull in voltage and pull in position were investigated.
Lil et al. [41] presented the modeling of micropump membrane with electrostatic actuator. MATLAB
platform was used for modeling. The resonant frequency of 635 Hz for silicon electrostatic actuating
membrane was calculated. Using FEM, 680 Hz frequency was reported.
2.2.3. Thermopneuamtic Micropumps
In thermopneumatic micropumps, the actuator is based on thermal expansion. The chamber is full
of air and thermopneumatic micropump is expanded and compressed periodically by the heater and
cooler. The periodic change in volume of chamber provides the membrane with a regular momentum
that results in fluid out flow. The pressure increase is expressed by the equation (7).
is pressure change,
is temperature change,
is thermal expansion,
is a percentage of
volume change.
The thermopneumatic type of micropump generates relatively strong pressure and displacement of
membrane. However, the driving power has to be constantly maintained above a certain level. The
schematic diagram of thermopneumatic micropump is shown by the Figure 5.
Figure 5. Thermopneumatic micropump.
The first thermopneumatic micropump based on microfabrication was proposed by
Van De Pol et al. [42]. The flow rate of 34 μL/min was observed at applied voltage of 6 V with
temperature around 30 ˚C. Jeong and Yang [43] reported a thermopneumatic micropump with
corrugated diaphragm. The flow rate of 14 μL/min was observed at applied voltage of 8 V with
frequency of 4 Hz. A thermopneumatic micropump consisting of a thin film heater, flow strictor and
two reservoirs has been proposed by Cooney and Towe [44]. The maximum flow rate of 1.4 μL/min
for 4.5 h was observed with an average power of 200 mW. Kim et al. [45] proposed a
thermopneumatic micropump with a glass layer, indium tin oxide heater, polydimethylsiloxane
(PDMS) chamber, PDMS membrane and PDMS cavity. The flow rate of 0.078 μL/min was achieved
at applied voltage of 55 V with frequency of 6 Hz. Jeong and Konishi [46] fabricated a peristaltic
Int. J. Mol. Sci. 2011, 12
micropump consisting of three cascaded thermopneumatic actuators and microfluidic channel
connecting two fluidic inlet/outlet ports. The flow rate of 73.9 μL/min was achieved for the de-ionized
(DI) water at zero backpressure. Chia et al. [47] proposed a novel thermopneumatic peristaltic
micropump comprised of two separate zones for air heating and fluid squeezing. The temperature
elevation of 2.0 K was reported on the fluid pumping area. Tan et al. [48] fabricated a peristaltic
micropump by bonding a PDMS part with microchannels to the PDMS/PMMA
(polymethylmethacrylate) part where PDMS/adhesive membrane worked like a pneumatic actuator.
The maximum flow rate of 96l μL/min was achieved.
2.2.4. Electromagnetic Micropumps
Electromagnet is a kind of magnet that is based on the combination of electric and magnetic fields.
When the current passes through the coils the magnetic field is produced. The strength of
electromagnet can be easily varied by changing the electric current flowing through the coils. The
force experienced by the point charge due to the electromagnetic field is known as the Lorentz force.
The Lorentz force can be expressed by equation (8).
Where, is force and is magnetic field.
Electromagnetic actuation is large and covers a longer distance as compared to electrostatic
actuation. It needs low voltage but an external source is required for actuation such as a permanent
magnet. On small scale, this type of actuation has no benefit because it is reduced by the cube of
scaling factor. The driving coils or permanent magnets bond directly with the membrane and provide a
magnetic field. However, at the same time, the size is compromised. Usually electromagnetic
micropumps have high power consumption and heat dissipation. A schematic of an electromagnetic
micropump is shown in Figure 6.
Figure 6. Electromagnetic micropump.
The first electromagnetic micropump with 7 μm thick Ni80Fe20 film electroplated on 17 μm thick Si
membrane was proposed by Zheng and Ahn [49]. The maximum flow rate of 20 μL/min was observed
at applied voltage of 3 V with 5 Hz frequency and 300 mA induced current. A plastic micropump with
electromagnetic actuation has been reported by Bohm et al. [50] that consisted of two folded valves
with a thin membrane in center, inlet/outlet at bottom and pump membrane at top. The maximum flow
Int. J. Mol. Sci. 2011, 12
rates of 40,000 μL/min for air and 2100 μL/min for water were observed with power consumption of
0.5 W. A four layer electromagnetic micropump was designed and its static/dynamic properties were
investigated by Gong et al. [51]. The membrane deflection by different magnetic driving forces was
analyzed by ANSYS FEM. The maximum flow rate of 70 μL/min was observed at frequency of 125 Hz.
Yamahata et al. [52] reported a PMMA micropump with electromagnetic actuation. The maximum
flow rate of 400 μL/min and back pressure of 1.2 kPa were observed at resonant frequencies of 12 Hz
and 200 Hz. Su et al. [53] reported the analysis and fabrication of a valveless electromagnetic
micropump with two parallel flexible diaphragms. The maximum flow rate of 6 μL/s and the
displacement of 0.30 mm were observed at 100 Hz frequency with 0.3 A induced current.
Balaji et al. [54] reported the design, fabrication and testing of a flat pump with millimeter thickness.
The maximum flow rate of 15 μL/min was observed at applied voltage of 2.5 V with 68 Hz frequency
and 19 mA current. Yu-feng et al. [55] reported a parallel dynamic micropump with valve, diaphragm
and electromagnetic coil. The maximum flow rate of 6 μL/s and the diaphragm displacement of 30 μm
were observed at 100 Hz frequency with 0.3 A of current. Shen et al. [56] fabricated and characterized a
reciprocating PMMA ball valve micropump with electromagnetic actuation. The micropump showed a
backpressure of 35 kPa and flow rate of 6 mL/min at 2 W electromagnetic actuation power with 20 Hz
resonant frequency. Halhouli et al. [57] worked on the design of a novel electromagnetic pump that
based on the rotation of two hard magnets kept in channel, with opposing polarity. The maximum flow
rate of 13.7 mL/min at 200 rpm and a pressure of 785 Pa at 136 rpm were observed.
2.2.5. Bimetallic Micropumps
Bimetal refers to an object that is composed of two different metals jointed together. The thermal
expansion coefficients of these metals are different. The deflection of a diaphragm made of bimetallic
materials is induced against thermal alternation as long as the two chosen materials possess adequately
discriminative thermal expansion factors. A block diagram of bimetallic micropump is shown in Figure 7.
Figure 7. Bimetallic micropump.
Zhan et al. [58] reported Si based bimetallic micropump with 10 μm thick layer of aluminum (Al)
on Si substrate. The flow rate of 45 μL/min and back pressure of 12 kPa were observed at applied
voltage of 5.5 V with 0.5 Hz frequency. Zou et al. [59] designed a micropump that operated on both
bimetallic thermal actuation and thermal pneumatic actuation mechanisms. When the bimetallic
actuator made of Al/Si membrane was heated, the membrane deformed in downward direction. At the
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same time, the gas in the air chamber expended due to the heat to support bimetallic actuation. The
flow rate of 336 μL/min was achieved when the open pressure was 0.5 kPa. A novel micropump
operated on bimetallic and electrostatic actuation mechanisms was reported by Pang et al. [60].
Experimental results showed that the on/off flow ratio of the micropump was 180. Yang et al. [61]
presented a bimetallic thermally actuated membrane micropump that consisted of two chips, pump
chamber, two bimetallic actuators and two check valves. The maximum flow rate of 43 μL/min was
achieved at applied voltage of 16 V and 0.9 Hz frequency. The forces generated through bimetallic
actuation are large and the implementation is simple. Usually the thermal expansion coefficients of
materials that are involved in bimetallic micropumps are small. That is why the diminutive deflections
are achieved in bimetallic actuation mechanism. The bimetallic micropumps require low voltage
values as compared to other micropump types. But the drawback of bimetallic micropumps is that they
are not suitable to work at high frequencies.
2.2.6. Ion Conductive Polymer Film (ICPF) Micropumps
ICPF actuator shows high speed response. However, the positioning control is difficult. The core
layer of ICPF is made of a sort of perfluorosulfonic acid polymer. Physically it looks like a ―sandwich‖
diaphragm between two thin films that are placed on both sides of the polymer. These two films have
high electrical conductivity. One end of the diaphragm is fixed and the ICPF diaphragm can be
controlled by bending in the direction of either upside or downside as long as an appropriate pair of
voltages is applied at the electrodes. The ICPF actuator is commonly called an artificial muscle
because of the large bending displacement, low actuation voltage and biocompatibility. A schematic of
ICPF and the bending principle is shown in Figure 8.
Figure 8. ICPF micropump.
ICPF actuators have been developed for various applications. Guo et al. [62] reported a new model
of micro catheter with active guide wire that had two bending degrees of freedom using ICPF actuator.
Tadokoro et al. [63] developed multi-degree-of-freedom (DOF) micro motion devices using ICPF soft
Int. J. Mol. Sci. 2011, 12
gel actuator. Guo and Asaka [64] proposed an underwater fish like microrobot using ICPF actuator as
the servo actuator swimming motion with three degrees of freedom. Nguyen et al. [65] reported the
design and fabrication of a flap valve ionic polymer/metal composite micropump with the diaphragm
supported by a flexible material. A maximum flow rate of 760 μL/min and backpressure of 1.5 kPa
were observed at the applied voltage of 3 V with 3 Hz frequency. Chen et al. [66] proposed the design
of an integrated sensory actuator. The polyvinylidene fluoride (PVDF) films were used for
simultaneous feedback of bending and force outputs of the actuator. Fang and Tan [67] proposed a
control oriented model to envisage the deformation of diaphragm and the flow rate. Experimental
results of the polypyrrole (PPy) actuated micropump showed that the maximum flow rate of
1260 μL/min was observed at the voltage of 4 V.
2.2.7. Phase Change Micropumps
The basic principle used in phase change type actuators and micropumps is the vaporization and
condensation phenomenon. In vaporization, the phase transition occurs from liquid phase to vapor
phase. While in condensation, the change of the physical state occurs from gaseous phase to liquid
phase. The phase change type micropump consists of a heater, diaphragm and working fluid chamber.
A schematic of phase change micropump is shown in Figure 9.
Figure 9. Phase change micropump.
Sim et al. [68] proposed a phase change type micropump consisting of a pair of Al flap valves and a
phase-change type actuator. The actuator comprised of a heater, working fluid chamber and silicone
rubber diaphragm. The diaphragm was actuated by the vaporization and the condensation of the
working fluid in the chamber of the pump. The maximum flow rate of 6.1 μL/min was achieved at
applied voltage of 10 V with 0.5 Hz frequency and 60% duty ratio for zero pressure difference.
Boden et al. [69] reported a high pressure micropump with polymeric paraffin actuation. The flow rate
of 74 μL/min was achieved at a low voltage waveform with water as a pumping fluid. When the
pressures up to 1MPa were applied on the valves, the micropump showed no leakage. Sim et al. [70]
reported the fabrication and testing of a micropump comprised of a pair of Al flap valves and a phase
change type actuator. The actuator was composed of a heater, diaphragm and fluid chamber. The
maximum flow rate of 97 μL/min was observed at applied voltage of 8 V with 70% duty ratio and 2 Hz
frequency for zero pressure difference.
Int. J. Mol. Sci. 2011, 12
2.2.8. Shape Memory Alloy (SMA) Micropumps
SMA are the metals which exhibit two very unique properties such as pseudo elasticity and the shape
memory (SM) effect. They have the capability of changing their shapes upon application of an external
stimulus. The SM effect involves a phase transformation between two solid phases. At high temperature
the phase is called austenite and at low temperature the phase is called martensite. SMA starts in
martensite phase and transforms into austenite phase after being heated. This property of materials is
useful to make SMA micropumps. A schematic of an SMA micropump is shown in Figure 10.
Figure 10. SMA micropump.
The first thin film SMA micropump with two different actuation configurations was reported by
Benard et al. [71]. The pump was driven by an electrical drive signal provided directly through the
Titanium/Nickel (Ti/Ni) thin films, resulting in Joule heating induced phase transformation that
initiated the SM effect. The maximum flow rate of 50 μL/min was observed at 0.9 Hz frequency.
Makino et al. [72] reported the development of SMA actuated micropump to use in micro analysis and
micro dosage systems. The maximum flow rate of 0.4 μL/cycle was observed at a bias pressure of 100 kPa.
Xu et al. [73] developed a micro SMA pump composed of a NiTi/Si composite membrane, pump
chamber and two inlet/outlet check valves. The flow rate of 340 μL/min and back pressure of 100 kPa
were achieved. Shuxiang and Fukuda [74] developed SMA micropump composed of SMA coil
actuator, two diffusers, pump chamber and a casing. The maximum flow rate of 500–700 μL/min was
achieved by changing the frequency. Zhang and Qiu [75] reported a Ti/Ni/Copper (Cu) shape memory
thin film micropump comprised of a TiNiCu/Si driving membrane, pump chamber and two inlet and
outlet check valves. The hysteresis width ∆T of 9 °C was observed. Setiawan [76] reported the
performance assessment of SMA spring as actuator for gripping manipulation. The SMA actuator was
a TiNi tensile spring with diameter of 50 mm wire and 350 gram hanging mass. SMA have many
attractive properties like high force to volume ratio, ability to recover large transformation stress and
strain upon heating and cooling processes, high damping capacity, chemical resistance and
biocompatibility. Usually the deformation of SMA cannot be precisely controlled and investigated due
to temperature sensitivity. Additionally, the designs based on TiNi film devices with more practical,
effective and complex characteristics, are required through multiple DOF and compact structures.
Recently reported mechanical micropumps are listed in Table 1.
Int. J. Mol. Sci. 2011, 12
Table 1. Recently reported mechanical micropumps.
Liu et al.
2010 [33]
Materials used
for fabrication
Polycarbonate (PC),
(Titanium) Ti
15 × 8 mm
200 Hz
22 kPa
6.23 ×
Insulin therapy
Not reported
for air,
for water
Drug delivery
Zhu et al.
2009 [77]
225 Hz for
17 Hz for
Kang and
2011 [78]
Si/Epoxy H31/
14.5 × 9 ×
1.1 mm
20-100 Hz
0-10 psi
et al.
2010 [57]
PC, Plexiglass
16 × 18
785 Pa
Shen et al.
2009 [79]
PDMS, Glass
24 × 40 ×
0.4 mm
12 Hz
70 mbar
Lee et al.
2009 [39]
7.3–3.3 kPa
Not reported
and Sani
2005 [80]
Si, Glass
7 ×4 ×1
50 Hz
Not reported
Drug delivery
Int. J. Mol. Sci. 2011, 12
Table 1. Cont.
Materials used
for Fabrication
Chia et al.
2010 [47]
Tan et al.
2010 [48]
Zou et al.
1997 [59]
Back Pressure/
Flow Rate
1.2 Hz
490 Pa
20.01 μL/min
10 Hz
138 kPa
96 μL/min
0.5 kPa
5.6 μL/s
Not reported
0.5 Hz
1.3 kPa
1260 μL/min
2 Hz
0 mm H2O
97 μL/min
Not reported
8 ×8 ×
1.8 mm
DI water
80 Hz
Not reported
235 μL/min
Not reported
PDMS, Glass
16 × 18
× 5 mm
Not reported
Al, Si, Glass
13 × 7 ×
2 mm
Fang and
Tan 2010
Stainless steel,
25 × 25
× 10
Sim et al.
2008 [70]
Phase Change
Al, Silicon,
Silicone rubber,
and Qiu
2006 [75]
Ti, Nickel (Ni),
Copper (Cu)
Int. J. Mol. Sci. 2011, 12
2.3. Non-Mechanical Micropumps
The non-mechanical micropumps have no moving mechanical part so that generally they need a
type of mechanism that can convert non-mechanical energy into kinetic momentum. In general, nonmechanical pumps do not need physical actuation components so the geometry, design and fabrication
of these micropumps are relatively simple and easy. These micropumps have certain limitations, such
as the use of only low conductivity fluids and the actuation mechanisms interfere with the pumping
liquids. A detailed description of non-mechanical micropumps is given below.
2.3.1. Electroosmotic (EO) Micropumps
EO flow is the motion of the liquid that is induced by an applied potential across a capillary tube or
microchannels. The fluid with electric conductivity feature is driven by appropriately exerting an
external electrical field upon the channel walls that are naturally charged. A schematic diagram of an
electroosmotic micropump is shown in Figure 11.
Figure 11. EO micropump.
Zeng et al. [81] fabricated an EO micropump that used DI water as working fluid. The maximum
flow rate of 3.6 μL/min and pressure of 2026.5 kPa were obtained at applied voltage of 2 kV.
Takemori et al. [82] reported an EO micropump with high pressure. The flow rate of 0.47 μL/min and
pressure of 72 kPa were observed at applied 3 kV. Hu and Chao [83] investigated the EO flow in EO
micropump with an overlapped electrical double layer (EDL). The results showed that the flow was
relatively different from the channel with a dimension greater than the EDL, which demonstrated plug
like flow properties. Good et al. [84] performed the mathematical modeling and experimental testing
of water activated micropump that was actuated using the osmotic effect. The maximum flow rate of
17 μL/min/mg of dry polymer particles, with a 355–425 μm diameter, was achieved. Ryu et al. [85]
proposed a biodegradable osmotic micropump for long use and controlled discharge of basic fibroblast
growth factor (bFGF). The release of bFGF was regulated at a rate of 40 ng/day for duration of four
weeks. Yairi and Richter [86] developed an EO micropump based on voltage control. The flow rate of
0.054 mL/min and pressure of 5.5 kPa were achieved. Borowsky et al. [87] fabricated a high pressure
EO micropump and tested the performance of fluid dynamic. The maximum flow rate of 85 μL/min
and pressure of 25 atm were achieved. Wang et al. [88] reported the general characteristics, fabrication
technologies and applications of EO micropumps. The transport of various solutions compositions into
Int. J. Mol. Sci. 2011, 12
capillaries can cause problems in the flow constancy of an EO pumped system in some applications.
Sometimes flow rates are modified due to adsorption of compounds from the samples or sample matrix
on the surfaces of the pumping elements. This problem can be solved by separating the pump fluid
from the sample and reagent solutions in the analytical system.
2.3.2. Electrowetting (EW) Micropumps
EW is a microfluidic phenomenon that is currently used as a driving mechanism for fluidic devices.
EW involves modifying the natural surface tension or capillary forces intrinsic to an oil and water
interface at small length scales. At less than 1 mm distance, the electrical and surface tension forces are
much stronger than gravity. The digital EW is applied to control the surface tension between solid
phase electrode and liquid phase droplet. A schematic of an EW micropump is shown in Figure 12.
Figure 12. EW micropump.
Yun et al. [89] reported a continuous EW micropump. For the actuation energy of micropump, the
surface tension induced motion of mercury drop in a microchannel filled with electrolyte was used.
The micropump consisted of a stack of three wafers bonded together. The flow rate of 70 μL/min and
pressure of 0.8 kPa were achieved at 2.3 V with frequency of 25 Hz and power consumption of
170 μW. Hoshino et al. [90] reported the pico liter liquid actuation in a microinjector by using a pulled
glass tube as the device structure. The tube caused pumping and ejection by EW on dielectrics.
500 picoliter water was pumped up at the maximum applied voltage of 1400 V. In pumping pressure,
an increase value of 0.6 Pa was calculated. Colgate and Matosumoto [91] reported a detailed model of
a test device showing liquid flow in a small channel for the study of EW. EW gives direct fluid
pumping without any moving mechanical parts that can be valuable in many application areas of
microelectronic devices. The initial results showed that EW might be used to get pressures on the order
of 0.01 MPa in a 10 μm radius channel. Chang et al. [92] reported the driving characteristics of the
EW-on-dielectric device with aluminum oxide (Al2O3) deposited by using the method of atomic layer
deposition. When the voltage was applied between control electrode and reference electrode then the
flow of 2 μL for water droplet in an air environment was achieved.
2.3.3. Electrochemical Micropumps
The most common feature of electrochemical micropumps is the generation of bubbles by
electrolysis in which the decomposition of water occurs into its constituents, such as hydrogen gas
Int. J. Mol. Sci. 2011, 12
(H2) and oxygen gas (O2), when the current is passed through water. During this mechanism, the key
component is a bubble reservoir filled with a redox electrolyte solution. The reaction of electrolysis
can be described by the equations (9) and (10).
At Anode
At Cathode
A schematic of electrochemical micropump is shown in Figure 13.
Figure 13. Electrochemical micropump.
Suzuki and Yoneyama [93,94] fabricated an electrochemical syringe pump by using micromachining
for low operating voltage and power consumption. A microfluidic system was developed by integrating
an on-chip micropump and check valves that worked through a H2 bubble generated electrochemically.
Thin film electrodes were used with a platinum black working electrode. PDMS substrate was used to
make flow channels and containers for electrolyte solutions. Two dye solutions were transported and
merged in a flow channel and sheath flows were observed. Yoshimi et al. [95] developed an artificial
synapse using the electrochemical micropump. The micropump consisted of a glass nozzle and two
blackened platinum electrodes filled with a neurotransmitter solution for the electrolysis process. To
drive the solution towards the neuron, a potential difference of 3.0 V was applied to the electrodes.
Kim et al. [96] reported a PPy-membrane microfluidic pump. The pumping action was stimulated by an
electrochemical actuated PPy-PDMS membrane. The check valves were used to control the direction of
flow. The maximum flow rate of 52 μL/min was obtained at ±1.5 V with input power of 55 mW.
2.3.4. Evaporation Micropumps
In evaporation micropumps, a controlled evaporation of liquid is used. Evaporation is a process in
which liquid is converted from its liquid form to vapor form. The reverse of this process is known as
Int. J. Mol. Sci. 2011, 12
condensation. The pumping principle of the evaporation type micropump is the same as the xylem
transport system in plants. A schematic of an evaporation micropump is shown in Figure 14.
Figure 14. Evaporation micropump.
Effenhauser et al. [97] reported the evaporation based disposable micropump for continuous
monitoring systems. The controlled evaporation of liquid was done through a membrane into gas space
that contained a sorption agent. In the gas chamber, the vapor pressure was kept lower than saturation.
During this process, the fluid evaporation from membrane was substituted by capillary forces that
resulted in a flow from the reservoir. The average flow rate of 0.35 μL/min was achieved.
Namasivayam et al. [98] reported the micropump based on the generally observed phenomenon of
transpiration in plant leaves for continuous very low flow rates. As the vapor diffused out due to
heating, a new transport of liquid was supplied into the channel from a reservoir for steady state
operation. Guan et al. [99] reported a micropump based on capillary-evaporation effects for a
microfluidic flow injection chemiluminescence system. The average flow rate of 3.02 μL/min was
achieved with an ambient temperature of 20–21 °C and relative humidity of 30–32% for fluctuation
within 2 h. Heuck et al. [100] reported the evaporation-based micropump integrated into a scanning
force microscope probe for the flow of liquid through its hollow cantilever and tip areas. A flow rate of
11 pL/s was obtained at room temperature.
2.3.5. Bubble Micropumps
The bubbles micropump is based on periodic expansion and collapse in the volume controlled by
voltage input. The volume change in chamber is incorporated with the diffuser/nozzle mechanism that
is used to determine the direction of fluidic flow. The bubbles are generated by heating process. A
schematic of the bubble micropump is shown in Figure 15.
Int. J. Mol. Sci. 2011, 12
Figure 15. Bubble micropump.
Tsai and Lin [101,102] reported a valveless thermal-bubble micropump. Later they developed a
microfluidic mixer system with a gas bubble filter using the bubble micropump. The maximum flow
rate of 5 μL/min was achieved at 250 Hz with applied periodic voltage, 10% duty cycle and power
consumption of 1 W. Lew et al. [103] developed a collapsing bubble micropump. The bubbles with a
radius of about 3–5 mm were investigated through the experimental set up that employed a low voltage
electrical spark of 55 V created with a capacitor for bubble generation. It was reported that the
proposed theory could also work with even smaller bubbles. Jung and Kwak [104] reported the
fabrication and testing of bubble type micropumps using an embedded microheater. The micropump
comprised of a pair of nozzle/diffuser, flow controller, microchannels and a pumping chamber. The
maximum flow rates of 6 μL/min at duty ratio of 60% for circular chamber and 8 μL/min at duty ratio
of 40% for the square chamber were achieved. Cheng and Liu [105] reported an electrolysis-bubble
micropump based on the roughness-gradient design in the microchannel. The electrolysis actuation and
the surface tension effect were used for the micropump. The maximum flow rate of 114 μL/min was
obtained at applied voltage of 15 V with a frequency of 4.5 Hz. Chan et al. [106] developed a bubble
type micropump with high frequency flow reversal using embedded electrodes in a closed microfluidic
microchannel. The micropump consisted of a microfluidic chamber and microelectrodes on a glass
substrate that was assembled by PDMS-sheet. The maximum flow rate of 37.8 μL/min was achieved at
voltage of 5 V.
2.3.6. Magnetohydrodynamic (MHD) Micropumps
MHD is a field in which the dynamics of electrically conducting fluids is studied. The Lorentz force
is the driving source perpendicular to the electric and magnetic fields for MHD type of micropumps.
The working fluid is selected to achieve conductivity of 1 s/m or higher, in addition to externally
providing electric and magnetic fields. The Lorentz force can be expressed by the equation (11).
Where, is force, is electric field, is instantaneous velocity of particles,
is electric charge of the particle.
A schematic of the MHD micropump is shown in Figure 16.
is magnetic field and
Int. J. Mol. Sci. 2011, 12
Figure 16. MHD micropump.
Jang and Lee [107] reported the MHD micropump. The pressure head difference of 18 mm at
38 mA and a flow rate of 63 μL/min at 1.8 mA were achieved with an inside diameter of 2 mm for
inlet/outlet tube and a magnetic flux density of 0.44 T. Zhong et al. [108] reported the fabrication of
MHD micropump using ceramic tapes. Experiments were performed using mercury slugs, saline
solutions and DI water. Eijkel et al. [109] developed a circular ac MHD micropump for
chromatographic applications. The device comprised of a glass-gold-laminate-glass sandwich structure
with the channel defined in the electroformed gold layer. Reversible flow rate of 40 μm/s was
achieved. Patel and Kassegne [110] reported a MHD micropump with EO-thermal effects using
3D-MHD equations. The use of a developed numerical framework, flow channel geometries, Joule
heating, effects of non-uniform magnetic/electric fields and EO in MHD micropumps were
investigated. Duwairi and Abdullah [111] developed a model to envisage the fluid flow in the MHD
micropump. By applying the finite difference method and the SIMPLE algorithm, the transient,
incompressible, laminar and flow equations were numerically solved. Kang and Choi [112] reported
the design and fabrication of MHD micropump with a mixing function in which the fluids were mixed
and pumped at the same time by coupling between Lorentz force and the moving force of an electric
charge in the electric field.
2.3.7. Flexural Planer Wave (FPW) Micropumps
The FPW micropumps are driven ultrasonically. The fluidic motion induced by traveling FPW can
be used for the transport of liquids. The liquid motion is in the direction of wave propagation and the
speed is proportional to the square of acoustic amplitude. Low operating voltage is required for
acoustic streaming. A schematic of the FPW micropump is shown in Figure 17.
Int. J. Mol. Sci. 2011, 12
Figure 17. FPW micropump.
Moroney et al. [113] reported the process of water pumping induced by 4.7 MHz ultrasonic Lamb
waves. The waves were moving in a composite membrane of silicon nitride and piezoelectric zinc
oxide with a thickness of 4 μm. The observed speed was 100 μm/s at the applied rf voltage of 8 V with
6.5 nm wave amplitude. Nguyen and White [114] reported the design and numerical model of an
ultrasonic FPW micropump and microfluidic system. The effects of channel height, wave amplitude,
and backpressure on the velocity and flow rate were studied. The influence of thermal transport of the
acoustic streaming was also investigated. Results showed that the micropumps with channel heights of
a few micrometers exhibited high-quality performance because the flow rate and hydraulic impedance
against backpressure were high. Nguyen et al. [115] reported a FPW micropump integrated with flow
sensor for in situ measurement. The FPW micropump and the flow sensor made a complex
microfluidic system capable of controlling the fluid flow in the device. Meng et al. [116] reported the
ultrasonic FPW micropump. The waves travelled along a thin membrane to stimulate an acoustic field
in the fluid that was in contact with the membrane. The micropump with a combination of radial
transducers and unidirectional fluid flow resulted in a flow speed of 1.15 mm/s. Jang et al. [117]
investigated the actuating frequency control of acoustic-streaming flow patterns in a diaphragm driven
microfluidic chamber. Microfluidic circulatory flow was achieved using the resonant vibration of
diaphragms. Experiments were performed to study in-plane velocity profiles near the interface of
circulations where the acoustic intensity was measured to be large. The proposed flow process was
reported to be useful for pumping, active mixing and particle focusing applications. Singh and
Bhethanabotla [118] studied the enhancement in the efficiency of acoustic-streaming. Microfluidic and
biosensing applications of surface-acoustic wave devices depend on the acoustic-streaming process
resulting from high intensity sound waves that interact with the fluid medium.
2.3.8. Electrohydrodynamic (EHD) Micropumps
In an EHD micropump, the force is generated by the interaction of electric field and mobile charges
in the fluid. These pumps have emitter and collector electrodes that are regularly spaced along a
microchannel and require no moving parts such as impellers, bellows or valves. The electrical charges
generated from the electrodes mobilize according to the direction of the electric field that is built up by
Int. J. Mol. Sci. 2011, 12
the electrodes and tract in the surrounding liquid molecules to move together by the ion dragging force.
The force acting on the fluid is given by the equation (12).
Where, is force on fluid, is current, is distance between electrodes, is ion mobility coefficient of
the dielectric fluid. A schematic of an EHD micropump is shown in Figure 18.
Figure 18. EHD micropump.
Ritcher and Sandmaier [119] fabricated the first dc charged injection EHD micropump comprised of
two electrically isolated grids. The flow rate of 15,000 μL/min and the pressure head of 1.72 kPa were
achieved at applied voltage of 800 V. Fuhr et al. [120] developed the first EHD micropump based on
travelling wave-induced electroconvection. The flow rates of 0.05–5 μL/min were achieved.
Darabi et al. [121,122] reported the EHD polarization micropump for electronic cooling and EHD ion
drag pump. The model devices exhibited a maximum cooling capacity of 65 W/cm2 with pumping
head of 250 Pa. Yang et al. [123] reported an ejection type EHD micropump using indium-tin-oxide
(ITO) planar electrodes to deal with the aging problem. The planar electrodes could drive the ethyl
alcohol with a flow rate of 356 μL/min at applied dc voltage of 61 V. Lin and Jang [124] reported the
numerical microcooling analysis for EHD micropump. The micropump offered the pumping power
using the dipole moment force generated from polarizing fluid molecules. The pressure head of
13 kPa and wall heat flux of 10 W/cm2 were observed at applied voltage of 500 V with pitch of
500 μm for parallel electrodes. Darabi and Rhodes [125] reported the computational fluid model of ion
drag EHD micropump. The micropump consisted of an array of interdigitated electrodes with the top
and bottom parts of the channel. Singhal and Garimella [126] reported induction based EHD
micropump for high heat flux cooling process. The numerical model was developed by solving the three
dimensional transient fluid flow and charge transport problem due to simultaneous actuation of EHD and
the vibrating diaphragm. Recently reported non-mechanical micropumps are listed in Table 2.
Int. J. Mol. Sci. 2011, 12
Table 2. Recently reported non-mechanical micropumps.
Chan et al.
Jung and
2007 [104]
et al.
2009 [127]
Material used
300 Hz
Not reported
Bubble type
PDMS, Glass,
DI water,
Bubble type
Si, Pyrex glass
DI water
0.5–2.0 Hz
Not reported
Carbon, Glass
23 Pa
373 kHz
Not reported
cooling system
Singhal and
2007 [126]
1500 ×
200 × 50
Lister et al.
2010 [128]
Glass, Platinum
buffer, DI water
1.6 kPa
Drug delivery
Xu et al.
2010 [129]
Glass, PDMS
Not reported
cell culture
Kang and
Choi 2010
Au (gold),
PBS solution
Not reported
Int. J. Mol. Sci. 2011, 12
Table 2. Cont.
Material used
for Fabrication
Lim and Choi
2009 [130]
Yun et al.
2002 [89]
Back Pressure/
Applied Pressure
Drug delivery
25 Hz
800 Pa
11 mbar
DI water
Not reported
11 pL/s
23.5 kPa
2–3 MHz
Not reported
Fluid delivery
Si, Pyrex glass,
40 × 25 ×
1 mm
Glass, Si,
Kim et al.
2008 [96]
Ppy, PDMS,
Heuck et al.
2008 [100]
Guan et al.
2006 [99]
1997 [131]
5.6 × 16 ×
26 mm
Pdms, PMMA,
Stainless steel
25 × 15 ×
3 mm
Si, Platinum,
Int. J. Mol. Sci. 2011, 12
3. Microneedles
Microneedles are very useful delivery devices. These devices provide an interface between the drug
reservoir and the patient’s body for releasing or extracting the fluid. The length of microneedles should
be long enough that it penetrates the epidermis and short enough not to reach the dermis, in order to
avoid pain. The concept of microneedles was proposed in the 1970s but it was not realized
experimentally until the 1990s when the industry of microelectronics provided the microfabrication
tools essential to make such small structures. The first microneedle arrays reported in the literature
were developed by etching the Si wafer for intracellular delivery [132]. These needles were inserted
into cells and nematodes to increase molecular uptake and gene transfection. After that a number of
attempts have been made by various researchers to develop the fabrication processes and different
designs of microneedles. MEMS technology is the most promising to fabricate the optimal design of
microneedles for particular applications. The typical diameter and length of MEMS-based
microneedles are in the range of micrometers. These microneedles are different from standard
hypodermic needles used in biomedicine. Generally, the length of the MEMS-based microneedles is
less than 1 mm. Thus microneedles are significantly smaller in length than ordinary needles [4,133].
Microneedles or microneedle arrays can be used as a stand-alone microfluidic device as well as part of
biological detection, fluid extraction or delivery system. Microneedles can be integrated with
micropumps, biosensors, microelectronic devices and microfluidic chips.
3.1. Categories of Microneedles
Different designs of microneedles have been reported in literature for various applications.
Microneedles can be classified in various ways such as according to structure, overall shape, tip shape,
length, array density, material used for fabrication and applications [3,4]. Details of microneedle
categories are shown in Table 3.
Table 3. Categories of microneedles.
Tip Shape
Material Used
Snake fang
Single crystal silicon
Silicon dioxide
Silicon nitride
Ti- alloy
Stainless steel
Drug delivery
Gene delivery
Blood extraction
Fluid sampling
Cancer therapy
Skin treatment
Cell surgery
Allergies diagnosis
Animal identification
Ink-jet printing
Sensing electrodes
Int. J. Mol. Sci. 2011, 12
3.1.1. Structure of Microneedles
Structure is the most important consideration for microneedles design and fabrication. Based on the
fabrication process, the microneedles are classified in two types.
In-plane microneedles
Out-of-plane microneedles
In in-plane microneedles, the microneedle shafts or lumens are parallel to the substrate surface. The
major advantage of in-plane microneedles is that the length of the microneedles can be easily and
accurately controlled during fabrication process. The limitation of in-plane microneedles is that it is
very difficult to fabricate microneedle arrays with 2D geometry. In out-of-plane microneedles, the
lengths of the microneedles protrude out of the substrate surface and it is easier to fabricate
out-of-plane microneedles in 1D or 2D arrays. However, fabrication of out-of-plane microneedles with
length and high aspect ratio structure is challenging [4,134]. A schematic illustration of in-plane and
out-of-plane microneedles is shown in Figure 19.
Figure 19. (a) In-plane microneedles; (b) Out-of-plane microneedles.
In-plane microneedles were developed in the 1980s [134] and not intended for drug delivery or
fluid transport. An implantable ten-channel microelectrode recording array with an on-chip signal
processing probe was fabricated for long term recording of neural bio-potentials. The length of probe
and thickness were 4.7 mm and 15 μm respectively. A 1D array of micro neural probes [135] and more
sophisticated 2D array have been developed [136]. After that various attempts have been made to
develop in-plane microneedles for different applications. The major drawback associated with in-plane
microneedles is the limited density. To overcome this limitation, out-of-plane microneedles have been
developed. One of the earliest out-of-plane microneedle array consisted of 100 microneedles with a
length of 1.5 mm was reported in 1991 [132].
Microneedles can also be categorized as solid or hollow according to the structure. Hollow needles
were invented in 1844 [137] and gained increasing importance in the biomedical field. There are no
Int. J. Mol. Sci. 2011, 12
other effective ways to transport the fluid into the human body [138]. Hollow needles have become
more important after the invention of microneedles. Hollow microneedles have an internal bore or
lumen which allows flow of fluid/drug through the microneedles. A combination of surface and bulk
micromachining techniques was used to fabricate hollow in-plane microneedles with 1-6 mm length
and fully enclosed channels of silicon nitride [139]. The channels were 9 μm in height. The solid
microneedles have solid lumens and exhibit more strength than hollow microneedles. Solid
microneedles can be further categorized into coated and dissolving microneedles. In coated
microneedles, the drug particles are coated on lumen surface and injected into patient body. The
microneedles are withdrawn from the body after dissolution of the coated drug. In dissolving
microneedles, the base is non-dissolving and withdrawn from the skin after dissolution of the
microneedles. Various types of solid coated and dissolving microneedles have been reported [19].
Coated Ti microneedles arrays with a length of 190 μm have been reported for the delivery of
parathyroid hormone (I—34) in human body for the treatment of osteoporosis by Zosano Pharma, Inc.
(formerly Macroflux®, ALZA Corp.) [140]. The successful delivery of drug depends on the methods
used for coating of microneedles [141,142]. References [143,144] fabricated the first out-of-plane
sharp solid microneedles for drug and gene delivery. A schematic of hollow and solid microneedles is
shown in Figure 20.
Figure 20. (a) Hollow microneedle; (b) Solid microneedle.
Hollow silicon dioxide (SiO2) microneedles have been fabricated using deep reactive ion etching
technique [145]. Reference [146] fabricated SiO2 microneedles which mimic a jagged mosquito’s
needle. In-plane hollow metallic hypodermic microneedles and microneedle array were reported using
electroplated palladium (Pd) alloys and Ni [147–149]. Using a combination of isotropic and an
isotropic etching process, sharp tip hollow out-of-plane single crystal Si microneedles were
fabricated [150]. One of the earliest solid microneedles design was in the form of pyramidal Si
microprobes [151]. Sharp Si solid microstructures with a height of 150 μm were fabricated with
anisotropic dry etching technique using SF6 and O2. Such type of solid microneedles was used to
increase permeability of human skin up to fourth order of magnitude in vitro. Solid microneedles for
TDD were reported for the first time in 1998 [144].
Int. J. Mol. Sci. 2011, 12
3.1.2. Shape of Microneedles
The shape of the microneedle is very critical and important during design and fabrication.
Microneedles can be classified on the basis of overall shape and tip shape. Different designs of
microneedles have been proposed and fabricated such as cylindrical, canonical, pyramid, candle, spike,
spear, square, pentagonal, hexagonal, octagonal and rocket shape [3,4]. Microneedles have also been
reported with various tip shapes like volcano, snake fang, cylindrical, canonical, micro-hypodermis
and tapered. Schematic illustrations of various designs of microneedles with respect to shape and tips
are shown in Figure 21.
Figure 21. Shapes of microneedles (a) Cylindrical; (b) Tapered tip; (c) Canonical;
(d) Square base; (e) Pentagonal-base canonical tip; (f) Side-open single lumen; (g) Double
lumen; (h) Side-open double lumen.
Rocket shape microneedles have been fabricated using two photon polymerization method [152].
Octagonal solid out-of-plane Si microneedle array has been fabricated for drug delivery [153]. Solid
Si-tip microneedles have been fabricated using wet etching technology [154]. Pyramidal out-of-plane
Si microneedle array has been fabricated by wet etching for transcutaneous drug delivery [155]. Side
opened sharp tip out-of-plane solid microneedle has been fabricated by hot embossing to improve skin
permeability for hydrophilic molecules [156]. Cylindrical hollow out-of-plane microneedles with
tapered tip using combination of ICP etching have been fabricated for TDD [4].
Int. J. Mol. Sci. 2011, 12
3.1.3. Materials Used for Microneedles
Microneedles can be classified on the basis of materials. Material selection is very important to design
and fabricate microneedles for any particular application. Many researchers used Si for microneedles
fabrication [4,157–162], which is a brittle material [163] and can be harmful to health. Different
researchers have understood this critical issue and used polymeric material instead. Most polymers have
a strong history of biocompatibility. They exhibit excellent mechanical and chemical properties [164]
that are suitable for microneedle fabrication. Fabrication of microneedles has been reported using various
polymers such as (Polyglycolic acid) PGA, (Poly-L-Lactide acid) PLLA, PC, PDMS, PMMA, etc.
Fabrication of polymeric microneedles has been reported by various researchers [165,166]. Some other
materials have also been reported such as glass, metal, alloy, etc. [4]. Glass hollow elliptical tip
microneedles have been fabricated using micropipette pulling technique for intrascleral delivery [167].
In-plane Ti microneedles have been fabricated using bulk micromachining for drug delivery [168].
Tungsten microneedles have been reported for nerve penetration [169].
3.1.4. Microneedles Applications
On the basis of applications, microneedles can be categorized into various types because different
types of microneedles are suitable for specific applications. The suitable length of microneedles for
drug delivery is 100 μm to 300 μm, but for blood extraction the appropriate length of microneedles is
1100 μm to 1600 μm [170]. Solid microneedles are suitable for cell surgery. Microneedles have been
reported for drug delivery, blood extraction, fluid sampling, cancer therapy, microdialysis, ink-jet
printing and sensing electrodes. Hollow Ti microneedles have been fabricated for blood extraction
using sputtering and deposition methods [171]. SiO2 hollow square microneedles have been reported
for flow delivery systems using electrochemical etching technique. Hollow out-of-plane SiO2
microneedles have been fabricated using lithography for cell surgery [162]. Stainless-steel hollow and
solid microneedles have been reported using surface micromachining and etching techniques for
dermal diphtheria and influenza vaccination [172]. Hollow out-of-plane Si microneedles have been
fabricated for TDD [4]. The extensive detail of materials used for microneedle’s designs, structure,
array size, fabrication techniques, analysis and application has been presented in Table 4, Table 5
and Table 6.
Int. J. Mol. Sci. 2011, 12
Table 4. Recent review of silicon microneedles.
Waseem et al.
2010 [4]
Chen et al.
2010 [173]
Zhang et al.
2010 [174]
Waseem et al.
2010 [3]
Zhang et al.
2009 [175]
Structure of
Shapes of
L = 200 µm
Di = 60 µm
L = 100 µm
D = 80 µm
5 ×5
Do = 150 µm
Star shape
Analysis type
Hollow/ Out-of
Solid/ Out-of
L = 200 µm
L = 200 µm
Di = 40 µm
tapered tip
Do = 425 µm
L = 200 µm
Side opened at
Di = 40 µm
Dt = 450 nm
30 × 30
10 × 10
6 ×6
Fluidic analysis
PLGA nano Particles
10 × 11
Fluidic analysis
4 ×4
Fluidic analysis /
9 ×9
6 ×6
Fluidic analysis
Not reported
Pig Skin
ICP etching
Deep reactive ion
etching (DRIE)
Transdermal drug delivery
Transdermal drug delivery
Human skin
film deposition
Transdermal drug delivery
Not reported
Potato skin/
Chicken skin
ICP etching
Bi-mask technique
Transdermal drug delivery
Drug delivery/fluid
L1 = 300 µm to
Ding et al.
2009 [172]
cut tip
900 µm
L2 = 300 µm
D2 = 200 µm
Mouse skin
L3 = 245 µm
Haq et al.
2009 [155]
Yu et al. 2009
Hollow/ Out-of
L1 = 180 µm
L2 = 280 µm
Human skin
Wet etching
Db = 180 µm
DP = 200 µm
D = 100 µm
Structural analysis
One-lead ECG
recording system
Transcutaneous drug
ECG measurement
Int. J. Mol. Sci. 2011, 12
Table 4. Cont.
Structure of
Shapes of
Coulman et
al. 2009
Chen et al.
2008 [177]
Roxhed et
al. 2008
Out-of plane
Donnelly et
al. 2008
Macro porous tip
Bhandari et
al. 2008
Not reported
Analysis type
L = 280 µm,
Db = 200 µm
Not reported
L1 = 310 µm,
L2 = 400 µm
Diffusion of nano particles
Human epidermal
Wet etching
Fluidic analysis
Pig skin
Fluidic analysis
Human skin
drug delivery
Transdermal drug
Transdermal drug
Square base canonical
Not reported
10 × 10
Not reported
Not reported
Blood sampling
Sharp 3D
Tip/Groovesembedded shaft
L = 270 µm,
Db = 240 µm
Fluidic analysis/Statistical
skin/Porcine skin
Wet etching
Photodynamic therapy
Drug delivery
of piglets
L1 = 800 µm,
Lee et al.
2008 [181]
Solid/ Outof plane
Db = 200 µm,
Dt = 20 µm,
3 ×3
Structural analysis
Not reported
L2 = 600 µm,
Wb = 300 µm
Notations: L = Length of needle, Wb = Base width, Do = Outer diameter, Di = Inner diameter, Db = Base diameter, Dt = Tip diameter, DP = Depth.
Int. J. Mol. Sci. 2011, 12
Table 5. Recent review of polymeric microneedles.
Structure of
Shapes of
Park et al.
Gomaa et al.
Solid/ Out-ofPMVE/MA
L = 600 µm
Donnelly et al.
Solid/ Out-of-
2010 [184]
Diffusion of trypan
cadaver skin
10 × 10
Wb = 250 µm
Analysis type
Fabrication techniques
Canonical/Square base
2010 [182]
2010 [183]
Array size/
Transdermal drug
UV lithography
L1 = 400 µm
11 × 11
L2 = 600 µm
14 × 14
Permeability with
L3 = 100 µm
19 × 19
microneedle density
Effect of Skin
Human skin
Laser micromachining
Drug delivery
Porcine skin
Molding process
Intradermal delivery
Structural/ Fluidic
Not reported
Not reported
L = 200 µm
Bodhale et al.
2010 [133]
Side opened/Sharp tip
Di = 30 µm
Hot embossing/UV excimer
25 × 25
Do = 150 µm
Drug delivery
Db = 300 µm
L = 500 to 1100 µm
Matteucci et
al. 2009 [185]
Rounded tip/Sharp tip
Bevel angle = 30°
10 arrays
Not reported
Not reported
Not reported
Lithography/Ni electroplating/ PDMS
replication/Hot embossing
to 40°
Han et al.
Sharp 3D Tip/
L = 880 ±20 µm,
Wb = 710 ±15 µm
T = 145 ±15 µm
Solid/ Out-of
2009 [186]
Jin et al. 2009
L = 200–1500 mm
Mouse skin and
Drug transportation
Transdermal drug
DXRL/Hot embossing
Oh et al. 2008
Solid/ Out-of
To improve skin
Sharp tip/ Spear
L = 200–500 µm
Mouse skin
Molding/Hot embossing
Emam et al.
2008 [188]
L = 500 μm,
Treatment of
Fluid analysis
Wb = 100 µm
hydrophilic molecules
Sharp tip
permeability for
Not reported
Int. J. Mol. Sci. 2011, 12
Table 5. Cont.
Structure of
Shapes of
Array size/
Analysis type
Fabrication techniques
Artificial skin of silicone
Not reported
Molding/KOH etching
L = 400 µm
Wb = 90, 120, 150,
Aoyagi et al.
Solid/ Out-of
230 µm
shape/Sharp tip
T = 115 µm
Drug delivery
Tip angle = 10°,
20°, 30°, 40°
Hsu et al.
SU-8 2050
2007 [189]
L1 = 236 µm
L2 = 350 µm
Not reported
Notations: L = Length of needle, T = Thickness, Wb = Base width, Do = Outer diameter, Di = Inner diameter, Db = Base diameter.
Table 6. Recent review of SiO2, glass, stainless-steel, and metallic microneedles.
Kim et al. 2010
Kato et al.
2010 [191]
Structure of
Shapes of
L = 700 µm
Solid/ In-plane
Hollow /Outof-Plane
2009 [172]
Jiang et al.
Spear/Sharp tip
Wb = 160 µm
T = 50 µm
Ding et al.
2009 [167]
Array size/
Analysis Type
Do = 5.5 µm
Elliptical tip
Not reported
Structural (Panitration)
L = 245, 300–900 µm
4 ×4
D2 = 200 µm
9 ×9
L = 3–4 cm
Not reported
Mouse skin
Infrared Laser
Vaccine delivery
Cellular function
Di = 3.5 µm
Tangentially cut tip
L = 77 µm
Circular Tip
Mouse skin
Human cadaver
image analysis
pulling technique
Int. J. Mol. Sci. 2011, 12
Table 6. Cont.
Jin et al.
2009 [187]
Hou et al.
2008 [192]
Structure of
Shapes of
Hollow/ Out-
Kolli and
Banga 2008
Not reported
2008 [194]
Analysis Type
L= 200–1500 mm
Not reported
Drug transportation
L = 120 µm
10 × 10
Fluidic analysis
L = 500 µm
27 needle per
Dt = 6 µm
L = 245, 300 µm
Verbaan et al.
Triangular tip/
D = 200, 300 µm
Tapered shaft
Beveled angle = 45°
Db = 250 µm
4 ×4
6 ×6
9 ×9
2007 [168]
Spare/ Sharp tip
1000 µm
Wb = 100 µm
Transdermal drug
Hot embossing
Not reported
Not reported
Mouse skin/
Drug transportation
Jacketed Franz
2007 [162]
Human skin
Transdermal drug
Transdermal drug
Transdermal drug
10 needles
Pressure Testing
Structural analysis
Drug delivery
L = 77 µm,
Do = 5.5 µm
diffusion cells
Waters HPLC
and serum
Tip taper angle = 60°
Shibata et al.
Mouse skin
L = 500, 750,
Parker et al.
Not reported
Structural analysis
10 × 10
Fluidic analysis
Not reported
Di = 3.5 µm
Cell surgery
L1 = 200 µm
Kim and Lee
2007 [195]
T = 10 µm
Tapered tip
L2 = 400 µm
T = 20 µm
SU-8 based UV
fluid sampling
Tapering angle < 5°
Tsuchiya et al.
2005 [171]
L = 1 mm
Di = 25 µm
Not reported
Fluidic analysis
Not reported
Sputter deposition
Do = 60 µm
Notations: L = Length of needle, T = Thickness, Wb = Base width, Do = Outer diameter, Di = Inner diameter, Db = Base diameter, Dt = Tip diameter.
Blood extraction
Int. J. Mol. Sci. 2011, 12
3.2. Forces Experienced by Microneedles during Penetration
Fluid is transported through hollow microneedles while solid microneedles are coated with
pharmaceutical materials to transfer the drugs into patient body. Microneedles are under the influence
of various forces during penetration such as bending, buckling, lateral, axial and resistive. To bear all
these forces, the design of microneedles is very important. Microneedles can break during penetration
into the skin because of these forces. An axial force is more dominant on the tip of microneedle during
insertion. This axial force is compressive and leads to buckling of the microneedle. The microneedles
also experience resistive force exerted by skin. Hence, in order to pierce the microneedle into skin, the
applied axial force must be greater than skin resistance. Due to uneven skin surface or human error
during needle penetration, bending may occur. So, it is very important to study the relation between
microneedle geometry and mechanical properties of the material for accurate microneedle design and
prediction of microneedles failure. The buckling force acting on the hollow microneedle during skin
insertion is given by [4,133,196,197].
Where, is Young’s modulus of material,
is moment of inertia of cylindrical section and
is length of the microneedle.
Moment of inertia (I) for hollow cylindrical section of microneedle is calculated by equation (14).
Where, is outer diameter and is inner diameter of hollow cylindrical section.
The bending force, which the microneedle can withstand without breaking is given by:
is the distance from vertical axis to the outer edge of the section [3,4,198].
The axial force (compressive force), which a microneedle can withstand without breaking is
given by:
is the yield strength of the material and
is cross-sectional area of the microneedle tip.
Microneedle experiences 3.18 MPa resistive forces exerted by the skin against penetration of
microneedle. To penetrate the microneedles into skin, the external applied force or pressure should be
greater than the resistive skin force. The resistive force offered by the skin before puncturing is given
by the following equation:
is the required pressure to pierce the microneedle into skin.
As the microneedle penetrates the skin, the resistive force falls drastically [199]. After the skin is
pierced by the microneedle, the only force that acts on the microneedles is the frictional force due to
contact of tissue with the microneedle.
Int. J. Mol. Sci. 2011, 12
3.3. Fabrication of Microneedles
Various fabrication techniques have been developed and used for microneedles fabrication such as
hot embossing [156], photolithography [162], micropipette pulling technique [167], surface
micromachining [172], bi-mask technique [175], laser micromachining [179], micro-molding [181],
deep x-ray lithography [200], DRIE [176], lithography, electroplating, molding (LIGA) [201], UV
excimer laser [202], coherent porous Si etching (CPS) [203], injection molding [204] and ICP
etching [3,4]. In these processes, Si and polymer can be used as substrate materials for
microfabrication. Each fabrication technique has its own advantages and limitations. A detailed
discussion on microneedles fabrication techniques will be presented in a subsequent paper.
Lithography and DRIE techniques are mostly used for fabrication of silicon microneedles.
Deposition and etching are the most important phenomenon during the development of microneedles.
Deep holes or free standing structures can be fabricated in silicon wafer with the help of anisotropic
etching process. These high aspect ratio structures are of considerable interest in developing micro
devices for various applications. The general steps of silicon microneedles fabrication are wafer
cleaning, photoresist coating, soft back, masking, exposure, development, hard back, and lift off. These
steps can be repeated according to requirements with desired parameters. For polymeric microneedles
molding, hot embossing, and laser drilling are promising fabrication techniques. The general steps for
polymeric microneedles fabrication are sheet preparation, mold preparation, heating and pressing
mold, de-molding and laser drilling for different lumens/reservoir.
3.4. Microneedles Testing
The concept of microneedles was introduced three decade ago but the first microneedle array for
TDD was fabricated in 1998. After that, various researchers have been involved in developing the most
suitable fabrication method and optimal design of microneedles for biomedical applications. After
2005, the interests of researchers changed and they shifted their attention towards the testing of
microneedles along with design and fabrication. Most of the research groups have been involved in the
structural analysis and skin penetration tests of microneedles in 2010. The testing of microneedles has
been reported on potato skin, chicken skin [175], mouse skin [156,172,180,187,193], cadaver
skin [167,182], pig skin [173,177], chicken leg, beef liver [204], and human skin [156,174,178,183].
Microneedle patches coated with solid state influenza vaccine have been reported to improve the
effectiveness of the vaccine when tested on mouse skin [190]. Dry coated microinjection arrays have
been developed to deliver HSV-2-gD2 DNA vaccine to sensitive regions of mouse skin [205]. The
pretreatment of skin by microneedles was combined with the use of highly water soluble pegylated
naltrexone for its transdermal delivery at different concentrations [206]. A new design of probe
integrated with hollow microneedles for atomic force microscope (AFM) has been developed to realize
cellular function analysis in a single living cell [176]. The geometry of microneedles affects its
strength. The shear strength of hollow silicon microneedles can be increased by variation in
microneedles geometry [207]. Using novel microneedle technology, hydrophobic dye called Nile red
has been delivered into porcine skin [169]. The effect of microneedle geometry has been studied on the
transport of a fluorescent dye into human skin [208]. To envisage the effect of microneedle geometry
Int. J. Mol. Sci. 2011, 12
and force of application, optical coherence tomography has been used by penetrating microneedles
arrays into neonatal porcine skin [209]. Short densely packed microinjection array has been developed
to see the effect of strain rate on the precision of penetration into human skin [210]. Reference [211]
has investigated that influenza virus like particles coated on microneedles can cause stimulatory effect
on langerhans cells in human skin. The super short microneedles have been fabricated using Si wet
etching technology and tested for TDD into human skin [212]. The separable arrowhead microneedles
have been introduced and tested for painless delivery of drugs and vaccines into human cadaver
skin [213]. A minimally invasive system has been developed using microneedle electrode array to
deliver macromolecular drugs to the deep skin tissues and tested on hairless rat skin [214]. Solid
silicon microneedles arrays have been used with different lengths and geometry to penetrate epidermal
membrane of human cadaver skin [215]. The microneedles coated with influenza virus like particles
have been used to test the immunogenicity and protective efficacy after vaccination of mice skin [216].
Microneedles rollers have been developed and tested on human and porcine skin to increase skin
permeability and to treat skin for cosmetic purposes [167]. Microneedles have been used to deliver
PLGA (poly-lactic-co-glycolic-acid) nanoparticles in the human skin [174]. Solid polymeric
microneedles have been developed to investigate the transepidermal water loss measurements of
dermatomed human skin [183]. The efficacy of transdermal delivery of insulin has been investigated
by using microneedles rollers on diabetic rats [217]. The administration of virus like particles influenza
vaccine has been studied using microneedles patch on lungs and bone marrow cells of mouse [218].
Hollow microneedles array with sonophoretic emitter has been used on pig skin to improve the
efficiency of drug delivery [173].
4. Discussion
MEMS and NEMS based microfluidic devices have many important characteristics that make them
attractive for biomedical applications. Microfluidic devices have the ability to control their physical
and chemical characteristics from a very small scale up to the nanometer range. These devices have
made it possible to meet critical medical needs such as nearly constant drug level at the site of action,
prevention of peak-valley fluctuations, site specific drug delivery, reduced side effects and increased
therapeutic effectiveness [4]. However, there are certain medical conditions for which constant drug
release pattern is not suitable. These conditions demand delayed release of drug. Such a release pattern
is known as pulsatile release. Recent research has shown that some diseases have a predictable cyclic
rhythm and the timing of drug release can significantly improve the outcome of a desired effect [219].
This condition requires release of drug as a pulse after a time delay. Some of the diseases where
pulsatile drug delivery devices are promising include duodenal ulcer, cardiovascular diseases, arthritis,
asthma, diabetes, neurological disorder, cancer, hypertension and hypercholesterolemia [4]. That is
why study of pulsatile flow is extremely important at small scales in microfluidic devices. Using
MEMS and NEMS technology, complex drug release patterns (such as simultaneous constant and
pulsatile release) can be achieved using integrated microfluidic systems. Microfluidic devices have
ability to control both release time and release rate. Micropumps and microneedles are essential
components for such biomedical systems. Micropumps are used for fluid transport and microneedles
provide interface between drug reservoir and patient body [4,13,6]. Material selection is a critical issue
Int. J. Mol. Sci. 2011, 12
in biomedical devices. Si has been widely used as material for such microfluidic devices, but
polymeric materials like PGA, PDMS, PMMA, PLLA, PLA, PC, etc. are replacing Si due to
biocompatibility, low cost, ease of fabrication and excellent structural properties. Various factors are
important during the selection of micropumps for particular biomedicine applications. Operating
voltage, pressure and flow rate of micropumps are critical issues to analyze the performance and
suitability of micropumps for certain medical applications. A schematic illustration of operating
voltages and flow rates of mechanical micropumps is shown in Figure 22.
Figure 22. Comparison of voltage versus flow rate for mechanical micropumps.
Piezoelectric and electromagnetic mechanical micropumps have been reported extensively for
microfluidic systems among the mechanical micropumps. The major limitation related to these types
of micropumps is very high operating voltage [224]. Electrostatic micropump is easy to fabricate on
integrated microfluidic systems but it also requires high operating voltage. The ICPF micropump has
an adequate flow rate at a relatively low operating voltage but with complex geometry. The schematic
illustration of operating voltages and flow rates of non-mechanical micropumps is shown in Figure 23.
Int. J. Mol. Sci. 2011, 12
Figure 23. Comparison of voltage versus flow rate for non-mechanical micropumps.
In non-mechanical micropumps, MHD micropump has gained more attention in recent years and
has been presented by many researchers for microfluidic systems. However, the electrochemical type
of micropump is more suitable for low voltage and high flow rate applications. Most micropumps and
microneedles have been reported in literature as an individual device for medical applications. Only a
few researchers have presented integrated devices [227]. Integration of micropumps and hollow
microneedles is a great challenge, but research on solid microneedles coated with nano-particles and
drugs has recently commenced in the biomedical field. However, hollow microneedles are more
attractive for fluid/drug transport. Microneedles can be integrated with micropumps or used as
stand-alone biomedical device. Various types of microneedles have been presented by different
researchers. In hollow microneedles, the side opened double lumen reservoir based microneedles are
more suitable for fluid transport. The pressure difference in the lumen regions is useful to avoid the
clogging effect. In solid microneedles, the sharper tip microneedles are more practical for drug
transport. Effectiveness of drug transport has also been presented in recent years by various researchers
using microneedles on mouse, pig, chicken and human. Various researchers have reported the structure
and fracture analyses of microneedles by applying force and pressure to predict the bending and failure
of microneedles. The schematic illustration of comparison of force and stress for microneedles has
been shown in Figure 24.
Int. J. Mol. Sci. 2011, 12
Figure 24. Comparison of force versus stress on microneedles.
5. Challenges and Future Aspects
In biomedical field, there are many challenges relating to the microfluidic devices (micropumps,
microneedles) such as design level issue, fabrication level issue, packaging level issues, use in
practical application, etc. At design and fabrication level, the most important issues and specifications
that must be fulfilled by the micropumps for particular applications are appropriate design for
maintaining a specific flow rate, control of back pressure, dosing accuracy, drugs resistive material
selection for fabrication, ease of fabrication, energy utilization, power supply at such small level,
bubble tolerance, durability and reliability. The most important issues relating to microneedles at
design and fabrication level are avoidance of clogging effect, suitable length, robustness, strength,
sharp tip to avoid pain, drugs resistance, less fabrication cost, reliability, biocompatibility, etc. Suitable
batch fabrication techniques need to be adopted to reduce cost of devices. Packaging of these devices
is very important consideration [229]. Packaging should be robust and strong enough to prevent
infection or damage of microfluidic devices. Simultaneously, the unintentional discharge of drug/fluid
during storage from the reservoir should be prevented. A protecting wrap may be possibly required to
secure such small size devices like micropumps and sharp tip microneedles. Mostly the micropumps
and microneedles reported in literature have been proposed as stand-alone devices. Integration of
micropump and microneedles is a great challenge that limits the use of these devices commercially for
biomedical applications. The final cost of these delivery devices should be affordable for the end
users/patients. The trend is now shifting towards the use of polymeric materials like PGA, PDMS,
PMMA, etc. for the fabrication of micropumps and microneeldes to overcome most of the above
described issues as these materials are cheap, biocompatible, exhibit excellent mechanical/chemical
Int. J. Mol. Sci. 2011, 12
properties, etc. [4,13,19,202]. Zosano Pharma [140] has developed the user friendly and simple TDD
patch system that can deliver vaccines, proteins, peptides and small molecules. 3 M has developed the
microneedles based transdermal system demonstrating good results on research level studies for
peptides, vaccines and protein [230]. Birchall et al. [231] conducted a survey to learn more about the
opinion of end users with regard to the convenience, efficacy and worth of microneedle technology.
Research on MEMS-based delivery devices shows that these devices are suitable for commercial
applications. However, the development of these devices is limited to research level due to some
factors such as investment, expertise for device development, marketing, awareness of public,
motivation, lack of collaboration between companies and research institutes, medical staff
training/recommendations, etc. Surveys, seminars, workshops, etc., need to be organized to promote
the benefits and convenience of using these delivery devices to end users. The availability of reliable
and manufacturable microfluidic devices will have a strong impact on the biomedicine field to meet
critical health care needs.
6. Conclusions
Fluid transport using microfluidic devices such as micropumps and microneedles is a relatively new
and attractive method that has many advantages. Microfluidic devices have received much more
attention in recent years due to their potential applications in the biomedicine field. Various types of
micropumps and microneedles structures using different materials like glass, silicon, metals, polymers,
etc., have been reported for biomedical applications, but Si has been mostly used as substrate material
for fabrication of microfluidic devices among other materials. Si is brittle and always some risk
involves for health care. Biocompatibility is very important for health and because of this reason the
trend is transferring towards polymeric materials. Most polymers, e.g., PGA, PDMS, PLA and PMMA,
are very suitable for biomedical devices due to their good biocompatibility, low cost, ease of
fabrication and excellent chemical and mechanical properties. Solid microneedles are easy to fabricate
and have more strength than hollow microneedles. However, the disadvantage of solid microneeldes is
the risk of fracture/breaking within the skin after being inserted. Drug particles can be coated in
restricted amounts. Hollow microneeldes are considered more suitable for TDD systems due to the
precise delivery of the desired amount of drug at a specific site with rapid action. ICPF and
electrostatic micropumps are considered suitable for drug/fluid delivery systems due to low operating
voltage and direct integration with electronic circuit respectively. The disadvantages of hollow
microneeldes and micropumps are complicated/costly fabrication, clogging effect, back pressure,
requirement of flow rate regulation, etc. Based on the presented literature review, the authors conclude
that MEMS based microfluidic devices for biomedical applications remain at the research level. Only a
few devices have been converted into commercial products due to some important issues like
complicated structure of microfluidic devices, difficulties in integration with other devices, investment,
expertise for device fabrication, marketing, public awareness, lack of collaboration between companies
and research institutes, medical staff training/recommendation and finally packaging. To present
microfluidic devices for practical applications in medical field, motivated researchers still need to
continue their work on the development of microfluidic devices.
Int. J. Mol. Sci. 2011, 12
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