Note : Les descriptions sont présentées dans la langue officielle dans laquelle elles ont été soumises.
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SURFACE MICROMACHINED MICRONEEDLES
BACKGROUND OF THE INVENTION
1. The Field of the Invention
S The present invention relates generally to microneedles for the injection
and
extraction of fluids. More specifically. the present invention relates to an
easily
fabricated, micromachined array of microneedles or a single microneedle which
can be
easily attached to a syringe and which are mechanically durable.
2. The Relevant Technolo~v
Micro instrumentation is a rapidly growing area of interest for a broad
spectrum
of engineering applications. One particularly fast growing area is biomedical
instrumentation where significant efforts are being made to develop micro
biochemical
analysis systems, physiological systems, and drug delivery systems. A variety
of
manufacturing technologies are used to fabricate these micro systems, many of
which are
categorized under the set of technologies known as micromachining. The number
of
biomedical applications for micromachining technologies is rapidly growing.
Since
micromachining technologies are relatively new, there is an increasing set of
manufacturing techniques and critical applications still to be addressed.
In many areas of biotechnology and medicine, there exists the need for fluid
injection on a microscale; either for injection into a precise location, or
for injecting small
amounts of fluid. It is advantageous to be able to perform an injection with a
minimal
amount of tissue damage, and also with a minimum amount of discomfort and pain
to
patients. Microneedles and microneedle arrays are capable of performing these
tasks.
Some of the smallest hollow needles that are currently available have inner
diameters of
over 200~m. Prior microsized (sizes on the order of microns, where 1 micron =
1 pm =
l0~bm) needles have been made, as in U.S. Patent No. 5,457,041 to Ginaven et
al., and
U.S. Patent No. 5,591,139 to Lin et al.
For some applications, it is desirable to inject small amounts of fluid;
however,
in other situations, larger amounts of fluid are required to be injected. Most
of the prior
systems do not have the capability to transmit large amounts of fluid into a
precise
location. One of the methods used to address this problem is to fabricate an
array of
needles, as in U.S. Patent No. 5,457,041 referred to hereinabove. The patent
to Ginaven
specifies an array of microneedles of about 20 needles by 20 needles, wherein
the length
of the needles is between 10 and 25 microns, and the spacing between needles
is between
about 5 and 20 microns.
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Problems with prior microneedles include relatively poor mechanical
durability.
This is mainly due to the fact that such microneedles have been made out of
etched
silicon or out of chemical vapor deposited polysilicon, both of which have a
tendency to
be brittle and break easily.
It would therefore be of substantial interest to develop a durable device
which is
capable of injecting precise quantities of fluids into specific locations with
a minimal
amount of tissue damage, and which overcomes the difficulties associated with
prior
microneedle devices.
SUMMARY AND OBJECTS OF THE INVENTION
It is a primary object of the invention to provide an array of microneedles
which
is capable of transmitting relatively large quantities of fluids with minimal
tissue damage
and which can be readily attached to a standard syringe.
A further object of the invention is to provide an array of microneedles or a
single
microneedle which can be easily and economically fabricated according to
standard
micromachining procedures.
It is yet another object of the invention to provide an array of microneedles
or a
single microneedle which has a high degree of mechanical durability.
To achieve the forgoing objects and in accordance with the invention as
embodied
and broadly described herein, surface micromachined microneedles are formed as
single
needles or in two- or three-dimensional microneedle arrays. The microneedles
are
fabricated on a substrate which can remain attached to the microneedles or can
be
subsequently removed. The two- or three-dimensional microneedle arrays can
have
cross-coupling flow channels which allow for pressure equalization, and
balance of fluid
flow within the microneedle arrays. Additionally, a plurality of mechanical
support
members can be integrated into the arrays for stability and to control the
penetration
depth of the microneedles.
In one embodiment of the invention, a microneedle array device includes a
substrate having a substantially planar surface, and a plurality of hollow non-
silicon
microneedles on the planar surface of the substrate. Each of the microneedles
has a
microchannel therethrough that provides communication between at least one
input port
at a proximal end of each of the microneedles and at least one output port at
an opposite
distal end that extends beyond an edge of the substrate.
A method of fabricating a microneedle device according to the present
invention
includes providing a substrate with a substantially planar surface, and
depositing a metal
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material on the planar surface to form one or more bottom walls for one or
more
microneedles. A top surface of the bottom walls is coated with a photoresist
layer to a
height corresponding to a selected inner height of a microchannel for each of
the
microneedles. A metal material is then deposited to form side walls and a top
wall upon
the bottom walls and around the photoresist layer. The photoresist layer is
then removed
from each microchannel to form the microneedles. The microneedles can be
released
from the substrate and used independent of the substrate, if desired.
The method of fabricating the microneedle device can include p' etch-stop
membrane technology, anisotropic etching of silicon in potassium hydroxide,
sacrificial
thick photoresist micromolding technology, and micro-electrodeposition
technology.
These and other objects and features of the present invention will become more
fully apparent from the following description, or may be learned by the
practice of the
invention as set forth hereinafter.
BRIEF DESC PTION OF THE DRAWINGS
In order to more fully understand the manner in which the above-recited and
other
advantages and objects of the invention are obtained, a more particular
description of the
invention briefly described above will be rendered by reference to specific
embodiments
thereof which are illustrated in the appended drawings. Understanding that
these
drawings depict only typical embodiments of the invention and are not
therefore to be
considered limiting of its scope, the invention will be described and
explained with
additional specificity and detail through the use of the accompanying drawings
in which:
Figure 1 is a schematic representation of microneedles in a two-dimensional
array
according to one embodiment of the present invention;
Figure 2A is a schematic representation of microneedles in a two-dimensional
array according to another embodiment of the present invention;
Figure 2B is a schematic representation of microneedles in a two-dimensional
array according to another embodiment of the present invention;
Figure 3 is a schematic representation of a single microneedle according to a
further embodiment of the present invention;
Figures 4A-4F depict schematically the fabrication process sequence for
forming
a microneedle array;
Figure 4G depicts an embodiment of a multilumen microneedle formed according
to the present invention;
Figures SA and SB depict alternative methods of assembling two-dimensional
needle arrays into three-dimensional needle array devices;
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Figures 6A-6C are vector magnitude plots showing the effects of fluid flow
rate
on needle coupling channels and needle channels;
Figure 7 is a graph of the flow rate as a function of pressure difference for
a
microneedle array of the invention; and
Figure 8 is a graph of the flow rate as a function of pressure for five
individual
microneedles of the invention.
DETAILED DESCRIPTION OF THE INVENTION
The present invention is directed to micromachined arrays of microneedles and
single microneedles which exhibit mechanical durability. The microneedles are
capable
of injecting precisely controlled amounts of fluid, and can be easily and
economically
fabricated according to standard micromachining procedures.
Referring to the drawings, wherein like structures are provided with like
reference
designations. the drawings only show the structures necessary to understand
the present
invention. Additional structures known in the art have not been included to
maintain the
clarity of the drawings.
A schematic depiction of a micromachined needle array 10 according to one
embodiment of the invention is shown in Figure 1. The needle array 10 is
formed in a
two-dimensional configuration on a substrate 12 having a substantially planar
upper
surface 14. The substrate 12 is preferably composed of a semiconductor
material such
as silicon, although other materials can be employed such as glass, metals,
ceramics.
plastics, and composites or combinations thereof.
A plurality of fluidly interconnected hollow microneedles 16 are formed on
upper
surface 14 of substrate 12. The microneedles 16 each have a bottom wall, two
opposing
side walls, and a top wall that define a microchannel therein. This provides
the
microchannel with a transverse cross-sectional profile that is substantially
rectangular.
Each bottom wall is formed partially on upper surface 14 of substrate 12. The
microneedles each have one or more input ports in an input shaft, and a
portion of each
microneedle 16 including the microchannels extends beyond an edge of upper
surface 14
of substrate 12 in a cantilevered section which terminates in a needle tip 18
with a
channel opening 20 therein. The microneedles 16 are preferably aligned
substantially
parallel to each other on substrate 12.
The microchannels in the microneedles 16 are preferably dimensioned to have a
width between sidewalls of less than about 100 Vim, and more preferably about
0 pm to
about SO p,m. When the width is zero between the sidewalls, the microneedles
16
effectively become one multilumen microneedle with a plurality of
microchannels. The
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height between the top and bottom walls of the microchannel is also preferably
less than
about 100 Vim, and more preferably about 2 pm to about 50 pm. The length of
each
rnicroneedle can be from about 0.05 ~m to about 5 mm, and the width of each
microneedle can be from about U.OS urn to about 1 mm. The center-to-center
spacing
5 between individual microneedles can be from about 50 ~m to about 200 pm. The
microneedles can also withstand flow rates of up to about 1.5 cc/sec.
The microneedle length extended from the substrate can be varied from less
than
about 50 pm (subcutaneous) to several millimeters for fluid
delivery/extraction. For
example, the distal end of each microneedle can extend beyond the edge of the
substrate
a distance from about 10 ~m to about 100 mm. The inner cross-sectional
dimensions of
the microchannels in individual needles can range from about 10 pm to about I
mm in
width and about 2 ~m to about 50 pm in height. Accordingly, the microchannel
in each
of the microneedles can have a cross-sectional area in the range from about 25
pm'- to
about 5000 pm'.
A needle coupling channel member 22 is also formed on upper surface 14 of
substrate 12 between microneedles 16. The coupling channel member 22 has a
bottom
wall, two opposing side walls, and a top wall that define a coupling
microchannel therein,
which provides for fluid communication between the microchannels of each
microneedle
16. The coupling channel member 22 also allows for pressure equalization and
balance
of fluid flow between microneedles 16.
A pair of structural support members 24 are formed on either side of coupling
channel member 22 on upper surface 14 of substrate 12. The structural support
members
24 mechanically interconnect microneedles 16 to provide rigidity and strength
to needle
array 10. The support members 24 also precisely control the penetration depth
of
microneedles 16.
The microneedles 16, coupling channel member 22, and support members 24 can
be formed from a variety of metal materials such as nickel, copper, gold,
palladium,
titanium, chromium, alloys or combinations thereof, and the like, as well as
other
materials such as plastics, ceramics, glass, carbon black, composites or
combinations
thereof, and the like. The microneedles can be in fluid communication with a
single fluid
input device or with multiple fluid input devices.
A micromachined needle array 30 according to another embodiment of the
invention is shown in Figure 2A. The needle array 30 has a two-dimensional
configuration with similar components as needle array 10, except that the
substrate has
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been removed from the array. Accordingly, needle array 30 includes a plurality
of
microneedles 32 with microchannels therein that are dimensioned as discussed
above for
needle array 10. The microneedles 32 each terminate at a needle tip 33 with a
channel
opening 34 therein. A needle coupling channel member 36 with a coupling
microcharmel
therein provides a fluidic interconnection between the microchannels of each
microneedle
32. A pair of structural support members 38 are formed on either side of
coupling
channel member 36 and interconnect with microneedles 32. One or more input
ports 37
and output ports 39 can be optionally formed in microneedles 32 to increase
fluid input
and output flow. The input ports 37 and output ports 39 can be formed in one
or more
of the bottom wall, side walls, or top wall of microneedles 32 by conventional
fabrication
processes such as etching.
A micromachined needle array 31 according to another embodiment of the
invention is shown in Figure 2B. The needle array 31 has a two-dimensional
configuration with similar components as needle array 30, except that needle
array 31 is
formed without a coupling channel member. Accordingly, needle array 30
includes a
plurality of microneedles 32 with microchannels therein that are dimensioned
as
discussed above for needle array 10. The microneedles 32 each terminate at a
needle tip
33 with a channel opening 34 therein. A pair of structural support members 38
interconnect with microneedles 32. One or more input ports 37 and output ports
39 can
be optionally formed in microneedles 32 to increase fluid input and output
flow.
Figure 3 is a schematic representation of a single hollow microneedle 40
according to a further embodiment of the present invention which can be
fabricated by
standard micromachining techniques. The microneedle 40 has a bottom wall, two
opposing side walls, and a top wall that defines a microchannel 42 therein.
The
microchannel 42 communicates with a flanged inlet end 44 for fluid input. The
flanged
inlet end 44 can include single or multiple fluid input ports and acts as a
structural
support for microneedle 40. The flanged inlet end 44 can control penetration
depth and
can be used to mechanically fix microneedle 40 to a surface that is
penetrated. The
microneedle 40 terminates at a needle tip 46 at an opposite end from inlet end
44. One
or more output ports 48 are formed in microneedle 40 between inlet end 44 and
needle
tip 46. The output ports 48 can be formed in one or more of the bottom wall,
side walls,
or top wall of microneedle 40 by conventional fabrication processes such as
etching. The
microneedle 40 can be formed from a variety of materials such as nickel,
copper, gold,
palladium, titanium, chromium, alloys thereof, and the like, as well as other
materials
such as plastics, ceramics, glass, carbon black, composites thereof, etc. The
microneedle
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40 can be on a substrate as shown for the array of Figure 1, or can be removed
from a
substrate as shown for the array of Figure 2A.
A method of fabricating a two-dimensional needle array according to the
invention is depicted schematically in Figures 4A-4F. As shown in Figure 4A, a
substrate
12 having a substantially planar surface 14 is provided, such as a silicon
wafer which is
polished on both sides. The wafer can have a thickness of about 1 pm to about
700 pm,
and is preferably about 150 pm thick. One side of the wafer is heavily doped
with boron
using high temperature thermal diffusion in order to form a 3-5 pm thick p+
layer. Next,
silicon nitride (Si3N4) is deposited on both sides of the wafer using plasma-
enhanced
chemical vapor deposition (PECVD). The silicon nitride on the undoped side of
the
wafer is patterned and etched employing photoresist as a mask, and then
isotropic etching
(a CF4 plasma for example) is used to etch the exposed silicon nitride to
define the area
upon which the microneedles are to be fabricated. After patterning the silicon
nitride
layer, the exposed silicon is anisotropicallv etched using a potassium
hydroxide (KOH)
solution. The p+ boron layer serves as an etch stop, resulting in a thin
sacrificial
membrane 52, as shown in Figure 4A. The sacrificial membrane 52 comprises the
surface upon which the microneedles are formed and then subsequently released
as
described below.
Next, a metal system of adhesion layers and an electroplating seed layer are
deposited (by electron beam evaporation, for example) onto the insulating
silicon nitride
film. The adhesion and seed layers are typically composed of, but not limited
to, titanium,
chromium, copper, or combinations thereof. Then, using a mask of the
appropriate
dimensions and standard photolithographic techniques, this metal multilayer is
patterned.
A metal material is then electroplated to form one or more bottom walls 54
(e.g., about
20 pm thick) for the microneedles, as shown in Figure 4B. Palladium is a
preferred metal
for the bottom wall since it provides high mechanical strength and durability,
is corrosion
resistant, provides biocompatibility for use in biomedical applications, and
is easily
electroplated to within precise dimensions. Other materials which fit the
criteria
mentioned could also be used equally well, such as copper, nickel, or gold.
When
performing the electroplating, the bath chemistry and the electroplating
conditions (such
as amount of applied current and time in the electroplating bath) should be
precisely
controlled for optimum results. The typical dimensions for bottom walls 54 are
about 10-
20 pm in thickness, and about 50 pm wide.
After bottom walls S4 have been formed, a commercially available thick
photoresist is deposited (e.g., about 20 pm thick) and patterned using
ultraviolet exposure
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and alkaline developer, resulting in sacrificial layers 56 as depicted in
Figure 4C. Next,
a metal seed layer such as gold is sputter deposited (e.g., about 800 ~ thick)
onto the
photoresist sacrificial layers 56. The metal seed layer serves as an
electrical contact for
the subsequent metal electroplating.
A metal layer is then electroplated (e.g., about 20 ~m thick) onto sacrificial
layers
56 to form the a plurality of side walls 5$ and top walls 60 of each
microneedle, as shown
in Figure 4D. The exposed metal seed layer is then removed using wet etching
techniques to expose the underlying photoresist. Once the metal seed layer has
been
etched. the wafer is placed in an acetone bath to remove the thick photoresist
from inside
the microneedle structures, thereby producing a plurality of hollow
microneedles 16, as
shown in Figure 4E.
In the final processing step, sacrificial membrane 52 is removed by an
isotropic
etching technique, such as reactive ion etching in a SF6 plasma. Thus, the
microneedle
ends are released from sacrificial membrane 52 and are freely suspended,
projecting
outward from substrate 12, as depicted in Figure 4F and in the embodiment of
Figure 1.
In addition, the needle arrays can be released from substrate 12 following
surface
fabrication on substrate 12, if desired, such as in the embodiment shown in
Figure 2A.
This method does not require sacrificial membrane formation or KOH etching.
Instead,
the needle arrays are released using wet etching of the seed metal from
underneath the
needle structures. If the seed metal is copper. for example, then this can be
done by a
selective etch of ammonium hydroxide saturated with cupric sulfate.
The support members and needle walls formed by micro-electroforming processes
provide increased structural integrity. In addition, the needle coupling
channels minimize
the effects of restricted needle passages by providing a redistribution point
for fluid flow
between the passages.
Arrays of two up to hundreds of microneedles can be easily and economically
fabricated in a package with dimensions on the order of millimeters according
to the
above procedure. A single microneedle such as shown in Figure 3 can also be
fabricated
according to the procedure outlined above. In addition, instead of having the
fluid outlet
at the tips of the microneedles, fluid outlet ports can be etched into the
side walls, top
walls, and/or bottom walls of the microneedle(s) if a larger amount of fluid
transfer is
desired.
Alternatively, the fabricating method outlined above can be used to form a
multilumen microneedle by forming bottom walls 54 with zero spacing
therebetween on
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substrate 12 and carrying out the remaining steps as described above. A
resulting
multilumen microneedle 26 with a plurality of microchannels 2$ is shown in
Figure 4G.
All of the microneedles and arrays of the present invention can be coated on
the
inside with biocompatible materials, such as silicon nitride, gold, plastics,
etc., by
conventional coating processes known to those skilled in the art.
Figures SA and SB depict alternative methods of assembling two-dimensional
needle
arrays into three-dimensional needle array devices. In the method depicted in
Figure SA,
a plurality of two-dimensional needle arrays 70 are provided which have been
released from
a substrate as shown for the array of Figure 2A. The needle arrays 70 are
positioned in a
stacked configuration with a plurality metallic spacers 72 therebetween to
define the distance
between any two microneedle arrays in the stack. The stacked needle array
configuration is
then subjected to flash electroplating to join needle arrays 70 with metallic
spacers 72 in a
fixed three-dimensional needle array device 74. The needle array device 74 can
then be
disposed in a machined interface structure 76 such as an acrylic interface.
allowing
I S connection to a dispensing means for injecting a liquid such as a syringe.
In the method depicted in Figure SB, a plurality of two-dimensional needle
arrays 80
are provided on substrates 82 such as shown for the array of Figure 1. The
needle arrays 80
are positioned in a stacked configuration, with substrates 82 acting as
spacers between the
arrays 80, to define the distance between any two microneedle arrays in the
stack. The
stacked needle array configuration is placed in a mold 84 such as an aluminum
mold for
plastic injection molding. The stacked needle array configuration is then
subjected to plastic
injection molding. This bonds needle arrays 80 together with a plastic molding
material 86,
thereby forming a fixed three-dimensional needle array device 88. The needle
array device
88 can then be disposed in an interface structure 76 allowing connection to a
dispensing
device such as a syringe.
In another alternative method of assembling two-dimensional needle arrays into
three-dimensional needle array devices, the two dimensional arrays are
manually assembled
under a microscope. The two-dimensional arrays are stacked with spacers or
with substrates
on the arrays and are bonded together with a polymeric adhesive such as a UV-
curable
adhesive to form a three-dimensional needle array device, which can then be
disposed in an
interface structure.
The fabricated three-dimensional needle array devices are typically
dimensioned to
have a length of about S mm, a width of about 5 mm, and a height of about 2
mm.
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The interface structures for connection to a syringe can be made from a
variety of
plastic materials such as acrylics, polystyrene, polyethylene, polypropylene,
and the like. The
interface structure typically accommodates a three-dimensional needle array
device having
up to about 25 two dimensional arrays. The interface structures are bonded to
the three-
s dimensional needle array devices with a polymeric adhesive such as a UV-
curable adhesive.
The interface structures are configured to accept direct syringe connection
via a connection
means such as a conventional Luer-lock connector. Alternatively. interface
structures can
be formed for a single two-dimensional array or a single microneedle so as to
accept direct
syringe connection via a connector such as a Luer-lock connector.
10 The micromachined microneedles of the invention have many benefits and
advantages. These include reduced trauma at penetration sites due to their
small size, greater
freedom of patient movement because of the minimal penetration depth of the
needles, a
practically pain-free drug delivery due to the smaller cross section of the
needle tip and
distribution of fluid force, and precise control of penetration depth from
needle extension
length. In addition, the microneedles have the ability to deliver drugs to
localized areas, and
are advantageous in their ability to be stacked and packaged into a three-
dimensional device
for fluid transfer.
The micromachined microneedles can be used in a wide variety of biomedical
applications. The microneedles can withstand typical handling and can
subcutaneously
deliver medication without the usual discomfort associated with conventional
needles. The
microneedles are minimally invasive, in that the microneedles only penetrate
just beyond the
viable epidermis, reaching the capillaries and minimizing the chance of
encountering and
damaging the nerves present in the area of penetration.
The following examples are given to illustrate the present invention, and are
not
intended to limit the scope of the invention.
Ex m 1
A two dimensional microneedle array was fabricated according to the procedure
outlined above to have to have 25 microneedles. The inner dimensions of the
hollow
microneedles were made to be of inner cross-sectional area of 40 x 20 ~.m'-
(width by height]
and outer cross-sectional area of 80 x 60 Vim'-. The needle coupling channels
were 100 pm
wide, and provided fluid communication between each needle channel. Two sets
of 60 x 100
pm'- structural supports were located 250 pm from each needle end. Each needle
channel
was 2 mm long, while the width of the 25 needle array was 5.2 mm. The center-
to-center
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spacing between individual needles was 200 pm. The needle wall thickness was
approximately 20 p.m of electroformed metal.
Needle arrays were fabricated from electroformed low stress nickel sulphamate,
gold
cyanide, and palladium electroforming solutions. The bath chemistry and
electroplating
conditions were selected and precisely controlled to allow formation of low
stress
depositions on top of a 3-5 pm sacrificial membrane. The surface roughness of
the
electroplated metals was found by Atomic Force Microscope (AFM) to be
approximately 15
nm, resulting in a relative roughness of 0.00056.
It is important to note the structural quality of the needle tips. The inner
dimensions
were approximately 30 x 20 p.m'-, outer dimensions were approximately 80 x 60
pm'-, and the
needle tips were formed with 45° angles for ease of penetration.
Example 2
Analytical and computer modelling were performed in order to assess how the
cross
flow design feature unique to this invention can be used to equalize the
pressure distribution
across the needle channels, to generally determine the relationships between
input pressure
and fluid delivery rates for the array, and to investigate the
interrelationships amongst the
physical dimensions of the needle array. The purpose of the model is (a) to
determine the
relationship between the input pressure and the fluid delivery rates for the
microneedle array
and (b) to investigate the relationships between the physical dimensions of
the microneedle
array (cross-sectional areas of the primary and coupling channels) and the
effect on restricted
flow within the microneedle array. Pressure equalization may become necessary
in the event
that some of the individual needle channels become obstructed. The model
described herein
can be used to characterize the fluid flow rates throughout the micromachined
needle array
and to determine the most effective design dimensions.
The general equation that governs the motion of a viscous compressible fluid
in
Cartesian coordinates can be expressed as:
,u~2V - ~p + pf = pV ( 1
where a superimposed dot indicates a material derivative, V is the fluid
velocity vector, f
represents the external forces (such as gravity, for example) ,u is the fluid
viscosity, and ~p
is the pressure gradient required to move the fluid. The equation is known as
the Navier-
Stokes equation. For the case of a Newtonian fluid with uniform physical
properties (water,
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for example) and negligible effects from external forces moving through a
rectangular
channel, the flow is two-dimensional and can be simplified to:
aZv- aZv i
-.~__op=o
aX 2
where x is the direction along the microchannel width and y is the direction
along the
microchannel height.
These equations were solved numerically using commercially developed ANSYS
software, which is based on the finite element method (FEM). The elements used
in this
numerical model are solved for flow distributions within a region, as opposed
to elements
that model a network of one dimensional regions hooked together. A segregated
sequential
solver algorithm is used; that is, the matrix system derived from the finite
element
discretization of the governing equation for each degree of freedom is solved
separately. The
degrees of freedom in this case are velocity, pressure. and temperature.
The flow problem is inherently nonlinear, and the governing equations are
coupled
together. The sequential solution of all the governing equations. combined
with the update
of any pressure dependent properties, constitutes a global iteration. The
number of global
iterations required to achieve a converged solution in this analysis was 15,
implying that the
model is stable.
The numerical model is solved for an array of 10 microneedles. Each needle
channel
is 40 ~m in width and 2 mm long, while the needle coupling channels which
interconnect
each needle channel are 100 pm in width. The vectors that run along the center
of each
needle channel depict the highest fluid flow rate ( 1.459 cc/sec), while the
small vectors in
the middle of each needle coupling channel depict a zero fluid flow rate. The
pressure drop
across the length of each individual needle was 16.2 kPa. The model was fairly
successful
at illustrating the effect that the needle coupling channels have on the rest
of the system. The
model indicates that an equilibrium in the fluid interaction within the needle
coupling
channels is obtained and that the fluid flow within the needle coupling
channels appears not
to interfere with the fluid flow in the needle channels.
The effect that the needle coupling channels have on the rest of the system is
shown
in the vector magnitude plot of Figure 6A. This plot shows the magnitude of
the fluid rate
components, with the vectors in the middle of each needle channel 92
representing the
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13
greatest magnitude (1.471 cc/sec), while the small vectors in the center of
each needle
coupling channel 94 represent a vector magnitude of zero. The fluid flow rates
remain static
within the needle coupling channels when there is no opposition (i. e., clog)
to the flow within
the needle channels. The input needle channels (indicated on the left in
Figure 6A) have the
same fluid flow rate as the output needle channels (indicated on the right of
Figure 6A).
The effects on flow characteristics when a needle becomes clogged can be seen
in
Figure 6B. Figure 6B is a vector plot of the flow in needle channels 92 when
one input is
obstructed. The model indicates that the flow in the channel adjacent to the
clogged needle
is augmented by the incoming fluid flow from other needle channels by means of
the needle
coupling channels. Flow within the needle coupling channel 94 is as high as
0.912 cc/sec.
The overall flow rate through the array remains constant due to conservation
of mass.
However, the flow rate within the individual output needle channels is
decreased to
compensate for the clogged passage. The flow rate in the input needle channels
is I .459
ce/sec, while the flow rate in the output channels is 1.276 cc/sec. This
indicates a 14.3
1 S percent decrease in the flow rate in individual needle channels. In
addition, a developing
flow region is present at the entrance of the output needle channel. extending
approximately
100 pm from the needle coupling channel. While this figure shows the
interaction between
only three needle channels, the needle coupling channels redistribute fluid
throughout the
entire array. It was found that all of the fluid flow rates were identical in
all of the output
needle channels with an accuracy of 0.0001 cc/sec.
In the event that an obstruction is encountered in one of the output needle
channels,
the result is similar, with the exception that fluid flow from the
unobstructed input needle
channels would then be redistributed to the unobstructed output needle
channels. Figure 6C
shows a vector plot illustrating the effect that the needle coupling channel
94 has on the
needle channels 92 when an output is obstructed. The flow in the needle
channels adjacent
to the clog act to augment the outgoing fluid flow to the other needle
channels by means of
the needle coupling channels. Flow within the needle coupling channels is as
high as 0.831
cc/sec. The overall flow rate through the needle array again remains constant
due to
conservation of mass. However, the flow rate within the individual output
needles is
increased to compensate for the clogged passage. The flow rate in the input
needle channels
is 1.246 cc/sec, while the flow rate in the output needle channels is 1.453
cc/sec. This
indicates a 16.6 % increase in the flow rate in individual needle channels. In
addition, a
developing flow region is present at the entrance of the output needle
channel. extending
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14
approximately 100 ~m from the needle coupling channel. The flow rates in the
output needle
channels are the same to within an accuracy of 0.0001 cc/sec.
The ANSYS model successfully characterized the flow in the microchannels using
the specified dimensions. In Figure 6A, the flow within the needle coupling
channels
remains static, while in Figures 6B and 6C, the flow within the needle
coupling channels
dynamically tends towards the needle channels of less resistance. Fluid flow
rates within the
needle coupling channels adjacent to the obstruction showed an increase of
90%, while the
fluid flow rates at the far end of the array showed an increase of less than
10%. The fluid
flow rates in individual needle channels show an increase or decrease
depending on the
location of the obstruction. The needle coupling channels, therefore, act to
redistribute the
fluid flow as necessary when needle inputs or outputs are clogged.
Exam le
Flows in channels of similar dimensions were compared experimentally. Figure 7
represents data for flow of water in 3000 x 600 x 30 Vim' (length x width x
height). The
microchannels in this example were fabricated according to the procedures
described
previously. The plot of the data in Figure 7 serves to show the agreement
between the data
and the predictions of the Navier-Stokes (N-S) theory and the model based on
the micropolar
fluid theory.
Example 4
Five packaged microneedles were fabricated according to the procedures
described
hereinabove. Fluid flow rate tests were conducted to determine the rate of
flow through
each packaged microneedle as a function of pressure. Each microneedle was
attached to a
standard 5 cc syringe filled with water and affixed vertically in a test
stand. The test stand
was positioned in proximity to an Instron (model 4400) load frame so that the
calibrated load
cell would come in contact with the syringe plunger. The load frame has the
capability of
maintaining a constant force on the syringe plunger over specified periods of
time. Prior to
each test, a sealed flask was placed on a scale and zeroed. The flask was
filled and kept at
a constant level to keep the surface area of the water nearly constant. This
method helps to
maintain a constant rate of evaporation, which was also monitored during the
tests. The load
frame was set to apply a constant force over a period of thirty minutes, while
the sealed flask
was used to capture the water dispensed by the syringe. After the thirty
minute period, the
load was removed and the flask was placed on a scale to measure the resultant
amount of
water that had accumulated. The scale (Sartorius model 1602 MP8-1) was
enclosed in a
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WO 00/16833 PCT/US99/21509
glass case and has a resolution of 100 fig. The evaporation rate was checked
prior to each
test by placing the sealed flask on the scale and noting the weight loss over
a period of thirty
minutes. The value obtained was then added to the resultant data from the
needle tests.
The dimensions of each needle shaft were 200 ~m wide and 60 ~m thick, and the
tip
5 dimensions were less than 1 S x 15 p.mz. The length of the tapered portion
was 1 mm and the
distance from the tip to the first output port was approximately 300 ~.m. The
total length of
the microneedles was 6 mm, with inner channel dimensions of 140 ~m wide and 20
~m high.
The wall thickness of each needle was approximately 20 ~m and the microneedle
output
ports were on the top and bottom with dimensions of 30 Vim'. The output ports
were
10 separated by 30 ~m and there were nine ports on the top and 12 ports on the
bottom.
The fluid flow experiments were performed on the five packaged microneedles at
pressures ranging from 1 to 70 psi. The fabricated needles were packaged into
an interface
using UV-curable epoxy. Pressures were applied to each microneedle at 30
minute intervals
while the weight of the water that had accumulated during each test was
recorded. The
I 5 weight data was converted into fluid flow rates for the needles based on
the density of water
at 25°C and the results are shown in the graph of Figure 8. The
microneedles demonstrated
flow rates in the range of 0.00384 to 2.67 ~L/min with applied pressures of 1
to 70 psi. The
packaged microneedles also exhibited the ability to withstand pressures
exceeding 100 psi.
The theoretical data were obtained by repeating the previously described
models for
the same input pressures that were used in the fluid flow tests. The
theoretical data are also
plotted in Figure 8; however, the resultant data do not accurately predict the
experimentally
obtained fluid flow rates. One possible explanation for the discrepancy in
flow rates is that
the presence of microscale surface effects, such as rotations of molecules,
variations of
viscosity, slip velocity, and capillary effects were not accounted for in the
theoretical
calculations.
The present invention may be embodied in other specific forms without
departing
from its spirit or essential characteristics. The described embodiments are to
be considered
in all respects only as illustrative and not restrictive. The scope of the
invention is, therefore,
indicated by the appended claims rather than by the foregoing description. All
changes
which come within the meaning and range of equivalency of the claims are to be
embraced
within their scope.
What is claimed is:
