Biomedical Engineering Reference
In-Depth Information
protein adsorption, infl ammation, and infection can interfere with the long-term use of implants.
The rationale for surface modifi cation of biomaterials is straightforward: retain the key physical
properties while modifying only the outermost surface to infl uence biointeraction [79]. The ulti-
mate goal of building such an interface is to improve the biocompatibility, reduce tissue-implant
reaction, and prolong the lifetime of a device.
As a general rule, nonspecifi c adsorption of proteins should be minimized and benefi cial mol-
ecules should become selectively adsorbed onto biomaterials as a result of modifi cations performed
before their implantation [80]. A hydrophilic coating with lubricious property on implants is mostly
preferred for prolonging the lifetime of the product. Application of LbL self-assembly to deposit
hydrophilic thin fi lms on implants has the following advantages: (1) broad selection of materials;
(2) precise control over coating thickness within nanometer range with designed fi lm structure; and
(3) no necessity of surface pretreatment in some cases. Ideally, alternation of only the outermost
molecular layer (3-10 Å) should be suffi cient, but thicker fi lms are required to ensure a full cover-
age on medical implant [79]. During LbL self-assembly, ordered multilayer fi lms can be built from
nanometer to micron range with defi nite knowledge of their molecular composition. A nonlinear
fi lm growth often occurs at the beginning of the alternating assembly process, and the fi rst two to
three layers have smaller amounts of adsorbed polyions [24,43]. The fi lm mass of subsequent layers
increases linearly with the number of adsorption cycles. The desired fi lm thickness depends on the
coating material, implant surface roughness, and biological environment of an implant. Appro-
priate fi lm thickness is important to provide both biointerface function and maintain the surface
morphology and the mechanical property of the implant.
LbL self-assembly has been applied to modify medical devices, and implants include contact
lens [81] and vascular stents [57,58]. Here, we are focusing on the thin fi lm coatings on stents. Stent
implantation is widely used for the treatment of occlusive blood vessel diseases with the reduction
of restenosis. However, stent implantation is also associated with excessive proliferation of vascular
SMCs, extracellular matrix synthesis, and chronic infl ammatory reaction, which are believed to
be initiated by deep vascular injury and further enhanced by the presence of a foreign metallic
device [82]. Modifi ed surface should have improved biocompatibility with less thrombogenic and
infl ammatory reactions. Polysaccharide-based nanocoating of either (PEI/heparin)
n
or (CH/HA)
n
on endovascular devices was recently developed through LbL self-assembly [57,58]. Both hyal-
uronic acid (HA) and heparin are glycosaminoglycans. HA is a naturally occurring linear, high-
molecular weight anionic polymer (p
K
a
=
2.9) consisting of alternating
N
-acetyl-β-d-glucosamine
and β-d-glucuronic acid residues linked (1
4), respectively. Its high-molecular mass
and numerous mutually repelling anionic groups make hyaluronate a rigid and highly hydrated mol-
ecule, which, in solution, occupies a volume
→
3) and (1
→
1000 times than in its dry state [83]. The inhibitive
effects of HA with respect to hyperplasia observed after either systemic or local delivery suggest
that the antiproliferative effects of HA may be associated with its antiinfl ammatory properties
[84]. Heparin is a variably sulfated glycosaminoglycan that consists predominantly of alternating
α(1
∼
4)-linked residues of d-iduronate-2-sulfate and
N
-sulfo-d-glucosamine-6-sulfate. It has an
average of 2.5 sulfate residues per disaccharide unit, which makes it the most negatively charged
polyelectrolyte in mammalian tissues and widely known as an anticoagulant agent [83].
PEI/heparin pairs were used in precursor establishment during LbL self-assembly of polyelec-
trolytes on stent materials NiTi and 316L stainless steel [57]. It has been found that in PEI/heparin-
paired multilayers, the contact angle alternatively shifted from
→
25° for a
heparin layer. It is a clear indication of fi lm buildup through LbL self-assembly. After immersing
the PEI/heparin multilayered fi lm in Tris-HCl buffer (pH 7.35) for 3 weeks, the coating was still
stable as verifi ed from electrochemical impedance measurements and contact angle studies. In
another study, the contact angle of a HA layer decreased from 40° to 30° when the number of HA
layers increased from one to four [58]. Further increase of HA layer numbers did not change the
contact angle value. It was suggested that in this case, at least four bilayers were necessary to mask
the substrate with respect to the properties of the multilayer-water interface. Linear fi lm growth
∼
59° for a PEI layer to
∼
