The interaction of blood proteins and platelets on surface

The interaction of blood proteins and platelets on

surface-attached poly (alkylacrylamide) networks

Dissertation zur Erlangung des Dokotorgrades

der Technischen Fakultät

der Albert-Ludwigs-Universität Freiburg im Breisgau

vorgelegt von

Chinnayan Kannan Pandiyarajan (M.Sc Chemistry)

Freiburg im Breisgau, Februar 2013

II

This research was carried out between January 2010 and January 2013 under the

extreme supervison of Prof. Dr. Jürgen Rühe in the laboratory of Chemistry and

Physics of Interfaces (CPI), Department of Microsystems and Engineerings (IMTEK),

University of Freiburg, Freiburg-79110, Germany.

Date of Submission: 12-03-2013

Date of Disputation: 11-04-2013

First referee: Prof. Dr. Jürgen Rühe

Second referee: Prof. Dr. Gerald Urban

III

Man needs his difficulties because they are necessary to enjoy success

- Dr. A.P.J. Abdul Kalam

IV

To my brother & mother

List of Charcaters

V

List of Characters

Α Linear swelling ratio

αsa Swelling of surface-attached networks

αuc Swelling of unconstrained networks (free gels)

Ε Permittivity constant

Θ Angle of incidence

Θc Critical angle of incidence

Φ Flory-Huggins interaction parameter

Ξ Linear distance between two crosslinks

Υ Number of crosslinks

φe Equilibrium volume fraction

φo Volume fraction of water

Τ Wall shear stress

µ Fluid viscosity

ΔEela Free elastic energy

ΔGads Free energy of adsorption

ΔHads Enthalpy of adsorption

ΔSads Entropy of adsorption

ΔSconf Entropy of configuration

ΔSmix Entropy of mixing

φw Volume fraction of water

φp Volume fraction of polymer

B chamber width

D thickness

H chamber height

K Boltzmann constant

List of Charcaters

VI

Mc Number average molecular mass between two crosslinks

Mn Number average molecular mass

n Refractive index

n1 Ratio of solvent molecule

n2 Ratio of polymer repeat unit

Nc Number of segments

Q Fluid flow rate

R Gas constant

Ra Roughness

Rg Radius of gyration

S Volumetric swelling

T Temperature

t Time

Z Number of carbon atom at N-substitution

List of Abbreviations

VII

List of abbreviations

AFM Atomic force microscopy

AIBN 4,4‟-Azobis(isobutyronitrile)

atm Atmosphere

ATR Attenuated total reflection

Au Gold

BAAm Butylacrylamide

BSA Bovine adult serum

CA Contact angle

Coll Collagen

Cr Chromium

DEAAm Diethyl acrylamide

DMAAm Dimethyl acrylamide

DOP Di-octylphthalate

EAAm Ethylacrylamide

EDTA Ethylenediaminetetraacetic acid

Fg Fibrinogen

Fn Fibronectin

FT-IR Fourier transform infrared (spectroscopy)

GPC Gel permeation chromatography

HUVEC Human umbilical vein endothelial cells

L929 Mouse fibroblasts

MAAm Methylacrylamide

MABP Methacryloyl-4-oxy-benzophenone

MB Microbubbles

NHDF Normal human dermal fibroblasts

NMR Nuclear Magnetic resonance (spectroscopy)

List of Abbreviations

VIII

OWS Optical waveguide spectroscopy

PAAm Propylacrylamide

PBS Phosphate buffered saline

PS Poly styrene

PU Polyurethane

PVC Polyvinylchloride

RPM Rotation per minutes

SEM Scanning electron spectroscopy

SPR Surface plasmon resonance spectroscopy

SSAz Styrenesufonylazide

Table of contents

IX

Table of Contents

1. Introduction ................................................................................................................. 1

1.1. Blood-contacting devices - Materials consideration ....................................................................... 1

1.2. Biomaterials and their limitations .................................................................................................... 6

1.3. Blood compatibility / Hemocompatibility ....................................................................................... 8

1.4. Thrombus formation ......................................................................................................................... 8

1.5. Coating methods for blood contacting devices ............................................................................... 9

1.5.1. Inorganic coatings .................................................................................................................... 11

1.5.2. Organic Coatings ...................................................................................................................... 11

1.6. Hydrogels ........................................................................................................................................ 14

1.6.1. Hydrogels from synthetic polymers ........................................................................................ 16

1.6.2. Responsive hydrogels ............................................................................................................... 17

1.7. Hydrogels in medical applications ................................................................................................. 21

1.8. Challenges in Biomedical applications .......................................................................................... 26

2. Goal and Strategy of the work ................................................................................ 27

2.1 Goal ................................................................................................................................................... 27

2.2. Strategy of the work ....................................................................................................................... 27

3. Synthesis and characterizations ............................................................................. 30

3.1. Synthesis of crosslinker (MABP) and anchor (Bp-Si) ................................................................... 30

3.2. Synthesis of alkylacrylamide copolymers ...................................................................................... 31

3.3. Surface-attached networks ............................................................................................................. 36

3.4. Contact angle and Topography ...................................................................................................... 41

3.5. Conclusion ....................................................................................................................................... 44

4. Swelling of the surface-attached poly (alkylacrylamide) networks ................ 45

4.1. Theory of swelling .......................................................................................................................... 45

4.1.1. Isotropic swelling ..................................................................................................................... 45

4.1.2. Anisotropic swelling ................................................................................................................ 47

4.2. Surface Plasmon resonance spectroscopy (SPR) ........................................................................... 48

Table of contents

X

4.3. Swelling in water ............................................................................................................................ 51

4.4. Swelling in humid air ..................................................................................................................... 60

4.5. Heat of hydration ............................................................................................................................ 82

4.6. Conclusion ....................................................................................................................................... 84

5. Adsorption of proteins on poly (alkylacrylamide) surfaces .............................. 87

5.1. Theory and mechanism of protein adsorption .............................................................................. 87

5.2. Competitive adsorption of proteins (Vroman effect) .................................................................... 89

5.3. Protein resistant surfaces ................................................................................................................ 90

5.4. Protein adsorption on surface-attached poly (alkylacrylamides) networks ................................ 93

5.4.1. Determination of mesh size and the size exclusion effect ................................................... 101

5.4.2. Entropic shielding .................................................................................................................. 104

5.5. Protein adsorption on terpolymer layers ..................................................................................... 106

5.6. Protein repellency of PDMAAm gels .......................................................................................... 108

5.7. Adsorption of lipids ...................................................................................................................... 110

5.8. Conclusion ..................................................................................................................................... 114

6. Platelet adhesion .................................................................................................... 116

6.1. Molecular mechanism of platelet adhesion ................................................................................. 116

6.2. Static platelet adhesion on surface-attached poly (alkylacrlamide) networks .......................... 118

6.3 Platelets adhesion on 2D structured surfaces ............................................................................... 122

6.4. Ultrasound contrast agent to detect platelets at shear flow condition ....................................... 124

6.5. Platelet adhesion under arterial flow condition .......................................................................... 131

6.6. Conclusion ..................................................................................................................................... 131

7. Cell adhesion on surface-attached poly (alkylacrylamide) networks .......... 133

7.1. Mechanism of cell adhesion ......................................................................................................... 133

7.2. Cell adhesion on surface-attached poly (alkylacrylamide) networks ........................................ 137

7.3. Cell adhesion of patterned substrates .......................................................................................... 142

7.4. Cell adhesion of copolymers ........................................................................................................ 145

7.5. Conclusion ..................................................................................................................................... 147

Table of contents

XI

8. Blood tube modification ....................................................................................... 149

8.1. Attenated Total Reflection spectroscopy (ATR) ......................................................................... 149

8.2. Modification process ..................................................................................................................... 150

8.3. Characterization of PDMAAm-SSAz coatings ............................................................................ 152

8.4. Conclusion ..................................................................................................................................... 155

9. Summary .................................................................................................................. 157

10. Zusammenfassung................................................................................................ 164

11. Experimental section ........................................................................................... 168

11.1. Materials ...................................................................................................................................... 168

11.2. Instrumentations ......................................................................................................................... 172

11.3. Synthesis and characterization of monomers and polymers .................................................... 176

11.4. Deposition of surface-attached networks .................................................................................. 182

11.5. Preparation of Protein solution for SPR .................................................................................... 184

11.6. Cell culture .................................................................................................................................. 184

11.7. Platelet extraction ....................................................................................................................... 185

12. Appendix ............................................................................................................... 187

Journal Publications ................................................................................................... 196

Referred conference proceedings ............................................................................ 196

References .................................................................................................................... 197

Acknowledgements.................................................................................................... 210

Abstract

IX

This thesis investigated the adsorption of blood proteins and platelets onto the

surface-attached poly (alkylacrylamide) networks that exhibit small, systematic variations of

the chemical composition. The polymer coatings were generated by depositing a thin layer of

benzophenone group containing copolymer onto a solid substrate, followed by photo

crosslinking and simultaneous surface-attachment. These surface-attached networks showed

anisotropic swelling that strongly depended on length of the alkyl substituent. A strong

correlation between swelling and protein adsorption was observed. The swollen surface-

attached layers were found to repel proteins both through entropic shielding or size

exclusion and as a consequence repel blood platelets and cells. Our results suggest that the

protein repellent coated materials were promising candidates for the generation of

hemocompatible surfaces.

In der vorliegenden Arbeit wurde der Einfluss der chemischen Zusammensetzung auf

die Adsorption von Blutproteinen und die Adhesion von Thrombozyten auf

oberflächengebundenen Poly(alkylacrylamid)-Netzwerken untersucht. Die

Polymerbeschichtungen wurden durch Abscheidung dünner, benzophenonhaltiger

Copolymerfilme auf festen Substraten erzeugt. Durch anschließende Belichtung mit UV-

Strahlung wurden die Polymerschichten vernetzt und an die Substratoberfläche gebunden.

Diese so erzeugten, oberflächen-gebundenen Polymernetzwerke zeigen ein anisotropies

Quellverhalten, welches sehr stark mit der Länge der Alkylsubstituenten korreliert. Die

Quellung der Netzwerke beinflusst dabei die Proteinadhesion an der Oberfläche der

Polymerschichten. Durch entropische Abschirmung und Größenausschluss wirken die

gequollenen, oberflächen-gebundenen Polymernetzwerke protein- und zellabweisend. Die

Ergebnisse zeigen, dass sich diese proteinabweisenden Polymerschichten zur Erzeugung

hämokompatibler Oberflächen eignen.

1.1.Blood-contacting devices

1

1. Introduction

1.1. Blood-contacting devices - Materials consideration

In the medical device industry, the choice of biomaterials employed in most devices

that come into direct contact with blood flow are based on stability, permeability, cost, ease

of sterilization, non-toxicity, and with a acceptable hemocompatibility.1-5 Table 1 lists some

of the commercially available blood-contacting devices that have been in frequent use.

Figures 1.1- 1.4 depicts the medical devices are listed in Table 1. Important examples for such

devices are ventricular assist devices (VADs), artificial blood vessels, vascular stents, and

artificial heart valves.6,7

A VAD is a mechanical pump that aids the function of damaged heart ventricles and

restores normal blood flow. Used as a bridge for transplantation, they also have applications

in treating patients with terminal heart failure.8 VADs are categorized by their mechanical

function as pulsatile, or continuous flow.9 Pulsatile assist devices are often reffered as first

generation pumps that are used clinically. They have large and multiple moving parts, which

help to pump the blood in a pulsatile fashion. The second-generation pumps are small and

posses a single moving part with a continuous flow. Third generation pumps are designed for

long-term applications, which uses the magnectic levitation (e.g. magnetically suspended

axial rotor or centrifugal rotor) for the non-pulsatile flow.8

A second way to catagerize VADs is based on their purpose, Left ventricular assist

devices (LVAD) are used when the left ventricles are damaged, and it helps left ventricles to

pump blood to aorta, which carries the oxygen rich blood from heart to the body. Right

ventricular assist devices (RVAD) are used when the right ventricle are not functioning.

RVAD‟s used to pump blood to the pulmonary artery from which the blood carries to lungs

for oxygen. Bi-ventricular devices (BVAD) are used when both ventricles are not

functioning.9-10


2

Figure 1.1: Heart Mate II left ventricular assists device that is currently approved by FDA.

Reproduced from Thoratec Corporation with permission © 2012.

Artificial blood vessels are intricate networks of hollow tubes that transport blood

throughout the entire body. For example, in heart bypass operation, surgeons remove veins

from patient‟s leg and replace the damaged arteries in the muscle of the heart.3 The use of

artificial blood vessels avoids such treatment and makes it simple for doctors and patients to

meet the criteria by replacing the natural blood vessel (Figure 1.2). Usually, these artificial

blood vessels are made from biodegradable materials, plastic mesh tube with human or

animal muscle cells.3

Figure 1.2: (A) Diagram showing catheters, which are tubes that can be inserted into the

body‟s duct or vessel depend on application and type of catheters. Figure adopted from

www.interfaceusa.com with permission. (B-D) An artificial blood vessel covered with

endothelial cells.3 (Figure adopted from www.materials.jhu.edu with permission).

(B)

(D)

(C)

(A)

http://www.interfaceusa.com/


3

Stents are prosthetic devices which are implanted in the lumen in order to provide

support and assure of patency of the lumen (Figure 1.3).4 Stents are implanted within the

vascular system to reinforce collapsing, particularely occuleded, weakened, or abnormally

dilated sections of blood vessels.4 Currently, stents are the most frequently used devices in

cardiac surgery. There are different types of medical stents that are available based on their

purpose such as angioplasty stents (for blood clots in the coronary vasculature), biliary stents

(use for problems associated with pancrease), carebral stents (for cerebral vasculature),

colonic stents (used when colon obstructed by tumor) and duodental stents (used for small

intestine that are obstructed by tumor). However, restenosis rate of 20-30% remains a major

chellage for all these medical stents. The biology behind the restenosis includes the plaque

redistribution, thrombosis and neointimal hyperplasia. Currently, there are four types of

stents such as bare metallic stents (BMS), coated metallic stents, biodegradable stents and

drug eluting stents (DES) are in usage.

Figure 1.3: (A) Human Coronary stent, and (B) Abdominal Aortic Aneuysms (AAA) stent

(Figures adoped from www.heartlinkplano.com with permission).

Artificial heart valves consist of an orifice, through which the blood flows, and a

mechanism that closes and opens the orifice.5 There are two types of artificial heart valves:

one is mechanical valves made from synthetic materials and other is from biological or tissue

valves made from animal or human tissues. Further, mechanical valves classified into three

http://www.heartlinkplano.com/


4

types based on the opening and closing mechanism. These mechanisms are a reciprocating

ball, a tilting disk, or two semicircular-hinged leaflets. Similarly, tissue valves can be divided

into homografts and xenografts or heterografts. Homografts signify the transport from the

same species (from human to human or from animal to animal) and xenografts denote the

transport from one species to another (from pig to human).7

Figure 1.4: (Left) Picture of a mechanical heart valve made from diamond like carbon with

flaps that can open and close with the pumping of blood, and (Right) A collapsible valve, first

introduced in 2007. It can be stretched-down to the width of a pencil in order to thread from

a blood vessel in the leg to the heart‟s failing valve and replace it. (Figures adopted from

www.scientificamerican.com with permission).

http://www.scientificamerican.com/


5

Table 1.1: Comparison of some commercially available blood contacting devices.1-11

Device Structure/ Biomaterial Used Application

Ventricular assist

devices (Figure.1.1)

mechanical circulatory pump

made of plastic/ Ti

replacing partial/ complete

function of failing heart 6-7

Catheters

(Figure.1.2A)

mostly thin flexible tubes made of

polymers such as, silicone, rubber,

latex or thermoplastic elastomers

delivery/ removal of body

fluids

monitoring cardiovascular

funtions

opening blocked conduits –

e.g. arterial angioplasty 2

Artificial blood vessels

(Figure.1.2B-D)

plastic/ mesh tube made with

human/ animal muscle cells

heart bypass surgery3

Vascular stents

(Figure.1.3)

mostly small flexible tubes made

of plastic/ wire mesh or memory

metal alloys

supporting patency of lumen

reinforcing weak, occuleded

or abnormally dilated sections

of blood vessels 4

Artificial heart valves

(Figure 1.4.)

mechanical or biological valves

with an orifice made from

synthetic materials or

human/animal tissues5

transport of blood between

same/ different species 7

Hemodialysis artificial membrane (filter) made

from regenerated cellulose

removing waste products in

blood, e.g. creatinine, urea,

and free water during kidney

failure9

Oxygenator

Membranes

thin gas permeable, hollow fiber

membrane made of micro-porous

polymers11

oxygenation, i.e. addition of

O2 and removal of CO2 from

blood for artificial life

support10

http://en.wikipedia.org/wiki/Silicone_rubber

http://en.wikipedia.org/wiki/Latex

http://en.wikipedia.org/wiki/Thermoplastic_elastomer

http://en.wikipedia.org/wiki/Creatinine

http://en.wikipedia.org/wiki/Urea

1.2.Biomaterials and their limitations

6

1.2. Biomaterials and their limitations

In general, biomaterials are materials that are used in medical devices or implants.

These can be polymers, metals, ceramics or composites of these materials. Table 1.2 lists a

few such examples.12-13 In 1982, the National Institute of Health (NIH) development on

clinical applications, first introduced biomaterials as, "any substance other than a drug or a

combination of substances, synthetic or natural in origin, which can be used for any period

of time wholly or as part of a system which treats, augments, or replaces any tissue, organ or

function of the body". Williams later modified this concept stating that, a biomaterial is a

synthetic or a modified natural material that interacts with parts of the body.14

Also in the early 1980‟s, biocompatibility, the most important among the

characteristics of biomaterials was equated to inertia15 although the idea was ambiguous since

there is no ideal inert material, which does not interact with body fluids or tissues.

Eventually by early 1990‟s Williams14-15 and Ratner 16 introduced definitions that were more

precise. According to Williams, biocompatibility is "the ability of the materials to perform

with an appropriate host response in a specific application”. Whereas according to Ratner,

biocompatibility is “the exploitation by materials of the proteins and cells of the body to

meet a specific performance goal”. In other words, the ability of an implant surface to: (a)

interact with liquids and cells of the biological system, and (b) cause exactly the same

reaction, which an analogue body tissue would bring about, i.e., the body‟s acceptance of

materials.16

Chemical and physical characteristics, such as, hydrophilicity, hydrophobicity, ionic

group, crystallinity, and surface topography have been used to describe a biocompatibility.17

However, these results remained unsatisfying and were often not understood. Hence, the

searches for physical properties that can be used to predict the biocompatibility of a material

are in the center of many research activities.

1.2.Biomaterials and their limitations

7

Table 1.2: Biomaterials and their applications.12

Classification Materials Applications

Metals,

Alloys

Steel

Titanium

Gold alloys

Silver

Fracture correction

Dental/ bone/ articular replacement

Pace makers

Antibacterial

Ceramics,

Glasses

Calcium phosphate

Bioactive glasses

Porcelain

Bone regeneration

Bone replacement

Dentures

Polymers Polyethelene

Polypropylene

Polytetrafluroethylene

Polyester

Polyurethane

Polyvinyl chloride

Polymethylmethacrylate

Polyacrylate

Silicon

Hydrogels

Articular replacement

Suture materials

Vascular grafts

Resorbable systems

Blood contacting devices

Tubes and bags

Intraocular lenses

Dental implants

Soft tissue replacement

Ophthalmology

1.3. Blood compatibility / Hemocompatibility

8

1.3. Blood compatibility / Hemocompatibility

A biomaterial is called as blood compatible or hemocompatible, provided its

interaction with blood does not instigate any damage to blood components, such as blood

cells and plasma proteins.18 Adsorption of proteins and other molecules initially occurs upon

exposure of the implant material to blood. Proteins typically adsorb onto the surface via non-

specific interactions.19 In general, hydrophobic surfaces will adsorb large amounts of protein

than hydrophilic surfaces. The abundant small plasma proteins in the blood are adsorbed fast

and sequentially displaced by larger proteins (more details provided in chapter 5.2) also

known as the Vroman-effect, governs the temporal pattern of protein adsorption.1,20 This can

affect the conformation of adsorbed proteins based on chemical and surface properties of the

biomaterials.

Additionally, the adsorbed protein layers will control the biological process such as,

cell adhesion, activation of enzyme cascades of coagulation and inflammation. These

inflammatory reactions end in an encapsulation of the biomaterial, i.e. the development of

scar tissue around the implants.21 This results in the formation of a thrombus, one of the

major complications in many medical devices, especially in the case of cardiovascular

application, where it affects the long-term stability of devices. Therefore, it is crucial to

understand the formation of thrombus to pave a new bioengineering path for

hemocompatible surfaces.

1.4. Thrombus formation

As mentioned earlier, the first event to occur after exposure of a foreign material to

mammalian blood stream is the quick adsorption of plasma proteins onto the material

surface, after which it triggers a cascade of complex reaction as shown in Figure 1.5.

1.4. Thrombus formation

9

The adsorbed proteins provide site for small disk-shaped cells called platelets, which

adhere to the interface. This leads to the release of a protein called thrombin, the key

enzyme for the formation of fibrin.22 Fibrin spontaneously assembles into fibrils, forming a

fibrin network/ matrix that stabilize the adsorbed molecules. Consequently, the adhered

proteins and cells form aggregates known as thrombi.23 The thrombus might obstruct the

blood flow at the point of formation or the blood current removes the thrombus from the

material‟s surface and the thrombus block an artery elsewhere. Hence, the main goal in the

design of blood compatible materials is to generate a surface that suppresses the non-specific

interaction with plasma proteins and blood.

Figure 1.5: Schematic diagram illustrating the formation of a thrombus (Figure adopted from

Tirrell et al.21). First, proteins from the blood plasma adsorb onto the material surface; then

the platelets specifically interact with proteins thereby adhering on top of them. The

adsorbed platelet releases fibrin, which stabilizes the adsorbed proteins and platelets leading

to the formation of thrombus.

1.5. Coating methods for blood contacting devices

Before approaching fabrication methods for blood compatible surfaces, it is necessary

to understand the physical and chemical surface properties (i.e. charges, topography, and


10

surface energy) influence hemocompatibility.24 Generally, hydrophobic surfaces adsorb more

plasma proteins than hydrophilic surfaces.25 However, this is not universal. The plasma

oxidation of hydrophobic surfaces, for example on polypropylene/polyurethane can cause an

increase in the wettability and proteins adsorption but reduces platelet adhesion.26 In

addition, negatively charged surfaces activate plasmatic coagulation,27 and positively charged

surfaces - enhance adhesion and activation of blood platelets.28-29

Textured surfaces show less thrombus formation. It has been hypothesized that the

cavities in such texturing entrap blood components to form a biological neointimal1 layer is

formed, which controls the thromboembolytic events.8,30 Excimer laser micromachining is

used to prepare a master negative mold of patterned cavities for the generation of a

polyurethane textured surface by a solvent casting. Figure 1.6 shows scanning electron

microscopic images of textured surfaces consisting regularly spaced and tapered with

polyurethane microfiberes of 25, 50 and 100 µm in length, and the spacing of approx. 100µm.

The textured surfaces showed a strong deposition of white cells than the non-textured

surfaces. However, one out of ten samples showed a thrombus formation and the origin of

this behavior is not well understood.30

Figure 1.6: Scanning electron micrographs of polyurethane textured surfaces with fiber

length of 25µm (left), 50µm (middle) and 100µm (right) were obtained using excimer laser

micromachining techniques.30 Figure adopted from Fujisawa et al.30

1neointimal – a new or thickened layer of arterial intima formed especially on a prosthesis or

in atherosclerosis by migration and proliferation of cells from the media (www.m-w.com).8,30


11

1.5.1. Inorganic coatings

Inorganic coatings especially metal nitrides, metal oxides and diamond like carbon

(DLC) are widely used in medical devices such as mechanical pumps and stents. In general,

these materials show excellent mechanical and chemical stability and comparatively high

inertness with body fluids and tissues. Titanium (Ti) and titanium alloy (TiN) are suitable

materials for bone implantations due to their excellent integration with bone, and for the last

two decades, it is one of the most successful coatings for heart valves and ventricular assist

devices.31 In addition, Ti oxides show improved hemocompatibility for many medical

devices.32 DLC also has an advantage of having high mechanical strength, low frictional

coefficient, chemical inertness, high thermal conductivity and excellent biocompatibility.

DLC is a meta-stable form of amorphous carbon that contains a combination of sp3 (diamond

like) and sp2 (graphite like) hybridization with some of the bonds terminated with hydrogen

(amorphous carbon and amorphous hydrogenated carbon).33 Cathodic arcs, pulsed laser,

direct ion beam can be used to deposit DLC film, as well as plasma enhanced chemical vapor

deposition and sputtering techniques are also being used. Among inorganic coatings, DLC

coated surface shows the least platelet adherence compared to metal oxides and nitrides.10

1.5.2. Organic Coatings

The use of organic compounds and several polymers as biomaterials has been

successful in medical devices for the last five decades. The advantage of organic coatings over

inorganic coatings is that they offer a wide range of chemical surface modification and an

abundance in the materials selection for desired applications.34 However, they are prone to

hydrolytic degradation in biological situations with time. The surface coatings include

Endothelial lining, Bioactive coatings, and Surface passivation. The following sections discuss

each of them in detail.


12

Endothelial lining:

This method mimicks the natural blood vessels of the human body with a layer lining

of endothelial cells (EC‟s) on the material surface of blood contacting device.35 The process

first includes depositing with biological molecules such as fibronectin, fibrinogen, peptides,

and growth factors at the devices. Then, endothelial progenitor is seeded and allowed to

differentiate into endothelial cells to form a monolayer of EC‟s (See Figure 1.7) However, the

application of EC coatings for blood devices is still in its infancy. Although less

thrombogenic, the generation of a uniform monolayer on the surface is difficult to achieve.36

Besides, EC‟s lining becomes less viable or loses its function of endothelial-dependent

relaxation and biological factor production.

Figure 1.7: Schematic drawing of a human blood vessel illustrating the Endothelial lining

which is a crucial characteristic of human blood vessel functions.Figure adopted from Tirrell

et al. 21

Bioactive coatings:

Heparin is one of the materials widely used in bioactive coating for blood contacting

devices. Its anticoagulant activity in plasma inactivates the formation of fibrin clots


13

inhibiting two principle procoagulant proteases, factor Xa and thrombin.[37 A heparin coating

is generated, first by preparing an intermediate layer of polyamine, and then the heparin

composed with antithrombin (which is a binding site for thrombin) is covalently attached to

the polyamine surface (which is already coated on the device surface, as shown in Figure

1.8.). Recent developments along these lines, employ polymeric coatings constituted of nitric

oxide (NO) chemistry and immobilized active heparin.38 A lipophilic N-diazeniumdiolate

(diazeniumdiolated dibutylhexanediamine (DBHD/ N2O2)), is doped into an underlying

polymeric layer of the coating and continuous release of NO was initiated upon exposure of

the layer to water. Consequently, the uptake of water molecules releases NO from the

polymeric surface. The top layer or the outer most layers also contain heparin attached with

the help of suitable spacers, as shown in Figure 1.8. This dual approach is very much similar

to that of natural EC lining of human blood vessels, where it meets the two principle

thrombotic mechanism of potent antiplatelet activity of NO and inhibition of two

procoagulant proteases as mentioned earlier.

Surface Passivation:

Surface passivation is a promising method to minimize non-specific interaction of

blood proteins and cells.39-41 Hydrophilic polymers such as poly ethylene glycol or PEG

related hydrogels and brushes have shown good protein repellency and biocompatibility.42-43

They have enhanced hemocompatibility and methods that can be used for surface

modification of these molecules are physisorption, covalent binding of functionalized

molecule are revieved elsewhere.44 Surface passivation was extensively used in the

application such as microencapsulation of cells, coatings on dialysis membrane, stents and

stents graft. Figure 1.8 shows an example for surface passivation of metal stents using

heparin-based coatings.

1.6. Hydrogels

14

Figure 1.8: Schematic showing prominent features of a heparin-coated stent. A, The stent is

coated with a polymer made of multiple layers of polyamine and dextran sulphate; B,

depolymerized molecules of heparin are covalently bound to this polymer and the –NH-CH2

(covalent) bond is described; C, pentasaccharide constituting the binding site for

antithrombin of each heparin molecule is shown; and D, continuous neutralization cycle of

thrombin is illustrated.38

1.6. Hydrogels

Hydrogels are three-dimensional polymeric networks, which has the ability to swell

in water without dissolving in it.45 Polymeric chains are the constitutents of the hydrogels

that are crosslinked either chemically or physically. If the polymer chains are not crosslinked

then the hydrophilic polymer can dissolve in water due to the theromodynamic

compatibility of the polymer and water. The presence of crosslinking points, the solubility of

the polymer is counter balanced by the retractive force of elasticity, induced by the

1.6. Hydrogels

15

crosslinking point of the network. In addition, the crosslinking in the network maintains the

three-dimentional intergrity of the hydrogels in a swollen state.

In physically crosslinked hydrogels, the chains of the polymer are connected through

the weak van der Waals interactions, ionic interactions, hydrogen bonding, or hydrophobic

interactions.46 These forces are comparatively weak so that the system can reversibly go from

soluble (sol) to a crosslinked state (gel) and vice-versa. For example, sodium alginate becomes

a gel in the presence of calcium ions (Ca+2) and turns into a sol in the absence of divalent

cations.47 Also these hydrogels are not homogenous in nature due to the clusters of molecular

entanglement or hydrophobically or ionically-associated domain formation.47

In chemically crosslinked hydrogels, the linear polymer chains are covalently bonded

with each other through a crosslinking agent (thermal or photocrosslinker). The swelling of

the hydrogels depend on the crosslink density of the network, which intern depends on the

molecular weight (MW) between two adjuscent crosslinks.45 The hydrogels also show non-

homogenicity due to the presence of regions that are poorly swelling in water caused by high

crosslinking density also known as clusters, which are dispersed within the regions of low

and high crosslink density.48-49 This is due to the aggregation of crosslinking agents, which

leads to the high crosslinking density clusters. Furthermore, the swelling of the chemical

hydrogels also depends on solvent, temperature, pH, salt concentration during the gel

formation. This may phase separate and lead to the formation of water filled „voids‟ or or

macropores.45

The physical characteristics of hydrogels resemble natural living tissues more than

any other synthetic material, due to their capability to hold large amounts of water and their

soft, rubbery tissue like consistency.50-54 Hydrogels can be classified into, homo/copolymeric

networks based on their method of preparation. The side groups of the hydrogels define the

polymer as neutral, acidic/anionic and basic/cationic networks.55-59 The following sections

1.6.1. Hydrogels from synthetic polymers

16

discusses the different classes hydrogels and their application in their utilities in the

medicinal field as biocompatible materials.

1.6.1. Hydrogels from synthetic polymers

Neutral hydrogels are generated using methacrylates, acrylamides, vinyl pyrrolidone,

and vinyl alcohols or ethylene glycol.60 Polymers synthesized from such monomers carry less

charge (or zero charge) in their structure. Among them polyethylene glycol (PEG) or

polyethylene oxide (PEO) is one of the most widely investigated and frequently employed

biomaterial in medical applications.60

PEG is a non-toxic, non-immunogenic and FDA approved polymer employed in

variety of medical applications. It also has application as a “stealth material” due to its inert

behavior against blood proteins and cells.61 It has been reported that the protein repellency of

PEG depends on their molecular weight or the length of their polymeric chains. For

example, PEG with shorter chains is prone to fouling, whereas longer chains show good anti-

fouling properties. PEG hydrogel can be generated through photo polymerization or using

photo or thermal crosslinkers. For example, PEG monomers terminated with acrylate or

methacrylate forms a network upon photo polymerization at appropriate conditions.62

The incorporation of hydrolytically degradable poly lactic acid (PLA) into the PEG

can generate biodegradable PEG gels. Further, block copolymers of PEG such as tri block

copolymers of polyethylene glycol (PEG), and polypropylene oxide (PPO - abbreviated as

PEG-b-PPO-b-PEG), and similar molecules were synthesized for specific applications.63 The

network structure of PEG related hydrogels and their applications in the medical field were

investigated in detail.

1.6.2. Responsive hydrogels

17

Another extensively investigated hydrogel is poly (2-Hydroxy ethyl methacrylate),

PHEMA because of its good mechanical and optical properties. PHEMA is widely used for

contact lenses64 and in drug delivery.65 Actually, the landmark paper from Wichterle and Lim

on poly (2-hydroxyethyl methacrylate) opened the door for hydrogels applications in

medicinal field.66 Since then the research on PHEMA and other hydrogels steadily increased.

Hydrated PHEMA can hold up to 50% water, and currently more than 10% of newly fitted

contact lenses are from PHEMA related compositions. Other polymers used in contact lenses

are methacrylic acids67 and N-vinyl pyrolidone.68

Poly (vinyl alcohol) (PVA) is another frequently used biomaterial having properties

very much similar to that of PEG polymers. The PVA must be cross-linked in order to be

used in biomedical applications. Therefore, both physical and chemical crosslinking is

employed to generate PVA hydrogels. One common way to crosslink PVA is to use

difunctional crosslinking agents such as glutaraldehyde, acetaldehyde, formaldehyde, and

other monoaldehydes.69


Responsive hydrogels have the ability to swell and de-swell as a response to changes

in the surrounding environment (e.g. pH or temperature). Accordingly, they are sometimes

called as „intelligent‟ or „smart‟ materials.70 Figure 1.9 describes the swelling and de-swelling

of the stimuli responsive gels with respect to environmental factors, which influences the

swelling of the hydrogels, for example pH or temperature.71 The responsive hydrogels could

be further classified as physical responsive hydrogels (pressure, temperature, ultrasound,

magnetic field, electric field and ligh), chemical responsive hydro gels (pH and glucose ) and

biochemical responsive hydro gels (antigen, enzyme and ligand). Among these, the most

widely investigated smart materials are those that are responsive to temperature72-73 and

pH.74-75


18

Thermoresponsive hydrogels

Thermo responsive hydrogels react to heat as trigger that determines their swelling

behavior.76 Upon, changes in the thermal energy they undergo a phase transition from a

homogeneous solution into a polymer richphase to a solvent rich phase. In thermoresponsive

hydrogels, the polymer-water interaction and the polymer-polymer interactions have a

comparable strength. Hence, even a small change in the thermal energy of the system can

drive the system across a coil-to-globule transition and induce microphase separation.

Figure 1.9: Volume collapse or volume phase transition of smart hydrogels in response to a

small change in environmental factors such as pH, temperature, light, salt type, solvent,

electric field, stress, pressure and ionic strength. Hydrogels may undergo dramatic increase

(dotted lines) or decrease (solid line) in the swelling ratio.

Further, they are catogerized into positive or negative temperature sensitive systems.

A positive temperature sensitive hydrogel has an upper critical solution temperature (UCST).

Such gels contract upon cooling below the UCST.70 Negative temperature sensitive hydrogels

have a lower critical solution temperature (LCST). Such gels contract upon heating above the

LCST.77 For example, copolymerization of NiPAAm with a more hydrophilic monomer

increases the overall hydrophilicity of the polymer, and stronger polymer-water interactions

lead to an increase in the LCST. Likewise, copolymerization with a more hydrophobic


19

monomer results in a lower LCST than PNiPPAm.78 In the crosskinked hydrogels, usually the

network chains undergoes phase separation above the LCST or below UCST, which makes

the system to collapse and thus the de-swelling.77-78

Table 1.3: Shows LCST of thermoresponsive polymers.

Abbreviation Polymer LCST (°C)

PNiPAAm [79]

PHiPAAm

PVCL [80-81]

PPO [83]

PVME [82]

MC

EHEC

PDMA [84]

PMOX

PEOX [80]

PEPA

HPC

Poly(N-isopropylacrylamide)

Poly(N-2-Hydroxyisopropylacrylamide)

Poly(vinylcaprolactone)

Poly (propyleneoxide)

Poly(vinylmethyl ether)

Methylcellulose

Ethyl (hydroxyethyl) cellulose

Poly (2-dimethylamino)ethyl methacrylate

Poly (2-Methyl-2-oxazoline)

Poly (2-Ethyl-2-oxazoline)

Poly (Ethoxypropylacrylamide)

Hydroxypropylcellulose

32

--*

31

10-20

33.8

50

65

50

--*

62

32

42

*soluble

PNiPPAm is a very important polymer in tissue engineering, as it possesses a unique

LCST (32°C) near the physiological temperature (37°C). The cells on PNiPAAm surface

enables one to easily remove or recover intact cell sheets without damaging the cells, by

simply lowering the temperature below LCST cells can be removed.78 In culture dishes (e.g.

polystyrene), the cells are recovered using proteases (e.g., Trypsin-EDTA). The recovery of

cells from PNiPAAm surface is much easier and cost effective. The LCST of NiPAAm and

other synthetic polymers having LCSTs close to the physiological temperature are


20

summarized in Table 1.3. However, it has been demonstrated that the presence of salts and

pH can also influences the phase transition temperature of these systems to a certain extent.79

pH responsive hydrogels:

pH responsive hydrogels/polymers contains acidic groups (e.g. carboxylic, sulfonic

acids) or basic (e.g. ammonium functional) groups, that either accept or donate protons upon

changing the pH of the surrounding environment.85 For example, poly acrylic acid (PAAc) is

an anionic polyelectrolyte that can be strongly ionized at high pH and consequently can be

soluble or swell (in case of crosslinked network). At low pH it is essentially neutral and no

longer soluble / swellable in water (Figure 1.11). Similarly, poly (diethylaminoethyl

methacrylate) (PDEAEMA) is a cationic polyelectrolyte that can be ionized at low pH and

consequently only soluble or swellable (crosslinked network) under such conditions.86 pH

sensitive hydrogels are frequently employed for oral drug delivery. Usually they are loaded

with an enzyme (typically glucose oxidase) that alters the pH of the local microenvironment

inside the hydrogels.87

Figure 1.11: pH-dependant ionization of polyelectrolytes. Top : Poly (acrylicacid), bottom:

Poly (N, N-diethylaminoethyl methacrylate).85-86

1.7.1. Drug delivery

21

1.7. Hydrogels in medical applications

Hydrogels are one of the fundamental materials employed in therapeutics such as

drug delivery, tissue engineering and medical devices.88 The popularity of hydrogels in

therapeutics is due to their satisfactory performance upon in vivo implantation in blood

contacting or tissue-contacting applications.89-90 The following section will discuss three

major uses of hydrogels, i.e. drug delivery, tissue engineering, and medical devices.


The change in the swelling behavior of hydrogels is an efficient tool to trigger drug

release.88 They provide promising properties for drug delivery systems because they allow

zero-order drug release, which is mandatory if drugs need to be delivered in a pulsatile

fashion.91 The oral drug delivery route is one of the most common methods in

pharmaceutical applications involving hydrogels. For example, in peroral administration,

hydrogels can deliver drugs to four major specific sites such as mouth, stomach, small

intestine and colon.92 By controlling their swelling properties or bioadhesive characteristics

in presence of biological fluid, hydrogels can be useful for releasing drugs in a controlled

manner at these desired sites.93 Additionally, they can also adhere to certain specific regions

in the oral pathway, leading to a locally increased drug concentration, and thus enhancing

the drug absorption at the release site.94 Other than oral delivery, rectal, transdermal,

subcutaneous deliveries are also of importance.95-97 Details are described elsewhere.92 The

release characteristics can be tuned by controlling factors like polymer composition,

molecular weight, crosslink density and network structure.

Environmentaly sensitive hydrogels have been the best candidates for drug delivery

and release due to their stimuli response against pH, temperature, light and other factors as

mentioned earlier.78-87 These gels swell or shrink upon changing surrounding environmental


22

conditions and can efficiently trigger drug release at specific sites in pulsatile fashion.

Another usual approach is to create micelles with the stimuli responsive gels as the outer

shell encapsulating the drug. For example, block copolymer comprised of PNiPAAm and

poly (methyl methacrylate) (PMMA) at 20°C forms micelles due to the amphiphilic nature of

the copolymer. These micelles can be loaded with the anti-inflammatory drug prednisone

acetate for the maximum drug release is obtained in 24h.98

Despite the vast number of in vitro investigations of PNiPAAm and other thermo-

responsive materials on advanced drug delivery, they are less promising in application of

regenerative medicine because they are not biodegradable.92-98 Another similar biodegradable

thermo-responsive polymer is poly [α/β-(DL-asparateisopropylamide)-co-(succinimide)]

(IPA-PSI) which contains the isopropylamide side groups, found in pNIPAM, as the source

of the thermo-responsive properties. 98, 99 However, these polymers have not been tested in

vivo for drug or cell loading/ release, and applications remain a challenge in biodegradable

thermo-responsive polymers for medical and biological technologies.100

Figure 1.12: Swelling behavior of poly (methacrylic acid -g -ethyleneglycol) hydrogels. Poly

(MAA-g-EG) as a function of the pH of the gastrointestinal tract.91

1.7.2. Tissue engineerings and medical devices

23

Figures 1.12 illustrate the hydrogel that responds to a pH change, protects protein

drugs such as insulin and calcitonin from the acidic pH of the stomach, and then release it

into the more alkaline pH of the intestine. Further, these bioadhesive polymers can protect

the proteins from degradation in the small intestine and temporarily open connections

between intestinal cells to allow proteins to penetrate into the intestine.100 However, one of

the major problems associated with these pH sensitive hydrogels is that they are not

biodegradable under physiological conditions and which limits their use in drug delivery

systems and related applications.46

1.7.2. Tissue engineering and medical devices

Tissue engineering aims to repair, replace, or regenerate its original tissue or organ

function to create artificial tissues or organs for transplantation.91 The earliest clinical

applications of tissue engineering revolved around the use of essentially flat materials

designed to stimulate wound healing.101 Later in the 1990s, a powerful cell-sheet technology

produced thin sheets of cells in cell culture. In both the applications, engineered tissue

equivalent is relatively easy to culture in vitro because oxygen and nutrient delivery to a

thin, essentially two-dimensional material is not challenging. The construct cultured ex vivo,

integration into the body is not a barrier for thin materials.102 Listed below are some pre-

requirement that need to be fulfilled for a material to be useful as scaffold in tissue

engineering:

Biocompatibility of the polymer: it should not cause any sorts of unwanted immune

response or cytotoxicity

It should have enhanced properties to mimic the natural extra cellular matrix (ECM)

and attribute cell adhesion, proliferation, migration and cell-cell interaction

Biodegradable

Easily to sterilize


24

Hydrogels are interesting candidates for successful scaffolds that possess excellent

biocompatibility because they fulfill most or all of the above criteria.103 Hydrogels are

designed in such way that, they have large pores to accommodate living tissues or dissolve or

degrade away, while releasing growth factors such as cytokines and hormones.104 That

generates pores in which cells can penetrate and proliferate comfortably.105 In summary

hydrogels have different functions in tissue engineering. They are scaffolds, deliver bioactive

molecules, and provide 3D structures that organize cells and present stimuli to direct the

formation of a desired tissue.90

Space filling scaffold

Space filling agents encompass scaffolds that function as bioadhesives also known as

biological glue.106 The basic requirement for a hydrogel is the ability to maintain a desired

volume and structural integrity for the required time.107 As a bulking material, these implants

are used to treat conditions such as urinary incontinence, vesicoureteral reflux, and are

needed for both plastic and reconstructive surgery.108

Scaffolds composed from natural polymers (see section 1.6.1) such as collagen,

alginate and chitosan are potential candidates employed as bulking agents. Implants of

porous scaffolds of RGD modified alginate and porous chitosan showed minimal immune

response in mice.109 However, the most common problem associated with bulking agents is

the successive injections to maintain the functionality. In addition, these scaffolds remain

relatively isolated from the surrounding tissue. Therefore, synthetic hydrogels are often

appropriate materials for use as anti-adhesives because cells lack adhesion receptors to them

and proteins often do not readily adsorb to them if designed appropriately.106 PEG and PVA

based hydrogels copolymerized or grafted with poly (L-lysine), (PLL) that would adhere to

the tissue on one side and provide non-adhesive brush like structure on the other side. Here,

the polymer liquid is dripped on the site of interest but not gelled.110


25

Scaffold for cell delivery

Hydrogel scaffolds are highly hydrated, three-dimensional networks that can provide

sites for cells to adhere, proliferate, differentiate, and provide chemical signals through

manipulation of the mechanical properties of the material. Hydrogel scaffolds currently help

to engineer a wide range of tissues, including cartilage, bone, muscle, fat, liver, and

neurons.106 Hydrogels have a macromolecular structure similar to the cartilage, which is a

highly hydrated tissue composed of chondrocytes and embedded in type II collagen and

glycosaminoglycans (GAG‟s). Photocrosslinked PEO and freeze dried chitosan scaffolds

posses high moduli and enhance the proliferation of cells. Both collagen and peptide

modified hydrogels have been used in the scaffolds for skeletal muscle engineering.111

Collagen has been widely used for engineering large blood vessels. Alginate hydrogels also

show potential as Schwann cell matrices in the area of nerve grafting and as scaffold to

promote hepatocyte function and liver specific proteins.91

The advantage of using hydrogels in tissue engineering is that, their tendency to hold

large amounts of water molecules in an aqueous environment, protect cells and fragile drugs

(like proteins, peptides, oligonucleotides and DNA). Hydrogels are good for transport of

nutrients to cells.90 Modified with cell adhesion ligands and injected in vivo as a liquid they

become gel at body temperature. However, poor mechanical strength, sterilization and

handling are common problems involved with hydrogels in tissue engineering.51

Medical devices

In the implantation of medical devices often requires complex surgery followed by

device implantation.112 However, with the development of minimally invasive surgery, it is

possible to place small devices inside the body using laparoscopes. These types of surgical

1.8. Challenges in biomedical applications

26

advances may create new opportunities to enable implantation of a bulky device into the

human body in a convenient way.113

1.8. Challenges in Biomedical applications

Numerous challenges remain in biomaterials development. These challenges include

targeting materials (containing drugs), to specific cells; designing materials that can sense

biochemical signals in the body; and in general developing materials with improved

biocompatibility.113

Biology and materials science need to address these challenges in an interdisciplinary

fashion. Investigation of the extracellular matrix biology, cell receptors, and immunology

will help to understand how the body responds to specific materials.114

Analogously, advances in biomaterials will create new opportunities to mimic entities

in the body (such as cells), and advances in materials characterization will aid in

understanding how materials interact with cells and tissues.1

A particular challenge in addressing material issues for biomedical engineering is that

the biological processes are not yet understood well enough to provide a clear set of design

parameters for advance specifications. Indeed, evolution of materials/ devices and knowledge

of biological processes occur simultaneously.

2. Goal and Strategy of the work

27


2.1 Goal

‘In biomaterials engineering we are missing a piece in the logical train’

Buddy.D.Ratner

One of the major complications that arise in blood-contacting medical devices is the

formation of a thrombus.1 The quick and non-specific adsorption of plasma protein is the

initial event of such a process.113-115 Consequently, the adsorbed proteins are ample enough to

trigger the necessary subsequent reactions that lead to thrombus accumulation at the surface

of the biomaterial.116 Therefore, the aim of this research was:

I. Generate a surface coating that suppresses the non-specific adsorption of plasma

proteins and their subsequent event such as blood platelet adhesion, activation, and

cellular adhesion.

II. Deduce a plausible mechanism for the protein adsorption and repellency of the

polymeric surfaces. More precisely, interrogate the previously developed models such

as size exclusion and entropic shielding to explain the protein repellency of the

surfaces.

III. Provide a detailed insight on the swelling of the surface-attached networks in water

and humidity. Draw a correlation between the swelling and the interactions of

biomacromolecules.

2.2. Strategy of the work

In view of the above, the ultimate strategy of the work was to understand and

develop the mechanisms that make surfaces or coatings bioinert and blood compatible in the


28

sense that they suppress any initial protein adsorption after implantation (Figure 2.1). The

focus is thereby on uncharged polymers and the initial hypothesis is that hydrogel (or

hydrogel coatings) fulfill this requirement due to the strong swelling in aqueous media.117

Various mechanism such as entropic shielding117 and size exclusion117,123 effect for the protein

repellency of such architectures were proposed and it was the goal to test these mechanism

by studying coatings of polymers, which are chemically similar but display rather different

swelling properties.

Figure 2.1: Bioinert surface-blood compatibility is provided by avoiding any initial unspecific

adsorption of proteins.

Another important goal of this thesis is to provide insights in mechanism that renders

many hydrogels or hydrogels coatings protein repellent. Building on previous results

obtained from PDMAAm coatings by Wörz et al., we decided to look at the correlation of

protein repellency and swelling by investigating the poly (alkylacrylamide) coatings, which

structurally similar but differ in their swelling behavior. Such materials can be obtained

through small variations in the length of the alkyl chain at N-substitution. This homologous

series will allow us to study systematically the influence of molecular structure on the


29

penetration of protein and the subsequent cell adhesion. The key parameters in the selected

materials are:

Monomers with different alkyl substituents are easy to synthesize and most of them

are commercially available.118-120

Readily prepared by free-radical polymerization along with crosslinking agent.121

Neutral with controlled hydrophilicity and water solubility.122

The hydrophilicity of the acrylamides can be modulated through length of the alkyl

chains at N-substitution. Hence, coatings of a homologous series of poly (alkylacrylamides)

namely methyl, ethyl, dimethyl, propyl, butyl and diethylacrylamides were chosen for these

studies. The coatings were prepared via an established photochemical technique, which

involves prepolymers of the desired alkylacrylamide containing few percent of a

photoreactive co-monomer.119

3.1. Synthesis of crosslinker (MABP) and anchor (Bp-Si)

30

3. Synthesis and characterizations

This chapter is organized in a way first, it discusses the synthesis of the crosslinking

agent 4-Methacryloyl-oxy-benzophenone (MABP) and the anchor silane that is used for

surface-attachment 4-[3-(Triethoxy silyl) propyloxy] benzophenone (Bp-Si). Then it

discusses the synthesis of copolymers and tricopolymers composed of alklyacrylamides with

the crosslinker MABP. The description of the surface-attachment procedure of synthesized

polymers is then followed by a section surface characterization with regard to water stability,

gel fraction, water contact angle and topography of the homologous series.

3.1. Synthesis of crosslinker (MABP) and anchor (Bp-Si)

Synthesis of 4-Methacryloyl-oxy-benzophenone (MABP)

MABP is readily available from methacryloyl chloride and 4-hydroxy benzophenone

via standard Schotten-Baumann esterification using triethylamine as an acid scavenger

(Figure 3.1). The raw product can be recrystallized from a mixture of ethylacetate and n-

hexane mixture. The final yields are typically between 80-85%.119,121 It has been

demonstrated previously that MABP was used to crosslink various polymers with an

excellent yield. 119, 123, 124

Figure 3.1: The reaction scheme of the synthesis of 4-Methacryloyl-oxy-benzophenone

(MABP).

3.1. Synthesis of crosslinker (MABB) and anchor (Bp-Si)

31

Figure 3.2: The reaction scheme of the synthesis of 4-[3(Triethoxy silyl) propyloxy]

benzophenone.

Synthesis 4-[3-(Triethoxy silyl)propyloxy]benzophenone (Bp-Si)

Bp-Si was employed in the preparation of surface-attached films. It has been

demonstrated that Si-(OEt)3 or Si-(Me)2Cl functionalities show increased affinity with silica

or glass substrates.117,123 Tri-ethoxysilane was allowed to react with alloyloxy benzophenone

in the presence of Pt-C, which was used as a catalyst in a nitrogen atmosphere. After

completion of the reaction, the catalyst was removed by filtration and the excess tri-ethoxy

silane was removed by vacuum distillation (Figure 3.2). Similarly, the reaction of dimethyl

chlorsilane and alloyloxy benzophenone resulted in 4-[3(Chlorodimethyl silyl) propyloxy]

benzophenone.124

3.2. Synthesis of alkylacrylamide copolymers

The series of alkylacrylamides: acrylamide (AAm), methyl acrylamide (MAAm),

ethylacrylamide (EAAm), dimethyl acrylamide (DMAAm), propyl acrylamide (PAAm),


32

butylacrylamide (BAAm), and diethyl acrylamide (DEAAm) are structurally similar to each

other except for the length of their alkyl substituent or the total number of carbon atoms at

N-substitution. The chemical structures of these monomers are shown in Figure 3.3.

Figure 3.3: The chemical structures of the monomers used in the preparation of N-

substituted poly (alkylacrylamide).

These monomers of homologous series were polymerized together with MABP

through free radical polymerization using 4, 41-Azobis isobutyronitrile (AIBN) as a radical

initiator (Figure 3.4). The MABP was typically adjusted to 3 mole percent. The conditions of

the polymerization of alkyl acrylamides are shown in Table 3.1.

AAm


33

PAAm (Z) Z = 1 Z = 2a Z = 2b Z = 3 Z = 4a Z = 4b

R1 Methyl Ethyl Methyl Propyl Ethyl Butyl

R2 H H Methyl H Ethyl H

Figure 3.4: The reaction scheme of the synthesis of poly (alkylacrylamides). Where, Z is the

combined number of carbon atoms at N-substitution or the length of the alkyl chains at N-

substitution.

Table 3.1: Summary of the polymerization conditions used to prepare polymers.

Polymer Monomer *Cross

linker

Initiator Solvent for

Polymerization

Solvent for

Precipitation

P(MAAm-co-MABP) MAAm MABP AIBN Methanol Diethyl ether

P(EAAm-co-MABP) EAAm MABP AIBN Methanol Diethyl ether

P(PAAm-co-MABP) PAAm MABP AIBN Methanol Diethyl ether

P(BAAm-co-MABP) BAAm MABP AIBN THF water

P(DMAAm-co-MABP) DMAAm MABP AIBN Methanol Diethyl ether

P(DEAAm-co-MABP) DEAAm MABP AIBN THF n-Hexane

*With 3 mole percent of MABP.


34

The copolymerization of AAm‟s with the higher content of MABP may leads to the

poor polymerization (also poor yield) and therefore the MABP content is kept relatively low

between 2-3 mole percent. In addition, higher amount of MABP (> 2.5 %) increases the

hydrophobicity and lowers the solubility of the respective copolymer.125-126 The spectroscopic

details of the synthesized copolymers are presented in the experimental section of the thesis

along with the synthetic procedures.

In an effort to fine-tune the hydrophilicity of the polymer, a number of tri-

copolymers were made from a mixture of a hydrophilic and a hydrophobic alkylacrylamides

along with MABP were synthesized. The two hydrophilic monomers EAAm and DMAAm,

and the two hydrophobic monomers BAAm and DEAAm were selected from the

homologous series (Figure 3.5). The hydrophobic and hydrophilic monomers were mixed in

the ratio of 75:25, 50:50, and 25:75 as shown in the Table 3.2. The detailed synthetic

procedure and spectroscopic details of the tri-copolymers can be found in the experimental

section of the thesis Table 10.3. The monomers that are employed in the synthesis of tri-

copolymers have the reactivity ratio closer to each other e.g. DMAAm r = 0.86 and DEAAm r

= 1.15.7 As outlined previously, the propagation step is depends on the reactivity (r ≈

DMAAm ≈ DEAAm ≈ 1±0.15) and structural properties of the monomers used in the

copolymerization


35

Figure 3.5: Reaction scheme of synthesis of tri-copolymers containing a mixture of

hydrophilic and a hydrophobic alkylacrylamides.

Table 3.2: Polymerization condition for the synthesis of tricopolymers.

*Polymer EAAm / DMAAm

(mmol)

BAAm / DEAAm

(mmol)

Poly(75%EAAm-co-25%BAAm-co-3% MABP) 21.80 7.30

14.55

21.80

7.30

14.55

21.80

Poly(50%EAAm-co-50%BAAm-co-3% MABP) 14.55

Poly(25%EAAm-co-75%BAAm-co-3% MABP) 7.30

Poly(75%DMAAm-co-25%DEAAm-co-3% MABP) 21.80



*75% = 73.5 %, 50% = 48.5% and 25% = 23.5% and MABP 3% = 0.90 mmol

3.3. Surface-attached networks

36

3.3. Surface-attached networks

Polymer networks are macromolecular systems in which all polymer chains are

interconnected with each other, which make them insoluble in solvents.127 One of the

widely employed methods is to use a crosslinker, which upon irradiation119 or heating leads

to the generation of networks.128 It is also essential to have stable coatings to the surfaces that

are applicable in medicine, especially when they are in contact with blood flow (e.g.,

ventricular assist devices). Hence, one plausible way to generate such coatings is to tailor

them through covalent binding.117 In this study, the photoactive MABP was used as a

crosslinker to prepare networks and for the covalent attachment to the surface.129-132

Accordingly, all the monomers that are shown in Figure 3.3 were copolymerized with 2-3

mole percent of MABP in order to achieve the sufficient crosslinking. It was demonstrated

that even a 2.5 mole percent of MABP was ample to achieve efficient crosslinking, which led

to the formation of stable surface-attached networks.129

Figure 3.6: Mechanism of the benzophenone based crosslinking process. Upon illumination

benzophenone forms biradical, which abstracts the hydrogen atom from a neighbouring C-H

group of a polymer leaving a two carbon radicals, which can recombine and establish

covalent bond.119,123

The mechanism of benzophenone based crosslinking is described in the literature by

Chang et al.121 and Prucker et al.119 Briefly, benzophenone absorbs UV light which results in

an n-π* or π-π* transition (depending on the wavelength) of the carbonyl group. The

3.3.1. Preparation of surface-attached poly(alkylacrylamide) networks

37

resulting biradicaloid triplet abstracts a hydrogen atom from a neighboring C-H group,

leaving behind two carbon-based radicals, which upon recombination generate a covalent

bond between the two involved molecules. When the recombination occurs with a surface-

group, surface attachment results and when the reaction occurs with another polymer chain,

the two chains become covalently linked to each other (Figure 3.6). Similarly, several groups

on a polymer chain reacts in this fashion and such a process leads to the crosslinking of

surface coating.119

Figure 3.7: Schematic diagram illustrates the simultaneous surface-attachment and

crosslinking process.119,123 The Bp-Si coated surface after deposition of polymer and after

irradiation yield the surface-attached network.

3.3.1. Preparation of surface-attached poly (alkylacrylamide) networks

The cleaned glass or silica substrates were spin coated with Bp-Si and annealed at

120°C overnight. Upon annealing, Bp-Si reacts with the surface and forms a covalent bond

with the substrate surface as depicted in the Figure 3.7. This Bp-Si covered substrate was spin

3.3.1. Preparation of surface-attached poly (alkylacrylamide) networks

38

coated with poly (alklyacrylamide-co-MABP) solution and after evaporation of the solvent;

the layers were illuminated with a total dosage of 4 J/cm2 from a UV-C-light source (254

nm). Upon irradiation, the benzophenone present in the precursor polymer formed a bi-

radical, abstracting the neighboring protons; thereby leading to simultaneous crosslinking

and surface attachment of the deposited polymer film.119 The uncrosslinked polymer

molecules were removed by solvent extraction. The condition that was employed in the

process of making networks for different polymers of the homologous series is summarized

in Table 3.3.

As mentioned in Table 3.3, polymers with shorter alkyl substituent such as PMAAm,

PEAAm and PDMAAm were found to have high solubility in methanol or ethanol. Polymers

such as PPAAm, PBAAm and PDEAAm were well soluble in propanol / isopropanol and

butanol. The used condition was sufficient to generate homogenous network and the film

thickness was controlled by the concentration of the polymer solution used (e.g. 10 mg/ml

and the film thickness was found to be between 70-90 nm) for the spin coatings.

Table 3.3: Conditions for preparation of surface-attached poly (alkylacrylamide) networks.

Polymer Solvent Concentration (mg/ml) Speed (rpm) Time (s)

P(MAAm-co-MABP) Ethanol 10 2500 60

P(EAAm-co-MABP) Ethanol 10 2500 60

P(PAAm-co-MABP) Butanol 10 2500 60

P(BAAm-co-MABP) Butanol 10 2500 60

P(DMAAm-co-MABP) Ethanol 10 2500 60

P(DEAAm-co-MABP) Ethanol 10 2500 60

3.3.2. Gel fraction of the poly (alkylacrylamide) networks

39

3.3.2. Gel fraction of the poly (alkylacrylamide) networks

As it was mentioned in the previous section, MABP showed absorptions at 365 nm

and 254 nm due to the carbonyl group of benzophenone. Schuh128, Bunte132 and Chang124

have extensively studied the kinetics of the photochemical crosslinking of MABP containing

polymers. They found that the rate of benzophenone crosslinking process followed the first

order reaction kinetics.124,132 The crosslinking was faster at 254 nm (UV-C light) than at 360

nm (UV-A light) due to the higher extinction coefficient of the π-π* than the n-π* transition

in the carbonyl functionality.132 Based on their results the films generated in this study

irradiated with UV-C light at 254 nm for four minutes. The films were then rigorously

extracted in solvent and the thickness before and after the extraction (gel content) were

compared to evaluate the conversion. The calculated gel contents are shown in Table 3.4.

Table 3.4: Gel fraction of the surface-attached poly (alkylacrylamide) networks.

Polymer Thickness before

crosslinking (nm)

Thickness after

crosslinking (nm)

Gel content

(%)

P(MAAm-co-MABP) 82 75 91

P(EAAm-co-MABP) 86 77 90

P(DMAAm-co-MABP) 72 68 95

P(DEAAm-co-MABP) 95 82 86

P(PAAm-co-MABP) 67 65 97

P(BAAm-co-MABP) 61 58 95

The calculated gel fractions demonstrate that the irradiation conditions were

sufficient to achieve high gel contents. Even, after 30 seconds of irradiation a gel content of

around 60% was found, which increased to 80% after 60 seconds of irradiation. A maximum

gel content of 90% was typically reached after four minutes. Further illumination of UV light

damages the film as demonstrated by Bunte et al.132 Thus, the deposited polymer layers were

3.3.3. Stability of the surface-attached poly (alkylacrylamide) layers

40

crosslinked for four minutes and extracted in the suitable good solvent to remove loosely

bound materials.

3.3.3. Stability of the surface-attached pol (alkylacrylamide) layers

The stability of the films is an important aspect for any long-term applications and

especially for medical uses during which the coatings have to withstand the high shear-stress

in blood flow. To investigate the stability of the films, the extraction experiments were

performed over longer period of time and the layer thickness of the samples were recorded.

Figure 3.8 shows for the different materials and demonstrates that no significant change

occurred during the experiment. All layers were rather stable even after one month of

exposure to a good solvent. Further extension of the experiment up to 6 months showed still

unaltered thickness values. This demonstrates that the surface-attached films were stable in

aqueous medium for a prolonged period of time. The results of investigation under shear-

stress condition will be discussed in Chapter 7.7.

Figure 3.8: Stability of surface-attached poly (dimethylacrylamide-co-MABP). The sample

was taken out of the aqueous medium at different time intervals, dried using nitrogen air and

the film thickness is measured using ellipsometry at variable angle mode. Where, Z=1, poly

3.4. Contact angle and Topography

41

(MAAm-co-MABP); Z= 2a, poly (EAAm-co-MABP), Z=2b, poly (DMAAm-co-MABP); Z= 3,

poly (PAAm-co-MABP); Z=4a, poly (BAAm-co-MABP), Z=4b, poly (DEAAm-co-MABP).

Light gray indicates the original film thickness and dark gray indicates the thickness after a

month.


In medical applications, an implantable biomaterial interacts with the surrounding

tissue (cells) critically through their interfaces. Such interactions largely depend on the

surface properties of the materials such as wettability, topography or roughness and charge

distribution on the surface.132 Understanding this phenomenon is crucial to the

comprehension of several fundamental queries related to protein-material (also cell-

materials) interaction for the development of biomaterials in the field of medicine.

Consequently, in this work, surface properties such as wettability and topography of the

surface-attached poly (alkylacrylamides) networks are investigated using contact angle

measurements and atomic force microscope, respectively.

3.4.1. Contact angle measurements

The wettability of the surface-attached poly (alklyacrylamide) networks was

characterized by contact angle measurements. Static contact angles were measured using the

sessile drop method though an Optical Contact Angle Meter (OCA 20) setup (Dataphysics,

Filderstadt, Germany), with a liquid dispenser and an image processing program which

provided an automatic determination of the contact angles. Two samples of each surface type

were measured at different points of the same sample and the average value was taken into

account as the static contact angle. All the measurements were performed at ambient

temperature and at 37°C using Milli-Q water. 1µl volume of water was dispersed onto the

surface at the rate of 0.5µl/s and contact angles were measured.


42

Figure 3.9 shows the static water contact angles values obtained on the various poly

(alkylacrylamide) coatings. As a reference, medical grade polyurethane was also used. It is

evident from the graph that most of these gels are moderately wettable surfaces having a

contact angle from 55-75°. The coatings made from poly (PAAm-co-3%MABP) and poly

(BAAm-co-3%MABP) are very hydrophobic with a advancing contact angle of 90-92°. The

important factor determining water wettability is the chemical structure of the polymeric

network at the interface.

Consequently, the water contact angles increases with length of the alkyl chain at N-

substitution or combined number of carbon atoms at N-substitution. However, the

differences between the coatings made from 4a (BAAm) and 4b (DEAAm) indicate that

further factors may also important. The longer butyl substituent of 4a may segregate to the

surface more efficiently than the two ethyl substituent of 4b. This might explain the

outstanding hydrophobic nature of the material of the series.

Figure 3.9: The water contact angles as a function of the combined number of carbon atoms

(Z). Where, Z=1, poly (MAAm-co-MABP); Z= 2a, poly (EAAm-co-MABP), Z=2b, poly

(DMAAm-co-MABP); Z= 3, poly (PAAm-co-MABP); Z=4a, poly (BAAm-co-MABP), Z=4b,

poly (DEAAm-co-MABP) and PU is polyurethane.


43

3.4.2. Topography

The topography or roughness may also influence the interaction of surface with

biomolecules. In this work, AFM (Nanoscope IIIa, Digital Instrument) was used to examine

the topography of the surface-attached alkylacrylamide coatings on silica substrate (see

section 3.3.1 of this chapter). The microscope was operated in tapping mode, using

commercial tips with a resonance frequency of ≈330kHz, and a spring constant of ≈42 N/m.

The AFM micrographs were recorded in air at room temperature.134

Table 3.5: Thickness and roughness of the surface-attached poly (alkylacrylamides).

Polymer Thickness (d) Roughness (Ra)

P(MAAm-co-MABP) 49 nm 0.5 nm

P(EAAm-co-MABP) 38 nm 6 nm*

P(PAAm-co-MABP) 60 nm 0.98 (≈1 nm)

P(BAAm-co-MABP) 59 nm 0.78 nm (≈1 nm)

P(DMAAm-co-MaABP) 42 nm 1.4 nm

P(DEAAm-co-MABP) 41 nm 1.68 (≈2 nm)

†PU --† 27 nm

*Sample quality was bad, †reference samples (medical grade polyurethane used in VAD)

The AFM results of surface-attached poly (alkylacrylamide-co-MABP) networks are

shown in Table 3.5. The thickness of the deposited films ranged from 40 nm to 60nm and the

roughness of the films was found always less than 2 nm, which indicates that there was no

significant difference in the topography among the homologous series (AFM micrographs are

presented in the appendix). In addition, the results of the poly (alkylacrylamides) were

----------------------------------

AFM experiments were performed by Mr. Xiaoqiang Hou, CPI-IMTEK, University of Freiburg, Germany.

3.5. Conclusion

44

compared with that of the reference polyurethane (medical grade polyurethane) used in

ventricular assist devices, showing that it is rougher (around 27 nm) than the homologous

series being investigated in this study.

3.5. Conclusion

A series of alkylacrylamide copolymers containing 2-3 % of photo crosslinker

(MABP) were successfully synthesized through a free radical polymerization. These

polymers were then deposited onto substrates and irradiated using UV-C-light (λ = 254 nm

or 4 J/cm2) for four minutes. This led to a simultaneous crosslinking and surface-attachment

of the films. The stability of the layers in water was investigated and it was found that these

layers were stable under these conditions for a prolonged period of time. Finally, water

contact angle experiments reveal that all alkylacrylamide coatings showed a moderate

wettability between 50-75°. Interestingly, BAAm was more hydrophobic with a contact

angle of 112° because the longer butyl substituent was segregated more efficiently than other

alkyl substituent of the homologous series. The roughnesses of the poly (alkylacrylamide)

coatings were examined using AFM and it was found that all coatings were rather small.

4.1.Theory of swelling

45

4. Swelling of the surface-attached poly (alkylacrylamide) networks

One of the most important properties of the hydrogels is their swelling behavior in

appropriate solvent.135 Several theoretical models were proposed to describe the swelling of

hydrogels. This chapter is organized in the following way; Firstly, it provides the background

of isotropic and anisotropic swelling of the polymeric coatings followed by a descripsion of

Surface Plasmon Resonance spectroscopy (SPR). SPR was used to measure the swelling of the

surface-attached poly (alkylacrylamide) networks in water and in humid air and the lastly

the results of these measurements were described and discussed. Finally, the swelling in

water and in humid air compared.

4.1. Theory of swelling

4.1.1. Isotropic swelling

Flory and Rehner developed a model to explain the swelling behavior of cross-linked

networks or hydrogels.136 A crosslinked polymer gel whilst immersed in water is allowed to

reach equilibrium with its surrounding is governed by two opposing forces: the

thermodynamic force of mixing, and the retractive force of polymer chains.

∆𝐺𝑡𝑜𝑡𝑎𝑙 = ∆𝐺𝑚𝑖𝑥𝑖𝑛𝑔 + ∆𝐺𝑒𝑙𝑎𝑠𝑡𝑖𝑐 (Eq.4.1)

At equilibrium, these two forces are equal, therefore the Gibbs free energy is written

as follows137

∆𝐺𝑚𝑖𝑥𝑖𝑛𝑔 = − ∆𝐺𝑒𝑙𝑎𝑠𝑡𝑖𝑐 (Eq.4.2)

Where ∆Gelastic is the contribution due to elastic retractive forces developed inside the

gel, and ∆Gmixing is the result of spontaneous mixing of water molecules with polymer

chains.125

∆𝐺𝑚𝑖𝑥𝑖𝑛𝑔 = 𝑘𝑇 𝑛1 ln𝜑1 + 𝜒 𝑛2𝜑2 (Eq.4.3)

4.1.1. Isotropic swelling

46

In the above equation 𝑛1 , is the ratio of solvent molecules, 𝜑1 is the volume fraction

of the solvent, 𝑛2 is ratio of the polymer repeat units, φ2 is the volume fraction of the

polymer, k is the Boltzmann constant and T is the temperature. χ represents the Flory-

Huggins interaction parameter.

The elastic free energy is represented purely by the entropic components. Therefore,

the expression for ∆Gelastic is equal to the change in the entropy, ∆Selastic due to the elastic

stretching of the polymer chains in the network and could be written as,

∆𝐺𝑒𝑙𝑎𝑠𝑡𝑖𝑐 = −𝑇∆𝑆𝑒𝑙𝑎𝑠𝑡𝑖𝑐 (Eq.4.4)

∆S𝑒𝑙𝑎𝑠𝑡𝑖𝑐 = − 𝑘𝜐

2 (3𝛼𝑠

2 − 3 − ln𝛼𝑠 3 ) (Eq.4.5)

Where k is the Boltzmann constant, T is the absolute temperature; υ is the number of

crosslinks in the network, and α is the elongation ratio, e.g. the linear deformation of the

network for isotropic stretching. Upon several rearrangements Eq.4.1-4.5 2-4, one obtains the

well known Flory-Rehner model:

ln 1 − φe + φe + χφe2 =

υ

Nav

V1

V0 φe

2 − φe

1/3 (Eq.4.6)

Here, the volume fraction of the polymer, φ2 is represented as, φe to emphasize the

equilibrium swelling and the prefactor on the right hand side of the equation, represents the

number of strands per unit volume in the original network. Now, the number of average

molecular weight between two adjacent crosslinks, Mc can be evaluated using the following

equation:138

𝑀𝑐 =2

𝑀𝑟−

𝜐

𝑉1 ln 1− 𝜑2 + 𝜑2+ 𝜒𝜑2

2

𝜑21/3

− 𝜑22

(Eq.4.7)

Where Mr is the molecular weight of the repeating unit and by using the Flory

polymer solvent interaction parameter, χ one can calculate the Mc. χ usually lies between 0

4.1.2. Anisotropic swelling

47

and 1. For a good solvent, χ is smaller than 0.5 whereas for a bad solvent it is greater than

0.5.139


When a hydrogel is attached to a surface, the chemical linkage of the network to the

surface prevents the swelling parallel to the substrate, effectively confining the volume

change to one dimension.122 This influences many important properties of the hydrogels such

as mechanical strength, dynamics and permeability of the network. Toomey et al. described

the differences in the swelling of unconstrained hydrogels and surface attached hydrogels

using Flory-Rehner model.136-137 Therefore, the free energy of mixing for anisotropic (one-

dimensional) swelling can be written as,

∆𝐺𝑚𝑖𝑥𝑖𝑛𝑔

𝑘𝑇 = 𝑛1 ln(1 − 𝜑2) + 𝜒 𝑛1𝜑2 (Eq.4.8)

Where k is the Boltzmann constant, T is the absolute temperature, n1 is the number

of solvent molecules, φ2 is the volume fraction of the polymer in swollen gel, and χ is the

Flory polymer-solvent interaction parameter. Moreover, the elastic free energy can written

as,

∆S𝑒𝑙𝑎𝑠𝑡𝑖𝑐

𝑘𝑇= 𝜐

𝑑

2 (𝛼𝑠

2 − 1 − ln𝛼𝑠 ) (Eq.4.9)

Here, υ is the number of cross-links in the network, and d is the number of

dimensions in which the network can swell. Therefore, the relationship between the linear

deformation at equilibrium and the degree of freedom of the network is given as,

𝛼 ≈ 1

(1

2 –𝜒)𝜑0𝑁𝑐

− 1

(𝑑+2)

(Eq.4.10)


48

The equation 4.10 reveals that the swelling is confined to fewer dimensions, α

depends strongly on the crosslink density (1/Nc). Here Nc is the number of segments

between two crosslinks. The volumetric swelling (S) is equal to the linear deformation α to

the power of the dimensions in which it can swell. The relation can be derived as follows

𝛼𝑢𝑐 ≈ 1

𝜑0𝑁𝑐 −

1

5 ; 𝑆𝑢𝑐 ≈

1

𝜑0𝑁𝑐 −

3

5 (Eq.4.11)

𝛼𝑠𝑎 ≈ 1

𝜑0𝑁𝑐 −

1

3 ; 𝑆𝑠𝑎 ≈

1

𝜑0𝑁𝑐 −

1

3 (Eq.4.12)

As mentioned earlier, surface attachment limits the swelling to one dimension. The

network experiences a higher osmotic pressure, which is partially relieved by further

stretching in its swelling direction perpendicular to the surface, thus leading to a higher

linear swelling than unconstrained network.

Therefore,

𝑆𝑠𝑎 = 𝑆𝑢𝑐 5/9

(Eq.4.13)

From the above equation, the volumetric degree of swelling in the surface-attached

network is approximately equal to the square root of the degree of swelling in the non-

attached network.122

4.2. Surface Plasmon resonance spectroscopy (SPR)

Surface Plasmon resonance spectroscopy is an optical method used for measuring

thickness and refractive indices of thin film on a metal surface.140 This technique is based on

the fact that at certain conditions, surface plasmon polaritons (i.e. propagating

electromagnetic waves at the interface of the metal dielectric layer) on a metallic film can be

excited at specific resonance frequency.141 The instrumentation and used configuration of the

4.2. Surface Plasmon resonance Spectroscopy (SPR)

49

SPR setup is given in chapter 10.2.8 of this thesis. There are two methods employed to excite

plasmons at the interface. They are the Kretschmann (Figure 4.1A) and Otto

configuration.142-143 The Kretschmann configuration is the most widely used geometry in SPR

and the following discusses briefly about the principle of this configuration.

The Kretschmann configuration143 is an extensively used geometry in many SPR

instruments. It excites the surface plasmon based on attenuated total reflection (ATR). Here,

a beam of p-polarized light (transversal magnetic, TM) is reflected from gold (approx. 50 nm)

through a prism or glass and the front side is in contact with air or solution (Figure 4.1A). If

the angular scans of reflectivity is measured (Figure 4.1B), when a incident angle (θ) is

greater than critical angle (θc) i.e. θ > θc, most of the incoming light is reflected, while a

small part of the light penetrates outside of the glass, which is called as the evanescent wave.

The surface plasmons are observed as a minimum in the reflectivity at the angle of θsp. The

critical angle (θc) is the critical angle at which the total internal reflection begins to occur.

By using such surface plassmon spectrum one can determine the optical components of the

dielectric layer (in this case polymer film) with the help of Fresnel formalism.144

If the thickness of the dielectric layer (e.g. polymer film) is larger than 300 nm,

optical waveguide modes are observed in addition to the surface plasmon resonance.147

Usually these waveguides are observed as a narrow dips measured at p-(TM) and s-(TE)

polarizations and the number of modes that are excited is purely depends on the film

thickness. The typical optical waveguide spectrum (OWS) is shown in Figure 4.2. By

simulating these waveguide spectra using Fresnel equation one can derive the film thickness

and refractive index of the dielectric layer.148

4.2. Surface Plasmon resonance Spectroscopy (SPR)

50

Figure 4.1: (A) Schematic diagram of Kretschmann configuration. (B) SPR reflectivity scan

measured for Au, where the reflectivity (R) is measured as a function of incident angle (θ).

The angle at which the total internal reflection starts is known as critical angle (θc). The

surface plasmons are observed as a minimum in the reflectivity of the incident light at θsp. 144-146

(B)

(A)

4.3. Swelling in water

51

Figure 4.2: Reflectivity scan obtained from poly (DMAAm-co-1% MABP) layer on gold

illustrating multiple waveguides in OWS. Square symbol represents the original data from

OWS and black line is the Fresnel calculations giving a thickness of 716 nm and refractive

index of 1.51.


The swelling behavior of surface-attached poly (alkylacrylamide-co-MABP) networks

was studied using optical waveguide spectroscopy (see 10.2.8 for instrumentation of SPR/

OWS) .149 The advantage in OWS is that multiple waveguides in the angle spectra of OWS

allow determining the layer thickness and refractive index of the surface-attached layer,

independent from each other.150 The swelling measurements can be divided into three major

steps:

1. Determination of gold layer thickness to characterize the blank substrate

2. Deposition and characterization of an anchor layer

3. Deposition and investigation of the poly (alkylacrylamide) layer in N2 and aqueous

environment.


52

Determination of Gold layer thickness

A LaSFN9 glass slides were used to which 1-3 nm of chromium was evaporated. This

layer serves as an adhesion layer for a 50 nm gold layer, which was also deposited through

evaporation.145 These slides are standard substrates for SPR and OWS and are reusable for a

minimum of three to four times. The optical properties of the gold layer were measured

using SPR in a nitrogen atmosphere. Typical resonance angle for this setup was found around

25°.The film thickness and refractive indices were computed according to Knoll et al.145 using

the software Winspall 3.1.0, which is based on Fresnel formalism. Typical gold layer

thickness of 50-51 nm were detected and (Figure 4.3 and Table 4.1).

Deposition and characterization of an anchor layer

A film of a water insoluble polymer such as polyurethane or polystyrene was

deposited to serve as an anchor layer to which the poly (alkylacrylamide) layers are later

linked. Typically, 10-15 nm of film were dip-coated on top of the gold (Table 4.1). Then the

polymer with photo crosslinker (MABP) deposited crosslinked and anchored to the

underlying polymer coated substrates as described in Chapter 3.

Investigation of the surface-attached poly (alkylacrylamide) layers

Finally, the thickness of the deposited films or surface-attached networks was

measured in nitrogen and water to determine the swelling. Figures 4.4 and 4.5 depict the

reflectivity scans of surface-attached poly (DMAAm-co-3%MABP), poly (DEAAm-co-

3%MABP) and poly (BAAm-co-3%MABP), respectively. The calculated film thickness (d)

and the refractive index (n) of the polymer films are shown in Table 4.2. These two are

selectively shown to represent the hydrophilic and hydrophobic polymers of the homologous


53

series. The reflectivity scans of other surface-attached networks are given in appendix of the

thesis (See Figures 11.1-11.4).

Figure 4.3: Reflectivity scan of a gold substrate before and after deposition of PU layer.

Where, the reflectivity is recorded as a function of angle of incident (θ). The minimum in the

reflectivity indicates the excitation of a surface plasmon. The solid lines are Fresnel

calculations both the samples. The calculated optical components are summarized in Table

4.1.

Table 4.1: Optical components obtained from SPR measurements of gold and PU.

Material *Film thickness,

d (nm)

Permeability constant

(real), 1

Permeability constant

(imaginary), 11

Glass 0 3.4036 0

Chromium 1.12 -6.3 18

Gold 50.91 -12.25 1.29

PU 13.7 2.2712 0

N2 0 1 0

*All the measurements were performed at room temperature (≈25°C) in N2 atm.


54

Table 4.2: Optical components derived from the reflectivity scans of poly (DMAAm-co-

3%MABP) and poly (BAAm-co-3%MABP) networks using a Fresnel formalism.

Medium Poly (DMAAm-co-3%MABP) Poly (BAAm-co-3%MABP)

d (nm) N *SF d (nm) n *SF

N2 334 1.52 3.50 411 1.49 1.18

H2O 1169 1.41 485 1.47

*Swelling factor, d is the film thickness and n is the refractive index.

Figure 4.4: Reflectivity scan of poly (DMAAm-co-3%MABP) abbreviated as PDMAAm.

Squares and circles symbol represent the measurement in N2 and in water, respectively. Solid

lines are the results of Fresnel calculations used to describe the measured data and to derive

the layer thickness and refractive index.

The optical components of the film were obtained by modelling the spectra using a

Fresnel equation.151 As noticed in the Figures 4.4 and 4.5, there are significant differences

between the measurement in nitrogen atmosphere and in water. For example, in nitrogen

atmosphere, one waveguide was observed for poly (DMAAm-co-3% MABP).


55

Whereas, in water multiple and narrow waveguides were observed. Analogously, for poly

(BAAm-co-3% MABP), the reflectivity scans showed only one waveguide in both these

media. However, the resonance angle of the waveguide in presence of water was shifted to a

higher angle of incidence that may be due to the change in interface from nitrogen to water

medium. From the respective thickness in the dry (dN2) and wet state (dw), the swelling factor

of the surface-attached network can be determined by using the following equation.122

𝛼 =𝑑𝑤

𝑑𝑁2 (Eq.4.14)

Additionally, the decrease in the refractive index (or dielectric constant, ε´) also

indicates that the network incorporates a significant amount of water from the aqueous

medium. However, this decrease depends on the degree of swelling of the polymers.

Figure 4.5: Reflectivity scan of poly (BAAm-co-3%MABP), abbreviated as PBAAm. Squares

and circles symbol represent the measurement in N2 and in water, respectively Solid lines are

the results of Fresnel calculations used to describe the measured data and to derive the layer

thickness and refractive index.


56

Figure 4.6 illustrates the results obtained from surface-attached poly (alkyl

acrylamide) layers. The linear swelling ratio (α) is plotted as a function of the length of the

alkyl substituent in the acrylamides. The results demonstrate that the polymers with shorter

alkyl substituent, PMAAm, PEAAm and PDMAAm, i.e. a combined number of carbon atoms

at N-substitution of one or two (Z = 1, 2a and 2b), swell to the factor of three to five

(PMAAm = 5.32, PEAAm = 4.6 and PDMAAm = 3.5). In contrast, the polymers with longer

alkyl substituent, PPAAm and PBAAm do not swell appreciably in water.

The volume fractions of polymer (φ2) and solvent (φ1) in the network can be derived

from the swelling factor α by using equations 4.15 and 4.16, respectively.

𝜑2 =1

𝛼=

𝑑𝑑𝑟𝑦

𝑑𝑤𝑒𝑡 (Eq.4.15)

𝜑1 = 1 −1

𝛼 = 1 − 𝜑2 (Eq.4.16)

Figure 4.7 shows the calculated volume fraction of the polymer and water of the

swollen network at room temperature (25°C) and physiological temperature (37°C). The

results show that the volume fraction of water in the swollen network, decreases as we go

along the homologous series from methyl (0.80) to butyl (0.20) and the volume fraction of

polymer increases accordingly in that direction. This means, that polymers with a lower

number of carbon units at N-substitution or shorter alkyl chains at N-substitution can retain

more water in the swollen state due to strong interaction or coordination with water

molecules and vice-versa.

The difference in the swelling behavior of the poly (alkylacrylamides) is a clear

consequence of the increasing hydrophobicity with increasing number of carbon atoms on

the alkyl substituent or with the number of alkyl groups present. When the N-alkyl-

substituted group is longer than an ethyl, attractive inter- segment interactions i.e.


57

hydrophobic association between alkyl groups become dominant, and the polymers become

insoluble in water, thereby exhibiting little or no swelling.

Figure 4.6: Swelling factor (α = dwet/ddry) of the surface-attached poly alkylacrylamide

networks. The swelling decreases as the length of alkyl chain increases.

4.3. Sweelling in water

58

Figure 4.7: (A) The polymer volume fractions (φ2) in the swollen network. (B) The water

volume fractions (φ1) of swollen network. Black dot and lines represent the volume fractions

measured at room temperature (25°C). Gray dot and lines represent the volume fractions

obtained at 37°C. The amount of water (volume fraction of water, φ1) in the network

decreases as the length of the alkyl group increases from methyl to butyl because the

swelling decreases.

4.3. Sweelling in water

59

Another interesting aspect is the difference in the swelling of mono-substituted

alkylacrylamides and di-substituted alkylacrylamides. PEAAm and PDMAAm have an equal

number of carbon atoms at N-substitution. The swelling of PEAAm was found to be ≈4.6, but

the swelling of PDMAAm was 3.35, because, in PEAAm both NH and C=O can participate in

the hydrogen bonding. This is not possible for PDMAAm and therefore binds less water than

PEAAm.

In addition, the swelling of the surface-attached poly (alkylacrylamide) networks

were measured at various temperature ranged from 15-55°C. The results are shown in Figure

4.8. Among the homologous series of poly (alkylacrylamides), PDEAAm showed a significant

difference in the swelling behavior with respect to the temperature than other polymers of

the series.

Figure 4.8: Swelling factor (α = dwet/ddry) of the surface-attached poly (alkylacrylamide)

networks as a function of temperature (15°C to 50°C). Figure A and B demonstrate that

swelling with respect to temperature remains more or less same for all the polymers except

for PDEAAm in which the hydrogel shrinks at around 30°C (B).

4.3. Sweelling in humid air

60

Accordigly, PDEAAm network swelled to the factor of around 1.7 at 25°C and

decreased to 1.3 at 37°C. The amount of water in the network was greatly reduced from 45%

to 23% as the temperature increased from 25°C to 37°C. It indicates that PDEAAm is a

thermo-responsive polymer, which has a trasition temperature of around 30°C. Above this

temperature, the chains of the network collapse / phase separate due to the strong interaction

of polymer-polymer than polymer-water. Consequently, the water molecules from the

surface-attached network were expelled and hence the swelling of the network decreased.

4.4. Swelling in humid air

Polymer networks that swell in water are typically hydroscopic and also swell in

humid air. It is important to consider this property where polymer layer thicknesses are to be

determined. Depending on variations in the environment this value might be strongly

overestimated.152 As a result, numerous studies of the humidity based swelling of a thin

polymer films, especially on polyelectrolyte layers were undertaken.153 For example,

Hernandez et al. reported that the volume fraction of the polymer film increases in

accordance with the raise of humidity in the environment. They concluded that such a

phenomenon mimics more of Flory sorption behavior than Langmuir process.154 However,

the clear understanding on this behavior and mechanisms of the adsorption is still lacking.

The motivation of this study is to interrogate the sorption behavior of surface-

attached poly (alkylacrylamide) networks in different humid environments, and to predict

the existing relationship between water and humidity swelling. A series PDMAAm networks

consisting of different amounts of photocrosslinker (1%, 2.5%, 5%, and 10%) was

synthesized for this investigation. PDMAAm is a hydrophilic polymer and exhibits good

swelling in water as discussed in the previous section. Furthermore, this particular series of

copolymers will allow us to look at the influence of crosslinking density on humidity

swelling. In the next step, we have studied the humidity swelling of the other poly

4.4.1. Experimental conditions for humidity swelling

61

(alkylacrylamides) to elucidate the influence of molecular structure (hydrophilicity-

hydrophobicity) of the on humidity swelling.

4.4.1. Experimental conditions for humidity swelling

Optical waveguide spectroscopy (OWS) was used to determine the swelling of the

surface-attached gels in contact with humid air. The sample was brought into contact with a

flow cell, which was kept at constant temperature and relative humidity. The humidity

inside the flow cell was maintained through a nitrogen stream humidified by a wash bottle

containing a saturated salt solution (Figure 4.9). The high salt content of the solution reduces

the vapor pressure of the water to a distinct value. The relative humidities of the saturated

salt solutions are summarized in Table 4.3.

Figure 4.9: Instrumentation of OWS used in the measurements of humidity swelling.

Saturated salt solutions were used to achieve the desired humidity in the measurement

region of the flow cell. The humidity within the system was determined using a humidity

sensor.

4.4.2. Results of PDMAAm networks in humid air

62

Table 4.3: Relative humidity of saturated salt solution at 25°C.152-154

Sat. salt solution Humidity (%) Sat. salt solution Humidity (%)

LiCl 25 NaCl 77

K2CO3 46 KCl 84

NaBr 60 KNO3 92


The polymer films are deposited on gold substrates by following the procedures

described previously in chapters 3.3 and 4.4 of this work. The thickness of the film was

adjusted to 450-650 nm and reflectivity scans are recorded as a function of the angle of

incidence (θ). The appearance of multiple waveguides in the reflectivity spectra endows

simultaneous estimation of thickness and refractive index of the given polymer film.150

A fundamental aspect in humidity swelling is the question of when equilibrium is

attained.35 To examine this, the kinetic measurement was performed, by selecting an angle

(θk) from reflectivity scans obtained from nitrogen atmosphere, typically last waveguide near

40° - 50°. Figure 4.10 illustrates the kinetic measurement of surface-attached gels, where the

reflectivity is monitored as a function of time (minutes). After minimal or no change in the

reflectance, kinetics is stopped and reflectivity scan is recorded at equilibrium to assess the

(swollen) film thickness.

The reflectivity scans of the surface-attached poly (DMAAm-co-1%MABP), poly

(DMAAm-co-2.5%MABP), poly (DMAAm-co-5%MABP) and poly (DMAAm-co-10%

MABP) networks are shown in Figures 4.11 - 4.14. A significant shift in the waveguides was

observed as the humidity was varied. The resonance angle of the waveguides typically shifted

to higher angles of incidence (Δθ ≈ 5-12°). In the case of poly (DMAAm-co-1%MABP), a new

4.4.2. Results of PDMAAm networks ín humid air

63

waveguide appeared which indicates that there is a strong change in the optical thickness of

the film.

Figure 4.10: Kinetics measurement of surface-attached poly (DMAAm-co-1%MABP) at

various relative humidity, where the reflectivity is measured as a function of time. In most

cases, it takes 30-35 minutes to reach equilibrium.

The change in the film thicknesses and refractive index of the surface-attached

PDMAAm networks were determined by modelling the data using Fresnel formalism.145

Accordingly, the swelling factor of the PDMAAm networks at different humidity was

determined by comparing the swollen layer thickness to the dry thickness.

The calculated swelling factors of the PDMAAm networks are summarized in Tables

4.4 - 4.7. In general, the thickness of the films has increased as the humidity of the system is

raised indicating that the gel undergo swelling. On the same token, the refractive index is of

the swollen layer decreased due to the water uptake. The refractive indices of water and

PDMAAm networks are 1.33 and 1.5, respectively. Hence, the refractive index of the

swollen layer depends directly on the water uptake.


64

Figure 4.11: OWS reflectivity scans obtained from a surface-attached poly (DMAAm-co-1%

MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%). Symbols

represent the measured reflectivity values and solid lines are the results of model calculations

based on Fresnel formalism.

Table 4.4: Optical parameters of poly (DMAAm-co-1%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 649 2.2762 1.50 1

25% 708 2.2794 1.5097 1.09

54% 740 2.2621 1.5040 1.14

74% 818 2.2367 1.4955 1.26

92% 882 2.2055 1.4851 1.36

*All measurements performed at 25°C.


65

Figure 4.12: OWS reflectivity scans obtained from a surface-attached poly (DMAAm-co-

2.5% MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%).

Symbols represent the measured reflectivity values and solid lines are the results of model

calculations based on Fresnel formalism.

Table 4.5: Optical components of poly (DMAAm-co-2.5%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 716 2.287 1.5123 1

25% 766 2.2685 1.5061 1.07

54% 802 2.2515 1.5005 1.12

74% 888 2.1959 1.4818 1.24

92% 952 2.1878 1.4791 1.33



66

Figure 4.13: OWS reflectivity scans obtained from a surface-attached poly (DMAAm-co-5%

MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%). Symbols

represent the measured reflectivity values and solid lines are the results of model calculations

based on Fresnel formalism.

Table 4.6: Optical components of poly (DMAAm-co-5%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 818 2.3027 1.5175 1

25% 867 2.299 1.5162 1.06

54% 905 2.2945 1.5147 1.106

74% 949 2.2845 1.5114 1.16

92% 984 2.2644 1.5048 1.203



67

Figure 4.14: OWS reflectivity scans obtained from a surface-attached poly (DMAAm-co-

10% MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%).



Table 4.7: Optical components of poly (DMAAm-co-10%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 475 2.3153 1.5216 1

25% 499 2.3382 1.5291 1.05

54% 513 2.3274 1.5255 1.08

74% 532 2.2937 1.5145 1.12

92% 561 2.2859 1.5119 1.18


4.4.3. Influence of layer thickness in humidity swelling

68


The motivation here is to address the influence of film thickness in swelling behavior

of a surface-attached network.The poly (DMAAm-co-2.5% MABP) was chosen for the

investigation due to the reason that it is hydrophilic and show strong swelling in water. The

deposited film thicknesses of 49 nm (thinner film) and 716 nm (thicker film) were prepared

on gold substrates by appropriate coating procedure. Such a large difference in thickness is

sufficient to study the influence of film thickness.

The reflectivity measurements of thin (49 nm) and thick (700 nm) films at different

humidity are shown in Figures 4.15 and 4.16 along with the Fresnel calculations. In both

graphs, significant changes in the angle θ are observed as the relative humidity of the

surrounding is increased. This demonstrates that both films (thin and thick) absorb water

molecules from the air and it increases with the increase in the humidity.

The calculated optical parameters of the films are summarized in Tables 4.8 and 4.9.

For the thinner film, the thickness was raised from 49 nm to 65 nm, which yields to the

swelling factor of around 1.33. Similarly, for the thicker film, the thickness was increased

from 716 nm to 974 nm indicating a swelling factor of 1.36. The variation in the swelling of

these two films are depicted in Figure 4.17, where the change in the film thickness and

relative swelling was plotted against the relative humidity in order to elucidate the influence

of layer thickness in swelling.


69

Figure 4.15: SPR reflectivity scans obtained from 49 nm surface-attached poly (DMAAm-co-

2.5% MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%).



Table 4.8: Optical components of 49 nm poly (DMAAm-co-2.5%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 49 2.2862 1.5120 1

25% 53 2.281 1.5130 1.08

54% 55 2.2741 1.5080 1.12

74% 60 2.2702 1.5066 1.22

92% 65 2.2302 1.4933 1.33



70

Figure 4.16: OWS reflectivity scans obtained from ≈700 nm surface-attached poly (DMAAm-

co-2.5% MABP) network at different relative humidity (RH ≈ 5%, 25%, 54%,74% and 92%).

Symbols represent the measured reflectivity values and solid lines are the model calculations

based on Freznel formalism.

Table 4.9: Optical components of ≈700 nm poly (DMAAm-co-2.5%MABP) at different RH.

Relative

Humidity (%)

*Thickness

d (nm)

Permittivity

Constant ( ε´)

Refractive

Index (n)

Swelling factor

(α)

0-5% 716 2.287 1.5123 1

25% 766 2.2685 1.5061 1.07

54% 802 2.2515 1.5005 1.12

74% 888 2.1959 1.4818 1.24

92% 952 2.1878 1.4791 1.33



71

Figure 4.17: Swelling factor of thin and thick films as a function of relative humidity.

The similarity in the shape of thickness or swelling curves in Figure 4.14 strongly

indicate that the gels are very sensitive and absorb a lot of water molecules from the

moisture. Howeveer, the slight variation between thin and thick layer occurs, may be due to

the difference in the film process and subsequent crosslink density. The estimated difference

in the swelling of these two films was α ≈ 0.03-0.04 and it is within experimental error of the

SPR and OWS measurements. Such a small variation is less pronounced when compared to

the net swelling of the gels. Therefore, the influence of layer thickness is neglected in the

case of humidity swelling.

Similar results were obtained for surface-attached poly (MAAm-co-3%MABP)

network using the film thickness of approx. 150 nm and 550 nm (results are summarized in

appendix) and the results are in agreement with the above statement that layer thickness is

less pronounced in the case of swelling in humid air.

4.4.4. Flory-Huggins sorption behavior

72


The swelling behavior can be explained by considering the modified Flory-Huggins

theory, where the water-polymer interaction becomes more favorable as the films swells

than between water molecules. Such observation is precedent in the swelling of hydrophilic

polymer gels in water and humid air, which suggests that it is energetically favorable for the

polymer structure to reorganize in response to the penetrant.155 According to the model, the

relative swelling of the surface-attached gels is described as

ln(𝜌

𝜌°) = ln𝜑1 + 1 − 𝜑1 + 𝜒(1 − 𝜑1)² (Eq.4.19)

Where ρ° is the saturation vapor pressure, ρ is the actual vapor pressure, φ1 is the

volume fraction of the small molecule in the polymer network and χ is the Flory-Huggins

solvent interaction parameter, which describes the free energy of mixing between water

vapor and polymer network. For a given volume fraction of a polymer (φ2), the smaller the

value of χ, the greater the rate at which the free energy of mixing of the solution decreases

with the addition of solvent molecules.

The volume fraction of solvent (φ1 = 𝑙/𝑙0)) and polymer (φ2 = 1- 𝑙/𝑙0)) is obtained

from the relative swelling ( 𝑙/𝑙0) of the surface-attached network.

ln ρ

ρ° = ln

𝑙

𝑙0 −

𝑙

𝑙0 + χ

𝑙

𝑙0 ² (Eq.4.20)

From equation 4.20, the relative thickness is exponentially increasing with increasing

relative humidity (Figure 4.18).156-158 The Flory-Huggins parameter (χ) for these gels is closer

to one (χ = 0.8-0.9) at the highest relative humidity of 90%, which indicates that the

interaction between the water molecules is stronger than between polymer and water.

Furthermore, even at high saturation (≈99 % relative humidity), the water vapor pressure at

25°C is 23 mbar and, the total amount of water in the air is less than 2.5%. Regardless of few


73

percentage of water in humid air, it is still a bad solvent for the poly (alkylacrylamide)

networks studied here.

Figure 4.18: Swelling factor of surface-attached PDMAAm networks as a function of relative

humidity.

Table 4.10: Swelling of surface-attached PDMAAm networks with different crosslink densities

in humid air and in water.

Polymer Swelling in humid air (αh)

0% 25% 60% 74% 92% Water

Poly (DMAAm-co-1% MABP) 1.0 1.09 1.14 1.26 1.36 4.0

Poly (DMAAm-co-2.5% MABP) 1.0 1.07 1.12 1.24 1.33 3.4



Generally, when polymer layers swell, one must in principle add an elastic

contribution to the above equation, which accounts for the anisotropic chain stretching of


74

the polymer network in the swollen state.21 The elastic energy can be obtained from the

equation 4.21

∆𝐺′

𝑉 = 𝑘𝑇

𝑑

2 (∝2− 1 − 𝑙𝑛 ∝) (Eq.4.21)

Where, ∆𝐺′ is the change in elastic energy, d is the dimention of the swelling (for

anisotropic swelling, d = 1), k is the gas constant, T is the temperature and α is the relative

swelling of the surface-attached network.

Table 4.11: Elastic free energy of the surface-attached PDMAAm networks normalized by

crosslink density (v).

Humidity (∆G'/v) x kT (Joules)

1% MABP 2.5% MABP 5% MABP 10% MABP

0% 0 0 0 0

25% 0.05 0.04 0.03 0.02

60% 0.1 0.1 0.1 0.05

74% 0.2 0.2 0.1 0.1

92% 0.3 0.24 0.13 0.1

water 7 4.7 4.3 0.6

k Boltzmann constant (1.38x10-23) and T is temperature in (298K)

As noticed from Table 4.10, the maximum swelling at highest humidity of 90% is only

to the factor of 1.36. Consequently, the change in elastic energy due to chain

stretching is around the ∆𝐺′ ≈ 0.3𝑘𝑇 per polymer chain of the surface-attached network

(Table 4.11). This contribution to Gibbs free energy of mixing is very small, when compared

to the contribution from chemical potential of the solvent molecules as shown in equation

4.27. Therefore, the elastic contribution is less significant or negligible in the case of

humidity swelling. However, in systems where the degree of swelling is significantly higher

4.4.5. Selective adsorption of water

75

like in the case of water, then the contribution from elastic energy will be utmost

important.150,152


As outlined previously, very similar results were reported for the swelling of

polyelectrolytes159 and surface grafted polystyrene160 in a mixtures of a solvent and a non-

solvent. Bunjes et al.161 have shown that toluene (good solvent) preferentially adsorbs in

polystyrene brushes over methanol for the reason that it has a better affinity to polystyrene

than the methanol (poor solvent). In addition, different compositions of toluene/methanol

mixtures were studied and under all circumstances, the volume fraction of toluene within

the swollen layer was considerably larger than in the solvent mixture. Hence, we have

plotted the volume fraction of water (φ1) in the network as a function of water in the air at

different relative humidity (Figure 4.19). As outlined previously, at maximum humidity of

≈100% the air contains ≈2.5% of water.152

From a principle point of view, when the surface is covered with water at a given

temperature (25°C), all the water molecules tend to evaporate until the maximum relative

humidity is reached. Subsequently, at low relative humidity (< 50%) or low water content

(<1.35%) in humid air, a comparable strong uptake of water is observed. This indicates that

water vapor is selectively extracted from air.159-161 The driving force for such selective

adsorption of water vapor is the polar amide groups present in the acrylamide gels.


76

Figure 4.19: Volume fraction of water (φ1) in the surface-attached network plotted as a

function amount of water in air at desired relative humidty (RH).

4.4.6. Correlation between water swelling and humidity swelling

In addition to the swelling of PDMAAm networks in humid air, we also measured the

swelling in water to interrogate the correlation between these two distinct phenomena. The

reflectivity scans of surface-attached PDMAAm networks in water and in nitrogen

atmosphere are shown in Figures 4.20 - 4.23. The reflectivity scans are quite different for

polymer layers in the presence of a nitrogen atmosphere and in water. In most cases, one

waveguide is observed for the measurement in nitrogen. While, an additional waveguide or a

strong shift in the resonance angle of the waveguides is found for the measurements in

water. This illustrates that there is a significant change in the interfacial properties of the

deposited films and this can be revealed if the optical parameters are derived from these

measurements. The calculated film thicknesses and permittivity constants of the surface-

attached PDMAAm networks are shown in Table 4.12. Accordingly, the swelling factor of

4.4.6. Correlation between water swelling and humidty swelling

77

the network is determined using the equation 4.16 (see page 11), and the results are shown in

Figure 4.24 along with the swelling in humid air to elucidate the correlation.

Table.4.12: Calculated optical components of surface-attached PDMAAm networks.

Polymer N2 Water *SF

d (nm) ε´ d (nm) ε´

Poly (DMAAm-co-1%MABP) 175 2.2299 700 1.9008 4.0

Poly (DMAAm-co-2.5%MABP) 334 2.318 1135 1.994 3.4



*Swelling factor

Figure 4.20: Reflectivity scan obtained from a poly (DMAAm-co-1%MABP) layer. Symbols

represent the measurement in N2 atmosphere and in water. Solid lines are the Fresnel

calculations of the measured data.


78

Figure 4.21: Reflectivity scan obtained from a poly (DMAAm-co-2.5%MABP). Symbols



Figure 4.22: Reflectivity scan obtained from a poly (DMAAm-co-5%MABP). Symbols




79

Figure 4.23: Reflectivity scan obtained from a poly (DMAAm-co-10%MABP). Symbols




80

Figure 4.24: (A) Swelling of the surface-attached PDMAAm gels in humid air as a function of

swelling in water. (B) Swelling of poly alkylacrylamide networks at maximum relative

humidity of ≈ 92 % and in water. Poly (PAAm-co-3%MABP) was used as a reference

polymer. In both the swelling (water and humid air), the shape of the curve follows the same

trend i.e. as we increase the MABP content, the swelling decreases.

According to Figure 4.25A, the polymer that showed significant swelling in humidity

also swell strongly in water. For example, surface-attached poly (DMAAm-co-1% MABP)

had a swelling factor of 1.36 in humidity (RH ≈ 92%) and swell to the factor of 4.0. While,

polymers which had poor swelling in humidity, were found to have weaker swelling in

water e.g. poly (DMAAm-co-10%MABP) showed swelling of 1.18 and 1.66 in humid and

water, respectively.

The driving force for the swelling behavior in both the case is the strong interaction

of polar amide moieties in the polymer with water molecules i.e. polymer-water interaction

and the only difference is the phase of the water molecules, vapor or liquid, involved in the

process. Consequently, the extent of swelling is based on the counteracting forces associated

4.4.7. Swelling of poly (alkylacrylamide) coatings in humid air

81

with the given condition, namely the heat of condensation for humidity and elastic

contribution. The heat of evaporation for water is ≈40 kJmol-1, which is quite larger than the

energy caused by the stretching of a polymer chain in solvent. Thus, the swelling in

humidity is weaker than in water. For example, swelling of poly (DMAAm-co-1%MABP) in

water, αw = 4.0 (≈4 times) and humid air (≈92% of relative humidity), αh = 1.36 (≈1.4 times).

However, for a qualitative understanding of the relationship between humidity and water

swelling requires thermodynamic parameter e.g. heat of hydration, which is measured using

solution calorimetry and discussed in the next section.


Similar to the previous investigation, we have explored the humidity swelling of

surface-attached poly (alkylacrylamide-co-MABP) networks. From a principle point of view,

these homologous series should imbibe water molecules from the moisture in a similar

fashion to that of PDMAAm gels. In addition, this investigation, allows us to interrogate the

influence of molecular structure or hydrophobicity of the gels upon contact with moisture.

The results are summarized in Table 4.13162 and the optical components of the layers at

different humidity can be found in the appendix of the thesis.

The polymers with shorter alkyl substituent such as MAAm and EAAm show

significant amount of swelling in presence of moisture. In contrast, polymers consisting

larger alkyl substituent such as PAAm, BAAm and DEAAm show weak or poor swelling

under such conditions. Furthermore, MAAm was found to extract the maximum amount of

water into the network (approx 30%) and BAAm and DEAAm were found retain less than

10% of water.

These gels also exhibit selective absorption of the water vapor upon contact with

moisture. The extent of this adsorption depends on the length of the alkyl chain at N-----------------------------------

Humidity swelling of poly (alkylacrylamide) coatings was performed by Mr.Fan Wu, CPI-IMTEK, University of

Freiburg, Freiburg, Germany.


82

substitution. Further details on swelling of poly (alkylacrylamides) in humid air can be found

in the thesis of Fan Wu.162

Table.4.13: Humidity swelling of surface-attached poly (alkylacrylamide) networks.162

Polymer Swellling at relative humidity (αh)

0% 25% 60% 74% 92%

Poly (MAAm-co-1% MABP) 1.0 1.07 1.17 1.26 1.42

Poly (EAAm-co-3% MABP) 1.0 1.07 1.13 1.15 1.23

Poly (PAAm-co-3% MABP) 1.0 1.04 1.08 1.10 1.12

Poly (BAAm-co-3% MABP) 1.0 1.02 1.05 1.06 1.08

Poly (DEAAm-co-3% MABP) 1.0 1.02 1.03 1.04 1.06

4.5. Heat of hydration

In order to calculate the heat released during the process of dissolution of precursor

polymers that are used in the study is investigated using solution calorimeter (Thermometric

precision solution calorimeter, LKB 8700, Sweden). In a typical run, 10-20 mg of polymer

powder are placed in a sealed glass ampoule (2222-150 for solid samples) and a reaction is

initiated by breaking the ampoule in the beaker containing 100 ml of de-ionized water, then

changes in the temperature of mixture are recorded.

The system was calibrated through electrically (i.e. by adding known amount of

energy (≈10J) to duplicate the effect of the thermal energy associated with chemical or

physical process) before and after the reaction in accordance with the standard protocol

described elsewhere.163 In addition, the heat exchange with the surroundings during the

process of the reaction and the heat released from stirring were adjusted mathematically

using respective obtained baseline temperatures before and after the reaction and finally the

heat of dissolution was calculated. The obtained calorimetric results are depicted in Figure

4.5. Heat of hydration

83

4.25, where, enthalpy of the solution is plotted as a function of mole percentage of MABP in

the precursor polymer.

Figure 4.25: Enthalpy of hydration (ΔH) as a function of MABP (%) in the Poly (DMAAm-

MABP) copolymers.

As noticed, the heat evolved from the system during the reaction is highly

exothermic for the polymers carrying low amounts of the crosslinker MABP. For example,

poly (DMAAm-co-1% MABP), showed ΔH ≈ -9 kJ/mol (excellent solubility in water) and for

poly (DMAAm-co-10%MABP), ΔH ≈ -1 kJ/mol (almost no solubility in water). This

demonstrates that, the solubility of the polymer decrease as the crosslink density of the

system increases because the polymers become more hydrophobic. The solubility of the

system is based on the assumption that „like dissolves like‟ principle. Hence, poly (DMAAm-

co-10%MABP) shows poor solubility in water due to a large energy difference in the

interaction between segments of the polymer and solvent molecules. In poly (DMAAm-co-

1%MABP) the segment-solvent interaction energy is closer to that of the interaction energies

between the segment-segment and solvent-solvent molecules, it is thus found to have a large

4.6. Conclusion

84

heat of hydration. To sum up the discussion, simply the hydrophobic polymers do not

dissolve appreciable in water.

The results of the heat of hydrations are in excellent agreement with the swelling of

polymers. The gel that showed strong hydration exhibits strong swelling in water and

consequently in humid air. Therefore, we conclude that the gels, which swell to the factor of

1.36 in humid air (RH ≈ 92%), can swell 4.0 times in water.

4.6. Conclusion

The swelling of the surface-attached poly (alkylacrylamide) networks was studied

using optical waveguide spectroscopy. The results demonstrated that these gels swell

anisotropically perpendicular to the surface as described by Toomey et al. 122 The swelling

was greatly affected by the hydrophobicity of the polymer, i.e., PMAAm (αw ≈ 5.4) > PEAAm

(αw ≈ 4) > PDMAAm (αw ≈ 3.5) > PDEAAm (αw ≈ 1.7) > PPAAm (αw ≈ 1.3) >PBAAm (αw ≈

1.17). An increase of the carbon units at N-substitution increases the hydrophobicity of the

polymer and thereby decreases the swelling of the surface-attached network. When the

length of the alkyl substituent is larger than ethyl (CH3-CH2-) group then the attractive

inter-segment interactions or hydrophobic association between alkyl groups become

dominant, the polymer becomes insoluble in water, and thus exibits little or no swelling in

aqeuous medium. Evidently, before crosslinking hydrophobic polymers such as PDEAAm,

PPAAm and PBAAm were also insoluble in water.

The thermo-responsibility of the surface-attached poly (alkylacrylamides) was

investigated by measuring the swelling at different temperatures varying from 15°C to 50°C.

Poly (DEAAm-co-3%MABP), was found to be the only thermo-responsive polymer in this

series of poly (alkylacrylamides). It showed a transition in the swelling behavior of between

25°C and 37°C and the transition temperature of the polymer was found to be ≈ 30°C. At this

4.6. Conclusion

85

temperature, the interaction of polymer-polymer dominates over polymer-water

interactions. Consequently, the chains of the polymer network tend to collapse and therefore

the water molecules are expelled from the network that results for the poor swelling of the

network.

Humidity swelling of the surface-attached poly (alkylacrylamide) networks was

studied using optical waveguide spectroscopy. Saturated salt solutions were employed to

control the moisture in the air.152 The temperature of all measurements were maintained as

25°C and the relative humidity was measured using a humidity sensor near the measurement.

The results suggest that the polymers have the tendency to swell in humid air upto the factor

of 1.4. More precisely, they can retain up to 30% of water in the networks. The crosslink

density and the hydrophobicity of the networks are the two dominating factors that affect

the swelling behavior in humid air. A weakly crosslinked poly (DMAAm-co-1%MABP)

showed significant swelling in humid air to the factor of 1.36 (RH≈92%) and strongly

crosslinked poly (DMAAm-co-10%MABP) swells 1.18 times (RH≈92%). Similarly,

hydrophilic poly (MAAm-co-3% MABP) swells upto 1.4 times and hydrophobic poly

(BAAm-co-3%MABP) barely swells (1.08, RH≈92%).

More significantly, the water volume fraction of the swollen network and volume

fraction of water in the moisture reveal that hydrophilic and weakly crosslinked polymers

selectively adsorb water vapor from the air at all humidy levels. The driving force behind

such phenomena are the strong interactions of the polar amide groups present in the

polymers that tend to extract water vapor from the air and binds rather strongly.

The correlation of water and humidity swelling suggests that the swelling in humidity

is much smaller than in water. Usually, hydrophilic polymer coatings swell four or five times

stronger in water than in humid air. In contrast, hydrophobic polymers show unappreciable

4.6. Conclusion

86

swelling in both cases. The extent of swelling in humid air depends on the enthalpy of

vaporization from the film and the volume fraction of water in the air. The enthalpic

components contain enthalpy of hydration (obtained from calorimetric measurements), the

enthalpy of segment-segment and water-water interactions and the entropy contains the

entropy of mixing and the entropy of stretching (elastic energy). Thus, the numerical

correlation between swelling in water and humid air is not straightforward. However, the

general trend is that, if the polymer swell appreciably in humid air (e.g. Poly (DMAAm-co-

1%MABP)) then they swell strongly in water and vice-versa. However, the quantitative

prediction is still a problem and requires more thermodynamic parameters.

5.1. Theory and mechanism of protein adsorption

87

5. Adsorption of proteins on poly (alkylacrylamide) surfaces

The biocompatibility of synthetic materials that are utilized in medical applications is

determined by the adsorption of blood plasma proteins.163 As outlined in chapter 1, it is a

widely held view that the adsorbed protein catalyses, mediates, or moderates subsequent

biochemical reactions that ultimately control biocompatibility.164 Hence, it is apparent that a

quantitative understanding of how proteins arrive at and adsorb to biomaterial surfaces is

essential for the design of biomaterials for advanced medical devices.165

As a consequence, the aim of the investigation described in the following was to gain

a fundamental understanding of protein adsorption on biomaterial surfaces with the goal to

develop blood compatible coatings. This chapter first provides a summary of the theory and

molecular mechanism of protein adsorption on organic and inorganic surfaces, followed by

the description of competitive adsorption of proteins. Then, protein resistant surfaces are

discussed. Finally, the protein adsorption on coatings of a homologous series of poly

(alkylacrylamides) is presented.


Proteins are biological heteropolymers comprised of twenty naturally occurring L-α-

amino acids that are joined together by amide bonds.166 Most of the proteins contain both

ionizable hydrophilic and non-ionizable hydrophobic functional groups. These

functionalities are randomly distributed in aqueous medium, which renders amphiphilicity

to the proteins in solution.167 Thus, a protein has the unique ability to exhibit patches and

domains that have a good solubility in water as well as patches and domains that have a poor

solubility in water (Figure 5.1). Those hydrophobic patches of proteins are expelled from

water and thus amphiphilicity is ultimately responsible for the adsorption of proteins from

solution onto the surface of any material.167


88

The adsorption of proteins is typically driven by electrostatic or hydrophobic

interactions between the proteins and the material surfaces as well as by a local increase in

entropy due to the displacement of water molecules and counter ions from the material

surface.168 Further important parameters are the adsorbate and any physical transformation of

the proteins upon adsorption.169

Figure 5.1: Schematic representation of the fibrinogen molecule. The three chains, Aα, Bβ,

and γ in the fibrinogen molecule are shown in red, green, and blue, respectively. FpA and

FpB are small fibrinopeptides in the central E region, which are cleaved by thrombin to

initiate fibrin polymerization. The peripheral D regions are composed of the βC and γC

domains. Figure adopted from Ivan et al.170

Norde et al. described the adsorption proteins as a Langmuir process from the

investigation of the systems BSA and lysozyme on silica and hematite.171 They related the

adsorption kinetics to the adsorption affinities between biomolecule and surface.

Interestingly, previously adsorbed BSA led a higher affinity than native BSA, which indicates

that the adsorption caused a physical transformation within the protein. No such effect was

reported for the lysozymes, which was attributed to the smaller size of these proteins.172

Schaaf el al. adopted the random sequential adsorption (RSA) model for protein

adsorption by accounting the irreversible process.172 The advantage of this theory is that it

considers the jamming limit of surfaces, beyond which no additional molecules can be

accumulated in the surface. In addition, this approach perceives the weakness of the

5.2. Competitive adsorption of ptoteins (Vroman effect)

89

Langmuir model, i.e. only one molecule can occupy one site of the adsorbate.

However, this is not appropriate for large protein molecules and this model failed to consider

the diffusion of protein molecules to the surface. Also, the work from Rebe et al. 173 showed

the adsorption of BSA diffuses over the acrylic polymers at the rate of 10-8cm2s-1. More details

on this topic can be found in recent reviews.174-176

Similarly, the distribution of protein molecules over the surface of the materials plays

a significant role in the adsorption kinetics of proteins.175 The adsorption of several systems

show an initial lag phase followed by auto acceleration and finally a phase of logarithmically

decreasing growth rate. This behavior is described as a fractal kinetics due to the non-

random distribution of the components in the surface region. The initial nucleation and

acceleration phases produce surface clusters of proteins on the surface, while in the later

phase leads to re-randomization.177

5.2. Competitive adsorption of proteins (Vroman effect)

Nearly three decades ago, Leo Vroman observed the preferential adsorption of

fibrinogen to silica and tantalum from blood plasma, ultimately called the „Vroman effect‟.178-

179 Figure 5.2 briefly describes the Vroman effect. A low molecular weight (LMW) protein

initially coveres the polymer surface, which is later displaced by a high molecular weight

(HMW) protein. This effect occur as it is more thermodynamically stable in nature when an

HMW protein replaces LMW proteins.180 However, the reverse sequence does not

occur, when HMW proteins covers the surface first, LMW proteins arriving later do not

displace the former (Figure 5.2).

Occasionally, the protein molecule at a given site of the surface (polymer surface)

may or may not be displaceable by a later arriving molecule of another protein, which

depends on how long time elapsed between the two events. The removability of the protein

5.2. Competitive adsorption of ptoteins (Vroman effect)

90

from the surface relies upon residence time since the adsorbed protein molecules undergo

molecular relaxation or a spreading process on the surface. A good example for such an effect

is how fibrinogen (Mol.wt ≈ 340kD) displaces the blood plasma protein kininogen (Mol.wt ≈

66-120kD) as described by Horbett et al. 181

Figure 5.2: A schematic diagram illustrating the Vroman effect. The low molecular weight

proteins adsorb first, which are then replaced by the high molecular weight proteins.182

All abundant proteins even of low affinity are expected to adsorb initially and will

later be replaced by proteins of high affinity. However, when adsorbed proteins reconfigure

into irreversibly adsorbed states, the situation is complicated as displacement is then more

difficult (Figure 5.2). The time of plasma dilution at which adsorption occur will be altered

by relaxation effects, and in some cases the relaxation is very rapid, and the protein becomes

irreversibly adsorbed before concentrations of competing proteins reach effective levels. In

addition, these adsorbed proteins will also change their conformation with a possible change

of function or of immunological properties.182 This early protein adsorption onto the surface

will inevitably have a strong impact on further biological processes.

5.3. Protein resistant surfaces

In 1965, Wichlle et al. first presented PHEMA hydro gels as protein repellent

materials.183 Since then many other materials have been discovered and a variety of models


91

have been proposed to explain this behavior. Some of these will be reviewed in the following

sections.

One of the classical models developed for the protein repellent surfaces is known as

mobility model.102 This model assumes that the protein requires certain time to adsorb onto

any given surfaces. According to this model, the surface-attached network and swollen

polymer chains posses high mobility and consequently the penetration of the protein

molecules on these surfaces are not possible i.e. it cannot penetrate inside the network.117

One can also name this as „wagging off‟ of the protein molecules cuased by the high chain

segment dynamics of the polymer networks.123

Perstin et al.184 studied the protein adsorption on oligoethyleneglycol self-assembled

monolayers (OEG-SAM) and proposed a model called „barrier model‟.184,123 This model is

based on the interaction of OEG-SAM‟s with water molecules, which depends on the

molecular confirmation of the SAM‟s and their ability to bind interfacial water molecules.

They reported that surfaces with high packing density are hydrophobic and the penetration

of proteins is facilitated through the hydrophobic interactions. When the surfaces are weakly

packed then the interface is mimic more water-like and hence the protein retains its natural

confirmation and moves freely from surfaces to medium withour adsorbing onto the surface.

Interestingly, Zhang et al.185 have studied the SAM‟s with terminal groups and found that the

large protein molecules can easily replace the loosely bound water molecules. While, tighly

bound water molecules are irreversible attached and consequently the interaction of protein-

polymer is restricted, therefore, the protein molecules are repelled from the surface.185,123

Jeon et al. and others demonstrated that poly (ethyleneoxide), (PEO) surfaces exhibit

protein repellency in solution.186 This behavior has been related to properties of these

molecules such as hydrophilicity, flexibility and the volume of exclusion exhibited by PEO

chains.187 These properties lead to a strong „steric repulsion‟ between PEO and proteins,


92

which dominate over the weak van der Waals interactions. However, the chain length of

PEOs, and spacing on the surface are crucial parameters that determine the proteins.188-189

Ivanchenko et al. 190 have shown that PEOs with shorter chain length are more prone to

resist the proteins (chains length from 4 – 24). However, the question of controlling the

chain length of PEO is somewhat obscured due to the fact that they are not directly grafted

onto the surface but are side chains of the acrylate polymers.191

Wittemann et al. 192 developed a model for weak polyelectrolyte brushes considering

counter ion release from brushes. Accordingly, protein adsorption is facilitated at low ionic

strength of the buffer medium, while the penetration is hindered at high ionic strength.193

The chains of polyelectrolytes are stretched to the maximum in solution. At low ionic

strength, the osmotic pressure within the chains is relatively high, since most of the counter

ions are confined inside the brush layer. Hence, adsorption of protein tends to release

counter ions from brushes and thus lower the osmotic pressure within the system.117 At high

salt concentrations, ionic strength inside and outside of the polymer layer is identical and

this effect is vanished.130

Recently, a model was reported by Wörz et al. that is entirely based on

thermodynamic considerations.123 The interaction between the protein and polymer is

governed by the Gibbs free energy, which has enthalpic and entropic components. The

enthalpic term is stronger for hydrophobic or ionic interactions between surfaces and the

respective components of the proteins. Under this circumstance, the protein adsorption will

always take place. For neutral hydrogels no hydrophobic interactions are possible and the

enthalpic contribution is considerably less or zero (i.e., there is no interaction between

polymer and protein). Therefore, the interaction of entropy is governing the adsorption

process.130, 194 The details of this model along with size exclusion effect (the large protein

molecules are able to penetrate inside the small mesh of the network) will be investigated in

the following sections using homologous series of surface-attached poly (alkylacrylamides).

5.3. Protein adsorption on surface-attached poly (alkylacrylamide) networks

93

5.4. Protein adsorption on surface-attached poly (alkylacrylamides) networks

Surface Plasmon Resonance spectroscopy (SPR) was used to study the protein

adsorption behavior of surface-attached poly (alkylacrylamide) networks. The experimental

details of SPR and sample preparation are described in Chapter 10.2 and 10.3. Briefly, a thin

film of polymer was spin or dip coated onto a gold substrate, and the dry film thickness was

measured in a nitrogen atmosphere. A subsequent measurement in PBS buffer yields a

swollen layer thickness. The kinetics measurements of adsorption process were performed at

an angle just below the resonance angle determined for the swollen layer. All of these

experiments were performed at 37°C and a typical result is shown in Figure 5.3.

Figure 5.4 depicts the SPR kinetic measurement obtained from PMAAm and PBAAm

coatings. The measurement was started with PBS buffer until a stable base line was achieved,

and then the buffer was replaced with a fibrinogen solution. The hydrophobic polymer

PBAAm, which does not swell in PBS buffer and showed an immediate rise in the reflected

intensity upon introduction of the fibrinogen solution. The measurement was continued

until a plateau of the reflectivity was reached which took approximately 30-40 min. Then

the protein medium was again substituted with PBS buffer to wash away any non-adsorbed

proteins (about 15 min). After the washing step, the reflected intensity did not return to the

original level of the initial measurement in PBS medium (Figure 5.4, black line), indicating

that a thin layer of protein was adsorbed during exposure to the fibrinogen solution.

Similarly, the experiment was performed on hydrophilic polymer PMAAm. Interestingly, no

change in the reflectivity was noticed upon introduction of Fg solution (Figure 5.4, gray line)

indicating that no adsorption of Fg had taken place. The SPR reflectivity scans of surface-

attached PBAAm and PMAAm before and after the protein adsorption are presented in

Figure 5.5 and Figure 5.6, respectively.


94

Figure 5.3: SPR reflectivity spectra obtained from Poly (BAAm-co-3% MABP) in PBS buffer

and in a nitrogen atmosphere. The solids lines are the Fresnel calculations. Gray arrow

indicate the angle chosen for kinetic measurement (θk).

Figure 5.4: Kinetic investigation of the adsorption of Fg to surface-attached PBAAm (black

line) and PMAAm (gray line). The experiment was performed at 37°C with Fg concentration

of 1mg/ml.


95

Figure 5.5: SPR reflectivity scans of surface-attached PBAAm in nitrogen air (A and B) and

PBS buffer (C and D). Symbols represent the measurement before (black) and after (gray)

kinetics. The solid lines are the Fresnel calculations.


96

Figure 5.6: SPR reflectivity scans of surface-attached PMAAm in nitrogen air (A and B) and

PBS buffer (C and D). Symbols represent the measurement before (black) and after (gray)

kinetics and the solid lines are the Fresnel calculations.

The SPR reflectivity scans of the hydrophobic PBAAm implies a significant difference

in the resonance angle i.e. angle θm is shifted to ≈1.1° (Δθ ≈1.1°). Additionally, the calculated


97

film thickness of the PBAAm using Freznel formalism was found to be 55 and 59 nm before

and after the Fg kinetics, respectively. The difference in the film thicknesses can be

attributed to the rise in the reflectivity upon exposure to Fg solution during the kinetics

experiment (and the ad-layer of Fg is approx. 4 nm). The hydrophilic PMAAm showed no

change in the resonance angle of the reflectivity after the Fg kinetics experiment i.e. Δθ ≈ 0°.

This would suggest that no protein is adsorbed on the surface upon protein exposure. In

addition, the calculated layer thickness of PMAAm coatings before and after the experiment

is identical (29 nm) and no ad-layer of fibrinogen has formed (Figure 5.6).

Figure 5.7: Fibrinogen ad-layer thickness on surface-attached poly (alkyl acrylamide)

networks. The film thickness was measured at physiological temperature (37°C) in a nitrogen

atmosphere after protein adsorption.

The results of the fibrinogen adsorption experiments on all other poly (alkyl

acrylamides) are presented in Figure 5.7. Layers of the hydrophobic polymers PDEAAm,

PPAAm and PBAAm (Z= 3,4a, and 4b) show adsorbed fibrinogen layer of approx. 3-4 nm.

All, hydrophilic and water swellable polymer layers (Z = 1, 2a and 2b) such as PMAAm,


98

PEAAm, and PDMAAm were found to be protein repellent and thus the ad-layer thickness is

zero.

Similarly, the adsorptions of other plasma proteins with different sizes are examined

on surface-attached poly (alkylacrylamide) networks. The molecular weight and size of the

various plasma proteins are summarized in Table 5.1. Here, two sets of experiments were

performed. First, the adsorption of proteins was examined individually on surfaces where the

experiment was started with water or buffer medium. After 5-10 minutes, the water was

replaced by a protein solution and a plateau was reached after approx. 20-30 minutes. Again,

the protein solution was substituted using buffer or water in order to remove any non-

adsorbed proteins. Typical SPR kinetics obtained from PBAAm and PDMAAm are presented

in Figure 5.8.

Table 5.1: Size and molecular weight of the plasma proteins used in the experiment.195-201

In the second experiment, the adsorption of various proteins was investigated

simultaneously by introducing them sequentially (aka. cocktail experiment). Similar to the

above experiment, the experiment was started with water until a stable base line was

achieved, then the water was replaced with a different protein solution one after another and

waited for approx. 20-30 minute until a plateau was reached. The protein solution was then

Protein Mol.wt (kDa) Rg (nm)

Fibronectin195-196 530 15,3± 0,3

Fibrinogen197-198 340-380 16,9

Vitronectin199 65-70 3 ± 0,6

BSA200 137 2,5

vonWillebrand factor201 500-387 16 (globular)


99

substituted with water to wash away any non-adsorbed proteins (approx. 15 min). Figure 5.9

depicts the typical SPR kinetic measurement obtained from PBAAm and PDMAAm coatings.

Figure 5.8: SPR Kinetics of the adsorption of the plasma proteins Von willebrand factor

(vWF), Fibrinogen (Fg), Collagen, and Fibronectin (Fn) on surface attached PBAAm (A) and

PDMAAm (B) coatings. The experiment was performed at 37°C and with the protein

concentration of 1mg/ml in corresponding buffer/water medium.

5.4.1. Determination of mesh size and size exclusion effect

100

Figure 5.9: SPR Kinetics of the adsorption of the plasma proteins Von willebrand factor

(vWF), Fibrinogen (Fg), Collagen, and Fibronectin (Fn) on surface attached PBAAm and

PDMAAm coatings. The experiment was performed at 37°C with a protein concentration of

1mg/ml in corresponding buffer/water medium.

As can be seen in Figures 5.8 and 5.9, the hydrophilic PDMAAm does not show any

change in reflectivity upon exposure to different protein solutions. While, the hydrophobic

PBAAm exhibit a sharp rise in the reflectivity upon introduction of protein solution and did


101

not return to the original level of the initial reflectivity as in water. This indicates that it had

strong adsorption of protein during the course of the kinetics. The calculated ad-layer of

proteins is presented in Table 5.2.

Table 5.2: Ad-layer of various proteins on PBAAm and PDMAAm coatings.

Polymer Collagen VWF Fn Fg Cocktail†

PBAAm 3.17 nm 4.13 nm 3.2 nm 4.1 nm 6.5 nm

PDMAAm --* --* --* --* --*

†Result obtained from Figure 5.10. *Ad-layer ≈ 0 nm.

The results are in agreement with that of the Fg analysis in which hydrophobic

PBAAm had a protein ad-layer of 3-6 nm and no ad-layer of protein was formed in the case

of the hydrophilic PDMAAm coatings. Obviously, the protein repellent character of the

latter coatings can be directly related to their swelling. This result supports the above-

mentioned model of entropic shielding.

5.4.1. Determination of mesh size and the size exclusion effect

The mesh size of the network plays a crucial role in many bio-medical applications

especially in the field of drug delivery.202 Here, the transport mechanism of drugs through

any swollen polymeric network is mainly controlled by the space available between two

crosslinks known as pore or mesh size.203 A crucial parameter that is frequently employed in

describing the size of the pores is the correlation length, ξ, which is defined as the linear

distance between two adjacent cross-links, and can be calculated using the following

equation:

𝜉 = 𝛼𝑠(𝑟02)1/2 (Eq.5.1)


102

Here, (𝑟02)1/2 is the root-mean-square, unperturbed, end-to-end distance of the

polymer chains between two neighboring cross-links and αs is the elongation ratio and it can

be related to the swollen polymer volume fraction, φ2, as

𝛼𝑠 = 𝜑2 −1/3 (For isotropic swelling) (Eq.5.2)

𝛼𝑠 = 𝜑2 −1 (For anisotropic swelling) (Eq.5.3)

The unperturbed end-to-end distance of the polymer chain between two adjacent

crosslinks can be calculated by using Eq.5.4

(𝑟02)1/2 = 𝑙(𝐶𝑛𝑛)1/2 (Eq.5.4)

𝑛 = 2𝑀𝑐

𝑀𝑟 (Eq.5.5)

Here, 𝐶𝑛 is the Flory characteristic ratio of the polymer, 𝑙 is the length of the bond

along the polymeric backbone (𝑙 = 1.25Å for PAAm‟s) and 𝑛 is the number of links/chain (C-

C bonds) between two crosslinks and the pre factor 2 comes from the C-C bonding of the

repeating unit (for acryl and vinyl polymers it is 2).204-205 Finally, 𝑀𝑐 is the number average

molecular weight between two cross-links and 𝑀𝑟 is the molecular weight of the repeat

unit.206

𝑀𝐶 = 𝑁𝐶𝑀𝑟 (Eq.5.6)

By combining the equations from 5.1 to 5.6, the correlation distance between two

adjacent crosslinks for anisotropically swollen network can be determined by the following

equation 207

𝜉 = (𝜑2)−1 2𝐶𝑛𝑀𝑐

𝑀𝑟

1/2

𝑙 (Eq.5.7)

The calculated mesh size of the surface-attached poly (alkyl acrylamide) networks

using the above equation in a swollen state is shown in Table 5.3. It was found that


103

hydrophilic networks have larger pore (or mesh size) compared to strongly hydrophobic

networks (PPAAm, PBAAm, PDEAAm). The hydrophobic polymers are tightly bound with

the adjacent polymer chain or greatly crosslinked. Hence, the end-to-end distance between

two crosslinks is lower.

Table 5.3: Mesh size of the surface-attached network in swollen state

Polymer α φ2 Nc Mc (g/mol) 𝝃 (nm)

Poly (MAAm-co-3%MABP) 5.34 0.187 193.2 17,283 48,8

Poly (EAAm-co-3%MABP) 4.6 0.217 124.3 12,809 37,8

Poly (DMAAm-co-3% MABP) 3.35 0.298 53.7 5,533 19,8

aPoly (DEAAm-co-3%MABP) 1.7 0.588 12.0 1,464 4

Poly (PAAm-co-3%MABP) 1.3 0.769 9.5 1,033 --*

Poly (BAAm-co-3%MABP) 1.18 0.847 10.7 1,311 --*

*For non-swollen gels, the mesh size is not a useful parameter

This mesh size of the gels (Table 5.3) seems to imply that all proteins in this study,

even the larger ones, (Fg, VWF and Fn) with radii of gyration of around 16 nm or a diameter

of 32 nm might be able to fit through such meshes (Table 5.1). However, the proteins are still

repelled by the hydrophilic surfaces indicating that the strong anisotropic swelling of the

surface-attached gels should be considered. Therefore, the size exclusion might only be

important only for larger proteins. This was further supported by the investigation on

PDMAAm polymers with different amounts of the crosslinker MABP, which yielded a

considerably lower crosslink density (i.e. 0.5% vs. 3% crosslinker), and a mesh size larger

than the size of proteins, show no unspecific protein adsorption within the experimental

error of the SPR method.

5.4.2. Entropic shielding

104


To explain of the observed differences within the homologous series of the surface-

attached poly (alkylacrylamide) networks, one needs to view the Gibbs free energy equation

for the overall adsorption process.208

∆Gads = ∆Hads – T (∆Sads) (Eq.5.8)

When hydrophobic patches of the protein interact with a hydrophobic polymer

attached to the substrate surface, the Gibbs free energy of adsorption (∆Gads) is dominated by

the enthalpy of adsorption (∆Hads) due to strong interactions with each other. The Gibbs free

energy of adsorption ∆Gads becomes negative and thus the adsorption of proteins is

thermodynamically favorable.209

The situation, however, is quite different for surface-attached hydrogels such as

PMAAm, PEAAm and PDMAAm. In the swollen state, these polymers do not show

significant interaction between proteins and polymers, i.e. the enthalpy of adsorption is very

small or even approaching zero (∆Hads = 0). Hence, the contribution of the entropic

component (∆Sads) is the crucial parameter, which determines the overall Gibbs free energy

of adsorption (∆Gads). ∆Sads contains two important factors, the entropy of mixing and the

entropy of configuration. The entropy of mixing (∆Smix) represents the uniform distribution

of molecules present in the system and is relatively low for most protein adsorption. The

entropy of configuration (∆Sconf) represents the significant chain stretching/ contraction of

polymers.123

∆Sads = - ∆Smix + ∆Sconf (Eq.5.9)

The surface-attachment of the polymer networks renders these systems to essentially

two-dimensions. Swelling is only possible perpendicular to the surface.122 It has been shown


105

that this leads to a strong stretching of molecules away from the surface, until an equilibrium

between the segment/ segment interactions (the osmotic pressure) and the energy required

for the elastic deformation of polymer chains is achieved.129 The extent of the swelling (and

chain stretching) can be described by the Flory-Rehner theory, provided that the

dimensionality of the system is taken into account (see Chapter 4.2). If now protein

molecules would diffuse into the layer this would require an even stronger swelling of the

network chains. Such an additional stretching reduces the chains entropy. As this entropy

loss is not compensated by any appropriate enthalpic gain, the Gibbs free energy of the

process becomes positive and protein penetration (or diffusion of any other macromolecule)

into the layers becomes energetically unfavorable.129, 210-212 One might call such a prevention

of unspecific protein adsorption caused by the entropic situation in the polymer film as

“entropic shielding” (Figure 5.10).

Figure 5.10: Schematic diagram illustrating the model entropic shielding. The penetration of

protein into the swollen network leads to a strong stretching of the chains. These chains,

however, are already stretched due to the swelling. Further stretching is entropically

unfavorable and hence, proteins are repelled.117,123, 130

5.5. Protein adsorption on terpolymer layers

106

Figure 5.11: Bovine serum (adult), BSA ad-layer thickness on surface-attached poly (alkyl

acrylamide) networks (homologous series). The film thickness was measured at physiological

temperature (37°C) in nitrogen atmosphere after protein adsorption.

In addition to model studies using plasma proteins, the adsorption study of whole

Bovine serum (adult) on homologous series elucidates „how the surface behaves when in

contact with more complex protein cocktails‟.123 In agreement with the model studies

described above, surfaces coated with polymers that exhibit strong swelling in aqueous

medium did not show any protein adsorption during the course of SPR measurements both

at room and physiological temperatures and the thickness of adsorbed protein = 0 nm.

Whereas, the polymer that do not swell in aqueous medium showed quick protein

adsorption and the thickness of adsorbed protein layer d = app.1-4 nm.


The terpolymers were investigated to explore the transition from protein repellent to

protein attractive surfaces or vice-versa. Accordingly, the copolymers comprised of

hydrophilic (EAAm) and hydrophobic (BAAm) at different composition were synthesized


107

and studied using the fibrinogen (Fg) as a model protein. Figure 5.12 demonstrates the SPR

kinetics of Fg obtained from hydrophobic poly (25%EAAm-co-75%BAAm-co-3%MABP)

and hydrophilic poly (75%EAAm-co-25%BAAm-co-3%MABP) coatings. As can be seen, the

polymer with a high content of BAAm (75%), which does not swell in water showed a quick

rise in reflectivity upon introduction of fibrinogen solution; indicating that a quick

adsorption of protein has taken place. The polymer with a low amount of PBAAm (25%) or

high content of EAAm (75%) showed no change in the reflectivity upon exposure to

fibrinogen solution (Figure 5.11). The calculated swelling and the ad-layer of Fg on surface-

attached tri-copolymer layers are summarized in Table 5.4.

Figure 5.12: SPR Kinetic measurement obtained from surface-attached poly (25%EAAm-co-

75%BAAm-co-3%MABP), (Black line) and poly (75%EAAm-co-25%BAAm-co-3%MABP),

(Gray line). All experiments were performed at 37°C and with the fibrinogen concentration

of 1mg/ml.

The argument from the previous section can be applied to explain the observed

difference between the copolymer layers. The hydrophobic polymer (with 75% BAAm)

adsorbs protein through a strong hydrophobic-hydrophobic interaction between protein and

5.6. Protein repellency of PDMAAm gels

108

polymer. The Gibbs free energy of the process is dominated by the enthalpy of adsorption

and thus the adsorption is thermodynamically allowed.

Table 5.4: Correlation of swelling ratio and fibrinogen adsorption on trip copolymers

Polymer α(dwet/ddry) ΓFg (nm)

Poly (BAAm-co-3%MABP) 1.0 4 ± 0.2

Poly (25%EAAm-co-75%BAAm-co-3%MABP) 1.05 2.6 ± 0.2

Poly (50%EAAm-co-50%BAAm-co-3%MABP) 1.2 1 ± 0.5

Poly (75%EAAm-co-25%BAAm-co-3%MABP) 2.86 0 ± 0.2

Poly (EAAm-co-3%MABP) 4 0 ± 0.2

The polymer with 75% EAAm renders the system more hydrophilic, as a consequence

it exhibit a strong swelling in water and therefore, the enthalpic interaction with is almost

closer to zero. As a result, the Gibbs free energy of adsorption is governed by the entropic

term, where the penetration of Fg is thermodynamically forbidden by the combined effect of

entropic shielding and size exclusion (refer sections 5.4.1 and 5.4.2).


The transport mechanisms of biomacromolecules (typically proteins) through swollen

polymeric networks have been discussed in the previous sections. Topological barriers such

as the number of crosslinks (junctions), branches, and crystallography in a network will have

an influence on the values of the diffusion coefficient. For example, it was shown that the

drug diffusion coefficient through hydrogels decreases as the crosslink density increases and

as the equilibrium degree of swelling decreases.213

Therefore, the motivation of this particular section was to understand how the

crosslink density and mesh size of surface-attached gels can affect the protein diffusion in


109

solution. Accordingly, DMAAm polymers with different amounts of photo cross-linker

(MABP) poly (DMAA-co-1%MABP), poly (DMAA-co-2.5%MABP), poly (DMAA-co-

5%MABP), poly (DMAA-co-7.5%MABP) and poly (DMAA-co-10%MABP) were used for

the investigation. The protein adsorption of the layers was measured using the procedure

outlined previously in Chapter 5.4.

Table 5.5 describes the physical properties of the surface-attached PDMAAm gels

with different amount of MABP. The results demonstrate that, gels with 1% of MABP are

weakly crosslinked, meaning that the number of monomers between two crosslinked points

are quite high and results in larger mesh size of the gel. On the other hand, gels with 10%

MABP are strongly cross-linked and thus the mesh size of the gel is smaller.

Table 5.5: The physical properties obtained from surface-attached PDMAAm networks with

different amounts of crosslinker (MABP).

Polymer Swelling (α) Mesh size,𝝃 (nm) Nc †Fg (nm)

Poly (DMAA-co-1% MABP) 4.0 24 85.33 0

Poly (DMAA-co-2.5% MABP) 3.50 18 59.5 0

Poly (DMAA-co-5% MABP) 2.5 11.3 47.7 0

Poly (DMAA-co-10% MABP) 1.66 10 11.7 0

‡Poly (BAAm-co-3% MABP) 1.18 --* 10.9 4

†Fibrinogen used as a model protein with concentration of 1mg/ml in PBS buffer at 37°C and

the radius of gyration of fibrinogen is 16.9 nm. ‡reference polymer. Nc is the segment density

of the surface-attached hydrogels. *for non-swollen gels the mesh size is not a useful

parameter

It is noted that, the protein repellency of strongly swollen systems e.g. PDMAAm

with 1% of MABP was dominated by the entropic shielding. Whereas, polymers that are

strongly crosslinked, the mesh size becomes smaller and the size exclusion is dominating

cause for the protein repellency surface-attached layers.

5.7. Adsorption of lipids

110


The motivation of this section was to interrogate the lipids adsorption of surface-

attached gels, especially onto protein repellent hydrogels. Accordingly, two distinct polymers

with different swelling properties were selected from homologous series as a model system.

One being neutral hydrophilic, water swellable poly (DMAAm-co-3% MABP), and the

other, neutral hydrophobic, not-swellable poly (BAAm-co-3% MABP). In addition, a

negatively charged poly (MAAc-co-10%SSAz), PMAAc was also used as a control, which

showed increased affinity to lipids (reference experiment). The phospholipid was chosen as a

model system for the investigation. It is constituted with the derivatives of phosphatic acid in

which the phosphate group is esterified with a choline residue, L-α-Phosphatidylcholine.

The chemical structure of the phospolipis is depicted in Figure 5.13.

CH3 N+

CH3

CH3

CH2 CH2 O P

OH

O

O CH2

CH O

CH2 O C

C

O

O

R

R

Figure 5.13: General structure of L-α-Phosphatidylcholine.215

Briefly, lipids are large organic ions that contain a single-charged entity at one end of

a long aliphatic chain or an amphiphilic molecules.214 These molecules are distinct due to

their aliphatic moieties, which will not interact with water yet tend to form hydrophobic

interactions with one another. Due to entropic driving forces, the aliphatic portions exclude

water and form van der Waals bonds, while the charged portions form hydrogen bonds and

electrostatic interaction with the aqueous medium.215 The effect of these different forces on

the structure of the solute, as well as its interaction with the solvent, is purely concentration

dependent. Thus, the concentration at which the change from a solubilized monomeric form

----------------------------------

Lipid adsorption experiments were performed by Mrs. Lisa Turnhoff, CPI-IMTEK, University of Freiburg,

Freiburg, Germany.


111

to a micelle occurs at is called the critical micelle concentration (Figure 5.14). Further basics

of lipids can be found in standard textbooks.214-215

The surface plasmon resonance spectroscopy was employed to study the lipid

adsorption on surface-attached polymer networks. As mentioned earlier, a thin film of

approximately 30 nm was generated on gold substrates and the layer thickness was measured

in nitrogen atmosphere and PBS buffer. The difference in layer thickness of pre- and post-

injection of lipid solution will reveal ad-layer thickness of lipid on polymer surface.

Figure 5.14: Arrangement of large organic ions into micelles. The polar head groups form the

outer hydrophilic surface that protects the hydrophobic aliphatic chains from having to

interact with water.215

Figure 5.15: Ad-layer thickness of L-α-Phosphatidylcholine (black column) and fibrinogen

adsorption (Gray column) on surface-attached polymer networks. Where, DMAAm

represents the poly (DMAAm-co-3%MABP), BAAm represents the poly (BAAm-co-3%


112

MABP) and MAAc represents the poly (MAAc-co-10%SSAz).The concentration 1mg/ml of

PC and Fg in PBS buffer were used for all the measurements at 37°C.

Figure 5.15 depicts the ad-layer thickness of L-α-Phosphatidylcholine on surface-

attached networks. The results demonstrate that hydrophobic PBAAm and negatively

charged PMAAc show lipid adsorption and consequently the ad-layer thickness of lipid

measured to be 4.6 nm (PBAAm) and 1.6 nm (PMAAc), respectively. At neutral pH, PBAAm

does not swell in PBS buffer. As a result, lipids adsorb through a strong hydrophobic

interaction, which is only the enthalpy of adsorption and simply dominates the Gibbs free

energy of adsorption and thus the adsorption is thermodynamically favorable. The PMAAc is

a weak polyelectrolyte and the extent of swelling is based on the pH of the medium.

Subsequently, the adsorption of lipids are depends on the ionic strength of the surrounding

medium.

Recently, Mossmann130 and Wörz et al. 117 revealed that poly acrylic acid (PAAc) is

susceptible to salt concentration. For example, PAAc show fibrinogen adsorbs at 0.2 mol/l of

NaCl, while repelling at 0.05mol/l NaCl solution. This is because at low ionic strength the

diffusion of fibrinogen into the PAAc network tends to release counterions, which enhances

the entropy of the system. In addition, fibrinogen acts like a multivalent counterion for

PAAc chains and both of them carry same negative charges. Hence, the change in entropy is

outrivaled by electrostatic repulsion and thus the fibrinogen adsorption is acquiesced. At

high ionic strength, the release of counter ions does not increase entropy due the same ionic

strength inside and outside of the PAAc network. Therefore, adsorption of fibrinogen is

prohibited.117,123

In our case, phospolipid (+ve) and PMAAc (-ve) are oppositely charged, hence the one

possible factor that could trigger lipid adsorption is ionic interaction. However, the

adsorption process is susceptible to the concentration of lipid (1mg/ml in PBS buffer) and salt

concentration (158 mM of NaCl). At a given situation PMAAc adsorb lipid through


113

electrostatic interactions and subsequently the measured layer is approx. 1.6 nm. The

adsorption is comparatively lower than PBAAm, which may be due to the charge

distribution around PMAAc chains, and it is weakly charged at neutral pH than at high pH.

Hence, the system is less charged and thus few molecules can interact with the chains of the

PMAAc and results in the low ad-layer thickness.

Hydrophilic PDMAAm does not show any adsorption during the course of lipid

kinetics. PDMAAm is a neutral polymer that shows strong swelling in PBS buffer and

consequently the chains of the polymer are surrounded with water molecules. In this case,

the enthalpy of adsorption (hydrophobic interaction) approaches zero and electrostatic

attraction is imperceptible due to the fact that the polymer does not carry any charge at pH

of 7.4 (for PBS buffer). It is conceivable that at higher salt concentration, polymer chains

tend to collapse and thus the adsorption may be possible. One should also consider the

energy required to expel bound water molecules larger than the energy required for lipid

interaction with PDMAAm chains.

Surface-attached PMAAc pre-adsorbed with phospolipid showed minor increment in

the reflectivity upon introduction of fibrinogen solution. The increased layer thickness was

approx. 0.9 nm, indicating that a larger fibrinogen molecule initially replaces the smaller

lipids before forming a monolayer on top of the lipid. However, the overall thickness

including lipid is approx. 4 nm (dlipid = 3 nm and dFg = 0.9 nm).

The preliminary results obtained from lipid adsorption implicit small insights on the

influence of protein adsorption. Accordingly, PDMAAm (neutral and hydrophilic) gel is

lipid and protein repellent due to their strong swelling in aqueous medium. While, PBAAm

(neural and hydrophobic) does not swell in water, and attracts both lipid and proteins

through hydrophobic interaction. To explore further, one has to consider several parameters

5.8. Conclusion

114

such as structural changes of incoming or penetrating molecules, salt content and lipid

concentration. These can be found in the reviews of Tonge and Schulz et al. 216-217

5.8. Conclusion

Protein adsorption of a homologous series surface-attached poly (alkylacrylamide)

networks was investigated using surface plasmon resonance spectroscopy. Hydrophobic

polymers (PPAAm, PBAAm and PDEAAm) which do not swell in water showed a strong

affinity to proteins. This non-specific adsorption is due to the strong hydrophobic interaction

between protein and polymers. Hence, the enthalpy of adsorption (ΔHads<0 i.e. -ve)

extravagantly makes the Gibbs free energy of adsorption as negative (ΔHads <0) and thus

adsorption of proteins is thermodynamically favorable.

The hydrophilic polymers (PMAAm, PEAAm and PDMAAm) that show strong

swelling in water were found to repel proteins. The enthalpy of adsorption is almost zero i.e.

there is no interaction between protein and polymer (ΔHads = 0). In such circumstances, two

important parameters such as size of the proteins in comparison with that of the mesh size of

gels the size exclusion effect, and the change in the entropy during the deposition of proteins

determines the adsorption process called as entropic shielding. The size exclusion effect is

more pronounced for the repellency of larger proteins due to the fact that mesh size is

smaller than the radii of gyration of protein (Fg, Fn and VWF). While, the latter was found

to be strongly dependant on the segment density of the swollen network were the extent of

swelling that a network can undergo in aqueous medium, independent of the chemical

composition of the system. The chains of a strongly swollen network were already stretched

to the maximum. Penetration of larger protein molecules will tend to stretch it further as a

result it lower the entropy of conformation and this loss is not compensated for by entropy

gain due to the adsorption. Therefore, the protein is repelled if there is no enthalpic

interaction with proteins.

5.8. Conclusion

115

PDMAAm gels with various amount of photo crosslinker (MABP) denote that the

mesh size of the gel is comparable to the size of the protein. In this case, it can be assumed

that both entropic shielding and size exclusion are equally important to determine the

protein resistance of hydrophilic PDMAAm networks. For example, PDMAAm with a

maximum of 10% MABP, swell to the factor of ≈1.7 (mesh size ≈ 10 nm, Rg of Fg ≈ 16 nm) in

water strongly is an ample indication for the above-mentioned statement. This also holds for

the explanation of other PDMAAm gels, e.g. poly (DMAAm-co-1% MABP), poly (DMAAm-

co-2.5% MABP) and poly (DMAAm-co-5% MABP).

In addition to the homologoues series and PDMAAm gels, protein adsorption studies

of copolymers containing both hydrophobic (PBAAm) and hydrophilic (PEAAm) moieties at

different composition were investigated. The results revealed that gels with more than 50%

BAAm tend to adsorb protein non-specifically. The gels less than 50% BAAm or over 50%

EAAm resist protein effectively. As noted earlier, the transitions between 50 to 75% of either

of them indicated that anything above 50% of BAAm leads to phase separation due to

collapse of polymer chains and consequently it showed a weak to zero swelling in water.

Hence, the hydrophobic pockets of polymer can interact with hydrophobic patches of

proteins. Above 50% of EAAm/DMAAm does not phase separate. Instead, it demonstrated

strong swelling in water and subsequently repelled proteins due to entropic shielding and

size exclusion effect. It was interesting to note that, even a minor change in the

hydrophobicity of polymers displayed a significant effect on the penetration of proteins. One

has to consider this critical parameter while designing the surfaces for medical applications.

Finally, lipid investigation on surface-attached PDMAAm and PBAAm illustrated

that as long as the gel is swollen in water, it can repel lipids adequately, provided it is not

undergoing any ionization. Hydrophobic and charged polymers show substantial adsorption

of lipids through electrostatic and hydrophobic interactions. However, more studies need to

be performed to have a clear understanding of the influence of lipids on protein adsorption.

6.1. Molecular mechanism of platelet adhesion

116

6. Platelet adhesion

The implantation of blood-contacting devices can trigger several undesirable clinical

complications due to platelet activation, adhesion or aggregation on biomaterial surfaces. As

mentioned in chapter 5, when an artificial material is exposed to blood, a layer of plasma

proteins (fibrinogen, fibronectin, vitronectin and von Willebrand factor ) quickly adsorbs to

the surface of the biomaterial. These adsorbed proteins have the tendency to bind to platelets

mediated through cell membrane receptors (e.g., GPIIb/IIIa also known as integrin IIb/IIIa).

Subsequently, the activated platelets adhere and trigger clotting events.116 The resulting

platelets aggregated in the vicinity of the implant, or dislodge and form emboli. The results

are often leads to life-threatening complications.

Consequently, the development of non-thromogenic surfaces has been a major focus

for the last three decades.116 A better understanding of the interactions of blood with

synthetic materials that stimulate clotting events is essential in order to design and generate

non-thrombogenic surfaces. Hence, the motivation of this chapter is to examine the

interaction of blood platelets with the surfaces generated from the homologous series of poly

(alkylacrylamide) networks. Furthermore, the platelets absorption is studied as a function of

the shear rate to mimic the blood flow.

6.1. Molecular mechanism of platelet adhesion

Platelet adhesion and activation is a multistep process involving several platelet-

receptor ligand interactions.218 Depending on rheological conditions, platelet adhesion occurs

through different mechanisms. Especially at high shear rates, vWF and Fg are essential

factors to decelerate fast-flowing platelets by reversible binding with platelet glycoproteins.

Among the different platelet adhesion receptors, glycoprotein GPIb and GPIIb/IIIa have the

6.2. Static platelet adhesion on surface-attached poly (alkylacrlamide) netwrorks

117

highest density on platelets. GPIb binds to adsorbed or immobilized vWF on the surface,

while the active form of GPIIb/IIIa crosslinks with Fg.219 Binding of Fg or other

glycoproteins containing Arg-Gly-Asp (RGD) sequences to activated GPIIb/IIIa triggers

platelet aggregation.220 Activated platelets also tend to release intracellular granules

containing coagulation factors VII and XI, the adhesion molecules P-selectin and vWF,

serotonin and platelet factor IV. All of these further affect the activation of other platelets,

coagulations and inflammation processes (Figure 6.1).

Figure 6.1: The activation of platelets by artificial surfaces: contact of platelets with artificial

surfaces leads to platelet activation in terms of ligand expression (GPIIb/IIIa). Activated

platelets either adhere (via proteins like fibrinogen) or aggregate to the surfaces.218

The mechanism of material-induced platelet activation is often presumed to happen

through the generation of thrombin due to the activation of intrinsic coagulation cascade or

the release of adenosine diphosphate (ADP) from damaged red blood cells or platelets. In

addition, even in the presence of heparin, small levels of thrombin are generated and thus

the activation of platelets is possible. However, the inability of thrombin and kellikrein


118

inhibitors that reduce platelet activation, suggest that other agonists at least partially mediate

platelet activation (Figure 6.2).221 Gorbet and Sefton described the correlation between

compliment activation and thrombocytopenia during dialysis.222 This eventually leads to

definitive platelet arrest and subsequent thrombus formation (Figure 1.6, Chapter 1).

Figure 6.2: Activation of the coagulation system is initiated by biomaterial-protein

interaction. Activation of factor XII is the initial step. Reciprocal activation and auto-

activation lead to amplification of activated factor XII, which in turn initiates the intrinsic

coagulation pathway via activation of factor XI, and finally towards production of fibrin.218

6.2. Static platelet adhesion on surface-attached poly (alkylacrlamide) networks

Platelet adhesion testing involves three individual steps: 1) the preparation of platelet

rich plasma from blood, 2) incubation of PRP on surface-attached hydrogels and 3) fixing of

adhered cells for electron microscopy. These steps are described in the following sections and

more details can be found in the experimental part of this thesis.

----------------------------------

Platelet adhesion experiments were performed with the collaboration of Prof. Barbara Zieger, Department of

Pediatrics and Adolescent Medicine, University hospital Freiburg, Freiburg, Germany.


119

Extraction of platelet rich plasma (PRP) from blood

Fresh human blood was collected from a healthy medication free donor and ACD

anti-coagulant is added. Platelet rich plasma (PRP) was harvested by centrifuging the blood

at 140g for 20 min at 25°C. The supernatant was again centrifuged at 250g for 15min at 25°C

to harvest the platelet poor plasma (PPP). The concentration of platelets was adjusted to 200

x 103 platelets/ µl by suspending PRP in platelet poor plasma (PPP).

Incubation of PRP on surface-attached networks

The glass slides covered with the surface-attached poly (alkyl acryl amide) networks

sterilized using 70% ethanol and dried for 12 hours‟ in a clean laminar flow hood. Then the

slides were transferred to 12-well plates and again sterilized with UV light (λ=254 nm for 1

min). After sterilization of the samples, 1 ml of PRP suspension was added to each well, and

incubated for 2 hours at physiological condition (37°C and 5% CO2). After incubation, the

non-adhered platelets were removed by rinsing with physiological saline. The morphology of

the adhered platelets was visualized using scanning electron microscopy (SEM).

Fixing of platelets on surfaces

The samples were incubated and rinsed with PBS buffer and then immersed in fixing

buffer containing 2.5% formaldehyde in PBS buffer. They were again rinsed three times with

PBS buffer. The fixed samples were dehydrated by incubation in ascending ethanol then

immersed in hexamethyl disilazane (HMDS), and dried overnight at room temperature. The

specimens were then covered with a 10-20 nm thick gold layer, and the morphology of

adhered platelets was examined by scanning electron microscopy (SEM).223-224


120

The SEM images of the surface-attached gels (homologous series of poly (alkyl

acrylamides) and polyurethane (used as reference/control) obtained after platelet adhesion

tests are shown in Figure 7.3. The hydrophilic gels containing shorter alkyl substituents

(PMAAm, PEAAm, and PDMAAm) show stromgly reduced or no platelets adhesion. Platelet

adhesion is extremely sensitive to adsorbed Fg or even to a very low amount of adsorbed

fibrinogen (Fg), which is sufficient to endow the full-scale platelet adhesion irrespective of

flow conditions,225 in other words, in absence of proteins, platelets will not attach onto the

surface.

The largely reduced platelet adhesion on these gels is directly attributed to the

plasma-protein repellent character of the gels and, hence, their strong swelling in plasma.

Gels with longer alkyl substituent (PAAm, BAAm, and PU) show extensive platelet

adherence and aggregation along with extended pseudopodia. These hydrophobic surfaces

allow Fg and VWf to adhere onto the surface via non-specific interaction and consequently

the specific interaction with platelets is supported. Presumably, the adsorbed Fg

predominately mediates the platelet adhesion in the experiments because they were

performed under static conditions.226

The amount of adsorbed platelet is not linearly proportional to the amount of

adsorbed Fg. However, the ability of adsorbed Fg to mediate platelet adhesion is rather

variable and depends on the adsorption conditions.226 The difference in this molecular

potency is caused by the conformational and orientational changes of the adsorbed (plasma)

proteins that can affect the availability of the binding domains of the platelets.227 The

conformational and orientation change of adsorbed proteins play a crucial role in

determining the ability of platelets to activate and adhere, but these factors may not be

important if Fg adsorption is sufficiently low. Gel coatings with short alkyl substituents such

as PMAAm, PEAAm and PDMAAm exhibit ultralow adsorption of Fg and vWF due to the


121

entropic shielding. Thus, it shows reduced or no platelet adhesion and consequently can

obstruct the generation of thrombi upon exposure to blood.

Figure 6.3: SEM micrographs of the platelet adhesion on poly (alkylacrylamide) network

coated glass: attached to (A) Z =1 (PMAAm), (B) Z = 2a (PEMAAm), (C) Z = 2b (PDMAAm),

(D) Z = 3 (PPrAAm), (E) Z=4 (PBAAm), (F) PU. Pictures were taken after 120 min of

incubation in human PRP at 37°C; 5% CO2 90% relative humidity.

(A) (B)

(C) (D)

(E) (F)

6.3. Plateletadhesion on 2D structures surfaces

122

The gels that are strongly swollen in aqueous medium repel proteins and blood

platelet and show great promise in improving the hemocompatibility of blood-contacting

applications. However, more in vivo and in vitro studies are obligatory to validate this

conclusion.

6.3 Platelets adhesion on 2D structured surfaces

To substantiate the platelet repellency of hydrogels we have investigated micro

structured surfaces. A simple photochemical method was used to achieve the patterned

PDMAAm surfaces and the used configuration is illustrated in Figure 6.4.

Figure 6.4: Fabrication process used for the preparation of the surface-pattern using a

chromium mask. PDMAAm was directly coated on a polyurethane glass substrate. The

chromium mask was placed on top of the sample and the polymer was crosslinked at 254 nm

for 4 minutes. The desired structures on the surface were then developed through washing

procedure. The white colored areas in the mask are glass in which light can penetrate and

provides the crosslinking layer underneath. The black colored area in the mask represents

chromium, where the light cannot penetrate and thus the PDMAAm layer below chromium

was removed upon washing in a solvent.

6.3. Platelet adhesion on 2D patterned surfaces

123

A 2D pattern of the PDMAAm hydrogel was directly prepared onto a polyurethane

covered glass substrate in order to evaluate the precision of the patterning technique, and to

demonstrate the contrast in the degree of platelet adhesion between uncoated and hydrogel-

coated polyurethane. In addition, this method allows the use of other plastic substrates, like

for e.g., polystyrene, polyurethane, or poly vinyl chloride, in the fabrication and yet enable

reliable thickness measurements. The PDMAAm (10 mg/ml in ethanol) is spin casted on top

of the polyurethane (PU) coated glass substrate (or silica wafer).The chromium mask with

desired pattern was placed above the sample and irradiated for 4 min at 254 nm. After cross-

linking, the sample was extracted and dried in ethanol and nitrogen air, respectively.

Further, the concentration precursor polymer and irradiation time can be used to govern the

thickness of the hydrogel layer.

Figure 6.5: SEM images of PDMAAm patterned surface after exposure to PRP. Extensive

platelet adherence was observed in the PU region and reduced or no platelet binding was

seen in the PDMAAm region.

The SEM images of the platelet adhesion on patterned surfaces are shown in Figure

6.5. The results indicate that, the platelet adhesion is drastically reduced in areas where the

surface is covered with the hydrogels. In contrast, the PU covered regions are full of adhered

platelets. This suggests that, the PDMAAm hydrogel is capable of showing strong inertness to

6.4. Ultrasound contrast angents to detect platelets at shear flow

124

non-specific adsorption of platelet adhering factors (Figure 6.1 and 6.2), as long as the

hydrogel is prepared with an adequate thickness. Hence, the PDMAAm gel, at least in the

period relevant to platelet adhesion events maintain a stable and well-defined surface

chemistry even after exposure to platelet rich blood plasma.

Furthermore, the patterned system thus allowed precise control on the mechanism of

the platelets adhesion, as well as, the lateral distribution of the adhering cells. The surface

chemistry described herein may be adopted as a platform for the development of a wide

range of cell microarrays with platelet assays as the primary application.

6.4. Ultrasound contrast agent to detect platelets at shear flow condition

Ultrasound contrast agents are micrometer-sized microbubbles consisting of a lipid

shell, and core contains a stabilizing gas, which can be injected into the blood stream. The

microbubble (MB) provide the contrast for ultrasound imaging and carry ligands from

outside to target specific locations.228 These agents enhance the contrast between different

types of tissues within the body and provide integrated information between molecular

events and the physiological events that takes place in the body.229

The uses of biomarkers linked to microbubbles (MB) that can specifically target and

elucidate molecular level interactions within the body provide a chance for earlier disease

detection. Targeted contrast agents are linked to chosen ligands, which selectively bind to

the molecules expressed on the plaque surface, like the activated GPIIb/GPIIIa receptor, the

scavenger receptor macrophages, or other molecules expressed by extracellular matrix,

thrombi, and neovasculature.230 Thus, molecular ultrasound imaging of aggregated or

adhering platelets may also offer a non-invasive diagnostic tool to identify thrombi within

unstable atherosclerotic plaques especially in-situ under flow condition.

----------------------------------

Microbubble targetting experiments were performed with the collaboration of Dr. Felix Günther, Universitäts-

Herzzentrum Freiburg, Freiburg, Germany.


125

The aim is to target the activated platelets using microbubbles (MB). The MB‟s used as

a contrast agent, which utilizes the P-selectin ligand - sialys lewisa (sLea) as the targeting

molecule. Platelet activation comprises series of casecade reactions such as spearding of cells,

release of α-granules and the formation of microparticles.231 Upon activation, the platelets

undergo shape change and become more spherical and form pseudopods on their surfaces.

This activated platelet has specific interaction with P-selectin ligand of the MB.

MBs are composed of a decafluorobutane gas core, which is encapsulated by a

phosphatidylcholine lipid shell. A brush of polyethylene glycol is attached onto the lipid

shell through a lipid anchor embedded in the lipid monolayer. Some of the polyethylene

glycol (PEG) molecules carry biotin at the distal tip.231 Targeting can be achieved by coupling

biotinylated sLea ligand (Sialyl Lewis-polyacrylamide-biotin, Glycotech, Gaithesburg, MD) to

the biotinylated MB via a streptavidin linker. This Biotin-PAA-sialyl Lewis has specific

interaction with P-selectin of the activated platelets as shown in Figure 6.6.

A standard PS culture dish was etched using 10% sulfuric acid solution and

thoroughly washed with de-ionized water. The dry dish was then spin casted (2500 rpm for

60 S) with PDMAA hydrogel (10 mg/ml in ethanol), and subsequent crosslinking and

extraction leaves a stable surface-attached PDMAAm network on PS (Figure 6.7). Both

coated and uncoated PS dishes were incubated with human fibrinogen solution for 30

minutes at physiological condition, and washed thrice with PBS buffer. Meanwhile, the

platelets in human blood plasma are extracted by washing PRP with sepharose and eluting

platelets and adjusted to a concentration of about 250000 platelets/ ml in PBS buffer and

activated by adding 20µM of adenosine diphosphate (ADP). The platelet coverage was

achieved by adding 1 ml of ADP-activated platelets medium on 35 mm fibrinogen-covered

dishes by incubating 30 minutes at 37° C. After incubation, dishes were carefully rinsed by

PBS.


126

Figure 6.6: Schematics of sLea-targeted microbubbles in the flow chamber. MB had

biotinylated PEG molecules on their shell and were coupled to biotinylated sialyl Lewis

polyacrylamide or a biotinylated control carbohydrate-freepolymer. Streptavidin was used as

a linker.231

The parallel plate flow chamber was employed to characterize the MB adhesion

efficiency to substrates to which fibrinogen and activated platelets were adsorbed. A silicon

gasket attached to a flow deck by vacuum grease, defined a rectangular flow path with a

width of 2.5 mm, a height of 0.127 mm, and a length of 20mm (Gasket A, Glycotech). The

wall shear stress in the chamber can be calculated by using following equation.

𝜏 =6𝜇𝑄

𝑏𝑕2 (Eq.6.1)

Where, τ denotes the wall shear stress (N/m2), µ is fluid viscosity (0.001N/s2 for water

at 20 °C), Q is fluid flow rate (m3/s), b is chamber width (m), and h is chamber height (m).

The wall shear stress is calculated in dynes/cm2 (one dynes/cm2= 1µbar = 0.1 Pa).


127

Figure 6.7: Modification of PS dish with PDMAA hydrogel. PS was etched with 10% sulfuric

acid and PDMAAm was deposited via spin coating at the speed of 2500 rpm for 1 minute. 200

ul of recombinant human Fg (100 ug/ml in PBS) were added to the dish, a plastic cover slip

was placed on top of the solution without damaging the polymer layer, which will help to

achieve uniform coverage of Fg. Then the dish was incubated overnight at 4° C, and washed

thoroughly with PBS buffer. After blocking non-specific binding sites by incubation of the

dishes with casein for 2 hours, dishes were rinsed again with PBS. Now, the platelet

concentrate (250,000 platelets /ml in PBS) was added to the dish, and incubated for 30

minutes. After incubation the sample was carefully washed with PBS buffer to remove any

non-adhered platelets.

A platelet-covered dish was carefully fastened in the flow chamber and the system

was vented completely to remove air bubbles. The center of the flow chamber was

monitored by a microscope attached to a CCD-camera and a PC. A withdrawal syringe pump

was employed and to achieve a wall shear stress of 40, 30, 20, 10 and 5 dynes/cm2,

respectively. The required refill rate of the syringe pump was calculated by using the above

equation (Eq.6.1). After a short delay, which was necessary for the stabilization of the MB

flux, a 60-second video was recorded for offline-analysis within a field view of 420 x 350 µm.


128

Figure 6.8: Illustration of the flow experiment using Microbubbles to visualize the platelets.

Micrographs of PS and PDMAAm coated dishes during platelet exposure are shown in

Figure 6.9. The PS dishes show a dense coverage with platelets. Hydrophobic polystyrene can

accommodate the Fg through a enthalpic interaction (also known as non-specific interaction,

as mentioned in the previous chapter protein adsorption). Consequently, the platelets adhere

to Fg covered surfaces which is specifically mediated by the glycoprotein‟s (GPIIb/GPIIIa).

Now, this scenario offers the feasibility for SLea MB to target the activated platelets,

specifically via P-selectin onto the activated platelet surface and thus, the extensive coverage

of SLea-targeted microbubbles were observed in the PS dishes (Figure 6.9C).

In addition, SLea MB have the tendency to bind to platelets even at high shear flow.

Apparently, after initial attachment of MB, they slowly roll on the surface of platelets along

with the flow direction (right to left, video), caused by continual detachment and immediate

reattachment of SLe to p-selectin. However, the capture efficiency is depending on the flow

rate and consequently decreases as the shear stress increased.229 The maximum capture

efficiency is achieved for 5 dynes /cm2, as 16.11% and the minimum capture efficiency

observed at 40 dynes/cm2, as 3.4%.231


129

Figure 6.9: Micrographs of MB binding obtained at 10 dynes/cm2. (A) PS dish covered with

platelets (white spots), yellow arrows inidcate the attached platelets. (B & D) PDMAAm

coated dish exhibits no platelet adhesion and the yellow arrow indicate the fringes in the

surface. (C) PS dish is covered with platelets that are expressing P-selectin on their surface;

black arrow indicates adhered SLea MB (black spots), which are rolling along the flow path

and immediately reattached. (E&F) PDMAAm patterned PS dishes, where the PS region is

covered with platelets and PDMAAm region show no platelet adherence. The field of view

all six images is about 420 x 350 mm.

6.4. Ultrasound contrast agent to detect platlets at shear flow condition

130

In contrast to the PS dishes, PDMAAm coated dishes show reduced or no platelet

adhesion (Figure 6.7B and 6.9D). This result can be attributed to their strong swelling in

aqueous medium that resists the adsorption of Fg through entropic shielding, and thus the

platelets (see Chapter 4. Protein repellency of PDMAAm gels). Hence, the surface observed

was plain without any MB attaching to the surface at any shear rate. Analogously, the

PDMAAm patterned PS dishes show, an extensive coverage of platelets (Figure 6.9E) and

binding of SLe MB (Figure 6.9F) in the region of PS. No platelet binding was observed in the

PDMAAm areas. These observations agree well with the previous results that PS is prone for

non-specific adsorption of proteins and subsequently platelets and that PDMAAm strongly

repels Fg and Platelets. More evidently, the different shear stress experiments on these

patterned surfaces also indicate that PDMAAm gels are very stable and render non-specific

adsorption at any given circumstance.

The specific adhesion of targeted SLe MB to P-selectin on activated platelets using

microscopy, along with a simple flow model was successfully demonstrated. To validate the

application of this system in targeted ultrasound imaging of activated platelets, more in vitro

and in vivo studies are need to be performed. Along this line, the modification of blood tubes

with PDMAAm hydrogel by implementing the above methods will be the first step to

visualize the thrombi inside the tube. Typically, the inner wall of the blood tube is modified

with PDMAAm hydrogels (see Chapter 8), followed by incubation in Fg solution and

platelets concentrate, and then will be subjected to SLe MB flow to capture the platelets.

From the principle point of view, platelet attached SLe MB will enhance the contrast for

ultrasound imaging and thus, it is possible to monitor the thrombi formation inside the tube,

later this would mean to be used in clinical applications. However, our preliminary results

(which are not shown) already confirm the bright future of this simple yet adequate and

cost-effective methodology to be commercialized.

6.5. Platelet adhesion under shear condition

131

6.5. Platelet adhesion under arterial flow condition

Flow is one crucial phenomenon in medical applications. It is well known to play a

vital role in haemostasis and thrombus formation by regulating the function of adhesive

proteins and blood cells.232 An increase in the shear rate increases the convective transport of

platelets and adhesive proteins to the vessel wall. Besides, flow also tends to remove and

dilute locally generated activators such as thrombin, and encourages the mixing with

inhibitors antithrombin III.232 It was already mentioned that Fg adsorption strongly promotes

platelet adhesion under static conditions. However, under flow condition, shear stress can

cause plasma proteins to change their conformation and exhibit a higher affinity to

platelets.233-234

By considering the above factors, the shear experiments were performed on

PDMAAm coated PS dishes using the flow chamber as described above (Figure 6.8). After

fixation of the modified dish in the flow chamber, the sample was perfused with Fg solution

(1 mg/ml in PBS buffer) for 60 minutes at 10dynes/cm2 (also 30 dynes/cm2), and then rinsed

with PBS for 2 minutes at the same shear rate. Similarly, the platelets solusion were also

perfused and washed using PBS buffer. The SLe MBs were permeated to visualize the

attached platelets through binding of SLe MB using microscope. The results demonstrate that

(appendix, video 11.1 and 11.2), spreading of platelets is observed only on the PS surface, and

it is utterly inhibited on the PDMAAm covered surfaces at all shear rates. This indicates that

the PDMAAm coating is stable at different (also arterial) shear rates. This provides a strong

evidence for the excellent anti-thromobogenicity of this coating also under flow condition.

6.6. Conclusion

In blood contacting devices, platelet adhesion is one of the important factors that

trigger the formation of a potentially life-threatening thrombus. Platelet adhesion is a

specific process, which caused by the adsorption of plasma proteins like Fg and vWF. These

6.6. Conclusion

132

proteins can bind to the cell membrane receptors such as GPIIb/ IIIa, which can lead to

platelet activation, adhesion and the formation of aggregates. Hence, it was an attempt of this

study to confirm this relation and to design a surface, which suppresses the platelet adhesion.

The stable surface-attached poly (alkylacrylamide) networks were prepared on glass

substrates, and the platelet adhesion was examined using fresh human platelet rich plasma

(PRP) under static and physiological condition. The gels that strongly swell in aqueous

medium (PMAAm, PEAAm and PDMAAm) were found to be platelet repellent. This was

due to the tendency of the gels to resist proteins that were present in the plasma. In contrast,

gels that do not swell in aqueous medium (PPAAm, PDEAAm and PBAAm) were platelet

attractive with extended adhesion and normal pseudopodia. These hydrophobic gels will

interact with plasma proteins and subsequently platelets will adhere on these surfaces

GPIIb/IIIa.

Furthermore, it was demonstrated that it is possible to target P-selectin of the

activated the platelets using SLea MB. The strong binding of SLe MB was observed on platelet

covered PS dishes at different shear rate (5 to 40 dynes/cm2). This methodology was used to

evaluate the platelet repellent of PDMAAm related hydrogels. The results were in excellent

agreement with that of the static experiments and no SLe MB binding was found on these

surfaces. The results from patterned dishes at different shear rate s, once again confirmed the

bio-inert character of PDMAAm related gels (PMAAm and PEAAm). Finally, from the above

investigation, PMAAm, PEAAm and PDMAAm were non-thrombogenic surfaces that might

be interesting for the fabrication of blood compatible devices. However, more in-vivo and

in-vitro studies are needed to prove long-term stability.

7.1. Mechanism of cell adhesion

133

7. Cell adhesion on surface-attached poly (alkylacrylamide) networks

The motivation of this chapter is to provide insights on cell compatibility of surface-

attached polymer networks.236 Normal human dermal fibroblast (NHDF), human umbilical

vein endothelial cells (HUVEC), and mouse fibroblasts (L929) are employed for the

investigation. Finally, the cell adhesion results are compared to that of swelling, and protein

adsorption surface-attached gels.


Cells are extremely sensitive to their surroundings.235-237 Typically, between 10 and

100 µm in diameter, cells responds to the environment at all length scales from the macro to

the molecular levels. The outer membrane of a cell is covered by specific carbohydrate

structures, and at least six different receptor systems that can be activated by interactions

with adjacent cells, ligands in the surrounding extra cellular matrix (ECM), and segregated

signaling molecules.238 Until now, hundreds of proteins are known to play a critical role in

stimulating the cell receptors, which in turn determines a plethora of responses, including

cell migration in the early embryo, coordinated organogenesis, and wound healing

throughout human life.239 Synchronically, these extrinsic factors make up a highly defined

and specialized cell microenvironment, which is necessary for correct tissue development

and continued functions.240


134

Figure 7.1: The process of integrin mediated cell adhesion comprises a cascade of four

different partly overlapping events:241 cell attachment, cell spreading, organization of the

actin cytoskeleton, and formation of focal adhesion sites. Firstly, in the initial attachment

step, the cell contacts the surface and some ligand binding occurs that allows the cell to

withstand gentle shear forces. Secondly, the cell body begins to flatten and its plasma

membrane spreads over the substratum. Thirdly, this leads to actin organization into

microfilament bundles, referred to as stress fibers. In the fourth step the formation of focal

adhesion sites occurs, which link the ECM to molecules of the actin cytoskeleton. Focal

adhesion sites consist of clustered integrins and more than 50 other transmembrane,

membrane associated and other cytosolic molecules. Among them e.g. tetraspanins, growth

factors receptors, syndecans, lipids, tensin, talin, vinculin, paxillin, and focal adhesion kinase

(FAK) were identified.242-245

Integrins are a diverse family of cell receptors that specifically binds to argininine-

glycine-aspartic acid (RGD) tri-peptides recognized in cell adhesive proteins such as laminin,

fibronectin, vitronectin and fibrinogen (Figure 6.1).246 These are heterodimers, which can

form in as many as 25 distinct pairings of their 18 α-subunits and 8 β-subunits, with each

7.1.1. Interaction with polymer surfaces

135

pair being sensitive to a unique set of ligands.247 For example, αvβ3 binds a wide range of

ECM molecules including fibronectin, fibrinogen, vonWillebrand factor, vitonectin, and

protelysed forms of collagen and laminin, while, integrin α5β1 specifically binds to

fibronectin.248 In addition to mediating the cellular adhesion to the ECM, integrins are

known to play a role as signal transducers, activating various intracellular signaling pathways

when activated upon ECM binding.248

7.1.1. Interaction with polymer surfaces

As outlined in chapter 5 (Protein adsorption), when a polymer is brought into contact

of biological medium, proteins (e.g. Fg, Fn, Vn and vWF) present in the medium tend to

cover the surface immediately.246 Subsequently, cell adhesion occurs through the mediation

of integrins.

One of the classical approaches that are used to enable the cell attachment on

polymeric surfaces is to deposit proteins such as fibronectin, collagen or laminin.249-251 This

approach, however, has some disadvantages in the medical perspective. At first, proteins

have to be isolated from other organisms and purified, which might lead to undesirable

immune responses and increases infection risks. In addition, proteins are susceptible to

proteolytic degradation and thus need to be refreshed occasionally.252 The conformational

change of proteins after adsorption on surfaces is another critical factor that influences the

cell adhesion.253-254 On hydrophobic surfaces, proteins tend to denature which often causes a

difference of the cell binding motifs.255 Hence, the wettability or water contact angles were

used to predict the cell adhesion. Most studies say that, polymers with moderate wettability

promote cell adhesion246 Lanza et al.247 demonstrated such wetting behavior of different

polymeric surface. The polymers with a contact angles in the range of 60 - 80° were found

have good cellular adhesion.247 However, the scattering in the plotted data indicates directly

that the water contact angle is not a suitable measure to classify the behaviors of the cells on

7.1.2. Biomimetic polymer surfaces

136

polymers. Other factors such as charges, surface topography and roughness that affect the

cell adhesion can be found elsewhere.254-255

7.1.2. Biomimetic polymer surfaces

A rather promising approach that is widely applied in stimulating cell adhesion is to

engineer surfaces with peptide sequences that are a part of the natural ECM. One common

example, which is often encountered in promoting the selective cellular adhesion, is to

employ RGD peptides.256 This is due of their wide spread distribution and use throughout the

organism, its ability to address more than one cell adhesion receptor, and its biological

impact on cell anchoring, behavior, and survival.257 Many attempts have been made to

develop RGD peptide based/ incorporated polymers.258-259

The covalent attachment of RGD sequence is achieved via functionalities like

hydroxyl-, amino-, or carboxyl groups.260 In the case of polymers, which do not contain such

functional groups, the RGD sequences are introduced by blending,261-263 copolymerization 264,

and chemical or physical treatment.265-266 However, it is well known that the RGD peptide

signal loses its affinity, and specificity when taken out of the context of a protein.267 For

example, RGD tripeptide exhibits virtually no effect in a cell detachment assay, but its

activity was retained by blocking the C-terminal carboxyl group or by adding flanking amino

acids according to the natural sequence: RGD (inactive) <RGD-NH2<RGDS<GRGDSP.268

Recently, Loschonsky269 and Petersen et al.268 have shown that a protein repellent

PDMAAm can promote cell adhesion upon modification with such RGD peptide sequences.

The adhesion of human fibroblasts in serum-free medium was investigated as a function of

the amount of peptide in the solution used for the film formation. It was noticed that, even

0.02 wt% of peptide in the film is enough to induce cell adhesion specifically to the regions

where the GRGDSP peptide is visible to macromolecules. Besides RGD, other important cell

7.2. Cell adhesion on surface-attched poly (alkylacrylamide) networks

137

adhesion motifs have been identified. Hence, the RGD sequence is not the universal cell

recognition motif, but it is nevertheless unique with respect to its broad distribution and

usage.266 It has been shown that factors other than the flanking amino acids are also

important such as the conformational presentation of the sequence.270 Other factors that

contribute to integrin-ligand binding affinity include the activation of integrins by divalent

cations and cytoplasmatic proteins. Further details about RGD and other stimulating factors

can be found in the literature of Jacob and Stevens et al. 271-272

7.2. Cell adhesion on surface-attached poly (alkylacrylamide) networks

Human umbilical vein endothelial cells (HUVEC), normal human dermal fibroblasts

(NHDF), and mouse fibroblasts (L929) are used in this study to evaluate the cell compatibility

of surface-attached gels. These are immortal cell lines, HUVEC is derived from normal

human umbilical vein, NHDF is derived from the dermis of juvenile foreskin or adult skin

from different locations like the face, the breasts, the abdomen, and the thighs and L929 is

derived from normal subcutaneous areolar and adipose tissue of a 100 days old male C3H/An

mouse. The complete description of sample preparation and other necessary details are

presented in the experimental section (see chapter 10.6).

Briefly, glass substrates (normal cover slips) were coated with surface-attached

polymer networks, sterilized by repeated immersion in 70% ethanol, and dried in a bio-safe

cabinet overnight. The samples were then transferred to 12-well plates and again sterilized

using UV light (λ = 254 nm) for 60 sec, followed by the addition of 1 ml of pre-warmed

medium to each well plate and incubation for 1 hour at 37°C and 5% CO2 atmosphere.

HUVEC, NHDF, and L929 cells (passage number of 5-7) were released from the culture

plates and washed using phosphate buffer. The cells were suspended in known amount of

basal medium, counted using a hemocytometer and diluted to the final concentration of

5x105 cells/ ml. 1ml of this cell suspension was added to each well plate that contained the


138

prepared substrates. The suspensions were mixed by gentle swirling and then incubated for 3

hour at 37°C and 5% CO2. After this time, the swimming cells were removed by changing the

medium after every 24 hours. Cell adherence was then investigated using an inverted

microscope 48 hours after initial seeding. The results of the cell adhesion experiment are

shown in Figures 7.2-7.4.

Figure 7.2: Micrographs of Normal Dermal Human Fibroblasts (NHDF) on surface-attached

poly (alkylacrylamide) networks. (A) PMAAm, (B) PEAAm, (C) PDMAAm, (D) PPAAm,

(E)PDEAAm and (F) PBAAm. All the micrographs were taken after four days of incubation.


139

Figure 7.3: Micrographs of Mouse Fibroblasts (L929) on surface-attached poly (alkyl

acrylamide) networks. (A) PMAAm, (B) PEAAm, (C)PDMAAm, (D) PPAAm, (E)PDEAAm

and (F) PBAAm. All the micrographs were taken after four days of incubation.


140

Figure 7.4: Micrographs of Human Umbilical Vein Endothelial cells (HUVEC) on surface-

attached poly (alkylacrylamide) networks. (A) PMAAm, (B) PEAAm, (C) PDMAAm, (D)

PPAAm, (E)PDEAAm and (F) PBAAm. All the micrographs were taken after four days of

incubation.

Figures 7.2, 7.3 and 7.4 show that the hydrophobic PDEAAm, PPAAm, PBAAm and

PU surfaces show an excellent cell adhesion with normal proliferation rates. The cell

adhesive proteins (e.g., fibronectin, vitronectin, laminin, and fibrinogen) present in the

serum adsorb onto the surface presumably through hydrophobic-hydrophobic interactions

and thus, form a monolayer on the surface of the polymer (network). The adsorbed proteins

7.3. Cell adhesion on patterned surfaces

141

enhance the interactions (via αβ integrins mediated cell adhesion) with adhesion sites on the

surface of the cells269, facilitating cell adhesion.

The highly hydrophilic hydrogel coatings based on PMAAm, PEAAm and PDMAAm

show no cell adhesion. This can be understood, as all these surfaces are strongly protein

repellent. It is interesting to note that the small difference between an ethyl- and a propyl

substituent is enough to generate a very strong change in the cell adhesion properties and

that the cell adhesion closely follows the swelling behavior of the systems.

Figure 7.5: NHDF cell growth of surface-attached poly (alkyl acrylamide) networks. The cells

only grow and proliferate on hydrophobic PBAAm, PPAAm and PDEAAm surfaces. The

hydrophilic PMAAm, PEAAm and PDMAAm show no cell adhesion even for prolonged

periods of times.

Figure 7.5 shows the cell adhesion profile of the attached cells over time on the

different polymer film surfaces. PMAAm, PEAAm, and PDMAAm are not supporting

cellular growth, as the count of cells remains almost unchanged between the first and the

seventh day of culture. Generally, anchorage-dependent cells such as L929, NHDF and

HUVEC need to adhere strongly to substrates in order to survive, spread, and transfer signals


142

into the cytosol to maintain cell homeostasis. Hence, the extent of cells spreading onto

substrates plays a crucial role in cell proliferation of the above-mentioned cells. Further, only

a few cells were firmly attached to the surfaces after 7 days of incubation, and these attached

cells showed poor spreading behavior. These results are consistent with previous findings by

Wörz that the swollen PDMAAm gel strongly repels the cells even for a period of up to 6

months. Similar findings have been observed for PMAAm and PEAAm gels.

In the case of hydrophobic PPAAm, PDEAAm and PBAAm coatings, the cells adopt a flat

and spreaded morphology, which results in the formation of a confluent cell layer. Cell

proliferation on these surfaces was significantly enhanced over time and can be positively

correlated to the hydrophobicity of the surface-attached gels. Accordingly, the most

hydrophobic PBAAm showed the maximum cell growth in the first few days and it keeps

increasing until the surface was completely covered with cells. PDEAAm and PPAAm also

showed a good proliferation rate time. All these gels support cellular adhesion and the results

are comparable to those obtained from standard PS culture plate.

7.3. Cell adhesion of patterned substrates

Cellular patterning is used in various aspects to control, modulate or understand the

behavior of cells on surfaces.273-275 In one example, an anisotropic solid micro-etching

procedure has been used for cellular patterning and for the analysis of microtubule

dynamics.275-276 The orientation of the division axis of cells has been found to be correlated to

the spatial distribution of adhesive micropatterns of a patterned extracellular matrix.277

Apoptosis of endothelial cells can be controlled by varying the size of adhesive islands micro-

patterned on gold surfaces and the differentiation of a stem cells can be modulated through a

variety of surface patterning techniques.278-280


143

A simple photochemical method wass employed to pattern the surface as described in

the section platelet adhesion. Accordingly, the cleaned glass Bp-Si coated cover slips were

first covered with hydrophobic PBAAm (or PPAAm or PDEAAm) and annealed at 70°C. The

sample was then crosslinked using UV-C-light at 254 nm for 4 min, followed by extraction

creating a stable surface-attached network. The above sample was then spin coated with

hydrophilic PDMAAm (PMAAm or PEAAm) and annealed at 70°C. The chromium mask

was placed on top of the polymer-covered sample and irradiated using UV-C-light at 254 nm

for 4 min, followed by extraction in a corresponding solvent that leaves a desired pattern on

the surface. The patterned samples were sterilized and seeded with cells as described above.

Micrographs were taken after seven days of incubation and the results are shown below.

Figure 7.6 shows that the protein attractive PBAAm areas on the sample promote cell

adhesion, while the protein repellent PEAAm region remains uncovered. The results are in

good agreement with the individual cell experiments presented in the previous section.

Evidently, the cells adhere and proliferate only in the region where the PBAAm is present,

while the PEAAm regions completely repel cells. In addition, the cell culture continued for

approximately four weeks, and it was found that the cells were not able to grow in the

PEAAm regions. In contrast, the cells were crowded in the PBAAm region, forming

aggregates, which then detached from the surface. Similar results were obtained for other

hydrogels, such as, PMAAm and PDMAAm on PBAAm. Overall, it can be concluded that

cell adhesion is completely controlled through the surface properties of the coatings.


144

Figure 7.6: Micrographs of NHDF (A) and L929 (B) cells on a patterned surface. The

hydrophobic and protein attractive PBAAm areas excellent cell adhesion. The hydrophilic

and protein repellent PEAAm regions repel cells. The experiments were continued for

approximately one month, and no change in the repellent behavior of PEAAm regions was

found.

7.4. Cell adhesion on copolymers

145

7.4. Cell adhesion of copolymers

Tri-copolymer composed of PEAAm and PBAAm with different ratios were

synthesized to explore the transition between cell attractive and repellent character

exhibited by the surface-attached poly (alkyl acrylamide) networks. In a way, these „zoom in‟

systems will also allow us to predict what is the minimum swelling required for a polymer

surfaces to repel proteins and the subsequent cellular adhesion. Those polymers with high

BAAm content (75 or 50%) are protein attractive and were also found to promote the cellular

adhesion with a usual proliferation rate. These gels are hydrophobic in nature due to the

presence of the BAAm and consequently exhibit poor to no swelling in aqueous medium (See

swelling and protein adsorption). Hence, all adhesion proteins (e.g. Fg, Fn and Vn) present in

the growth medium can bind non-specifically and mediate cell attachment.

Polymer layers with low content of BAAm (25%) or high content of EAAm (50% and

75%) show reduced cellular adhesion. This can be attributed to their hydrophilic character

due to the domination of EAAm content over BAAm. Hence, the polymer was found to swell

more in aqueous medium to the factor of 2.86 (≈ 3) and consequently repels proteins (see

section 5.6 and table 5.2). The cellular adhesion on polymeric surfaces is completely

mediated by the proteins. When there are no proteins on the surface, then the interaction of

cell receptors are strictly inhibited and thus the cell adhesion becomes impossible. Therefore,

the protein repellent polymers (75% EAAm ) showed no cellular adhesion. Similar results

were observed for tri-copolymers composed from DMAAm and DEAAm (Figure 7.7 D, E and

F).

7.4. Cell adhesion on copolymers

146

Figure 7.7: Micrographs of L929 (Mouse fibroblasts) cellular adhesion of Tri-copolymer

layers: (A) poly (25%BAAm-co-75%EAAm-co-3%MABP), (B) poly (50%BAAm-co-

50%EAAm-co-3%MABP), (C) poly (75%BAAm-co-25%EAAm-co-3%MABP), (D) poly

(25%DEAAm-co-75%DMAAm-co-3%MABP), (E) poly (50%DEAAm-co-50%DMAAm-co-

3%MABP), and (F) poly (75%DEAAm-co-25%DMAAm-co-3%MABP). Hydrophobic and

protein attractive polymers promote substantial cellular adhesion ((B, C and F). Hydrophilic

and protein repellent polymers inhibit the cellular adhesion (A, D and E). All the

micrographs were taken after four days of incubation.

7.5. Conclusion

147

7.5. Conclusion

The cell adhesion phenomenon of surface-attached poly (alkylacrylamides) was

investigated using three different cell lines, namely, human fibroblast (NHDF), human

endothelial (HUVEC) and mouse fibroblast (L929) cells. The results were in excellent

agreement with those of the swelling studies and the investigation of the protein adsorption

of the gels. The gels that are strongly swollen in aqueous medium repel the proteins and

consequently they resist cell adhesion. In contrast, gels that are hydrophobic and not

swellable in aqueous medium exhibit proteins adsorption and cellular adhesion mediated by

αβ integrins. Figure 6.8 illustrates the correlation between the swelling, protein adsorption

and cellular adhesion. Evidently, PMAAm, PEAAm and PDMAAm are cell repellent and

PPAAm, PDEAAm and PBAAm are cell attractive.

To substantiate and illustrate the results, the patterned samples were prepared using

protein repellent PEAAm and protein attractive PBAAm (Figure 6.8). Evidently, the cells

adhere and spread in the area covered by PBAAm and no cellular adhesion was observed in

the region covered with PEAAm. This is a strong indication that only protein attractive

surfaces will promote cellular adhesion. This was further supported by the tri-copolymers,

where the swellable gels repel cells and non-swellable gels attract cells.

It was noticed from the results of tri-copolymers that, even a slight increase in the

concentration of BAAm or EAAm in the systems changes the swelling phenomenon of the

polymer dramatically and consequently the protein and cell adhesion. Furthermore, these

results actually substantiate the transition between ethyl and propyl (transition point) in the

homologous series. Even an increase of one methyl group in the side chain has made the

system hydrophobic and accordingly the subsequent events of protein adsorption and cell

adhesion.

7.5. Conclusion

148

Figure 7.8: Schematic illustrating the correlation between swelling, protein adsorption and

cell adhesion. The maximum cell adhesion was taken as four and accordingly the graph was

plotted using the original protein and swelling data obtained at physiological temperature.

PMAAm, PEAAm and PDMAAm show strong swelling and no protein or cellular adhesion.

PDEAAm, PPAAm and PBAAm show poor swelling and strong protein and cellular

adhesion.

By combing the above observation, one can estimate the cell adhesion behavior of a

polymer. Charged polymers or polyelectrolytes, such as, polyacrylicacid (PAAc, negatively

charged), and polyethyleneimine (PEI, positively), promote cell adhesion due to their

tendency to adsorb protein through ionic interactions as described by Wörz et al.123 The

protein molecules can absorb through a hydrophobic interaction i.e. non-specific

interactions. For example, if the polymers are hydrophobic as in the case of PBAAm, PPAAm

and PDEAAm, this will exhibit poor swelling (weak interaction with water) in water. Hence,

the protein-polymer interaction is stronger and thus cells adhered via integrins. On the other

hand, if the polymer strongly swells in water, i.e. PMAAm, PEAAm and PDMAAm, there

are no polymer–protein interactions; therefore, proteins are repelled due to entropic

shielding and as well as the cells.

8.1. Attenuated Total Reflection spectroscopy (ATR)

149

8. Blood tube modification

The objective of this chapter is to develop an effective coating methodology that can

prevent the interaction between proteins and cells. Tubes that are currently being used in

the ventricular assist devices (VAD‟s) are made from polyvinylchloride (PVC). The product

(Medos® cannulae, size: 3/8 x 3/2.2 m) was purchased from Medos-Medizinintechnik AG

through the University hospital, Freiburg, Germany. The tube was coated with PDMAAm

through appropriate conditions and examined using surface-analytical tools, such as, contact

angle measurement and Attenuated Total reflection spectroscopy (ATR).

8.1. Attenated Total Reflection spectroscopy (ATR)

ATR was developed synchronously by Harrick280 and Fahrenfort281 It is a type of

internal reflection spectroscopy in which the sample is placed in contact with an internal

reflection element (IRE) of high refractive index (e.g. Zinc-Selenide or Germanium or

Titanium crystal). An Infrared beam is focused onto the edge of the IRE (in this case, the Zn-

Se crystal) and it is reflected through the IRE. The beam is then directed to a suitable

detector (Figure 9.1). At the sample/IRE interface, a complete internal reflection occurs and

that creates an evanescent wave. This evanescent wave penetrates a short distance into the

sample depending upon the wavelength of the incident light, where it can be absorbed.282

Consequently, an absorption spectrum of the sample in contact with the IRE or Zn-Se crystal

can thus be obtained. The spectrum is dependent on three important parameters such as the

angle of incident (θ), refractive index of the IRE (n1) and the sample (n2). The angle of

incidence or critical angle, and the penetrating depth (dp) can be obtained by using the

following equations 8.1 and 8.2, respectively.

𝜃 = 𝑠𝑖𝑛−1 𝑛2

𝑛1 (Eq.8.1)

𝑑𝑝 =𝜆

2𝜋𝑛1 𝑠𝑖𝑛2𝜃− 𝑛2 𝑛1 2 (Eq.8.2)

8.2. Modifications process

150

Figure 9.1: Schematic diagram of Attenuated Total Reflection (ATR) cell equipped with the

sample. I and R are the incident and the reflected beams, respectively.

The incoming infrared beam passes through an optically dense Zn-Se crystal and gets

reflected at the surface of the sample. When the propagating wave approaches through the

non-absorbing medium, i.e. the IRE or Zn-Se crystal, it forms a standing wave perpendicular

to the total reflecting surface. If the sample absorbs light (or energy), the propagating wave

interacts with the sample and becomes attenuated. This evanescent wave protrudes only a

few microns (0.5µ - 5µ) beyond the crystal surface and into the sample. Therefore, an

intimate contact between the sample and the surface of the crystal is utmost important for a

good spectrum.

8.2. Modification process

Polymer of interest

A thermal crosslinker styrene sulfonylazide (SSAz) developed by Kerstin Schuh was

used for the coating process. The copolymer of poly (DMAAm-co-5%SSAz) was synthesized

by following the procedures established by Schuh et al.283 The detailed mechanism of the

crosslinking process was revealed in the literature.284 Briefly, upon heating the sulfonyl azide

groups at 120°C, the nitrogen (N2) is removed, and it forms a nitrene. The resultant nitrene

8.2. Modification process

151

then reacts with the C-H-bond of a neighboring polymer chain or it abstracts the proton

from the C-H bond and recombines to give a cross-linked network (Figure 8.2).

Figure 8.2: (A) Chemical structure of the poly (DMAAm-co-5%SSAz) abbreviated as

PDAAm-SSAz. (B) Schematic diagram illustrates the thermal crosslinking of a polymer281.

Coating of tubes

The blood tube (Medos® cannulae) was cut into 10-15 cm lengths and annealed at

60°C for a minimum of 60 min. The tube (placed as shown in the setup, Figure 8.3) was

supported by hose clamps at an angle of 15° below the horizontal plane. The tube was rotated

with the help of a mechanical stirrer at the speed of 100rpm (rotation cycles per minute).

While rotating the tube, app. 20 ml of PDMAAm-SSAz solution (50 mg/ml in Isopropanol)

was slowly added through a glass pipette connected with tube using a hose clamp (Figure

10.1). The polymer solution moves through a tube by forming an „O‟ ring and thereby

generates a surface coating.

The polymer solution was added until a uniform wetting was observed in the tube

and it continues to be rotated for another 5 minutes. The tube was then crosslinked at 120°C

for 2 hrs. The tube was cooled down to room temperature and extracted thoroughly in

(A)

(B)

8.3. Characterization of PDMAAm-SSAz coatings

152

Isopropanol for two times. The tube was rinsed with water and dried using a flow of nitrogen

air.

Figure 8.3: Schematic diagram showing the instrumentation of tube coatings.


Attenuated Total Reflection Infrared spectroscopy (ATR-IR) employed as an

alternative tool for the characterization of the PDMAAm-SSAz coated tubes. The

background of the sample (only Zn-Se crystal) is scanned in the presence of nitrogen

atmosphere at an incident angle of approx. 45°. The modified tube is cut into a desired size as

per the requirements of the ATR sample holder (as shown in Figure 8.3). The sample was

then placed on top of the Zn-Se crystal with an intimate contact with each other and

scanned in nitrogen atmosphere to obtain the spectrum of the polymers. The resulted spectra

of the PVC containing DOP (Di-octylphthalate) and PDMAAm-SSAz coated PVC tubes are

shown in Figure 8.4 and the characteristic bands of the polymers are summarized in Table

8.1. ----------------------------------

Tube coatings were performed together with Mr. Joshua Marcum, California polytechnic State University, San

Luis Obispo, California, USA.


153

Figure 8.4: ATR spectra of PVC (A) and PDMAAm-SSAz coated PVC tube (B). DOP (Di-

octylphthalate) is a plasticizer present in the PVC tubing.


154

Table 8.1: ATR-FTIR frequencies of PVC and PDMAAm-SSAz.

Polymer Characteristic ATR-FTIR absorption frequencies

PVC

2964-2915 and 2853 cm-1 (s, CH and CH2 stretching), 1726 cm-1 (s, C=O

stretching of DOP), 1426, 1329 and 1253 cm-1 (w, CH2 and CH deformation),

1095 (w, C-C stretching), 964 cm-1 (w, CH2 deformation) and 686, 634 and

607 cm-1 (s, C-Cl stretching).

PDMAAm-

SSAz

2982 and 2873 (s, CH and CH2 stretching), 1641 cm-1 (s, –N-C=O stretching),

1528, 1496 and 1399 cm-1 (w, CH and CH2 bending), 1140- 1092 (C-O

stretching) and 750 cm-1 (w, CH bending).

w-weak, s-strong

As noticed in Figure 8.4 and Table 8.1, the results demonstrate that these two

polymers show their own characteristic bands based on the functional group present in the

moiety. PVC exhibits a strong carbonyl starching frequency at 1726 cm-1 (C=O streching of

DOP). Upon modification of PVC with PDMAAm-SSAz, the absorption intensity of the peak

(1726 cm-1) is lowered and a strong band is observed at 1645 cm-1 due to the

stretchingfrequency of carbonyl amide group of the PDMAAm-SSAz. This is ample evidence

that the PVC tube is completely covered with PDMAAm-SSAz.

In order to investigate the quality of the coatings, several samples at different

positions of the tubes were measured. In all those cases, the strong band at 1645 cm-1 was

obtained demonstrating that the tube was uniformly covered with PDMAAm-SSAz.

Furthermore, the stability of the coatings in water was examined by immersing a sample in

water for an extended period of time. The ATR-FTIR spectrum of the tube was recorded at

different time intervals and the results are shown in Figure 8.5.


155

Figure 8.5: ATR-FTIR spectra of PDMAAm-SSAz coated tube at different period of time.

The results depict that the characteristic band of PDMAAm-SSAz at 1625 cm-1 (-N-

C=O) remains constant over a period of one week. The intensity of the carbonyl stretching of

PVC at 1725 cm-1 also remains constant without any significant change in the increase in the

peak of the spectrum. Moreover, the spectrum of the PDMAAm-SSAz after soaking 8 days in

water exactly matches with that of the freshly prepared spectrum from the first day. All

these results demonstrate that PDMAAm-SSAz coating is very stable in the presence of

water.

8.4. Conclusion

PVC is one of the commercially employed materials for blood-contacting devices,

especially in VAD‟s, blood tubes and connecting joints. However, the non-thromogenicity of

the material is still an unanswered question. Hence, an attempt was made in this chapter to

modify the PVC tubes using blood compatible hydrogels (see chapter 7). The inner surface

8.4. Conclusion

156

(or wall) of the tubes were successfully coated with PDMAAm-SSAz according to the

experimental setup as shown in Figure 8.2.

The crosslinker, styrene sulfonyl azide (SSAz) allows a simple and efficient cross-

linking upon heating the tube at 120°C for 2 hrs.283 The coating was very much similar to that

of surface-attached network on flat substrate surfaces (e.g. silica or glass substrates).The

coatings were characterized using ATR-FTIR spectroscopy. The results illustrated that the

tube contains a uniform layer of PDMAAm-SSAz. In addition, the stability of the PDMAAm-

SSAz modified tubes was examined by immersing a piece of tube in water for a prolonged

period of time and the ATR spectra was recorded at different time intervals. It was found

that the PDMAAm-SSAz network was stable in water. However, the exact homogeneity of

the coatings cannot be proved due to the restriction of the tube curvatures. More

sophisticated methods should be utilized to evaluate the coatings.

The above study is ample evidence for a modification in the blood contacting devices

or more precisely tubes or „T‟ or „Y‟ joints. Furthermore, it was successfully demonstrated

using ATR-FTIR spectroscopy. Our next step along this line of research is to perfuse with

whole blood and evaluate the non-thrombogenicity of the alkylacrylamide hydrogels at high

shear-stress conditions.

9. Summary

157

9. Summary

The lack of hemocompatibilty at the surface of blood contacting devices is still

remains a major challenge in the medical device industry.16,102,285-287 Hemocompatibility is

primarily affected by blood-material interactions which results from a multi-step process

having strong interdependencies at the individual steps, which include protein adsorption

and platelet activation.12, 288 These interactions strongly affect the short-term and long-term

thrombotic response of the employed materials. When blood is exposed to any artificial

material, the first event is an instantaneous adsorption of proteins present in the plasma.1, 115,

227, 289-290 The resulting protein layer interacts strongly with the platelet receptors on the

surface of the contacting blood platelets, anchoring the platelets onto the surface, thereby

causing platelet activation. This process of primary and secondary homeostasis finally leads

to the formation of a thrombus.26 Hence, minimizing protein adsorption is an important goal

to avoid platelet activation and the subsequent thrombus formation in blood contacting

devices.

Numerous attempts were made in an ongoing effort to develop new blood-contacting

materials.1 So far, hundreds of polymers and swellable hydrogel surfaces have been

evaluated, by taking into consideration, various aspects of their interactions with plasma

proteins (Chapter 1). Hydrogels are a promising candidate and were studied extensively due

to their excellent properties that resist proteins and platelets. However, the overall

relationship between the polymeric structure and hemocompatibility is still lacking.

Therefore, it was aim of this thesis to provide insights on the perspectives of the polymeric

structure that resist proteins and platelet cells, followed by the hemocompatibility

assessment of the surfaces.

Previously, it was shown that the poly (dimethyl acrylamide) hydrogels could resist

proteins and cells under physiological conditions.123 To stretch the thread along these lines,

9. Summary

158

we have synthesized a series of poly (alkyl acrylamides) that were structurally very similar

but differ in their number of carbon atoms at N-substitution or length of the alkyl chain at

N-substitution. All these polymers contained 2-3 mole percent of the photocrosslinker

MABP, which was sufficient to obtain the crosslinked and surface-attached networks

simultaneously. The stability of these coatings was investigated by immersing a substrate

covered with the poly (alkyl acrylamide) networks in water for a prolonged period of time.

Then the sample was taken out from water at various time intervals and the film thickness

was measured using ellipsometry in nitrogen. There was no significant change observed in

the film thickness for six months, which indicates that these coatings remained stable in an

aqueous medium.

Optical waveguide spectroscopy (OWS) was used to determine the swelling of

surface-attached poly (alkylacrylamide) networks. The film thicknesses of a series of samples

were measured in nitrogen and in aqueous medium. The polymers with shorter alkyl

substituent such as PMAAm, PEAAm and PDMAAm showed a strong anisotropic swelling in

water (swelling factor ≈ 3-5), whereas, polymers with longer alkyl substituents like PPAAm,

PBAAm and DEAAm do not swell appreciably. An increase in the carbon units at N-

substitutions increases the hydrophobicity and the swelling of the surface-attached network

decreases, accordingly. In addition, the swelling of the networks was measured as a function

of temperature in the range of 15 to 55°C. All the polymers showed almost no significant

change in the swelling with respect to the temperature except the poly (diethyl acylamide).

This polymer showed a swelling and de-swelling transition between 25°C and 37°C

indicating a LCST around 32°C.

The swelling of the surface-attached poly (alkylacrylamide) networks in humid air

was also investigated. It was found that the swelling purely depended on the hydrophilicity,

crosslink density and the amount of moisture in the environment. Analogously, the

hydrophilic coatings (PMAAm, PEAAm and PDMAAm) swelled to the factor of 1.4 i.e. it

9. Summary

159

retained approx. 30% water in their networks, when compared with that of hydrophobic

coatings (PPAAm, PBAAm and DEAAm) which were not swellable. More importantly, the

estimated volume fractions of water in the swollen networks and the volume fractions of

water in the moist air revealed that the hydrophilic and weakly crosslinked networks adsorb

water vapor from the air. The driving force for this phenomenon was the strong interactions

of the polar amide groups that are capable enough to extract water vapor from the air and

binds rather strongly.

As mentioned previously, the adsorption of plasma proteins on biomaterials plays a

important role in triggering platelet adhesion, platelet activation and the subsequent

thrombus formation.227, 289-290 Hence, the adsorption of plasma proteins such as fibrinogen

(Fg), fibrinoectin (Fn), von Willebrand factor (VWF) and vitronectin (Vn) on surface-

attached poly (alkylacrylamide) coatings was investigated using SPR spectroscopy. SPR offers

the possibility to monitor the protein adsorption in-situ and the protein ad-layer thickness

can be measured with an accuracy of less than 1 nm.123 The SPR results revealed that the

hydrophilic and water swellable coatings made from polymers containing shorter alkyl

substituents (PMAAm, PEAAm and PDMAAm) were absolutely protein resistance. The

hydrophobic and water non-swellable coatings with longer alkyl substituent (PPAAm,

PBAAm and PDEAAm) showed a stronger protein adsorption. Such discrepancies in the

protein adsorption of these coatings can be explained by considering the Gibbs free energy

adsorption process and the mesh size of the surface-attached gels.

When hydrophobic patches of the protein interact with a hydrophobic polymer

attached to the substrate surface, the Gibbs free energy of adsorption (∆Gads) is dominated by

the enthalpy of adsorption (∆Hads) due to strong interactions with each other. As a result, the

Gibbs free energy of adsorption ∆Gads becomes negative and thus the adsorption of proteins is

thermodynamically favorable.

9. Summary

160

However, the situation is quite different for the hydrophilic and swollen surface-

attached networks (PMAAm, PEAAm and PDMAAm) where the enthalpy of adsorption is

almost equals to zero i.e. no interaction takes place between the polymer and the proteins.

Under these circumstances, the contribution from the entropy of adsorption (∆Sads)

dominates the overall Gibbs free energy adsorption process. The ∆Sads contains the entropy of

mixing (∆Smix) and the entropy of configuration (∆Sconf). The ∆Smix represents the uniform

distribution of molecules present in the system and is relatively low for most protein

adsorptions. The ∆Sconf represents the significant chain stretching or contraction of polymers.

The surface-attachment of the networks renders the systems to swell only in one

direction, i.e. perpendicular to the substrate surface. Due to this reason, the swelling is

anisotropic and at equilibrium; all chains of the swollen networks are stretched to the

maximum. The diffusion of protein molecules into this layer would require an even stronger

swelling of the network. Such an additional stretching of the layer would further reduce the

entropy of the chain. As this entropy loss is not compensated by any gain in the enthalpy due

to adsorption, the Gibbs free energy of the process becomes positive and the protein

adsorption into these layers becomes energetically unfavorable. This prevention of unspecific

protein adsorption driven by the entropic situation is termed as „entropic shielding‟.

Therefore, the strongly swellable coatings PMAAm, PEAAm, and PDMAAm repel proteins

effectively. However, the entropic shielding alone may not always be the major cause for the

observed protein repellency of these swollen networks. Larger protein molecules may simply

not fit into the meshes of the networks. These proteins may be exclused by size.

The mesh size of the surface-attached poly (alkylacrylamide) networks was calculated

in the swollen state. The pores of the hydrophilic networks (PMAAm, PEAAm, and

PDMAAm) were larger (ξ = 48, 38 and 20 nm) than the hydrophobic coatings (PPAAm,

PBAAm, PDEAAm) which had pores less than 4 nm. However, the mesh is not a useful

9. Summary

161

parameter for non-swollen networks since they have strong hydrophobic interactions with

proteins. The mesh size of the swollen gels imply that some of the proteins in this study, i.e.

the larger ones, (e.g. Fg, VWF and Fn) with radii of gyration of around 16 nm or a diameter

of 32 nm might also be exclused by size. The small proteins however, which are still being

repelled by the swollen layers must be kept out by mechanism such as entropic shielding.

Platelet adhesion is an important factor governing the hemocompatibility of the

materials that are applicable in the blood contacting devices. Platelet adhesion and

aggregation would lead to the formation of thrombus. If such an event occurs at surface of

the implant, it may cause even a severe life-threatening problem. Hence, to evaluate the

hemocompatibilty of the surface-attached poly (alkylacrylamide) networks, platelet adhesion

experiments were performed using platelet rich plasma (PRP). The protein attractive

alkylacrylamide coatings showed strong platelet adhesion with normal pseudopodia. The

protein repellent coatings showed reduced or no platelet adhesion. Usually, the platelet

adhesion is mediated through glycoproteins such as GbIIb/ GbIIIa. Especially, under static

conditions the adsorbed Fg enhances the platelet activation and adhesion. In the absence of

Fg or any other plasma proteins the platelet activation is not possible and that was the reason

the protein resistance surfaces, such as, PMAAm, PEAAm and PDMAAm did not allow

platelet adhesion.

However, most of the medical devices are in contact with blood flow. Hence, the

evaluation of hemocompatibility of the biomaterials under flow condition is important for

any blood contacting device application. Accordingly, the platelet adhesion experiments

were performed at different shear rates (5, 10, 20 and 40 dyn cm-2) on surface-attached

networks. Before the perfusion of platelets, the samples were pre-incubated with Fg or VWF.

These proteins were identified to enhance the activation and adhesion of platelets. Again, the

swollen coatings did not show platelet adhesion at any shear rate, while the hydrophobic

coatings had a strong platelet adhesion. These results were in good agreement with results

9. Summary

162

obtained at static conditions. Hence, the same explanation can be applicable here as well.

Nevertheless, these coatings were not damaged even at a very high shear rate of 40 dyn cm-2

is an added evidence for their brilliant stability under flow situations.

The cell compatibility of the surface-attached poly (alkylacrylamide) networks was

examined using three different cell lines, Human Umbilical Vein Endothelial Cells

(HUVEC), Normal Human Dermal Fibroblasts (NHDF) and Mouse Fibroblasts (L929). The

hydrophobic polymers such as PPAAm, PDEAAm and PBAAm showed an excellent cell

adhesion and proliferation with respect to time. The proteins present in the cell growth

medium quickly adsorbed onto the surface, which later mediated the cells through αβ

integrins. In contrast to that the hydrophilic coatings made from PMAAm, PEAAm and

PDMAAm did not show any cellular adhesion even after a prolonged period time in culture

medium. This is again a consequence of the swelling of the coatings in aqueous medium and

as a result, they repel proteins for entropic reasons. As there were no proteins on these

surfaces, also cellular adhesion is not possible. To explore further, the cell culture

experiments were performed on patterned glass substrates fabricated using hydrophilic and

hydrophobic alkyl acrylamides. Again, the cells were able to grow and proliferate only in the

region where the surface was covered with hydrophobic polymers, while it was repellent in

the area covered with hydrophilic polymers. Furthermore, the experiment on the PDMAAm

patterned sample was continued for a period of six months and it was found that no cellular

adhesion was observed in the area coated with PDMAAm polymers, which is an added

evidence for the strong repellent character of these coatings.

Finally, the results of this study suggest design rules for the identification of potential

candidates for blood contacting non-thrombogenic surfaces. A suitable layer will consist of

neutral, water swellable surface-attached polymer networks. Any changes to the layer, varies

the hydrophilic/hydrophobic balance, which in turn will make the polymer more

hydrophobic. This will also induce cell or platelet adhesion and render the surface blood

9. Summary

163

incompatible. In conclusion, the results demonstrate tha the interaction of a material with a

biological environment requires a precise control of the surface physical properties and a

careful choice of parameters used for evaluation.

10.Zusammenfassung

164

10. Zusammenfassung

Aufgrund der fehlendenden Hämokopatibilität der Oberflächen implantierter

Geräte, steht die Untersuchung blutkompatibler Materialien im Fokus der Forschung der

Medizintechnik.16,102,285-287 Die Hämokompatibilität wird hauptsächlich von den

Wechselwirkungen zwischen Blut und Materialoberfläche beeinflusst. Es handelt sich

hierbei um einen Prozess, der aus mehreren Einzelschritten besteht, z.B. Proteinadsoption

und Thrombozytenaktivierung.12, 288 Die Wechselwirkungen beeinflussen die Kurz- und

Langzeitreaktion der Thrombozyten auf das eingesetzte Material sehr stark. Werden die

Oberflächen künstlicher Materialen Blut ausgesetzt, kommt es zu einer plötzlichen

Adsorption von Plasmaproteinen.289-290 Die so entstandene Proteinschicht reagiert nun mit

Rezeptoren an der Oberfläche der Thrombozyten. Diese haften an der Materialoberfläche

und werden aktiviert. Dieser Prozess der primären und sekundären Homeostase führt

schließlich zur Bildung von Thromben.26 Die Minimierung der Proteinadsorption sollte

somit die Aktivierung der Thrombozyten und damit auch die Bildung von Thromben

verhindern.

Bisherige Studien zeigen, dass Polydimethylacrylamidhydrogele zell- und

proteinabweisende Eigenschaften unter physiologischen Bedingungen aufweisen.1

Ausgehend von diesen Erkenntnissen wurde ein Reihe unterschiedlicher

Polyalkylacrylamide synthetisiert. Diese weisen eine sehr grosse strukturelle Ähnlichkeit

auf, unterscheiden sich aber dennoch in der Anzahl bzw. der Länge der Kohlenstoffketten

am Amidstickstoff. Zur Erzeugung beständiger, oberflächengebundener Hydrogelschichten

wurden alle hergestellten Polymere mit 2-3 % des Photovernetzers MABP versehen.

Die Quellungseigenschaften der oberflächengebundenen Polymernetzwerke wurde

mittels Oberflächenplasmonenspektroskopie (SPR) untersucht. Dazu wurden die

entsprechenden Schichtdicken der Polymerfilme im trockenen und im gequollenen Zustand

10.Zusammenfassung

165

bestimmt. Aus diesen Werten wurde dann der entsprechende Quellgrad ermittelt. Polymere

mit kurzen N-Alkylresten, wie PMAAm, PEAAm und PDAAm, zeigen eine starke

anisotrope Quellung in Wasser (Quellgrad ≈ 3-5). Dagegen zeigen die Polymere mit längeren

N-Alkylresten, wie PPAAm, PBAAm und DEAAm, kein ausgeprägtes Quellverhalten. Eine

Verlängerung der Kohlenstoffketten am Amidstickstoff führt damit zur Vergrösserung der

Hydrophobizität und damit zu einer Verminderung der Quellungsfähigkeit. Dies resultiert

aus den verstärkten hydrophoben Wechselwirkungen der Alkylgruppen und der daraus

resultierenden schlechteren Wechselwirkung mit Wasser.

Weiterhin wurde die Adsorption von Plasmaproteinen (z.B. Fibrinogen (Fg),

Fibrinoectin (Fn), von Willebrand Faktor (VWF) und Vitronectin (Vn)) auf

oberflächengebundenen Polymerfilmen mittels SPR-Spektroskopie untersucht. Die

hydrophilen und in Wasser quellbaren Polymerschichten (PMAAm, PEAAm und

PDMAAm) zeigten eine vollständige Resistenz gegen Proeinadsorption, wohingegen die

hydrophoben und nicht in Wasser quellbaren Polymerschichten (PPAAm, PBAAm und

PDEAAm) eine ausgeprägte Proteinadsorption aufweisen. Diese ist auf starke

Wechselwirkung zwischen dem hydrophoben Polymer und hydrophoben

Proteinsegmenten zurückzuführen. Die proteinabweisenden Eigenschaften der hydrophilen

Polymere sind ein Resultat der starken Quellung. Durch die Quellung in Wasser sind die

Polymerseitenketten sehr stark gestreckt. Unter diesen Bedingungen würde das Eindringen

von Proteinen oder Proteinsegmenten eine weitere Streckung der Polymerketten erfordern.

Thermodynamisch betrachtet würde diese weitere Streckung einen sehr großen

Entropieverlust bedeuten. Der entsprechende Enthalpiegewinn ware dann vergleichsweise

klein, sodass eine Proteinadsorption nicht stattfinden kann. Diese entropiegetriebene

proteinabweisende Eigenschaft quellbarer Polymerschichten wird als entropische

Abschirmung (oder entropic shielding) bezeichnet. Neben dieser entropischen Abschirmung

10.Zusammenfassung

166

kann ebenfalls die Größe der eindringenden Moleküle, verglichen mit der Porengröße der

Netzwerke in Betracht gezogen werden.

Um die Hämokompatibilität der Polyakrylamidschichten zu untersuchen wurden

Thrombozytenadhäsionsexperimente mit thrombozytenreichem Plasma (PRP) durchgeführt.

Erwartungsgemäss zeigten die proteinattraktiven Polyacrylamidschichten eine starke

Thrombozytenadsorption, wohingegen die proteinabweisenden Polymerschichten eine

reduzierte bzw. keine Thrombozytenadsorption zeigten. Die eingesetzten Thrombozyten

wurde durch Zugabe von Glycoproteinen, wie GbIIb/ GbIIIa, aktiviert. Speziell unter

statischen Bedingungen unterstützt adsorbiertes Fg die Thrombozytenaktivierung bzw. –

adhäsion. Ohne jegliche Plasmaproteine ist eine Thrombozytenaktivierung unmöglich.

Damit erklären sich die proteinresistenten Eigenschaften der hydrophilen und stark

quellbaren Polymeroberflächen (PMAAm, PEAAm und PDMAAm).

In Anlehnung an die Proteinadsorptions- und Hämokompatibilitätsstudien wurde

ebenfalls die Zellkompatibilität der Polymerschichten untersucht. Dazu wurden 3

verschiedene Zellinien eingesetzt: humane Hautfibroblasten, Endothelzellen und

Mausfibroblasten. Die hydrophoben Polymerschichten, wie PPAAm, PDEAAm und PBAAm

zeigten eine hervorragende Zelladhäsion und -proliferation. Dagegen zeigten die

hydrophilen, stark quellbaren Oberflächen aus den Polymeren PMAAm, PEAAm und

PDMAAm, auch nach längeren Zeit im Zellkulturmedium keinerlei Zelladhäsion aufgrund

der entropischen Abschirmung. Zelladhäsion lässt sich damit als ein Abfolge von

Einzelschritten beschreiben, wobei die Ablagerung von Proteinen auf Oberflächen eine

tragende Rolle spielt. Können sich keine Proteine auf einer Oberfläche ablagern, wird das

Zellwachstum vollständig verhindert.

10.Zusammenfassung

167

Aus diesen Erkenntnissen lassen sich damit Richtlinien für die Entwicklung von

blutkompatiblen und thrombusfreien Oberflächen ableiten. Eine geeignete Oberfläche sollte

aus einem neutralen, wasserquellbaren und oberflächengebundenen Polymernetzwerk

bestehen. Veränderungen der Oberfläche führen zu einer Verschiebung des Gleichgewichts

zwischen hydrophilen und hydrophoben Wechselwirkungen der Oberfläche mit dem

Umgebungsmedium. Dies kann unter anderem zur Adsorption von Lipidmolekülen führen,

welche wiederrum für eine erhöhte Hydrophobizität de r Oberfläche sorgen. Dieser Vorgang

führt dann ebenfalls zu Zell- bzw. Thrombozytenadhäsion, resultierend in einer

Blutinkompatibilität. Die in dieser Arbeit hergestellten Polymerschichten stellen also einen

Ausgangspunkt für die Entwicklung medizinisch nutzbarer, blutkompatibler Oberflächen

dar.

11.1. Materials

168

11. Experimental section

11.1. Materials

11.1.1. Chemicals, solvents and reagents

Acetone (p.a) Roth

Acrylamide (>98%) Fluka

Allylbromide Fluka

4-aminobenzophenone Fluka

4,4‟-Azobis(isobutyronitrile) Fluka

4-Benzoylbenzoic acid Fluka

1-Butanol (p.a) Roth

Butylacrylamide ABCR GmbH&Co KG, Germany

Chloroform-d Roth

Chloroform (p.a) Roth

Calcium chloride Sigma-Aldrich

Dichloromethane (p.a) Roth

Diethylether (p.a) Roth

Diethyacrylamide ABCR GmbH&Co KG, Germany

Dimethyacrylamide Sigma-Aldrich

Dimethylchlorosilane Fluka

Dimethylformamide (p.a) Roth

Dimethylsulphoxide (p.a) Roth

1,4-Dioxane (p.a) Roth

Ethanol (abs or 100%) Sigma-Aldrich

Ethanol (p.a) Roth

Ethyl acetate (p.a) Roth

11.1. Materials

169

Ethylacrylamide (p.a) ABCR GmbH&Co KG, Germany

n-Hexane(p.a) Roth

Hydrochloric acid Fluka

Hydrogen peroxide Sigma-Aldrich

4-Hydroxybenzophenone Fluka

N-Isopropylacrylamide Sigma-Aldrich

Lithoumchloride Sigma-Aldrich

Methanol (p.a) Roth

Methylacrylamide ABCR GmbH&Co KG, Germany

Methacrylolychloride,98% Fluka

Polyurethane (medical grade) University Hospital Freiburg

Potassium bromide (FT-IR) Sigma-Aldrich

Potassium bromide Sigma-Aldrich

Potassium chloride,99% Sigma-Aldrich

Potassium nitrate,99% Sigma-Aldrich

Propylacrylamide ABCR GmbH&Co KG, Germany

2-Propanol (p.a) Roth

Sodium bromide Sigma-Aldrich

Sodium chloride Sigma-Aldrich

Sodium hydroxide Sigma-Aldrich

Sodium sulfate Merck

Sulfuric acid Merck

Toluene (p.a) Roth

Triethylamine Roth

Triethoxysilane Fluka

Water (Millipore,18MΩcm-1) --

11.1. Materials

170

11.1.2. Substrates

Glass substrates (19 mm diameter), Thickness 0.28 - 0.32

mm

Plano GmbH, Germany

*Silicon wafer, single side polished (oxide layer ≈1.7)

Diameter 125 mm, Resistivity 1-5Ωcm, Thickness

600±25 nm

Si-Mat, Kaufering, Germany

*Silicon wafer, double side polished (oxide layer ≈1.7)

Diameter 125 mm, Resistivity 1-5Ωcm, Thickness

600±25 nm

Si-Mat, Kaufering, Germany

LaSFN9 substrates coated with chromium (0.5-2nm) and

gold (45-50nm) for SPR

RES-TEC GmbH, Germany

ATR crystal, ZnSe, dimension 56x10x4 mm/ 45° RESULTEC, Germany

*Ordered by Miss. Wibke Hartleb, and Dr. Katrin Moosman (Currently at BASF,

Ludwigshafen, and Germany), Department of Microsystems Engineering, University of


11.1.3. Materials for cell experiments

12-well plates Nunk TM

*Adenosine diphosphate (ADP) MöLaboratory, Germany

Anti-P selectin antibody R&D Systems, Menneapolis, MN

*Biotinylated microbubbles Sigma-Aldrich

Centrifugation tubes (15 ml and 30 ml) Nunk TM

Cryo-SFM (Freezing medium for HUVEC and NHDF) PromoCell GmbH

Culture flask -25 cm2(Polystyrene) Nunk TM

Culture flask -75 cm2(Polystyrene) Nunk TM

Dulbecco‟s MEM (L929 growth medium) BioChrom

11.1. Materials

171

Dulbecco‟s phosphate buffered saline w/o Ca2+ and Mg 2+,

PBS

BioChrom

Endothelial cell growth medium PromoCell GmbH

Ethanol, absolute Sigma-Aldrich

Fetal Bovine serum (FBS) Superior BioChrom

Fibroblasts growth medium PromoCell GmbH

Flow chamber Glycotech

Formaldehyde Fluka

Glutaraldehyde Fluka

Haemacytometer VWR international GmbH

Hexamethyldisiloxane (HMDS) Sigma-Aldrich

Hepess buffer PromoCell GmbH

*Human Fibrinogen Enzyme Research Laboratory,

Swansea, United Kingdom

Human umbilical vein endothelial cells (HUVEC) PromoCell GmbH

L-Glutamine BioChrome

Micro tips sterilized (100μl, 200μl and 1000μl) Nunk TM

Normal human dermal fibroblasts (NHDF) PromoCell GmbH

Penicillin streptomycin solution Sigma-Aldrich

Pipettes (1ml, 5ml, 10ml and 25 ml) Nunk TM

*Platelet cells (Fresh) ‡Volunteer

†Platelet rich plasma (PRP) ‡Volunteer

*Polystyrene culture dish (35mm) Corning, NY

*Sialyl Lewisa-polyacrylamide biotin Glycotech,Gaithersburg,MD

*Streptavidin Sigma-Aldrich

Trypsin/EDTA Sigma-Aldrich

Tryptan Blue (0.4% (w/v)) Sigma-Aldrich

11.2. Instrumentations

172

Trypsin Neutralizing solution (TNS) PromoCell GmbH

Trypsin/EDTA solution PromoCell GmbH

*Ordered by Dr. Felix Guenther, MD, Department of Human genetics, University hospital

Freiburg, Freiburg, Germany. †Prof. Barbara Zieger, Department of pediatrics, University hospital Freiburg, Freiburg,

Germany. ‡C.K Pandiyarajan, Department of Microsystems and engineering, University of Freiburg,

Freiburg, Germany.


11.2.1. Gel permeation chromatography (GPC)

The molecular weight of the synthesized precursor copolymer and tricopolymers

were investigated using GPC with the model Agilent 1100 from polymer standard service

(PSS). Polystyrene and polymethylmethacrylate (PMMA) standard were employed with the

concentration of 1-3 mg/ml in DMF or THF column at the flow of 1 ml/min.

11.2.2. Attenuated Total internal Reflection (ATR) & Fourier Transform -Infrared

spectroscopy (FT-IR) Spectroscopy

ATR and FTIR spectra were recorded using a BioRad Excalibur FTS 3000

spectrometer. For ATR-FTIR spectroscopy, Zn-Se (Res-Tech) was used for the measurement.

FT-IR spectroscopy, potassium bromide pallets and sodium chloride crystals were used for

solid and liquid samples, respectively. The recorded spectra were in the range of 4000-450

cm-1 with the resolution of 2cm-1.


173

11.2.3. Nuclear Magnetic Resonance (NMR) spectroscopy

1HNMR (Proton NMR) and 13CNMR (Carbon NMR) spectra were acquired using

Avance 250MHz spectrometer from Brucker. For 1HNMR, the sample was dissolved in CDCl3

or d6-DMSO with the concentration of 10-20 mg/ml. For 13CNMR, the concentration of

solution was around 100-150 mg/ml.

11.2.4. Dip coating and spin coatings

The samples for SPR and Ellipsometry were prepared through dip coating using Z 2.5

tension testing machine from Zwicki GmbH. The precursor polymer with the concentration

5-20 mg/ml in a good solvent was prepared. The substrate sample, which was held using a

sample holder was dipped or immersed at the velocity of 100 mm/min and pulled at 100

mm/min. All of the dip coating was performed at an ambient room condition with the

temperature of around 25°C.

The spin coating of the samples was achieved by B.L.E. Delta 10 Spin coater at a

rotation speed of 2500 rpm for 60S at room temperature of around 25°C. The spin coating

was mainly employed to prepare thick polymer layers, like in the case of swelling and often

for cell experiments with extra care.

11.2.5. Contact angle measurements

Static contact angles of the surface-attached networks were measured using the sessile

drop method though an Optical Contact Angle Meter (OCA 20) setup (Dataphysics,

Filderstadt, Germany), with a liquid dispenser and an image processing program which

provides an automatic determination of the contact angles. The temperature during the

measurement was controlled by using peltier - 50.


174

11.2.6. Ellipsometry

The deposited film thickness of the polymer layers were measured using auto-nulling

ellipsometry from Nanofilm (EP3) operated at the wavelength of λ= 532 nm at variable

angular modes from 45-75°. The thickness of the film was fitted using software called EP3

view with the help of the self-made model. All the thickness analysis was performed with an

object of 5x and at four zone measurements.

11.2.7. Atomic force microscopy (AFM)

AFM micrographs were recorded using Nanoscope IIIa microscope (digital

instrument, Santa Barbara, CA) and the microscope was operated in tapping mode, using

commercial tips with a frequency of ≈330kHz, and spring constant of ≈42 N/m. All the AFM

micrographs were obtained in ambient condition at room temperature of around 25°C.

11.2.8. Surface Plasmon Resonance (SPR) and Optical waveguide spectroscopy (OWS)

SPR and OWS measurements were performed with an instrument constructed with

Kretchmann configuration from RES-TEC. Where, He-Ne having the wavelength of 633 nm

(λ = 632.8 nm) was used as an efficient light source. The optical components of the films

were obtained by using the Fresnel simulation with the help of software winspall 3.02

developed by the Max Planck Institute for Polymer Research at Mainz. The real (ε1) and

imaginary (ε11) permittivity constants of the metal layers are summarized in the table 10.1.


175

Figure 11.1: (A) Instrumentation of SPR/OWS. (B) Kretchmann configuration used in the SPR.

Table 11.1: Optical components materials that are employed in the SPR measurements.

Layer Thickness (nm) ε1 ε11

LaSFN9 --* 3.4036 0

Cr 0.3-2 -6.3 17-20

Au 45-50 -12.3 1.29

Polymers --† 2.2-2.4 0

H2O 1.769 0

PBS 0 1,769 0

N2 0 1 0

*Glass substrate with high refractive index of around 1.8

†Thickness varied from 400 nm to 800 nm.

(B)

(A)

11.3. Synthesis and characterizations of polymers

176

11.2.9. Peristaltic pump

The protein and PBS buffer solution was injected to the SPR flow cell using peristaltic

pump from Ismatec®, Type 1SM596D, Switzerland. The pump was operated in flow mode

with the flow rate of 100μl s-1. The machine was calibrated prior to the usage by measuring

the amount of liquid with respect to the time using the standard measuring cylinder.

11.2.10. Humidity sensor

The humidity and temperature of the SPR chamber was measured using Testo© 635-

1/2 (USA) humidity sensor. It has the detection resolution of 0.1% and can measure from 0 to

100% humidity. The accuracy and performance of the sensor was checked by measuring the

humidity and temperature of the standard cell culture incubator, which has automated with

the humidity sensor.

11.2.11. Calorimetry

The heat of hydration of the polymers were measured using solution calorimetry from

Thermometric Precision Solution Calorimeter (Thermometric AB) along with the software

called SolCal.

11.3. Synthesis and characterization of monomers and polymers

11.3.1. Synthesis 4-[3-(Triethoxy silyl)propyloxy]benzophenone

5g of 4-Allyloxy benzophenone was dissolved in ca.20 ml of freshly distilled dimethyl

chloro silane. 70 mg of Pt-C was then added to the reaction mixture slowly which was then

refluxed at 120°C for 5 hrs under nitrogen. After completion of reaction, the excess triethoxy

silane was removed by evaporation under vacuum. The catalyst was removed by filtration of


177

the solution under nitrogen. This solution was then directly used to prepare the surface

attached polymer networks.

FT-1NMR: δ = 0.8 ppm (t, 2H, CH2-*CH2-Si), δ = 1.2 ppm (m, 9H, *CH3-CH2-O), δ = 1.9 ppm

(p, 2H, O-CH2-*CH2-CH2-Si), δ = 3.8 ppm (q, 6H, CH3-*CH2-O), δ = 4.0 ppm (t, 2H, O-*CH2-

BP), δ = 7.8-6.9 ppm (m, 9H, C-Haro)

11.3.2. Synthesis of 4-Methacryloyl-oxy-benzophenone (MABP)

In a typical run, 11.5 g of methacryloyl chloride in 40 ml of dichloromethane (DCM)

was added to a solution containing 19.8 g of 4-hydroxybenzophenone and 11.12 g of

triethylamine in 200 ml of DCM, while the reaction mixture was cooled in an ice bath. The

mixture was allowed to warm to room temperature and stirred overnight. After completion

of reaction, the mixture was thoroughly washed with 5% HCl, NaHCO3 solution and water

(Three times each). Then, the solution (organic phase) was added into a beaker containing

cold n-hexane (1:10) and kept in a freezer at -20 °C overnight to complete precipitation:

Subsequently, the precipitate was filtered off at 0°C, dried in vacuum and re-crystallized

from 10% ethyl acetate- hexane mixture. (1:9).

FT-IR: 3049(=C-H), 2980-2927(-C-H), 1732(-C=O), 1653(-HN-C=O), 1596(-C=C-) and 1446(-

C-H) cm-1

FT-1NMR: δ = 2.1 ppm (s, 3H, *CH3), δ = 5.7-6.3 ppm (2s, 2H, *CH2=), δ = 7.1-7.9 ppm

(various m, 9H, C-Haro)

11.3.3. Synthesis of Poly (alkylacrylamide-co-MABP)

A 100 ml Schlenk tube was dried using vacuum and charged with the 97 mmol of

acrylamides , 3 mmol of MABP and 14 mg of AIBN. The above mixture was dissolved in 20-

25 ml of MeOH for MAAm, DMAAm, EAAm and PAAm and THF for BAAm and DEAAm.


178

Then, the schenk tube was closed carefully and degassed under nitrogen through three freeze

and thaw cycles: Polymerization was carried out at 60°C overnight in a thermostated water

bath. After completion of polymerization the polymer was precipitated in diethyl ether and

repeated for three times.

Figure 11.1: Molecular structures of the monomers used in the study.

Table 11.2: Synthetic details of Poly (alkyacrylamide-co-MABP)

Polymer Monome

r

(mmol)

MABP

(mmol)

Solvent for

polymerization

Solvent for

precipitation

*Mol.wt

(g/mol)

P(MAAm-co-MABP) 97 3 Methanol Ether 82000

P(EAAm-co-MABP) 97 3 Methanol Ether 70000

P(PAAm-co-MABP) 97 3 Methanol Ether 66000

P(BAAm-co-MABP) 97 3 THF Water 49000

P(DMAAm-co-MABP) 97 3 Methanol Ether 91000

P(DEAAm-co-MABP) 97 3 THF n-Hexane 52000

*Mw (weight average molecular weight)

Poly (MAAm-co-MABP)

FT-IR: 3314 (-HN-), 3019(=C-H), 2941(-C-H), 1727(-C=O), 1651(-HN-C=O), 1546 (-

C=C-) and 1410(-C-H) cm-1


179

FT-1NMR: δ = 1.9-1.4 ppm (*CH2-CH), δ = 2.5 ppm (*CH3-NH), δ = 3.3 ppm (*CH-

CH2), δ = 7.8-7.3 ppm (C-Haro)

Poly (EAAm-co-MABP)

FT-IR: 3319 (-HN-), 3018 (=C-H), 2972- 2934(-C-H), 1743(-C=O), 1646(-HN-C=O),

1531(-C=C-) and 1448 (-C-H) cm-1

FT-1NMR: δ = 1.0 ppm (*CH3-CH2-NH), δ = 1.9-1.5 ppm (*CH2-CH), δ = 3.0ppm (*CH-

CH2), δ = 3.3 ppm (*CH2-NH), δ = 7.8-7.2 ppm (C-Haro)

Poly (PAAm-co-MABP)

FT-IR: 3314(-HN-), 3081(=C-H), 2962-2933(-C-H), 1741(-C=O), 1650(-HN-C=O),

1536(-C=C-) and 1459 (-C-H) cm-1

FT-1NMR: δ = 0.9 ppm (*CH3-CH2-CH2-NH), δ = 1.5 ppm (CH3-*CH2-CH2-NH), δ =

2.3-1.7 ppm (*CH2-CH), δ = 2.9 ppm (*CH-CH2), δ = 3.15 ppm (*CH2-NH), δ = 7.9-7.5

ppm (C-Haro)

Poly (BAAm-co-MABP)

FT-IR: 3317(-HN-), 3080(=C-H), 2958-2931(-C-H), 1741(-C=O), 1647(-HN-C=O),

1539(-C=C-) and 1462(-C-H) cm-1

FT-1NMR: δ = 0.9(*CH3-CH2-CH2-CH2-NH), δ = 1.3(CH3-*CH2-CH2-CH2-NH), δ = 1.5

ppm (CH3-CH2-*CH2-CH2-NH), δ = 2.2-1.7(*CH2-CH), δ = 2.7(*CH-CH2), δ = 3.2 ppm

(CH3-CH2-CH2-*CH2-NH), δ = 7.9-7.5(C-Haro)


180

Poly (DMAAm-co-MABP)

FT-IR: 3061(=C-H), 2928-2873(-C-H), 1743(-C=O), 1641(-N-C=O), 1496(-C=C-) and

1399 (-C-H) cm-1

FT-1NMR: δ = 1.8-1.2 ppm (*CH2-CH), δ = 2.94-2.9 ppm (*(CH3)2-N), δ = 3.1 ppm

(*CH-CH2), δ = 7.9-7.4 ppm (C-Haro)

Poly (DEAAm-co-MABP)

FT-IR: 3065(=C-H), 2971-2933(-C-H), 1741(-C=O), 1636(-N-C=O), 1451 (-C=C-) and

1377(-C-H) cm-1

FT-1NMR: δ = 1.0ppm (*CH3-CH2-N), δ = 1.9-1.5 ppm (*CH2-CH), δ = 3.0 ppm (*CH-

CH2), δ = 3.3 ppm (*CH2-N), δ = 7.8-7.2 ppm (C-Haro)

11.3.4. Synthesis of Tri-copolymers

100 ml schlenk tube was dried with vacuum and back filled with N2 for three times.

The schelenk tube was then charged with calculated amount of ethylacrylamide or dimethyl

acrylamide, butylacrylamide or diethylacrylamide and mole percent of MABP and 0.014g of

AIBN (See table 3.2). The above mixture was dissolved in 20 ml of MeOH and 5 ml of 1, 4-

Dioxane followed by degassed with N2 using freeze and thaw (3 times). Polymerization was

carried out at 60°C for 16-18 hrs. After the completion of reaction, the solution was diluted

in the solvent used for polymerization and precipitated in not solvents (see table 10.3). Re-

precipitation was repeated thrice and dried under vacuum at 25°C.

Note: In the case of poly (25%DMAAm-co-75%DEAAm-co-3% MABP), after completion of

the polymerization, the solvent was evaporated and dissolved in THF and precipitated in

petroleum ether.


181

Table 11.3: Synthetic details of Tri-copolymers

*Polymer Solvent for

polymerization

Solvent for

precipitation

Mol.wt

(g/mol)

Poly(75%EAAm-co-25%BAAm-co-3% MABP) Methanol †Ether 250000

Poly(50%EAAm-co-50%BAAm-co-3% MABP) Methanol †Ether 56000

Poly(25%EAAm-co-75%BAAm-co-3% MABP) THF Water 230000

Poly(75%DMAAm-co-25%DEAAm-co-3% MABP) Methanol †Ether 150000

Poly(50%DMAAm-co-50%DEAAm-co-3% MABP) Methanol Cyclohexane 86000

Poly(25%DMAAm-co-75%DEAAm-co-3% MABP) Methanol ‡Pet. ether 120000

*75% = 73.5 %, 50% = 48.5% and 25% = 23.5% , †Diethyl ether, ‡Petroleum ether

Poly (75%EAAm-co-25%BAAm-co-3% MABP)

FT-1NMR: δ = 0.92 ppm (*CH3-CH2-CH2-CH2-NH), δ = 1.12 ppm (*CH3-CH2-NH), δ =

1.34-1.47 ppm (*CH2-CH), δ = 3.0ppm (*CH-CH2), δ = 3.18 ppm (CH3*CH2-NH), δ =

3.35 ppm (CH2*CH2-NH)δ = 7.8-7.5 ppm (C-Haro)



1.36-1.49 ppm (*CH2-CH), δ = 3.0ppm (*CH-CH2), δ = 3.21 ppm (CH3*CH2-NH), δ =

3.35 ppm (CH2*CH2-NH)δ = 7.8-7.5 ppm (C-Haro)



1.47-1.34 ppm (*CH2-CH), δ = 2.3ppm (*CH-CH2), δ = 3.18 - 3.32cppm (CH3*CH2-NH),

δ = 7.8-7.46 ppm (C-Haro)

11.4. Deposition of surface-attahced networks

182

Poly (75%DMAAm-co-25%DEAAm-co-3% MABP)

FT-1NMR: δ = 0.92 ppm ((*CH3-CH2)2-N), δ = 1.88-1.66ppm (*CH2-CH), δ = 2.92-2.85

ppm (*(CH3)2-N), δ = 3.12 ppm (*CH-CH2), δ = 3.47-3.42ppm (CH3*CH2-NH) δ = 7.80-

7.50 ppm (C-Haro)


FT-1NMR: δ = 1.09 ppm ((*CH3-CH2)2-N), δ = 1.83-1.62ppm (*CH2-CH), δ = 3.0-2.83

ppm (*(CH3)2-N), δ = 3.12 ppm (*CH-CH2), δ = 3.42-3.33ppm (CH3*CH2-NH) δ = 7.82-

7.47 ppm (C-Haro)


FT-1NMR: δ = 1.08 ppm ((*CH3-CH2)2-N), δ = 1.87-1.56ppm (*CH2-CH), δ = 2.81 ppm

(*(CH3)2-N), δ = 3.2 ppm (*CH-CH2), δ = 3.56-3.42ppm (CH3*CH2-NH) δ = 7.80-7.50

ppm (C-Haro)

11.4. Deposition of surface-attached networks

11.4.1. Silanization on Silica wafer/Glass surface

Glass slides were cleaned by immersing them in piranha solution (30 ml of conc.

H2SO4: 10 ml of H2O2) for 1 hr followed by thorough washing with DI water and drying with

N2. Bp-Si (30 mmol) was freshly prepared in toluene and spin cast onto the cleaned substrate

at 2500 rpm for 20 seconds. The samples were then covered with aluminum foil, and

annealed at 120°C overnight, followed by extraction in toluene and drying over N2 gas,

which leaves a monolayer of Bp-Si (4-6 nm).

11.5. Preparation of protein solusion for SPR

183

11.4.2. Preparation of surface attached poly (alkylacrylamide) network

Table 11.4: Experimental conditions for spin coating.

Precursor polymer Solvent Concentration

(mg/ml)

Speed

(rpm)

Time

(s)

P(MAAm-co-3%MABP) Ethanol 10 2500 60

P(EAAm-co-3%MABP) Ethanol 10 2500 60

P(PAAm-co-3%MABP) Butanol 10 2500 60

P(BAAm-co-3%MABP) †Butanol 10 2500 60

P(DMAAm-co-3%MABP) Ethanol 10 2500 60

Poly(75%EAAm-co-25%BAAm-co-3% MABP) Ethanol 10 2500 60

Poly(50%EAAm-co-50%BAAm-co-3% MABP) Ethanol 10 2500 60

Poly(25%EAAm-co-75%BAAm-co-3% MABP) Butanol 10 2500 60

Poly(75%DMAAm-co-25%DEAAm-co-3% MABP) Butanol 10 2500 60

Poly(50%DMAAm-co-50%DEAAm-co-3% MABP) Ethanol 10 2500 60

Poly(25%DMAAm-co-75%DEAAm-co-3% MABP) Ethanol 10 2500 60

Poly(DMAAm-co-1% MABP) Ethanol 10 2500 60

Poly(DMAAm-co-2.5% MABP) Ethanol 10 2500 60

Poly(DMAAm-co-5% MABP) *IPA 10 2500 60

Poly(DMAAm-co-7.5% MABP) Butanol 10 2500 60

Poly(DMAAm-co-10% MABP) †Butanol 10 2500 60

*Isopropanol, †Hot (50-60°C)

Solutions of the precursor polymer were prepared with a concentration of 10mg/ml

and filtered (0.45μm filter). 50 µl of these solutions were used to spin cast a thin polymer

layer onto the BP-Si modified glass surface (2500 rpm for 60s). After evaporation of the

solvent, the obtained layers were illuminated with a total dosage of 4 J/cm2 with a UV-A-

11.6. Cell culture

184

light source (Stratalinker 2400, Stratagene, CA, USA). The samples are then extracted in

solvent and with nitrogen gas.

11.5. Preparation of Protein solution for SPR

The Proteins like 10 mg of Fibrinogen or Fibronectin were dissolved in 10 ml of PBS

buffer (pH=7.4). The above mixture is gently mixed to obtain the homogenous solution

without air bubbles. The solution is then filtered with 0.45μm filter to remove the unwanted

particles and used for the experiment. Similarly, other proteins vitronectin and von

willebrand factor is dissolved in de-ionized water with the concentration of 1 mg/ml.

11.6. Cell culture

11.6.1. Normal Dermal Human Fibroblasts (NHDF)

NHDF cells were cultured in a 75-cm2 culture containing 15 ml of growth medium at

37C, 5% (V/V) CO2 and steam saturated atmosphere. When the cell confluent reached 60-

80%, it was viewed under the microscope. The culture flask was brought into a laminar flow

cabinet from the incubator and the growth medium was carefully removed using a sterile

pipette. The medium was replaced with 7.5 ml of Hepes buffer, washed for 30S by swirling

gently, and removed carefully with a sterile pipette. The trypsin / EDTA solution was added

to the flask (100µm/cm2) and gently swirled for 30S and was examined under microscope.

When more than 50% of the cells were detached from the surface, then 7.5 ml of trypsin

neutralizing solution was added and the flask was washed thoroughly using a sterile pipette

and transferred to an 15 ml falcon tube without any air bubble. The cells were centrifuged at

about 220 x g for 4 minutes at room temperature, followed by the removal of the clear upper

layer. 5 ml of medium was then added and the cells were re-suspended. Cells were counted

using haemocytometer and the concentration was adjusted to 0.5x 106 cells/ml, distributed to

11.7. Platelet adheison

185

the culture flask and placed in an incubator. The culture medium was changed for every 24

hrs and the confluence of cells was examined under the microscope.

11.6.2. Human Umbelical Vein Endothelial cells (HUVEC)

---- Refer 10.6.1 ----

11.6.3. Mouse fibroblasts (L929)

L929 cells were cultured in a 75-cm2 culture containing 15 ml of growth medium at

37C, 5% (V/V) CO2 and steam saturated atmosphere. When the cell confluent reached 60-

80%, it was viewed under the microscope. The culture flask was brought into a laminar flow

cabinet from the incubator and the growth medium was carefully removed using a sterile

pipette. The medium was replaced with 7.5 ml of PBS buffer, washed for 30S by swirling

gently, and removed carefully using a sterile pipette and repeated once more. 1.5 ml of

trypsin / EDTA solution was then added to the flask and gently swirled and removed

immediately. The flask was incubated for 90s, 10 ml of the growth medium was added to the

flask, and the cells were suspended and transferred to a 15 ml falcon tube. The cells were

centrifuged at about 220 x g for 4 minutes at room temperature and the clear upper layer was

removed. The cells were suspended in 5 ml of the growth medium and counted using

haemocytometer. The cell suspension was then distributed to the culture flask and placed in

an incubator. The culture medium was changed for every 24 hrs and the confluence of cells

was examined under the microscope.

11.7. Platelet extraction

11.7.1. Preparation of platelet concentrate

50-60 ml of Blood from healthy volunteers under no medication was anti-coagulated

with citric acid and centrifuged at 150 x g for 10 minutes. The resulting platelet rich plasma

(PRP) was separated and again centrifuged at 2000 x g for 10 minutes. The resulting platelet

11.6. Platelet extraction

186

repellet was carefully separated from the platelet poor plasma (PPP). The platelet cells were

then suspended in phosphate buffer (PBS) in a concentration of about 250,000 platelets ml.

The platelets were activated by adding 20 µM of adenosine diphosphate (ADP) and used for

the flow experiments.

11.7.2. Preparation of Microbubbles (MB)

The Microbubbles (MB) were prepared using the procedure established by Sakamoto

et al.291 Briefly, MB‟s are composed of decafluorobutane gas core encapsulated by a

phosphatidylcholine lipid shell. A brush of polyethylene glycol was attached onto the lipid

shell via a lipid anchor embedded in the lipid monolayer. Some of the polyethylene glycol

molecules contained biotin at the distal tip.40.291

Targeting was achieved by coupling biotinylated sLea ligand to the biotinylated MB

via a streptavidin linker (Figure 7.6). Biotinylated MB‟s were washed with phosphate

buffered saline by centrifugation (400 x g for 3 minutes) to remove excess of unincorporated

lipid, and incubated with streptavidin for 10 minutes at room temperature (3 µg streptavidin

per 107 MB). Again, excess streptavidin was removed by centrifugation. The MB‟s were

incubated either with biotinylated sLealigand or with a biotinylated control polyacrylamide

(control-PAA) with the same backbone without carrying the functional sialyl Lewis groups

for 10 minutes at room temperature (1.5 g sLealigand or control-PAA per 107 MB).

MB were washed again to remove unreacted ligand and stored at 4°C. The size

distribution of the MB‟s was determined using scanning electron microscopy (SEM)

observations. Aggregation of MB as the result of ligand coupling procedure was minimal.

Mean MB size was 2.71±0.05µm.291

12. Appendix

187

12. Appendix

11.1. Swelling of surface poly (alkylacrylamide) networks

Figure.12.1: Reflectivity scan of poly(MAAm-co-3%MABP). Solid lines are the Fresnel

formalism.

Figure 12.2: Reflectivity scan of poly (EAAm-co-3%MABP). Solid lines are the Fresnel

formalism.

12. Appendix

188

Figure 12.3: Reflectivity scan of poly (PAAm-co-3%MABP). Solid lines are the Fresnel

formalism.

Figure 12.4: Reflectivity scan of poly (DEAAm-co-3%MABP). Solid lines are the Fresnel

formalism.

12. Appendix

189

Table 12.1: Optical components of poly (MAAm-co-3%MABP) at different humidity.

Poly (MAAm-co-3%MABP)

Salt RH 5

(%)

Thickness

(nm) Permittivity

Refractive

index Swelling factor

N2 air 0 547,1 2,32 1,523 1

LiCl 25 589,1 2,312 1,520 1,08

K2CO3 46 624,1 2,295 1,515 1,14

NaBr 62 643 2,281 1,510 1,17

NaCl 75 691,3 2,254 1,501 1,26

KCL 86 722,6 2,2269 1,492 1,32

K2SO4 96 780,8 2,197 1,482 1,43

Table 12.2: Optical components of poly (EAAm-co-3%MABP) at different humidity.

Poly (EAAm-co-3%MABP)

Salt RH 5

(%)

Thickness

(nm) Permittivity

Refractive


Dry 0 544,2 2,296 1,515 1

LiCl 25 580,5 2,29 1,513 1,07

K2CO3 46 597,5 2,28 1,509 1,09

NaBr 62 617 2,266 1,505 1,13

NaCl 75 624,3 2,262 1,503 1,15

KCL 86 649,5 2,246 1,498 1,19

K2SO4 96 667,4 2,233 1,494 1,23

12. Appendix

190

Table 12.3: Optical components of poly (PAAm-co-3%MABP) at different humidity.

Poly (PAAm-co-3%MABP)

Salt RH 5

(%)

Thickness

(nm) Permittivity

Refractive


Dry 0 624,46 2,2925 1,514 1

LiCl 25 649,16 2,2831 1,510 1,04

K2CO3 46 663,95 2,2783 1,509 1,06

NaBr 62 674,14 2,271 1,506 1,08

NaCl 75 686,21 2,264 1,504 1,10

KCL 86 701,71 2,2521 1,500 1,12

K2SO4 96 711,49 2,2451 1,498 1,14

Table 12.4: Optical components of poly (DEAAm-co-3%MABP) at different humidity.

Poly (PDEAAm-co-2.5%MABP)

Salt RH 5

(%) Thickness (nm) Permittivity

Refractive


Dry 0 562,82 2,2837 1,5111 1

LiCl 25 573,75 2,2811 1,5103 1,019

K2CO3 46 576,71 2,2805 1,5101 1,02

NaBr 62 582,07 2,28 1,5099 1,03

NaCl 75 587,3 2,2788 1,5095 1,04

KCL 86 594,6 2,2748 1,5082 1,06

K2SO4 96 598,1 2,271 1,5069 1,062

12. Appendix

191

Table 12.5: Optical components of poly (BAAm-co-3%MABP) at different humidity.

12.2. Synthesis and characterizations

12.2.1. Alfrey and Price Q, e Model.125,139

Monomer -A is Dimethyl acrylamide (DMAAm)

r1= 0.8653 (monomer reactivity ratio)

Q 1= 0.41 (reactivity)

e = -0.26 (polarity)

Monomer -B is Dimethyl acrylamide (DMAAm)

r1= 1.1527 (monomer reactivity ratio)

Q 1= 0.48 (reactivity)

e = -0.31 (polarity)

Monomer conversion can obtained from the following equation:

𝑋1 =𝑓𝐴 𝑟1𝑓𝐴+𝑓𝐵

𝑓𝐵 𝑟2𝑓𝐵+𝑓𝐴 (Eq. 11.1)

Poly (PBAAm-co-2.5%MABP)

Salt RH 5

(%)

Thickness

(nm) Permittivity

Refractive


Dry 0 741,6 2,253 1,500 1

LiCl 25 757 2,247 1,499 1,02

K2CO3 46 774,1 2,243 1,497 1,04

NaBr 62 780 2,242 1,497 1,05

NaCl 75 785,8 2,241 1,497 1,06

KCL 86 798,1 2,239 1,496 1,08

K2SO4 96 805,1 2,235 1,495 1,09

12. Appendix

192

The calculated converted amount is shown in the Figure 11.5, where the converted amount

is plotted as a function of added amount.

Figure 12.5: Q e model calculations obtained from monomer DMAAm and DEAAm

conversation.

12.2.2. AFM Images

Figure 12.6: AFM image of Poly (MAAm-co-3%MABP)

12. Appendix

193

Figure 12.7: AFM image of Poly (DMAAm-co-3%MABP)

Figure 12.8: AFM image of Poly (PAAm-co-3%MABP)

Figure 12.9: AFM image of Poly (DEAAm-co-3%MABP)

12. Appendix

194

Figure 12.10: AFM image of Poly (BAAm-co-3%MABP)

12. Appendix

195

12.3. Cell adhesion

Figure 12.5: Micrographs of NHDF cellular adhesion on surface-attached poly (alkyl

acrylamide) networks with time.

Journal publications and conference proceedings

196

Journal Publications

1. Pandiyarajan, C.K.; Prucker, O.; Zieger, B.; Rühe, J., Influence of the molecular

structure of surface-attached poly (alkylacrylamide) coatings on blood platelet

adhesion, Macromol. Biosci., 2013.

2. Pandiyarajan, C.K.; Prucker, O.; Zieger, B.; Rühe, J., Poly (alkylacrylamide) based

hydrogel coatings for blood devices, PMSE, 2012, 106, 349

3. Pandiyarajan, C.K.; Prucker, O.; Rühe, J., Swelling of surface-attached poly (alkyl

acrylamide) network in humid air, (2013 – expected).

Referred conference proceedings

1. Pandiyarajan, C.K.; Prucker, O.; Zieger, B.; Rühe, J., Poly (alkylacrylamide) based

hydrogel coatings for blood devices, 243rd ACS National Meeting and Exposition, San

Diego, CA, United States, Mar 25-Mar30, 2012. Poster presentation.

2. Pandiyarajan, C.K.; Prucker, O.; Zieger, B.; Rühe, J., Blood compatible hydrogel

coatings, Bayreuth Polymer Symposium-2011 (International symposium), University

of Bayreuth, Bayreuth, Germany, Sep 09- Sep12, 2011. Poster presentation.

3. Pandiyarajan, C.K.; Prucker, O.; Goto, S.; Zieger, B.; Rühe, J., Blood compatible

hydrogel coatings for VAD, IRTG summer school, Mittelwihr, France, July 10-14,

2011. Poster and oral presentations.

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Acknowledgements

I would like to express my sincere gratitude to Prof. Jürgen Rühe for giving me a

wonderful opportunity to work in his research group. I thank him for his excellent

supervision and many fruitful discussions that helped me to learn a proper science rather

than producing some results. I also thank him for his mentorship and help for my career.

I also sincerely thank Dr. Oswald Prucker for his guidance in this project. From the

beginning, he was a great support and very handful in all aspects of my research. Especially,

his help on humidity swelling measurements were highly appreciated. I also thank him for

his timely effort on correcting my thesis.

I take this opportunity to acknowledge my project collaborators Prof. Barbara Zieger

and Dr. Felix Günther, The University Hospital, Freiburg, Germany; for granting me the

opportunity to work in their lab and to learn more on platelet adhesion.

I greatly appreciate Mr. Joshua Markum, Mr. David Kroman, Mrs. Lisa Turnhoff, and

Mr. Fan Wu for doing their Bachelor‟s or Master‟s thesis with me and for their significant

contributions in the success of this research.

It is a great pleasure to acknowledge Dr. Josephine Selvaraj, Carson city, Nevada,

USA; for reading and correcting my thesis. Special thanks to Dr. Micael Henze for translating

my short summary to german (Kurzezusammenfassung). I am also grateful to Mrs. Maya

Narayanan Nair, Dr. Jon Green, Dr. A.G.Venkatesh and Mr. Nils Korf for their useful

suggestions and corrections in the thesis.

I thank Mrs. Petra Hettich, Mrs. Gudrun Condrad and Mrs. Waltraud Hanzer for

their help regarding administration procedures. My special thanks to Mrs. Petra Hettich for

helping many international students, especially for those who are new to the group. She is

one good assert in the group and deserves a high level of appreciation.

Acknowledgements

211

I thank Dr. Yi Thomann and Mr. Jonas Groton for their help regarding SEM images,

Mr. Xiaoqiang Hou for his timely help on AFM images, and Mrs. Natalia Schatz for synthetic

help and GPC measurements. I thank Mr. Martin Schoenstein for calorimetric and Mr.

Vitaly Kondrashov for contact angle measurements.

I thank Mr. Martin Marazita for sharing hood with me and being a wonderful friend

in the group. I greatly appreciate the company of Dr. Anke Würz, Dr. Chrishan Schuh, Dr.

Christine Bunte, Dr. Kerstin Schuh, Dr. Jacob Belardi, Mr.Gregor Osterwinter, Mrs. Melanie

Eichhorn and Mr. Vitaly Kondrashov for sharing office with me.

Finally, I thank and all members of CPI who provided me a wonderful work

environment that felt like home. Thank you one and all.

Documents

The interaction of blood proteins and platelets on surface