On-demand Wireless Infusion Rate Control in an Implantable
Micropump for Patient-tailored Treatment of Chronic Conditions
Roya Sheybani-EMBS Student Member, and Ellis Meng, IEEE Senior Member

Abstract—
Wirelessinfusion
ratecontrolandprogrammabilityfor
animplantable,
lowpower,electrochemical micropump is presented. Flow rate control wasachieved through adjustment of the wiper position of a current
potentiometer in the wireless receiver (0.6-3.2 mA output
current with a resolution of 0.2 mA per step). An off-the-shelf
Bluetooth module and Basic Stamp microcontroller kit wasused to initiate amplitude-shift keying (ASK) modulation of theinductive power signal. Accurate flow control of two model
regimens was achieved on benchtop. Wireless transmission(power transfer and control) was not affected by simulated
tissue material placed between the transmitter and receiver.

I. INTRODUCTION
Effective drug therapy is essential in the treatment andmanagement of chronic conditions such as hypertension,
respiratory disease, and diabetes [1]. As of 2005, 46% of
American adults suffer from at least one chronic illness and75% of health care dollars are spent on the treatment ofchronic diseases [2]. Implantable drug delivery devices may
improve treatment of chronic conditions by bypassing
physiological barriers to enable delivery of new compounds
such as biologics, biosimilars, and other small molecules
directly to the target tissue [3]; and maximizing therapeutic
efficacy of the drug and limiting toxic side effects bydelivering the correct amount of drug in the vicinity of thetarget cells, while reducing the drug exposure to the nontarget cells [4]. Most commercial implantable pumps only
provide constant infusion rate. Certain chronic conditions,such as drug addiction and menopause, have associatedstatic therapeutic windows and may be treated with
continuous flow drug administration [5], however, many
chronic conditions have chronobiological pattern in theirpathogenesis [6-9]. More importantly, drug dosing affects
the therapeutic window. Following successful drug
administration, the clinical need for additional
pharmacological intervention for certain conditions
diminishes transiently while the ratio of risk to benefit
dramatically increases [10]. Controlled patient-tailored drug
administration avoids the peak and trough time course ofdrug concentrations in the plasma between successive doses
and the corresponding peak and trough pattern of drug
action, leading to therapies that mimic chronobiologicalpatterns [3, 11, 12].
Implantablepumps are advantageous for chronic*Research supported in part by NSF through the PFI: AIR
Technology Translation program (award number IIP-1343467) andNIH (award number R21GM104583) funding sources.
E. Meng, and R. Sheybani are with the Department of Biomedical
Engineering, Viterbi School of Engineering, University of Southern
California, Los Angeles, CA 90089-1111 USA (corresponding author:213-821-3949, e-mail: ellis.meng@usc.edu).

treatment; eliminating transcutaneous wires and catheters
permits improved patient mobility and allows drug
administration outside of clinical settings [13]. Despiteadvances in miniaturization, batteries are still large incomparison to microelectromechanical systems (MEMS)
pump actuators and can significantly increase device size.
Also, batteries have limited lifetime and may pose healthrisks (leakage or malfunction) [14]. Wireless powering andcontrol of drug delivery devices eliminates the battery,
reduces device dimensions, and can improve system lifetime.
A wired programmable and implantable drug infusion
micropump system, capable of delivering a diverse
assortment of liquid drug formulations with high accuracy
within a wide dynamic range of dose volumes and flow rates,
as well as wireless powering at a single flow rate, werepreviously demonstrated by our group [15, 16]. Here, wepresent, for the first time, a completely wireless infusion
micropump system with flow rate adjustability capable ofachieving patient-tailored therapy. Our research prototype iscurrently scaled for small animal (rodents) research (Fig. 1).
The animal with the implanted micropump system is placed
inside a cage set atop the transmitting coil. The user-defineddelivery parameters (duration and infusion rate) are enteredon-demand or stored as a program on a laptop that
communicates via Bluetooth with the transmitter circuit tocontrol operation of the implanted pump.

Figure 1: Schematic diagram of system setup for small animal research.

II. DESIGN AND FABRICATION
A. System Requirements
The micropump consists of an electrolysis driven actuatorhoused in a drug reservoir. Interdigitated Nafion®-coatedplatinum (Pt) electrodes electrolyze the water in the actuationchamber into hydrogen and oxygen upon electric current
application. The resulting volume expansion is harnessed toinflate a Parylene bellows and expel fluid in the adjacent drugchamber through a catheter to the target site [16]. The
current-controlled actuator requires 0.66-51.31 mW of power(for 0.1-13 mA applied current) to achieve a wide range offlow rates (0.33 - 141.9 µL/min, respectively). Here, we limit

the flow rate range to 2.0 - 25.0 µL/min (0.6-3.2 mA appliedcurrent) which is suitable for most small animal research
applications. The overall device footprint should also be
minimized to allow for implantation in mice (Ø < 22 mm,height < 10 mm). Wireless inductive powering was chosen tosatisfy the low power and small footprint requirements, aswell to maximize the usable lifetime of the device.
Under constant current conditions, production and masstransfer of electrolysis gases is proportional to the square rootof time [17-19]. Thus, constant current is the preferred modefor electrolysis actuation, producing an infusion rate linearly
proportional to current. For a single infusion rate, the powerreceived is rectified and applied to a current regulator toprovide a single constant current to the actuator. Variablecurrent (and therefore variable infusion rates) could beachieved by changing the wiper position of a current setting
programmable digital potentiometer in the receiver circuit toalter the output current. The potentiometer should be chosen
to allow for the selection of the full range of desired currents(0.6 -3.2 mA) with a resolution of 0.2 mA per step. The wiper
position could be changed using a data signal wirelesslytransferred through the skin to the receiver. Two methods ofdata transmission were considered for this purpose: infraredand ASK. ASK was chosen due to its simplicity, lower
number of additional components on the transmitter andreceiver circuits, and lower susceptibility to environmental
noise.
B. System Architecture
Using the ASK data transmission method, a low
frequency data signal is carried by the power signal andsubsequently demodulated on the receiver. Pulse width of thedata signal determines the number of incremental steps ofpotentiometer wiper position, altering the potentiometer
resistance, and changing the output current set by thepotentiometer. The transmitter output is controlled by theEasy Bluetooth module and Basic Stamp microcontroller kit(Parallax Inc., Rocklin, CA) and custom user interface
program using the Basic Stamp editor. The overall systemarchitecture is presented in Fig. 2.

Figure 2: System architecture of ASK modulation circuit to enable wirelessflow control.

C. Circuit Design and Layouta)Transmitter

A 2MHz clock oscillator (ECS -2100, ECS international,Olathe, KS) was used to create the power signal. A quadbilateral switch (CD4016BC, Fairchild Semiconductor, SanJose CA) controlled by the Easy Bluetooth module and Basic
Stamp microcontroller kit (Parallax Inc., Rocklin, CA)
switched between the power signal or the modulation signal.
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 x 140 mm).
b)

ReceiverLitz wire (6 turns, 50/54 SPN/SN Litz Wire, Wiretron,Volcano, CA) was used for the receiving coil (Ø 22 mm).
The circuit components (including the current regulator)
require a direct current (DC) power signal. Therefore, thereceived alternating signal was fully rectified using two
Schottky diodes (BAT54A and BAT54C, Fairchild
Semiconductor, San Jose, CA). The modulation signal wasseparated by half-wave rectification (BAT54A) and thenfiltered using a low pass RC design. Two zener voltage
regulator diodes (BZV55, NXP Semiconductors, Eindhoven,Netherlands) regulate the output voltage of the power anddemodulated signal. AD5227 (Analog Devices, Norwood,MA) was used as the current setting potentiometer to set theoutput current of the current regulator (PSSI2021SAY, NXP
Semiconductors, Eindhoven, Netherlands). A low frequency

Figure 3: Schematic diagram of the receiver circuit.

oscillator (LTC 6991, Linear Technology, Milpitas, CA) set
the potentiometer frequency to 500 Hz. An LED was addedin series with the infusion pump to provide visualconfirmation of power supply to the actuator (Fig. 3). The
receiver was implemented on a flexible printed circuit board(PCB) to allow it to be wrapped around the roundmicropump. Small surface mount components were chosen
to minimize the overall circuit footprint (Fig. 4).

and 19.8 g, respectively). 1.98 g of ProClin® 950 preservative(Sigma-Aldrich, St. Louis, MO) was added as a replacementfor Dowicil 75® (1-(3-chloroallyl)-3, 5, 7-triazalazoniaadamantanechloride). A 4 cm thick slab of gel-likematerial was created for testing. The pump and circuit werethen placed on top of the simulated brain tissue. Differentcurrent values were applied using the Basic Stamp editor
program and the infusion flow rate was calculated bymeasuring the fluid front movement in a calibrated 100 µL
micropipette. The simulated tissue material was thenremoved and the experiment repeated with only air separating
the transmitter and receiver.
IV. RESULTS & DISCUSSION
A. Testing with “Dummy” LoadFigure 4: Mircopump with integrated receiver. Inset: Receiver flexible PCB,
front and back.

III. EXPERIMENTAL METHODSThe transmitter and receiver circuits were first tested
using discrete components on breadboards and then on PCBs.
Preliminary benchtop testing using a “dummy” 1 kΩ load inplace of an actuator was performed to ensure feasibility andrepeatability of the approach. The circuit performance wasthen tested with a micropump on benchtop. In order to study
the electric and magnetic effects of tissue on the wirelesstransmission, testing across simulated brain tissue material[20] was performed.
A. Testing with “Dummy” Load
The receiver current output was increased incrementally(10 steps) and the values achieved were compared totheoretical values calculated based on equations provided in
the potentiometer and current regulator datasheets. The
output current was calculated by measuring the voltage
across a 1 kΩ load. Repeatability of the output current for a
specific modulation program was also tested by varying thecurrent in 4 steps: 0.6, 2.6, 0.7, and 2 mA applied
successively for each run (n=5).
B. Testing with Infusion Pump
An infusion pump with a 1 mL reservoir [15] wasconnected to the receiver circuit (Fig. 4). Two different
regimens were selected to demonstrate wireless operation at
multiple flow rates. First, the current output was increased in0.5 mA steps from 0.6 - 2.1 mA. For the second regimen, thecurrent was increased from 0.6 to 1.2 mA, then decreased to
0.6 mA and increased to 2.2 mA. For both experiments,
double distilled (DD) water was used as the drug model, andthe infusion flow rate was calculated by measuring the fluid
front movement in a calibrated 100 µL micropipetteconnected to the outlet of the pump.
C. Testing in Simulated Brain Tissue Material
In order to mimic the effects of wireless transmission
through tissue, brain tissue was simulated according to therecipe described in [20]. Briefly, sugar, NaCl salt, andNatrosol® (hydroxyethylcellulose, Ashland Inc., Covington,KY) were dissolved in 800 mL of DD water (1108.9, 49.5,

The results for the incremental increase in receiver current
output are shown in Fig. 5. The measured value closely
followed the expected value for all the modification steps.
Repeatability in achieved current output for the same
modulation program was also confirmed (Fig. 6).
B. Testing with Infusion Pump
Fig. 7 shows the variable current results with the infusion
pump. The measured flow rate closely followed the changes
in current initated using the interface program. Measured
flow rate values were comparable to those achieved for
wired actuation of the infusion pump [15].

Figure 5: 10 step incremental increase in receiver current output measured
across a 1 kΩ resistor (at each modification step, a specific pulse width wasapplied to achieve the desired change in the wiper position; the values werecalculated based on equations provided in the potentiometer and current
regulator datasheets)

Figure 6: Current output repeatability for a specific modulation program.

C. Testing in Simulated Brain Tissue Material
Successful current modulation was achieved through thesimulated brain tissue material (data not shown; no
significant difference was observed between infusion flow
rates for wireless transmission through air vs. simulated
tissue; one way analysis of variance, p < 0.05). However, inboth cases, the achieved flow rate was smaller than thatobserved for when the pump was placed directly above thetransmitter coil (15% decrease). As previously mentioned,
flow rate is linearly dependent on applied current, andtherefore, dependent on received power. As a result,decreased power transfer due to coil distance leads todecreased flow rate. The decrease in power transfer with
distance between the coils can be mitigated by increasing thepower output of the transmitter [13].

increasing the power transmitted and inclusion of a second
receiver coil. Future work includes integration with anelectrochemically based dose tracking system capable ofreal-time tracking, confirmation of delivery [21].
ACKNOWLEDGMENT
The authors thank Dr. C. A. Gutierrez and the members of
the USC Biomedical Microsystems Laboratory for theirassistance.
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Figure 7: Wireless flow variation in (a) incremental steps, and (b) random
fashion.

V. CONCLUSION
A completely wireless infusion micropump system withflow rate adjustability (2.0 – 25.0 µL/min for 0.6-3.2 mA
applied current) capable of achieving patient-tailoredtherapy was presented. A decrease in power transmission
was observed when transmitter and receiver coil distanceand angle foveation were increased. This resulted in a flowrate decrease. Performance can be improved if necessary by

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