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Universidade de Lisboa
Faculdade de Farmácia
Antifouling surfaces as a strategy to prevent
microorganism’s adhesion on medical
devices
Ana Beatriz Lopes Roque dos Santos Marques
Mestrado Integrado em Ciências Farmacêuticas
2017
Universidade de Lisboa
Faculdade de Farmácia
Antifouling surfaces as a strategy to prevent
microorganism’s adhesion on medical
devices
Ana Beatriz Lopes Roque dos Santos Marques
Monografia de Mestrado Integrado em Ciências Farmacêuticas
apresentada à Universidade de Lisboa através da Faculdade de Farmácia
Orientador: Professora Doutora Isabel Ribeiro, Professora Auxiliar
2017
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Resumo
As infeções associadas aos cuidados de saúde, ou infeções nosocomiais, ou ainda
infeções hospitalares são um grande problema clinico, com impacto significativo na
mortalidade e morbilidade e representam um elevado custo económicos para os
sistemas de saúde. Na europa, o número estimado de doentes com infeções nosocomiais
em 2013 foi de 4.2 milhões. Algumas das infeções hospitalares estão associadas ao
usado de dispositivos médicos, como tubos endotraqueais, cateteres urinários e
venosos. As práticas da medicina moderna levaram a um aumento da utilização destes
dispositivos médicos invasivos. O aumento da sua utilização representa um fator de
risco para o desenvolvimento destas infeções. A contaminação dos dispositivos
médicos pode ocorrer por inoculação de microrganismos provenientes das mãos de staff
médico e clinico. Contudo, é mais provável que ocorra por inoculação de
microrganismos provenientes da pele ou mucosas do paciente aquando da implantação
do dispositivo. Estas infeções resultam da interação entre fatores dos microrganismos,
dispositivo médicos e do hospedeiro. No entanto, a capacidade dos microrganismos
para aderir à superfície dos materiais e promover a formação de um biofilme é o fator
mais importante na patogénese da infeção.
Os biofilmes são comunidade estruturadas de microrganismos que conseguem aderir e
crescer em superfícies bióticas e abióticas. Desta forma, estão protegidos do ambiente
externo, incluindo componentes do sistema imunitário e antibióticos.
Quando se suspeita de uma infeção associada a dispositivos médicos, o tratamento
passa pela remoção ou substituição do dispositivo e/ou iniciar terapêutica
antimicrobiana. Contudo, a suscetibilidade dos microrganismos aos antibióticos está
diminuída. Assim, é mais lógico apostar em estratégias que previnam a formação de
biofilmes.
Apesar dos fatores associados aos microrganismos serem os mais importantes na
patogénese da infeção, os fatores associados ao dispositivo médico são os mais
modificáveis. Fatores como o tipo e forma do dispositivo podem favorecer a adesão
microbiana. Assim, a investigação tem se focado em estratégias que visam a
modificação do dispositivo de forma a prevenir a adesão de microrganismos,
colonização e consequente formação de biofilmes. As estratégias preventivas podem
ser distinguidas em estratégias anti-adesivas e antimicrobianas. As estratégias
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antimicrobianas baseiam-se na ligação superficial ou incorporação de substâncias
antimicrobianas, como antibióticos, metais ou compostos de amónio quarternário.
Dispositivos médicos funcionalizados com agentes antimicrobianos estão disponíveis
comercialmente, contudo, a sua utilização deve ser monitorizada devido ao risco de
toxicidade de desenvolvimento de resistência aos agentes utilizados. As estratégias anti-
adesivas e baseiam-se na alteração das propriedades físico-químicas dos dispositivos,
de modo a modificar as interações especificas e inespecíficas entre os microrganismos
e a superfície do dispositivo. A adesão dos microrganismos às superfícies é influenciada
pelas propriedade físico-químicas da superfície, como a carga e a hidrofobicidade. É
também influenciada pelas condições do meio, como presença de proteínas e o pH.
Assim, é necessário que estas superfícies consigam, também, inibir a adesão de outras
biomoléculas.
As propriedades anti-adesivas podem ser conferidas pela modificação química ou física
(da topografia) da superfície. Ainda assim, não é possível desenvolver superfícies com
adesão zero.
Esta monografia foca-se maioritariamente no estudo das diferentes estratégias anti-
adesivas, obtidas por modificação química, usadas em dispositivos médicos.
Como estratégia anti-adesiva têm vindo a ser utilizados os polímeros anti-adesivos têm
sido utilizados, os polímeros poli-hidrofílicos e os poli-zwiteriónicos. Ambas as classes
de polímeros são compostas por polímeros eletricamente neutros, que conseguem
diminuir as interações electroestáticos entre proteínas carregadas e os dispositivos
médicos. Além disso, formam uma camada de hidratação à superfície que funciona
como uma barreira física e energética, que previne a adsorção de proteínas. A camada
de hidratação é formada por ligações de hidrogénio entre as moléculas de água
presentes no meio e os grupos funcionais do dispositivo, no caso dos polímeros poli-
hidrofílicos. Contudo, no caso dos polímeros poli-zwiteriónicos esta é formada por
interações electroestáticas. As superfícies anti-adesivas podem ser obtidas por
diferentes métodos, tais como monocamadas auto-organizadas, revestimento com
hidrogel e polímeros em escova.
O polietilenoglicol é um polímero neutro e hidrofílico e é o polímero mais usado e
investigado para construir superfícies anti-adesivas. Este é utilizado para prevenir a
adsorção de proteínas e a adesão de bactérias, tendo demonstrado conseguir inibir
5
também a adesão de leveduras. O sucesso da sua ação depende do método de produção
utilizado e da arquitetura do polímero. Apesar do PEG ser considerado uma referência
no que diz respeito aos polímeros anti-adesivos, a sua ação, quando em contacto com
fluidos fisiológicos, é diminuída ou mesmo perdida, uma vez que sofre degradação
oxidativa quando na presença destes fluidos. Assim, o PEG não é apropriado para uma
utilização a longo prazo. Por este motivo, outros polímeros e materiais têm sido
estudados como eventuais alternativas. Algumas dessas alternativas passam pelo uso
de polivinilpirrolidona, polibetaína, poli-oxazolinas, acrilatos polihidroxifuncionais,
poliacrilamida, heparina e fosforilcolina. Estes revestimentos têm demonstrado, em
testes realizados in vitro, bons resultados a inibir a adsorção de proteínas e a adesão de
microrganismo e consequentemente a formação de biofilmes em testes realizados in
vitro.
O PVP é um polímero biocompatível, altamente hidratado e consegue repelir proteínas.
Revestimentos com PVP conseguem inibir a formação de biofilme de algumas bactérias
como E. coli. Devido às suas propriedades anti-adesivas, este polímero é utilizado como
revestimento em dispositivos médicos, nomeadamente em cateteres urinários e
cateteres venosos centrais.
Por sua vez, polímeros de acrilatos polihidroxifuncionais também demonstraram a
capacidade de reduzir a formação de biofilmes de algumas estirpes bacterianas,
nomeadamente S. epidermidis e P. aeruginosa. Uma outra vantagem deste polímero é
ser capaz de reduzir a adsorção de fibrinogénio e plaquetas. Além disso, apresenta
elevada estabilidade mecânica e biológica. Devido a estas características, revestimentos
como estes polímeros são utilizados em dispositivos médicos como cateteres urinários.
As POXs possuem propriedades anti-adesivas semelhantes ao PEG, no entanto estas
apresentam uma maior estabilidade em meio fisiológicos e oxidativo. Testes in vitro
demonstraram que as POXs conseguem reduzir a formação de biofilmes de E. coli.
Devido às suas propriedades anti-adesivas entes polímeros podem ser utilizados como
revestimento de dispositivos médicos como próteses, implantes, entre outros.
Em diferentes testes in vitro, dos acrilatos polihidroxifuncionais demonstraram que
conseguem reduzir a adsorção de proteínas, como o fibrinogénio, e reduzir a adesão de
bactérias. Estes polímeros podem ser aplicados a diversas superfícies, podendo,
nomeadamente, ser aplicado como revestimentos de lentes intraoculares.
6
Os polímeros de poliacrilamida demonstraram boas propriedades anti-adesivas em
testes realizados em condições in vitro. Estes revestimentos conseguem inibir a
adsorção de proteínas e células e diminuir a formação de biofilmes tanto de bactérias
como de fungos. Polímeros de poliacrilamida podem ser utilizados em dispositivos
médicos como lentes de contacto, cateteres urinários e próteses vocais.
A heparina é um polissacarídeo natural, que devido às suas propriedades
anticoagulantes e anti-adesivas têm sido amplamente utilizada para revestir dispositivos
médicos, nomeadamente em dispositivos que estão em contacto com o sangue, como
cateteres e enxertos vasculares.
A fosforilcolina pode ser usada em polímeros anti-adesivos, uma vez que consegue
reduzir a formação de biofilmes de várias estirpes bacterianas e fúngicas. Polímeros
com fosforilcolina podem ser aplicados a diversos dispositivos médicos, tais como
dispositivos de fixação óssea, stents coronários, pulmões artificiais, entre outros.
Considerando que os polímeros anti-adesivos não conseguem inibir totalmente a adesão
de microrganismos, é possível que estes que consigam aderir à superfície dos
dispositivos formem um biofilme maduro. Assim, a conjugação de polímeros com
propriedades anti-adesivas com agentes microbianos é uma abordagem promissora para
o desenvolvimento de superfícies antibacterianas. Existem vários revestimentos que
conjugam ambas as propriedades, nomeadamente polímeros de PEG funcionalizados
com antibióticos.
Os vários materiais abordados nesta monografia conseguem formar dispositivos
médicos com superfícies anti-adesivas e assim ajudar a combater as infeções associadas
a dispositivos médicos. No entanto estes têm características distintas entre si e por isso
o material escolhido para revestimento um dispositivo médico deve ser selecionado
tendo em consideração a sua aplicação especifica.
Palavras-chave: Infeções nosocomiais; Dispositivos médicos; Adesão bacteriana;
Biofilmes; Superfícies anti-adesivas.
7
Abstract
Healthcare-associated infections are a major clinical problem with significant impact
on mortality and morbidity and represent an economic burden of health systems. The
number of patients with a HAIs, in 2013, in Europe, was estimated at 4.2 million. The
use of invasive medical devices, such endotracheal tubes, urinary catheters, central
venous catheters, mechanical heart valves and others represent a risk factor for these
infections. Such infections are related to biofilm formation onto medical device
surfaces. Biofilms are a structured community of microorganisms that can attach and
grow on abiotic and biotic surfaces and are protected against antimicrobials and
immune system components. Thus, the best approach against such infections is the
prevention of biofilm formation.
This monographic work aims to study the different antifouling strategies used to prevent
biofilms formation on medical devices.
The preventive approaches can be divided into antifouling and antimicrobial strategies.
Antimicrobial strategies are based on the incorporation or surface bonding of
antimicrobial substances, such as antibiotics. Antifouling strategies are based on the
modification of basic polymers to develop polymers with new surface properties. PEG
is the gold standard of the antifouling polymers however, has been reported that is not
appropriate to long-term applications. Thus, several materials such PVP, POXs,
polybetaines, poly (hydroxyfunctional acrylates), PAAm, heparin and
phosphorylcholine have been studied for anti-adhesive applications and showing to be
able to reduce protein adsorption, cell adhesion and bacterial attachment and
consequently reduce biofilm formation. Nevertheless, more studies are necessary to
assess their efficacy under real physiological conditions.
Various antifouling polymers are already being applied in medical devices such as
catheters, orthopedic implants, dental implants, vascular grafts, contact lenses, among
others. The drawback is that despite these polymers being able to reduce/delay biofilm
formation on medical device surfaces, a few microorganisms can eventually adhere and
form a biofilm. Another interesting approach is the possible association of theses
coatings with antimicrobial agents to improve the antibiofilm properties of the material.
8
Antifouling coatings are different from each other, thus the type of coating should be
adapted for a specific application.
Keywords: Healthcare associated infections; Medical devices; Bacterial adhesion;
Biofilm; Antifouling surfaces.
9
Abbreviations
AM - Ampicillin
Amino-PPX – poly(o-amino-p-xylylene-co-p-xylylene)
ATP – Adenosine 5’- triphosphate
ATRP - Atom transfer radical polymerization
BFG – Bovine fibrinogen
BSA – Bovine serum albumin
CHT – Chitosan
CLSM – Confocal laser scanning microscope
CSF – Cerebrospinal fluid
DFS – Dynamic force spectroscopy
ECDC – European Centre for Disease Prevention and Control
EGDA – Ethylene glycol diacrylate
EPS – Extracellular polymeric substances
ePTFE – Expanded polytetrafluoroethylene
FDA – US Food and Drugs Administration
Fg – Fibrinogen
GEN – Gentamicin
HAIs – Healthcare Associated Infections
HP - Heparin
ICU – Intensive care unit
IOLs – Intraocular lenses
LbL – Layer-by-layer
Lyz – Lysozyme
MA - Maleic anhydride
10
MAPC-co-BMA - Poly [(ω-methacryloyloxyalkyl phosphorylcholine-co-n-butyl
methacrylate)]
MPC – 2-methacryloyloxyethyl phosphorylcholine
OD – Optical density
OEG – Oligo (ethylene glycol)
PAAm – Polyacrylamide
PBS – Phosphate-buffered saline
PC – Phosphorylcholine
PCBMA – Poly (carboxybataine methacrylate)
PDMS – poly(dimethylsiloxane)
PEG – Poly (ethylene glycol)
PEN – Penicillin
PEO – Poly (ethylene oxide)
PEOXA – Poly (2-ethyl-2-oxazoline)
PET – Polyethylene terephthalate
PHEAA – Poly (N-hydroxyethylacrylamide)
PHEMA – Poly (2-hydroxyethyl methacrylate)
PHPMA – Poly N-(2-Hydroxypropyl methacrylamide)
PLL-g-PEG – Poly(L-lysine)-graft-poly (ethylene gylcol)
PLL-g-PMOXA – Poly(L-lysine)-graft-poly(2-methyl-2-oxazoline)
PMB – poly (MPC-co-n-butyl methacrylate)
PMMA – Poly (methyl methacrylate)
PMOXA – Poly (2-methyl-2-oxazoline)
POEGMA – Poly (oligo (ethylene glycol) methyl ether methacrylate)
POXs – Poly (2-oxazoline)s
PP – Polypropylene
11
PPO – Polypropylene oxide
PPOXA – Poly (2-phenyl-2-oxazoline)
PRP – Platelet rich plasma
PS – Polystyrene
PSA – Poly (3-sulfopropyl methacrylate potassium)
PSBMA – Poly (sulfobetaine methacrylate)
PU – Polyurethane
PVDF – Poly (vinylidene fluoride)
PVP – Poly (vinylpyrrolidone)
SAM – Self-assembled monolayer
SE – Silicone elastomers
SEM – Scanning Electron Microscopy
SPR – Surface plasma resonance
TMC – N-Trimethyl chitosan
TT – Tympanostomy tubes
UV-VIS – Ultraviolet - Visible
XPS – X-ray photoelectron spectroscopy
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Index
1. Introduction .......................................................................................................... 13 1.1 Healthcare-associated infections .................................................................. 13
1.1.1 Medical device-associated infections ...................................................... 13 1.2 The role of biofilms in device-associated infections ................................... 14 1.3 Strategies to control bacterial biofilm ................................................................ 15
1.4 Strategies to prevent biofilm formation ............................................................. 15 2. Objective .................................................................................................................. 17 3. Materials and methods ............................................................................................. 17 4. Results ...................................................................................................................... 17
4.1 Antifouling surfaces ........................................................................................... 17
4.1.1 Poly (ethylene glycol) and PEG-based materials........................................ 19
4.1.2 Poly (vinylpyrrolidone) ............................................................................... 23 4.1.3 Polybetaine .................................................................................................. 25
4.1.4 Poly(2-oxazoline)s ...................................................................................... 28 4.1.5 Poly (hydroxyfunctional acrylates) ............................................................. 30 4.1.6 Polyacrylamide ........................................................................................... 32 4.1.7 Heparin ........................................................................................................ 34
4.1.8 Phosphorylcholine....................................................................................... 36 4.2 Antifouling and antimicrobial surfaces .............................................................. 38
4.2.1 PEG-Antibiotics .......................................................................................... 38 4.2.2 Heparin/Chitosan multilayer film ............................................................... 40 4.2.3 MEO2MA based copolymer ........................................................................ 42
5. Discussion ................................................................................................................ 43
6. Conclusion ............................................................................................................... 45 7. Bibliography ............................................................................................................ 46
Index of Figures:
Figure 1. Stages of biofilm formation…………………………………………………14
Figure 2. Antifouling coating…………………………………………………………16
Figure 3. Examples of surface functionalization methods……………………………18
Figure 4. Structure of Poly (ethylene glycol)………………………………………….19
Figure 5. Structure of Poly (vinylpyrrolidone)………………………………………..23
Figure 6. Structure of sulfobetaine and carboxybetaine, respectively………………..25
Figure 7. Structure of Poly(2-oxazoline)s…………………………………………….28
Figure 8. Structure of pHEAA, pHEMA and pHPMA, respectively…………………30
Figure 9. Structure of Polyacrylamide………………………………………………..32
Figure 10. Structure of Heparin……………………………………………………….34
Figure 11. Structure of Phosphorylcholine……………………………………………36
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1. Introduction
1.1 Healthcare-associated infections
Healthcare-associated infections (HAIs), also mentioned to as “nosocomial” or
“hospital infection”, are infections that patient acquire during the course of receiving
treatment for other conditions, in a hospital or another healthcare facility, and are not
present or incubating at the time of admission. (1,2) Such infections are the most
frequent adverse event in care delivery, having a significant impact on mortality,
morbidity, and quality of life, leading to an economic consequence for health systems.
(1,3) According to the European Centre for Disease Prevention and Control (ECDC)
surveillance report, the annual number of patient with a HAIs, in European acute care
hospitals, was estimated at 4.2 million, in 2013. The most frequent HAIs were
respiratory tract infections (33.6%), symptomatic urinary tract infections (22.3%), and
skin infections (21.4%). The most common microorganism associated with HCAIs
were Escherichia coli, Staphylococcus aureus, Enterococcus spp., Pseudomonas
aeruginosa, Klebsiella spp., Coagulase-negative staphylococci and Candida spp.. (4)
An important issue is that some of these infections are related with the use of medical
devices. (4)
1.1.1 Medical device-associated infections
Medical devices are widely used in medical modern practice for diagnostic and
therapeutic purposes. Nevertheless, the increasing use of this invasive devices
represents a significant risk factor for the development of HAIs. (5,6) Medical devices
such as urinary catheters, central venous catheters, endotracheal tubes, mechanical heart
valves and others are associated with high rate of infections. (2,7) According to ECDC
surveillance report, in 2013, (intensive care unit) UCI-acquire pneumonia was device-
associated in 92% of the cases, UCI-acquired urinary tract infections and UCI- acquired
bloodstream infections were reported catheter-related in 96.7% and 43.3%,
respectively. (4)
The contaminations of medical devices can occur by inoculation with microorganisms
from the hands of clinical or medical staff. However is more likely occur by inoculation
of pathogens from the patient’s skin or mucous membrane, during implantation of the
medical device. (6,8) There are many risk factors for the development of HAIs, such
14
invasive surgery, prolonged use of invasive medical devices, long hospital stay,
immunocompromised patients and other underlying patient conditions. (1,2)
The medical device-associated infections result from the interaction of three factors:
microorganism, device and host factors. The ability of bacteria to adhere to materials
and promote the formation of a biofilm is probably the most important factor in the
pathogenesis of infection. (6,7)
1.2 The role of biofilms in device-associated infections
A biofilm can be defined as a structured community of microorganisms that can attach
and grow on abiotic and biotic surfaces. They produce extracellular polymeric
substances (EPS) which protects them from the external environment, including
antimicrobials and immune system components. (2,5) In many cases antimicrobial
resistance is associated with long-term persistence of biofilm infections. (6,9) Biofilms
can be composed of single species or multiple species of microorganisms. (10)
Biofilms formations include different stages (Figure 1): reversible attachment,
irreversible attachment, colonization, maturation and dispersion. Different
microorganisms have different and specific mechanisms that allow initial surface
attachment, development and detachment from the biofilm.Figure 1 (2,11)
Figure 1. Stages of biofilm formation
15
First, free microorganism cells (planktonic cells) adhere to a surface by reversible
bonds, such a van der Waals force, hydrophobic and electrostatic interactions. (6,11,12)
The initial attachment is reversible, whereas is followed by an irreversible attachment.
The permanent attachment to the surface is stimulated by features of the microbial cell
surface (i.e. flagella, pili, fimbriae, glycocalyx and cell adhesion molecules). Once
adhered to the surface, microorganisms continue to multiply and accumulate in
multilayered cell clusters and begin to build an extracellular matrix. The matrix, along
with other microbial components, stabilizes biofilm structure and holds the biofilm
together. Finally, the aggregated cells may detach from the biofilm and proceed to
colonize other sites or cause systemic infection. (2,6,9,11)
1.3 Strategies to control bacterial biofilm
When a medical device-associated infection is suspected, a general decision is to
remove or replace the device and/or initiate antimicrobial therapy. (6) The clinical
experience with HAIs demonstrates that defence mechanisms of the host are incapable
of handle the infection. (8) Compared with free living cells, biofilms are more difficult
to eliminate because they are better protected against immune system components and
antibiotics. (10,13) The levels of susceptibility to antimicrobial agents are decreased
10-1000 times less. (6) The failure of antimicrobial therapy can result in chronic
infection. (14) Therefore, inhibiting biofilms formation is a more reasonable approach
than the possible dispersal of the formed biofilms. (10)
1.4 Strategies to prevent biofilm formation
As was mentioned above, these infections result from the interaction between
microorganism, device and host factors. Microorganisms factors are the most important
in the pathogenesis and some stages of biofilms formations can be possible targets of
preventions strategies. However, device factors are the most modifiable. Several
device-related factors may favor microorganisms adhesion such as the type of device
material, source of device material, surface and shape of the device. (7) Recently
research has been focused on the modification of the medical device, in order to prevent
bacterial attachment, colonization and, consequently, biofilm formation. (2) The
16
preventive approaches can be distinguished in antifouling and antimicrobial strategies.
(15,16) The antimicrobial strategies are based on the incorporation or surface bonding
of antimicrobial substances, such as antibiotics, metals, quaternary ammonium and
disinfectants. Medical device with antimicrobial agents are commercially available and
are already used in clinical applications. However, the use of this devices must be
monitored due to the risk of toxicity and development of antimicrobial resistance
against the agents included. (6,10,17) Another preventive approach are the antifouling
strategies. These are based on the modification of basic polymers to develop polymers
with new surface properties, such as antiadhesive polymers (Figure 2). (6,17)
Microorganisms adhesion to surfaces is influenced by the microorganism and
physicochemical properties of the surface, such as hydrophobicity and surface charge.
It is also affected by environment properties, like the presence of proteins and pH. Thus,
it necessary that antifouling surfaces can also prevent non-specific interactions with
proteins and other biomolecules. (12,18,19)
Since, the intrinsic properties of a material influences microbial adhesion is possible to
improve the surface of the device for development infection-resistant materials. Anti-
adhesive properties can be conferred by modification of the topography or chemistry of
the surface. Nevertheless, is not possible develop surfaces with zero adhesion. (6,12)
Another effective approach can also be the combination of the two previous approaches,
i.e., conjugate both antiadhesive and antimicrobial properties. (11)
These surface modification strategies, namely antifouling polymers, are a growing area,
with applications in many fields beyond medical devices and have been target of several
research studies. (12,20,21)
Figure 2. Antifouling coating
17
2. Objective
The aim of this monograph is to study the different antifouling strategies used to prevent
biofilms formation in a medical device. This work will focus mainly on antifouling
strategies by chemical modification.
3. Materials and methods
The present review is a research at antifouling surfaces used to prevent microorganism’s
adhesion on medical devices. To write this monograph, different data bases were used,
such as Web of Knowledge, PubMed, Science Direct and B-on. First, the research was
focused in a review article in the area published over the last five years. For this search,
terms such as “antifouling strategies for medical devices”, “antifouling coatings”,
“antiadhesive coatings”, “antibiofouling polymers”, “anti-infectious surfaces”,
“antibacterial coatings” and “polymers for biomedical applications” were used. After
this, a more specific research was made to find studies for each strategy selected. Thus,
as an example the search terms such as “PEG polymers for biomedical application” and
“PEG coating preventing microbial adhesion” were selected. Articles published within
ACS, Springer or Elsevier were favoured. This monograph, intends to make a review
on different studies published in literature mentioning antifouling strategies to prevent
medical device-associated infections.
4. Results
4.1 Antifouling surfaces
Antifouling strategies are based on the modification of basic polymers to develop
polymers with new surface properties. (6,17) These modifications lead to a change in
nonspecific and specific interactions between the microorganism and the material
surface. Thus, is possible to develop medical devices resistant against microbial
adhesion. (6) Two major classes of antifouling polymers coatings have been used for
repelling protein and bacteria. These classes are based on polyhydrophilic and
polyzwitterionic materials. Since, these polymers are electrically neutral, they can
reduce electrostatic interactions between charged protein domains and device surfaces.
18
The antifouling properties of these coatings are also related to the formation of a
hydration layer, near the surface. This layer prevents protein adsorption, by acting as
an energetic and physical barrier. The hydration layer, in hydrophilic surfaces, is
formed by hydrogen bonds between water molecules in the environment and functional
groups on the medical device surface. The strength of this barrier affects the
effectiveness of the hydrophilic material and depend on the physicochemical properties
of a material (e.g. molecular weight) and surface packing (e.g. polymer chain). The
hydration layer in zwitterionic surfaces is strongly bound through electrostatic
interactions. The stronger the interaction, the more resistance to adhesion is the
material. (12,15)
Manny antifouling strategies, based on these polymers, have been applied to design
medical devices, implants and biosensors with new surface properties. (22) Antifouling
surfaces can be obtained by various methods, and examples are self-assembled
monolayer (SAM), polymer brush coatings and hydrogel coatings (Figure 3). (19)
SAMs represent a highly organized molecular assembly formed spontaneously by
chemisorption followed by self-organization of long chain molecules, on the surface.
These molecules consist of three domains, i.e. head group, alkyl chain and
functionalized end region. (23,24) On the other hand, Polymer brushes are polymer
chains covalently tethered on a surface. This coating can be obtained by grafting-from
and grafting-to techniques and applied onto varied materials (e.g. silicon, titanium).
(25,26) Another technology for surface modification is hydrogel coatings. A hydrogel
is a hydrophilic polymer that can reduce microbial adhesion by formation of a soft
surface. Furthermore, hydrogel can absorb water, which also contribute to reduce
biofilm formation. (2)
Figure 3. Examples of surface functionalization methods
19
4.1.1 Poly (ethylene glycol) and PEG-based materials
Poly (ethylene glycol) (PEG) (Figure 4), a neutral hydrophilic polymer and is widely
studied as an antifouling polymer. (26) The antifouling properties of PEG or PEG-based
materials are due to the hydrophilic nature of these polymers that leads to the formation
of a hydration layer which inhibits protein attachment and microorganisms adhesion.
(22)
In fact, these coatings have demonstrate to reduce adhesion of yeasts and bacteria in
vitro. (27) The success of PEG as an antifouling polymer depends, in general, of the
polymer architecture and the surface grafting method. (22) PEG can be attached to
surfaces by different grafting techniques such as “grafting from”, “grafting to” block
copolymerization, chemical functionalization, among others. (22,24)
As mentioning above, PEG is one of the most studied antifouling polymer, one of these
studies was performed by Fernandez et al. (27), that developed a multi-component
cross-linked PEG-based polymer coating to inhibit non-specific biomolecular
adsorption, cell and protein attachment. This coating is formed by three core coating
components (i.e. the active, the matrix-forming and the intermolecular cross-linking
component) and applied to surfaces from a volatile carrier solvent, in a single step. (27)
The first component is a hetero-bifunctional PEG molecule, terminated with a
succinimidyl ester (functional group) and an alkoxysilane terminus (reactive
crosslinking group). The second component is a polyoxyethylene sorbitan tetraoleate
and the third is an azidosilane. (27)
This polymer was characterizatied by X-ray photoelectron spectroscopy (XPS), water
contact angle and streaming potential and the results showed that the surface was
hydrophilic and negatively charged. The microbial adhesion was evaluated in a parallel
plate flow chamber, with a bare glass as a control. For this study five bacterial strains
Figure 4. Structure of Poly (ethylene glycol)
20
were used, Streptococcus salivarius, Staphylococcus epidermidis 3399,
Sthaphylococcus epidermidis HBH 276, Escherichia coli and Pseudomonas
aeruginosa. Obtained results revealed that the coating reduces adhesion of many
clinically isolated bacterial strains in vitro. An exception was P. aeruginosa that
adhered similarly to uncoated and coated glass. (27) In another study was reported that
P. aeruginosa (adhesive strains) releases surface-active substances, such biosurfactants,
that can penetrate the coating. This may cause decreasing of poly (ethylene oxide)
(PEO) interfacial properties and increasing attractive interactions between bacteria and
PEO brush. (28) The small effect of the multi-component coating in inhibiting
attachment of P. aeruginosa could be related with this bacterial factor. (27)
Also the effectiveness and stability of the coating in physiological fluids were
evaluated. The polymer continues stable and effective against S. epidermidis 3399
adhesion after seven days’ exposure in urine. However, the antifouling effectiveness
decrease over time in protein-rich physiological fluids, like saliva. An additional
advantage of this three component coating is its compatibility with varied substrates,
such as metal oxide, glass and numerous plastics. (27)
The antifouling properties of this multi-component crosslinked PEG-based polymers
have also been investigated by other authors (29). The results showed that this multi-
component polymer reduced adhesion of non-specific proteins. Furthermore, it also
decreased adhesion of fibrinogen (Fg) and lysozyme (Lyz). The inhibition of bacterial
adhesion was assessed with two bacterial strains, S. aureus and K. pneumoniae, in a cell
flow. Results demonstrated that the adhesion of these bacteria to the uncoated surfaces
(i.e. bare glass and bare indium tin oxide) was strong and mainly irreversible (>60%
and >80% of the microorganisms remains in the surface, respectively). However,
almost all the bacteria cells were removed from the coated surfaces. (29)
The antifouling properties of another grafting technologies for PEG and PEG-based
materials also have been investigated. Roosjen et al (30) developed a (PEO)-brush
coating, covalently attached to silica and glass by reaction in a polymer melt. The
presence of a coating was evaluated by X-Ray photoelectron spectroscopy, and contact
angle measurement showed that coated surfaces were more hydrophilic that uncoated
surfaces. Its antifouling properties were assessed in a parallel plate flow chamber,
utilizing various microbial strains, S. epidermidis, Staphylococcus aureus, S. salivarius,
E. coli, P. aeruginosa, Candida albicans and Candida tropicalis, and compared to
21
adhesion to bare glass. After 2 h of incubation, PEO-coated side, did not have any
bacteria or yeast adhered, however the uncoated glass was covered by microorganisms.
After 4 h, the results showed a decrease in adhesion in the PEO-coated side when
compared to uncoated glass. This reduction was more than 98% for S. epidermidis, S.
aureus and E. coli and a reduction of 88% was observed with S. salivarius.
Nevertheless, the results of P. aeruginosa adhesion were not statistically significant.
This coating also reduced the adhesion of C. albicans and C. tropicalis, in 81% and
75%, respectively. (30)
In conclusion, this study showed that bacterial adhesion to PEO-brushes was greatly
decreased with respect to adhesion to glass, except when more hydrophobic bacteria
were tested, such P. aeruginosa. Comparing to bacterial adhesion, the extent of
reduction of yeasts strains adhesion was smaller. Based on these outcomes, the authors
proposed that PEO-brushes decreased the microbial adhesion by reduction of van der
Waals interactions, between microorganisms and the surfaces. (30)
In another study, the same authors investigated the influence of the polymer molecular
weight and temperature in microbial adhesion. The adhesion of S. epidermidis,
P.aeruginosa, C. albicans and S. tropicalis in PEO chain of 526, 2000 and 9800 Da
was compared to their adhesion to bare glass. The adhesion was evaluated in a parallel
plate flow chamber at 20 ºC and 37 ºC. The results show, after 4 h of incubation, that
adhesion of P. aeruginosa and C. tropicalis was only decreased in higher molecular
weight brushes due to the hydrophobic character of these microorganisms. On the other
hand, the three PEO-brushes could reduce adhesion of S. epidermidis and C. albicans.
In this study, the retention of microorganisms was also assessed by a passage of an air
bubble. The results demonstrated that almost all microorganisms adhering to the PEO
brush were removed, independently of brush length. However, the retention of the
microorganisms on uncoated glass was greatly higher. The experiments were realized
at 20 ºC and 37 ºC however, no significant differences in adhesion or retention were
found. (31)
The polymer brush-coating with PEG-based materials was also assessed by Nejadnik
et al, (32) that studied its effect on the kinetics of bacterial growth. For this, initial
attachment and succeeding 20 h growth was evaluated with three different bacterial
strains (i.e. S. aureus, S. epidermidis and P. aeruginosa) on uncoated and coated
silicone rubber, in a parallel plate flow chamber. They used a tri-block copolymer of
22
poly (ethylene oxide)-polypropylene oxide (PPO) – poly (ethylene oxide)
(PEO99PPO65PEO99), a low-density polymer brush coating, physically adsorbed to
hydrophobic silicone rubber. Initial attachment of S. aureus and S. epidermidis (after
30 min) decreased from 5.2x106 cm-2 and 4.1x106 cm-2 on uncoated silicone to 0.4x106
cm-2 and 0.1x106 cm-2 on coated silicone, respectively. On the other hand, initial
attachment of P. aeruginosa was similar on coated and uncoated silicone. The biofilms
grew on both surfaces, however, developed more slowly on brush-coatings. The strains
of Staphylococcus demonstrated a lag-time of 3 h, on uncoated silicone, after which
biofilm initiated to grow exponentially. After 8 h of growth, the entire surface was
covered by biofilm. However, on coated silicone, the lag-time was approximately 8h
and after 20 h only 71% and 60% of the surface was covered by S. aureus and S.
epidermidis biofilm, respectively. In the case of P. aeruginosa, the kinetics of bacterial
growth was completely different and the surface coverage was smaller than 24% in both
surfaces. (32)
To assess the antifouling properties after growing of a biofilm, an induced detachment
was performed by shear stress. Biofilm of Staphylococcus strains was completely
removed from the coating surface, whereas the removal from uncoated silicone was
incomplete, only 58% and 71% of the biofilm of S. aureus and S. epidermidis was
successfully removed, respectively. Biofilm of P. aeruginosa could not be stimulated
to remove, on both surfaces, by shear stress. So, the authors concluded that biofilms on
coatings developed more slowly and were more easily removed by high fluid shear.
(32)
Sustainability of the antifouling properties of the coating was also assessed. The results
demonstrated that formation and succeeding removal of S. epidermidis biofilm did not
affect the non-adhesiveness of the coating. However, part of the antifouling properties
was lost after removal S. aureus biofilm. Nonetheless, brush coating continues to be
effective against Staphylococcus strains, compared with the adhesiveness to uncoated
silicone. The loss of non-adhesive functionality could be explained by the presence of
nitrogen containing bacterial components that remain on the brushes after biofilm
removal. A higher amount was found after removal of S. aureus biofilm when compared
to S. epidermidis. These findings can be clinically significant, once the slower biofilm
formation on polymers brushes coatings, can permit more time for treatment with
antibiotics before a more resistant biofilm developed. (32)
23
Polymer coating with PEG can be used in biomaterials, and an example is titanium
surfaces. Titanium and titanium alloys are widely used as hard tissue replacements in
dental and orthopedic implants (artificial joint and bones). Are also used in prosthetic
heart valves and artificial hearts. Polyurethane (PU) is a synthetic polymer used to made
cardiovascular implants, such vascular grafts. PU surfaces can be modified using PEG
and its derivatives for additional hydrophilicity. (33–36) PEO can also be used in coated
silicone contact lenses. (37)
The drawback is that although, PEG coatings have been shown to reduce
microorganism’s adhesion in vitro, after contact with physiological fluids in vitro or in
vivo, the reduction is generally lesser or even lost. This can be due to the continuous
bulk protein assault that leads to an eventual overwhelming of the surface or coating
degradation (e.g. chain cleavage, hydrolysis, surface removal). (27,38) Furthermore, it
has been described that, under in vivo conditions, the terminal hydroxyl group of PEG
might be oxidized to an aldehyde and PEG chains, in aqueous solution, shows weak
hydrophobic interactions with proteins. Consequently, PEG may not be appropriate as
a long term antifouling surface, and other polymers are being explored as potential
alternatives. (12)
4.1.2 Poly (vinylpyrrolidone)
Poly(vinylpyrrolidone) (PVP) (Figure 5) is a highly hydrated, biocompatible, water
soluble and protein-repellent synthetic polymer. (22)
Due to its properties, PVP have been studied to originate antifouling surfaces. Antonelli
et al. (39), studied the influence of (PVP)-coated tympanostomy tubes (TT)
(commercially available) on biofilm formation of P. aeruginosa and S. aureus. Silicone
TTs with or without PVP coating were exposed to blood or phosphate-buffered saline
(PBS), for 24 h, and biofilm development was evaluated by quantitative bacterial counts
Figure 5. Structure of Poly (vinylpyrrolidone)
24
and scanning electron microscopy (SEM). The results demonstrate that P. aeruginosa,
after 4 days in culture, formed a biofilm on both PVP-coated and uncoated TTs.
However, the number of bacteria was lower in PVP-coated TT. On the other hand, PVP
coating did not demonstrate a significant effect on preventing biofilm formation of S.
aureus. Thus, they conclude that PVP coating reduced biofilm formation of some
bacteria under certain conditions. (39)
The effectivity of PVP at repelling both protein and bacteria was evaluated by Sun et
al. (40). For thus, PVP crosslinked with ethylene glycol diacrylate (EGDA) was grafted
to surfaces (i.e. polyvinylidene fluoride (PVDF) membrane), via vapor-based grafting.
They have created coatings with a different crosslinking degree. The antifouling
properties of this coating was tested using bovine serum albumin (BSA) and E. coli, as
protein and bacteria models, respectively. The results of protein adsorption test show
that uncoated PVDF membrane have the major adsorption of BSA. The PVP-co-EGDA
coating reduced BSA absorption more than 90% compared to the uncoated membrane.
The adhesion of E. coli was tested on glass surfaces coated with PEGDA and PVP-co-
EGDA. The surface coated with PEGDA exhibited similar bacteria adhesion as the
uncoated glass surface. For surfaces coated with PVP-co-EGDA with different
crosslinking degrees , it was observed that the decrease in the crosslinking degree
contributed to reduce adhesion of E. coli. (40)
Other authors have also evaluated the antifouling properties of the PVP, and an example
was Liu et al. (41), that developed a PVP-grafted poly(dimethylsiloxane) (PDMS)
surfaces by surface-initiated atom transfer radical polymerization (ATRP). The
antifouling properties of the PVP-grafted PDMS surfaces was tested by the
determination of the protein adsorption and the cell and bacteria adhesion. Fg was
chosen as a model protein and its adsorption was decreased by 86% on the PVP90-
grafter PDMS when compared with the uncoated PDMS surface, even after 30 days at
ambient conditions. All other PVP-grafted PDMS surfaces tested showed good
resistance to protein adsorption. The resistance to bacterial fouling was also tested by
cell and bacterial adhesion tests, using L929 cells and E. coli, respectively. The results
demonstrated that practically no cell or bacteria adhered to PVP90-grafted PDMS
surfaces when compared to uncoated PDMS surfaces, even after 30 days in ambient
conditions. All of the other PVP-grafted PDMS surfaces tested showed similar results.
(41)
25
Surfaces functionalized with PVP have also been selected improve the blood
compatibility of PU central venous catheters. (42) PVP hydrogels have been used on
indwelling urinary catheters, since this coating has showed reduce bacterial attachment.
(43) It has also been applied in biosensors and biochips, haemodialysis and blood
purification units. (24) PVP can be used to increase the water content of polymers
included into soft contact lenses, such as poly (2-hydroxyethyl methacrylate) (PHEMA)
, poly (methyl methacrylate) (PMMA) or silicone. (37)
Furthermore, PVP hydrogel also has demonstrated to reduce surface roughness and
decrease bacterial adhesion and protein absorption on central venous catheters. (43)
4.1.3 Polybetaine
Zwitterionic polybetaines have been studied as antifouling coatings due their
hydrophilic properties. (26) They are named according to the negatively charged group
and include sulfobetaine and carboxybetaine (Figure 6). (33)
One of these studies was reported by Cheng and co-workers (44) that have studied the
antifouling properties of long-chain zwitterionic poly(sulfobetaine methacrylate)
(PSBMA) grafted in surfaces (i.e. gold surfaces) using ATRP. They investigated their
resistance to short-term and long-term bacterial adherence and biofilm formation with
two bacterial strains (i.e. S. epidermidis and P. aeruginosa), using a laminar flow
chamber. For comparison, poly (oligo(ethylene glycol) methyl ether methacrylate)
(POEGMA) grafted surfaces (i.e. gold surfaces) were also investigated. Additionally,
they quantified how surface grafting methods would influence the long-term resistance
to bacterial attachment and the coating stability. In that way, were prepared SAMs of
alkanethiols with shorter-chain oligo (ethylene glycol) (OEG) and mixed
Figure 6. Structure of sulfobetaine and carboxybetaine, respectively
26
SO3−/N+(CH3)3 (SA/TMA) terminated groups. Methyl (CH3) SAMs grafted on gold and
bare glass was also prepared as a reference. (44)
The short-term adhesion (i.e. after 3 h) of S. epidermidis showed a significantly
decrease on POEGMA, PSBMA, OEG SAMs and SA/TMA SAMs surfaces as
compared to CH3 SAMs and bare glass. On the other hand, short-term attachment of P.
aeruginosa to OEG SAMs was reduced by 87% when compared to the glass surface.
Adhesion of P. aeruginosa on PSBMA and POEGMA was reducing in 80% and 75%,
respectively, compared to OEG SAMs. The lower adhesion on PSBMA and POEGMA
surfaces than on SA/TMA and OEG SAMs are due to the longer molecular chains of
the antifouling groups. They found that S. epidermidis and P. aeruginosa adhered more
to SAMs that POEGMA or PSBMA. This could be a result of degradation of the SAMs
with the time and/or thickness of the coating produced using brushes. The long-term
accumulation of two bacterial strains on these surfaces was studied, utilized bare glass
and CH3 SAM on gold were selected as control. After 24 h or 48 h, qualitative images
showed that PSBMA and POEGMA reduced biofilm formation of P. aeruginosa and
S. epidermidis as compared to the controls. Both PSBMA and POEGMA reduced
bacterial adhesion, thus PSBMA (zwitterionic material) can be used as an antifouling
material. (44)
In another work, Cheng et al. (45), reported a study on the resistance to long-term
biofilm development of zwitterionic poly(carboxybetainemethacrylate) (PCBMA). The
influence of various conditions, such a temperature, bacterial strain and zwitterionic
group (PCBMA and PSBMA) on the antifouling properties were investigated. The
PCBMA and PSBMA brushes were grafted from glass and gold surfaces by ATRP. The
antifouling performance of PCBMA coating was further investigated by parallel flow
cell system using two bacterial strains (i.e. P. aeruginosa and P. putida). The results
showed that PCBMA coating decreased long-term biofilm formation of P. aeruginosa
up to 10 days in 95% at 25 ºC and up to 2.5 days in 93% at 37 ºC, compared to
unmodified glass. The coating also reduced the biofilm formation of P. putida up to 8
days in 95% at 30 ºC. PSBMA surface was also evaluated by parallel flow cell, using a
strain of P. aeruginosa at 25 ºC, and compared with PCBMA surface. Results showed
that PSBMA coated surfaces could decrease biofilm formation over 9 days, in other
words, the capability of PSBMA coating to resist biofilm development was similar to
27
PCBMA. They concluded that PCBMA surfaces could significantly delay biofilm
formation. (45)
In order to investigate the blood compatibility of these polymers, Jiang et al. (46),
grafted sulfobetaine onto PU and assessed the protein absorption with bovine fibrinogen
(BFG) and platelet adhesion with platelet rich plasma (PRP). The quantification of
absorbed BFG on the both coated and uncoated surfaces was made using 125I-labeled
protein, with a radioimmunoassay. Water contact angle results demonstrated that
uncoated PU surfaces was more hydrophobic that sulfobetaine coated surfaces (78 º ±
3º, 60º ± 3º, respectively). The results showed that the adsorption of BFG onto coated-
PU was reduced by 8.8%. The results of platelet adhesion were observed by SEM and
showed that the number of platelets adhered to the surface decreased in coated-PU when
compared to uncoated-PU, after 1 h and 3 h. These tests demonstrated that films grafted
with sulfobetaine could reduce protein adsorptions and resist to platelet adhesion. (46)
The biological and mechanical stability of carboxybetaine zwitterionic hydrogel was
investigated by Hsiang-Chieh Hung et al.(47). In this study, three main terminal
sterilization techniques (i.e. ethylene oxide gas, steam autoclave and gamma
irradiation) were used. The results showed that carboxybetaine hydrogels were stable
at an oxidative gas environment and high pressure and temperature without altering
their antifouling properties. (47)
PSBMA coating was studied as a grafted layer to modify cellulose and PU surface due
to their anti-thrombogenicity properties. Nevertheless, medical devices with PSBMA,
commercially available, are rarely seen. US Food and Drug Administration (FDA) have
approved two PSBMA modified peripherally inserted central catheters. (43) Vascular
catheters of PU surfaces can be modified using polybetaine polymers, such as
carboxybetaine, phosphobetaine and carboxybetaine. (33) Furthermore, PSBMA can
be also applied to endotracheal tubes, orthopedic devices and urinary catheters. (43,48)
28
4.1.4 Poly(2-oxazoline)s
Poly(2-oxazoline)s (POXs) (Figure 7), such as poly(2-ethyl-2-oxazoline) (PEOXA),
poly(2-methy-2-oxazoline) (PMOXA) and poly(2-phenyl-2-oxazoline) (PPOXA) have
been investigated as PEG alternatives. (22)
POXs and PEG demonstrate similar antifouling properties if the optimum density of
the brush is well chosen. However, POXs have numerous advantages such as lower
viscosity, higher stability, and a less demanding synthesis. Manny surface attachment
techniques are available, such as “grafting from” and “grafting-to”. The surface
immobilization process plays a significant role in the stability and efficiency of the
coating. Two types of POXs coatings can be distinguished, depending on the structure
of the used polymer and the immobilization method (e.g. brush-like POXs or linear
POXs). Various parameters influence the properties of the coatings, such as side chain
grafting density, the film thickness and the surface density. (49,50)
The non-adhesive properties of PMOXA was compared to PEG in a study performed
by Moeller and co-workers. (51) They used niobium surfaces coated with poly(L-
lysine)-graft-poly (ethylene gylcol) (PLL-g-PEG) (2kDa) and poly(L-lysine)-graft-
poly(2-methyl-2-oxazoline) (PLL-g-PMOXA) (5kDa) of optimal grafting densities (α
= 0.29 and 0.22, respectively) and assessed the bacterial attachment of E. coli at 37 ºC
in a physiological buffer. The result demonstrated that bacterial attachment was reduced
by 99% in PLL-g-PEG or PLL-g-PMOXA coated surfaces compared to PLL coated
surfaces and uncoated surfaces. Thus, they concluded that both PMOXA and PEG-
based surface coatings could prevent bacterial attachment to a similar extent. (51)
Other authors have been investigated the antifouling properties of PMOXA. The
microbial adhesion of two genetically engineered E. coli strains to poly(L-lysine)-graft-
poly(2-methyl-2-oxazoline) (PLL-g-PMOXA) in niobium surfaces, was assessed by
Pidhatika et al (52). The results showed that dense short brushes (grafting density =
Figure 7. Structure of Poly(2-oxazoline)s
29
0.33) decrease bacterial adhesion, in greater extent, compared to PPL-g-PMOXA with
different grafting density and uncoated surfaces. This copolymer reduced adhesion of
fimbriated and non-fimbriated E. coli bacteria in both low and high (physiological
level) ionic strength however, less dense brushes did not prevent bacterial adhesion to
the same level. (52)
In another study, Pidhatika et al (50)., compared the stability of the PMOXA coating
with PEG coating in niobium oxide surfaces. In a way to perform a direct comparison,
they selected a PLL-PMOXA and PLL-PEG graft copolymers. The antifouling
properties were assessed from protein resistance test, before and after stability tests.
They concluded that PMOXA copolymer was more stable than PEG copolymer and
kept their antifouling properties under different environmental conditions (oxidative
solution, solution that mimics the ionic strength of body fluids and a solution that
mimics the physiological solution with oxidative substance). (50)
POXs polymers have been studied for biomedical applications and could, ultimately,
be applied to medical devices such as catheters, prosthesis, implants and wound
dressings. (53,54)
30
4.1.5 Poly (hydroxyfunctional acrylates)
Hydroxyfunctional acrylate, such as Poly (N-hydroxyethylacrylamide) (PHEAA),
PHEMA and poly N-(2-Hydroxypropyl methacrylamide) (PHPMA) (Figure 8) are
neutral, hydroxyl-rich monomers and have been studied as antifouling coatings. (26,55)
The antifouling properties of these polymers have been determined in different studies.
Zhao et al. (55), developed a PHEMA and PHPMA polymers brushes grafted on gold
surfaces via surface-initiated ATRP. The stability of these polymers was determinated
in function of incubation time in PBS. PHEMA polymer with ~32 nm of film thickness,
remained practically unchanged in 20 days, demonstrating a very high structural
stability. PHPMA polymer also showed a very high structural stability in 40 days in
PBS. The protein adsorption from single-protein solution (i.e. Fg, BSA, and Lyz in PBS
buffer) was also investigated and the results showed that PHEMA and PHPMA brushes
have low protein adsorption. Additionally, they also investigated the protein adsorption
from human blood plasma and serum. The results of surface plasma resonance (SPR)
showed that these polymers brushes have very low fouling to 10% human blood plasma
and serum. However, PHPMA only exhibit low fouling protein with appropriated film
thickness. Antiadhesive properties were evaluated by short-term static bacteria
adhesion. Thus, bacterial adhesion of Cytophaga lytica to PHEMA and PHPMA, after
2 h at 25 ºC, was measured. The results demonstrated that both PHEMA and PHPMA
had very low bacteria adhesion as compared to the bacterial attachment on controls (i.e.
PSA (poly(3-sulfopropyl methacrylate potassium), PSBMA and bare gold). Antifouling
properties of PHEMA were better than PHPMA and this could be explain by the fact
that PHEMA surface was more hydrophilic that PHPMA surface. (55)
Figure 8. Structure of pHEAA, pHEMA and pHPMA, respectively
31
In another work, Zhao et al. (56), developed a PHEAA grafted on gold surfaces by
ATRP and studied the antifouling properties in different complex biological media,
such as undiluted human blood plasma and serum, different single protein solutions,
and bacteria. They studied the resistance of nonspecific protein adsorption from
undiluted serum and plasma on PHEAA brushes with different film thickness by SPR
spectroscopy. The results showed that adsorption was practically undetectable on the
PHEAA brush at a film thickness of 12 nm (selected to the bacterial adhesion test). To
assessed bacterial attachment resistance of PHEAA brushes two bacterial strains were
used, E. coli and S. epidermidis, in a static bacterial adhesion assay. They also studied
the adhesion of these two strains in PHEMA and PHPMA, for positive control and bare
gold for negative control. The results demonstrate that attachment of S. epidermidis on
PHEAA, after 66 h, was reduced in approximately 59% compared to PHEMA and
PHPMA. The adhesion of E. coli on PHEAA was reduced 79%, compared to PHEMA
and PHPMA brushes. The author concluded that PHEAA brushes had better and
extended resistance to both protein adsorption and bacterial adhesion (approximately 3
days) when compared to PHEMA and PHPMA brushes. The better antifouling
properties of PHEAA could be due to the fact that this polymer contains more
hydrogen-bond donors which enhance surface hydration. This was confirmed by water
contact angle test, which demonstrate that PHEAA was more hydrophilic that PHEMA
and PHPMA (water contact angle was 15º, 31, 33º, respectively.) (56)
Poly (hydroxyfunctional acrylate) can be grafted onto various inorganic substrates, such
as silica, quartz, graphite and silver. These materials have been widely used in implants,
tissue engineering scaffolds and contact lenses. PHEMA can also be used in artificial
organs, blood-contacting implants, contact lenses and intraocular lenses (IOLs).
(22,37,56)
PHEMA hydrogel are used to coat ventricular catheters, since they reduce cell
attachment, in vitro, when compared to uncoated catheters. (57)
PHEMA hydrogel can also be used to entrap nebulized antimicrobial solutions or as a
drug delivery system in PVP endotracheal tubes and urinary catheters. (58,59)
32
4.1.6 Polyacrylamide
Polyacrylamide (PAAm) (Figure 9) is a stable, hydrophilic, biocompatible, electrically
neutral and polar polymer. (12,60) PAAm can form low-fouling surfaces due to the
surface hydration layer formed by hydrogen bond between water and amide groups.
(61)
The antifouling properties of PAAm polymers has been investigated in different
studies. One of these studies was performed by Fundeanu et al (62)., that assessed the
antifouling properties of amino-poly(o-amino-p-xylylene-co-p-xylylene) (PPX)-
PAAm brush grafting to silicone rubber by ATRP. The bacterial adhesion was studied
using two bacterial strains, S. aureus and E. coli, in a parallel plate flow chamber. The
results demonstrated that initial deposition rates and adhesion were greatly reduced for
both strains in amino-PPX-PAAm brush-coated silicone rubber compared to uncoated
silicone rubber. After 4 h, the bacterial adhesion of E. coli and S. aureus on coated
silicone rubber was reduced in 99% and 93%, respectively, compared to adhesion on
uncoated silicone rubber. (62)
In another study, a PAAm brush coating grafting on silicone rubber surfaces (water
contact angle of 109º) by ATRP and the efficacy of the coating was evaluated against
three microbial strains (i.e. S. aureus, S. salivarius and C. albicans), by parallel plate
flow adhesion. The functionalization of the silicone rubber surfaces with PAAm
allowed to create a more hydrophilic surface (water contact angle of 28º). The results
showed that microbial adhesion on coated surfaces was reduced when compared to
uncoated silicone rubber. Adhesion of S. aureus, S. salivarius and C. albicans was
reduced in 58%, 52% and 77%, respectively. The detachment of all adhering bacteria
and yeast from the brush-coating confirmed the weak adhesive forces between adhering
microorganisms and the coating. This coating had good hydrolytic stability and did not
degrade upon incubation in saliva for one month. (63)
Figure 9. Structure of Polyacrylamide
33
The influence of PAAm polymers at inhibiting protein adsorption and cell adhesion was
also assessed. For this, Liu et al. (61), grafted PAAm on gold surfaces by surface-
initiated ATRP and tested protein adsorption, cell adhesion and bacterial attachment.
The water contact angle of PAAm grafted surfaces was 14.8º. The protein adsorption
resistance was assessed by SPR, with three model proteins, Fg, BSA and Lyz. The
results showed that PAAm-grafted surfaces resisted to protein adsorption, compared to
uncoated gold surfaces. The results also demonstrated that PAAm surfaces suppressed
protein adsorption in a similar way to PEG or pSBMA. The PAAm surface also resisted
to protein adsorption from serum and plasma and cell attachment. The bacterial
attachment was assessed with two bacterial strains, P. aeruginosa and S. epidermidis.
Results demonstrated that PAAm-grafted surfaces resisted to the initial bacterial
attachment, leading to more than 97% reduction for both strains compared to uncoated
gold surfaces. (61)
PAAm brushes have good hydrolytic stability, can resist to protein adsorption, cell
adhesion, bacterial attachment and are stable on physiological fluids. (61,63,64)
Some authors have reported studies of PAAm grafted to silicone rubber surface with
promising results for in vivo applications. Silicone rubber is widely used as a biomedical
polymer for intravenous and urinary catheters, voice prostheses, oxygenators and
contact lenses. (60,62,63)
34
4.1.7 Heparin
Heparin (HP) (Figure 10) is a naturally occurring polysaccharide and consists of two
sulfated sugar monomers, which occur as repeating disaccharides units. (65)
Heparin has been used as an antiadhesive coating, since it can increase hydrophilicity
by formation of a hydrated layer between the surface and the bacteria. It also used to
increase the blood compatibility of medical devices due it’s anticoagulants properties.
The exposure of biomaterials to physiologic fluids cause the activation of blood defence
mechanisms leading to activation of the coagulation cascade, cellular inflammatory
mechanisms, complement system and platelets. These activations may compromise the
performance of medical devices. (65,66)
The influence of HP on the biding of S. epidermidis to fibronectin and consequently
their influence in biofilm formation was investigated by Arciola et al. (66) by dynamic
force spectroscopy (DFS). They investigated the attachment of S. epidermidis to a
fibronectin-functionalised gold-plated disk, incubated with HP solution or with
antihuman fibronectin monoclonal antibody solution. Results showed a decrease of
bacteria adhered to disk surfaces, in both cases. So, HP specifically inhibiting the
interaction of S. epidermidis adhesins and fibronectin. (66)
This influence was also tested by Lundberg and co-authors in IOLs under in vitro flow
conditions. The IOLs, fabricated with PMMA, with or without HP coating, were
incubated with human cerebrospinal fluid (CSF) for 1h or with CSF plus 0.50% plasma
for 12h. In this study two S. epidermidis strains were used, and bacterial adhesion was
assessed by bioluminescence. The results demonstrated that bacterial adhesion was
decreased on IOLs coated with HP compared to uncoated IOLs, with both strains, after
incubation with CSF plus 0.50% plasma for 12h. So, results suggested that IOLs with
HP coating were less predisposed to bacterial adhesion than uncoated IOLs. (67)
García-Sáenz et al. (68), also studied the in vitro adhesion of S. epidermidis to IOLs.
Figure 10. Structure of Heparin
35
They used two strain of S. epidermidis, slime-positive and slime-negative and compared
its adhesion to IOLs of PMMA, silicone, acrylic and HP-modified PMMA. The results
showed that slime-positive strains adhered at a higher level than slime-negative strains.
The results also demonstrated that bacterial adhesion to HP coated PMMA was lower
than adhesion to uncoated PMMA, because this coating formed a hydrated surface that
reduced bacterial adhesion. (68)
Heparin coating can be applied on diverse surfaces, used in medical devices, such
polyvinylchloride, polymethylmethacrylate, polyurethane, polyethylene and silicone.
This coating has been studied to reduce bacterial adhesion to catheters, artificial lenses,
and vascular grafts of PU. Heparin can also be used in orthopaedic devices, such
fracture fixation, knee replacements, tendon and ligament reconstruction and others.
(19,33,69)
Heparin-coated blood contacting medical devices are widely used in a clinical
application and can be divided into eluting and non-eluting technologies. These
technologies can be applied to medical devices such as cardiopulmonary bypass,
catheters, haemodialysis catheters, vascular grafts, extracorporeal circulation devices,
coronary stents, stents-grafts and others. (65)
36
4.1.8 Phosphorylcholine
Phospholipid polymers with phosphorylcholine (PC) (Figure 11) head groups are
appropriate for biomedical use, once they are biocompatible and can inhibit bacteria
and proteins adhesion.Figure 11 Phospholipids coating can also reduce platelets
adherence once theirs structure mimics biological cell membranes. These polymers
have antifouling properties and decrease thrombus formation by the formation of a
hydration layer. (70)
Polymers with phosphorylcholine have demonstrated to increase the hydrophilicity of
silicone elastomers (SE), (e.g. the water contact angle of unmodified SE was 100.8º
whereas, water contact angle of 2-methacryloyloxyethyl phosphorylcholine (MPC) -
modified SE was 20º). Additionally, this coating also showed to increase the
hemocompatibility of SE once it could reduce protein adsorption and platelets
adherence. (70)
The antifouling properties of phospholipids polymers surface was studied by Patel et
al. (71), under dynamic flow conditions, in PBS and 20% human serum. Thus,
polyethylene terephthalate (PET) surfaces were coated with Poly [(ω-
methacryloyloxyalkyl phosphorylcholine-co-n-butyl methacrylate)] (MAPC-co-
BMA), with two different methylene chain length. The results showed that
phospholipid polymer surfaces in PBS significantly reduced bacterial adhesion of S.
epidermidis when compared to uncoated PET surfaces. They also demonstrated that the
length of the methylene chain on MAPC unit also influenced bacterial adhesion since a
shorter chain length was more effective at reducing bacterial adhesion. Bacterial
adhesion in the presence of 20% human serum was markedly reduced to <10% when
compared to PBS. (71)
Other studies have been realized to assess the antifouling properties of phospholipids
polymers. One example was the work performed by Fujii et al. (72), that have
investigated the effect of poly (MPC-co-n-butyl methacrylate) (PMB) coating grafted
Figure 11. Structure of Phosphorylcholine
37
on stainless steel plates by radical polymerization, on bacterial adhesion. The formation
of a biofilm of S. aureus, S. epidermidis and P. aeruginosa was assessed by the
observation with a culture experiments for confocal laser scanning microscope (CLSM)
and the number of bacteria was measured by ATP (Adenosine 5’- triphosphate) assay.
The results demonstrated that PMB-coated surfaces reduced bacterial adhesion of the
three strains under study when compared to uncoated surfaces. Additionally, to the
resistance of the bacteria enveloped in the biofilm antibiotics were added. The
application of antibiotic decreased the number of S. epidermidis and S. aureus on both
uncoated and coated surfaces. In the case of P. aeruginosa, the addition of gentamicin
decreased the number of adhered bacteria whereas, the addition of cefazolin increased
that number on both coated and uncoated surfaces. This variance is explained by the
fact that P. aeruginosa was not susceptible to cefazoline. Thus, they conclude that PMB
coating prevent biofilm formation in in vitro conditions. (72)
The influence on biofilm formation was also tested by Hirota and co-authors (73), that
have studied the influence of MPC polymers on adherence of 4 microbial strains, S.
aureus, S. mutans, P. aeruginosa and C. albicans, to PET surfaces. The results
demonstrated that bacterial adhesion of all four strains, after 1h of incubation, were
significantly lower in MPC-coated surfaces (water contact angle = 10º) compared to
uncoated surfaces (water contact angle = 72º). Thus, application of an MPC coating on
medical devices surfaces could prevent microbial retention and biofilm formation. (73)
The PC coating can be used for medical devices such bone fixation devices, implantable
artificial hearts, artificial blood vessels, coronary stents, vascular grafts of PU, artificial
lungs, intravascular stents, intravascular guide wires, oxygenators and soft contact
lenses. (33,72–74) These coating can also be used in urological devices, such urinary
catheters and ureteral stents. (48,75)
38
4.2 Antifouling and antimicrobial surfaces
In some situations, the conjugation of antimicrobial and antifouling properties may be
advantageous. Since an antifouling coating cannot completely prevent initial
microorganism attachment, a few adhering microorganisms cells can ultimately form a
mature biofilm. (12) An ideal antibacterial surface should be capable of preventing
initial attachment, kill all microorganism that can attach and lastly remove dead
microorganisms. (38)
4.2.1 PEG-Antibiotics
Hydrophilic polymers like PEG are commonly used as spacers for the immobilization
of bioactive molecules such antibiotics to create surfaces with both antifouling and
antimicrobial properties. PEG is one of the most selected antifouling polymers and is
also used to immobilize several antibiotics, such penicillin (PEN), gentamicin (GEN)
and ampicillin (AM), on polymeric substrates. (38)
One example is the modification of ePTFE (expanded polytetrafluoroethylene) surfaces
with penicillin using PEG as a spacer. To evaluated antibacterial properties of the
coating, ePTFE surfaces were immersed into bacterial cultures of P. aeruginosa and S.
aureus, for 3-4 h at 37 ºC. The antibacterial activity was evaluated by measuring the
absorbance by ultraviolet-visible (UV-VIS) spectrometer. The results showed that %
absorbance for S. aureus was lower in the presence of the PEN-PEG-MA-ePTFE
compared to uncoated ePTFE, PEG-MA-ePTFE or PEN-ePTFE. Nevertheless, for P.
aeruginosa % absorbance in presence of PEN-PEG-MA-ePTFE was higher than in
presence of uncoated ePTFE. They concluded that PEN-PEG coating is highly effective
against S. aureus. (76)
In another study, an antibacterial effectiveness of this coating was determined. The
same authors incubated ePTFE, maleic anhydride (MA)-ePTFE, PEG-MA-ePTFE,
PEN-PEG-MA-ePTFE, and PEN-ePTFE surfaces into S. aureus culture, at 37 ºC, for
3, 6, 9, 12 and 24 h. At those times, an aliquot was collect and spread on agar plates to
grow the colonies. After incubation CFUs were determined. This study demonstrated
that antimicrobial efficiency of PEN-PEG-MA-ePTFE surface remained substantial
after 24 h of exposure to S. aureus. They also monitored the ester linkages between
PEG and PEN by ATR-FTIR spectrometry to assessed the hydrolytic stability of the
39
coating. The results revealed that this coating was still effective even with a 32% loss
of PEN. (77)
The attachment of penicillin to ePTFE is only effective against gram-positive bacteria.
To create an antibacterial coating that inhibits proliferation of an extensive range of
bacteria, Aumsuwan et al., developed a PEG-ampicillin coating. Ampicillin is a broad-
spectrum antibiotic with activity against gram-positive and gram-negative bacteria. To
determine an antibacterial effectiveness of this coating, AM-PEG-MA-ePTFE surfaces
are immersed into bacterial cultures of gram-positive bacteria (i.e. S. aureus, E. faecalis,
B. thuringiensis,) and gram-negative bacteria (i.e. P. putida, S. enterica and E. coli) at
37 ºC for 5 h. Antibacterial properties were assessed by measuring the optical density
(OD). The results demonstrates that AM-PEG-MA-ePTFE surface reduces bacterial
grown when compared to ePTFE, MA-ePTFE and PEG-MA- ePTFE in all bacterial
strains. Thus, attachment of ampicillin prevented the development of microbial films.
(78)
Another different approach can be the attachment of two different antibiotics to surfaces
(e.g. polypropylene (PP)). Thus, is possible eliminate the growth of gram-positive and
gram-negative bacteria, simultaneously. PEN-PEG-MA-PP and GEN-PEG-MA-PP are
effective against S. aureus and S. putida, respectively. An GEN/PEN-PEG-MA-PP
surface was developed and was possible to create an antibacterial PP surfaces with
different antimicrobial strengths, by altering the antibiotics concentration. (79)
Expanded polytetrafluoroethylene is a synthetic polymer widely used in medical
devices such as mitral valve tendon replacements, vascular grafts and soft tissue in
plastic and reconstructive surgeries. (76)
40
4.2.2 Heparin/Chitosan multilayer film
Chitosan (CHT) is a natural biocompatible cationic antimicrobial molecule. The
mechanism of antibacterial action of CHT is not completely understood. However, it is
probably related to the interactions between negatively charged microbial cells
membranes and positively charged CHT. These interactions result in alterations of the
cell surface, causing leak of intracellular components of microbial cells that leads to
cell death. (38,80–82) As describe above, HP is an anionic antifouling molecule with
antiadhesive and anticoagulants properties. The construction of multilayer film with
these two polysaccharides can form a film that combine antifouling and antimicrobial
properties. (82)
A heparin/chitosan multilayer can be prepared via layer-by-layer (LbL) technique,
resulting in a coating with antifouling and antimicrobial properties. (38,82,83)
Several studies about the antimicrobial properties of this multilayer films have been
reported. Fu et al. (82), developed a multilayer HP/CHT grafter into PET surfaces. The
initial adhesion of E. coli on PET substrates was determinate by SEM. PET substrates
were immersed in an E. coli suspension for 4 h at 37 ºC. The results showed that the
number of bacteria adhered onto HP/CHT multilayer films was lower than onto the
uncoated PET. The in vitro antibacterial activity of the multilayer film was assessed
and the results revealed that the coating could kill the bacteria successfully. The authors
observed that 46%-68% of the microbial cells were no longer viable after 7 h of
exposure to the multilayer-modified PET films and that this reduction was enhanced
with time. On the other hand, the number of viable cells only decreased 7% after 7 h
of contact with PET films. The surface properties of this multilayer film was pH
dependent, once pH influences the amount of chitosan on the surface. The number of
viable bacteria at pH 3.8, 2.9 and 6.0 was 68%, 58% and 46%, respectively. (82)
Another studied was performed by Follmann and co-workers (84), that developed a
multilayer film by LbL technique of N-Trimethyl chitosan (TMC)/HP grafting on
polystyrene (PS) chemically modified surfaces. TMC is a quaternary polycation
obtained from the N-methylation of CHT and have similar properties to the CHT but, a
higher antimicrobial effect. TMC/HP multilayer with different quaternization degree
(TMC20 and TMC80) and PS-modified surfaces (control) were incubated in a bacterial
41
suspension of E. coli for 4 h at 37 ºC. The initial adhesion was tested and the results
showed that bacterial cells was only detected in the control, once this surface has low
hydrophilicity compared to TMC/HP multilayer surfaces. The in vitro antibacterial
activity of the both TMC/HP multilayer prepared in this study was compared.
TMC80/HP assembled at pH 3.0 exhibited no significant influence on the viability of
bacterial cells as compares to control. However, TMC20/HP assembled at pH 3.0 and
TMC80/HP prepared at pH 7.4 it reduced the number of viable-cells in 32.7% and
64.6%, respectively. TMC80/HP assembled at pH 7.4 showed good antibacterial
properties, so it is possible to use it as an antimicrobial coating. (84)
Wang et al., development a (heparin/chitosan)10− (poly(vinylpyrrolidone)/poly (acrylic
acid))10 [(HP/CHT)10−(PVP/PAA)10] multilayer film by LbL self-assembly. The
multilayer film was formed by a top-down degradable PVP/PAA film that can resist
bacterial attachment for 24 h. The multilayer film is also constituted by a subjacent
HP/CHT film, with antibacterial properties. The antimicrobial properties of the
multilayer film were assessed by water-borne assay with S. aureus. Some samples were
chosen for the assay: silicon wafer, (HP/CHT)10 multilayer film, (HP/CHT)10-
(PVP/PAA)10 multilayer film that thermal cross-linked at 170 °C for 4 h and
(HP/CHT)10-(PVP/PAA)10 multilayer film that thermal cross-linked at 110 °C for 16 h.
These substrates were exposed to an S. aureus suspension. In the case of the
(HP/CHT)10 multilayer film, a few bacteria were observed on the surface after 24 h.
The (HP/CHT)10−(PVP/PAA)10 multilayer film with thermal cross-linked at 110 °C for
16 h, exhibited few bacteria adhered at 24 h. On the other hand,
(HP/CHT)10−(PVP/PAA)10 multilayer films that were cross-linked at 170 °C for 4 h,
exhibited several bacteria adhered. After 24 h, the top-down was completely degraded
and the (HP/CHT)10 was exposed, providing a contact-killing surface. To assess it, the
(HP/CHT)10−(PVP/PAA)10 multilayer film cross-linked at 110 °C for 16 h and then
degraded at 37 °C. After 24 h, the multilayer film and bare glass (control) was exposed
to S. aureus (previously stained). Fluorescent microscopy images demonstrate that the
bacteria on uncoated glass were alive. However, almost all of the adhered bacteria on
the multilayer film had been killed (85)
The HP/CHT multilayer film could be potentially applied to medical devices such as
cardiovascular implants. (82) On the other hand, [(HP/CHT)10−(PVP/PAA)10]
multilayer film has potential to be applied in implants. However, this films can only be
42
applied into medical devices that can be thermally treated at a high temperature such as
bone nails and steel stents. (85)
4.2.3 MEO2MA based copolymer
Copolymer brushes based on 2-(2-methoxyethoxy)ethyl methacrylate (MEO2MA) have
also antifouling properties. Furthermore, these copolymers have reactive hydroxyl
groups which allowed controlled loading of antimicrobial peptides. Glinel et al.,
prepared copolymer brushes based on MEO2MA and oligo(ethylene glycol)
methacrylate (OEGMA). This copolymer was prepared by surface-initiated ATRP, on
silicon surfaces, and was subsequently functionalized by magainin I, a natural
antibacterial peptide. The antibacterial activity of this coating was tested against two
bacterial strains (i.e. L. ivanovii and B. cereus). Coated silicon surfaces were incubated
in bacterial suspension for 3 h and observed by optical microscopy. The results revealed
the absence of bacteria for both strains on the nonfunctionalized copolymer brushes.
However, a few bacteria were detected on functionalized brushes (coverage did not
exceed 1% of the surface). These results demonstrated that bacterial cells interacted
with magainin I peptide. CLSM images showed that the adhered cells (previously
stained) corresponding to dead bacteria. Thus, the antifouling properties of the brushes
combined with the antimicrobial peptide lead to a delay of bacterial biofilm formation.
(20)
The antibacterial effect of these coatings was also assessed by Blin et al., that developed
hydrophilic copolymer brush (poly-(MEO2MA-co-HOEGMA)) grated to silica-based
microparticles by surface-initiated ARTP, subsequently functionalized with magainin
I. The antibacterial effect was assessed against L. ivanovii by incubating the modified
surface in a bacterial suspension for 3 h at 25 ºC. The viability of the cells, previously
stained, was assessed by CLSM and fluorescence microscopy. The observations reveal
that most of the cells adhered to coated silica were dead. (86)
These copolymers can be applied to the biomedical field, in particular to implants. (20)
43
5. Discussion
Medical device-associated infections represent a significant risk for patient’s health.
Thus, different strategies have been used to prevent these infections. One of these
strategies is the surface modification of biomaterials with antifouling polymers.
PEG is the most studied and used antifouling polymer, due to its noncytotoxic,
nonimmunogenic and biocompatibility properties. One advantage of PEG is that this
polymer can be applied to varied biomaterials and can be used as antifouling polymer
in orthopedic and dental implants, catheters of PU, among others. However, has been
reported that PEG has reduced stability due to its oxidative degradation in vivo. Despite
PEG representing the gold standard in this respect, it not appropriate for a long-term
application. Thus, it is necessary investigate new polymers and materials with
antifouling properties, biocompatible and stables at physiological conditions. Examples
of other polymers, that have been studied for their antifouling applications, can be PVP,
POXs and PAAm. (22,50,52,87)
Medical devices with heparin coatings are already used, namely in blood-contacting
devices. POXs polymers have demonstrated antifouling properties similar to PEG but,
a higher stability and the possibility of keeping their antiadhesive properties even in
physiological environments. This polymer can be used in some medical devices such
catheters, prosthesis and others. PC coating it widely used in medical devices, for
example was used for developed glaucoma drainage device, intraocular lenses, contact
lenses (omafilcon A) based on the copolymers of PHEMA and MCP. PC polymers can
also be applied in coronary stents, ventricular assist device and vascular grafts. This
polymer also has been studied for orthopedic devices usage. PSBMA polymers have
been applied to endotracheal tubes, orthopedic devices and peripherally inserted central
catheters (FDA approval). (43)
Theses polymers have demonstrated to be able to reduce biofilm formation of some of
the most common microorganisms associated with HCAIs, such E. coli, S. aureus,
among others. Additionally, polymers like PEG and PAAm are also able to reduce
biofilm formation of yeast. (30,51,63,72)
There are limitations to compared these different strategies and assessed the best one
since the tests are performed under different conditions. Furthermore, most of the
44
studies only focus the initial stage of the biofilm formation. Thus, a standardized test is
essential, in order to directly compare the studies. (18,19)
Despite the good results obtained under in vitro conditions, there are difficulties in
demonstrating the effectiveness of these surfaces under real clinical conditions. The in
vitro test does not reflect the environment that medical device will be exposed.
Furthermore, there are not enough studies for testing the stability, toxicity, and safety
of these polymers under in vivo conditions. Thus, it necessary an in vivo test for an
appropriate study of the medical device. These difficulties create limitations in
correlating the in vitro, in vivo preclinical and clinical data. (12,18,19)
A conjugation of antifouling and antimicrobial surfaces can be a promising approach,
once an antifouling coating cannot completely prevent biofilm formation. Ideally, these
surfaces are able to prevent initial microorganism attachment, kill microorganism that
might attach and remove the dead microorganisms. (12)
45
6. Conclusion
HAIs are a major clinical problem with significant impact on mortality and morbidity.
The best approach against such infection is the prevention of biofilm formation.
Consequently, several surface modification techniques have been developed. Polymers
such PEG, PVP, polybetaines and others have demonstrated good anti-adhesive
properties under in vitro condition. However, in some cases it is necessary studies to
assess their efficacy under physiological conditions. These antifouling coatings are
different from each other and have advantages and disadvantages. Thus, the type of
coating should be adapted for a specific application. Some of these polymers are already
applied in medical devices such catheters, orthopaedic implants, dental implants,
vascular grafts, contact lenses and other.
Despite these polymers being able to reduce microorganism attachment on medical
device surfaces, a few adhered bacteria can eventually form a biofilm. Therefore, the
association of theses coatings with antimicrobial agents is an interesting approach.
46
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