Characterization of a Wireless Implantable
Infusion Micropump for Small Animal Research
Under Simulated In Vivo Conditions
A.M. Cobo, IEEE Student Member, H.M. Tu, R. Sheybani, IEEE Student Member, and E. Meng, SMIEEE
Department of Biomedical Engineering
University of Southern California
Los Angeles, USA
acobo@usc.edu
Abstract— A wirelessly-operated implantable site specific drugdelivery micropump with on-demand control of dosing forchronic infusion applications in small laboratory animals ispresented. The system includes an electrolysis-based actuator,
refillable drug reservoir, wireless powering, and a check valve.
Micropumps were fabricated, assembled, and tested in simulated
in vivo conditions. As expected, power transfer decreased whenvarying the distance and angle between the transmitter andreceiver coils resulting in a drop in flow rate; these conditions
replicate variations in pump orientation with respect to an
external transmitter coil as would be the case in a device placedin a freely moving animal. Stationary micropumps operated
wirelessly successfully delivered 30 μL doses daily for seven days
as required by an anti-cancer application. The average flow rate
across the pumps was 1.93 ± 0.35 μL/min. Flow rate performance
was not significantly affected by simulated in vivo conditions suchas back pressure of up to 20 mmHg and delivering solutions withviscosities of up to 6 cP.
Keywords— drug delivery; implantable mircopumps; wirelesspowering; MEMS electrochemical bellows actuator; check valve

I.

INTRODUCTION

Rodents are the most widely used animal model in thestudy of human disease [1] and drug administration
technologies suitable for use in rodents are therefore criticallyimportant in the development of new treatments. Oral andintravenous routes are commonly used for intermittent
delivery but require frequent animal handling which induces
stress. Stress in turn triggers undesirable physiologicalresponses which may alter experimental outcomes. Tetheredinfusion systems provide greater accuracy, but are impracticalfor chronic studies as the tether restricts movement, poses
entanglement risks, and also induces stress in the animal.
Implantable drug delivery devices eliminate frequent
handling, tethers, and stress induction while providing site
specific and controlled drug administration [2, 3]. Althoughimplantable pumps are commercially available, none aresuitable for controlled chronic dosing experiments in mice.
These commercial pumps possess one or more of thefollowing deficiencies: too large, limited drug payload, limited
lifetime due to single-use battery, single fixed flow rate, orThis work was supported in part by the NIH/NIGMS under award number
R21GM104583, NSF PFI: AIR Technology Translation program under award
number IIP-1343467, and the Viterbi School of Engineering Ph.D. MeritFellowship.
E. M. has a significant financial interest in Fluid Synchrony LLC(Alhambra, CA, USA).

inability to adjust flow rate after implantation. For example,
the Alzet osmotic pump is the only commerical pump smallenough to be used in mice, but provides a single preset flow
rate, has a non-refillable drug reservoir, operates continuously,
and suffers from poor control over the initiation of delivery[4]. No commercial pumps are wirelessly operated.
Previously, we reported on the development of a wired
implantable micropump for acute, site specific drug delivery
in small laboratory animals (e.g. mice) [5-7]. The micropump
was actuated using electrolysis which offers low power
consumption, large driving force, low heat generation, andelectronic flow rate control via magnitude of applied current(flow rate linearly related to current) [7]. Here, we present awireless implantable drug delivery system based on the same
pumping principle that is suitable for chronic infusions.
Wireless operation enables chronic dosing in a freely moving
animal which significantly expands the utility of themicropump. This wireless drug administration method caneliminate the induced stress from animal handling or tether
and wired drug administration systems that could potentiallyconfound study results [8]. This pump has been further
miniaturized from the previous wired version to minimize
dead volume and realize a more compact form factor.
II.

DESIGNThe micropump system is based on the previously
developed electrolysis actuation mechanism [5-7]. For
application in small animals, in particular mice, the total
device mass should be ≤ 10% of the weight of the animal(typically adult mice are 30 g) to allow free movement. Major
components of the pump include the electrochemical actuator,drug reservoir, check valve, catheter, refill ports, and wirelesselectronics. The pump is implanted subcutaneously and thereservoir can be periodically refilled via ports consisting ofsilicone rubber plugs.
The electrochemical actuator consists of a pair ofinterdigitated electrodes in contact with an electrolyte (water)
encased in a polymer bellows. The bellows membrane
separates the drug in the reservoir from the electrolyte. An
applied electrical current to the electrodes generates oxygenand hydrogen gas and produces a corresponding increase inpressure. This in turn deflects the bellows, displaces thesurrounding fluid, and activates a passive check valve toachieve site specific delivery at the catheter outlet. The gases
recombine once the current is turned off, eventually allowing
the bellows to reset to its initial position [7].
The bellows possesses two rectangular cross-section
convolutions having 6 mm inner diameter, 9.5 mm outer
diameter, and 13.5 μm wall thickness. The bellows candisplace 185 μL of fluid contained within the adjacent drug
reservoir without undergoing plastic deformation. Compared
to previous pump prototypes, this new design greatly reduced
dead volume by 46% to achieve a more compact footprint.
This improvement is largely attributed to the reduced reservoir
size and modified assembly process matching the bellows
actuator to the reservoir.
In this system design, the pump is normally off until
wirelessly activated. A Class E inductive powering system
consisting of a 9 V power supply, an external transmitting
circuit and coil placed underneath the animal cage, and animplanted receiving circuit and coil packaged with themicropump, was designed (Fig. 1). A 2 MHz clock oscillator(ECS -2100, ECS international, Olathe, KS) was used to create
the power signal. The generated signal was then amplified in
two stages before being applied to a tuned transmitting coil (8turns of 20 AWG single strand wire, size: 310 mm x 140 mm).
Litz wire (6 turns, 50/54 SPN/SN Litz Wire, Wiretron,Volcano, CA) was used for the receiving coil (Ø 17 mm). The
received alternating signal was fully rectified using two
Schottky diodes (BAT54A and BAT54C, Fairchild
Semiconductor, San Jose, CA). A resistor was used to set theoutput current on a 3-terminal adjustable current source to0.33 mA. This output current powers the pump actuator andachieves a single flow rate. Flow rate selection is possible bymodifying the circuitry to allow for user adjustable driving
current and is currently under development.

Fig. 1. Schematic diagram of in vivo testing setup for small animalresearch.

III.

FABRICATIONThe electrolysis actuator consisted of a pair ofinterdigitated Pt/Ti electrodes (100 μm wide elements
separated by 100 μm gaps, 8 mm diameter footprint, 300Å/2000 Å Ti/Pt) and Parylene C (Specialty Coating Systems,Indianapolis, IN) bellows filled with Milli-Q water. Electrodes
were supported on a borosilicate glass and Nafion®-coated(Dupont DE521Solution, Ion Power, INC, New Castle, DE) to
improve electrolysis efficiency [9, 10]. Parylene C bellows
were fabricated using a lost-wax two-part molding process as
detailed in [11] and were securely mounted over the electrodes
with adhesive.

The reservoir dome and base were produced using
stereolithography in WaterShed® XC 11122 polymer(FineLine Prototyping, Inc., Raleigh, NC). This material is
optically clear, durable, and water resistant. The reservoir
refill ports were filled with medical grade silicone rubber
plugs (MDX-4 4210; Factor II, Lakeside, AZ). The pumphousing was coated with 5 μm of Parylene C to improvebiocompatibility and barrier properties. The bellows actuator
was situated in the reservoir assembly, and the dome and baseinterface was secured with biocompatible epoxy (EPO-TEK®
730 unfilled, Epoxy Technologies, Billerica, MA).
A medical grade silicone check valve was integrated intothe pump housing in the output path (2.0 mm duckbill;
Minivalve, Cleveland, OH). Valves were prescreened using acustom acrylic fixture to ensure consistent performance andlow reverse leakage before integration into the pump. Valves
were affixed to the reservoir housing using biocompatible
epoxy. The delivery catheter consisted of silicone,
polytetrafluoroethylene and polyurethane tubing segments
joined together to reduce from the larger diameter valve
housing to a smaller diameter catheter outlet. The secondary
receiver circuit was wrapped around the pump body. The final
micropump assembly was encapsulated with 5 μm Parylene C
followed by medical grade silicone (Fig. 2). The total mass ofthe prototype pump with a filled reservoir was 3.8 grams.

Fig. 2. Schematic diagram of (a) drug delivery system and (b) a fully
assembled wireless micropump.

IV.

EXPERIMENTAL METHODS

Detailed characterization of the pump actuator
performance was described in [5-7]. Here, wireless operationof the pump system as a function of coil separation distanceand angular misalignment was investigated. This replicates thepotential positions of a pump implanted in freely moving
animal with respect to the transmitter coil below the cage (Fig.
1). Then, the ability of the pump to provide dosing following apredetermined regimen for an experimental study of a genetherapy against radiation resistance in cancer was examined.
These experiments adhered to the dose and flow raterequirements for delivery of a siRNA-gold nanoplex drug thatwill be carried out in vivo in xenograft tumors in nude mice inplanned studies. This application requires a low flow rate(~μL/min), 30 μL daily dose volume, and infusion duration ofapproximately two weeks with one refill per week. Finally, aseries of experiments were performed to evaluate micropump
performance when delivering against a physiologically
relevant back pressure and solutions of differing viscosity.
These experiments elucidate the operating range addressed by
the micropump. All experiments were performed at room
temperature since the flow rate only increases slightly whenmonitored at body temperature (37 °C) [7].
A. Coil separation distance and angular misalignment
Power transfer between inductive transmission coils is
optimal when concentric transmitter and receiver coils areplaced in parallel and as close to each other as possible [12].
The micropump is intended to be implanted in a moving
subject. Therefore, it is expected that the axial separation
distance and angular misalignment between the transmitting
and receiving coils will vary over the course of administration.
The separation distance (parallel coils, 2 – 4 cm, 0.5 cm steps)
and angle (0, 30, 45, and 60°) between concentric transmitterand receiver coils was varied to simulate practical operating
conditions encountered in moving subjects and study theirimpact on pump performance.
B. Daily dosing
Four stationary (2 cm separation distance and 0° angular
misalignment between concentric coils) micropumps wereoperated daily for a period of one week to simulate the dosing
regimen planned for the anti-cancer application. A constantcurrent of 0.33 mA was supplied to the actuator wirelessly fora duration that allowed 30 μL of Milli-Q water to be delivered
daily. Water displacement in a 100 μL calibrated micropipette(VWR International, Radnor, PA) connected to the catheter
outlet was recorded and forward flow rate was calculated.
Reverse leakage between dosing events, if any, was recorded.
C. Back pressure
To mimic the presence of finite pressure within the various
compartments of the body, micropumps were operated againstback pressures varying from 0 to 20 mmHg applied at thepump outlet. The upper limit of the back pressure was selected
because it exceeds the pressure anticipated for subcutaneous
drug delivery in mice [13]. The catheter outlet was connected
to one end of a 100 μL calibrated micropipette while the otherend was connected to a custom pressure setup. The flow rate
was determined as previously described during trials lasting 5
minutes with 2 minute rest periods between trials. Four trials
were conducted at each back pressure and the flow rate
averaged.
D. Viscosity
Since delivered solutions may possess a range ofviscosities, pump performance using glucose solutions
calibrated to different viscosities ranging from 1 to 6 cP wasinvestigated. Anhydrous D-glucose (VWR International,Radnor, PA) was dissolved into Milli-Q water in five different
concentrations. Solutions were chosen based on a calibrated
curve determined using a Cannon-Fenske viscometer (data notshown). The flow rate was determined as previously described

during trials lasting 5 minutes for each solution. Four trials
were conducted at each viscosity and the flow rate averaged.
When changing solutions, reservoirs were flushed with MilliQ water to prevent changes to the viscosity of the newsolution.
V.

RESULTS AND DISCUSSION

A. Coil separation distance and angular misalignment
Changing the axial separation distance between coils andangular misalignment between transmitter and receiverresulted in a flow rate drop of 64.1% and 42.86% for a 3.5 cm
distance between coils and for 45° misalignment between thetransmitter and receiver coils, respectively. Strategies toameliorate the fluctuations in power transmission during
animal movement and maintain consistent pump performance
include increasing the power transmitted and including asecond receiver coil.
B. Daily dosing
The wirelessly operated stationary micropumps
consistently delivered 30 μL doses daily over the course of
one week. Fig. 3 shows representative results from the daily
dosing regimen of one micropump. Similar data was obtained
for all four pumps tested. The average flow rate across thepumps was 1.93 ± 0.35 μL/min. Since we previously
demonstrated consistent pump actuator (consisting ofelectrode pair, electrolyte, and bellows) performance [7],slight variations in mean flow rate were attributed to
packaging variations in the manually assembled pumps, valve
performance, and receiver coil fabrication. Also, despiteprescreening parts, the valves consistently permitted up to 10
μL of back flow when the pump was turned off. The durationof applied current was increased with each successive dosing
event as the pumps were not refilled during the one week
study. This delay is expected and can be attributed to therecombination reaction between electrolysis actuation cycles[7].

Fig. 3. Representative data from one pump wirelessly operated once a
day for seven days (2.79 ± 0.10 μL/min).

C. Back pressure
Minimal variation in flow rate compared to the value at 0mmHg back pressure was observed up to a back pressure of 20mmHg. Flow rate decreased slightly as the back pressureincreased, but remained within 10% of the baseline flow rateof each pump recorded at 0 mmHg. No significant differenceswere observed between micropump flow rates whennormalized to 1, as confirmed by statistical analysis (ANOVA,p < 0.05). The data was plotted using normalized flow rate
values (Fig. 4).

Fig. 4. The variation in flow rate observed in four micropumps when
pumping against physiological back pressures (up to 20 mmHg). Flowrate was normalized across pumps at a back pressure of 0 mmHg.

D. Viscosity
The same four micropumps were tested with solutions ofvarying viscosities and no significant differences in flow rateswere observed. The baseline flow rate is defined by the value
measured by delivering Milli-Q water with a 1 cP viscosity atroom temperature. Flow rates were within 10% from thebaseline flow rate for each pump tested. While the mean flow
rates across the pumps were different as previously explained,
the normalized flow rates between micropumps were notsignificantly different as confirmed by statistical analysis(ANOVA, p < 0.05; Fig. 5).

VI.

We developed an implantable wireless drug delivery
micropump for use in chronic infusions into small research
animals. Wireless control of operation permits a user to selectthe desired onset of delivery and delivery regimen in a freely
moving animal that is not tethered by wires or catheters. We
tracked the pumping performance by monitoring flow rate as afunction of axial separation distance between the transmitting
and receiving coils as well as during angle misalignment at a
fixed separation distance. As expected, a decrease in powertransmitted resulted in a decrease in flow rate as coilseparation distance and angular misalignment were increased.
Performance can be improved if necessary by increasing thepower transmitted and inclusion of a second receiver coil.
Stationary micropumps accurately delivered a single bolus
each day, as might be required in an experimental regimen forevaluating anti-cancer drugs. We also demonstrated that theflow rate is not significantly affected by back pressures of up
to 20 mmHg (when delivering solutions of 1 cP) or deliveringsolutions with viscosities of up to 6 cP. We are now
performing additional testing to study flow rate performancein response to additional simulated in vivo conditions. Future
work will include flow rate adjustment capabilities after
implantation and integrated dosing sensors to control
activation time.
ACKNOWLEDGMENT
The authors thank the members of the USC BiomedicalMicrosystems Laboratory for their assistance with this project.
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Fig. 5. The variation in flow rate observed in four micropumps when
delivering calibrated glucose solutions of differing viscosity. Bloodviscosity is 3 to 4 cP at body temperature (1cP = 1mPa·s). Flow rate wasnormalized across pumps at a viscosity of 1 cP.

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