Download 5Electric Sensitive Release Systems

5
Electric Sensitive Release Systems
5.1 Introduction
Numerous synthetic polymers have been used to make electroactive hydrogels, such
as polyvinyl alcohol/poly(sodium maleate-co-sodium acrylate) [1], acrylic acid/vinyl
sulfonic acid copolymer [2] and sulfonated polystyrene [3]. In addition to synthetic
polymers, some natural polyelectrolytes have been blended with synthetic polymers
to prepare such hydrogels. For example, alginate/poly(methacrylic acid) [4], chitosan
(CS)/polyaniline [5] and hyaluronic acid/poly(vinyl alcohol) (PVA) [6] hydrogels have
also been reported in the literature. However, few reports were found on comparing
pure natural polymer-based electroactive hydrogels with those made from synthetic
and synthetic/natural polymer blend materials. It is well accepted that natural polymers
have better biocompatibility and less latent toxic effect than most synthetic polymers
[7–9], so pure natural polymer-based hydrogels would be more suitable for application
in biomedical fields. Electrically responsive delivery systems are prepared from
polyelectrolytes (polymers that contain a relatively high concentration of ionisable
groups along the backbone chain) and are thus pH-responsive as well. Under the
influence of an electric field, electroresponsive hydrogels generally deswell or bend,
depending on the shape and orientation of the gel. The gel bends when it is parallel to
the electrodes, whereas deswelling occurs when the hydrogel lies perpendicular to the
electrodes. Several technologies are currently under investigation for the development
of new drug-delivery devices that can achieve chronotherapy, specifically iontophoresis,
infusion pumps and sonophoresis [10, 11].
In particular, electric-sensitive hydrogels with biocompatibility and biostability are
more attractive as they are able to implement an isothermal energy conversion from
chemical free energy directly to mechanical work, actuated by an external electric
stimulus. As such, the mechanical energy is triggered by an electric signal. This results
in one of their important functions as biosensors/actuators, in which the electricstimulus responsive hydrogels with fixed-charge groups can bend reversibly when they
are subjected to an externally applied electric field [12]. Usually the electric-sensitive
hydrogels are composed of electrolytes and a swellable and insoluble crosslinked
polymer network with fixed-charge groups, as shown in Figure 5.1. When the
hydrogels are immersed into a bathing solution under an externally applied electric
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
field, the hydrogels deform. They swell and bend due to the pressure difference of
both hydrostatic and osmotic pressures, which are caused by the ionic concentration
difference between the internal hydrogels and the external solutions.
+
+
Mobile Ion
–
Fluid Filled
Region
Fixed Charge
–
–
+
+
–
–
–
Polymer Chain
+
Crosslink
Undissociated
+
Ionizable Group
Figure 5.1 Micro view of the hydrogel structure
5.2 Theories of Electrosensitive Release System
5.2.1 Donnan Equilibrium Theory
The optimum condition for quick bending was determined by a simple method that
identified the initial conditions based on Donnan equilibrium theory. The study [13,
14] investigated a parameter that determines an optimum condition of the content
of the ionic group and the concentration of outer solution for high-performance
electrodriven polymer hydrogel membranes.
The electrolytic cell shown in Figure 5.2 was used to investigate the bending
behaviour of poly(vinyl alcohol)-poly(2-acrylamide-2-methylpropane sulfonic acid) gel
membrane. The membrane (20 mm long and 5 mm wide) was suspended at the centre
between a pair of Pt-electrodes in the cell filled with a Na2SO4 aqueous solution. The
membrane was equilibrated in the Na2SO4 aqueous solution (1–2 valent electrolyte
solution) before the bending experiment. The time course of bending angle of the
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Electric Sensitive Release Systems
membrane was measured through monitoring the bending behaviour by a video
camera. The bending angle θ of the membrane was calculated using Equation 5.1:
 y
θ = 2 tan −1  
 x 
(5.1)
where x and y are the positions of the free ends of the membrane on the x–y coordinate
shown in Figure 5.2.
D.C. power supply
Pt electrode
30 mm
0
y
θ
20 mm
x
(x, y)
θ = 2 tan –1 (y/x)
40 mm
20 mm
Hydrogel Membrane
Figure 5.2 Schematic diagram of electrolyte cell for measuring the bending
behaviour of the gel membrane. Reproduced with permission from W.Y. Gu,
W.M. Lai and V.C. Mow, Journal of Biomechanical Engineering, 1998, 120, 169.
©1998, ASME [15]
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
The concentrations of ions in the membrane and in the outer solution are related by
Equation 5.2, based on Donnan equilibrium theory.
Cig
F Δφ 

zi
= exp  − zi
=K
Cis
RT


(5.2)
where K, Δφ, F, R, T and z denote Donnan ratio, Donnan potential, Faraday constant,
gas constant, absolute temperature and valence of ion, respectively. Cig and Cis denote
the concentrations of ions in the gel membrane (g) and in the outer solution (s). The
subscript i means the species of ion [i = A (counterion), B (co-ion), H (proton), OH
(hydroxide ion)]. The electrically neutral condition should be satisfied simultaneously
in both phases (in the gel membrane and the outer solution), as expressed by Equations
5.3 and 5.4, respectively: the outer solution):
z A C Ag + CHg − − z B CBg + COHg + CM
(5.3)
z A C As + CHS = − z B CBs + COHs (5.4)
where CM is the concentration of ionic group within the gel membrane suspended in
the electrolyte solution. The chemical equilibrium relation of water is:
K w = CHs COHs = CHz COHg
(5.5)
pH in the outer solution is 7 in this system under study (CHs = COHs). The Donnan
ratio K in a membrane-electrolyte solution system is calculated by Equation 5.6,
derived from Equations 5.2, 5.3 and 5.6:
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Electric Sensitive Release Systems
(− K
If
ZA
1

+ K ZB ) Z A Z B CS +  K −  CHS − CM = 0
K

z A = zB
1
 ZA
 K + ZB
K

1


 z A CS +  K −  CHS − CM = 0
K


(5.6)
where Cs is the concentration of outer electrolyte solution. Experiments showed the
bending behaviour of the electrodriven hydrogel membranes and provided a method
to evaluate optimum conditions for designing a high-performance membrane. The
following results were obtained from the experimental part:
• The bending rate depends on the initial condition of the system for a long period,
which means that the initial condition of the system dominates the bending
behaviour. The bending rate is linearly controlled by the intensity of the applied
electric field.
• Each membrane has a maximum bending rate at a specified concentration of outer
solution. The concentration of outer solution at the maximum bending rate shifts
to a higher concentration region of outer solution with increasing concentration
of ionic group.
• The bending rate is correlated with the thickness of membrane, by the following
equation beyond 2.5 V/cm:
Based on these experimental results, the inverse of the Donnan ratio (1/K) is the
effective simple parameter to determine the optimum condition of the concentration
of the outer solution Cs and the content of ionic group CM for the quick bending. The
bending rate of the electrodriven polymer hydrogel membrane can be predicted by
both the changeability of the swelling volume of the membrane and the conductivity
in electrolyte solution at the initial condition of the system. The electrolyte containing
a univalent counterion is effective for the quick bending of the membrane.
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
5.2.2 Mixture Theory
A mixture theory [15] was developed to model the mechano-electrochemical
behaviours of charged-hydrated soft tissues containing multielectrolytes. The mixture
is composed of n + 2 constituents (1 charged solid phase, 1 non-charged solvent phase
and n ion species).
The three main objectives of the study are:
• To develop a more general mixture theory to describe the transport of multielectrolytes through a charged-hydrated tissue, based on the triphasic theory of
Lai and co-workers [16].
• To study the electromechanical kinetic properties of such tissues associated with
a multielectrolyte transport process.
• To investigate the effects of ion exchange on tissue swelling and solvent
transport.
It is emphasised that the theory has been developed specifically to address an important
hypothesis in the articular cartilage literature, i.e., the kinetics of transport of multiions through such tissues is governed by the tissue fixed-charge density. Since many
biological membranes have similar physicochemical characteristics, the theory can be
used to derive the well-known formula [17] for the resting cell membrane potential.
In order to simplify the discussion, we confine ourselves to the study of hydrated,
isotropic and soft tissues under infinitesimal deformation, although the theory can
be easily extended to describe anisotropic and inhomogeneous materials undergoing
finite deformation [16].
5.2.3 The Generalised Triphasic Theory
The theory presented in this section is an extension of the triphasic theory for articular
cartilage developed by Lai and co-workers [16]. In the following model a chargedhydrated tissue is a continuum mixture consisting of three phases:
• a solid phase,
• a solvent phase, and
• n ion phases (species).
Each ionic species could be multivalent. In general, the tissue has a total of n + 2
constituents. Each phase is assumed to be intrinsically incompressible. The solid phase
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Electric Sensitive Release Systems
and ion phase are charged while the solvent phase is electroneutral. The tissue as a
whole must be electrically neutral.
A generalised triphasic theory composed of n + 2 species has been developed for
describing mechano-electrochemical behaviours of charged-hydrated soft tissues
with multielectrolytes. This theory is consistent with and subsumes many previously
derived mixture and phenomenological transport theories pertaining to specific
aspects of charged-hydrated tissues [16–21]. From general analyses using this theory,
it is shown that three types of force are involved in the transport of ions and solvent
through such materials:
• a mechano-chemical force (∇P; hydraulic and osmotic pressures)
• an electrochemical force (∇Nα; gradient of Nernst potential)
• an electrical force (∇Ψ)
These general results indicate that mechanical force (part of mechano-chemical force)
plays an equally important role in modulating transport process and membrane
potential as do the electrochemical forces and the electrical force in charged permeable
tissues. Further, the analyses also provide the results of the three types of mechanoelectrokinetic coefficients required to characterise solvent and ion transport through
charged-hydrated tissues:
• the hydraulic permeability (ko)
• the mechano-electrochemical coupling coefficients (bα)
• the ionic conductance matrix (hαβ)
These material coefficients governing the transport processes are nonlinearly related
to the tissue, fixed charge density (FCD), tissue hydration, ion concentrations and
the frictional coefficients (or diffusion coefficients). These general results support our
hypothesis that tissue FCD is a primary determining factor governing solvent and ion
transport through articular cartilage. Finally, the problem of Na+ and Ca++ exchange
through cartilage (an important physiologic process; [22]) was solved specifically
to illustrate how multielectrolyte solution transport occurs within a finite-thickness
specimen with swelling. The strain, concentration, flux and pressure fields within
the tissue were also calculated. Many important phenomena associated with passive
transport of ions and solvent in charged-hydrated soft tissues, such as membrane
potential, streaming potential and current, and anomalous osmosis can be described
by this generalised triphasic theory. The theory can also be applied to describe the cell
membrane potential and cell membrane selectivity in terms of ion conductance, though
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
the present continuum theory is unable to provide more information concerning the
mechanism of this selectivity at a microscopic level (e.g., ion channels).
5.2.4 Refined Multi-effect Coupling Electric-Stimulus (rMECe) Model
A multiphysics model is presented by Li [23–28] for simulation of kinetics of the smart
hydrogels subject to an externally applied electric field, especially for analysis of the
transient deformation of the hydrogel. The model termed the multi-effect coupling
electric stimulus (MECe) takes account of the coupled chemo-electromechanical
multiphysics domains and the multiphase effect of polymeric network and interstitial
liquid as well as ionic species.
5.2.4.1 Theory and Formulation
Due to mass conservation, the change in number of moles of diffusive species k (in a
volumetric space) with respect to time t can be characterised by the difference between
the fluxes entering and leaving the reference volume. The Nernst–Planck type of
continuity equations can be written as [29] as seen in Equation 5.7:
∂ck
∂c
+ div (J k ) = k +
∂t
∂t

z F


div − [Dk ].  grad (ck ) + k ck grad (ψ )   = 0
RT



(k = 1, 2,....., N )
(5.7)
where Jk, [Dk], ck and zk are the flux (mM/s), diffusivity tensor (m2/s), concentration
(mM) and valence number of the kth diffusive ionic species. N is the number of
diffusive ionic species and c the electrostatic potential. F, R and T are the Faraday
constant (9.6487 × 104 C/mol), universal gas constant (8.314 J/mol K) and absolute
temperature (K), respectively.
To describe the spatial distribution of the electric potential in the domain, the Poisson
equation, derived from Gauss Law, is written as seen in Equation 5.8:
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Electric Sensitive Release Systems
∇ 2ψ = −
F  N
∑ zk ck + z f c f
εε 0  k

,

(5.8)
where ε is the relative dielectric constant of the surrounding medium, ε0 the vacuum
permittivity or dielectric constant (8.85418 × 10–12 C2/Nm2) and cf is the density of
the fixed charge groups bound to the polymeric chains of the hydrogel.
The first contribution of the present rMECe model is the reformulation of the fixed
charge density for incorporating the effect of externally applied electric field. Based
on the triphase mixture theory [16] a hydrogel is defined as a three-dimensional
crosslinked polymeric hydrophilic network consisting of solid matrix, interstitial
water and ionic phases. The corresponding saturation equation can be written as
seen in Equation 5.9:
∑
α = s , w, k
φ α = 1,
(5.9)
where φα (α = s, w, k) represents the volume fractions of the solid, water and ion
phases, respectively. The volume fraction of the water phase is given by Lai and coworkers [16], see Equation 5.10:
φw = 1−
φ0s
,
1 + tr ( E ) (5.10)
where E is the elastic strain vector of solid matrix phase and φs0 is the volume fraction
of the solid phase at reference configuration. Compared with φs and φw, φk is negligibly
small. The saturation Equation 5.9 can be simplified to:
φ s + φ w = 1,
(5.11)
and the corresponding saturation equation at reference configuration is written as:
as seen in Equation 5.12:
φ0s + φ0w = 1.
(5.12)
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
Based on Equations 5.10 and 5.12, the volume fraction of the current water phase
is obtained by Equation 5.13:
φ0s
1 − φ0w
tr ( E ) + φ0s
φ = 1−
= 1−
=
1 + tr ( E )
1 + tr ( E ) 1 + tr ( E ) w
(5.13)
By the triphasic theory [16] and using Equation 5.7, the fixed charge density is finally
derived as follows in Equation 5.14:
cf =
φ0w c 0f
=
φ0w c 0f
φ w [1 + tr ( E ) ] φ0w + tr ( E )
=
c 0f
1 + tr ( E ) / φ0w
(5.14)
As the second contribution of the rMECe model, the mechanical governing equation
is treated in finite deformation, instead of small deformation. In consideration of
chemo-electromechanical coupling effects, the electric-stimulus-responsive hydrogels
undergo large deformation when the applied electric voltage is relatively high, in
which case the linear theory does not provide sufficiently accurate computation. This
results from the fact that the difference between the initial (reference) and deformed
configurations cannot be neglected as is done for the case of linear elasticity. For
geometrically nonlinear analysis, the governing equations of large deformation using
a total Lagrangian description are given as seen in Equations 5.15-5.17:
∇.P = ∇.( FS ) = 0
in Ω, u = G in Γ g ,
P.N = H
in Γ h , (5.15)
(5.16)
(5.17)
where F is the deformation gradient tensor, G is the specified displacement vector on
the boundary portion Γg, H is the surface traction vector on the boundary Γh, N is the
unit outward normal vector, u is the displacement vector from the initial configuration
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Electric Sensitive Release Systems
X to the deformed configuration x. P is the first Piola–Kirchhoff stress tensor that is
a kind of expatriate, living partially in the deformed (current) configuration x and
partially in the reference configuration x, where x = X + u.P). As P is immeasurable
and asymmetrical, the second Piola–Kirchhoff stress tensor S is required because S is
symmetric and is often used as the stress measure as seen in Equation 5.18:
S = CE − posmotic I .
(5.18)
where C is the material tensor, posmotic is the osmotic pressure, I is the identity tensor
and E is the Green–Lagrangian strain tensor used as strain measure. Substituting
Equation 5.18 into Equation 5.15, the large deformation governing equation of the
hydrogel is obtained by Equation 5.19:
∇.[F (CE − posmotic I ) ] = 0.
(5.19)
However, when the applied electric voltage is low, the hydrogels have small
deformation. Then the linear theory satisfies sufficiently the computational accuracy
requirement, and Equation 5.19 is thus reduced to Equation 5.20:
∇.σ = ∇.(λs tr ( E ) I + 2 µ s E − posmotic I ) = 0,
(5.20)
where σ is the Cauthy stress tensor. λs and μs are the Lamé coefficients of solid
matrix.
5.2.4.2 Boundary and Initial Conditions
− for
The computational domain of the ionic concentrations c−k and electric potential ψ
Equations 5.20 and 5.21 is defined as covering both the hydrogel and surrounding
− are imposed
bath solution. As such, the boundary conditions of the unknown −
c k and ψ
at the two electrodes located at the ends of the bath solution as illustrated in Figure
5.3:
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
D
h
Hydrogel
a
0
b
X
Solution
Xgel
0
L
Cathode
Anode
Figure 5.3 Diagram of a hydrogel strip immersed in a bath solution under
applied electric field, where the whole computational domain is defined as the
X-coordinate system and the hydrogel domain as the X gel-coordinate system.
Reproduced with permission from W.Y. Gu, W.M. Lai and V.C. Mow, Journal of
Biomechanical Engineering, 1998, 120, 169. ©1998, ASME [15]
_
_
C | Anode = C |Cathode = C *
ψ | Anode = 0.5Ve
and ψ |Cathode = −0.5Ve (5.21)
(5.22)
where c−* is the initial salt concentration of the bath solution and v−k is the applied
voltage. However, the computational domain of the fluid pressure p− and hydrogel
displacement u− for Equations 5.22 and 5.23 covers the hydrogel region only. The
corresponding boundary conditions are required at both the interfaces between the
hydrogel and surrounding solution. Based on the assumption that, at equilibrium
state, the chemical potentials of fluid and ion phases inside the hydrogels should be
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Electric Sensitive Release Systems
equal to those outside the hydrogels, the boundary condition of the fluid pressure
−
P at the hydrogel–solution interfaces is given as:
p int erface = RT (C
+
in −int erface
+C
−
in −int erface
−C
+
out −int erface
−C
−
out −int erface
) − p0
(5.23)
where c−kin-interface and c−kout-interface (k = +, −) are the ionic concentrations within the
hydrogels near the interfaces and within the surrounding bath solution near the
interfaces, respectively. p0 denotes the fluid pressure at reference configuration. In
terms of the mechanical state of the mixture phase, the boundary condition of the
hydrogel displacement u−at the hydrogel–solution interface can be written as seen in
Equation 5.24:
(3λs + 2 µ s )
∂u int erface
= β RTcref p int erface
∂x
(5.24)
To implement the transient simulation for kinetics of the electric-sensitive hydrogels,
initial conditions are required. It is assumed here that the initial hydrogel is in the
equilibrium state when the effect of the bath solution is considered only and no external
electric field is applied. This equilibrium state will be taken as initial condition for
transient simulation. Corresponding steady-state computational results can thus be
used as the initial conditions as seen in Equations 5.25-5.28:
transient
c initial
ψ initial
p initial
u initial
transient
transient
transient
steady
= c v =0
(5.25)
(5.26)
(5.27)
steady
= ψ v =0
steady
= p v =0
steady
= u v =0
(5.28)
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
c
steady
steady transient
transient transient
steady transient
where
,
,
and
initial
v =0
initial
initial
v = 0 initialv = 0
ψ= c
p= ψ
=up
steady
= u v =0
represent the steady-state computational
results without the externally applied electric field.
5.2.4.3 Discretisation of the Transient Governing Equations of the MECe
Model
A meshless numerical method, termed the Hermite-cloud method [30, 31], is used for
transient solution of the nonlinear partial differential governing equations of the MECe
model. The Hermite-cloud method constructs the Hermite-type interpolation functions
and employs the point collocation for discretisation of the governing equations to
directly compute the approximate solutions of both unknown functions and firstorder derivatives [30, 31]. By the Hermite-cloud method, an unknown continuous
real function f(x,y) can be expressed approximately by Equation 5.29:
Ns
NT
NT


f ( x, y ) = ∑ N n ( x, y ) f n + ∑  x − ∑ N n ( x, y ) xn M m ( x, y ) g xm
n =1
m =1 
n =1

NT
Ns


+ ∑  y − ∑ N n ( x, y ) yn  M m ( x, y ) g ym
m =1 
n =1

(5.29)
where Nn(x,y) and Mn(x,y) are defined as the shape functions of the unknown
function f(x,y) and corresponding first-order differential functions gx(x,y) and gy(x,y),
respectively, which are simply polynomials in x and y. fn denotes the unknown point
value of f(x,y) at the nth discrete point, gxm and gym the unknown point values of
gx(x,y) and gy(x,y) at the mth discrete point. NT and NS are total numbers of discrete
points scattered within the computational domain.
By the θ-weighted finite difference scheme (0.5 < θ < 1.0) [32], the equation is
discretised first in time t domain as:
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Electric Sensitive Release Systems
k
k
C ( n +1) − C ( n )
k
  2 k
∂C ( n +1) ∂ψ ( n +1)
∂ 2ψ ( n +1)  
k
∂
C
( n +1)
k
k
θ 

+
+
α
α
z
z
C
.
.
n
+
(
1)
2
2
 

∂ x.∂ x
∂x
ΔtDk   ∂ x

= 2
k

2
Lref 
 ∂2 C k

∂C ( n ) ∂ψ ( n )
k ∂ ψ (n)
(n)
k
k
 +(1 − θ ) 

+
+
α
z
α
z
C
.
.
(
)
n
2

 ∂ x2
∂ x.∂ x
∂ x  


(5.30)
where the subscript n denotes time t = tn and Δt = tn+1 – tn is time step. The model
is composed of coupled nonlinear partial differential governing equations, and it
has been examined for kinetic study by comparison with published experiments.
The kinetics of diffusive ionic species concentration is simulated well and is in
good agreement with the published steady-state simulation. Influences of important
parameters are discussed in detail, including the externally applied electric voltage,
initially fixed charge density and surrounding bath solution concentrations as well as
the displacement of the hydrogel. By the present transient simulations, it is concluded
that the smart hydrogels have the capability of responding to the externally applied
electric field in short time.
5.3 Measurement of Bending Angle
Jing Shang and co-workers [33] prepared a natural amphoteric polyelectrolyte
hydrogel film by solution blending of CS and its derivative carboxymethyl chitosan
(CMCS) and crosslinked with glutaraldehyde, compared to the CS/carboxymethyl
cellulose (CMC). CS/CMCS hydrogels show a faster bending rate, a larger bending
angle and a higher tensile strength, which suggests great potential for their application
in natural polymeric electroactive hydrogels. The authors previously prepared CS/
CMC hydrogels [34], which exhibit a good electromechanical response over a wide
range and could change their bending direction depending on the pH of the buffer
solution.
A schematic diagram of the equipment used for studying the electrical response of
hydrogels is shown in Figure 5.4. All hydrogel strips (10 mm long and 2 mm wide)
were first swollen to equilibrium in an electrolyte solution (Britton–Robinson pH
buffer solution) to be used later in the bending experiments; the thickness of the
hydrogel in a swollen state was 0.145 mm (pH = 6). Two parallel carbon electrodes,
50 mm apart, were immersed in the electrolyte solutions and the hydrogel strip
under investigation was mounted centrally between them. Upon application of a
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
direct current (DC) electric field, the degree of bending, q, was measured by reading
the angle of deviation from the vertical position. The authors define the value of
the bending angle as being positive when the hydrogel bends toward the anode and
negative when it bends toward the cathode. The bending behaviour was recorded
with a digital camera.
carbon electrodes
electrolyte solution
θ
hydrogel
bending angle
Figure 5.4 Schematic diagram of the apparatus for testing the bending behaviour
of hydrogels
The CS/CMCS hydrogels exhibit better overall mechanical properties and electrical
sensitivity, suggesting their great potential for microsensor and actuator applications,
especially in the biomedical field. Natural amphoteric CS/CMCS hydrogels were
prepared by a simple solution blending of two natural polymers, CS and its derivative
CMCS. Such an amphoteric hydrogel gave a fast mechanical response in an electric
field (bending toward an electrode) in different electrolyte solutions and over a wide
pH range. The bending behaviour of the hydrogels was different in different pH buffer
solutions, i.e., when pH ≤ 7, it bent toward the anode, but when pH > 7 it bent toward
the cathode. The equilibrium bending angle of CS/CMCS hydrogels was influenced
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Electric Sensitive Release Systems
by pH, ionic strength, electric strength, the thickness of the hydrogel film and content
of crosslinking agent in the hydrogel. The equilibrium bending angle of the hydrogel
reached a maximum that was over 90° in the buffer solution at pH = 6 with 0.2 M
ionic strength. In addition, the authors found that thin hydrogel films and optimised
crosslinking were favourable to enhance the electrical sensitivity. The CS/CMCS
hydrogel shows better overall mechanical properties and electrical sensitivity than
the CS/CMC hydrogel, which indicates its great potential for application in artificial
muscles, microsensors and actuators, particularly in biomedical fields because all the
materials are from CS, an excellent natural material.
5.4 Application of Electrosensitive Release System
The poly(3,4-ethylene dioxythiophene)/poly(styrenesulfonate) (PEDOT/PSS) entangled
in the polyacrylamide (PAAm) network by the formation of a semi-interpentrating
network (semi-IPN) on hydrogel performance. In addition, PAAm networks with
good mechanical properties have been obtained [35]. The presence of PEDOT/
PSS increased the ionic conductance of swollen semi-IPN hydrogels substantially.
Electrochemical experiments demonstrated that PAAm-PEDOT/PSS hydrogel is
electrochemically stable and presents reversible responses to electrochemical stimuli.
Scanning electron microscopy images revealed that the morphology of conducting
hydrogels changed significantly with the entrapment of PEDOT/PSS. The data from
Wu [7] and mechanical properties revealed that the presence of PEDOT/PSS within
the hydrogels changes the hydrophobicity and, consequently, the compaction of the
network at lower crosslinking density. These conducting hydrogels presented good
mechanical properties, which were driven by the amounts of acrylamide (AAm) and
N,N-methylene bisacrylamide (MBA) in the hydrogel synthesis feed solution. On the
other hand, the presence of PEDOT/PSS increased ionic conductance substantially,
and it is more significant in highly swollen semi-IPN hydrogels. It was suggested that
the water molecules and the PEDOT/PSS chains in the hydrogel help ionic diffusion
through the polymer networks by increasing ionic conductance. In addition, as the
amounts of AAm and MBA used in the hydrogel synthesis feed solutions may be
controlled, this process allows tailoring hydrogel ionic conductance and mechanical
properties. The results shown in this work indicate that the hydrogels obtained
have potential for further tests as conducting hydrogel membranes, e.g., sensors and
electrodes. Electrochemical experiments have demonstrated that PAAm-PEDOT/PSS
hydrogel has good electrochemical stability and reversible response to electrochemical
stimuli. These attributes make it attractive for capacitor applications.
Ramanathan and Block [36] evaluated and characterised the use of CS gels as
matrices for electrically modulated drug delivery. In electrification studies, release-time
profiles for neutral (hydrocortisone), anionic (benzoic acid) and cationic (lidocaine
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
hydrochloride) drug molecules from hydrated CS gels were monitored in response
to different milliamperages of current as a function of time. Likewise, chondroitin
4-sulfate hydrogels were examined by Jensen and co-workers [37] as potential matrices
for the electrocontrolled delivery of peptides and proteins. Kim and co-workers [38]
synthesised an interpenetrating polymer network (IPN) hydrogel composed of PVA
and CS which exhibited electrosensitive behaviour. They investigated the response
of the IPN hydrogel in electric fields. A swollen PVa/CS/IPN was placed between a
pair of electrodes and bending behaviour in response to the applied electric field was
noted. The bending angle and the bending speed of the PVA/CS IPN increased with
increasing applied voltage and concentration of NaCl aqueous solution. Synthetic as
well as naturally occurring polymers, either alone or in combination, have also been
used. Examples of naturally occurring polymers include hyaluronic acid, chondroitin
sulfate, agarose, xanthan gum and calcium alginate. The synthetic polymers are mostly
methacrylate and acrylate derivatives such as partially hydrolysed polyacrylamide,
polydimethyl aminopropyl acrylamide, and so on.
Electrically induced anisotropic gel deswelling was first explained by Tanaka and coworkers [39], who suggested that the electric field produces a force on both the mobile
counterions and the immobile charged groups of the gel’s polymeric network. For
example, in partially hydrolysed polyacrylamide gels fixed to one of the electrodes,
the mobile H+ ions migrate toward the cathode while the negatively charged immobile
acrylate groups in the polymer networks are attracted toward the anode. The anionic
polymeric gel network is pulled toward the anode. The pull creates a uniaxial stress
along the gel axis, the stress being greatest at the anode and smallest at the cathode.
This stress gradient is thought to contribute to the anisotropic gel deformation. When
the gel is not fixed to either electrode, the attractive forces between the immobile
negative charges of the polymer network and the anode can result in translational
movement of the gels towards the anode. In a recent study by Bajpai and co-workers
[40], polyaniline was impregnated into a macromolecular matrix of PVA-g-poly(acrylic
acid) (PAA) and electrical conductivity and electroactive behaviour of the resulting
nanocomposite were studied.
Kishi and co-workers [41] prepared drug-delivery devices responding to electrical
stimuli to alter gel-swelling behaviour. The gels responded to the on-off stimulus
of electrical currents, which induced the gel swelling-shrinking that contributes
to the release of drug molecules. The researchers prepared poly(sodium acrylate)
microparticle gels containing pilocarpine as a model drug. During application of
a direct electric current, the pilocarpine release increased in a current-dependent
manner. However, the pilocarpine release did not stop upon termination of the
electrical stimulus, since the prepared gels themselves maintained a highly swollen
state. Thus, a complete on-off release regulation of drugs cannot be achieved with
this polymer gel system.
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Electric Sensitive Release Systems
The electromechanical behaviour of crosslinked strong acid hydrogels was investigated
in different salt solutions and in a 1.6 V/cm DC electric field [3]. The sulfonated
polystyrene-based copolymer networks were observed to bend toward the cathode
when they were preequilibrated in the corresponding salt solutions. A polythiophenebased conductive polymer gel actuator was reported to show expansion/contraction
behaviour in response to an applied potential [42]. The axial pressure generated by
the expansion of the gel against a fixed wall was measured, which demonstrated that
the generated closure pressures could be utilised as a small actuator valve. In most
cases, electric field responsive polymer networks are based on electrically driven
motility. On the other hand, some neutral polymer gels were reported to show electric
field sensitivity. Weakly crosslinked poly(dimethyl siloxane) gels containing finely
distributed electric field sensitive particles (TiO2 particle) showed a significant and
quick bending in silicon oil [43]. The mechanism of this responsiveness was that the
forces acting on the particles, which could not leave the gel networks, directly caused
the locomotion or the deformation of the gel.
Maleic anhydride was used to prepare polyvinyl alcohol/poly(sodium maleate-cosodium acrylate) hydrogels by a repeated frost–defrost process because of its higher
charge density and potential electric stimuli sensitivity [1]. The bending angle was
measured in a non-contact electric field using carbon as plate electrodes. It was found
that the bending angle was dependent on various factors, including composition of
hydrogel, concentration of NaCl solution, types of electrolyte solution and electric
voltage. It exhibited that the bending angle increased when the concentration of
NaCl solutions and the electric voltages increased. An abnormal bending direction
was observed, and it was affected not only by the kinds of hydrogels, but also by the
exterior variations. The hydrogel showed good reversibility in on-off electric field
and could be a candidate for practical application.
Kwon and co-workers [44–46] prepared crosslinked poly(2-acrylamide-2methylpropane sulfonic acid-co-butyl methacrylate) (P(AMPS-co-BMA)) hydrogels
and evaluated the applicability of these hydrogels for electrical stimuli-responsive
drug-delivery devices. They used a cationic drug molecule, edrophonium chloride,
within the negatively charged hydrogel. Rapid drug release from the hydrogels resulted
from an application of electric fields through the ion exchange between positively
charged drug molecules and protons at the cathode. The squeezing effects arising
from the electric field application induced rapid drug release from the gels, which
increased as the voltages increased in a dose-dependent manner. Using the P(AMPSco-BMA) hydrogels, an on-off drug release regulation was achieved under an on-off
application of electric current.
Kwon and co-workers [47] further investigated the electric current-induced release
of anionic heparin from a positively charged polyallylamine polyion complex. Rapid
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
structural changes and an apparent dissociation of the polyion complex occurred
upon application of an electric current. During the electric current application,
the positively charged poly allylamine was neutralised at the cathode owing to the
microenvironmental pH changes, and apparent dissociation of the polyion complex
occurred. Although bioactive heparin was released by electric current application,
polyallylamine was also released. Such positively charged molecules are incompatible
in vivo, and these complications must be addressed before clinical applications can
be considered.
Hsu and Block [48] studied electrokinetic phenomena in three types of anionic gels
(agarose, agarose-carbomer 934P and agarose-xanthan gum) under an applied electric
current for electrically modulated drug delivery. It was shown that electrical current
strength and gellant content could influence both syneresis and drug migration. These
findings reported by the two research groups may give a practically useful insight
into the application of hydrogels to transdermal delivery assisted by electroportion,
iontophoresis or sonophoresis.
Electrosensitive hydrogels, which are basically pH-sensitive hydrogels, are able to
convert chemical energy to mechanical energy [49]. These systems can serve as
actuators or artificial muscles in many applications. All living organisms move by
the isothermal conversion of chemical energy into mechanical work, e.g., muscular
contraction, and flagellar and ciliary movement. Electrically driven mobility and
ciliary movement has been demonstrated using weakly crosslinked poly(2-acrylamido2-methylpropanesulfonic acid) hydrogels. In the presence of positively charged
surfactant molecules, the surface of the polyanionic hydrogels facing the cathode is
covered with surfactant molecules reducing the overall negative charge. This results in
local shrinkage of the hydrogel leading to bending of the hydrogel. Application of an
oscillating electrode polarity could lead the hydrogels to quickly repeat its oscillatory
motion, leading to a worm-like motion [50].
Reversible change in optical properties of ionic polymeric gels, 2-acrylamido2-methylpropane sulfonic acid and PAA plus sodium acrylate crosslinked with
bisacrylamide, under the effect of an electric field is reported. The shape of a cylindrical
piece of the gel, with flat top and bottom surfaces, changed when affected by an
electric field [51]. The top surface became curved and the sense of the curvature
(whether concave or convex) depended on the polarity of the applied electric field.
The curvature of the surface changed from concave to convex and vice versa by
changing the polarity of the electric field. By the use of an optical apparatus, focusing
capability of the curved surface was verified and the focal length of the deformed gel
was measured. The effect of the intensity of the applied electric field on the surface
curvature and, thus, on the focal length of the gel are tested. Different mechanisms are
discussed; either one or a combination of these may explain the surface deformation
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Electric Sensitive Release Systems
and curvature. Practical difficulties in the test procedure and the future potential of
the electrically adaptive and active optical lenses are also discussed. These adaptive
lenses may be considered as smart adaptive lenses for contact lenses or other optical
applications requiring focal point undulation.
Toshihiro Hirai and co-workers [52] reported that non-ionic gel of PVA swollen with
dimethyl sulfoxide responds to an electric field rapidly with a large and reversible
deformation, which includes a spherical bending motion to the anode side and a
contraction in the direction of an electric field. The electrically induced strain in
the polymer gel has been suggested to be due to electrically induced unidirectional
movement of solvent in the gel. The chemically crosslinked polymer gels used here,
which had only about 2 wt% of polymer content and a much larger amount of solvent
in the gel networks, were a good elastic body to suffer a large mechanical deformation.
Here, we discussed mechanical properties of the gel under and out of the application
of an electric field, and also the interaction between the polymer network and solvent.
The electrically induced deformation was discussed, as well, on the proportionality
of strain to the square of an electric field for both the bending motion to the anode
side and the contracting motion in the field direction. Furthermore, we drove out the
electrically induced solvent force in the gel under various conditions, and demonstrated
that the electroactive non-ionic gel can be used in some mechanical devices.
Composites of monodomain nematic liquid crystal elastomers and a conducting
material distributed within their network are shown to exhibit large deformations,
i.e., contraction, expansion, bending with strains of over 200% and appreciable force,
by Joule heating through electrical activation [53]. The electrical activation of the
conducting material induces a rapid Joule heating in the sample leading to a nematic to
isotropic phase transition where the elastomer of dimensions 32 mm × 7 mm × 0.4 mm
contracted in less than a second. The cooling process, isotropic to nematic transition
where the elastomer expands back to its original length, was slow and took 8 seconds.
The material studied here is a highly novel liquid crystalline co-elastomer, invented
and developed by Heino Finkelmann and co-workers at Albert-Ludwigs-Universität in
Freiburg, Germany. The material is such that the mesogenic units are in both the side
chains and the main chains of the elastomer. This co-elastomer was then mechanically
loaded to induce a uniaxial network anisotropy before the crosslinking reaction
was completed. These samples were then made into a composite with a conducting
material such as dispersed silver particles or graphite fibres. The final samples were
capable of undergoing more than 200% reversible strain in a few seconds. Sun and
Mak [54] prepared a smart hydrogel fibre based on CS/poly(ethylene glycol). The
dynamics of this hydrogel fibre in response to electric stimulation is reported. The
effects of a number of factors have been systematically studied, including the fibre
diameter, concentration of the crosslinking agent, electric potential imposed across
the fibre, pH and ionic strength of the bath solution. Fibre deformation is expressed
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Stimuli Responsive Drug Delivery Systems: From Introduction to Application
in terms of the curvature at the mid length of the fibre for various times. The amount
of bending to a given extent within a given time period is used to describe the rate of
cyclic deformation. Our experimental results show a stable reversibility of bending
behaviour under the applied electric field. The bending curvature is proportional to
the intensity of the applied electric potential. Although adequate mechanical properties
are maintained, the rate of deformation can be improved via the adjustment of a
number of the aforementioned extrinsic factors. These observations are interpreted
in terms of fibre stiffness, fixed charge density and swelling pressure, which depend
on the hydrogel equilibrium states in different pH and ionic environments along with
the electrochemical reactions under the electric field.
PVA and PAA composite fibres were prepared via solution spinning and, subsequently,
semi-IPN were obtained by crosslinking the fibres with glutaraldehyde [55]. The
hydrogel fibres exhibited bending behaviour under DC electric stimulation. The effects
of a number of factors have been systematically studied, including the PAA content
within the network, electric voltage imposed across the fibre, the fibre diameter,
concentration of the crosslinking agent, pH and ionic strength of the bath solution.
Our experimental results show a stable reversibility of bending behaviour under the
applied electric field. The degree of bending at equilibrium and the bending speed
of the hydrogel fibre increased with the intensity of the applied electric voltage and
the PAA content having negatively charged ionic groups within the semi-IPN. The
electroresponsive behaviour of the present semi-IPN hydrogel fibre was also affected
by the aforementioned extrinsic factors. These observations are interpreted in terms
of fibre stiffness, fixed charge density and swelling pressure, which depend on the
hydrogel equilibrium states in different pH and ionic environments together with the
electrochemical reactions under DC electric field.
Semi-interpenetrating polymer networks (semi-IPN) composed of CS and PAAm
hydrogels have been prepared, and the effect of changing pH, temperature, ionic
concentration and applied electric fields on the swelling of the hydrogels was
investigated [56]. The swelling kinetics increased rapidly, reaching equilibrium within
60 min. The semi-IPN hydrogels exhibited a relatively high swelling ratios of 385–
569% at T = 25 °C. The swelling ratio increased with decreasing pH below pH = 7
due to the dissociation of ionic bonds. The swelling ratio of the semi-IPN hydrogels
was pH, ionic concentration, temperature and electric field dependent. Differential
scanning calorimetry was used to determine the volume of free water in the semi-IPN
hydrogels, which was found to increase with increasing PAAm content.
IPN hydrogel composed of PVA and poly(diallyldimethylammonium chloride)
(PDADMAC) exhibited electrical-sensitive behaviour [57]. The PVA/PDADMAC IPN
hydrogel was synthesised using the sequential IPN method. The swelling behaviour
of the IPN hydrogel was studied by immersion of the gel in aqueous NaCl solutions
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Electric Sensitive Release Systems
at various concentrations and pH. The swelling ratio decreased with an increase in
NaCl concentration. The stimuli response of the IPN hydrogel in electric fields was
also investigated. When swollen IPN hydrogel was placed between a pair of electrodes,
the IPN exhibited bending behaviour upon the application of an electric field. The
IPN also showed stepwise bending behaviour depending on the electric stimulus.
The bending mechanism of an electrodriven ionic polymer hydrogel membrane was
investigated in terms of the concentration of ionic groups and the valence of ionic
species within the hydrogel membrane. It was elucidated that the distribution of
the ionic concentration played an important role in the bending behaviour of the
membrane [58].
IPN composed of PVA and PAA exhibited electrical-sensitive behaviour [59]. PAA
as an initial network was prepared inside a PVA solution using ultraviolet (UV)
irradiation; then, PVA networks as a secondary network were formed by a repetitive
freeze–thawing process. Their mechanical properties were influenced by the swelling
ratio, crosslinking by UV radiation and a freeze–thawing process, and intermolecular
force by hydrogen bonding. When a swollen PVA/PAA IPN was placed between a
pair of electrodes, the IPN exhibited bending behaviour upon the applied electric
field. The equilibrium bending angle (EBA) and the bending speed of the PVA/PAA
IPN increased with the applied voltage and the content of the PAA network having
negatively charged ionic groups within the IPN. The electroresponsive behaviour of
the present IPN was also affected by the electrolyte concentration of the external
solution. Particularly, IPN37 showed a maximum EBA when the critical ionic strength
was 0.1. Anisotropic deswelling of the IPN was observed in a direct contact with a
pair of electrodes under aerobic conditions. The PVA/PAA IPN also showed stepwise
bending behaviour depending on the electric stimulus.
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