Biomed Microdevices (2012) 14:483–496
DOI 10.1007/s10544-011-9625-4

An implantable MEMS micropump system for drug delivery
in small animals
Heidi Gensler & Roya Sheybani & Po-Ying Li &Ronalee Lo Mann & Ellis Meng

Published online: 25 January 2012
# Springer Science+Business Media, LLC 2012

Abstract We present the first implantable drug delivery
system for controlled timing and location of dosing in small
animals. Current implantable drug delivery devices do notprovide control over these factors nor are they feasible for
implantation in research animals as small as mice. Our
system utilizes an integrated electrolysis micropump, isrefillable, has an inert drug reservoir for broad drug compatibility, and is capable of adjustment to the delivery regimen while implanted. Electrochemical impedance
spectroscopy (EIS) was used for characterization of electrodes on glass substrate and a flexible Parylene substrate.
Benchtop testing of the electrolysis actuator resulted in flow
rates from 1 μL/min to 34 μL/min on glass substrate and up
to 6.8 μL/min on Parylene substrate. The fully integrated
system generated a flow rate of 4.72±0.35 μL/min underapplied constant current of 1.0 mA while maintaining a
power consumption of only ~3 mW. Finally, we demonstrated in vivo application of the system for anti-cancer drugdelivery in mice.
Keywords Drug delivery . Electrochemical bellows
actuator . Implants . Micropump

H. Gensler : R. Sheybani : R. L. Mann : E. Meng (*)
Department of Biomedical Engineering,
Viterbi School of Engineering, University of Southern California,
1042 Downey Way, DRB-140,
Los Angeles, CA 90089-1111, USAe-mail: ellis.meng@usc.edu
P.-Y. Li : E. Meng
Ming Hsieh Department of Electrical Engineering,
Viterbi School of Engineering, University of Southern California,
3651 Watt Way, VHE-602,
Los Angeles, CA 90089-0241, USA

1 Introduction
Laboratory animal research plays an essential role in the
study and understanding of human diseases. Rats and mice
are the most commonly used models and comprise an estimated 95% of laboratory animals used in research (Trull andRich 1999). The rat was the first mammal domesticated forscientific research, and historically has been preferred overthe mouse for physiological and behavioral studies due tothe convenience of the rat’s larger size and more complex
behavior (Benedikz et al. 2009; Petit-Zeman 2004). The
mouse, however, has been preferred by geneticists because
it is easier to genetically manipulate (Benedikz et al. 2009).
Advances in genetic techniques in the rat and improved
physiological sampling techniques in the mouse are enabling propagation of both models as a means to better
understanding of human disease (Worley et al. 2007). In
particular, both rodent models offer a relatively inexpensive
way to screen for and investigate the most appropriatedelivery of novel therapeutics.
It is well established in the drug delivery literature thatboth dosage and timing are critical to the study of drug
effects on the body (pharmacodynamic responses) (Urquhart
et al. 1984; Smolensky and Peppas 2007; Youan 2004), as
drug function is often tied to biological rhythms. Chronotherapeutics entails drug therapy that is tuned to biologicalrhythms which play a role in disease activity and thus
impact drug efficacy (Smolensky and Peppas 2007). In
treating deep vein thrombosis with the anticoagulant drug
heparin, the same dosage can either exceed or fall below thetherapeutic level depending on the time of administration(Smolensky and Peppas 2007). Cancer studies have shownthat administration of therapy at a particular time in a tumor
cell cycle can increase effectiveness while reducing toxicity
to normal cells. In the case of diabetes, control over both

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basal levels of insulin as well as carefully timed boluses at
mealtimes is required in order to mimic natural levels(Youan 2004).
Another important drug delivery parameter is location.
Localized delivery, or administration of drug directly to thedesired site of action (“site-specific”), is preferred in cases
where the drug is intended to act on a specific location andcould have serious side effects if delivered systemically (e.g.
oral and intravenous routes). Some forms of injection are
localized but are undesirable in applications where dosing
occurs at high frequency. In some tissues, such as the eye,frequent injection can lead to severe trauma (Ambati et al.
2000). In addition, frequent injection (and related handling)in small animals could induce stress effects that confoundstudy results (Urquhart et al. 1984; Kramer et al. 2001).
Localized delivery is not to be confused with “targeted” delivery, which generally refers to molecular systems and nanoparticles that accumulate preferentially at or only act on a
specific site, but are actually delivered systemically (Farokhzad
and Langer 2009). Some success has been found with thesetargeted systems, even at the clinical stage, but each drug andcarrier molecule combination requires significant laboratory
and clinical testing to ensure safety and efficacy. siRNAbased therapies for cancer treatment are one example of an
application supporting localized delivery. siRNA has a shorthalf-life in vivo (~sec to min). With chemical modification,
the half-life can be extended to minutes or days, but atthe cost of decreased potency (Dykxhoorn et al. 2006).
Having a means of delivering the drug directly to thesite of action without delays of systemic administration
and without need for chemical modification could provide greater efficacy. In the case of pain management,
administration directly to the afflicted site could reduce
systemic side effects and decrease the amount of drug
required to achieve reduction in pain level (Ruan 2007).
Although there is ample evidence of dose and temporaldependence of drug actions, there are limited means fordelivering drugs with control over both dose and timing.
The need for localized delivery places further constraints on
the method of administration. Conventional drug delivery
methods include tether infusion (animal is connected toinfusion pump via tether) and vascular access ports (subcutaneously implanted injection ports), but these require frequent handling of laboratory animals and constant care toprevent infection and other complications. As mentioned
previously, tethered and restraint-based systems induce
stress that may confound study results (Urquhart et al.
1984; Kramer et al. 2001). Implantable pumps offer analternative to tethered systems, injection ports, and frequent
animal handling. However, there are no commercially available implantable pumps suitable for periodic infusion of adrug in small animals (rodents) with control of dosage,timing, and delivery site in chronic experiments. The

Biomed Microdevices (2012) 14:483–496

requirements for this application include: refillability, small
size, adjustable flow rate, on/off operation, and low power
consumption. For subcutaneous implantation to be a minor
procedure, the surgery should involve less than 10% of theanimal’s surface area (Urquhart et al. 1984). The single-use
osmotically-driven Alzet® pumps (Alzet 2011) are of a form
factor suitable for use in mice but provide only continuous
infusion; the flow rate is predetermined by the water permeability of the pump’s semipermeable membrane and is fixedat the time of manufacture. Drug payload lasts from 1 day to
6 weeks and is therefore not adequate for extended chronic
studies lasting months. The non-refillable Med-e-Cell InfuDisk™ (Med-e-cell) is too large for practical implantation inmice (~25 to 30 g), as the smallest available reservoir is
5 mL and weighs ~12 g when empty. The actuation isachieved with an electrochemical cell module that provides
only continuous drug delivery at a factory set flow rate.
Even at the lowest flow rate, 0.03 ml/h, the maximum
delivery duration for the 5 ml reservoir pump is less than aweek, making this pump unsuitable for chronic studies
lasting longer than 7 days. The Primetech iPrecio® pumps(iPrecio 2011) are programmable prior to implantation, but
the regimen cannot be modified afterwards. In addition, thisbattery-powered pump is too large for use in mice, is singleuse only, and the flow rates are limited to a maximum of
30 μL/h.
The issues of controllable delivery in a small platform
can be addressed by employing microelectromechanicalsystems (MEMS) technology, which utilizes fabrication
techniques borrowed from the semiconductor industry to
produce miniaturized structures, sensors, actuators, and systems. Many MEMS pumps have been reported and arereviewed elsewhere (Nguyen et al. 2002; Laser and Santiago
2004; Woias 2005; Tsai and Sue 2007; Amirouche et al.
2009; Yun and Yoon 2006), but few of these devices feature
an integrated reservoir and show feasibility for implantationin small animals. MEMS-based micropumps are generally
divided into dynamic (non-mechanical) and displacement(mechanical) types. Non-mechanical micropumps typicallyare limited in flow rate (maximum of 10 μL/min) (Amirouche
et al. 2009), have relatively slow response compared tomechanical micropumps, and often require interaction witha working solution with particular electrical properties, suchas conductivity (e.g. electrokinetic pumps) (Laser andSantiago 2004; Tsai and Sue 2007; Amirouche et al.
2009). Examples of dynamic actuation include electroosmosis, electrowetting, electrophoresis, electrohydrodynamics, and magnetohydrodynamics.
Displacement actuation schemes offer distinct advantages for drug delivery applications in small animals, including operation independent of the fluid properties (suchas conductivity), higher flow rates, and faster response times(Amirouche et al. 2009). There are numerous displacement-

Biomed Microdevices (2012) 14:483–496

type actuation methods including electrostatic, piezoelectric,
thermopneumatic, shape-memory alloy (SMA), bimetallic,and electromagnetic, and ionic conductive polymer films(ICPF). Of these, piezoelectric pumps are the most common.
Smits (1989) described a three piezoelectric valve peristaltic
pump designed for insulin delivery. Maillefer et al. presented a piezoelectric ceramic disk-actuated micropump(1999; 2001) that includes a reservoir and battery intended
for transcutaneous delivery in humans and weighs over 20 g
(Debiotech 2011). Evans et al. (2011) reported an intrathecal
drug delivery system with lead zirconate titanate (PZT)
valves that consumes low power, but is also intended forhuman use. Other piezoelectric micropumps have beenreported and comparisons between them can be found in(Nguyen et al. 2002; Laser and Santiago 2004; Woias 2005;
Tsai and Sue 2007; Amirouche et al. 2009). Piezoelectric
pumps offer high forces and fast response times, but disadvantages include high voltage requirements and mounting
procedures of PZT disks.
Electrostatic pumps were introduced and demonstrated
by Judy et al. (1991) and Zengerle et al. (1992) starting in
the early 1990’s. Although electrostatic actuation allows fast
response times and low power consumption, only small
stroke volumes and forces are achieved (Amirouche et al.
2009). In thermopneumatic and phase-change type actuators, large pressures and deflection are achieved, but at thecost of high power consumption, slow response times, andthe added fabrication difficulty of a fluid-filled chamber thatmust be sealed. SMA micropumps feature high deflection,
but require high power and special materials. Bimetallic
micropumps feature relatively simple fabrication and lower
driving voltage, but deflection is small and operation at highfrequency is limited (Tsai and Sue 2007). Electromagnetic
actuation offers faster mechanical response than thermopneumatic (Woias 2005) and large forces, but suffers fromhigh power consumption, heat generation, and requirementof an external magnet (Woias 2005; Amirouche et al. 2009).
In ICPF pumps, applied voltage to metal electrodes deposited on either side of a perfluorosulfonic acid polymer
membrane induces bi-directional bending used in pumping.
Although there is concern of repeatability issues related tobatch fabrication, this scheme boasts low driving voltage,
fast response, biocompatibility, and operation in aqueous
environments (Tsai and Sue 2007; Amirouche et al. 2009).
In addition to micropumps, there are several passive
MEMS drug delivery systems consisting of membranesealed microreservoir arrays for controlled release (Santini
et al. 2000; Grayson et al. 2003; Li et al. 2004; Elman et al.
2009; Staples 2010). Although microreservoir technology
has been demonstrated in acute studies in rats (~150 to
200 g) (Li et al. 2004), there are still some limitations interms of chronic drug delivery in small animals. Microreservoirs allow for storage and release of one or multiple

485

drugs, but the reservoirs are not refillable. Thus, there must
be a compromise between duration of the implant and sizeof the implant to minimize surgeries for removal and/or
replacement of the device. In the case of (Li et al. 2004),
membrane dissolution is sensitive to and must be modified
for in vivo environmental conditions, as protein adsorption
affects the gold surface of the membrane and impedes
corrosion. The microreservoir seals required a cleaning cycle and at least 10 min of a square wave cycle for themembranes to open consistently.
Despite abundant work in micropumps and drug delivery,
there is still an unmet need for localized, controllable delivery in small animals such as mice. Previously, we demonstrated a passive reservoir system (Lo et al. 2009), an
electrochemically based system (Li et al. 2008), and anelectrochemical bellows actuator (Li et al. 2010) for applications in drug delivery. Lo et al. (2009) reported the development of the first manually-actuated refillable MEMS drugdelivery device with an integrated check valve and demonstrated intraocular delivery ex vivo in enucleated porcine
eyes and in vivo in rabbit eyes. Electrolysis was chosen as
the actuation method in (Li et al. 2008) for its low power
consumption (~μW to mW), low heat generation, large
driving force, and flow rate control through adjustment of
the applied current (Neagu et al. 1996). An electrochemically driven implantable pump with transscleral cannula (Liet al. 2008) was operated on benchtop and showed a broad
operation range (from pL/min to μL/min) and ability topump against several physiologically relevant backpressures. Feasibility of the electrochemical device for ocularapplications was also demonstrated ex vivo in enucleatedporcine eyes and in vivo in rabbit eyes. A Parylene bellows
electrochemical actuator introduced by Li et al. (2010) separated the electrolysis operation from the drug reservoir in abenchtop testing fixture and achieved large deflection under
low power. Various electrode geometries were investigatedto optimize pump performance in terms of gas generationand efficiency.
Here we present, for the first time, a refillable implantablemicropump system (RIMS) with an integrated electrochemical bellows actuator in a form factor suitable for use in small
animals (e.g. rodents). RIMS provides repeatable, controllable
drug delivery to a specific site. The system is made of biocompatible materials, has a refill port, and features an inert
drug reservoir to enable broad drug compatibility. Our fully
implantable system is capable of adjustment to the delivery
regimen while implanted, a feature not available in commercial devices on the size scale of our system. With the adjustability in flow rate offered by our system, a multitude ofpharmacodynamic responses can be studied in a way thathas not been possible until now. We also demonstrate, forthe first time, an implantable MEMS micropump system forcontrolled dose, timing, and location of drug delivery in mice.

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2 Design
The bellows electrolysis actuator consists of a pair of interdigitated platinum (Pt) electrodes contained within a bellows filled with electrolyte (water). Electrical current
applied to the electrodes splits the water molecules intohydrogen and oxygen gases (Fig. 1). The bellows expands
under the imposed pneumatic driving force of the gases
generated by the electrolysis reaction, which then pushes
the fluid (drug) surrounding the bellows out of the reservoir
and catheter to the delivery site. When the current is turnedoff, the Pt electrode acts as a catalyst for the gases toFig. 1 Illustrations and photograph of the refillable implantable micropump system (RIMS) (a) The main features are an electrochemical
bellows actuator, refill port, and check valve (b) Current applied to apair of interdigitated electrodes splits water (electrolyte) into hydrogenand oxygen gases. The phase change induces a pressure increase within
the bellows and drives fluid out of the reservoir and through thecatheter (c) The reservoir can be refilled through the built-in refill port(d) Photograph of the RIMS

Biomed Microdevices (2012) 14:483–496

recombine as water (Li et al. 2010). The electrolysis actuatorcan be set to on/off states and provide adjustable flow rate
(through applied current magnitude) to achieve a desireddelivery regimen. Our actuation method allows multiple
dosing cycles, which combined with a refillable reservoir,
makes our RIMS appropriate for chronic use.
Pt was chosen as the electrode material due to its biocompatibility, role as the catalyst for recombination, and its resistance to oxidation and corrosion (Meng et al. 2008). Theinterdigitated electrode design was chosen to reduce the resistive path through the electrolyte, which translates to improved
efficiency and lower heat generation (Belmont and Girault1994). The dimensions of the electrode features (100 μm wide
interdigitated fingers, 100 μm spacing) were chosen based on
improved robustness seen with increased finger width. In
consideration for a drug delivery reservoir implanted next toa rounded body feature, such as the eye, it is desirable to havea flexible electrode substrate that coupled with a curved reservoir design so as to better conform to the anatomy of theimplantation site (Lo et al. 2009; Li 2009). To this end, weinvestigated Parylene C as a flexible substrate material. Electrodes on glass substrate were ~8 mm in diameter in terms ofoverall footprint, but the electrodes on Parylene C were fabricated with a smaller overall diameter of ~4 mm due to theintended application in the eye (less space for implantation).
A bellows actuator format (Li et al. 2010) was adopted toseparate the drug from the electrolyzed fluid and prevent associated drug oxidation, an undesirable effect reported in ourprevious work (Li et al. 2009). Parylene C was selected as thebellows material for its biocompatibility, inertness, very low
permeability, and low Young’s modulus to achieve high deflections with low applied pressure. A flexible bellows configuration was chosen over corrugated or flat membranes for theirhigher deflection with lower material stress under pressureloading (Li et al. 2010). The dimensions of the bellows werechosen based on the size of the electrodes and the height of thereservoir. The bellows volume occupies significant space within the reservoir and therefore its dimensions must be appropriately selected to minimize dead volume while still maintaining
the ability to deliver sufficient drug volumes. A bellows with 2
convolutions was chosen over 1 and 3 convolutions to balance
the requirements of deliverable volume against dead volume
and overall device dimensions (Gensler et al. 2011).
The drug reservoir was made of biocompatible silicone
rubber, chosen for its ease and low cost of fabrication, andhas rounded features to reduce irritation and erosion oftissue around the implanted system. The dimensions werechosen based on bellows actuator dimensions and constraints on total implanted volume due to the use in small
animals, and can be easily modified for different applications. For subcutaneous implantation to be a minor procedure, less than 10% of the animal’s surface area should be
involved in the creation of a subcutaneous pocket (Urquhart

Biomed Microdevices (2012) 14:483–496

et al. 1984). The catheter material was biocompatible silicone rubber and its dimensions were chosen to minimize
flow resistance while taking into consideration placement in
the animal. This was in contrast to the Parylene cannula
used earlier in (Li et al. 2008), which had clogging issues,
was not robust enough to withstand surgical handling andimplantation (kinking), and did not include an in-line valve.
The silicone rubber catheter here has a larger, circular crosssection to avoid clogging.
Integration of a check valve in the drug delivery catheter
ensures safe dosing as well as separation of the reservoir
contents from fluids at the implantation site (Lo and Meng2009). The integrated check valve described in (Lo et al.
2009) had a relatively high cracking pressure (~10 kPa or
76 mmHg) and was integrated into a rectangular cannula,which is difficult to secure and seal with sutures. Here wechose to use an in-line check valve with high pressure
shutoff as previously described (Lo and Meng 2009,2011). The valve was designed to open at low forward
pressures during drug delivery, but to be closed under backpressure and at excessive forward pressures greater than apreset threshold to prevent accidental dosing under pumpmalfunction or other unexpected conditions.

3 Methods and materials
3.1 Electrodes
3.1.1 Fabrication
Previously we described a dual-layer photolithography andliftoff process for platinum (Pt) electrode fabrication on a
glass substrate (Li et al. 2010) and flexible Parylene C
substrate (Li 2009; Gensler et al. 2010) (Fig. 2). An undercut formed on the photoresist sidewall assisted with metal
liftoff. A soda lime wafer was treated with 1,1,1,3,3,3-hexamethyldisilazane (HMDS, MP Biomedicals, LLC, Solon,
OH) for improved adhesion between the wafer and photoresist. A layer of photoresist (AZ1518, AZ Electronic Materials, Branchburg, NJ) was spun onto the treated soda lime
wafer at 4 krpm for 30 s and globally exposed at 90 mJ/m2.
A second layer of photoresist (AZ4400, AZ Electronic
Materials, Branchburg, NJ) was spun on at 4 krpm for
30 s and UV patterned with 90 mJ/m2. After descummingfor 60 s in oxygen plasma (100 mW, 100 mTorr), Pt(2,000 Å) or Pt/Ti (2,000 Å/500 Å) electrodes were deposited by e-beam evaporation. Ti acted as an adhesion layer
between the Pt and glass. Electrodes (100 μm wide interdigitated fingers, 100 μm spacing) were then defined bystandard liftoff by immersion in acetone, isopropyl alcohol,and deionized water. Kynar™ wire-wrap wires (30 AWG,
Jameco Electronics, Belmont, CA) were affixed to487

electrodes using conductive epoxy (Epo-tek® H20E, Epoxy
Technology, Billerica, MA), and further coated with nonconductive epoxy (5-Minute Epoxy System, ITW Performance Polymers, Riviera Beach, FL) for insulation.
For the electrodes on Parylene C, a silicon wafer (supporting substrate used only during processing) was coatedwith 5 μm Parylene C (Specialty Coating Systems, Indianapolis, IN). Native oxide was not removed prior to coatingin order to facilitate release of the Parylene C from thesilicon support. The photolithography and liftoff steps were
the same as those for the glass electrodes. Electrodes were
cut with a razorblade and peeled from the supporting Si
substrate, then mounted with epoxy (5-Minute Epoxy System, ITW Performance Polymers, Riviera Beach, FL) on a
flexible polyetheretherketone (PEEK) backing for subsequent packaging.
3.1.2 Characterization
Electrochemical impedance spectroscopy (EIS) is an electrochemical technique where the application of current or
voltage across a broad range of frequencies can provide
information about electrochemical reactions and characteristics of the electrode. EIS can be used to study surfacechanges due to corrosion, effects of dielectric coatings,
electrochemical cleaning, surface activation, or electrolytic
solution conductivity (Parthasarathy 1992; Bard 2001;
K’Owino and Sadik 2005). We performed potentiostatic
EIS (voltage potential between working and reference electrode kept constant) and measured the current value between
the working and counter electrodes, which was then used toplot impedance as a function of frequency. The change inshape of the impedance curve for an electrode before andafter the electrode has been used in electrolysis provides
insight on the status of the electrode surface, indicating, forexample, whether corrosion or delamination have occurred.
Electrodes were characterized using EIS (Gamry Reference
600 Potentiostat and EIS300 Software, Gamry Instruments,
Warminster, PA) performed in 1X phosphate buffered saline
before and after application of current. 0.2, 0.6, and 1.0 mA
were applied successively to the same Pt/Ti electrodes on
Parylene C, and additionally at 10.0 mA for electrodes on
glass.
Electrodes were fixed in a custom laser-machined (Mini/
Helix 8000, Epilog, Golden, CO) acrylic test fixture filled with
deionized water. The application (2 min) of constant current(2,400 Sourcemeter, Keithley, Instruments Inc., Cleveland,OH) induced electrolysis, which split water into hydrogenand oxygen gases. This phase change was accompanied byan increase in pressure, which then dispensed the fluid throughthe fixture output and into a small weigh boat (measured with a
1 mg-resolution digital balance) for volumes >100 μL or into acalibrated micropipet (Accu-Fill 90, Becton, Dickinson and488

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Fig. 2 The dual-layer photoresist process for electrode fabrication on glass (left) or Parylene
C (right, photograph taken prior
to detachment from silicon
support)Company, NJ) for delivered volumes <100 μL. Electrodes on
glass were tested from 1.0 mA to 5.0 mA and electrodes on
Parylene C were tested from 0.2 mA to 2.0 mA (Fig. 3).
3.2 Bellows
Parylene C bellows with 2 convolutions were fabricatedusing a polyethylene glycol (PEG 1,000 Mn, SigmaAldrich, St. Louis, MO) molding process described previously (Li et al. 2010; Gensler et al. 2011). The process usedhere and in (Gensler et al. 2011) improved upon the processdeveloped for (Li et al. 2010). The silicone rubber molds inthe new process are reusable and several steps were eliminated to significantly reduce the time required for the molding process from 1 week to 1 day. Additionally, the Mn of
PEG was reduced from 14,000 to 1,000 in order to obtainsmoother and less brittle PEG forms. The process beganwith punching 9 and 6 mm holes in 0.4 mm thick silicone
rubber sheets (10:1 base-to-curing agent ratio Sylgard 184,
Dow Corning, Midland, MI). These were then layered on
Fig. 3 A sourcemeter provided
current to the electrodes forelectrolysis in a testing fixture,then fluid was dispensed into(a) a weigh boat for volumes>100 μL or (b) a calibrated
micropipet for volumes<100 μL

glass slides as shown in Fig. 4 to form modules. Each
module was filled with molten (~50°C) PEG, taking careto ensure that no air bubbles were introduced to the molds.
After several minutes at room temperature the PEG was
solidified and the silicone molds were peeled away, leaving
solid PEG modules that could be stacked to form the final 2
convolutions PEG form. The modules were fused together
by slightly moistening the appropriate water-soluble PEG
faces of each module with water. A 13.5 μm layer ofParylene C (Specialty Coatings Systems, Indianapolis, IN)
was deposited over the PEG form, and then the PEG was
dissolved out of the bellows by soaking in water at room
temperature.
Load-deflection tests were performed on a 2 convolution
Parylene C bellows (13.5 μm wall thickness). Bellows were
mounted in a custom laser-machined (Mini/Helix 8000,
Epilog, Golden, CO) acrylic test fixture and pressures ranging from 0.0 psi to 0.5 psi (3.45 kPa) were applied inincrements of 0.05 psi (0.34 kPa) using a computercontrolled pressure regulator (900X, ControlAir, Inc.,Biomed Microdevices (2012) 14:483–496

Fig. 4 (a) Fabrication of the bellows consisted of molding a PEG form
using PDMS, deposition of Parylene C over the PEG mold, andremoval of the sacrificial PEG (b) Photograph of the 2 convolution
Parylene C bellows after removal of PEG

Amherst, NH). Deflection was recorded using a compound
microscope (PSM-1000, Motic China Group Co., Xiamen,
China) with a 100× objective lens (1 μm vertical resolution).
The Parylene C bellows was filled with room temperaturedeionized water and adhered to the electrode substrate withdouble-sided pressure sensitive adhesive film (Tape 415, 3M,
St. Paul, MN). The seal was further reinforced with epoxy (5Minute Epoxy System, ITW Performance Polymers, Riviera
Beach, FL). After assembly of the bellows and the electrode,the bellows actuator was mounted in the same fixture forstand-alone electrodes and flow rates were measured at currents ranging from 0.2 mA to 1.0 mA (Li et al. 2010).

489

heated at a rate of 1.5°C to 215°C, the temperature for the heat
shrink process recommended by the FEP tubing manufacturer.
After 30 min of baking, the temperature was ramped down at a
rate of 1.5°C to avoid thermal stress on the SU-8. The FEP
tubing shrank to hold the valve components securely without
the need for adhesives (Fig. 5). The outer diameter of the valve
and FEP packaging was ~1.23 mm, which is significantly
smaller than either of the commercial valves and allows integration into the system without significantly adding to theoverall device size.
The valve was characterized with a custom pressure setup.
Pressurized water was applied to the inlet of the valve incrementally from 0 kPa to 266.62 kPa (2,000 mmHg or
38.67 psi). Flow rate was measured in a calibrated micropipet(Accu-Fill 90, Becton, Dickinson and Company, NJ) attachedto the valve outlet. The valve was kept dry prior to the first
application of pressurized water, but was submerged in water
before the second run to prevent stiction, which was observed
previously and described in (Lo and Meng 2011).
The MEMS check valve was compared to two commercial valves (Fig. 6) using the same pressure setup. The firstwas a one-way disc valve (240270524, Halkey Roberts, St.
Petersburg, FL) constructed of a Class VI biocompatible
polycarbonate housing and silicone seal disc and chosen
for its low cracking and sealing pressures ( <0.25 psi or
1.72 kPa and ≤0.50 psi or 3.45 kPa, respectively, according
to manufacturer). The second, a normally closed in-line
check valve (80031, Qosina, Edgewood, NY) made ofacrylic and ethylene propylene diene monomer, has an evenlower cracking pressure of <0.05 psi or 0.35 kPa according
to the manufacturer.

3.4 Refillable Implantable Micropump System (RIMS)
Rapid prototypes of the reservoir were made using silicone
rubber casting in custom CNC milled acrylic molds. Medical grade silicone rubber (MDX-4-4210, Dow Corning,

3.3 MEMS check valve
The four-part MEMS check valve (Lo and Meng 2011) consisted of a valve seat, pressure limiter, and spacer plate madeof SU-8 2100 (MicroChem, Newton, MA), and an s-shaped
arm valve plate made from MDX-4-4210 (10:1 base-to-curing
agent ratio, Dow Corning, Midland, MI). The components
were aligned and held in a custom fixture and covered with
1.3:1 shrink ratio fluorinated ethylene propylene (FEP) heatshrink tubing (22 G, Zeus Industrial Products Inc., Orangeburg, SC). The entire fixture assembly was placed in a vacuum
oven (Model V0-914A Lindberg Blue, Asheville, NC) and

Fig. 5 Side view of the check valve packaged in FEP tubing

490

Fig. 6 Commercial one-way disc valve (left) and in-line check valve(right) used for comparison to the fabricated check valve

Midland, MI) was thoroughly mixed (Super Mixer AR-250,
Thinky Corp., Tokyo, Japan) in a 10:1 base-to-curing agent
ratio and then degassed in a vacuum chamber (Model V0914A Lindberg Blue, Asheville, NC) for 45 min. The uncured solution was poured into the molds and baked in anoven (EC0A Environmental Chamber, Sun Electronic Systems Inc., Titusville, FL) for approximately 1 h at 80°C tocure the polymer. Class VI silicone tubing (VWR International, Radnor, PA) was set in the mold during curing. A slot
was made on the back reservoir wall for inserting the bellows actuator. A small amount of uncured MDX-4 was used
to adhere the bellows actuator to the inner base of thereservoir and seal around the slit and lid of the reservoir.
The fully packaged device (Fig. 7) was then cured at 80°C
for approximately 45 min. A 30-gauge non-coring needle
was used to fill the reservoir with deionized water for testing
as previously demonstrated (Lo et al. 2009).
The FEP-packaged valve inlet was connected to the integrated silicone catheter and its outlet was connected to acalibrated micropipet as previously described for bellows actuator testing. Constant current of 1 mA was applied (2400
Sourcemeter, Keithley Instruments Inc., Cleveland, OH) for
15 min, then turned off for 45 min. This on/off cycle wasrepeated four times for both valved and non-valved systems
to simulate periodic dosing in a chronically implanted device.
For all of the valved systems, valves were soaked overnight inwater prior to testing. The commercial valves were eachconnected to the outlet of the RIMS catheter in place of theintegrated MEMS check valve and tested similarly.

Biomed Microdevices (2012) 14:483–496

As expected for two metal (reference and working) electrodes
in solution, the overall impedance trend decreases with increasing frequency in all cases (Fig. 8). For electrodes on
glass, after the first application of current, the impedance
decreases at all frequencies and exhibits only slight change
for increasing current. The first application of current (0.2 mA
in this case) acts as a cleaning cycle (Sawyer 1974, p.78) that
increases the effective active surface area of the Pt catalyst,
which can be seen in the impedance shift after the applicationof 0.2 mA in Fig. 8. The electrode surface becomes more
active and stabilizes upon the first application of current. Thischange does not cause any significant effect on the flow rate in
the first versus subsequent tests. Electrodes on glass were able
to withstand up to 10.0 mA of current without any degradation. Stand-alone electrodes were tested at the benchtop prior
to integration with reservoirs. A broad range of currentdependent flow rates was successfully demonstrated. Constantcurrents from 1.0 mA to 5.0 mA applied to Pt/Ti electrodes ona glass substrate resulted in flow rates from ~1 μL/min to
34 μL/min, respectively, in an approximately linear relationship (Li et al. 2010).
Pt electrodes on Parylene C, however, delaminate evenfor modest (>1 mA) applied currents (Fig. 9). The dimensions and surface area of the electrodes on glass were larger
than those of electrodes on Parylene C, thus the impedance
magnitudes of the EIS should not be directly compared.
However, the overall (and expected) trend of decreasing
impedance with increasing frequency was also present for
the electrodes on Parylene C.
Electrodes on Parylene having a Ti adhesion layer werealso fabricated and a subset were annealed (205°C). There isevidence that annealing may help stabilize Ti-polymer interfaces (Faupel et al. 1998). However, we found that at currents above 1 mA delamination was delayed, but notprevented, for Pt-Ti electrodes on Parylene C following
annealing (at 150 and 205°C) and a more aggressive descum
prior to metal deposition. With annealing, electrode delamination occurred on the order of seconds. Overall, the greatest improvement in electrode adhesion under applied current
was attained by inclusion of the Ti adhesion layer but
without subsequent annealing of the metal/Parylene C composite film (Fig. 10). For electrodes on Parylene with the Ti
adhesion layer but no annealing, electrodes did not show
delamination for up to 30 min at constant current less than
1 mA. Above 1 mA, signs of degradation occurred in less10 min.

4 Results

4.2 Bellows and bellows actuators

4.1 Electrodes

Load-deflection tests were performed on a 2 convolution
Parylene C bellows (13.5 μm wall thickness) as described in(Gensler et al. 2011) and a typical result is shown in Fig. 11.
A deflection of approximately 1.08 mm was achieved at

EIS was performed on electrodes on glass and Parylene C
substrates before and after the application of constant current.

Biomed Microdevices (2012) 14:483–496

491

Fig. 7 Assembly of the RIMS.
The bellows actuator is inserted
into the cast silicone rubber
reservoir, then PDMS is used toseal the slit and adhere a lid tothe reservoir. Photograph(bottom) of the MEMS checkvalve integrated with the fully
packaged micropump system(left) and a miniature prototype(right)3.45 kPa (25.85 mmHg or 0.5 psi), demonstrating large
deflection under relatively low applied pressure. This is incomparison to the bellows with a 10 μm wall thickness
described in (Li et al. 2010), which deflected 1.78 mm at

3.45 kPa. Deflection followed a nonlinear relationship withwall thickness, decreasing for thicker walls, and increased
with additional convolutions. The deflection exhibited a
cubic relationship with pressure and at 13.5 μm operates
within the elastic region for our pressure range (Gensler et
al. 2011). Plastic deformation for 13.5 μm wall thickness
typically occurs above 6.89 kPa (51.71 mmHg or 1.0 psi).
Bellows were then adhered to glass electrodes and tested
similarly to the stand-alone electrodes at currents between
0.2 and 1.0 mA, and demonstrated a linear relationship withflow rate (Li et al. 2010), which was consistent with ourprevious observations. An additional test was performed todetermine the maximum deliverable volume before bellows
failure (leaks or pops). For a 2 convolution bellows of
13.5 μm wall thickness, the maximum deliverable volume
was 228 μL and was set to provide a safety margin toprevent plastic deformation of the Parylene C bellows.
4.3 MEMS check valve

Fig. 8 EIS results of electrodes on glass show an overall decrease in
impedance with increasing frequency, as expected for two metals
electrodes (reference and working) in solution

The check valve cracking pressure was approximately
1.31 kPa (9.82 mmHg or 0.19 psi) and the (forward) closingpressure 93.29 kPa (699.64 mmHg or 13.53 psi). The valve
remained closed and no leaks were observed up to the highest

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Biomed Microdevices (2012) 14:483–496

Fig. 9 EIS results of electrodes
on Parylene show an overall
decrease in impedance withincreasing frequency, as
expected for an electrochemicalcell. Delamination was seen
above 1 mA for Pt-Ti electrodes
deposited on a Parylene C substrate. As fabricated (top right)
and delaminated (bottom right)
electrodes

forward pressure tested at 266.62 kPa (1999.63 mmHg or
38.67 psi). Keeping the valve hydrated did not significantly
change the cracking pressure, but the (forward) closing pressure of the valve increased. The MEMS check valve wascompared to two commercial valves and the cracking pressures are shown in Table 1.

Constant current of 1.0 mA was applied for 15 min tobellows actuators (Pt/Ti electrode on glass, 2 convolution
bellows) in a non-valved and three valved versions of the
RIMS. The duration of current application was based on
total dose requirements and desired flow rate for a specificapplication of the system in radiation dose reduction incancer. The system had a power consumption of ~3 mW

when operated at 1 mA. Current was turned off for
45 min in between delivery periods to allow recombination to occur and the bellows to deflate. The nonvalved RIMS (Fig. 12) produced a flow rate of 4.72±0.35(mean ± SE, n04) μL/min with current ON. There was an
approximate 5-minute delay until significant positive fluid
flow for the non-valved RIMS. When the current was removed, the gases recombined and the bellows deflated. Without a valve and with the outlet of the catheter (or micropipet)
exposed (to air or an uncovered container with liquid), fluid
moved along the catheter in reverse towards the reservoir.
Even after recombination had ceased, the position of the fluidin the catheter did not fully return to the original starting
location.
Neither of the commercial valves nor the MEMS checkvalve fully prevented backflow during the recombination

Fig. 10 Flow rates (mean ± SE, n 04) for Pt/Ti electrodes on a
Parylene C substrate show higher variability than those on glass andexhibit delamination, particularly at currents above 1.0 mA

Fig. 11 Load-deflection results for a 2 convolution, 13.5 μm wall
thickness bellows under applied pressure

4.4 Refillable implantable micropump system

Biomed Microdevices (2012) 14:483–496
Table 1 Comparison ofvalve cracking pressures

493

Valve

Cracking pressure(psi / kPa / mmHg)

MEMS checkvalve
One-way discvalve
In-line checkvalve

0.19 / 1.31 / 9.82
<0.10 / 0.690 / 5.17
<0.05 / 0.345 / 2.59

period (Fig. 13). The delay to forward flow in the RIMS with avalve was only slightly longer (seconds) than the non-valved
RIMS, but the valves added resistance to the flow path. Flowrates were slightly reduced to 4.28 μL/min (MEMS checkvalve), 4.67 μL/min (one-way disc valve), and 4.49 μL/min(in-line check valve). In all cases, the fluid front in the catheter
flowed in reverse during recombination, but did not return
completely to the original starting position.

5 Discussion
Electrode delamination was one of the primary concerns forthe system’s electrochemical actuator when fabricated on
the flexible polymer substrate. Three electrode fabrication
methods were investigated for the polymer substrate case: Pt
on Parylene, Pt-Ti on Parylene, and Pt-Ti on Parylene withannealing. We found that a Ti adhesion layer without subsequent annealing resulted in the most robust electrodes on
Parylene C substrate, but for higher and longer durations ofFig. 12 Constant current of
1.0 mA was applied for 15 min
and turned off for 45 min for 4
cycles in a RIMS with no valve.
After current was removed, andwithout a valve, a significant
portion of the delivered volume
flows in reverse duringrecombination (inset)

applied current delamination was still observed. Generalmechanisms for metal-polymer adhesion are physisorption
due to Lifshitz-Van der Waals forces, molecular interactions,chemical interaction, and electrostatic interaction (Faupel etal. 1998; Lee 1991). Improved adhesion with a Ti adhesion
layer is likely due to double bonds that form between Ti andcarbon in the Parylene C as suggested in (Vasenkov 2011).
Lee et al. (2004) demonstrated that annealing increased
crystallinity in Parylene C films and led to improved adhesion to Au. Wu et al. (1996), however, found conflicting
results and attributed decreased adhesion after annealing to
two possible mechanisms: the thermal stress induced by adifference in the coefficients of thermal expansion (CTE)
between Parylene N (44–69 ppm/K) and Cu (16.6 ppm/K),
and phase separation that transforms the intermixed interfacial layer (Parylene and metal) into an abrupt interface. In
our experience, annealing did not improve adhesion of Pt-Ti
to Parylene. For Pt electrodes on Parylene without theadhesion layer, some improvement in adhesion was seen
with annealing at 205°C. Parylene C has a CTE of
35 ppm/K, whereas Pt and Ti have a CTE of 8.8 ppm/K
and 8.4–8.6 ppm/K, respectively.
The bellows prevented the drug from being affected by
the electrolysis reaction, but did not significantly impede
actuation. We previously reported that varying the bellows
convolution number (and thus bellows volume) (Sheybani etal. 2011) did not significantly affect actuator performance interms of flow rate. Recombination is essential to being able
to deliver multiple drug boluses, because the bellows should
only be operated within the limit of plastic deformation.

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Biomed Microdevices (2012) 14:483–496

and it remains a challenge to develop reliable MEMS checkvalves for this application; further research on check valves isrequired. Valved RIMS also experienced a slight delay toforward flow due to the added resistive element and pressurebuild up required to overcome the valve cracking pressure.
However, the slopes of the curves for non-valved andcommercially-valved RIMS were similar as the flow rate
was coupled to gas generated by the applied constant current.

6 Application in cancer therapy

Fig. 13 Constant current of 1.0 mA was applied for 15 min and turned
off for 45 min in a RIMS with no valve (n04 consecutive runs), a
commercial one-way disc valve (n04 consecutive runs), a commercial
in-line check valve (n01), and a MEMS check valve (n01)

After determining the most robust electrode fabrication
process, the actuator with electrolyte-containing bellows
was integrated into the drug reservoir. Electrolysis enables
current-controlled conversion of electrical energy to mechanical energy. Electrolysis does not involve any side
reactions (Neagu 2000), making electrolysis efficient interms of low heat generation. Some delay to forward flow
during the current application period was present in all
RIMS tested (no valve, two commercial valves, and one
MEMS check valve) and was attributed to the compliant
nature of the silicone reservoir and catheter, low current
applied (Neagu 1998), and slow diffusion of hydrogen andoxygen gases away from the electrode (Böhm et al. 1999).
Recently, Nafion®-coating of Pt electrodes demonstrated
great improvement in actuator efficiency, linearity, and repeatability (Sheybani and Meng 2011). Gases are moresoluble in Nafion® than water, and therefore gases quickly
diffuse away from the electrode to prevent bubble occlusionof the reaction sites. The coating also improves recombination, such that recombination can occur on the order ofminutes. Future RIMS will employ Nafion®-coated actuators for even greater accuracy and reduced delays in pumping onset after current application.
For in vivo studies, preventing reverse flow is desired toprevent biological fluids from mixing with the drug from thereservoir. The commercial valves and our fabricated valve
require a finite sealing pressure (> 6.89 kPa for one-way discvalve, >1.03 kPa or 7.76 mmHg for in-line check valve)
which exceed that produced in RIMS, resulting in significant
reverse flow during recombination. From these results it isclear that a valve with a much lower sealing pressure (noreverse leakage) is necessary for our system. Commerciallyavailable valves could not provide the desired performance

A potential application of the RIMS is for controlled, localized
delivery of short half-life compounds in vivo, such as siRNA(Dykxhoorn et al. 2006). In preliminary studies, we demonstrated RIMS for subcutaneous delivery of anti-cancer drugs totumors (Gensler et al. 2010). An siRNA-gold nanoplex drug
(HNB-001) was developed that reduces radiation resistance in
certain types of tumors expressing a gene for the protein
sphingosine-kinase-1 (SPK-1), but lacked a suitable sitespecific delivery method (Dykxhoorn et al. 2006; Gensler et
al. 2010). HT29 cells were implanted subcutaneously in bothsides of the flank of nude mice to induce tumors, with the left
tumor receiving treatment and the right acting as a control withno treatment (Gensler et al. 2010). All procedures were performed in accordance with protocols approved by the University of Southern California’s Institutional Animal Care and
Use Committee. A RIMS with Pt-Ti electrodes on Parylene
C (no valve) delivered the nanoplexes directly to the tumor sitein mice under constant current of 0.78 mA. The RIMS (active
pumps) delivered a bolus of ~50 μl daily for 2 weeks, withreservoir refills twice per week. Passive pumps were assembled and used to deliver control treatments: phosphate bufferedsaline, gold nanorods, free siRNA, and nanoplexes. These
units did not contain an electrochemical bellows actuator or
valve and consisted only of the silicone reservoir cavity andcatheter. Without the actuator, these implanted pumps provided
drug by diffusion from the reservoir through the catheter.
Injections of phosphate buffered saline, free siRNA, and nanoplexes were also conducted at the same rate and with the same
bolus volume as the RIMS. A subset of the treated tumors were
exposed to 100 rad of radiation. The active delivery of nanoplexes by the RIMS in combination with radiation treatment
showed significant tumor regression (~50%) over both diffusion and injection, with the greatest effect occurring adjacent tothe catheter tip (Gensler et al. 2010).

7 Conclusions and future work
We developed a refillable implantable micropump systemwith great potential for a multitude of applications, fromdrug development to highly flexible pharmaceutical studies

Biomed Microdevices (2012) 14:483–496

in research animals as small as mice. We demonstrated theability to administer drug to small animals with control overdelivery rates and volume, with an example of chronicdelivery of a novel gene therapy drug. We have shown
preliminary success in using gene therapy in combination
with radiation to combat cancerous tumors, and that active
delivery was essential to this outcome. Our next steps include using Nafion®-coated actuators in RIMS to further
improve repeatability of the micropump, using a rigid material for the reservoir to eliminate compliance, improved
catheter designs to ensure optimal delivery to the site, anddevelopment of a valve with no reverse leakage. Future
devices will be powered wirelessly, allowing the animals
to move freely without tethers or wires and enabling a newparadigm in laboratory research with small animals.
Acknowledgements This work was supported in part by the Wallace
H. Coulter Foundation Early Career Translational Research Award,
National Institutes of Health/National Eye Institute under award number R21EY018490, and National Institutes of Health/National Instituteon Drug Abuse under award number R21DA026970. H.G. was supported by a National Science Foundation Graduate Research Fellowship. The authors would like to thank Dr. Ken-Tye Yong, Dr. Indrajit
Roy and Dr. Paras N. Prasad of University of Buffalo, The State
University of New York, for providing the gold nanoparticles; Dr.
Rizwan Masood and Dr. Uttam K. Sinha for providing the HNB-001
siRNA-gold nanoplexes and their surgical expertise; and Sutao Zhu of
the University of Southern California for her surgical expertise. We
would also like to thank Dr. Donghai Zhu and the members of the USC
Biomedical Microsystems Laboratory for their assistance with thisproject.

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