Structural and microfluidic analysis of hollow side-open polymeric microneedles for transdermal drug delivery applications

RESEARCH PAPER

Structural and microfluidic analysis of hollow side-open polymericmicroneedles for transdermal drug delivery applications

Dhananjay W. Bodhale Æ Asim Nisar ÆNitin Afzulpurkar

Received: 8 April 2009 / Accepted: 8 June 2009 / Published online: 7 July 2009

� Springer-Verlag 2009

Abstract In this paper, we present a new design of hol-

low, out-of-plane polymeric microneedle with cylindrical

side-open holes for transdermal drug delivery (TDD)

applications. A detailed literature review of existing

designs and analysis work on microneedles is first pre-

sented to provide a comprehensive reference for researchers

working on design and development of micro-electrome-

chanical system (MEMS)-based microneedles and a source

for those outside the field who wish to select the best

available microneedle design for a specific drug delivery

or biomedical application. Then, the performance of the

proposed new design of microneedles is numerically

characterized in terms of microneedle strength and flow

rate at applied inlet pressures. All the previous designs of

hollow microneedles have side-open holes in the lumen

section with no integrated reservoir on the same chip. We

have proposed a new design with side-open holes in the

conical section to ensure drug delivery on skin insertion.

Furthermore, the present design has an integrated drug

reservoir on the back side of the microneedles. Since

MEMS-based, hollow, side-open polymeric microneedles

with integrated reservoir is a new research area, there is a

notable lack of applicable mathematical models to analy-

tically predict structural and fluid flow under various

boundary conditions. That is why, finite element (FE) and

computational fluid dynamic (CFD) analysis using ANSYS

rather than analytical systems has been used to facilitate

design optimization before fabrication. The analysis has

involved simulation of structural and CFD analysis on

three-dimensional model of microneedle array. The effect

of axial and transverse loading on the microneedle during

skin insertion is investigated in the stress analysis. The

analysis predicts that the resultant stresses due to applied

bending and axial loads are in the safe range below the

yield strength of the material for the proposed design of

the microneedles. In CFD analysis, fluid flow rate and

pressure drop in the microneedles at applied inlet pressures

are numerically and theoretically investigated. The CFD

analysis predicts uniform flow through the microneedle

array for each microneedle. Theoretical and numerical

results for the flow rate and pressure drop are in close

agreement with each other, thereby validating the CFD

analysis. For the proposed design of microneedles, feasible

fabrication techniques such as micro-hot embossing and

ultraviolet excimer laser methods are proposed. The results

of the present theoretical study provide valuable bench-

mark and prediction data to fabricate optimized designs of

the polymeric, hollow microneedles, which can be suc-

cessfully integrated with other microfluidic devices for

TDD applications.

Keywords Computational fluid dynamic analysis �Drug delivery � Hollow, polymeric microneedle �Hot embossing � Transdermal drug delivery

1 Introduction

Transdermal drug delivery (TDD) refers to the movement

of pharmaceutical compound across the skin to reach the

systemic circulation for subsequent distribution in the

D. W. Bodhale (&) � A. Nisar � N. Afzulpurkar

School of Engineering and Technology, Asian Institute

of Technology (AIT), Pathum Thani 12120, Thailand

e-mail: [email protected]; [email protected]

A. Nisar

e-mail: [email protected]

N. Afzulpurkar

e-mail: [email protected]

123

Microfluid Nanofluid (2010) 8:373–392

DOI 10.1007/s10404-009-0467-9

human body. TDD systems cover a wide range of non-

invasive and minimally invasive technologies for deliver-

ing drugs and vaccines across the skin (Barry 2001;

Prausnitz 2004; Schuetz et al. 2005). Transdermal drug

delivery devices can be divided into active and passive

devices, based on the technologies used for skin perme-

ation. In passive devices, the methods used for skin

permeation are chemical enhancers, emulsions and lipid

assemblies, as well as biological methods such as peptides

(Prausnitz 2004; Schuetz et al. 2005; Schreier and Bouwstra

1994; Karande et al. 2004). A recent trend toward

increasing use of active methods for skin permeation has

been reported by Brown et al. (2006). The most common

active methods of skin permeation are jet injectors, ionto-

phoresis, electroporation, ultrasound, microneedles, pow-

der injection, ablation and tape stripping (Arora et al.

2008). One of the major drawbacks of TDD systems has

been their inability to deliver the drugs through the skin

within the desired therapeutic range. To overcome this

limitation, many studies have been conducted on new drug

delivery methods using emerging micro- and nanotech-

nologies. The major focus of MEMS for drug delivery has

been toward the development of microneedles for mini-

mally invasive TDD applications. Many developments in

the use of micron-size needles have been reported recently

to dramatically increase transdermal delivery, especially

for macromolecules. Using MEMS technology, micronee-

dles with various sizes, shapes and materials have been

fabricated. In vitro and in vivo studies have shown suc-

cessful applications of solid microneedles in terms of skin

permeability for a broad range of molecules, DNA vaccines

and nanoparticles. In addition, hollow microneedles have

also been developed for insulin delivery.

Most common traditional methods to deliver drugs are by

means of oral delivery and with the help of syringes and

hypodermic needles. The use of syringes and hypodermic

needles has been utilized to deliver drug for over 150 years

(McAllister et al. 2000). Hypodermic syringes were first

developed by C. Pravaz and A. Wood in the 1850s

(McGrew and McGrew 1985). These hypodermic syringes

are mostly used for body fluid extraction, vaccination and

medication. Hypodermic syringes with pointed, hollow

needles have been fabricated with various materials using

different methods; however, the fundamental design of

these needles remains the same since their first use. There

are some disadvantages of conventional oral drug delivery

and injections, such as oral drug delivery can create drug

absorption problems in the gastrointestinal tract or side

effects of liver damage. In addition, the syringes and

hypodermic needles penetrate into the deep dermis layer of

the skin and create pain and infections. Moreover, con-

ventional hypodermic needles pose various health hazards,

such as transmission of blood-borne pathogens (e.g., HIV

(human immunodeficiency virus) and hepatitis; Jagger et al.

1988). As needles pierce the skin easily for drug delivery,

injection is still a prominent method. To avoid all the lim-

itations of conventional drug delivery, microneedles can

play an important role for drug delivery without causing

pain and infections. Currently, the size of some of the

smallest needles for injection available in the market is in

the range between 30 and 31 gauge, with outer diameters of

305–254 lm (Zachary and Nicholas 2005). The micronee-

dles must have enough strength to pierce the outer epider-

mis of the human skin and should not have any type of side

effect on the skin. Hence, it becomes important to study the

properties of the skin while designing the microneedle.

Human skin is composed of three layers: epidermis, dermis

and subcutaneous tissue. The epidermis can be divided into

three important layers: the outer most stratum corneum

composed of fibrous keratinized dead cells, a layer of living

epidermis and the basal layer or stratum germinativum.

Stratum corneum is about 10–15-lm thick. The living

epidermis is approximately 150-lm thick. Stratum corneum

forms a protective barrier layer that prevents the loss of

body fluids and blocks the entry of any external material

into the body. Hypodermic needles cause pain during

injection because they penetrate deep into the skin and

make excessive contact with sensory organs. Microneedles,

on the other hand, are fabricated using MEMS technology

that penetrates only the epidermis layer for minimally

invasive transdermal drug delivery and fluid sampling.

Sharp projections of microneedles that penetrate the stratum

germinativum, and not beyond, are pain free, since there are

no sensory organs or nerve endings that could trigger any

sensation of pain. In addition, such projections cannot cause

bleeding since blood capillaries cannot be reached. Various

researchers have reported integration of microneedle arrays

with other microfluidic devices such as micropumps for

TDD applications (Zahn et al. 2004; Ma Bin et al. 2006;

Roxhed et al. 2007; Yang Ruoting et al. 2006).

MEMS-based microfluidic devices have dimensions in

micrometers. A microneedle is different from standard

hypodermic needles used in medical applications, as gen-

erally the length of the MEMS-based microneedles are less

than 1 mm. Thus, microneedles are significantly smaller in

length than ordinary needles. A significant distinction that

can be made in microneedles is whether microneedles are

solid or hollow. Hollow microneedles have an internal bore

or lumen, which allows flow of fluid or drug through the

microneedles. Solid and hollow microneedles have been

fabricated with the following design variations: solid

microneedles that pierce the skin to make it more permeable,

solid microneedles coated with dry powder drugs or vac-

cines for dissolution in the skin, polymer microneedles

with encapsulated vaccine for release in the skin, and

hollow microneedles for direct fluid injection into the skin.

374 Microfluid Nanofluid (2010) 8:373–392

123

Microneedles can also be classified according to the

fabrication process involved. Based on the fabrication

process, the microneedles are classified as in-plane and out-

of-plane microneedles. Many different designs of out-of-

plane and in-plane, solid and hollow microneedles have

been developed. In-plane microneedles are fabricated with

the microneedles parallel to the plane of the substrate. Out-

of-plane microneedles are fabricated with the microneedles

normal to the plane of the substrate. A schematic illustra-

tion of in-plane and out-of-plane microneedles is shown in

Fig. 1. The limitation of in-plane microneedles is that it is

difficult to fabricate microneedle arrays with two-dimen-

sional geometry and hence cannot be used for biomedical

applications to deliver the drugs in the desired amount.

(Griss et al. 2002; Gerstel and Place 1976). However, the

length of the in-plane microneedle can be controlled very

accurately than the out-of-plane microneedles (Roxhed

2007). Out-of-plane microneedles are suitable for fabrica-

tion in two-dimensional array by wafer-level processing

and are desirable for transdermal drug delivery (Griss et al.

2002; Henry et al. 1998a, b). In addition, the density of out-

of-plane microneedle array plays an important role in

controlling flow resistance. The flow resistance can be

decreased by increasing the density of the microneedles in

an array (Griss et al. 2002). However, to control the length

and high aspect ratio is a challenge in the out-of-plane

fabrication process (Roxhed 2007). The different catego-

ries of microneedles are summarized in Table 1.

Material selection is an important issue, as microneedles

may break during skin insertion and can cause undesirable

effects on health. Many research groups (Griss et al. 2002;

Khumpuang et al. 2003; Gardeniers et al. 2003; Stoeber

and Liepmann 2000, 2002; Paik et al. 2004; Choi et al.

2006; Mukherjee et al. 2004; Shibata et al. 2007; Wilke

et al. 2005a, b) used silicon as material for microneedle

fabrication. As silicon is a brittle material (Runyam and

Bean 1990), it can break during skin penetration and cause

harmful problems due to particles remaining in the skin. In

some cases, the silicon particles can pass via blood directly

to the heart and cause blockage of the artery. The patient

may die within a few minutes (Khumpuang et al. 2003).

Thus, silicon is not an established biomaterial (Runyam

and Bean 1990). Many polymer materials have a long

history of biocompatibility, possess excellent mechanical

properties (Ambrose and Clanto 2004) and are cost

Masking material

Silicon Substrate

Top view

Top view

Side View

Top viewOut-of-Plane

structure

In-plane structure

Silicon substrate

Fig. 1 A schematic illustration of in-plane and out-of-plane

microneedles

Table 1 Categories of microneedles

Geometry

In plane Out of plane Application

Structure Solid Silicon, metal, polymer, silicon

dioxide

Silicon, metal, polymer,

silicon dioxide

Electrodes for sensing,

stimulation and perforation

Hollow Silicon and nitride/oxide Silicon, metal Injection, sampling of fluids

Metal Silicon dioxide

Polymer Silicon

Polymer

Advantages Ease of integration with electronics

on a single substrate

High density arrays of

microneedles can be realized

Microfluid Nanofluid (2010) 8:373–392 375

123

effective. Polymeric microneedles can be fabricated from

novel biocompatible engineering materials such as plastics,

biodegradable polymers and water-soluble polymers (e.g.,

polycarbonate, polyglicolide acid (PGA) and carboxy-

methylcellulose). Some research groups (Park et al. 2005;

Aoyagi et al. 2007; Yu 2005; Jeong et al. 2008) have used

polymers as material for microneedle fabrication. Park

et al. (2005) and Kuo and Yukon (2004) successfully

fabricated and characterized polymeric microneedles for

various drug delivery applications.

For skin piercing, the microneedle must have enough

strength and should not fracture during insertion. The skin

resistance forces acting on the microneedle depend directly

on the cross-sectional area of the microneedle tip (Aggar-

wal and Johnston 2004). Hence, it is necessary to decrease

the area of the tip as much as possible. However, the area

cannot be reduced beyond a certain limit in case of a

hollow microneedle. There is also a risk of clogging in

hollow microneedles with openings at the top of the mi-

croneedles (Griss et al. 2002). Most of the research groups

(Khumpuang et al. 2003; Gardeniers et al. 2003; Stoeber

and Liepmann 2000; Mukherjee et al. 2004; Shibata et al.

2007; Yu 2005; Wilke et al. 2005a, b; Davis et al. 2005)

have developed microneedles with openings at the tip,

designed symmetrically and asymmetrically. Griss et al.

(2002) designed microneedles with a side opening to

increase the area for drug exposure. The geometry of the

microneedle decides the strength. Aggarwal and Johnston

(2004) reported that the cylindrical microneedles can

withstand more load than square or rectangular micronee-

dles with the same area.

To design the microneedle for drug delivery, certain

other parameters need to be taken into account, such as

pressure, pressure drop, velocity of fluid and flow rate.

Drug delivery applications require precise flow rates

between 10 and 100 lL/min (Prausnitz et al. 2000).

Applied pressure between 10 and 100 kPa is recommended

for various designs of micropumping devices available in

the market that can be integrated with microneedles

(Khumpuang et al. 2003). To get the desired and continu-

ous flow rate, the dimensions and position of the hole of the

microneedle also plays an important role. The flow rate

increases with decrease in length and increase in the inner

diameter of the microneedle (Khumpuang et al. 2003). In

addition, a flat, hollow tip may cause obstruction to the

fluid flow due to punched skin tissues (Gardeniers et al.

2003). Hence, it is necessary to keep the hole off-centered.

The forces created from the flow resistance can cause

breaking of the device and hence the flow resistance should

be controlled (Stoeber and Liepmann 2000). The flow

resistance decreases with increase in microneedles density

per unit area (Griss et al. 2002; Stoeber and Liepmann

2000).

A variety of technologies and fabrication methods have

been developed to fabricate microneedles for drug delivery

applications. These include silicon microfabrication and

polymeric micromolding techniques (Varadan 2006). Very

few research groups have used polymeric fabrication

techniques. Among them Aoyagi et al. (2007) and Park

et al. (2005) used injection-molding technique for poly-

meric microneedle fabrication. To fabricate hollow mi-

croneedles with a side opening, it is very important to

develop a good technique of fabrication. Aoyagi et al.

(2007) used ultraviolet (UV) excimer laser fabrication

technique to create holes in polymeric microneedles.

The microneedle design analysis involves strength

modeling and CFD analysis. Many research studies have

been conducted for structural and microfluidics (flow)

analysis of the microneedles. Shibata et al. (2007) con-

ducted numerical and experimental studies for mechanical

stability analysis of hollow cylindrical microneedle using

silicon dioxide (SiO2) material. Wilke et al. (2005a, b) used

numerical analysis for fluidic system to ensure uniform

release of the fluid from each hollow microneedle.

Aggarwal and Johnston (2004) used finite element modeling

(FE) analysis for prediction of various forces acting on the

microneedle during skin insertion. Aoyagi et al. (2008)

used FEM simulation and confirmed that the stress con-

centration occurred severely at the tip area of the micro-

needle, and the sharp tip angle of the microneedle could be

easily inserted into the skin. Paik et al. (2004) used in-plane

single crystal silicon microneedle array for drug delivery

and performed FEM simulation and experiments on

microneedle for stress analysis by applying bending load at

the tip of the microneedle. Stoeber and Liepmann (2002)

developed a syringe using MEMS technology and designed

fluid mechanical model to avoid the clogging of the

microneedle at the inlet by increasing the internal diameter

of the syringe in two steps. Stoeber and Liepmann (2002)

performed the CFD simulation using CFD ACE? Software

to calculate the shear rate in the flow for the appropriate

scaling of the stepped internal lumen diameters. The study

of the fluidic system of the microneedle was also done

experimentally by many researchers (Griss et al. 2002,

Griss et al. 2003; Gardeniers et al. 2003; Matteucci et al.

2008). A review of design and analysis of various cate-

gories of microneedles along with types of fabrication

method, applications and performance characteristics are

summarized and referenced in Tables 2 and 3, respectively.

The first and foremost requirement that any design of

microneedle must meet is its ability to penetrate the skin

without breaking. Metallic microneedles are strong enough

to penetrate the skin. However, polymeric materials for

microneedles must be selected to have adequate mechani-

cal strength. Another important factor in microneedle

design is the shape of the microneedle. The sharp tip of the


123

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(spec

ific

appli

cati

on

N/r

)


123

microneedle greatly affects the ability of the microneedle

to penetrate the skin. Other dimensional parameters that

can influence the insertion force of the microneedles into

the skin are microneedle length, width and overall profile.

In the present study, the authors conducted theoretical

structural and microfluidic analysis of hollow polymeric

microneedles with side-open holes for transdermal drug

delivery applications, such as treatment of cardiovascular

and hemodynamic dysfunctions and insulin delivery for

diabetic patients. The existing TDD devices have many

limitations, such as material biocompatibility, low strength

of microneedles, disproportionate drug delivery, etc.

Hence, there is an emerging need for the development of

microneedles array with biocompatible material having

enough strength to pierce the skin and achieve propor-

tionate and continuous drug supply. All the previous

designs of hollow microneedles have side-open holes in the

lumen section (Griss et al. 2002) with no integrated res-

ervoir on the same chip. The authors have proposed a new

design with side-open holes in the conical section to ensure

drug delivery upon skin insertion. Furthermore, the present

design has an integrated drug reservoir on the backside

of the microneedles. The authors have also proposed a

fabrication process that will facilitate the fabrication of

microneedles, as well as the drug reservoir at the same

time.

Various numerical studies have been conducted previ-

ously on the design and analysis of microneedles as men-

tioned above. However, all the previous studies cover

either structural or fluid analysis only. In the present study,

the authors have performed both the structural and CFD

analysis and have validated the numerical predictions with

theoretical calculations. Since MEMS and microfluidics are

a new research area, there is a notable lack of applicable

mathematical models and material information to predict

fluid flow through micron-sized structures. There are no

published accurate mathematical models for hollow side-

open polymeric microneedles with integrated reservoir to

analytically predict structural and fluid flow under various

boundary conditions. That is why, numerical analysis

involving structural and microfluidic analysis has been

performed to facilitate design optimization before fabri-

cation. The performance of the microneedle is character-

ized in terms of its strength and flow rate at applied inlet

pressures. The proposed design is based on polymeric

fabrication technology using biocompatible and biode-

gradable PGA (polyglicolide acid) material. To achieve

better skin permeation and structural strength at desired

flow rates, authors designed cylindrical microneedles with

ultra-sharp tip and T-shaped lumen. To avoid problem

of clogging at the inlet due to shear rate (Stoeber and

Liepmann 2002), the design was modified by increasing the

lumen diameter at the inlet. Our proposed microneedleTa

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123

arrays are designed to be integrated with micropumping

devices, which can provide 10–100 kPa of pressure at the

outlet. Although the present study discusses only theoreti-

cal design and analysis of hollow out-of-plane polymeric

microneedles with side-open holes, the authors have also

proposed appropriate fabrication process to highlight that

the proposed fabrication process will facilitate integration

by fabricating microneedles with integrated drug reservoir

at the same time, which has not been done previously. The

proposed fabrication process combines micro-hot emboss-

ing and excimer laser to fabricate hollow polymeric

structures and integrated drug reservoir.

2 Proposed design and fabrication process for hollow

microneedles

2.1 Mechanical design specifications

A schematic illustration of the proposed microneedle

design with dimensions is shown in Fig. 2a and b.

The proposed hollow polymeric microneedle has an

ultra-sharp tip for easy skin insertion. The microneedle has

two geometrical sections; the first one is a cylindrical

section, which provides enough strength to the micronee-

dle, and other one in conical section, which provides ease

of skin insertion. The length of the microneedle is 200 lm

and the outer diameter of the microneedle is 150 lm to

increase the bending strength. As shear rate during fluid

flow is responsible for clogging (Stoeber and Liepmann

2000), to reduce the shear rate, the lumen is designed in

two steps to avoid clogging at the inlet. The proposed

design becomes strong enough to pierce the skin easily due

to side-open lumen. In extreme cases, there will not be any

effect on drug delivery after the damage to the microneedle

tip, as the lumen is quite away from the tip area.

2.2 Microfluidic model design

The constant flow rate through the microneedles in skin is

important for drug delivery application. Many researchers

proposed the tip-open, hollow microneedle designs for

drug delivery. However, during skin insertion, the micro-

needle tip may be damaged and can cause resistance to the

drug flow. Hence, it is a good idea to keep side-open holes

in microneedles to avoid resistance to the fluid flow inside

the skin. Fluid model of our proposed microneedle design

for CFD analysis is shown in Fig. 3.

2.3 Fabrication process

Micro-hot embossing technique and ultraviolet (UV) ex-

cimer laser beam technique have been proposed for

microneedle array and lumen fabrication. The micro-hot

embossing has many advantages, as this process includes

comparatively simple machines, low mechanical influence

on mold insert, small thermal cycles approximately 40�C

and high-quality replication (Becker and Heim 1999). In

addition, this technique provides the ability to fabricate

extremely small structures with thin polymer foils. The

mold insert for hot embossing can be fabricated from sil-

icon and nickel, which are the most used materials for

molds (Becker and Heim 1999). It can be fabricated in a

number of ways from CNC machining of stainless steel for

microstructures of 100 lm range, for LIGA technologies,

and of microstructures with dimensions of few micrometers

range and high aspect ratios (Becker and Heim 1999).

Some researchers have used microinjection molding and

spin coating technique to fabricate polymer microneedles

(Aoyagi et al. 2007; Park et al. 2005; Yu 2005). Micro-hot

embossing technique is different from the above techniques,

as in hot embossing the heating temperature is just above the

glass transition temperature (Tg) (Varadan 2006). In addition,

we can use polymer films as starting materials instead of

pallets of polymers used in injection molding (Varadan 2006).

The process flow for hot embossing is shown in Fig. 4

(Deodhar et al. 2005). Due to small thermal cycles, this

method provides low internal stresses with high structural

replication accuracy and flexibility (Becker and Heim 1999).

UV excimer laser beam is a very useful and reliable

micromachining technique for organic materials such as

polymers, plastic, etc. UV excimer laser does not burn or

melt the material like infrared (IR) lasers and ablates the

organic material by breaking the individual molecules by

converting from solid state to gaseous state (Faulkner

2006). The complex geometric polymeric structures can be

easily ablated using excimer laser, which is difficult byFig. 2 a Schematic illustration of 3D microneedle with base and

b dimensions of the structure (all dimensions are in lm)


123

conventional lithographic process (Yang and Pan 2003).

Some working parameters, such as laser energy, shot

number, scanning velocity and pulse repetition rate are need

to be set for excimer laser fabrication (Yang and Pan 2003).

The ablation depth in the polymer material depends on the

shot number. With increase in shot number, the ablation

into the polymer material increases (Yang and Pan 2003).

The drug flow rate through the microneedles should not

be more than a certain limit. The number of microneedles in

an array also depends on the overall size of the microfluidic

device. Therefore, the present design proposes 25 micro-

needles per array. The designed structure of the micronee-

dle array chip is 3,600 9 3,600 lm2. Each microneedle is

separated by 600 lm of center-to-center distance. The step-

by-step proposed fabrication process is shown in Fig. 5. By

preparing polymer material layer, hot embossing process is

carried out by heating the pressing tools (stamp and mold

insert) and the polymer layer. After the embossing process,

the pattern of microneedle array chip is obtained. The

ablation is carried out by UV excimer laser for drug reser-

voir. With the same technique, step 1 and step 2 lumens are

drilled in the microneedle array. Finally, the horizontal

side-open lumens are drilled with UV excimer laser.

2.4 Theoretical analysis

The design of the side-open microneedle with sharp

tip considered for strength modeling and CFD analysis

is shown in Fig. 6. The inlet is defined as in Sect. 1.

Sections 2 and 3 are the outlet sections. d is the lumen

diameter and D represents the outer diameter. L is the

length of the microneedle; L1 represents the length of the

lumen, X and Y represent the lengths of the lower cylin-

drical section and the upper conical section, respectively.

P1, P2 and P3 are the pressure values at each section and

q is the flow rate in the lumen.

2.5 Microneedle mechanics

During skin insertion, possible failure of the microneedle

may occur due to bending or buckling. An axial force acts

on the microneedle tip during insertion. This axial force is

compressive and causes buckling of the microneedle. The

axial force can be reduced by decreasing the tip area of

the microneedle. As buckling is directly related with

compressive force, which acts during insertion, sharp

microneedle tip reduces buckling and skin insertion

Fig. 4 Process flow of hot embossing process for microfabrication

Fig. 3 a Complete fluid model for microneedle. b Microneedles fluid lumen design


123

becomes easy. The microneedle experiences resistive

force exerted by skin during insertion, hence in order to

pierce the microneedle into the skin, the applied axial

force has to be greater than the skin resistance force. On

the other hand, bending may occur due to uneven skin

surface or human error during skin insertion. Hence, it is

important to find the relation between microneedle

geometry and mechanical properties of the material for

proper microneedle design and to predict failure of the

microneedle. The axial force (compressive force), which

the microneedle can withstand without breaking is given

by:

FCompressive ¼ ryA ð1Þ

where ry is the yield stress of the material and A is the

cross-sectional area of the microneedle tip.

The buckling force acting on the hollow microneedle

during insertion is given by (Zahn et al. 2000; Gere and

Timoshenko 1997).

FBuckling ¼p2E I1 þ I2ð Þ

L2ð2Þ

where E is young’s modulus of the material, I1 (m4) is the

moment of inertia of the lower cylindrical section and I2

(m4) is for the upper conical section. L(m) is the length of

the microneedle.

Moment of inertia for the hollow cylindrical section in

Eq. 2 is I1 ¼ p64

D4 � d4ð Þ where D is the outer diameter and

d is the inner diameter of the hollow cylindrical section.

Moment of inertia for the conical section in Eq. 2 is

I2 ¼ Dy3

396(assuming solid cone) where y is the length of the

conical section.

Fig. 5 Proposed process flow for the fabrication of out-of-plane microneedle array


123

As the microneedle penetrates the human skin, it expe-

riences resistive forces exerted by the human skin. To

penetrate the human skin, the outside force or pressure

must be greater than the resistive skin force. The resistive

force offered by the skin before the skin is punctured is

given by the following equation:

Fresistance ¼ PpierceA ð3Þ

where Ppierce is the required pressure to pierce the

microneedle into the skin and A is the cross sectional area

of the tip of the microneedle. As the microneedle pene-

trates the skin, the resistive force falls drastically (Frick

et al. 2003). After the skin is pierced by the microneedle,

the only force that acts on the microneedle is the fric-

tional force due to contact of the tissue with the

microneedle.

The bending force, which the microneedle can withstand

without breaking is given by:

FBending ¼ry I1 þ I2ð Þ

cLð4Þ

where c = D/2 is the distance from the vertical axis to

the outer edge of the section (Aggarwal and Johnston

2004).

2.6 Microfluidic analysis

It is important to model fluid flow characteristics because

microneedles are used to inject drug solutions into the skin.

In hollow microneedles, the drug flows through the lumen,

thereby allowing increased drug transport. The pressure

drop required in the microneedle for fluid flow is dependent

on a number of factors, such as microneedle geometry,

fluid viscosity and microneedle density. Microneedles have

micron-sized dimensions; therefore, there is significant

resistance to flow. In designing microneedles for trans-

dermal drug delivery, measurements and predictions of

fluid dynamics are important, so that the microneedle are

small enough to avoid pain, yet sharp enough to penetrate

the skin easily and large enough to achieve the desired flow

rate.

To determine fluid flow through the microneedle,

Poiseuille’s law of fluid flow in a cylinder is considered:

q ¼ pd4 DPð Þ128l L1ð Þ

ð5Þ

where d is the inner diameter of the microneedle and is

constant throughout the lumen, q is the flow rate, DP is the

pressure drop across the microneedle lumen. l is the vis-

cosity of fluid for water at 25�C, and L1 is the length of the

lumen of the microneedle.

The Reynolds number indicates the type of flow and is

given as:

Re ¼ qdV

lð6Þ

where q is a density of the fluid and V is the fluid velocity

in the lumen. At the micron level, the flow becomes lam-

inar. The flow is considered to be laminar if Re is less than

2,100, otherwise the flow is considered to be turbulent.

The total flow rate through the microneedle can be

calculated by using the continuity equation:

q ¼ q2 þ q3 ð7Þ

where q is the total flow through the microneedle in vertical

section (Fig. 6). q2 and q3 are the flows through outlet 2

and outlet 3 of the microneedle.

The microneedle section shown in Fig. 6 is modeled by

using Modified Bernoulli equation (Janna 1998). By con-

sidering the minor losses and friction losses due to 90�bend in side-open lumen, the pressure loss by the fluid is

calculated as:

P1

qgþ V1

2gþ z1 ¼

P2

qgþ V2

2gþ z2

� �þ P3

qgþ V3

2gþ z2

� �

þ fL

dþ V2

2gþXKV2

2gð8Þ

where P1 is the inlet pressure, P2 and P3 are the outlet

pressures, V1 is the inlet velocity, V2 and V3 are the outlet

velocities, q is the density of water, z1 and z2 are the

distances from data (z1 = 0 and z2 = L1) and f is the

Fig. 6 Cross section of the proposed microneedle design


123

friction factor. Since the above section is symmetrical

about a vertical axis, the outlet pressure, velocity and head

(z1 and z2) remain the same. Hence Eq. 8 can be modified

as:

P1

qgþ V2

1

2gþ z1 ¼

Poutlet

qgþ V2

2gþ z2

� �þ fL

dþ V2

2g

þXKV2

2g: ð9Þ

The pressure drop can be calculated from Eq. 9:

DP ¼ l128qðL1Þ

pd4þ q

8q2

p2d4K1 þ K2ð Þ ð10Þ

where K1 and K2 are loss coefficient factors for square edge

inlet and T-joint outlet for branched flow. The values of K1

and K2 for the proposed design are taken as 0.5 and 0.69,

respectively (Janna 1998). The friction factor for laminar

flow is given as f ¼ 64Re (Janna 1998). The first term on the

right-hand side in Eq. 10 describes the pressure drop due to

viscous shear force (Batchelor 1967) and the second term

describes the inertial effect at the inlet and exit (Stoeber

and Liepmann 2000).

3 Numerical simulation

3.1 Structural analysis

The structural analysis of the single microneedle was

conducted using finite element software ANSYS. The

microneedle experiences a bending force, axial force and

many other forces during skin insertion. The single out-of-

plane microneedle was modeled with fixed base and free

tip end for bending and axial stress analysis. The structural

model was built using element SOLID 191, which is

mostly used to model polymeric materials. Linear isotropic

material properties of polyglicolide acid (PGA) were used

for analysis (Young’s modulus of 10 GPa, Poisson ratio of

0.2 and yield strength of 90 MPa were considered (Ratner

et al. 1996)).

In simulation study, bending stress analysis was per-

formed by applying transverse force from 0.5 to 1 mN

along the ?x-axis on the tip of the microneedle. The range

for transverse force was chosen based on previous study

(Shibata et al. 2007), where it has been reported that

applied transverse force would generate bending moment

due to incorrect axial insertion. In the present study, the

simulation result shows the effects of bending forces during

skin insertion on the microneedle tip section and lumen

section. Similar types of analysis have been performed

previously for different designs of microneedles fabricated

from different materials (Shibata et al. 2007; Paik et al.

2004). It was found in the numerical simulation shown in

Fig. 7 that the stress exceeds the yield strength of the

material at a transverse force greater than 0.73 mN at the

tip of the microneedle. The stresses generated along

the length of the microneedle due to applied bending force

at the tip are negligible.

The axial stress (compressive stress) analysis was also

performed. Due to skin resistance, the axial force domi-

nates during microneedle insertion. The pressure required

to pierce the human skin is given as 3.183 9 106 Pa

(Henry et al. 1998a, b). The required piercing pressure

decreases with decrease in the area of the tip. In the pro-

posed design, the tip is ultra-sharp and hence the pressure

required to pierce the human skin will be lesser than the

given piercing pressure. The pressure of 3.183 9 106 Pa

was applied at the tip of the microneedle. The simulation

result for axial stress analysis is shown in Fig. 8. The

maximum stress occurs at the tip, which is less than the

yield strength of the material, and hence the microneedle

can penetrate the skin without failure. The simulation result

also shows negligible tip deflection due to applied pressure

at the tip.

Fig. 7 Bending stress analysis

Fig. 8 Axial stress analysis


123

The numerical analysis in the present study predicts

that there is no effect on the lumen area as well as at the

bottom of the microneedle due to bending or axial load

at the tip. The analysis is helpful to predict that the

proposed design of microneedles is strong enough to

penetrate the skin without breaking. Generally, metallic

microneedles are strong enough to penetrate the skin.

However, polymeric materials for microneedles must be

selected to have adequate mechanical strength and the

present analysis is helpful in selecting suitable polymeric

material such as polyglicolide acid (PGA) for micro-

needle fabrication.

3.2 CFD analysis

CFD analysis was conducted to predict pressure, velocity

and flow rate of the fluid flowing through the microneedles.

The simulation is done by applying pressures from 10 to

100 kPa at the inlet of the microneedle array. Pressure

between 10 and 100 kPa is recommended due to various

micropumping device specifications available in the market

(Khumpuang et al. 2003) that can be easily integrated with

microneedle array for drug delivery applications. Water at

25�C was considered as an isothermal fluid domain. The

simulation result shows uniform distribution of pressure

and velocity across the microneedles. The values of flow

rate and pressure drop across the microneedle were com-

puted using simulation results and compared with theo-

retical results.

The pressure was applied at the inlet of the reservoir.

The outlet pressure was assumed to be 0 kPa, because

the previous study (Khumpuang et al. 2003) showed

negligible pressure at the outlet of the microneedle dur-

ing skin penetration. The simulation result shows that

pressure distribution is uniform along the microneedle

array and decreases along the length of the microneedle.

This uniform distribution of pressure throughout the

array is important to deliver the drug in the proper pro-

portion. With reference to the Poiseuille’s law (Eq. 5),

the pressure drop increases with increase in length and

decrease in diameter of the pipe. At the side-open lumen

area, the pressure variation is believed to be due to

variation in pressure loss coefficient and flow separation.

Figure 9a, b shows the simulation result for pressure

distribution in a microneedles array at 30 kPa of applied

inlet pressure.

For the applied pressure at the inlet, the velocity dis-

tribution along an array was found from simulation result,

which is uniform throughout the array. Figure 10a, b

shows the uniform distribution of the velocity of fluid

flow through microneedles and variation in velocity along

the lumen. The velocity of the fluid is dependent on the

area of the lumen, viscosity, pressure and density of the

fluid. The velocity increases with decrease in the area. In

addition, the losses due to shear force, friction force

between lumen wall and fluid, and cross-sectional varia-

tion in the lumen are responsible for the velocity varia-

tion. The velocity at the wall is lesser than that at the

center region, due to friction between the wall and the

fluid. The velocity in the side-open lumen area is not

uniform due to flow separation and variation in loss

coefficient.

Fig. 9 Pressure distribution using CFD analysis. a Uniform pressure along the microneedle array. b Pressure distribution along single

microneedle


123

4 Results and discussion

4.1 Stress analysis

The first set of simulation results explains the mechanical

strength of the microneedle. According to the theoretical

calculation, maximum compressive force experienced by

the microneedle before microneedle failure is 0.282 mN.

The calculated skin resistive force was found to be 10 lN

and the calculated maximum bending force was 0.1732 N.

After applying the theoretically calculated bending force of

0.17 N at the tip of the microneedle, the FEM simulation

showed that the maximum stress occurred at a very small

portion of the tip, which exceeded the yield strength of the

material. The average stress at the bottom of the micro-

needle section was 72.4 MPa, which was below the yield

strength of the material. Figures 11 and 12 show the

maximum stress for the applied bending load at the tip and

the bottom of the microneedle, respectively.

Due to the conical shape, the stress concentration occurs

at the smallest area of the cone, i.e., at the tip. The simu-

lation is done in the two stages for bending stresses. In the

first set of simulation, the bending force reported in ref-

erence (Shibata et al. 2007) was applied at the tip of the

microneedle, and in the second set of simulation, theoret-

ically calculated bending force was applied at the tip of the

microneedle. Table 4 summarizes the results for these two

sets of simulations.

The deflection of the microneedle along its length due to

bending forces, shown in Table 4, was found by using

numerical analysis. Both the results are compared in

Fig. 13. The deflection at the cylindrical section and con-

ical section gradually increases with increase in distance

from the root of the microneedle. As the bending force is

applied directly at the microneedle tip, maximum deflec-

tion occurs at the tip of the microneedle. The main focus of

the study is also to determine the overall microneedle

structure strength and to predict the deflection on the rest of

the microneedle section, other than the conical section.

Except at the tip section, deflection in the entire micro-

needle structure is very less, compared with the other

microneedle designs presented by other researchers (Lee

et al. 2008; Paik et al. 2004; Park et al. 2005; Shibata et al.

2007). This proposed conical section design gives enough

Fig. 11 Simulation result for applied bending load at the theoreti-

cally calculated load

Fig. 12 Stress distribution at the bottom of the microneedle

Fig. 10 Uniform velocity across the plane. a 2D view of microneedle array for uniform velocity distribution. b Velocity distribution of fluid

through a single microneedle


123

strength to the microneedle to withstand bending and axial

forces with less deflection, which avoids fracture at the

bottom of the microneedle.

The variation in stress and deflection of the microneedle

due to applied bending force is in the range of 0.53–1 mN,

as reported by Shibata et al. (2007) is shown in Fig. 14.

The stress and deflection increases linearly with applied

bending force. The bending force is applied in the range of

0.5–1 mN for failure analysis of the microneedle. The

simulation result shows that the failure of the microneedle

occurs at the tip (Fig. 8), when the applied force exceeds

0.73 mN. The maximum stress of 90 MPa, which is the

yield strength of polyglycolic acid, is found at the tip of the

microneedle at an applied force of 0.73 mN. The maximum

deflection without failure is 0.14 lm, as shown in Fig. 14.

During microneedle insertion, the skin offers resistance

to the microneedle due to its elastic property. The micro-

needle should puncture the human skin without failure and

with less deflection. Theoretical calculation of the designed

microneedle shows that 10 lN of axial force is required to

puncture the skin. For the theoretical axial force, the sim-

ulation result shows that 37.6 MPa of maximum stress

occurs at the tip with negligible deflection. The resultant

stress is less than the yield stress of the material and hence,

the microneedle is safe at the applied theoretical axial

force. The microneedle can puncture the human skin with

applied skin piercing pressure of 3.18 MPa (Wilke et al.

2005a, b; Wang et al. 2006) at the tip of the microneedle

without failure. For the applied skin piercing pressure, the

maximum stress of 2.3 MPa with negligible deflection was

found at the tip, which is less than the yield stress of the

material. From the above results, it is predicted that the

proposed microneedle design can withstand the bending as

well as axial forces during skin insertion.

4.2 Flow analysis

The comparison between theoretically and numerically

calculated flow rates at the applied range of pressure

between 10 and 100 kPa is shown in Fig. 15. The numer-

ical results show that the pressure and velocity distribution

is uniform in the microneedle array (Figs. 9, 10). This

uniform distribution is important to deliver the drug in the

Table 4 The bending stress results at the tip and on the body of the microneedle

Applied force

at tip

Maximum

deflection (lm)

Bending

moment (Nm)

Failure occurrence Material

Bending force reported by

Shibata et al. (2007)

0.73 mN 0.14 1.46 e-7 [1.46 e-7

Nm. only at the tip

PGA

Theoretical bending force 0.17e3 mN 112

At tip

3.4 e-5 Microneedle fracture

at the tip

PGA

Fig. 14 Variation in stress and deflection of the microneedle with

applied force

Fig. 15 Comparison of theoretical and simulation result of flow rate

variation for the applied pressure

Fig. 13 Deflection of needle from root to the tip position for applied

loads of 0.17 N and 0.73 mN


123

proper proportion. The depicted results are for single

microneedle and not for an array. The flow rate by

numerical analysis is less compared to flow rate from

theoretical analysis, because theoretical results for flow rate

are calculated using Poiseuille’s law inside circular pipe

(Eq. 5) and do not consider any frictional losses, while

frictional losses were considered in the lumen for numer-

ical analysis by defining friction loss coefficient factors for

inlet (K1 = 0.5) and T-section outlet (K2 = 0.66) (Janna

1998).

The comparison between theoretically and numerically

calculated pressure drops with flow rate in the lumen section,

as shown in Fig. 16. The numerical and theoretical results

Fig. 17 Pressure variation from inlet to outlet of the microneedle array. a Pressure variation from the inlet to the step-2 lumen. b Pressure

variation from step 2 lumen to side-open holes (outlets)

Fig. 16 Theoretical and numerical results of pressure drop over the

flow rate


123

are depicted for an array and not for single microneedle.

The total pressure drop is the sum of pressure drop due to

viscous shear force of Poiseuille’s law and inertia effects

(Stoeber and Liepmann 2000; Batchelor 1967), which are

considered for theoretical analysis (Eq. 10). CFD analysis

has been validated by theoretical analysis showing that

the boundary conditions used in the analysis are reason-

ably accurate to predict the pressure drop for various flow

rates.

The pressure variation along the microneedle length is

shown in Fig. 17a and b. Initially, the pressure remains

constant throughout the reservoir. Pressure drop takes place

in step 1-lumen (Refer Fig. 3a), due to reduction in the area

along the length. Similarly, the pressure drops in step-2

lumen due to further reduction in the area along the length.

Due to the sudden impact of fluid at the T-junction, the

pressure increases suddenly at the end of the step-2 lumen

up to 25 kPa. Figure 17b shows the pressure variation in

the side-open zone. The pressure reduces rapidly after the

impact of fluid at the top mid-region of the T-section.

The small back flow occurs at the top mid-region of the

T-section in the microneedle and flow separation takes

place in the horizontal lumen; hence, the intensity of fluid

velocity decreases. Due to pressure difference, the fluid

flows out from the microneedle side-opened holes at low

velocity, which is sufficient for drug delivery applications.

The relationship between pressure and velocity of the

fluid flow along the vertical (steps 1 and 2) and horizontal

side-open lumen is shown in Fig. 18a, b. The pattern of

pressure and velocity variation is the same throughout the

microneedle array. Hence, the pressure–velocity relation-

ship was analyzed only in the individual lumen. At 100 kPa

Fig. 18 Pressure velocity relationship along the horizontal lumen area. a Pressure–velocity plot from the inlet to step-2 lumen. b Pressure–

velocity plot from step-2 lumen to side-open hole (outlet)


123

of applied inlet pressure, the result (Fig. 18a, b) shows that

the velocity increases from the inlet to the outlet region,

and the fluid pressure decreases with increase in distance

from the inlet to the outlet. There is velocity variation

across the lumen due to varying cross-sectional area. Ini-

tially at the inlet, the velocity is very small, but it increases

rapidly at step-1 lumen. Due to the same cross-sectional

area in step-1 lumen, the velocity becomes constant and

again increases rapidly due to decrease in thecross-sec-

tional area at step-2 lumen. Simulation result shows that

the fluid velocity along step-2 lumen is constant, and at T-

section shows small amount of backflow due to the sudden

impact with the lumen wall (Fig. 10). Hence the velocity of

the fluid becomes negligibly small and again increases

gradually due to pressure difference of 25 kPa (outlet

pressure = 0) (Fig. 18b). In the horizontal lumen, the

intensity of fluid velocity is low compared to step-2 lumen.

Only one side of the lumen is considered in Fig. 18b, due

to dimensional symmetry.

5 Conclusions

This paper first presents an in-depth review of design and

analysis of MEMS-based various categories of micronee-

dles, fabrication methods, performance parameters and their

medical applications where reported. Then, theoretical

design and analysis of new, hollow, out-of-plane polymeric

microneedles with cylindrical, side-open holes and inte-

grated reservoir is presented for TDD applications such as

treatment of cardiovascular and hemodynamic dysfunctions

and insulin delivery for diabetic patients. Suitable fabrica-

tion techniques such as micro-hot embossing process and

ultraviolet excimer laser drilling are proposed for side-open

polymeric hollow microneedles. Theoretical calculations

are done to validate the numerical analysis. The analysis

predicts that the performance of the microneedle in terms of

mechanical strength and flow rate has a crucial impact on

transdermal drug delivery applications. Theoretical and

simulation results are summarized below:

1. The maximum deflection occurs at the tip of the

microneedle due to the very small conical section. At

0.73 mN of force, the microneedle tip does not fail as

the maximum stress is below 90 MPa (yield strength

of PGA). Hence, there is no chance for the micronee-

dle to fail at the base, but for the theoretically

calculated load of 0.17 N, the buckling occurs at the

tip and the tip fails. However, the stress at the base of

the microneedle is below the yield strength of the

material. Thus, the designed polymeric microneedles

have less possibility of failure, as stresses generated

due to applied loads are less than the yield strength of

the material.

2. At a theoretically calculated axial force of 10 lN, the

simulation result shows that 37.6 MPa of maximum

stress occurs at the tip with negligible deflection. The

resultant stress is less than the yield stress of the

material and hence the microneedle design is safe at

the applied theoretical axial force.

3. The velocity decreases in a small region at the top of

the T-section due to the sudden impact, but fluid flows

out from the microneedle side-open holes, due to

pressure difference, with a sufficient flow rate that is

suitable for drug delivery application. Hence, it is

predicted that the effect of backflow of fluid from the

T-section will have no significant effect on flow rate

through the microneedles.

4. In the numerical solution, a maximum flow rate of

270 lL/min and a minimum flow rate of 32 lL/min

are achieved through a single microneedle at 100 and

10 kPa of inlet pressure, respectively. The range of

flow rates achieved from the analysis of the proposed

design of microneedle is suitable for drug delivery

application. The numerically and theoretically pre-

dicted values of flow rate are in close agreement,

thereby validating the CFD analysis results.

The proposed novel design of the polymeric micronee-

dle array, with side-open holes in the conical section and

integrated reservoir, is suitable to be integrated into various

categories of micropumping devices operating in the range

of 10–100 kPa of outlet pressure. Based on the extensive

literature review and design and analysis of microneedles

presented in this paper, the authors conclude that the

overall commercialization of MEMS microneedles in drug

delivery and biomedical application is still in its beginning.

A lot of technical information is available for a number of

microneedle concepts. However, many of the novel

microneedle designs reported in literature for drug delivery

and other biomedical applications still need to be incor-

porated into practical devices. Our future work will involve

fabrication of the proposed design of microneedle and

integration with other microfluidic devices, such as mi-

cropump, for integrated drug delivery applications. The

fabrication results of the polymeric microneedle array shall

be presented in a subsequent paper.

References

Aggarwal P, Johnston CR (2004) Geometrical effects in mechanical

characterizing of microneedle for biomedical applications. Sens

Actuators B 102:226–234


123

Ambrose CG, Clanto TO (2004) Bioabsorbable implants, review of

clinical experience in orthopedic surgery. Ann Biomed Eng

32(1):171–177

Aoyagi S et al (2007) Laser fabrication of high aspect ratio thin holes

on biodegradable polymer and its application to a microneedle.

Sens Actuators A 139:293–302

Aoyagi S et al (2008) Biodegradable polymer needle with various tip

angles and consideration on insertion mechanism of mosquito’s

proboscis. Sens Actuators A 143:20–28

Arora A et al (2008) Micro-scale devices for transdermal drug

delivery. Int J Pharm 364:227–236

Barry BW (2001) Novel mechanisms and devices to enable successful

transdermal drug delivery. Eur J Pharm Sci 14:101–114

Batchelor GK (1967) An introduction to fluid dynamics. University of

Cambridge, Cambridge

Becker H, Heim U (1999) Silicon as tool material for polymer hot

embossing. In: Proceedings of IEEE: MEMS’99, IEEE, Piscat-

away, NJ, USA, pp 228–231

Brown MB et al (2006) Dermal and transdermal drug delivery

systems: current and future prospects. Drug Deliv 13:175–187

Choi JW et al (2006) Insertion force estimation of various micronee-

dle array-type structures fabricated by a microstereolithography

apparatus. In: SICE-ICASE, International Joint Conference

Davis SP et al (2005) Hollow metal microneedles for insulin delivery

to diabetic rats. IEEE Trans Biomed Eng 52(5):909–915

Deodhar A et al (2005) An AI based controller for temperature

control for a Hot Embossing System. In: Proceedings of the 13th

Mediterranean Conference on Control and Automation Limas-

sol, Cyprus

Faulkner LL (2006) Microengineering, MEMS, and interfacing:

a practical guide

Frick TB et al (2003) Resistance forces acting on suture needles.

J Biomech 34:1335–1340

Gardeniers HJGE et al (2003) Silicon micromachined hollow

microneedles for transdermal liquid transport. J Microelectro-

mechan Syst 12(6)

Gere J, Timoshenko S (1997) Mechanics of materials, 4th edn. PWS

Publishing Co., Boston

Gerstel MS, Place VA (1976) Drug delivery device. US Patent No.

3,964,482

Griss et al (2002) Novel, side opened out-of-plane microneedles for

microfluidic transdermal interfacing. In: The fifteenth IEEE

international conference on micro electro mechanical systems,

pp 467–470

Griss P et al (2003) Side-opened out-of-plane microneedles for

microfluidics transdermal liquid transfer. J Microelectromech

Syst 12(3):296–301

Henry S et al (1998a) Micro machined needles for the transdermal

drug delivery of drugs. In: Proceedings of IEEE Workshop

MEMS, pp 494–498

Henry S et al (1998b) Micro-fabricated microneedles: a novel method

to increase transdermal drug delivery. J Pharm Sci 87:922–925

Jagger J et al (1988) Rates of needle-stick injury caused by various

devices in a university hospital. N Engl J Med 319:284–288

Janna WS (1998) Design of fluid thermal system, 2nd edn. PWS

Publications, Boston

Karande P et al (2004) Discovery of transdermal penetration

enhancers by high-throughput screening. Nat Biotechnol

22:192–197

Khumpuang S et al (2003) Design and fabrication of coupled

microneedle array and insertion guide array for safe penetration

through skin. In: International symposium of micromechatronics

and human science

Kuo Shyh-chi, Yukon Chou (2004) A novel polymer microneedle

arrays and PDMS micro-molding technique. Takang J Sci Eng

7(2):95–98

Lee JW et al (2008) Dissolving microneedles for transdermal drug

delivery. Biomaterials 29:2113–2124

Ma Bin et al (2006) A PZT insulin pump integrated with silicon

microneedle array for transdermal drug delivery. Microfluid

Nanofluid 2:417–423

Matteucci M et al (2008) A compact and disposable transdermal drug

delivery system. Microelectron Eng 85:1066–1073

McAllister DV et al (2000) Microfabricated microneedles for gene

and drug delivery. Annu Rev Biomed Eng 2:289–313

McGrew R, McGrew M (1985) Encyclopedia of medical history.

McGraw Hill, New York

Mukherjee EV et al (2004) Microneedle array for transdermal

biological fluid extraction and in situ analysis. Sens Actuators A

114:267–275

Paik SJ et al (2004) In-plane single-crystal-silicon microneedles for

minimally invasive microfluidic systems. Sens Actuators A

114:276–284

Park JH et al (2005) Biodegradable polymer microneedles: fabrica-

tion, mechanics and transdermal drug delivery. J control release

104:51–66

Prausnitz MR (2004) Microneedles for transdermal drug delivery.

Adv Drug Deliv Rev 56:581–587

Prausnitz MR et al (2000) Micro-fabricated microneedles for

transdermal drug delivery. In: Rathbone MJ, Hadgraft J, Roberts

MS (eds) Book chapter 42, Modified-release drug delivery

technology, Published by Informa Health Care, 2003. ISBN

0824708695, 9780824708696

Ratner BD et al (1996) Biomaterial science: an introduction to

materials in medicine. Academic, New York

Roxhed N (2007) A fully integrated microneedle-based transdermal

drug delivery system. A PhD thesis, Microsystems Technology

Laboratory, Royal Institute of Technology, Sweden

Roxhed N et al (2007) Membrane-sealed hollow microneedles and

related administration schemes for transdermal drug delivery.

Biomed Microdevices 10:271–279

Runyam WR, Bean KE (1990) Semiconductor integrated circuit

processing technology. Addison-Wisley, New York

Schreier H, Bouwstra J (1994) Liposomes and niosomes as topical

drug carriers: dermal and transdermal drug delivery. J Control

Release 30:1–15

Schuetz YB et al (2005) Emerging strategies for the transdermal

delivery of peptide and protein drugs. Expert Opin Drug Deliv

2:533–548

Shibata T et al (2007) Fabrication and mechanical characterization of

microneedle array for cell surgery. In: Actuators and Microsys-

tems Conference, pp 719–722

Stoeber B, Liepmann D (2000) Fluid injection through out-of-plane

microneedles. In: Microtechnologies in Medicine and Biology,

1st Annual International, Conference, Berkeley, CA 94704, USA

Stoeber B, Liepmann D (2002) Design, fabrication and testing of a

MEMS syringe. In: Proceedings of Solid state sensor and

actuator workshop

Varadan VK (2006) Smart material systems and MEMS: design and

development methodologies. Wiley, London

Wang X et al (2006) A novel fabrication approach for microneedles

using silicon micromachining technology. In: 1st IEEE Interna-

tional Conference on NEMS, pp 545–549

Wilke N et al (2005a) Silicon microneedle electrode array with

temperature monitoring for electroporation. Sens Actuators A,

123–124:319–325

Wilke N et al (2005b) Process optimization and characterization of

silicon microneedles fabricated by wet etch technology. Micro-

electron J 36:650–656

Yang H, Pan CT (2003) Analogous micro-optical components

fabricated using excimer laser ablation. Amkang J Sci Eng

6(3):145–150


123

Yang Ruoting et al (2006) Dynamic modeling and control of a micro-

needle integrated piezoelectric micropump for diabetes care. In:

Sixth IEEE Conference on nanotechnology, pp 146–149

Yu H (2005) Fabrication of a Hollow metallic microneedle array

using scanning laser direct writing. In: 9th international confer-

ence on miniaturized systems for chemistry and life sciences,

Boston, MA, USA

Zachary JH, Nicholas AP (2005) Microfabricated drug delivery

devices. Int J Pharm 306(1–2):15–23

Zahn JD et al (2000) Microfabricated polysilicon microneedles for

minimally invasive biomedical devices. Biomed Microdevices

2:295–303

Zahn JD et al (2004) Continuous on-chip micropumping for micronee-

dle-enhanced drug delivery. Biomed Microdevices 6(3):183–190


123

Structural and microfluidic analysis of hollow side-open polymeric microneedles for transdermal drug delivery applications

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