Download Influence of electrostatic interaction on fibrinogen adsorption on gold

Journal of Colloid and Interface Science 283 (2005) 477–481
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Influence of electrostatic interaction on fibrinogen adsorption on gold
studied by imaging ellipsometry combined with electrochemical methods
Yong Yu, Gang Jin ∗
Laboratory of National Microgravity, Institute of Mechanics, Chinese Academy of Science, Beijing 100080, People’s Republic of China
Received 31 March 2004; accepted 15 September 2004
Available online 5 November 2004
Abstract
Imaging ellipsometry was combined with electrochemical methods for studying electrostatic interactions of protein and solid surfaces.
The potential of zero charge for gold-coated silicon wafer/solution interfaces wad determined by AC impedance method. The potential of the
gold-coated silicon wafer was controlled at the potential of zero charge, and the adsorption of fibrinogen on the potential-controlled and noncontrolled surfaces was measured in real time at the same time by imaging ellipsometry. The effect of electrostatic interaction was studied by
comparing the difference between the potential of controlled adsorption and the potential of noncontrolled adsorption. It was shown that the
rate of fibrinogen adsorption on the potentiostatic surface was faster than that on the nonpotentiostatic surface. The electrostatic influence on
fibrinogen adsorption on the gold-coated silicon wafer was weak, so the hydrophobic interaction should be the major affinity.
 2004 Elsevier Inc. All rights reserved.
Keywords: Electrostatic interaction; Protein adsorption; Imaging ellipsometry; Electrochemistry
1. Introduction
Many studies have been carried out to investigate the effect of interactions which include H-binding, electrostatic,
and hydrophobic interactions on protein adsorption [1–40].
The electrostatic interaction, which is among the interactions
between the proteins and surfaces, was important in protein
adsorption, which was studied by a lot of researchers [1,2,
19–40]; even so, no clear consensus have been developed as
to electrical charge effect [1]. Studies such as Guo et al. [26]
indicated that the adsorption of negatively charged albumin
was highly accelerated by the application of a positive potential to the solid surface, and Bernabeu et al. [24] found
that negatively charged proteins could be adsorbed onto a
negatively charged surface and the adsorption increased with
increasing negative surface charge. Choosing a fit method is
the key to studying electrostatic interactions of proteins and
the solid surfaces. The first problem is how to control the
* Corresponding author. Fax: +86-10-62631816.
E-mail address: [email protected] (G. Jin).
0021-9797/$ – see front matter  2004 Elsevier Inc. All rights reserved.
doi:10.1016/j.jcis.2004.09.021
charge of the protein and the solid surface. The pH variation can control the total charge of the protein; on the other
hand, the solid surface is modified chemically [2,20,21] or
controlled by electrochemical methods [22–34] to change
the charge. When the surface charge is changed by chemical modification, the H-bonding and hydrophobic interaction
will also change more or less at the same time. Using electrochemical methods may change surface charge arbitrarily
without influencing other interactions. At present, the impedance method or cyclic voltammetry has been utilized to
study protein adsorption. Single electrochemical methods do
not allow measurement of protein adsorption in real time
and control of surface charge simultaneously. Thus a technique should be developed to combine with electrochemical
methods to measure protein adsorption in real time. Some
sensitive optical techniques could be chosen to combine with
electrochemical methods, such as ellipsometry [41], reflectometry [42], optical waveguide methods [43–45], and total
internal reflection fluorescence [46].
In this paper, imaging ellipsometry [41] was combined
with electrochemical methods. Ellipsometry is an optical
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Y. Yu, G. Jin / Journal of Colloid and Interface Science 283 (2005) 477–481
technique for characterization of an interface or film between two media. This technique is based on exploiting the
polarization transformation that occurs as a beam of polarized light is reflected from or transmitted through the
interface or film. Its nonperturbing character and remarkable sensitivity are the principal advantages of ellipsometry.
Imaging ellipsometry has been developed recently and used
in many applications. The novelty of this technique lies in
the use of combined null and off-null ellipsometry on a substrate with a high refractive index. The measurement principle is to generate an ellipsometric image of the surface
under investigation by using a CCD camera as a detector in
an ellipsometric configuration. One image containing more
than 105 sampling points (pixels) is possible to record in
less than 1 s on an area of the order of square centimeters.
It has the advantages of high sensitivity to layer thickness
(0.5 nm) and high lateral resolution obtainable with a CCD
camera (pixel resolution better than 5 µm in this mode).
Protein adsorption in liquid can be measured in real time.
Imaging ellipsometry is applied compatibly in the fields of
biophysics, molecular biology, and medicine. In this paper,
impedance measurement was utilized to determine the potential of zero charge for gold-coated silicon wafer/solution
interfaces. At the potential of zero charge, the surface of
the conductor had no electrostatic charge. Then the potential of the gold-coated silicon wafer was controlled near
the potential of zero charge, and the charge of the protein
was controlled by pH variations. The adsorption of fibrinogen on the gold-coated silicon wafer was measured in real
time by imaging ellipsometry. The effect of electrostatic interaction was studied by comparing the difference between
potential controlled adsorption and potential non-controlled
adsorption. The effect of electrostatic interaction on protein
adsorption was fully studied by using this combined technique.
2. Experimental procedures
2.1. Reagents and solutions
The fibrinogen was purchased from Sigma. The supporting electrolyte was a 0.1 mol/L KH2 PO4 –Na2 HPO4 buffer
solution (pH 7.80). All other reagents were of analytical
grade. Distilled water was used.
2.2. Experimental equipment
The experimental setup is shown in Fig. 1. It is based
on a PCSA ellipsometer combined with an electrochemical
instrument [41]. The PCSA ellipsometer has a collimated expanded beam and a CCD camera as a detector. In the design
of the optical system, special considerations that should be
taken are as follows:
Source intensity and detector linearity. Operation in both
null and off-null ellipsometric modes is required. Null el-
Fig. 1. The schematic diagrams of an imaging ellipsometer combined with
an electrochemical instrument and a sample cell.
lipsometry is independent of detector linearity and source
intensity fluctuations. However, off-null ellipsometry is not,
and therefore the stability of the light source and the linearity
of the sensor elements in the CCD camera are important.
Parallel light. The expanded probe beam must have parallel light to ensure that the angle of incidence is the same
over the entire test surface. This is essential for the quantification of the surface thickness distribution.
Beam uniformity. The intensity over the cross section of
the probe beam must be constant, because the intensity distribution over the beam area after reflection is a measure of
the thickness distribution of the layer under study. A nonuniformity of the probe beam intensity leads to errors in the
thickness distribution.
In our experimental setup, a 75-W xenon arc lamp and a
collimating system are used as a light source to provide an
expanded parallel probe beam with a diameter of 25 mm. To
increase the ellipsometric contrast, a wavelength-selective
filter (632.8 nm) with a bandwidth of 10 nm is placed in the
incident optical path. The CCD camera is a Sony XC-73CE
CCD B/W video camera module; it is used as an intensity
detector.
The structure of the cell is shown in Fig. 1. It was divided into two parts. There is a 1.5-mm slot at the joint of
the two parts. The small quantity of protein added from sample input was first mixed equally in the left part (in Fig. 1)
and then diffused towards the working electrode through the
slot. The aim of this design is to avoid disturbance in transportation.
Y. Yu, G. Jin / Journal of Colloid and Interface Science 283 (2005) 477–481
Impedance data were modeled using a CHI 650A (CH
Instrument, USA). The potential was controlled by an HDV7C constant potential instrument (Sanming Second Wireless
Instrument Factory, China). A three-electrode system, consisting of a gold-coated silicon wafer as a working electrode,
a saturated calomel reference electrode, and a platinum black
counterelectrode, was used.
The gold coating in the middle of small pieces of goldcoated silicon surface was removed to expose the silicon
substrate to divide the gold coating into two areas. Then
the gold coating was washed in H2 SO4 + 30% H2 O2 solution (7:3), ethanol, and distilled water ultrasonic for 3 min.
After being rinsed in distilled water, the gold coating was
connected with a wire to be used as a working electrode.
The working electrode was made in an electrochemical wash
using cyclic voltammetry from 1.5 to −0.3 V (vs SCE) in
0.5 mol/L H2 SO4 . Silica sol was spread on the silicon substrate and solidified.
charge (PZC) is the potential corresponding to the smallest
value of Cd .
2.3.2. Protein adsorption measurement
A quantity of 18 ml of 0.1 mol/L KH2 PO4 –Na2 HPO4
buffer solution (pH 7.80) was purged with purified nitrogen for 30 min to remove oxygen and then was added to
the cell. The purified nitrogen was imported to prevent oxygen movement through the nitrogen hole. The chip was
rinsed with distilled water and dried with nitrogen gas. Imaging ellipsometry was adjusted after the gold-coated silicon
wafer was added in the cell. Then the potential was controlled at −0.599 V (vs SCE). A quantity of 1 ml fibrinogen (19 mg/ml) dissolved in 0.1 mol/L KH2 PO4 –Na2 HPO4
buffer solution (pH 7.80) was added though the sample input
hole. The fibrinogen adsorption was measured by imaging
ellipsometry in real time.
2.3. Methods
3. Results and discussion
2.3.1. Measurement of potential of zero charge (PZC)
If the counterelectrode was a platinum black electrode,
and a high-frequency alternating current was added, the
equivalent electrical circuit of the electrode is shown in
Fig. 2. R is the ohmic resistance between the working and
the reference electrode, Cd is the capacitance represented by
the constant-phase element. The impedance transfer function
for the one-time constant circuit can be written as
3.1. Measurement of the potential of zero charge (PZC)
Z=R−j
1
.
2πf Cd
(1)
Z expresses a real number and Z an imaginary number,
so that
Cd =
1
.
2πf Z (2)
The value of Cd can be calculated [23,24]. In 0.1 mol/L
KH2 PO4 –Na2 HPO4 buffer solution (pH 7.80), the impedance of gold coating was determined in a different potential. The frequency is 962 Hz. The values of Cd were
calculated to draw the V –Cd graph. The potential of zero
479
The potential of zero charge (PZC) was determined by
AC impedance in a different potential. Fig. 2 is the potential
V –Cd graph. The potential corresponding to the smallest Cd
value is −0.60 V (vs SCE), and so it means that the potential of zero charge (PZC) of gold-coated silicon is about
−0.60 V (vs SCE).
3.2. Measurement of open circuit potential
The open circuit potential of the gold-coated electrode
was 0.10 V (vs SCE), which was higher than the PZC. This
result indicated that the gold-coated electrode was positively
charged in 0.1 mol/L KH2 PO4 –Na2 HPO4 buffer solution
(pH 7.80).
3.3. Contact angle measurement
Water contact angles were measured at 25 ◦ C for a dried
gold-coated silicon wafer with the sessile drop method.
Fig. 2. The potential V –Cd graph and the equivalent electrical circuit 0.1 M KH2 PO4 –Na2 HPO4 (pH 7.80), 962, amplitude 5 mV.
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Y. Yu, G. Jin / Journal of Colloid and Interface Science 283 (2005) 477–481
Deionized water (4 µl) was gently dropped onto the surfaces
and the contact angle was read directly with a goniometer.
The contact angle for the gold-coated silicon surface was
about 75◦ . This means that the gold-coated silicon wafer was
hydrophobic.
3.4. Measurement of fibrinogen adsorption
According to the method mentioned above, the adsorption of fibrinogen was measured on a potentiostatic surface
and a nonpotentiostatic surface in real time at the same time.
Fig. 3 shows plots of gray scale vs adsorption time. The gray
scale indicates the density of protein on the surface. Two
tangents were made in Fig. 3, whose slope denoted the rate
of fibrinogen adsorption on the surface. It was shown that
the rate of fibrinogen adsorption on the potentiostatic surface was faster than that on the nonpotentiostatic surface.
An electrostatic double layer exists on the solid/solution
interface. The gold-coated surface is positively charged in
0.1 mol/L KH2 PO4 –Na2 HPO4 buffer solution (pH 7.80).
Therefore the anion should be on the other side of the double layer. The fibrinogen is negatively charged in 0.1 mol/L
KH2 PO4 –Na2 HPO4 buffer solution (pH 7.80). When the
surface was potentiostatic near PZC, the positive charge of
the gold-coated surface was reduced to make the electrostatic double layer weak. The rate of fibrinogen adsorption on
the potentiostatic surface was faster than that on the nonpotentiostatic surface. The main reason for fibrinogen adsorption may be that the protein could rearrange easily to a more
stable position onto an uncharged surface. The reduction of
the electrostatic charge on the surface could accelerate the
adsorption on gold-coated silicon wafers, but the accelerated
effect was weak, so that the hydrophobic interaction should
be the major affinity for the adsorption of fibrinogen on the
gold-coated silicon wafer.
Fig. 3. The gray (surface concentration of fibrinogen) versus the adsorption time 0.1 mol/L KH2 PO4 –Na2 HPO4 buffer solution (pH 7.80),
Cfibrinogen = 1.0 mg/ml, −0.599 V (vs SCE). Curve A shows the gray of
the potentiostatic surface; curve B shows the gray of the nonpotentiostatic
surface.
4. Conclusion
The gold-coated electrode was positively charged in
0.1 mol/L KH2 PO4 –Na2 HPO4 buffer solution (pH 7.80).
The reduction of the electrostatic double-layer could accelerate adsorption on gold-coated silicon wafers, but the
accelerated effect was weak, and so the hydrophobic interaction should be the major affinity for the adsorption of
fibrinogen on a gold-coated silicon wafer.
Acknowledgments
The Chinese Academy Sciences (KSCX2-SW-322) are
acknowledged.
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