Plasma based methods to produce hydrophobic coatings (repels water)

Received: 22 March 2018
| Revised: 4 July 2018
| Accepted: 6 July 2018
DOI: 10.1002/ppap.201800059
REVIEW
Low pressure plasma modifications for the generation of
hydrophobic coatings for biomaterials applications
Kim S. Siow
Institute of Microengineering and
Nanoelectronics, Universiti Kebangsaan
Malaysia, 43600 Bangi, Selangor D.E.,
Malaysia
Correspondence
Kim S. Siow, Institute of Microengineering
and Nanoelectronics, Universiti Kebangsaan
Malaysia, 43600 Bangi, Selangor D.E.,
Malaysia.
Email: kimsiow@ukm.edu.my
Funding information
Universiti Kebangsaan Malaysia Research
Grant, Grant number: GUP-2015-039;
Malaysia Ministry of Education Research
Grant, Grant number: Hi-COE Bio-MEMS
AKU95
This review focuses on low-pressure plasma modification methods to produce
hydrophobic coatings and surface modifications on biomaterials. Plasma-deposited
fluoropolymer, siloxane, and diamond-like carbon (DLC) coatings are reviewed in
terms of process developments, monomers used, stability and aging properties, and
their behavior in adsorption of proteins, cell attachment, and bacterial adhesion.
These hydrophobic coatings are stable with correct selection of monomers and
process conditions, but the plasma polymerized siloxane and
fluorocarbons have been mainly
applied in biochip and test kits
rather than in blood-contact
applications. Similarly, the sur-
face characteristics and interfa-
cial bonding of DLC coatings
play a crucial role in their
successful implementation.
K E Y W O R D S
aging, diamond-like carbon (DLC), fluoropolymers, hydrophobic coatings, siloxane coatings
1 | INTRODUCTION
Low temperature, low-pressure (p) gas plasmas offer versatile
and convenient approaches for modifying the surface
chemistries and properties of materials with a high degree
of process control and reproducibility. Therefore, these
plasma approaches have been explored extensively in
materials surface engineering research,[1,2]
with a number
of successful translations to commercial products such as
Sumitomo “Solrox” membrane,[3]
Ciba-Vision “Day and
Night” contact lenses,[4,5]
and Becton-Dickinson “Pure-
CoatTM”
cultureware as well as Altrika “Myskin®
and
Cryoskin®
” cell-based skin regeneration therapies.[6]
Plasmas
can be classified according to whether they modify the surface
chemistry without substantial alteration to the mass and
thickness of the material being treated (plasma surface
treatment), or lead to the deposition of thin organic-polymeric
coatings (plasma polymerization), or to the ablation of a
significant amount of substrate material (plasma etching,
which will not be discussed here).
Plasmas can also be categorized in terms of the surface
properties that result from the process of polymerization. One
outcome of plasma processing can be the insertion of various
chemically reactive surface groups,[7–10]
which can then be
used to perform conventional chemical reactions for the
surface attachment of desired molecules, such as specific
biologically active entities such as proteins or aptamers, that
could otherwise not be attached onto the surfaces of solid
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https://doi.org/10.1002/ppap.201800059

materials via aqueous chemical reactions. Another aim of
plasma processing can be the fabrication of hydrophilic,
usually hydrated surfaces and coatings that are chemically
inert under the intended usage conditions[11,12]
; such passive
hydrated surfaces are of interest for various biomedical and
biotechnology applications. Examples of these applications
comprise coatings containing polyethylene glycol, sulfonate
or sulfate surface groups, and N-isopropylacrylamide, all of
which confer biocompatibility and can elicit desired bio-
interfacial interactions such as non-fouling or temperature-
dependent cell adhesion.[11]
Another type of plasma
modification is the generation of surfaces and coatings that
are concurrently chemically inert and highly hydrophobic,
which can be desirable for specific biomaterials applications.
This review will focus primarily on low-p plasma
polymerization (ppt) and, to a lesser extent, plasma surface
treatments, used to generate surfaces that are hydrophobic or
super-hydrophobic, or converted to hydrophilic via careful
selection of processing conditions and monomers, as shown
in Figure 1. As seen in Figure 1, SiOx plasma polymer (pp)
was produced from hexamethyldisiloxane (HMDSO) with
co-polymerization from oxygen to produce the hydrophilic
SiOX coatings.[13]
Several researchers have specified this
critical WCA as 90°[14]
or 65°[15]
as the threshold between
hydrophobic from hydrophilic surfaces based on the sessile
drop technique. However, this static WCA measurement is
only meaningful if both advancing and receding contact
angles are reported for plasma modified surfaces due to the
absence of stable and “equilibrium” contact angles on such
surfaces which have suffered from chemical heterogeneities
and/or topographical features.[16,17]
Hence, this review uses
the term hydrophobic as a comparative or relative adjective to
mean “water repellent,” similar to those previously de-
scribed.[17]
For those who prefer a quantitative definition of
water repellent in terms of surface energy, this reviewer
suggests a low surface energy that varies from a few to
20 mJ m−2
.[14]
This review is restricted to polymer film deposition and
surface treatment approaches performed under low-p plasma
conditions, a topic not covered in the recent reviews of
atmospheric-p plasma approaches.[15,18,19]
In addition to the
plasma polymers (pps) derived from organosilicone and
fluorocarbon shown in Figure 1, this review will include
hydrocarbon process vapors; in the case of hydrocarbons,
much interest has been focused on diamond-like carbon
(DLC) coatings. A focus on these areas of main interest serves
to bring out the key ideas and principles that also underpin
less extensively researched approaches. Fluorocarbon-based
pps are differentiated from those produced from organo-
silicone and hydrocarbon vapors by the lower degree of
hardness-elastic modulus and higher hydrophobicity of the
plasma fluoropolymer (pfp). Siloxane pps has a hardness
range of 0.3–1.0 GPa[20,21]
while DLC has a reported range of
35–60 GPa, depending on the processing technique.[22]
Therefore, the cell attachment studies reported in the
literature are influenced not only by the chemical functionali-
ties on the DLC or siloxane coatings but also by their
differences in hardness and roughness. Unfortunately,
hardness and chemical functionalities of these coatings
cannot be controlled independently; an increase of hardness
or elastic modulus often results in a change of chemical
functionalities.[23]
Instead, most studies on the relationship
between mechanical properties and cell attachment have been
carried out with model substrates (e.g., polyacrylamide and
polydimethylsiloxane [PDMS]) that are easily tunable
without changing surface chemistries.[24,25]
Due to the rich literature in this field already richly
reviewed,[26–33]
this review is selective in its choice of key
findings to be mentioned in the respective sections on
organosilicone, fluorocarbons, and DLC and in the final
summary. Recurring themes in the literature include the
complexity of the biological phenomena, the lack of common
definition and accepted test protocol, and the nature of
biocompatibility (e.g., blood or tissue compatibilities), all
obstacles that continue to hamper the proliferation and
commercialization of these technologies.[26,34]
A more
practical definition of biocompatibility is adopted in this
review: biocompatibility is the “the exploitation by materials
of the proteins and cells of the body to meet a specific
performance goal.”[35]
Note, however, that others acknowl-
edge that biocompatibility requirements are material, site, and
application specific.[36]
Furthermore, this review also shows that the continuous
improvement and availability for the past two decades of
advanced surface characterization (e.g., x-ray photoelectron
spectroscopy [XPS]) has led to greater understanding of these
plasma modification technologies. This understanding is
illustrated in this review on related process development,
aging properties, and the performance of these plasma
FIGURE 1 Hydrophobic and hydrophilic properties of
polycarbonate (PC) sheets after plasma modified with different
monomers, that is, fluorocarbon, siloxane, nitrogen (N2), and silica
(SiOx)[13]
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modifications in protein adsorption and bacterial or mamma-
lian cell adhesion for these three types of coatings. Aging
properties of pps and plasma-treated substrates are crucial
because many of the intended applications, such as products
for cardiovascular or orthopedic use, after unpredictable
periods of shelf life are exposed to biological fluids for
extended lengths of time with complex in situ movements.
Dynamic environments can induce changes in physico-
chemical and mechanical properties of plasma-modified
surfaces and coatings because such nanoscale surface cues
determine material-host tissue interactions.[37]
2 | SILOXANE PLASMA POLYMERS
Organosilicone (siloxane) polymers are well known for
their hydrophobic, water-repellent properties. A popular
method to generate siloxane polymer chemistry, as thin film
coatings, on various substrates is by ppt. Various monomers
have been used with varying degrees of success for
biomaterials applications; some of these monomers are
listed in Table 1.
In ppt of these monomers (Table 1), an increase in the
ratio of discharge power (P) to flow rate (FR) results in
extensive fragmentation, thereby producing an inorganic/
organic hybrid structure.[45,46]
The organic structural element
is similar to conventional PDMS, while the inorganic
structural element is similar to amorphous silica (SiOx).[47]
This inorganic nature is reflected in a higher polar surface
tension than in conventional PDMS polymers because of the
presence of OH groups and Si-O bonds on the surface.[45,46,48]
A decrease in the ratio of P to FR tends to retain a higher
extent of the precursor structure, which is organic in
nature.[46]
C─O bonds have been found to be formed in
siloxane pps due to reactions between trapped radicals and
atmospheric oxygen.[46]
The presence of C─O as well as
CO bonds was confirmed by Inagaki et al. when they
analyzed a range of silane─siloxane compounds at reduced
P.[45]
They concluded that at high P, the diverse constituents
of the plasma, produced by fragmentation of the organo-
silicon process vapor molecules, determine the structure of
the pp coatings while at low P it is the structure of the
monomer that influences the structure of the pp.
Methyl abstraction has been found to be the key step in the
ppt of hexamethyldisiloxane (HMDSO), with a high extent of
retained Si─O─Si structures incorporated intact into the
growing film.[45,49,50]
An increase in P also increases the
cross-linking density of HMDSO pps with extensive
formation of Si─O─Si bonds and reduced organic carbon
content.[49]
The same outcome can be achieved by pulsing the
plasma.[51]
Others have proposed that the radical surface
recombination produced stable species such as (CH3)3SiH,
TABLE 1 Types of silicon-containing monomers and studies used to evaluate the biocompatibility of pps containing silicon and oxygen
Monomer Chemical formula Application Ref
Hexamethyldisiloxane [(CH3)3Si]2O Platelet adhesion studies
Ex vivo baboon shunt
Ex vivo dog shunt
In vivo mouse model
[38]
[39]
[40]
Hexamethyldisilazane [(CH3)3Si]2NH Neurological electrode [41]
Hexamethylcyclotrisiloxane Ex vivo dog shunt
Platelet adhesion studies
[42]
[43]
Methyltrimethoxylsilane CH3─Si(OCH3)3 Platelet adhesion studies [43]
Phenyltrimethoxysilane C6H5─Si(OCH3)3 Platelet adhesion studies [43]
N-Trimethylsilylimidazole Platelet adhesion studies [43]
Tetramethylhydrocyclotetrasiloxane In vivo sheep model [44]
Tetramethylorthosilicate Si(OCH3)4 Platelet adhesion studies [43]
Tetraethylorthosilicate CH─Si(OC2H4)3 Platelet adhesion studies [43]
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pentamethyldisiloxaneand CH4 to form (CH3)xSiH during low
P deposition, while high P deposition resulted in a large
decrease in the Si(CH3)3 group, instead favoring the formation
of ─Si─CH2─Si─ bridges in the final HMDSO pps.[52]
On the other hand, however, some researchers have
reported a decrease in Si─O─Si structures with an increase of
P, basing their conclusion on Fourier transform infra-red
(FTIR) study of hexamethyldisilazane (HMDSN) pps.[53]
Figure 2 shows that other moieties such as Si─CH2─Si,
Si─CH3, Si─N─Si, Si─C, and Si─O─Si also were reduced
with an increase of P for HMDSN pp.[53]
Although methyl abstraction was considered to be the key
step in “continuous wave” (as opposed to pulsed) ppt for
monomers such as HMDSO and HMDSN, [50,53,54]
the pulsed
ppt of hexamethylcyclotrisiloxane showed complete absence
of methylene groups in the resultant pps.[55]
The hexame-
thylcyclotrisiloxane pp incorporated Si─(CH3)2 elements as
part of the growing pp network during pulsed polymeriza-
tion.[55]
Hence, it is important to also consider the effect of the
molecular structure of the monomer on the final pp.
In addition to the influence of the ratio of P to FR, other
factors which control the proportion of organic to inorganic
elements in siloxane pps are discharge frequency (f),[47]
substrate temperature (Ts),[54]
post-deposition heat treatment
(ht),[56]
and addition of gases, such as ammonia[54]
or
oxygen.[54,57,58]
An increase in Ts during plasma deposition
creates a silica-like surface.[54]
The reaction pathways in the
plasmas are likely to differ between silazane monomers, such
as HMDSN and hexamethylcyclotrisilazane (HMCTSN), and
siloxane monomers, such as HMDSO, because of the
presence of Si─NH─Si bonds in the silazane monomers[54]
;
silazane structures are reactive to water and oxygen, as
opposed to the inert nature of Si─O bonds under conditions
applicable to usage of biomedical devices and biotechnology
products. It has been observed that during deposition of
silazane pps at high Ts, Si─NH─Si bonds underwent thermal
scission of N─H bonds and formed Si─N and Si─C inorganic
structures (nitrides and carbides).[54]
In addition, the intensity
of Si─CH3 moieties decreased when the substrates were
heated to 200°C, compared to that found at room temperature,
because of thermal activation to form Si─N inorganic
linkages.[53]
In the case of HMDSO pp, the intensity of
Si─O─Si adsorption bands increased with increasing Ts,
which led to a denser structure, while the deposition rate
decreased compared to that on a substrate at room
temperature.[54]
Addition of oxygen during ppt promotes the formation of
O─Si─O structures in HMDSO pp.[51,57–60]
This effect was
amplified when the plasma on-time was reduced during
pulsed ppt of HMDSO.[51]
When the FR of oxygen was
increased during ppt of HMDSO, the pp became more silica-
like with significant quantities of O─H bonds.[59,60]
The CH3
group of HMDSO reacted with oxygen in the plasma to
produce volatile products, such as CO, CO2, H2O, OH,
HOSiCH3CH3, (HO)2SiCH3, Si(CH3)4, and Si(CH3).[57]
Some of these volatile products were detected by Lamendola
et al. with actinometric optical emission spectroscopy.[58]
Overall, the carbon content was found to be reduced,[57,58]
and carbon–oxygen functionalities were completely elimi-
nated in the presence of oxygen during ppt of HMDSO.[58]
Separately, the carbon atoms in the HMDSO pp combined
with silicon, such as C─Si,[57]
or among themselves, to form
new C─C bonds, to contribute 48% (C─Si bond) and 44%
(C─C bond) of this residual carbon for the HMDSO pp
deposited at an equivalent FR of O2 and HMDSO
monomer.[59]
The influence of plasma excitation f on the chemical
properties of siloxane pps has been investigated with
HMDSO and HMDSN monomers.[47]
An increase in the f
to the microwave range was found to encourage the formation
of silica-like products, such as Si─O, Si─N, and Si─C bonds,
in the pps. Similar to heat-treating the substrate during
deposition, a 490-fold increase in the f increased the density
of the pps by almost 46% based on analysis of gravimetric and
thickness data. This density increase was confirmed by NMR
and FTIR analyses with microwave plasma polymerized
tetramethyldisiloxane (TMDSO) with different ratios of O2
added as a concurrent process gas; the chemistry of the
resultant siloxane pps consisted mainly of ternary and
quaternary Si─O bonding.[61]
2.1 | Aging properties of siloxane plasma
polymers
The surface chemistry and surface properties of pps and
plasma-treated surfaces can undergo slow “aging” changes
FIGURE 2 Infra-red spectra of HMDSN pps deposited with
increasing discharge power at 30, 100, and 230 W on a substrate at
room temperature.[53]
In the figure, a refers to (Si─CH2─Si); b refers
to (Si─N─Si); c refers to (Si─CH3); d refers to (Si─C); e refers to
(Si─CH3); f refers to (Si─O─Si)
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with time when they are stored under ambient conditions after
plasma processing. An example of such aging is the
observation that a freshly deposited HMDSO pp had a
hydrophobic surface with an air/water contact angle of
100°,[40,48,62]
a value similar to conventional Silastic®
polymer surfaces,[39]
but after soaking in phosphate buffer
solution for 2 weeks at 37°C, HMDSO pp was found to be less
hydrophobic with a corresponding increase of the oxygen
content on the surface.[40]
The percentage of water intake
depends on the dominant bondings in this HMDSO pp: 1–2%
for the polymer-like pp and 5–13% for the silica-like pp.[60]
In the case of aging in air, siloxane pps have been
investigated with hexamethyldisilane and hexamethylcyclo-
trisilazane.[63]
Although the choice of monomers affected the
exact mechanism or reaction pathway, the aging processes
were characterized by the “formation of OH, CO,
Si─O─Si, and Si─O─C groups with the decay of Si─H
bonds.”[63]
This decay of Si─H and Si─OH bonds in the
siloxane pp has also been confirmed elsewhere.[61]
The role of
Si─O─Si was also emphasized by Hegemann et al., who
reported that ppt carried out above the critical activation
energy produced an HMDSO pp that was stable up to a year of
air aging during storage.[64]
In this stable region was found a
“high degree of linearization growth of Si─O─Si,” while the
high concentration of the methyl group was maintained.[64]
Stable water contact angles (WCAs) have also been recorded
for HMDSO─O2 pp produced from the flow ratios of
HMDSO:O2 equivalent to 1:1, 1:2, 1:5, and 1:10, but which
rapidly turned hydrophobic upon aging in air, for the pp with
flow ratio of HMDSO:O2 equivalent to 1:15.[48]
Although these researchers did not speculate on their
results, this review proposes that the role of critical activation
energy in forming O─Si─O could have played a role in the
stability of HMDSO─O2 pps produced at the lower FR of O2.
This speculation is confirmed in a separate NMR analysis of
air-aged TMDSO─O2 pp, shown in Figure 3.[61]
In
Figure 3a and 3b, the intensities of the resonance lines
assigned to Q-type bonding increased, albeit to different
degrees, for the pps produced from the TMDSO-to-O2 ratio of
0.3 and 0.18 during their air-aging period of 8 weeks.
Similarly, the insignificant difference in the intensity of Q-
type bonding was also reported for the pps with the TMDSO-
to-O2 ratio of 0.05 (Figure 3c). Q-type bonding indicates the
number of Si atoms connected to the four oxygen atoms in the
cross-linked O─Si─O structures. In other words, a coating
with a large signal of Q-type bonding is highly cross-linked
because these Si atoms are connected into the network, as
opposed to Si atoms terminated with a CH3 or a H atom.
Other monomers also tended to form the Si-O-Si bonds in
the pps during oxidation.[65]
Although Inagaki et al. did not
carry out any long-term aging studies, their systematic
variation of the “x” group in their selected monomer chemical
structure of (CH3)3Si─x─Si(CH3)3 produced pps with ease of
oxidation in the following order: bis(trimethylsilyl)meth-
ane > hexamethyldisilane > HMDSN > HMDSO.[65]
Their
deposition rates did not differ much among the four
monomers, while the Si atoms in all the pps oxidized to
Si─O─C and Si─O─Si.[65]
The role of surface restructuring on the aging behavior of
siloxane pp has been discussed by Gengenbach and Griesser
in their study of HMDSO and HMDSN pp.[50]
Their angle-
resolved XPS analysis suggested that the carbon enrichment
at the surface of both pps could be attributed to methyl group
migration to reduce the interfacial enthalpy during long-term
aging studies.[50]
During the air aging study, the HMDSO pp
also suffered from the abstraction of methyl groups, resulting
in the reduction of the C/Si ratio, and from oxidation, which
increased the O/Si ratio (Figure 4).[50]
Instead of a rapid
increase in oxygen in the first few days of aging, there was no
measurable oxidation increase because of efficient binding
between Si and Si radicals. These Si-Si bonds underwent a
variety of reactions, such as UV-induced homolysis and
hydrolysis to form the Si─O─Si bonds.
FIGURE 3 Changes in intensities of resonance lines in 29Si CP/
MAS NMR spectra over 8 weeks for TMDSO-to-O2 ratio coatings of
(a) 0.3, (b) 0.18, and (c) 0.05. For panels (a) and (c), measurements
were taken after storage periods of 24 h, and at 1, 3, and 8 weeks. For
panel (b), measurements were taken at 24 h, and at 2 and 8 weeks[61]
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In the case of HMDSN pp, their O/Si ratio increased from
0.17 to 1.15, while the N/Si ratio decreased from 0.36 to 0.05,
signaling the oxidation and the elimination of silazane groups in
the pp, respectively (Figure 5).[50]
The reduction of silazane was
evidentbasedonthereductionofN1moietiesassociatedwiththe
Si─N (Figure 5b), while some residual N2 moieties associated
with the amide bond (Figure 5b) remained in the HMDSN pp.[50]
These changes transformed the HMDSN pp to a silicone-like
surface found in a typical HMDSO pp.[50]
Separately, Figure 1
also shows that the silicone-like HMDSO pp is more resistant to
air aging than the SiOx pp based on the measured WCAs.[13]
In terms of thermal annealing, heat treatment of HMDSO
pp increased the O/Si ratio, resulting in a silica-like and highly
cross-linked structure.[56]
It was discovered in a separate
study that most of the monofunctional methylsiloxanes were
also converted to tri-functional and tetra-functional Si groups
during successive ht cycles at 100, 200, 300°C based on their
Si29
nuclear magnetic resonance with magic angle spinning
and cross-polarization analysis (NMR-MAS-CP).[56]
Simi-
larly, high-temperature ht of dichloro(methyl)phenylsilane pp
resulted in an increase of oxygen and a reduction of carbon
concentration when the ht temperature increased from room
temperature to 427°C.[66]
The vaporization of low molecular
weight methyl or phenyl and the additional crosslinking
within the siloxane bonds produced the multifunctional
silicones in the siloxane pp.[66]
However, such high thermal
deviation is neither experienced nor expected by the siloxane
plasma polymerized medical device during manufacturing or
subsequent implant in the host.
2.2 | Bio-interfacial reactions on siloxane
plasma polymers
Early cell studies on siloxane pps did not use surface-sensitive
analytical techniques, such as XPS, to characterize their
surfaces. Instead, WCA measurements and FTIR spectros-
copy were used to correlate their results with their in vitro cell
and platelet adhesion studies as well as ex vivo animal
models. The micron-deep analysis and low resolution of this
early FTIR spectroscopy, especially in the low wavenumber
regions, posed some doubts as to the positive conclusions
tabulated in Table 1. While WCA is a sensitive surface
analysis with only depth of 0.5–1.0 nm, the reported WCA
measurement in the literature is often incomplete, with sessile
drop method as the only parameter for discussion; this sessile
drop method does not provide equilibrium values due to the
chemical heterogeneity on such plasma-modified surfaces.[16]
During in vitro platelet adhesion studies, Ishikawa et al.
multiplied the number of adhering platelets with the amount
of released ATP to derive a performance indicator for the
FIGURE 4 XPS O/Si (▪) and C/Si (○) ratios as a function of
storage time for HMDSO[50]
FIGURE 5 XPS X/Si ratios as a function of storage time for
HMDSN pp. (a) O/Si (▪) and C/Si (○), (b) N1/Si (▴) and N2/Si (Δ)[50]
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different substrates.[43]
They attributed the 20–50% improve-
ment of thrombo-resistance to the chemical structure and
physicochemical heterogeneity of siloxane pps.[43]
A similar
positive result in a platelet adhesion study has also been
reported by Kiaei and Hoffman, for their plasma polymerized
HMDSO on PET coverslip.[38]
Encouraging results for
siloxane-coated polymeric substrates have also been reported
for their studies with ex vivo animal models implanted with
plasma polymerized devices produced from hexamethylcy-
clotrisiloxane[42]
and HMDSO.[38]
Although different hemo-compatibility methodologies
and medical devices have been tested with different ex vivo
animal models and in vitro tests and discussed herein, the
conclusions have been similar, that is, reduced platelet
adhesion and lack of morphological changes in the few
platelets which attached to the siloxane plasma-polymerized
surfaces. This non-thrombogenicity could be attributed to
the albumin from the whole blood adsorbed on the
substrates. Since albumin does not have the peptide
sequence to interact with platelets or enzyme receptors in
the coagulation cascade, albumin adsorption renders the
surfaces less thrombogenic. Elsewhere, in vitro multiple
protein adsorption tests demonstrated that the pre-adsorbed
albumin on siloxane surface could not be displaced by
fibrinogen or immunoglobulin because of the small size and
tenacious binding of albumins to siloxane surfaces.[67,68]
Recent studies albumin binding to hydration stratified pp
matrix (50 nm of a silica-like hydrophilic base layer with
the dosed addition of O2 gas, followed by a hydrophobic
cover layer of HMDSO pp of varying thickness) have
shown that nano-confined hydration of the deeper silica-like
sub-surface layers also influenced the albumin adsorption
and related conformation.[69]
However, Lin and Cooper have reported that the low
density polyethylene plasma polymerized with HMDSO and
bare LDPE had similar platelet adhesion and fibrinogen
adsorption results during their ex vivo animal model tests.[39]
Lin and Cooper attributed their negative results to a higher
percentage of oxygen in their HMDSO pp, but the exact
bonding sites of additional oxygen groups was not provided in
their report. Others have speculated that siloxane pp (organic-
like) are more cell-friendly than silica-like (inorganic-like)
pps; siloxane (organic)-like pps supported rat aortic smooth
muscle cell proliferation while silica (inorganic)-like siloxane
pps had a cell proliferation rate similar to that of bare 316L
stainless steels.[62]
Another siloxane compound which has been evaluated for
blood contact applications is tetramethylhydrocyclotetrasi-
loxane.[44]
A twofold reduction in thrombus formation in an in
vivo sheep model was reported 14 days after implantation.
The backbone of the pps was made of ─O─Si─O─Si─ bonds,
while the C atoms were also bonded to the Si atoms.[44]
The
correlation of molecular structure to the performance of this
siloxane treatment was neither discussed nor reported for their
in vivo test. A similar lack of detailed surface analysis for a
separate positive result on the siloxane pps has also been
reported by Tang et al.[40]
They merely reported their pp to be
made of 23% Si, 18% O, and 59% C, without any further
information on the bonding between these atoms.[40]
However, they have reported that their siloxyl-terminated
pps showed the best results among the four moieties (i.e., OH,
NH2, and CF3) tested in the chronic fibrotic responses during
their in vivo Swiss Webster mice model.[40]
In summary, the main constituents of siloxane pps are the
silica-like (SiOx) and polymer-like (O─Si─(CH3)2) groups,
which together determine the aging and bio-interfacial
properties of siloxane pps. Various process parameters like
P, FR, f, Ts, ht, and addition of co-monomer like ammonia
and oxygen play a major role in determining these
constituents of silica- and polymer-like groups.
3 | PLASMA FLUOROPOLYMER (pfp)
Generally, fluorine groups are deposited on the substrate
surfaces by either plasma treatment or plasma polymeriza-
tion. In terms of molecule structure, a monomer with a high
fluorine to carbon ratio etched the substrate, whereas a
monomer with an F/C ratio of less than 2 instead polymerized
and coated the substrate.[70]
A monomer with F/C ratio of 3
(e.g., C2F6) also favored etching over deposition during the
plasma modification.[71]
This observation has been attributed
to the polymerization route for pfp in which scission of the
C─C bonds with minimum contribution from the fluorine
detachment formed the coating, resulting in monomers with
high F/C ratios not able to plasma polymerize.[72]
Copolymerization of fluorocarbon-based monomers with
H2 has also been carried out successfully.[73]
If the feed
composition consisted of 80% H2–20% C2F6, the polymeri-
zation rate increased with increase of P to produce a highly
cross-linked structure.[73]
The polymerization rate for a
monomer with ratio of 80% H2–20% C2F6 reached a
maximum at the P of 60 W, after which the F atoms increased
by orders of magnitude to promote etching at the P of more
than 60 W.[73]
The same observation has been reported at the
critical P of 40 W for hexafluoropropylene pp.[74]
Higher P
also reduced the formation of CF3 moieties by extracting the
fluorine atoms to form CF2 during the etching process.[74]
Likewise, reduction of P retained the CF3 moieties in the pfp.
The concentration of CF3, instead of the atomic concentration
of F (at%), was found to determine the hydrophobicity or
surface energy of the surfaces, hence the importance of
controlling this CF3 concentration.[75]
In short, the chemical
composition, microstructure and cross-linking density of this
pfp depended on the plasma P, as illustrated by their two main
classes in Figure 6.[76]
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| 7 of 19

In a similar mixture of H2 and C2F6, sp3
hybridized carbon
atoms formed part of these pfp when the percentage of H
increased from 80% to a range between 88 and 95% and
electrical biasing of the substrate was set between −100 and
−150 V.[77]
At this biasing range, F atoms played a role in
stabilizing the sp3
bonds, but further biasing to −200 V
transformed the bonding to graphitic (sp2
) hybridization.[77]
Although most fluorocarbon monomers produced hydro-
phobic surfaces during ppt, chlorine- (e.g., CF3Cl), or
bromine- (e.g., CF3Br) containing fluorocarbon monomers
produced hydrophilic pps.[78]
These Cl- and Br-based
fluorocarbon monomers could be used to tailor the pps to
have varying degrees of surface energy suitable for cell
biology studies.
The influence of Ts on the surface chemical properties has
been studied with hexafluoropropylene oxide as the mono-
mer. It was found that Ts of less than 20°C favored the
formation of CF3 over C─F or C─CF moieties, while the
reverse preference in chemical moieties was deposited at the
Ts of 126°C.[79]
Similarly, the formation of CF3 moieties in
the pp was promoted at the lower Ts (temperature not
disclosed) because of the reduced fragmentation in the
hexafluoro-2-propanol, producing a coating resembling the
monomer molecule structure.[80]
The concentration of CF2
moieties also reduced as the Ts increased from −26 to 126°C
because of the dominance of the ion bombardment steps
during ppt when the Ts was elevated to 126°C.[79]
In the case of pulse-polymerization with a mixture of
CHF3 with Ar and H2O, the density of CF3 and CF2 moieties
has been shown to increase on the pfp with increasing length
of on-time[81]
while others have shown that it was the short
on-time that favored the formation of CF2 for plasma
polymerized hexafluoropropylene oxide or 1H,1H,2H-per-
fluoro-1-decene.[79,82]
These findings further differ from
those produced by perfluoro-2-butyl-tetrahydrofuran, which
did not show any simple relationship between deposition
rates, density of CF2 and CF3 moieties and duty cycles.[83]
These different findings only reflect the complexity of pulsed
polymerization that depends on the choice of monomers and
the hydrodynamics of the plasma reactors that need to be
discerned only with the rarely reported method, that is, optical
emission spectroscopy.
Besides pulsed plasma, four other plasma methods, that is,
capacitively coupled plasma, inductively coupled plasma
(ICP), plasma source ion implantation/inductively coupled
plasma (PSII/ICP) and self-ignition plasma were used to
produce pp from octafluoropropane and acetylene.[84]
Among
these five techniques, ICP produced the highest level of CF3
and CF2 moieties, followed by PSII/ICP techniques because
of their higher plasma density and lower plasma potential
which led to less dissociation than did other methods.[84]
The influence of substrate type on the pfp was also evident
from the higher F/C ratios for SF6 and CF4 plasma-treated
polypropylene (PP) compared to that of a polyethylene (PE)
substrate.[71]
This preference has been attributed to the easier
abstraction of tertiary hydrogen from PP than from PE. This
influence differed from the plasma treatment on stainless steel
that usually involves a plasma-etching step to increase the
adhesion of ppt of C2F6 and H2 onto the substrate.[85]
The
downside of such etching is the reduced resistance to aging
because of the thinned protective chromium oxide on the
original stainless steel substrate.[85]
3.1 | Aging properties of plasma
fluoropolymer
Aging studies of pfps have been carried out at ambient[86]
or
under accelerated conditions.[87,88]
Regardless of the envi-
ronment, the aging mechanisms were similar in certain
respects. In ambient conditions, the rapid initial oxygen
uptake of plasma polymerized perfluoro1,3 dimethyl cyclo-
hexane (C8F16) within the first day was similar to those
observed for hydrocarbon pps produced from alkanes and
alkylamines monomers.[86]
This observation suggests that
these different pps had a similar density of radicals capable of
reacting with in-diffusing O2 molecules.[86]
Subsequent
oxidation steps were markedly reduced for pfps because of
the absence of hydrogen in the pfps that hampered the
conversion of peroxy to hydroperoxide radicals.[86]
The WCA
for C8F16 pp was reduced only for the first 2 months, but the
XPS spectra continued to show uptake of oxygen for
2 years.[86]
Depending on the starting monomers, pfps showed
different durations of stable WCA ranging from 20 days
for hexaflourobenzene pfp to 120 days for perfluorohexane
pfp.[89]
The stability of the density of radicals within these
plasma polymers at 100°C has also been confirmed elsewhere
with electron spin resonance study for 18 h.[90]
In terms of
long-term air aging, this stability result implied that surface
FIGURE 6 Schematic diagrams of Teflon-like coatings (a)
network structure with variable F/C ratio (2 ≥ F/C > 0) and high
crosslinking, (b) ordered chain structure with ─CF3 surface groups,
high F/C ratio, very low crosslinking and surface energy[76]
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| SIOW

restructuring was restricted to the sub-surface regions beyond
the detection range of the WCA measurement. Although the
surface was continually enriched with CF3, defluorination,
especially of CF2 groups, proceeded, albeit at a slow rate. This
minimum surface restructuring showed that this pfp was well-
suited for applications needing long-term stability.[86]
Defluorination behavior was also shown by C3F6 pp when
subjected to aging at 65°C/85% RH for 700 h.[87]
A critical P
during ppt influenced the oxidation resistance of this pp. If the
P was above this critical value, the monomer would be
extensively fragmented, resulting in a high density of active
sites, which encourage extensive oxidation.[87]
Alternatively,
the O/C ratios did not change significantly when the P was
below this critical condition because the residual radicals
were terminated by the available higher mass species.[87]
Although this critical P was said to be 100 W in this study, it
differed from one plasma reactor to another as well as
according to the types of monomer used.
The loss of hydrophobicity of the pfps has also been
studied with plasma co-polymerization of perfluoropropane
(C3F8) and 3,3,3 trifluoropropylmethyldimethoxylsilane
coated on silicon substrates that had been soaked in different
media and temperatures, that is, methanol (25°C), propylene
glycol (60°C), or water (60°C).[88]
Despite the different aging
environments, the contact angle decreased rapidly because of
oxidation of residual carbon free radicals trapped in the pps.
However, the researchers claimed that diphenylamine and
heptafluorobutyric anhydride could be used to arrest this
hydrophilic recovery by inhibiting the free radicals and
acylating the OH groups produced with fluorinated anhy-
drides to prevent this oxidation from taking place.[88]
Others have used a vacuum annealing step at 100°C to
slow the aging of the perfluorocyclobutane pp.[91]
Vacuum
annealing cracked the outermost CF3 of the pfp to CF2 and CF
moieties to reduce over-layer formation and maintain their
high surface energy.[91]
In order to eliminate the influence of
atmospheric oxygen, an in-situ XPS analysis has also been
carried out directly on as-deposited octafluorocyclobutane
and trifluoromethane pps.[92]
Both pfps suffered from
defluorination, that is, decrease in CF3 and CF2 concen-
trations during the x-ray irradiation duration of more than
4000 min, although CF2 concentration and F/C ratio of
trifluoromethane pp decreased less than did those of the
octafluorocyclobutane pp.[92]
This decrease differed from
direct x-ray irradiation of polytetrafluoroethylene (PTFE)
substrate, which showed an increase of CF3 moieties
produced by chain scission in the PTFE substrate.[93]
Other research on the stability of this pfp simulated the
sterilizing procedure used in the intraocular lenses (IOL)
industry, that is, 120°C, 1.5 bar, 21 min.[94]
After this
sterilizing procedure and drying at 50°C for two days, the
stability of this pfps has been found to depend on the plasma
deposition conditions and type of monomer used; microwave
plasma polymerized perfluoroethane has been found to be
more stable than RF plasma polymerized perfluoropropane
coatings deposited on poly(HEMA-co-MMA) IOLs.[94]
This
review suggests that such ambiguity might be attributed to the
mobility of the different segments of the fluorinated
substrates when the segment compositions were elucidated
from the component fitting of the XPS high-resolution spectra
taken at different incident angles.
3.2 | Bio-interfacial reactions on plasma
fluoropolymers
Early studies on the biocompatibility of pfp showed favorable
results, as shown in Table 2. Positive results were reported for
the pfp tested with the ex vivo baboon femoral shunt
model[38,95]
and in vitro platelet studies.[38]
The lowest
platelet adhesion count occurred on the plasma fluoropol-
ymers that adsorbed the highest amount of denatured
fibrinogen.[96]
Tight binding of fibrinogen affected its ability
to communicate with the platelet receptors. This finding
agreed with those of fibrinogen adsorption studies on
perfluorohexene pp, which also showed the highest percent-
age of sodium dodecyl sulfate (SDS) non-elutable fibrino-
gen.[40]
Although perfluorohexene pp was not considered
super-hydrophobic, this pp had a WCA of more than 125°.[40]
This finding concurred with the earlier theoretical work
suggested by Ikada, who postulated that zero work of
adhesion could be achieved either by extreme super-
hydrophilicity or by hydrophobicity.[97]
Others have claimed that hydrophobicity was not the
controlling factor, but that an optimum surface energy of 20–
30 mJ m−2
would reduce the protein absorption on any bio-
interfacial surface.[107]
However, such assertions are compli-
cated by orientation and concentration of protein adsorption at
the biomaterial interface that included attachment, detach-
ment and conformational changes in an aqueous environment.
Despite similar WCA and surface chemistry, the pfps still
possesseddifferentroughnessandmorphology(Figure7).[108]
In the case of protein adsorption and cell adhesion, the surface
roughness and morphology played a bigger role than WCA or
surface chemistry because the osteoblast cells adhered more
favorably on the rougher surface of ribbon-like morphology
than on the surfaces with other morphologies.[108]
Unfortu-
nately, the influence of protein adsorption per se on these
different morphologies has not been reported.[108]
Another
study reported that the molecular structure of the monomer
influenced the surface roughness; the double bonds in
perfluoro(2-methylpent-2-ene) and perfluoro(4-methylpent-
2-ene) increased their surface roughness significantly over
that of the pp produced by perfluorohexane.[109]
Regarding cell adhesion, the preferential adsorption of
albumin from serum appeared to block the matrix protein
deposition or mask its recognition by adhering cells (e.g.,
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| 9 of 19

fibroblast cells), leading to reduced adhesion and prolifera-
tion.[81,103,105,110]
In the case of epithelial cells (e.g.,
RINm5f), the cell's inability to adsorb Ca2+
resulted in
non-adherence on the CHF3 plasma-treated surfaces.[103]
However, these non-adhesive cells were also affected by the
duration of the testing; a shorter test of 4 h resulted in better
adhesion than the 48-hr test, which resulted in non-adhesive
cells.[81]
This time dependency has been attributed to the
conformation change of the proteins and to the expression or
suppression of the extra-cellular matrix during the cell
adhesion test.[81]
In an in vivo test, Clarotti et al. carried out the ppt of
perfluorohexane on a polyhydroxybutyrate (PHB) substrate,
with two separate carrier gases, Ar and Ar-H2.[106]
When
these plasma fluoropolymerized PHB and untreated
PHB substrates were implanted in the peritoneum of
Wistar rats, scanning electron microscopy (SEM) and
anatomic-pathological analyses showed that the “bio-com-
patibility” of PHB pp deteriorated slightly more than did that
of untreated PHB.[106]
These results were independent of the
type of carrier gas used during the plasma deposition. The
presence of hydrogen in the Ar carrier gas produced a rough
surface (roughness not quantified in the literature) with F/C
ratios of 0.25–0.3, much lower than the F/C ratios of 1.5–1.6
detected for perfluorohexane pp produced with pure Ar as the
carrier gas.[106]
Similar fibrotic response tests have also been reported by
Tang et al.[40]
with their in vivo test with Swiss Webster mice
implanted for 2 weeks, shorter than the 3-month tests
conducted by previous researchers.[106]
Considering the
amount of implant-associated hydroxyproline, Tang et al.
showed that pfp-deposited PET was comparable with
untreated PET.[40]
Therefore, the beneficial effect of the
pfp over that of the underlying substrate was not apparent in in
vivo tissue studies. PHB was itself considered an inherently
biocompatible surface that did not benefit from the pfp while
the plasma fluoro-polymerized PE did show an improvement
over the untreated PE in the earlier test. However, crucial
information on the surface chemical (e.g., F/C ratio, CF3
concentration) and morphological properties (e.g., roughness
and morphology) have not been reported for these PE
substrates.[106]
Furthermore, recent in vivo implant studies in
rabbit with pfp from tetradecafluorohexane did not provide
detailed surface analysis beyond speculative results obtained
from ATR-FTIR spectroscopy.[111]
Hitherto, it is this lack of conclusive study on the
biocompatibility of pfp that has led to its uncommercializ-
ability,tothebestofmyknowledge,fortissue-relatedorblood-
contact applications.[112]
This uncertainty can be attributed to
the types of tests and choice of markers that have led to
different conclusions. For example, when Sefton et al.
investigated a series of biomaterials with different surface
chemistries for their hemocompatibility, they found that
TABLE 2 Fluorine-containing monomers and studies evaluating the biocompatibility of plasma fluoropolymers
Monomer Chemical formula Application Ref
hexafluoropropylene oxide/tetrafluoroethane C3F6O/C2F4 Osteoblast cell adhesion studies [98]
Tetrafluoroethylene CF2 = CF2 Protein adsorption on vascular graft [95,99]
Tetrafluoroethene C2F4 Cell adhesion studies [100]
Tetrafluoromethane CF4/(H2) Bacterial adhesion
Platelet adhesion test
[101,102]
Trifluoromethane CHF3/Ar Cell adhesion studies [81,103]
Perfluoropropane C3F8 Protein adsorption and cell adhesion studies [94,104,105]
Perfluorohexane C6F14 In vitro protein adsorption, cell and
blood compatibility studies
In vivo tissue compatibility studies
[106]
Perfluorohexene C6F12 In vitro protein adsorption studies
In vivo cell compatibility studies
[40]
FIGURE 7 Static water contact angle for three different plasma
fluoropolymers as a function of root mean square roughness[108]
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| SIOW

CF4-treated PE and PEU-F had poor thrombus resistance as
indicated on the C3A complement test, but other tests failed to
reveal the lack of biocompatibility of fluoro-related sub-
strates.[113] 1
As a result, technology commercialization of pfp
has focused on the bio-chip testing sector; the strong binding of
a protein to pfp could serve as a method for immobilizing an
antibody to a substrate for immunoassay purposes.[114]
For
example, two groups, studying specific cell interactions,
plasma-polymerized pentafluorophenyl methacrylate with 1,7
octadiene before introducing biotin-streptavidin conju-
gates[115]
or peptide IKVAV[116]
to the surface.
In the case of monocytes or macrophages like BMMO,
IC-21, RAW264.7, J774A.1, pfp with a WCA of 114°
supported cellular adhesion and proliferation during a long-
term test of more than 24 h.[105]
These adhesion results were
similar whether the pfp were tested directly in serum cell
culture or preadsorbed with serum/pure protein before the
bacteria adhesion test. The differences between fibroblast and
macrophages has been attributed to the different cell adhesion
receptors, integrins and matrix proteins that facilitated the
adhesion and proliferation process by these two cell lines.[105]
When the hydrophobicity of the pfp increased to super-
hydrophobic range, with a WCA of 156°, the surfaces became
resistant to bacterial adhesion because their nano-textured
surfaces could trap air that reduced the surface areas for
protein adsorption and subsequent bacterial adhesion.[117]
This lack of adhesion was also visible in the pfps that offered
few sites for bacterial adhesion in the dynamic flow test.[102]
In summary, the bio-interfacial and aging properties of
pfp depend on process parameters, such as P, Ts, co-
monomers (e.g., H, Ar, H2O), duty cycles, and type of
substrates. The stability of pfp was enhanced when it was
produced below certain critical P during plasma polymeriza-
tion. Furthermore, pfp could also be stabilized by inhibiting
its radicals with fluorinated anhydrides to prevent oxidation.
Vacuum annealing also maintained pfp stability by cracking
the outermost CF3 of the pfp to CF2 and CF moieties to
maintain the pfp's high surface energy. The bio-interfacial
properties of pfp have shown early promise as hemo-
compatible coatings but have been unable to yield conclusive
results in clinical trials. Instead, pfp has found success in
biochip applications as an immobilization platform.
4 | DIAMOND LIKE CARBON (DLC)
COATING
Although diamond-like carbon (DLC) has been under study
since 1971,[118]
its active application in the biomaterial field
was relatively short.[119]
The motivation to investigate DLC
as a biocompatible coating arose from its inert nature,
superior wear resistance, lubricant effect, and corrosion
resistance, all of which are essential for arthroplasty and
cardiovascular (particularly heart stent) applications.
DLC does not have any specific composition, instead
consisting of crystalline and amorphous phases with sp2
and
sp3
bonding. If hydrogen is present in DLC, the coating is
known as an amorphous hydrogenated alloy.[120]
DLC is also
known as amorphous carbon, ion-bombarded carbon,
diamond-like hydrocarbon, hydrogenated amorphous carbon
or amorphous hydrogenated carbon or amorphous carbon
hydrogen film. If the percentage of sp3
bonding in the
amorphous carbon or amorphous carbon hydrogen film is
very high, these forms of DLC are usually known as
tetrahedral amorphous carbon or hydrogenated tetrahedral
amorphous carbon, respectively.[120]
This review has focused on the relationship between the
chemical bondings of DLC and their aging properties for the
cell or bacterial adhesion and protein adsorption studies on
DLC. Furthermore, in addition to the general reviews
mentioned in section 1, other reviews have been written on
their processing parameters,[120]
such as ion bombardment,
biasing condition, deposition[121]
or Ts,[122]
httemperature, UV
or ion-beam irradiation and design of equipment.[120,123–126]
In terms of deposition technologies, Figure 8 shows various
techniques used to deposit DLC on biomaterials, but not all
methods can be used to deposit DLC on polymeric biomaterials
because of their high deposition temperatures. Other metallic
and ceramic biomaterials, however, may benefit from these
different processing routes. For the polymeric biomaterials, the
1
Meanwhile, fluoro-related bulk polymerized medical device has found
more commercialization success for hemo-compatible application than
their plasma fluoropolymerized device has (Ref: http://www.
interfacebiologics.com). In the bulk polymer approach, the fluoroligomer
surface-modifying additive, known as Endexo™ technology, is currently
used by AngioDynamics Inc. to build their FDA-approved peripherally
inserted central catheter (BioFlo PICC) and implantable port (BioFlo
Port). Arkis Biosciences used the same Endexo™ technology to build
their ventricular drainage catheter, CerebroFlo™. (Ref: https://www.
prnewswire.com/news-releases/arkis-biosciences-achieves-fda-clearance-
of-its-cerebroflo-evd-catheter-with-endexo-technology-300522094.html)
At this stage, it is unclear what could be the technological reasons for the
different commercialization outcome between the pfp and bulk-polymer-
ized “Endexo™” technologies.
FIGURE 8 Plasma-based techniques used to deposit DLC on
substrates. PIII refers to plasma immersion ion implantation
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| 11 of 19

operating temperature reduction was made possible by
controlling the duty cycle of the pulsed discharge[127]
or using
ion beam-assisted deposition.[128]
Plasma-assisted CVD or ppt
is one of the few proven routes to consistently produce DLC on
a polymeric substrate.[127]
Unlike other pps that used complex monomers, the
deposition of DLC coatings used simple hydrocarbons, such
as acetylene[102,119,129]
with Ar[130–132]
or H2,
[102]
or He,[133]
butane,[119]
propane,[119]
hexane,[134]
a mixture of methane
and helium,[135]
or a mixture of methane and hydro-
gen.[127,136]
Besides liquid monomers, Ar+
sputtering of
graphite targets has also been used to create DLC by
bombarding graphite with a CHn+
beam from methane
precursors.[128,137]
DLC has been doped with other elements,
that is, fluorine,[138–140]
silicon,[138,141–143]
titanium,[144]
vanadium,[144]
CaO,[145,146]
and nitrogen,[143,147,148]
for
various applications to alter the atomic and chemical structure
to attain the bio-compatibility of DLC.
4.1 | Aging properties of DLC coatings
Surprisingly, the aging behavior of DLC coating has not been
investigated thoroughly, possibly because of its perceived
inert nature. As-deposited DLC is hydrophobic,[130,132,147]
but it is easily tunable to a different degree of hydrophobicity-
hydrophilicity by adding elements like nitrogen,
oxygen, silicon, and fluorine during the DLC deposition
(Figure 9).[149,150]
Furthermore, Garguilo et al. also used
nitric acid etching to oxidize the nitrogen-doped DLC to
further decrease the WCA.[147]
As mentioned in the
introduction, the approach using static WCA alone over-
simplifies or even misleads the readers about its accuracy.
Therefore, this review includes these data (Figure 9) to serve
as a guide on the influence of these elements on the surface
energies of DLC coatings. Among the different techniques
mentioned in Figure 8, the PIII approach produced the most
stable DLC because the penetrating ions deposited their
energy in collisions with electrons and atoms to create highly
reactive chemical groups, that is, radicals, to form a densified
and cross-linked DLC.[151]
Similarly, Ostrovskaya et al. oxidized and increased the
surface energy of DLC by air-annealing the coating at 500°C
for 30 min.[136]
Others showed an increase of hydrophobicity
in the DLC coating after vacuum annealing, also at 500°C for
30 min.[132]
One possible explanation for the difference in
surface energy for these two coatings is that the air-annealing
step introduces O2 to confer hydrophilicity to the DLC
coating produced by Ostrovskaya et al. On the other hand,
vacuum annealing promoted film graphitization and hydro-
gen effusion that has been reported to be the cause of
hydrophobicity due to the formation of sp2
bonding in
DLC.[136,147]
Besides accelerated aging and ht studies, researchers have
also studied the air aging behavior of titanium (Ti)-containing
DLC coating under ambient conditions for 80 days.[152]
Oxidation was reported for this DLC coating based on the
emergence of a CO bond in their component-fitted XPS
spectra. TiO2 and TiC0.6 were also detected in the 7 at% Ti-
DLC coatings.[152]
Similarly, VC and V2O5 were also formed
in the V-DLC coating when the coating was exposed to
ambient air.[144]
In the case of aqueous aging, the interfacial shear strength
of DLC coatings lessened when immersed in bio-fluids for
1 month; the greatest reduction in interfacial shear strength
occurred for the coating immersed in artificial salivas,
followed by those immersed in phosphate-buffered saline
(PBS) and finally, those immersed in 50% fetal calf serum in
PBS.[141]
This decrease in strength has been attributed to fluid
penetration through the nano-pores in the DLC coating that
was not detected by atomic force microscopy (AFM) or
scanning electron microscopy analysis.[141]
4.2 | Bio-interfacial reactions on DLC coatings
A number of cell adhesion and protein adsorption studies have
investigated the biocompatibility of the DLC coatings, but the
lack of physical and chemical analysis performed on those
coatings in early studies resulted in inconclusive outcomes.
The literature shows that DLC coating has found two main
applications, namely, blood contact implants (e.g., heart stent
and valve), and load- or wear-reduction applications (e.g.,
joints). In the earlier reported studies, research involved other
cells and bacteria; Thomson et al. showed that DLC-coated
and uncoated Linbro culture plates had comparable levels of
macrophage and fibroblast cell activity for 7 days.[119]
Roughness or morphological analysis was not reported on this
relatively thick coating in spite of the 1-h long plasma
deposition.
Others have shown the influence of substrate on the
morphology of DLC deposition. AFM investigation has
shown the presence of “woven” morphology on DLC-coated
FIGURE 9 Change in water contact angle with addition of
elements in DLC coatings (at%)[149]
12 of 19
| SIOW

polystyrene (PS) and poly(methyl methacrylate)
(PMMA),[127]
but not on DLC deposited at a higher
processing temperature onto another stainless steel sub-
strate.[153]
Elsewhere, SEM examination also has shown
evidence of etching on a polycarbonate membrane during
DLC deposition.[154]
However, cell adhesion was not affected
by the different morphologies of those DLC-coated sub-
strates[127,154]
because the cells could easily penetrate into
very shallow (≤1 μm) or wide (≥5 μm) microgrooves,[155]
when these DLC-coated stainless steels (10 ± 2 nm[130]
) and
polycarbonate (16–40 nm[156]
) were relatively smooth.
Although cell attachments were higher on the DLC-coated
substrate than on the uncoated one, their growth rates were
similar.[127,154]
Similarly, non-toxic behavior was exhibited
by DLC-coated titanium alloy with fibroblast cells, as per
ISO10993-5 standard.[129]
Further transmission electron
microscopy analysis showed no difference in cell growth
morphology between the DLC-coated and uncoated
polycarbonates.[154]
Another group used an immunofluorescence technique to
study monocyte and macrophage growth on DLC-coated
glass coverslips because these cells offered the advantage of
studying and imaging the cytoskeletal elements within the
cells; no significant difference was reported for the DLC-
coated and uncoated substrates.[135]
Similar results were
obtained for the cell tests conducted using the Alamar blue
assay, MTT assay, and measurement of the production of
hydrogen peroxide to indicate the metabolic activity of the
cells on the different substrates.[157]
Similarly, there were no
significant differences between DLC-coated and polyure-
thane-coated stainless steels.[157]
In order to mimic the host environment, Schaub et al.
tested DLC-coated titanium with an in vitro parallel plate
flow chamber, inspecting their results in “real time” with
fiber optics and fluorescence microscopy to quantify the
platelet adhesion.[158]
Their results showed that the number
of adhering platelets on DLC-coated Ti lay between those
reported for Ti alloy and pyrolytic carbon. Dynamic flow
has also been used to study bacterial Staphylococcus
epidermis adhesion on the DLC-coated polyvinyl chloride
(PVC) substrate.[102]
Bacterial adhesion on DLC-coated
PVC was further reduced by the addition of silver, a known
anti-microbial element.[102]
The positive evaluation of
DLC-coated polyurethane was repeated in another static
bacteria Escherichia coli adhesion test.[133]
These research-
ers attributed these encouraging findings to the optimum
thickness and defined refractive index, which was shown to
depend on the favorable ratio of sp3
and sp2
carbon bonds in
the DLC coating.[133]
However, several factors like F and
surface roughness could also have played a role in
modulating bacterial adhesion on DLC-coatings.[102,133]
Others have suggested that DLC coatings with reduced
Raman ID/IG spectra would also have reduced platelet
adhesion, though no reasons were provided in their
report.[159]
Platelet and granulocyte adhesion tests showed a
reduction on DLC-coated PMMA intraocular lenses
(IOL).[128]
In the same IOL study, the researchers reported
that granulocyte and platelet adhesions decreased with
increasing proportion of sp3
bonds in the DLC-coating.
These findings agreed with those of other researchers who
conducted platelet adhesion studies on annealed[132,141]
and
highly biased [131]
DLC coatings. These two different
processing steps induced the formation of sp2
bonds, causing
an increase in platelet adhesion.[131,132]
In other words, the
decrease in hemocompatibility has been attributed to the
“increase of electrical conductivity” induced by the sp3
bonding within the graphite of the DLC coating.[131,132]
The
influence of roughness was ruled out by their AFM analysis
that showed minimum changes, regardless of high bias
deposition or subsequent high ht temperature.[131,132]
Hauert et al. showed that the addition of F or Si into the
DLC did not affect fibroblast cell proliferation because the
state of Si and F as Si─C and C─F bonds in the amorphous
matrix of DLC neutralized their toxic effects.[138]
Others have
postulated a silicon oxy-carbide bonding state for these
elements, which resulted from natural oxidation, but such
characteristics also were found to depend on deposition
route.[160]
Similar positive platelet adhesion tests results have
been reported for F-doped DLC[139]
and Si-doped DLC.[141]
The good properties of Si-DLC have been attributed to the
increased formation of sp3
bonds,[160]
although an upper
saturation limit of Si concentration was found for this
coating.[141]
The increased formation of sp3
also resulted in
reduced hardness in Si-DLC.[161]
This reduction in hardness
correlated with a reduction in residual stresses, but also in
increased beneficial adhesion to the Si-DLC coatings.
The presence of Si atoms in DLC also negated the
influence of ht, thus rendering it suitable for adhesion of
human microvascular endothelial cells (HMEC).[141]
Human
retinal pericytes, on the other hand, showed similar growth
behavior on Si-DLC and tissue culture polystyrene.[142]
It
should be mentioned here that primary cell culture is more
adhesion-sensitive than subsequent cell lines.[162]
Hence,
caution should always be exercised when comparing findings
of cell adhesion studies from different publications for
different cell lines and cell types.
Although studies of amorphous hydrogenated silicon (a-
Si:H) coating were not within the scope of this review, we
note that a-Si:H was often used as an interlayer to promote the
adhesion of DLC with the underlying substrate.[163]
Insignif-
icant differences in the lactate dehydrogenase assays were
detected between this DLC-(a-Si:H) coated composite
structure and uncoated glasses during in vitro cell adhesion
tests.[163]
Other coatings which showed promise as interlayers
for the DLC coatings were TiN and TiC.[164]
Although these
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composite coatings did not show any adverse effects in the
hemo-compatibility test, TiN and TiC showed slight
thrombus formation toward the end of incubation tests.[164]
Others used functional gradient interphases to promote
adhesion of the DLC coating to the Ti substrate to avoid
any sharp interface which deteriorated the adhesion
strength.[165]
Osteo-integration is one of the key factors to be
investigated in the study of DLC coatings; osteoblast cells
were found to have thrived better on DLC compared to on
their base silicon substrates.[137]
The introduction of nitrogen
into DLC increased adhesion of fibroblast cells[148]
and
endothelial cells[143]
over the level of adhesion on un-doped
DLC coatings. These researchers attributed their findings to
the polarization of C─N and N─H bonds in the DLC coatings
that bonded electrostatically to the proteins and cells, though
no surface analysis was carried out to confirm these
bonds.[148]
Others have attributed the excellent properties
of nitrogen-doped DLC to the optimum ratios of sp3
/sp2
and H
concentration.[143]
Elsewhere, it has been demonstrated that
PIII-produced DLC maintains a radical-rich carbonized
surface layer that immobilizes bioactive protein molecules
covalent.[151]
The cell-surface interaction was slightly different for Ti-
and V-incorporated DLC coatings because of their subse-
quent oxidation to TiO2 and V2O5 at the surface,[144]
although their carbide equivalence was also detected in
DLC matrices. While the incorporation of Ti into DLC
enhanced osteoblast differentiation and reduced bone
resorption, the addition of V inhibited the activity of
bone marrow cells. This difference has been attributed to
the leaching of V ions from V-DLC into the cell culture
media, while Ti-DLC did not suffer from any ionic Ti
leaching. However, this study did not investigate the
influence of V or Ti on the formation of sp2
and sp3
, which
could have also influenced DLC biocompatibility.
Besides solitary elements, compounds such as
CaO─H2O have also been co-deposited with the DLC
coatings, seeming to encourage the formation of sp2
crystallites to promote the viability of fibroblast cells.[145,146]
These findings compare previous results showing the
importance of sp3
bonds in improving cell interac-
tions.[128,131,132]
Although the surface roughness between
CaO-doped and undoped DLC was comparable, the formation
and role of CaCO3 in the CaO-doped DLC has not been fully
investigated within the context of an optimum ratio of sp3
/sp2
bonds to confer biocompatibility properties to DLC coatings.
There is considerably less information for in vivo testing
of DLC coatings. One in vivo study involved the implantation
of DLC-coated stainless steels into chest muscles and tibia
bones of guinea pigs for 52 weeks.[166]
Their substrates were
electrolytically polished before mplantation. Although corro-
sion products and patho-morphological changes were not
noticed in the animals, the implant showed typical bio-inert
reaction, that is, encapsulation by connective tissue built
from fibrocytes and collagen fibers.[166]
In another study,
Tang et al. implanted free-standing DLC and control
samples, such as Ti and stainless steel into the intra-
peritoneal regions of mice.[167]
Seven days after implanta-
tion, the DLC showed the minimum inflammatory response,
comparable to the responses seen on the stainless steel and
Ti implants.[167]
In the sample preparation steps, DLC
samples were etched with a mixture of H2SO4 and H2O2
solution to dissolve the silicon substrate before implanting
the samples into the animal models. Etching was found to
oxidize the DLC coatings, increasing their surface energy
above that found in their intrinsic properties. The surface
preparation steps of these early in vivo DLC tests may have
caused the findings not to reflect the intrinsic biocompati-
bility of DLC coatings.
In order to avoid these treatment-related artefacts,
Dowling et al. implanted as-deposited DLC-coated and
untreated stainless steel cylinders into bone and muscle sites
of sheep, as per the ISO/CEN 10993-6 standard.[129]
Examinations were carried out after 4 and 12 weeks.
Histological evaluation showed that the DLC coating did
not elicit any inflammatory reaction. Allen et al. found
similar positive results when they implanted their DLC-
coated and un-coated cobalt-chromium alloy in the trans-
cortical sites of a sheep and into intramuscular locations of
several rats for 90 days.[134]
Other positive results have also
been demonstrated with DLC-coated zirconium implant and
F-DLC-coated stainless steels implanted in Wistar rats for
30[168]
and 84 days, respectively.[140]
In human body
implantation, the success of DLC-coated steel in assisting
the healing of bone fracture without eliciting any
inflammation for 7 months has also been demonstrated,
but not fully understood.[169]
Although the in vitro and in vivo results published in the
literature appeared encouraging, a DLC-coated femoral
head failed at a significantly higher rate than those coated
with alumina during clinical trials with 202 patients because
of interfacial delamination between the DLC coatings and
the substrates during their follow-up period of 8.5 years.[170]
Other studies attributed the delamination failure of this DLC
coating to the slow bio-corrosion process, that is, crevice
corrosion and stress corrosion cracking, of the adhesive
interlayer in the DLC coatings.[30]
A similar result was also
reported for the heart stent application; no significant
differences were found in restenosis rate between DLC-
coated heart stent and stainless steel of similar design in 347
patients (520 lesions) during their 6 months of follow-up
check.[171]
Such results may not reflect the lack of benefits
for DLC-coated biomaterials, but instead the need to control
the processing conditions of the DLC coatings to ensure
excellent interfacial adhesion, as well as the need to
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characterize the atomic structures of these DLC coatings
that may differ across processing conditions, hence, the
importance of having an interlayer coating to increase the
adhesion of DLC to the substrates.[163]
In summary, the aging property of DLC coating depends
on the process technique employed, alloying elements
(e.g., V, Ti), post-deposition annealing temperature and
environment. PIII produced the most stable, densified and
cross-linked DLC. Air and vacuum annealing produced
different surface energy on the DLC coating arising from
differences in level of oxidation of the DLC coatings in the
presence of atmospheric oxygen. The inert nature of DLC
coating has been found suitable for blood-contact and wear-
reduction applications, but its successes have been limited by
the interfacial adhesion properties of DLC on the substrates of
the medical devices.
5 | CONCLUSIONS AND OUTLOOK
This review has focused on the processing conditions, bio-
interfacial interactions and aging properties of plasma-
polymerized organosilicone, pfp and DLC coatings produced
by plasma polymerizing and plasma treatment of various
substrates. Although these three hydrophobic coatings can be
produced easily with existing processes and equipment, their
reliability and stability have depended on the careful selection
of monomer, processing routes and parameters, such as P, Ts,
ht conditions, co-monomer, and deposition conditions.
The siloxane pps consist of mixtures of silica-like (SiOx)
and polymer-like (Si─C─Si) components that confer unique
chemical properties and stability to these coatings. HMDSO
and HMDSN are probably the most researched monomers to
produce siloxane pp; their aging properties differ slightly
because of the labile Si-N bonds in the latter, but both pps
aged to become silicone-like surfaces.
The hydrophobicity of pfp depends on the morphology,
roughness and density of the CF3 moieties instead of on the
fluorine concentration (F at%) per se. Hence, it is important to
use the relevant surface analytical technique, such as XPS and
AFM, to characterize these properties during process
development. The aging behavior of pfp was somewhat
similar to those of hydrocarbon pps with an initial uptake of
oxygen, but reduced at a later stage because the absence of
hydrogen hampered the conversion of peroxy to hydroperox-
ide radicals. The stability of these pfp also has been found to
depend on the critical P controlling the termination of residual
radicals by the higher mass species.
Although the biocompatibility tests, such as platelet
adhesion, on the siloxane pp and pfps were favorable, their
field applications have been focused on biochip and test kits
instead of blood-contact implants. Siloxane pp and pfp
probably derive their initial anti-thrombogenic properties as
the preferential adsorption of albumin became non-stable
during long-term implantation.
The biocompatibility of the DLC coatings also derives
from the chemical bondings in the DLC coatings with their
inert and smooth surfaces. While various chemical factors
such as the ratio of sp2
to sp3
and Raman ID/IG spectra have
been postulated to be the source of their biocompatibility,
contradictory results have also been widely reported in the
literature. During what is to the best of my knowledge the
only widely reported field trial, the DLC failed at the
interfacial bonding to the substrate, not because of any bio-
chemical properties of the DLC coating itself. Furthermore,
existing information has suggested the DLC coatings to be
susceptible to oxidation upon exposure to air aging, high-
temperature ht or acidic etching. While the DLC interfacial
strength also decreased when exposed to prolonged bio-
fluid incubation, some early success in using multiple Ta
layers (i.e., Ta(CoCrMo)0.5–2.0/alphaTa/Ta carbide) as the
interlayer to promote adhesion and to reduce bio-corro-
sion = induced delamination has been reported.[172,173]
However, the chemical bonding and microstructure of these
DLC coatings have, sadly, not been reported in most open
literature to provide insights into their failure mechanisms
to enable improvement in the next generation of DLC
products.
Another issue demanding industry attention is the
influence of mechanical properties on cell attachments.
While the relationship between mechanical properties and
cell attachments is quite established for model substrates
like polyacrylamide, the same cannot be said of DLC
coatings because of the influence of dopants like Si or SiOx;
insignificant differences in cell attachments were observed
in the DLC coatings whose hardness varied from 11 to
16 GPa[161]
although others have reported otherwise with
different testing conditions.[174]
Hence, this review has
emphasized the importance of surface physical-chemical-
mechanical analysis in the development of any surface-
modified biomedical devices for implant application.
ACKNOWLEDGMENTS
The author acknowledges financial support from Malaysia
Ministry of Education research grants Hi-COE Bio-
MEMS AKU95 and Universiti Kebangsaan Malaysia
Research Grant GUP-2015-039 for this work. The author
also thanks Alena Sanusi for editorial comments on the
manuscript.
ORCID
Kim S. Siow http://orcid.org/0000-0003-2519-780X
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| SIOW

K. S. SIOW is a research fellow at the
Institute of Micro-Engineering and
Nanoelectronics, as well as an asso-
ciate fellow at the Center for
Collaborative Innovation, Universiti
Kebangsaan Malaysia (UKM). His
multi-disciplinary research interests
are related to plasma surface modi-
fication, sintered silver bonding and patent circumvention.
Before joining UKM, he worked as a materials engineer in
multi-national companies and National University of
Singapore, as well as a technology transfer officer at the
commercialization arm of Singapore A*STAR research
institutes. Besides materials engineering education at
University of South Australia (PhD) and Nanyang
Technological University (MASc and BASc (Hons)), he
also completed his Master of Laws in Intellectual Property
at the University of Turin-WIPO program. In addition,
K. S. Siow is a registered Chartered Engineer (UK
Engineering Council) with Project Management Profes-
sional PMP® and International TRIZ Association
(MATRIZ) Level 3 certifications.
How to cite this article: Siow KS. Low pressure
plasma modifications for the generation of
hydrophobic coatings for biomaterials applications.
Plasma Process Polym. 2018;e1800059,
https://doi.org/10.1002/ppap.201800059
SIOW
| 19 of 19

Plasma based methods to produce hydrophobic coatings (repels water)

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