Sensors and Actuators A 239 (2016) 18–25

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A wireless implantable micropump for chronic drug infusion against
cancerAngelica Cobo a , Roya Sheybani a , Heidi Tu a , Ellis Meng a,b,∗a
Department of Biomedical Engineering, Viterbi School of Engineering, University of Southern California, 1042 Downey Way, DRB-140, Los Angeles, CA
90089-1111, USAb
Department of Electrical Engineering, Viterbi School of Engineering, University of Southern California, 3651 Watt Way, VHE-602, Los Angeles, CA
90089-0241, USAa r t i c l e

i n f oArticle history:
Received 24 July 2015
Received in revised form
12 November 2015
Accepted 5 January 2016
Available online 12 January 2016
Keywords:
Drug delivery
Wireless powering
Implantable micropumps
MEMS electrochemical bellows actuator
Check valve
Cancera b s t r a c t
We present an implantable micropump with a miniature form factor and completely wireless operationthat enables chronic drug administration intended for evaluation and development of cancer therapies infreely moving small research animals such as rodents. The low power electrolysis actuator avoids the needfor heavy implantable batteries. The infusion system features a class E inductive powering system thatprovides on-demand activation of the pump as well as remote adjustment of the delivery regimen without
animal handling. Micropump performance was demonstrated using a model anti-cancer application inwhich daily doses of 30 ␮L were supplied for several weeks with less than 6% variation in ﬂow rate within
a single pump and less than 8% variation across different pumps. Pumping under different back pressure,viscosity, and temperature conditions were investigated; parameters were chosen so as to mimic in vivo
conditions. In benchtop tests under simulated in vivo conditions, micropumps provided consistent andreliable performance over a period of 30 days with less than 4% ﬂow rate variation. The demonstrated
prototype has potential to provide a practical solution for remote chronic administration of drugs toambulatory small animals for research as well as drug discovery and development applications.
© 2016 Elsevier B.V. All rights reserved.

1. Introduction
Each year, approximately 8 million people around the world
die from cancer. Research suggests that one-third of these deaths
can be prevented with appropriate technologies and services [3].
Treatment typically involves surgery to remove cancerous tissues, followed by long-term chemotherapy or radiation therapy,
among other types of therapies [28]. Chemotherapy and radiotherapy are used to initiate intrinsic apoptotic (cell death) pathways,
but these type of treatments, when administered systemically, arenon-speciﬁc and affect non-cancerous tissues as well as lead tosevere side effects. Delivery to a speciﬁc tissue facilitates the goalof achieving therapeutic drug levels while reducing harmful side
effects, the amount of drug needed, and the frequency of dosing

Abbreviations: siRNA, small interfering ribonucleic acid; Pt, platinum; PBS, phosphate buffered saline; PDMS, polydimethylsiloxane; PEG, polyethylene glycol; USP,
United States Pharmacopeia; SL, stereolithography; EI, electrochemical impedance.
∗ Corresponding author at: Department of Biomedical Engineering, Viterbi Schoolof Engineering, University of Southern California, 1042 Downey Way, DRB-140, Los
Angeles, CA 90089-1111, USA. Fax: +1 213 821 3897.
E-mail address: ellis.meng@usc.edu (E. Meng).
http://dx.doi.org/10.1016/j.sna.2016.01.001
0924-4247/© 2016 Elsevier B.V. All rights reserved.

events [19]. Additionally, studies have shown that periodic biological ﬂuctuations such as circadian rhythms may affect drug kinetics
and dynamics [37]. Administration of chemotherapy at a speciﬁctime in a tumor cell cycle can increase treatment effectivenesswhile decreasing toxicity to healthy tissue [43]. A study of morethan 30 anti-cancer drugs showed that the toxicity and efﬁcacy ofthese drugs varied by more than 50% depending on the time of dosing [21]. Patient tolerance to medication and clinical outcome arealso greatly affected by the dosing time [21]. Greater control overdelivery rates allows tailored therapies and drug regimen changes
during the course of therapy. Consequently, there is a great need fortherapies combined with delivery methods to allow for long-term,
localized, and patient tailored anti-cancer therapies.
In drug development for disease management, the scientiﬁccommunity relies heavily on animal studies to provide basis
for human clinical trials [31]. The mouse genome has been
completely sequenced allowing accurate comparisons to othermammal genomes such as humans and its physiologic, metabolicand pathologic processes have been well documented [24]. As
such, human drug therapies are initially evaluated in small animal models; therefore, drug administration technologies for smalllaboratory animals are critically important for drug discovery anddevelopment. The most commonly used drug administration routes

A. Cobo et al. / Sensors and Actuators A 239 (2016) 18–25

for laboratory animals are oral, topical, or intravenous which
require high drug dose concentration, leading to severe side effects,
and frequent animal handling. In chronic studies, frequent animal
handling or restraint induces stress, which can confound study
results [17].
Current technologies for chronic dosing in laboratory animals
include tethered infusion systems and vascular access ports. In
tethered systems, the animal is connected to a precision infusion
pump with a tether as shown in Fig. 1a [40]. The use of the external pump provides accurate dosing, but is impractical for chronic
studies since restriction of movement by the tether induces stress.
Vascular access ports are chronically implanted subcutaneously
minimizing animal movement restrictions [41]. However, vascular ports require regular cleaning, maintenance, and appropriate
protective measures to prevent infection or damage by the animal.
Implantable infusion pumps further minimize animal handlingand permit natural movement after the initial healing period following implantation. The Med-e-Cell Infu-Disk is non-reﬁllable, hasa preset constant ﬂow rate, and a maximum lifetime of 1 week
[26]. This pump is only suitable for drug delivery in animal models
larger than rats due to its large size. The Primetech iPrecio and theAlzet pumps are the only commercial pumps scaled for mice. The
iPrecio peristaltic pump allows programmable infusion proﬁles,but is battery powered, only provides up to 52 days of operation,cannot be powered off, and is single use [16]. The Alzet osmotic
pump provides a preset continuous ﬂow rate, a drug payload ofup to 6 weeks, and is non-reﬁllable [1]. These technical limitationsof current technology prohibit ﬂexibility in designing administration regimens for drug discovery and development research. Whilecommercially available pumps are not suitable for chronic drug
infusion in small animals, recent advances in microelectromechanical systems (MEMS) technology have led to the miniaturizationand improved efﬁcacy of implantable drug delivery systems suitable for small laboratory animals [27]. Several ongoing efforts arebeing carried out in research laboratories to develop remotely activated drug delivery microsystems. Detailed descriptions of these
efforts are available in [36,4].
The micropump presented in this work was actuated using
electrolysis which offers low power consumption, large driving
force, low heat generation, on-demand activation, and postimplantation ﬂow rate control through the applied current [33].
Wired implantable MEMS micropumps scaled for use in small laboratory animals (e.g., mice) were previously demonstrated [10,8].
Preliminary in vivo studies with this wired system demonstrated
the potential of using micropumps to combat radiation resistanceof tumors in mice by improving the exposure of siRNA through sustained localized delivery compared to needle injections [10,8,25].
However, the use of transcutaneous wires for powering the pumprequired animals to be restrained during dosing, prohibited chronic
studies, and placed restrictions on experimental throughput. Also,
the lack of reliable, small form factor check valves resulted in a slow
diffusion of drug from the reservoir into the animal and thereby limited the accuracy of delivery. The absence of a reliable check valve
could also lead to ﬂuid to ﬂow back through the delivery catheter
to the reservoir as a result of the reverse pressure gradient createdduring recombination of electrolysis gases.
Here, we present a redesigned implantable micropump thatincorporates wireless power telemetry and an integrated miniaturecheck valve, suitable for chronic drug delivery in untethered, freely
moving small laboratory animals such as mice (Fig. 1b, Table 1). Thiswireless drug administration method can eliminate the induced
stress from animal handling or tethered and wired drug administration systems that could potentially confound interpretation ofdata [17]. Wireless operation was facilitated by improving electrolysis actuation efﬁciency which further reduced actuator power
consumption (1 mW from 3 mW). A miniature check valve with

19

Table 1Primary differences between the previous generation reﬁllable implantable micropump system (RIMS) and our current implantable micropump for drug delivery insmall animals.

Operation

One way valve
speciﬁcations

HeightFootprint
Dead volume
Powerconsumption

RIMS Pump [8]Current Pump

Wired, external
constant current power
supplyBridge-type membrane
valve
0.515 × 1.23 mm, L × OD
1.31 kPa cracking
pressure
10 mm
21 mm OD
1087 ␮L
3 mW

Wireless power
telemetry
Duckbill valve
3 × 2 mm, L × OD
0.69 kPa cracking
pressure
8.1 mm20 × 15 mm, L × W
96 ␮L
1 mW

small form factor and improved performance enabled appropriate ﬂow regulation during pumping and resting states. The checkvalve contributes to accurate dosing and prevents drug contamination with biological ﬂuids during the recombination reaction
between electrolysis actuation cycles [33]. Acute characterizationof pump performance was previously presented [5]. For this work,
the pump performance characterization was extended to simulate
in vivo chronic conditions, including those relevant to future studies to deliver agents to combat radiation resistance in a mouse
xenograft cancer model. A redesigned drug reservoir achieved amore compact footprint by reducing dead volume (inaccessibledrug) by nearly 91% and check valve dimensions, which were essential to accommodate the addition of wireless circuitry. Integrated
reﬁll ports allow long term use of the micropump for chronic studies.
2. Design
Our micropump system utilizes a low power electrolysis actuation mechanism. The electrochemical actuator consists of anelectrolyte (water) encased by a Parylene bellows and a pair ofinterdigitated platinum (Pt) electrodes supported on a rigid glass
substrate. An applied electrical current to the electrodes dissociates
liquid water into oxygen and hydrogen gas and induces an increasein pressure. This deﬂects the bellows, activates a one-way checkvalve, and displaces the surrounding ﬂuid out of the rigid reservoir
through the outlet catheter to the delivery site. The gases recombine to water once the current is turned off, allowing the bellows
to return to the original rest position.
The interdigitated electrode design reduces the solution resistive path resulting in improved electrolysis reaction [23]. Pt was
selected as the electrode material due to its biocompatibility, good
corrosion and oxidation resistance, and ability to catalyze recombination [15,14]. Electrolysis efﬁciency was further improved bycoating the electrodes with Naﬁon® , a solid polymer electrolytethat provides higher solubility of oxygen and hydrogen gases thanwater [33]. Naﬁon® -coated electrodes also provided faster andmore repeatable recombination, decreased electrode delamination
or damage at high current operation, and increased ﬂow rate [34].
Bellows possess the ability to achieve large deﬂection with low
applied pressure compared to corrugated or ﬂat diagrams [23].
Parylene C was selected as the bellows material due to its low
Young’s modulus, biocompatibility, inertness and low permeability[22,23,12]. The bellows consisted of two convolutions each having
a 6 mm inner diameter, 9.5 mm outer diameter, 0.4 mm height and13.5 ␮m thickness. This bellows can safely displace 185 ␮L of ﬂuid
contained within the adjacent drug reservoir without experiencing
plastic deformation [7].

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A. Cobo et al. / Sensors and Actuators A 239 (2016) 18–25

Fig. 1. Schematic diagram of in vivo testing setup for small animal research with (a) commercial tethered infusion system, and (b) wireless implantable drug delivery pump.

The micropump design achieves compact footprint by reducing
dead volume between the reservoir wall and bellows. The reservoir
inner diameter was reduced by 46% resulting in a substantial reduction in dead volume from approximately 1100 to 100 ␮L compared
to previous micropump prototypes [9]; this new volume combinedwith the capability for reﬁll is adequate for most small animal drugdelivery applications. The reduction in reservoir footprint allowed
incorporation of the wireless components while maintaining a total
device mass of approximately 10% of the weight of the animal (typically adult mice are 30 g) to allow free, unhindered movement[39].
The pump is normally off until wirelessly activated by a class E
inductive powering system (Fig. 2). Inductive transmission is mostsuitable for transcutaneous powering compared to other formsof wireless powering (radiation, conductive, and capacitive) [20].
For this system the power transmission frequency was chosen tominimize heat generation [13]. It relies on magnetic ﬁeld coupling and transmission frequency can be chosen to minimize power
dissipation. The class E inductive powering transmitter is placed
underneath the rodent cage. A current regulator on the subcutaneously implanted receiver controls the application of current to
the electrolysis actuator; here, a single constant current magnitude was supplied corresponding to a single ﬂow rate. The volumedelivered is controlled by the duration of pump activation. Heatgeneration from the wireless system is negligible in part due tochosen operation frequency [36]. High efﬁciency off the shelf components were chosen to minimize power dissipation [42].
3. Fabrication
3.1. Electrolysis actuator
The actuator’s interdigitated electrodes (100 ␮m wide elements
separated by 100 ␮m gaps, 8 mm diameter footprint, 300 Å/2000 Å
Ti/Pt) were fabricated on a Boroﬂoat 33 glass substrate (University
Wafer, Boston, MA) by liftoff following e-beam deposition of thethin ﬁlm metal using a previous method described in [8,34]. The
electrodes were potentiostatically cleaned at ±0.5 V (Gamry Reference 600 Potentiostat, Warminster, PA) in 1× phosphate bufferedsaline (PBS). The electrodes were then dip coated in Naﬁon®
(Dupont DE521 Solution, Ion Power, Inc., New Castle, DE). KynarTM
nickel plated copper wires (30 AWG, Jameco Electronics, Belmont,CA) were soldered to the electrode contact pads. The joint was reinforced and insulated with nonconductive marine epoxy (Loctite,
Westlake, OH) [34].
Parylene bellows were fabricated using a lost-wax two-partmolding process previously described in [7]. Polydimethylsiloxane (PDMS, Sylgard 184; Dow Corning Corp., Midland, MI) sheets(0.4 mm thick) were perforated with 6 mm and 9.5 mm holes tomold the inner and outer diameters of the bellows, respectively. The
sheets were visually aligned and stacked together on glass slides toform mold modules. These modules were ﬁlled with molten (50 ◦ C)
polyethylene glycol (PEG; 1000 Mn, Sigma–Alrich, St. Louis, MO).

Once the PEG solidiﬁed, PDMS sheets were peeled. The modules
were stacked and fused together by moistening connecting interfaces with water. Completed PEG molds were coated with 13.5 ␮m
United States Pharmacopeia (USP) class VI Parylene C (Specialty
Coating Systems, Indianapolis, IN), followed by soaking in deionized water at room temperature to remove sacriﬁcial PEG.
The Parylene C bellows were ﬁlled with double deionized waterand attached to the electrode using double-sided pressure sensitive adhesive ﬁlm (Tape 415, 3M, St. Paul, MN). The joint was thenreinforced with marine epoxy.
3.2. Check valve
A one-way duckbill valve made of medical grade silicone(2.0 mm duckbill; Minivalve, Cleveland, OH) with a measured
cracking pressure of less than 0.69 kPa (5.17 mmHg) was selected.
Valves were prescreened for proper closure and then incorporated
into the pump by attaching the base of the valves to the reservoir
housing output path using biocompatible epoxy (EPO-TEK® 730
unﬁlled, Epoxy Technologies, Billerica, MA).
3.3. Wireless powering system
A 2 MHz clock oscillator (ECS-2100, ECS International, Olathe,KS) was used to create the power signal from a 9 V power supply (0.3 W). The generated signal was then ampliﬁed in two stages
before being applied to a tuned transmitting coil (8 turns of 20 AWG
single strand wire, size: 310 mm × 140 mm). Litz wire (6 turns,
50/54 SPN/SN Litz Wire, Wiretron, Volcano, CA) was used for thereceiving coil (Ø 17 mm). The received alternating signal was fullyrectiﬁed using two Schottky diodes (BAT54A and BAT54C, Fairchild
Semiconductor, San Jose, CA). A resistor was used to set the output
current on a 3-terminal adjustable current source (LM334, Texas
Instruments, Dallas, TX) to 0.33 mA for driving the pump.
3.4. System packaging and assembly
The reservoir domes and bases were produced using stereolithography (SL) in USP class VI DSM Somos® WaterShed XC 11122
material (FineLine Prototyping, Inc., Raleigh, NC). This material is
optically clear, durable, water resistant, and rigid [38,6].
Medical grade silicone rubber septa (10:1 base-to-curing agent
ratio USP class VI MDX-4 4210; Factor II, Lakeside, AZ) were fabricated based on results of a leak puncture test allowing for at least
one dozen punctures with a maximum applied pressure within
the dome of 775 mmHg (15 psi). The septa were then inserted into
the reﬁll ports of the SL dome and the perimeters reinforced withmarine epoxy. This assembly was encapsulated with 5 ␮m Parylene
C to further improve biocompatibility and barrier properties. Thenthe electrolysis actuator was situated in the reservoir assembly andsecured with biocompatible epoxy.
The wireless receiver was mounted on the micropump body andencapsulated with 5 ␮m Parylene C, followed by a thin coating of

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21

Fig. 2. Circuit schematic for the class E wireless inductive powering system. Link efﬁciency for concentric transmitting and receiving coils with zero angular misalignment
was calculated to be approximately 0.82.

Fig. 3. Schematic diagram of (a) drug delivery system, and (b) fully assembled wireless micropump.

Fig. 4. Experimental setup for ﬂow rate characterization of wireless micropumps.

medical grade silicone to provide a softer interface with tissue. The
total mass of a ﬁlled prototype pump was 3.8 g (Fig. 3).
4. Experimental methods
Our speciﬁc application of interest is a micropump for asiRNA-based therapy against radiation resistant cancer in which
an electrostatically-coupled siRNA-gold nanorod drug is directly
delivered to xenograft tumors in nude mice. This anti-cancer application requires low ﬂow rate (∼␮L/min), 30 ␮L daily dose volume,and infusion duration of at least two weeks with one reﬁll per week.
The micropump’s ability to administer the desired regimen under
wireless operation was investigated on benchtop. In addition, therepeatability of performance across different wireless receivers andimpact on performance of changes in coil orientation were examined. Power transfer can be affected by coil orientation changes.
The wireless micropump is intended to be implanted in a freely
moving animal. Therefore, it is important to investigate the effectsof varying the axial separation distance and angular misalignment
between the transmitting and receiving coils that can result overthe course of drug administration in an animal study. Previously,
we presented preliminary results of successful micropump performance in simulated in vivo conditions including delivery against aphysiologically relevant back pressure and delivering solutions ofdiffering viscosities [5]. Here, we extend this simulated in vivo studies to analyze micropump performance when delivering at bodytemperature and long term pump functionality when subjected toa simulated in vivo environment (saline soak at body temperature).
4.1. Wireless testing
Flow rate measurements were taken on a single micropumpwith ﬁve different wireless receivers. Flow rate was calculated
by measuring ﬂuid front displacement in a 100 ␮L calibrated
micropipette (VWR International, Radnor, PA) connected to thecatheter outlet as shown in Fig. 4. The axial separation distance(parallel coils, 2–4 cm, 0.5 cm steps) and angular misalignment (0,30, 45, and 60◦ ) between concentric transmitter and receiver coils(2 cm distance between coils) was varied to study the effects onpump ﬂow rate performance.

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A. Cobo et al. / Sensors and Actuators A 239 (2016) 18–25

Fig. 5. The effects of (a) axial separation distance between concentric transmitter and receiver coils, and (b) angular misalignment between concentric coils on pump
performance (at 0.33 mA constant current).

Fig. 6. Representative data from one stationary micropump wirelessly operated. (a) Bolus delivery from a micropump operated once a day during week two (at 0.33 mA
constant current) and (b) ﬂow rates for the three weeks dosing.

4.2. Daily dosing
Successful dose delivery of six stationary (2 cm distance and0◦ angular misalignment between concentric coils) micropumps
over a period of one week was obtained. Here, two of the previous
micropumps were operated daily for an additional week to simulate the two week dosing regimen planned for the anti-cancer
application. Daily dosing regimen over a period of three weeks wasstudied in one of said micropumps. A constant current of 0.33 mA
was supplied to the actuator wirelessly for a duration that allowed
30 ␮L of double deionized water to be delivered daily. The pumps
were only reﬁlled once per week and the duration of applied current
was increased with each successive dosing event. The micropump
catheter was connected to a calibrated micropipette for ﬂow ratemeasurements as previously described. The ﬂuid front position wasrecorded several times during each delivery at discrete time points.
Reverse leakage between dosing events, if any, was recorded. To
avoid confounding results from evaporation, the ﬂuid front being
tracked was followed by a small air pocket (10 mm), followed by acolumn of water resulting in negligible evaporation between dosing
events.

4.3. Room temperature versus body temperature
In implantable drug delivery devices, the device will be
operated at body temperature. Three fully assembled and encapsulated micropumps were ﬁlled with double deionized water,
ﬂush of any air bubbles and soaked in a room (23 ◦ C) or bodytemperature (37 ◦ C) deionized water bath for 30 minutes. Flowrate test was performed while the pump was submerged in atemperature-controlled water bath. The ﬂow rate was determined
as previously described during four cycles each lasting 5 minutes
after valve opening.

4.4. Physiological enviroment simulation
System performance in simulated in vivo conditions was investigated to verify that Parylene C and silicone rubber encapsulation
layers provide adequate protection for the electronics for the duration of the anticipated in vivo study. Three fully assembled andencapsulated micropumps were soaked in 1× PBS at body temperature (37 ◦ C). Periodically, the micropumps were removed fromthe bath, reﬁlled with double deionized water, and subjected to
ﬂow rate testing. Valve opening times were recorded and ﬂow ratemeasurements were taken for 5 minutes after valve opening.

5. Experimental results
5.1. Wireless testing
Benchtop testing of a single micropump with ﬁve different
receivers resulted on an average ﬂow rate of 2.66 ± 0.16 ␮L/min(mean ± SE, n = 5), a 6% variation between receivers. The resultsfrom changing the axial separation distance between coils andangular misalignment between transmitter and receiver showeda drop in ﬂow rate by 64.1% and 42.86% for a 3.5 cm separation
distance between coils and for 45◦ angular misalignment between
transmitter and receiver, respectively as shown on Fig. 5.

A. Cobo et al. / Sensors and Actuators A 239 (2016) 18–25

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Fig. 7. Representative data from one wireless micropump ﬂow rate performance
when increasing the environmental temperature from room (23 ◦ C) to body temperature (37 ◦ C) (at 0.33 mA constant current).

Fig. 8. Representative ﬂow rate performance results for a micropump under simulated in vivo conditions for 30 days (at 0.33 mA constant current).

5.2. Daily dosing

after a high number of punctures under high reservoir pressure.
The new redesigned rigid reservoir signiﬁcantly decreased device
footprint by decreasing dead volume by 91%. The total ﬁlled encapsulated device mass including wireless power components was
3.8 g, slightly exceeding the mass requirements for a mouse experiment (3 g or approximately 10% of an adult mouse weight of30 g). The encapsulation was identiﬁed as the heaviest component.
Therefore, the pump mass can be reduced by eliminating the encapsulation material and instead using a low density polymer outercasing to enclose the current system.
Fully integrated wireless stationary micropumps successfully
delivered 30 ␮L daily boluses at low ﬂow rates (∼␮L/min) for several weeks, a regimen relevant for the administration of siRNAtherapy to a radiation resistance cancer. The 2.8 fold increase in
ﬂow rate observed during the third week of the daily dosing study
was attributed to a faulty receiver. Due to this systematic error,
ﬂow rate variation was calculated after the receiver was replaced(4%). Since we previously demonstrated consistent actuator performance [33], slight variations in mean ﬂow rate were attributed tovariations due to manual assembly of the packaging and receivercoils and inconsistent performance in the commercially produced
valves. Also, valves permitted up to 33% of dose volume in back ﬂow
when the pump was turned off. Some delay in forward ﬂow during current application was present in all micropumps tested and isdue to the low current application [29], and valve cracking pressure.
This delay was further increased in daily dosing regimens where thereservoir was not reﬁlled and higher pressure build up was requiredin order to deliver the same dose volume [33]. As discussed in [11],the forward ﬂow delay prior to reﬁll follows a predictable patternand can be accounted for by increasing the current application time,thus our ﬂow rate variation of 4% does not include the forward
ﬂow delay. The forward ﬂow delay in our previous prototypes was
minimized by replacing the compliant reservoir material with arigid material and utilizing higher efﬁciency electrolysis electrodes[34]. The dosing regimen can be altered depending on the intended
application. A diverse assortment of liquid drug formulations canbe delivered within a wide dynamic range of dose volumes (nL–␮L)and ﬂow rates (0.33–141.9 ␮L/min) [33]. The bellows can safely
deliver 180 ␮L. Once this volume is reached the actuator shouldbe allowed to fully recombine prior to the next dosing event. Fullrecombination is achieved within several hours [35]. While thepump can store medication for long term, it is important for the
drug to be stable when stored at body temperature between reﬁll
events.
Previous prototypes of this pump included commercial valves
with high sealing pressures [8] that showed signiﬁcant reverse ﬂow
during recombination. For in vivo studies, reverse ﬂow needs to be
avoided in order to obtain safe and accurate dosing. The check valve

Two wirelessly operated micropumps consistently delivered
30 ␮L doses daily over the course of two weeks as required by theanti-cancer application. These micropumps showed similar ﬂow
rate performance during the two week testing with less than 8%
variation. Fig. 6a shows representative results from the daily dosing regimen of one micropump for one week. The data is presented
as accumulated volume to study the effects of increasing pressurebuild up within the reservoir during recombination periods. Similar data was obtained for all pumps tested. The shown micropump
delivered 30 ␮L doses daily for three consecutive weeks. Mean ﬂow
rates during the ﬁrst two weeks are not signiﬁcantly different as
conﬁrmed by statistical analysis (ANOVA, p > 0.05) (Fig. 6b). Valve
performance varied across the micropumps, but it consistently limited back ﬂow to a maximum of 10 ␮L when the pump was turned
off. The slight volume drop seen before each dosing event correlates to the back ﬂow allowed by the valve. Some valves showed
improved sealing performance in each consecutive week of testing.
5.3. Room temperature versus body temperature
Fig. 7 shows representative results of the ﬂow rate performanceat room temperature (23 ◦ C) and body temperature (37 ◦ C) for one
micropump (mean ± SE, n = 4). As expected, an increase in ﬂow rate
was apparent at body temperature. For two micropumps, changes
in temperature resulted in signiﬁcant ﬂow rate differences as conﬁrmed by performing a one way analysis of variance (ANOVA,
p < 0.05).
5.4. Physiological enviroment simulation
A 30 day trial with three fully encapsulated and packaged pumps
immersed in a 1× PBS bath at body temperature (37 ◦ C), showed
continued accurate and reliable ﬂow rate performance with a standard error of less than 4% for each pump. Fig. 8 shows representative
results of the ﬂow rate performance for one micropump (mean ± SE,
n = 4). This micropump was previously used for the three week daily
dosing study, but has a new receiver. Mean ﬂow rates between
weeks for each pump are not signiﬁcantly different as conﬁrmed
by statistical analysis (ANOVA, p > 0.05).
6. Discussion
The pump components were redesigned to optimize pump performance and ease of use. The new reservoir dome features two
reﬁll ports as oppose to one to facilitate ﬁlling and ﬂushing of
the reservoir. The reﬁll ports are robust enough to prevent leaks

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incorporated into this system was the only small commercial valve
with low sealing pressure and cracking pressure. Unfortunately, thevalve manufacturing process leads to variability and inconsistency
in valve performance. Thus, valves were screened before incorporation into the pump and only 50% of the valves passed screening(reverse leakage <20% of dose volume). Variable and inconsistent
performances of the valve resulted in additional variability among
pump performance. Therefore, high quality valves with repeatableperformance are now being developed.
A class D wireless power system was developed to power anelectrochemical actuator in Sheybani et al. [8]. While it was simple in design and provided large currents, this class D system had
several drawbacks including difﬁculty in ﬁne tuning the ampliﬁer
and inability to generate the required power without overheating
the transmitter circuit components. A class E system combines theadvantages of the class D system with its series-tuned output andparallel-tuned ampliﬁer tanks. Class E systems can cope with the
rigorous demands of weakly coupled links, produce smooth signals,
be insensitive to low switching time, and operate over several tensof MHz. This pump was powered with a class E powering system,
which allows higher efﬁcacy when a considerable amount of power
is to be transferred [30]. Pump ﬂow rate is linearly dependent on
applied current to the actuator [34]. A current regulator was chosen as part of the implanted receiver circuit to ensure the current
applied to the actuator to be relatively constant (within certain limits of distance and angular misalignment between the transmitting
and receiving coils) regardless of slight changes in the power. Current regulators are inherently lossy components, as such, power
efﬁciency is partially sacriﬁced to ensure proper pump operation.
Flow rate performance for an individual pump varied when it was
powered with different receivers. Variations in ﬂow rate remained
within less than 6%. This slight variation in performance can be
attributed to the manual assembly of receiver coils, but can be alleviated by calibrating each pump prior to in vivo studies. As expected,
a decrease in power transferred due to increased axial separation
distance and angular misalignment between concentric transmitter and receiver coils resulted in decreased in ﬂow rate. Pump ﬂow
rate stability can be further improved by increasing the transmitted
power [20]. The addition of multiple coils at either the transmitter or receiver side can improve power transfer to mitigate angle
misalignment. Another potential solution is to move the external
receiving coil along with the moving animal by tracking the movement of the internal transmitting coil with a similar setup to theone proposed by Kilinc et al. [18].
The effects of environmental factors on pump performance were
investigated. Changing the environmental temperature from room
temperature (23 ◦ C) to body temperature (37 ◦ C) for our wirelessmicropumps resulted in increased ﬂow rate. Improved electrolysis efﬁciency can be attributed to the increased permeability ofNaﬁon® to hydrogen and oxygen when the ambient temperature
is increased [2]. This is in agreement with [33], which found slightincrease in ﬂow rate at body temperature for electrochemical actuators. Physiological environmental simulation test showed that all
three pumps were functioning consistently and as expected after
30 days of soaking in 1× PBS at body temperature (37 ◦ C) exceeding the intended in vivo study duration by 100%. This veriﬁes that
the Parylene C and silicone rubber encapsulation layers provide
adequate protection for the electronics for the duration of the anticipated in vivo study. This packaging scheme may also be suitable forlonger term in vivo studies beyond 30 days.

7. Conclusion
We demonstrated the characterization of a fully implantablewireless infusion micropump for use in chronic drug delivery

therapies in small research animals. Wireless control of operation
permits a user to select the desired delivery regimen in a freely
moving animal. Wireless powering of the actuator with a class E
inductive powering system allowed repeatable delivery with lessthan 6% variation in ﬂow rate. Pump performance was affectedby increased axial separation distance and angular misalignment
between concentric transmitter and receiver coils. If necessary,
this can be alleviated by increasing the power transmitted andinclusion of a second receiver coil [20]. Wireless powering with
an implantable device requires inductive coupling link through theskin and biological tissue. As such, it is essential to characterizedevice performance in vivo. Future work will include characterization of system operation in simulated tissue material to mimic
device operation in animals.
Stationary micropumps successfully delivered a single bolus
daily for several weeks and achieved ﬂow rates on the order of␮L/min, an experimental regimen relevant for an anti-cancer drugdelivery application. Valves signiﬁcantly decreased reverse ﬂow
compared to a valve less system. We demonstrated that ﬂow rate
was not signiﬁcantly affected by back pressures of up to 20 mmHg
(when delivering solutions of 1 cP) or delivering solutions with viscosities of up to 6 cP (data shown in [5]). As expected, micropump
ﬂow rate increased with increased environmental temperatures.
Consistent micropump performance was also demonstrated in simulated physiological environmental conditions with less than 4%
ﬂow rate variation.
Drug delivery micropumps need to incorporate closed-loop
feedback systems to enable pump performance monitoring, andincrease effectiveness of therapies. Control over delivery proﬁle canbe achieved through the incorporation of physical sensors that provide information on pressure, ﬂow rate, delivered volumes, forward
ﬂow delay and real-time state of the pump. A variety of sensing
techniques are available for MEMS ﬂow sensors, such as electrical,
nuclear, thermal, mechanical, optical, and acoustic. We are currently developing an electrochemical impedance (EI) dose tracking
system capable of tracking accidental release of reservoir content,
blockage and reﬁll detection or lack of delivery in real time [32].

Acknowledgements
This work was funded 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 ofEngineering Ph.D. Merit Fellowship (AC). The authors would like tothank the members of the USC Biomedical Microsystems Laboratory for their assistance with this project.
E. M. has a signiﬁcant interest in Fluid Synchrony LLC (CA, USA).

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Biographies

Angelica M. Cobo received her B.S. degree in Biomedical
Engineering from Florida International University in 2008.
She was a research assistant in the Bioengineering Laboratory at the Diabetes Research Institute where she focusedon the development of a bioactive scaffold material foruse in cellular replacement therapies to treat Type 1 Diabetes Mellitus. In 2012, she received her M.S. in Biomedical
Engineering from University of Southern California (USC).
She is currently a PhD student in Biomedical Engineering at USC where she is working on the development of
implantable MEMS drug delivery devices for managementof chronic diseases in preclinical research.
Roya Sheybani received her B.S. (2008), M.S. (2009), PhD(2015) degrees in biomedical engineering from the University of Southern California (USC), Los Angeles, where
she focused on developing closed-loop implantable wireless MEMS drug delivery devices for management ofchronic diseases. She is currently a Hibbitt Engineering Postdoctoral fellow at Brown University, working
towards developing “smart” biomedical devices that combine designer drug eluting biomaterials with biosensors toenable closed-loop therapies.

Heidi Tu received her B.S. in Biomedical Engineering and
B.A. in German from The University of Texas at Austin.
Heidi has a Ph.D. in Biomedical Engineering from the University of Southern California, where she developed an
implantable drug delivery pump. She is currently a Product Design Engineer at a private healthcare technology
company in the San Francisco Bay Area. Her professional
goal is to develop medical technology to enable better
treatment and care for patients.

Ellis Meng received her B.S. degree in engineering andapplied science and her M.S. and Ph.D. degrees in electrical engineering from the California Institute of Technology(Caltech), Pasadena, in 1997, 1998, and 2003, respectively.
She is a Professor of Biomedical Engineering and Electrical Engineering at the University of Southern California,
Los Angeles where she has been since 2004. She directs
the Biomedical Microsystems Laboratory which specializes in biomicroelectromechanical systems, implantablebiomedical microdevices, neural interfaces, and microﬂuidics. She is a recipient of the NSF CAREER, the Wallace H.
Coulter Foundation Early Career Translational Research,and the ASEE Curtis W. McGraw Research Awards. In 2009,
she was recognized as one of the TR35 Technology Review Young Innovators under
35. E. M. has a signiﬁcant ﬁnancial interest in Fluid Synchrony LLC (CA, USA).