Conducting Polymer Based Active Catheter for Minimally Invasive
Interventions inside Arteries
Tina Shoa, John D. Madden, Niloofar Fekri, Nigel R. Munce, Victor XD. Yang
Abstract— An active catheter intended for controllable
intravascular maneuvers is presented and initial experimental
results are shown. A commercial catheter is coated with
polypyrrole and laser micromachined into electrodes, which
are electrochemically activated, leading to bending of the
catheter. The catheter’s electro-chemo-mechanical properties
are theoretically modeled to design the first prototype device,
and used to predict an optimal polypyrrole thickness for the
desired degree of bending within ~30 seconds. We compared
the experimental result of catheter bending to the theoretical
model with estimated electrochemical strain, showing
reasonable agreement. Finally, we used the model to design an
encapsulated catheter with polypyrrole actuation for improved
intravascular compatibility and performance.
ONDUCTING polymer (CP) actuators are a class of
“artificial muscle” actuators, with
characteristics which are advantageous for minimally
invasive surgical and diagnostic tools. Some of the
characteristics include low actuation voltage (<2 V), ease of
fabrication, relatively high strain (typically 2% or greater),
and biocompatibility [1]. One of the interesting applications
of these types of actuators is to convert passive catheters into
active ones, hence providing controllable catheter
manipulation inside the body [2]. Catheters are extensively
used in many medical procedures such as angiography, stent
deployment, intravascular ultrasound, coiling of cerebral
aneurysms, and treatment of thromboembolic diseases.
Catheters can be used to provide a channel for fluid passage
or an entry for a medical device. In angioplasty and
stenting, for instance, catheters are employed to guide a
therapeutic device to open a blockage inside a vessel.
Traditionally, guide wires manipulated external to the
Manuscript received April 7, 2008.
Tina Shoa is with Advanced Materials and Process Engineering
Laboratory at the University of British Columbia, Vancouver BC Canada.
(Corresponding author: phone: 604-657-8651 e-mail: [email protected]).
John D. Madden is with Advanced Materials and Process Engineering
Laboratory. He is an associate professor in the Dept. Electrical and
Computer Engineering of the University of British Columbia, Vancouver
BC Canada (email: [email protected]).
Niloofar Fekri is with Advanced Materials and Process Engineering
Laboratory at the University of British Columbia, Vancouver BC Canada.
Nigel R. Munce is with Imaging Research, Sunnybrook Health Science
Centre and Dept. of Medical Biophysics at the University of Toronto,
Toronto, Canada.
Victor X. D. Yang is with Imaging Research, Sunnybrook Health
Science Centre. He is an assistant professor in the Dept. Electrical and
Computer Engineering of Ryerson University, Toronto, Canada.
patient are used for guidance of the catheter by combinations
of push-pull and torque motions [3]. Limitations of the
current catheter and guidewire designs include long
procedural time, lumen or vessel wall damage, and the
subsequent medical complications. These issues become
more critical when dealing with narrow and complex
passages such as blood vessels of the brain and tertiary
bronchi of the lung.
Recently, advanced catheter designs exploit active tip
bending for more controllable and efficient minimally
invasive medical procedures. Although various active
catheters driven by shaped memory alloys (SMA) [4-7],
piezoelectric materials and MEMS based devices [8, 9] have
been presented, no active catheters are in wide spread use.
Conducting polymer actuators have shown attractive
properties, which make them promising to be employed
extensively in active catheter application. IPMC (ionic
polymer metal composites), another type of artificial muscle
actuators, has been used in steerable catheters [10-12]. These
actuators can generate large displacements at relatively low
voltages (<10 V) and are typically faster than conducting
polymers; however, their manufacturing process is often
relatively expensive, stresses are smaller and unlike
conducting polymers, additional energy is usually consumed
to hold the actuator in place. Conducting polymers offer
higher stiffness than IPMCs, an attribute which is often
important in catheter design [13]. Their efficiency can be
higher as maintaining deflection does not require energy
In this paper we investigate the feasibility of producing
small radius of bending curvature using polypyrrole
actuators in order to maneuver a catheter through highly
curved vessels. The goal is to access the lesions in narrow
and curved blood vessels where a conventional guide wire
cannot reach. In our approach a catheter with an active tip is
guidewire/catheter insertion techniques. The catheter tip is
then bent actively to access the lesion. A flexible wire is
subsequently inserted through the lumen of the catheter to
navigate through the lesion (See Fig. 1). The whole process
is monitored using x-ray angiography from the outside and
intravascular imaging from the inside simultaneously. The
same catheter can be used to inject the radio opaque x-ray
contrast agents.
upon reduction.
Vessel wall
Active catheter
Radius of curvature
Active Catheter
Flex wire
Fig. 1. Schematic of the device
The catheter used to demonstrate the bending is Micro
Therapeutics Inc. (Irvine, CA) UltraflowTM HPC (0.5 mm
OD/0.28 mm ID), which is a neurointerventional catheter
with a thin, flexible tip that allows maneuvering through
narrow and curved arteries. The tip of this catheter is coated
with conducting polymer. Coating the flexible tip of the
catheter with polypyrrole actuator substantially increases the
stiffness at the tip providing a sufficient axial rigidity to
prevent the catheter from buckling during insertion and yet
enabling passage through highly curved regions (Fig. 1).
Using our initial design, which is similar to a design
suggested by Mazzoldi and De Rossi [2], we demonstrate
that the polymer can provide sufficient force and
displacement to achieve the desired bending. We have
achieved a minimum radius of curvature of 9.8 mm in 30 s,
which is sufficient for maneuvering through most vessels’
branches (e.g. the smallest radius of curvature of the left
anterior descending coronary artery is 10 mm [14]). The
results of our experiments and modeling are presented here,
and a new encapsulated design is suggested which should be
able to be used in-vivo to achieve the desired bending in a
reasonable time.
The polypyrrole coated catheter is a trilayer structure,
where two polypyrrole layers are the electromechanically
active materials. Fig. 2a shows schematics of the structure
which consists of a flexible tube (i.e. catheter) with inner
and outer diameters of b=0.28 mm and a=0.5 mm. The
coated polypyrrole is divided into two segments that run the
length of the tip. As shown in Fig. 2b, application of a
voltage between opposing halves in an electrolyte leads to
deflection of the catheter towards one side or the other,
depending on polarity. The goal is to achieve a 10 mm
radius of curvature in less than 30 s.
When a positive potential is applied across the two
polypyrrole electrodes inside an aqueous solution of NaPF6
salt, the polypyrrole film on one side is oxidized and the one
on the other side is simultaneously reduced. During
oxidation mobile negative ions enter the polymer from a
surrounding electrolyte to balance charge. In this case
positive ions are considered to be too large to be transported
within the polymer. The insertion of anions from the
electrolyte generates a stress tending to expand the polymer
and produces a bending moment. The process reverses itself
Polymer Electrodes
Fig 2. (a) Schematics of the polymer coated catheter modeled as a trilayer,
(b) Actuation of the trilayer by applying voltage inside an electrolyte
Equation 1 describes the stress generated, σ, within the
oxidized polypyrrole;
σ =
yE p
− EP ε
where Ep is the Young’s modulus of the polypyrrole, R is the
radius of curvature; y is the distance from the center of the
trilayer structure and ε is the polypyrrole active strain that
would occur in response to ion flux under no load. To first
order the strain is proportional to the number of ions per unit
volume inserted, and equivalently the charge per unit
volume, ρ, via the relationship;
ε = αρ .
α is an empirically determined strain to volumetric charge
ratio [15-17]. In the same way, polypyrrole contraction
occurs during the reduction process because anions leaving
the polymer cause a stress gradient opposite to the one
described for the oxidation process (see Fig. 3). The net
torque leads to a bending of the structure which is opposed
by the stresses induced in the catheter walls. Considering
the force and torque balance to achieve zero bending
moment [18] at the tip of the structure (i.e. cantilever beam
condition), and assuming that the electrodes wrap around the
catheter almost completely, the radius of curvature of the
trilayer structure “R” and the total time of actuation “τ” are
found to be:
t 2p , (3)
3 Eb − E p (a + 2t p ) − Ea + E p a (2) and
E pε ( a )3 − (a + 2t p )3
Here a and b are the outer and inner diameters of the
catheter, tp represents the thickness and of the polypyrrole
electrodes, E and Ep are the Young’s moduli of the catheter
and the polymer respectively. τ represents the time of ion
diffusion through the thickness of the polymer, which is also
the catheter bending time. D is the effective diffusion
coefficient of ions and was estimated to be 5.5×10-11 m2/s
from the experiment. The Young’s modulus of an
electrochemically grown polypyrrole film and a
commercially available catheter were measured to be 300
and 75 MPa respectively. The free polypyrrole strain, ε, was
measured to be ~ 0.8% during actuation of films between
±0.2 V, increasing to 4% at ±0.8 V vs. Ag/AgCl reference
electrodes in aqueous 1 M NaPF6 electrolyte.
Removal of ions
Insertion of ions
Fig. 3. Stress distribution upon insertion and removal of ions from the
polypyrrole layer.
The response of the actuator to a step input of voltage was
measured to determine the minimum radius of curvature
achieved using the initial design and compared with model
predictions. According to the model a thickness of > 40 µm
is required to achieve the minimum radius of curvature on
the UltraflowTM HPC catheter used in this experiment.
Electroless deposition was first used to deposit an initial
layer of polypyrrole, onto which 40 µm thick polypyrrole
(doped with PF6) was electrodeposited. The polymer coating
was then divided into two electrodes as shown in Fig. 4.
A. Curvature improvement
Curvature can be increased by applying a higher voltage.
According to Equation 2 the radius of curvature, R, is
inversely proportional to the polymer electrochemically
induced strain. As was mentioned, strain is proportional to
the number of charges per unit volume inserted. In steady
state the polymer behaves like a capacitor or battery and a
capacitive model works quite well, where charge density is
expressed in terms of applied voltage, V, and the volumetric
capacitance of the polymer, Cv [21]. Hence strain can be
written as ε = αρ = α CvV . α was estimated (by measuring
the amount of charge transferred) to be 3.6 × 10 -11 C/m3 and
Cv to be 0.3 × 109 F/ m3. We increased the voltage, V, in
order to increase the strain (see Fig. 6) and decrease the
radius of curvature. Further increases in voltage were found
to lead to larger strains, but shorter lifetimes, and were thus
avoided [19].
Radius of Curvature (mm)
Discharging Charging
±0.5 V
±0.8 V
Strain (%)
Fig. 6. Experimental and simulated bending radius versus strain for 40 µm
thick polymer
Fig. 4. a) SEM image of the catheter coated with polypyrrole and
micromachined into two polymer electrodes.
The analytical model of Equation 2 is compared to
experimental results to verify that it provides a reasonable
description of the response, and then is used to predict
conditions under which small radius of curvature can be
achieved. In this experiment a radius of curvature of R= 16.8
mm was obtained in about 32 s by applying step potential of
±0.5 V to the polymer electrodes versus Ag/AgCl reference.
Fig.5 shows the initial and final positions of the catheter.
Fig.5. Experimental result showing the catheter at two states: initial, bent.
Fig. 6 shows the experimental and model-predicted values
of radius of curvature for different strains obtained using 40
µm thick polymer layers on the catheter shown in Fig. 2.
The model is an extension of Equation 2 in which the width
of the spacing between polypyorrole electrodes, 100 µm, is
also taken into account. As shown in this figure, a radius of
curvature of R=9.8 mm was achieved by applying ±0.8 V
across the two polymer electrodes. According to Equation 3,
the time needed for ions to diffuse through 40 µm thick
polymer is ~ 30 s which is in close agreement with the
measured value 32 s for total time of bending.
The presented results indicate that the desired degree of
bending is achievable using polypyrrole actuators and our
initial prototype served as a “proof of concept”. However,
electrochemical actuation of this device involves using
electrolytes and electrical currents which in most medical
applications mandate encapsulation. A more practical
design is presented in Fig. 7, where the active catheter is
encapsulated inside a tube. Four polypyrrole actuators are
considered in this design to provide bending moment in two
directions. The tube is slit into quadrants and polypyrrole is
deposited on the inside surfaces. The catheter coated with a
gel electrolyte is then embedded in the center of the tube
lumen. A medical adhesive such as Dymax “CTH” UV
curable catheter bonding adhesive will be used to attach the
quadrants together and the central catheter to the tube wall,
creating a completely encapsulated device that can be bent in
either of two directions. The model was modified for the
proposed geometry and suggested the following parameters
in order to achieve the bending radius of 10 mm: tube
Young’s modulus of 20 MPa with an outer and inner
diameter of 1mm and 0.8 mm respectively, electrolyte gel
Young’s modulus of 1MPa with a thickness of 0.3 mm,
catheter tip with a Young’s modulus of 75 MPa and an outer
and inner diameter of 0.5mm and 0.28 mm respectively
(UltraflowTM HPC). According to the model prediction a
minimum bending radius of 10 mm can be achieved in ~ 30
s with a polymer electrode thickness of 40 µm capable of
generating an electrochemical strain of 4%.
Encapsulation tube
Catheter tip
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Conducting polymer actuators are good candidates for use
in active catheters due to their biocompatibility, low cost,
large strain, low actuation voltage and ease of fabrication. In
addition these actuators offer high rigidity, which is often
important in catheter design.
In this paper the bending of a catheter using polypyrrole
actuators is demonstrated. The device was electrochemically
actuated in an aqueous solution of NaPF6 by applying a step
potential of between -0.8 V to +0.8 V and a bending radius
of 9.8 mm was achieved in 30 seconds. The predicted
bending radius shows relatively good agreement with
Since electrochemical actuation of these devices involves
using ionic electrolytes, encapsulation is often required. An
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