• TABLE OF CONTENTS
HIDE
 Title Page
 Acknowledgement
 Table of Contents
 List of Tables
 List of Figures
 Abstract
 Introduction
 Background
 Pulsed laser ablation deposition...
 Nanoindention and nanoscratch measurements...
 Empirical model for ablation of...
 Functionalization of stainless-steel...
 Conclusions and future work
 References
 Biographical sketch














Title: Surface modification of biomaterials by pulsed laser ablation deposition and plasma/gamma polymerization
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 Material Information
Title: Surface modification of biomaterials by pulsed laser ablation deposition and plasma/gamma polymerization
Physical Description: Book
Language: English
Creator: Rau, Kaustubh R
Publisher: University of Florida
Place of Publication: Gainesville Fla
Gainesville, Fla
Publication Date: 2000
Copyright Date: 2000
 Subjects
Subject: Materials Science and Engineering thesis, Ph. D   ( lcsh )
Dissertations, Academic -- Materials Science and Engineering -- UF   ( lcsh )
Genre: government publication (state, provincial, terriorial, dependent)   ( marcgt )
bibliography   ( marcgt )
theses   ( marcgt )
non-fiction   ( marcgt )
 Notes
Summary: ABSTRACT: Surface modification of stainless-steel was carried out by two different methods: pulsed laser ablation deposition (PLAD) and a combined plasma/gamma process. A potential application was the surface modification of endovascular stents to enhance biocompatibility. The pulsed laser ablation deposition process had not been previously reported for modifying stents and represented a unique and potentially important method for surface modification of biomaterials. Polydimethylsiloxane (PDMS) elatomer was studied using the PLAD technique. Cross-inked PDMS was deemed important because of its general use for biomedical implants and devices as well as in other fields. Furthermore, PDMS deposition using PLAD had not been previously studied and any information gained on its ablation characteristics could be important scientifically and technologically. The studies reported here showed that the deposited silicone film properties had a dependence on the laser energy density incident on the target. Smooth, hydrophobic,silicone-like films were deposited at low energy densities (100-150 mJ/cm2). At high energy densities (> 200 mJ/cm2), the films had an higher oxygen content than PDMS, were hydrophilic and tended to show a more particulate morphology.
Summary: ABSTRACT (cont.): It was also determined that: 1) the deposited films were stable and extremely adherent to the substrate, 2) silicone deposition exhibited an 'incubation effect' which led to the film properties changing with laser pulse number and 3) films deposited under high vacuum were similar to films deposited at low vacuum levels. The mechanical properties of the PLAD films were determined by nanomechanical measurements which are based on the Atomic Force Microscope (AFM). From these measurements, it was possible to determine the modulus of the films and also study their scratch resistance. Such measurement techniques represent a significant advance over current state-of-the-art thin film characterization methods. An empirical model for ablation was developed for the 248 nm laser irradiation of silicone. The model demonstrated a good fit to the experimental data and showed that silicone underwent ablation by a thermal mechanism. In addition to PLAD studies, functionalization of stainless steel was carried out by a combined plasma/gamma method involving deposition of a hexane plasma polymer by RF plasma polymerization, followed by gamma radiation graft polymerization of methacrylic acid. The hydrograft modified surfaces were further modified by chemisorption reactions with poly(ethylene imine) to produce amine-rich surfaces. Bovine serum albumin was then bound via amino groups using glutaraldehyde coupling. A streaming potential cell was also built and used to measure the zeta potential of these ionic surfaces.
Summary: KEYWORDS: surface modification, pulsed laser deposition, silicone, plasma polymerization, stents
Thesis: Thesis (Ph. D.)--University of Florida, 2000.
Bibliography: Includes bibliographical references (p. 209-220).
System Details: System requirements: World Wide Web browser and PDF reader.
System Details: Mode of access: World Wide Web.
Statement of Responsibility: by Kaustubh R. Rau.
General Note: Title from first page of PDF file.
General Note: Document formatted into pages; contains xv, 221 p.; also contains graphics.
General Note: Vita.
 Record Information
Bibliographic ID: UF00100817
Volume ID: VID00001
Source Institution: University of Florida
Holding Location: University of Florida
Rights Management: All rights reserved by the source institution and holding location.
Resource Identifier: oclc - 50751272
alephbibnum - 002678739
notis - ANE5966

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Table of Contents
    Title Page
        Page i
        Page ii
    Acknowledgement
        Page iii
        Page iv
    Table of Contents
        Page v
        Page vi
        Page vii
    List of Tables
        Page viii
    List of Figures
        Page ix
        Page x
        Page xi
        Page xii
        Page xiii
    Abstract
        Page xiv
        Page xv
    Introduction
        Page 1
        Page 2
        Page 3
        Page 4
        Page 5
        Page 6
    Background
        Page 7
        Page 8
        Page 9
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    Pulsed laser ablation deposition of silicone
        Page 46
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    Nanoindention and nanoscratch measurements on PLAD films
        Page 91
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        Page 123
        Page 124
        Page 125
    Empirical model for ablation of silicone at 248nm
        Page 126
        Page 127
        Page 128
        Page 129
        Page 130
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    Functionalization of stainless-steel by a combined plasma/gamma process
        Page 149
        Page 150
        Page 151
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    Conclusions and future work
        Page 191
        Page 192
        Page 193
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        Page 207
        Page 208
    References
        Page 209
        Page 210
        Page 211
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        Page 213
        Page 214
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    Biographical sketch
        Page 221
Full Text










SURFACE MODIFICATION OF BIOMATERIALS BY PULSED LASER ABLATION
DEPOSITION AND PLASMA/GAMMA POLYMERIZATION.

















By

KAUSTUBH R. RAU


A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL
OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT
OF THE REQUIREMENTS FOR THE DEGREE OF
DOCTOR OF PHILOSOPHY

UNIVERSITY OF FLORIDA


2000




























Copyright 2000

by

Kaustubh R. Rau















ACKNOWLEDGMENTS

I thank Dr. Goldberg, my thesis advisor, for supporting me in my graduate

education. His scholarship, vast knowledge and openness to my ideas have made

working in his lab a pleasant and learning experience. Dr. Singh, my thesis co-advisor

readily agreed to be a part of the PLAD research, for which I am grateful. His thin film

expertise and knowledge of materials science issues have helped me numerous times.

My thesis committee members of Drs. Batich, Brennan and Dickinson have also been

extremely helpful with their comments and suggestions whenever I have bothered them.

The work reported here would not have been completed without the help and

support of key people. Drs. Drew Amery and Chris Widenhouse got me started in the lab

and taught me the "art" of surface modification. In particular, Drew helped me master the

various characterization techniques and Chris helped me out with his materials expertise.

Dr. Emmanuel Biagtan trained me to use the RF plasma chamber. Ahmad (Bob) Hadba

trained me on the SEM and worked with me on several projects most notably the

phospholipid and NVP projects. He was also good to discuss ideas with and helped me

out with his chemistry knowledge. Dr. Jim Fitz-Gerald and Mike Ollinger trained me to

use the excimer laser and vacuum chamber. Mike Grumski worked with me on the

PLAD project and refined the data acquisition software developed by Drew for the XPS.

Uday Mahajan introduced me to the wonderful world of nanoscale measurements and

trained me on the instrumentation. Drs Emmanuel Biagtan and Gwen Clark helped me in









the initial development of the streaming potential cell while Jaime Serrano and Elizabeth

Rees worked on calibrating it as part of their undergraduate research.

To past and present members of the Goldberg research group which include: Dr.

T-L. Lin, Dr. James Marotta, Dr. Lynn Peck, Dr. James Kirk, Paul Martin, Brett Almond,

Lauri Jenkins, Brian Cuevas, Josh Stopek, Dan Urbaniak and Haseeb Jabbar, I would just

like to say a collective thanks for helping me out from time-to-time. I would also like to

thank several other people for their friendship, love and support. They include: Nandini

Mandlik, Raj araman Kalyanraman Kaushik, Arul Chinnappan, Jennifer Russo and Sharad

Mathur.

Finally, I thank my parents, Ravi and Alaka Rau, for giving me the means and

opportunity to come to the United States for further education. While being generous in

their advice and support, more importantly they were unwavering in their belief in me.
















TABLE OF CONTENTS

page

A C K N O W L E D G M E N T S ................................................................................................. iii

LIST OF TABLES .............................................. viii

L IST O F F IG U R E S .................................................... ............................................... ix

ABSTRACT ................................................. ...... .............. xiv

CHAPTER

1 INTRODUCTION ................................ ........... ........................ ..... 1

2 B A C K G R O U N D ............................................................................ .. ... ... .............. 7

Stent D esigns, M materials and C oatings...................................................... ...............9...
Polym er C oated Stents ..... ...................................................... .... .. ....... ....... 11
N on-Polym eric C oatings ................................................................... .............. 16
R radioactive C oatings ... ... ......................................... ....................... . . ..... ..... 17
Pulsed L aser A blation of Polym ers.......................................................... .............. 18
General Characteristics of Laser Ablation. ............... ... ................20
Pulsed Laser Ablation for Polymer Thin Film Deposition.....................................28
Surface Modification by Laser Ablation.................................................................40
Silicone F ilm s and C oatings......................................... ........................ ............... 43


3 PULSED LASER ABLATION DEPOSITION OF SILICONE............................... 46

In tro d u ctio n ................................................................................................................ .. 4 6
M materials .................................................................................................... . .......... 4 6
Methods ......................................................... ............... 47
C h aracterizatio n ........................................................................................................... 4 9
R e su lts ............... ... ................................................................................... . ........ .. 5 1
E effect of E energy D ensity.......................................... ........................ ............... 5 1
PL A D Film Stability T testing ..................................... ...................... ................ 72
E effect of L aser Pulse N um ber............................................................. ............... 72
Depositions under High Vacuum.................................................................... 78
D isc u s sio n ................................................................................................................. ... 7 9
S u m m ary ...................................................................................................... ........ .. 8 8









4 NANOINDENTION AND NANOSCRATCH MEASUREMENTS ON PLAD
F IL M S .......................................................................................................................... 9 1

In tro d u c tio n .................................................................................................................. 9 1
T h e o ry ......................................................................................................... . ........ .. 9 2
Nanoindentation....................... .. ........... ..................................... 92
N an o scratch .............................................................................................................. 9 6
B ack g ro u n d ............................................................................................................... 10 0
M ate rials ............................................................................................................. . .. 1 0 3
M eth o d s............................................................................................................... .. 1 0 3
R e su lts. ...................................................................................................... . ........... 10 5
D isc u ssio n ............................................................................................................... .. 12 2
S u m m a ry ................................................................................................................ . .. 12 4


5 EMPIRICAL MODEL FOR ABLATION OF SILICONE AT 248nm.................. 126

In tro d u ctio n ............................................................................................................... 12 6
Results and Discussion.... ................................................. .............. 130
S u m m a ry ................................................................................................................ . .. 14 8


6 FUNCTIONALIZATION OF STAINLESS-STEEL BY A COMBINED
PLASMA/GAMMA PROCESS ...................................................... .............. 149

In tro d u ctio n ............................................................................................................... 14 9
P lasm a P olym erization................................................................. .................... 149
Surface Modification of Stainless steel........................................ 151
M ateria ls ..................................................................................................................... 1 5 3
M eth o d s ............... ................. ........ .......................................................... .......... 15 3
Plasma Polymerization of Hexane ............................................................... 153
Gamma Radiation Initiated Grafting of Methacrylic Acid ...............................155
Chemisorption with Polyethylene imine (PEI) .......................... .................. 156
Bovine Serum Albumin (BSA) Grafting...... ... ....................................... 156
C h aracterizatio n ......................................................................................................... 15 8
R results ................. ..... .. ............................................................................ . . . 158
Surface Characterization of Hexane Plasma Polymer.................. ................. 158
Surface Characterization of Functionalized Steel Surfaces. ............................ 166
D isc u ssio n ............................................................................................................... .. 1 8 0
S u m m a ry ................................................................................................................ . .. 1 9 0


7 CONCLUSIONS AND FUTURE WORK ...... ......................... 191

Pulsed Laser Ablation Deposition....................................................................... 191
Functionalization of Stainless-steel by Plasma/Gamma Processes......................... 194









APPENDIX

DESIGN, SETUP AND CALIBRATION OF A STREAMING POTENTIAL CELL. 195

In tro d u ctio n ................. ............................................................................................. 19 5
Streaming Cell Design and Setup................................................................... 196
Flow Analysis of the Streaming Potential Cell...... .... ................................... 202
Stream ing C ell C alibration...................................... ........................ ................ 203


LIST OF REFEREN CE S .................... ............................................................... 209

BIO GRAPH ICAL SK ETCH ...................................... ........................ ............... 221
















LIST OF TABLES


Table Page

2.1 Stent specifications and coatings for major designs and manufacturers............... 10

2.2 Operating parameters of different excimer lasers................................ .............. 19

3.1 FTIR absorbance bands observed for PLAD films along with literature
a ssig n m en ts. ............................................................................................................. 5 2

3.2 Elemental compositions for PLAD films deposited at different laser pulse
n u m b er..................................................................................................... . ........ .. 7 7

3.3 Contact angles and XPS elemental % for PLAD films deposited under high
v a cu u m .................................................................................................................... 7 8

4.1 Literature results for modulus values from nanoindentation of different polymers. 102

6.1 XPS analysis of plasma polymers synthesized at different input power............... 160

6.2 Average contact angles for hexane plasma polymers. ................... ...... ........... 161

6.3 Deposition rates for hexane plasma polymers at different input power................ 165

6.4 XPS results for functionalized steel surfaces....... ... ................................... 169

6.5 Average contact angle values for different steel surfaces................................. 178

6.6 Zeta potentials of different steel surfaces........................................ 178

6.7 Effect of storage time on elemental composition of a hexane plasma polymer
fi lm ...................................................................................................... . . .......... 18 1

A. 1 Physical parameters along with units used in the zeta potential calculation........ 201















LIST OF FIGURES


Figure Page

1.1 Palmaz-Schatz stent prior to and after expansion by a balloon catheter. Ref. [2] ..... 1

2.1 Schematic diagrams of several major stent designs in use today..............................9

2.2 Schematic representation of the laser ablation process.....................21

2.3 Characteristic graph of ablation depth/pulse vs. Energy density for polymers..........22

3.1 Curing reaction for preparing silicone elastomer targets.............. ................47

3.2 Schematic diagram of the vacuum chamber setup for the PLAD process..............50

3.3 Reflectance absorption FTIR spectra of PLAD films deposited at: (a) 100, (b) 250
and (c) 510 mJ/cm2....................... .... .......... ........................ 53

3.4 Attenuated Total Reflectance -FTIR spectrum of a silicone target........................55

3.5 Expanded FTIR spectra of PLAD films at different energy densities: (a) 100-, (b)
250- and (c) 510 m J/cm 2. ........... .. ............. ............................................. 56

3.6 Elemental composition of PLAD films as a function of energy density................. 58

3.7 High resolution Elemental XPS scans of PLAD films deposited at different
en erg y d en site s.......................................................................................................... 6 1

3.8 High resolution elemental XPS scans for a silicone target...................................63

3.9 Si2p peak location from XPS for PLAD films...................................... ................ 65

3.10 Contact angles of PLAD films as a function of energy density..............................66

3.11 Deposition rate as a function of energy density for PLAD films...............................68

3.12 Effect of pressure on deposition rate of PLAD films ...................... ..................... 69

3.13 Scanning Electron Micrographs of PLAD films deposited at different enery
d e n site s ...................................................................................................................... 7 0









3.14 XPS elemental % for PLAD films, pre- and post chloroform wash.......................73

3.15 Micrographs of PLAD films pre- and post chloroform immersion........................75

3.16 Schematic diagram of the different regimes for silicone ablation at 248 nm ..........87

3.17 Cross-sectional SEM of a PLAD film coated steel substrate. The film was
deposited at 400 m J/cm 2 .................................................................. ................ 89

3.18 Electron Micrograph of a coated stent after expansion .........................................90

4.1 Load-displacement curve in a nanoindentation experiment with points of interest...93

4.2 Profile of an indention. The top picture shows the tip in the surface at maximum
load. The lower picture is at the end of an indentation cycle. Adapted from ref
[10 4 ] ........................................................................................................ . ....... .. 9 4

4.3 Schematic diagram of the stages in a nanoscratch test. A ramp force profile is
sh o w n ................................................................................................................... . . 9 8

4.4 AFM scans of a PLAD film surface deposited at 100 mJ/cm2 (a) pre scratch, (b)
post 500 [tN 5 |tm ram p scratch ........................................................... ................ 99

4.5 Force-displacement curves for the sample shown in fig. 4.4. Adhesive Failure is
seen at 100 nm which corresponds to the film thickness. (Shown by the solid
line) ......................................................................................... .......... 100

4.6 Normal force and lateral displacement profiles as a function of time used for the
nanoscratch testing. ..................................... ........ ...................104

4.7 Modulus vs. contact depth for a 100 |tm PET film......................... ...................106

4.8 Force-displacement curves for a 100 |tm PET film........................ ...................106

4.8 Modulus vs. contact depth for a 100 |tm cross-linked silicone film ........................107

4.9 Force vs. displacement curves for a 100 |tm silicone film. (Compare scale with
fi g u re 4 .5 )............................................................................................................... ... 1 0 7

4.10 Force-displacement curves for a PLAD film deposited at 100 mJ/cm2.................. 109

4.11 Force-displacement curves for a PLAD film deposited at 250 mJ/cm2. ...................109

4.12 Force-displacement curves for a PLAD film deposited at 500 mJ/cm2. ...................110

4.13 Modulus vs. contact depth for PLAD films synthesized at different energy
d en sities.................................................................................................... . ........ .. 1 10

4.14 Time response of a PLAD film deposited at 100 mJ/cm2....................................111









4.15 Average modulus values for PLAD films deposited at different energy densities. ... 112

4.16 AFM scan of a PLAD surface. Arrows indicate indent sites...............................113

4.17 Normal displacement of a PLAD film deposited at 100 mJ/cm2, subjected to a
ramp force scratch of 500 [[N. The graph shows four scratch profiles...................115

4.18 Time profiles of Normal force and lateral force/normal force for a PLAD film
synthesized at 100 mJ/cnm2. A yield point is seen on the normal force profile ..........116

4.19 Atomic Force Microscope (AFM) scan of a PLAD film surface deposited at 100
mJ/cm2, with a 500 [[N, 5 |tm scratch. A 250 [[N ramp scratch is seen to the left
of the main scratch. ............ .. ................. .... ....... ..... ............... 116

4.20 Normal displacement of a PLAD film deposited at 250 mJ/cm2, subjected to a
ramp force scratch of 500 [[N. The graph shows four scratch profiles which show
a variety of responses. ............................ ........................................... 117

4.21 Time profiles of normal force and lateral force/normal force for a PLAD film
synthesized at 250 m J/cm2. ......................................................... 118

4.22 Atomic Force Microscope (AFM) scan of a PLAD film surface deposited at 250
mJ/cm2, with a 500 [[N, 5 |tm scratch. ...... ...........................1... 18

4.23 Normal displacement of a PLAD film deposited at 500 mJ/cm2, subjected to a
ramp force scratch of 500 [[N. The graph shows four scratch profiles which show
a variety of responses with the arrows indicating points of film failure..................119

4.24 Time profiles of normal force and lateral force/normal force for a PLAD film
synthesized at 500 m J/cnm2. ......................................................... 120

4.25 Atomic Force Microscoe (AFM) scan of a PLAD film surface synthesized at 500
mJ/cm2, with a 500 [[N, 5[tm scratch. ...... ... ........................... 120

4.26 Normal displacement plots for a cross-linked silicone film subjected to a 500 |[N,
5 |tm scratch test......................................................................................................... 12 1

4.27 Normal force and lateral force/normal force vs. time profile for a single scratch
test for a cross-linked silicone film ...... ..... ...... ....................... 122

5.1 Scanning Electron Micrographs of silicone target surfaces ablated at 100 mJ/cm2
(5kV x2000) (a): 100 pulses, (b): 5000 pulses....... ... ...................................... 131

5.2 Scanning Electron Micrographs of silicone target surfaces ablated at 250 mJ/cm2
(5kV, x500) (a) 100 pulses, (b) 2000 pulses, (c) 5000 pulses...............................132

5.3 Scanning Electron Micrographs of silicone target surfaces ablated at 250 mJ/cm2
(5kV, x2000) (a) 100 pulses, (b) 2000 pulses, (c) 5000 pulses................................133









5.4 Scanning Electron Micrographs of silicone target surfaces ablated at 420 mJ/cm2
(5kV, x500) (a) 100 pulses, (b) 2000 pulses, (c) 5000 pulses). ................................134

5.5 Scanning Electron Micrographs of silicone target surfaces ablated at 420 mJ/cm2
(5kV, x2000) (a) 100 pulses, (b) 2000 pulses, (c) 5000 pulses................................135

5.6 Scanning Electron Micrograph (5kV, x500) of a silicone target showing the edge
of the ablated area. Arrows indicate the presence of microcracks...........................137

5.7 Scanning Electron Micrograph of an ablated region on a silicone target showing
the resultant etch-pit. The arrows indicate the edges of the pit. (Scale bar = 2
m m ). ................................................................................................. ........ . . ....... 1 3 9

5.8 Scanning Electron Micrograph (5 kV, xlOO) showing the cross-section of an etch
pit. The lines indicate the depth of the pit....... ... ......................................... 140

5.9 Plot of the ablation rate vs. incident fluence. ...... ... ........................................ 143

5.10 Data from figure 5.9 plotted according to equation 8. The line represents the best
fi t. ....................................................................................................... . . ......... 14 6

6.1 Chemical structure of Poly(ethylene imine)...... .... ....................................... 153

6.2. Schematic diagram of the plasma reactor and accompanying instrumentation .........155

6.3. Schematic diagram of the surface functionalization process for stainless steel.........157

6.4 Reflectance absorption IR spectra of hexane plasma polymers synthesized at
different input power ... ................................................................ 159

6.5. Low resolution XPS survey spectra of steel surfaces (a) unmodified and (b)
hexane plasm a polym er coated. ......................................................... 162

6.6. High resolution XPS spectra of unmodified steel and a hexane plasma polymer
coated steel surface (50W power input) (a) CIs and (b) Ols. ................163

6.7. High resolution scans of hexane plasma polymers synthesized at 25-, 50- and 100
W (a) Cls and (b) Ols ........... ........... ..... ... ............... 164

6.8. Scanning electron micrographs of steel surfaces coated with hexane plasma
polymers synthesized at different input power. ......................................167

6.9. Reflectance absorption IR spectra of steel surfaces after successive modifications
with methacrylic acid, polyethylene imine and bovine serum albumin.................. 170

6.10. Low resolution XPS survey spectra of functionalized steel surfaces.......................171

6.11. High resolution XPS spectra of a methacrylic acid modified steel surface, (a) C Is
and (b) O ls .............................................................................. ............... .................. 174









6.12. High resolution XPS spectra of polyethylene imine and bovine serum albumin
m odified steel surfaces.. ................................................................................ 175

6.13. High resolution XPS spectrum of the S2p peak for a bovine serum albumin
m odified steel surface... .................................................................... .............. 177

6.14. Scanning electron micrographs of methacrylic acid plasma/gamma modified steel
surfaces at (a) x500 and (b) x2000 ....... ........ ....... ...................... 182

6.15. Scanning electron micrographs of polyethylene imine chemisorbed steel surfaces
at (a) x500 and (b) x2000 ............. ............................................................. 183

6.16. Scanning electron micrographs of bovine serum albumin modified steel surfaces
at (a) x500 and (b) x2000 ............. ............................................................. 184

A. 1. Top block of the streaming cell showing the associated features ........................... 197

A .2. B bottom block of stream ing cell. ...... .......... ........... ......... ............198

A.3. Schematic of the experimental setup for measuring streaming potential.................199

A.4. Cell block, sample and gasket assembly for the streaming potential cell................200

A.5. Schematic representation of the parallel plate flow system and the velocity profile. 202

A.6. Volumetric flow rate as a function of driving pressure for the streaming................204

A.7. Reynold's number as calculated from equation 2 for the streaming potential cell
indicating that the flow is laminar under all experimental conditions .....................205

A.8. Entrance length as a function of Re as calculated from equation 3..........................206

A.9. Zeta potential vs. pH trend for fused silica substrates..................... ................... 208
















Abstract of Dissertation Presented to the Graduate School
of the University of Florida in Partial Fulfillment of the
Requirements for the Degree of Doctor of Philosophy

SURFACE MODIFICATION OF BIOMATERIALS BY PULSED LASER ABLATION
DEPOSITION AND PLASMA/GAMMA POLYMERIZATION

By

Kaustubh R. Rau

December, 2000
Chairman: Eugene P. Goldberg
Major Department: Materials Science and Engineering

Surface modification of stainless-steel was carried out by two different methods: pulsed

laser ablation deposition (PLAD) and a combined plasma/gamma process. A potential

application was the surface modification of endovascular stents to enhance biocompatibility.

The pulsed laser ablation deposition process, had not been previously reported for modifying

stents and represented a unique and potentially important method for surface modification of

biomaterials. Polydimethylsiloxane (PDMS) elatomer was studied using the PLAD

technique. Cross-linked PDMS was deemed important because of its general use for

biomedical implants and devices as well as in other fields. Furthermore, PDMS deposition

using PLAD had not been previously studied and any information gained on its ablation

characteristics could be important scientifically and technologically.

The studies reported here showed that the deposited silicone film properties had a

dependence on the laser energy density incident on the target. Smooth, hydrophobic,

silicone-like films were deposited at low energy densities (100-150 mJ/cm2). At high energy









densities (> 200 mJ/cm2), the films had an higher oxygen content than PDMS, were

hydrophilic and tended to show a more particulate morphology. It was also determined that:

1) the deposited films were stable and extremely adherent to the substrate, 2) silicone

deposition exhibited an 'incubation effect' which led to the film properties changing with

laser pulse number and 3) films deposited under high vacuum were similar to films deposited

at low vacuum levels.

The mechanical properties of the PLAD films were determined by nanomechanical

measurements which are based on the Atomic Force Microscope (AFM). From these

measurements, it was possible to determine the modulus of the films and also study their

scratch resistance. Such measurement techniques represent a significant advance over

current state-of-the-art thin film characterization methods.

An empirical model for ablation was developed for the 248 nm laser irradiation of

silicone. The model demonstrated a good fit to the experimental data and showed that

silicone underwent ablation by a thermal mechanism.

In addition to PLAD studies, functionalization of stainless steel was carried out by a

combined plasma/gamma method involving deposition of a hexane plasma polymer by RF

plasma polymerization, followed by gamma radiation graft polymerization of methacrylic

acid. The hydrograft modified surfaces were further modified by chemisorption reactions

with poly(ethylene imine) to produce amine-rich surfaces. Bovine serum albumin was then

bound via amino groups using glutaraldehyde coupling. A streaming potential cell was also

built and used to measure the zeta potential of these ionic surfaces.














CHAPTER 1
INTRODUCTION

Since their introduction in the early 1980s, metallic endovascular stents have

revolutionized the treatment of coronary heart disease [1]. When used in conjunction

with balloon angioplasty, stenting has a greater than 80% chance of success. The

procedure is minimally invasive leading to less patient trauma, faster recovery times and

less expense. It involves introducing a stent--an expandable metallic mesh-like tube--

mounted on a balloon catheter to the site of vascular disease. The catheter is then blown

up, expanding the stent and propping open the blocked blood vessel. Figure 1.1 shows

the stent prior to and after expansion.









Figure 1.1 Palmaz-Schatz stent prior to and after expansion by a balloon catheter. Ref.
[2].



Although significant advances have been made in stent designs, processing and

finish, several problems still persist [3]. In the short term, thrombosis or formation of

a blood clot is a major concern. The thrombogenicity is caused by the vascular

response provoked by the stent's metallic surface. In the long term, cellular ingrowth

from underlying smooth muscle cells and other associated events leading to arterial

thickening, a phenomenon known as restenosis, remains the major cause for repeat









procedures. This cellular growth is the natural remodeling of the arterial wall in

response to the vascular injury elicited by the angioplasty. The restenosis response is

quite complex and involves several different growth factors, proteins and cells.

Although the use of stents has led to reduced restenosis rates in balloon angioplasty

procedures, the figure is still around 15-30% within six months of stenting.

Surface modification of stents with polymer coatings has been seen as one possible

solution to these problems. A polymer coated stent could present a less thrombogenic

surface to blood and invoke a benign response. Also, with the use of anti-proliferative

and anti-inflammatory drugs loaded into polymer coatings it may be possible to

mitigate the restenosis response.

Surface modification has been a widely studied technique to improve the

performance of biomaterials in a physiological environment [4]. It offers the

opportunity to tailor the surface properties for specific applications without

compromising the bulk properties of an implant. Surface modification of metals is

however challenging. Commonly used techniques like chemical grafting or gamma-

radiation initiated grafting are not readily successful due to interfacial bonding

problems. High-energy techniques like vapor deposition and ion implantation cannot be

applied to produce polymer coatings. The two methods most widely studied are

polymer solution coating and radio-frequency plasma treatments.

Polymer solution coating, spray- or dip-coating as it is widely known, is a simple

technique wherein a polymer solution is applied to a substrate and dried to produce a

coating. This technique, although easy, has disadvantages associated with it. Several

processing steps are needed to build up the coating. The coatings produced can have









non-uniform thickness and coverage. The adhesion of the coating to the implant can

also be tenuous unless coupling agents are used and the coupling chemistry needs to be

tailored for each metal surface.

Radio-frequency plasma treatment is a well-developed technique for modifying

metal, ceramic and polymer surfaces. It can produce coatings which are not otherwise

obtained by conventional chemical methods. An initiating species is not required and

the process usually produces crosslinked polymer films. However, growth rates can be

slow and process optimization remains a challenge to date.

Implants such as stents also offer unique problems due to their small size and

complex geometry. Apart from developing a coating, assessing the quality and

performance of the coating is a challenge. Current methods used for surface

modification of stents have usually involved solution coating methods. Apart from the

problems noted above with such techniques, polymer webbing between the metal

struts may also occur. Surface modification using plasma treatments have been

applied only recently to stents and their efficacy is not yet known. Therefore, a

method which can apply a coating in a single step, reproducibly and with improved

adhesion, would prove beneficial and would also be a significant advance over the

current state-of-the-art.

The research reported here investigated the use of pulsed laser ablation deposition

(PLAD) of polymers for developing coatings on metal surfaces such as stents. It was

hypothesized that the PLAD method could produce coatings by complete transfer of

polymer from the target to the steel surface Thus the desirable properties of the

polymer would be reproduced in the coating. The process would use commercially









available polymers for coatings thus obviating the need for monomers, or problematic

reagents and conditions. This would be an advantage as compared to solution coating

and plasma processes and might provide better control over the coating properties.

Stronger film adhesion was also likely due to the energetic nature of the coating

process. The PLAD method had not previously been reported for the surface

modification of biomedical implants and represented a unique and potentially

important new approach for the surface modification of biomaterials.

The polymer primarily invesitgated here to produce coatings by the PLAD method

was polydimethyl siloxane (PDMS), commonly known as silicone. The choice of

material was dictated by two reasons. First, the good haemopatibility of silicone has

been proven in several studies and this made it a good choice as a material for a blood-

contacting stent coating. Crosslinked PDMS can also serve as a drug delivery depot

which is a significant advantage. Additionally, the pulsed laser ablation of cross-

linked silicone elastomer had not previously been reported and the results would add

to our scientific understanding of the polymer ablation and deposition process.

Chapter 2 provides the background and literature review on the topic of stent

coatings. The different coatings studied to date are discussed therein. An introduction

to the subject of pulsed laser ablation of polymers is also presented. Among the

various aspects discussed are the nature and characteristics of the process and

applications to surface modification. Further, current methods to produce silicone

coatings and thin films are briefly mentioned.

Chapter 3 presents the results of the pulsed laser ablation deposition of silicone on

stainless-steel. The variation in surface properties of the deposited silicone films as a









function of the incident energy density is examined in detail. Other aspects examined

include the deposition rate, effect of the laser pulse number on the film chemistry,

coating stability after solvent immersion and effect of high vacuum during deposition.

The results of nano-indentation and nano-scratch experiments on PLAD silicone

films using methods based on the Atomic Force Microscope (AFM) are presented in

Chapter 4. These experiments were carried out to determinethe mechanical strength

and adhesion of the coatings. Nanomechanical testing is a new but promising

technique for studying the mechanical properties of thin films. The experiments

reported here provide a basis for further investigation of the method for characterizing

polymer films and coatings.

Chapter 5 is devoted to the development of a model to explain the ablation

characteristics of silicone. This empirical model is applied to experimental results and

is seen to give a good fit. Since research on pulsed laser ablation of silicones is

limited, this chapter serves to extend current theories of polymer ablation to silicon

based polymers.

Immobilization of bioactive molecules is a further important aspect of biomaterial

surface modification. The attachment of proteins, peptide sequences and drugs to

biomaterial surfaces can lead to an improved outcomes by achieving increased

localized bioactivity. However, for stents the challenge has been to attach these

biomolecules to metallic implant surfaces. Thus, the second part of this research

investigated the surface modification of stainless steel to produce functionalized

surfaces. This was done by using a novel plasma/gamma radiation graft method









developed in this laboratory. Biofunctionalization of steel surfaces was achieved by

this method by coupling of bovine serum albumin to a modified steel surface.

Chapter 6 presents the results of the surface modification of stainless steel by the

plasma/gamma method as well as results for the coupling of bovine serum albumin to

the functionalized steel surfaces. The charge on the functionalized, ionic surfaces was

measured by a streaming potential cell which was built for this purpose. The design,

setup and operation of this cell are given in the appendix.

Chapter 7 presents the conclusions which have been reached from this research

and provides recommendations for future research directions.














CHAPTER 2
BACKGROUND

Percutaneous transluminal coronary angioplasty (PTCA) has proven to be a

successful technique for treating coronary heart disease. The procedure is minimally

invasive and allows patients to regain active function earlier. However a large number

of these procedures suffer from the problem of restenosis, a phenomenon in which

smooth muscle cells from underlying cell layers proliferate causing gradual thickening

of the vessel wall and ultimately occlusion. Stents have been shown to address the

problem of restenosis in PTCA procedures successfully. Two recent trials reported that

stents reduced the need for revascularization due to restenosis by approximately 30%

[5],[6]. Other studies have shown that restenosis rates were reduced from up to 50%

for PTCA to 15-30% with stenting in the first 6 months after surgery. Their success has

led to the ever increasing use of stenting in PTCA procedures. It was estimated that in

1998 more than 500,000 stent procedures were carried out worldwide on coronary

blood vessels [7]. It is expected that stents will be used in 56% of all PTCA procedures

worldwide by the year 2001.

However the use of stents in these procedures also has its problems. One of the

main problems has been short-term thrombosis or clotting [8],[9]. This is due to the

adverse response elicited by the stent surface on implantation. Patients need a longer

hospital stay and a rigorous anticoagulant regimen to prevent the acute thrombosis of

the stent. Currently all stents on the market are made from stainless-steel, tantalum or a

nickel-titanium alloy (i.e. nitinol). These metals are inherently thrombogenic and









induce platelet adhesion and clot formation upon implantation. With current

thrombosis rates around 3-5%, up to 20,000 people could experience life-threatening

events each year[10].

Research efforts to improve the blood contact properties of stents have

concentrated on modifying the outer layer of the stent surface rather than changing the

stent material itself. This lets the mechanical properties of the stent remain unaltered

while presenting a favorable surface to the blood vessel. Efforts have also been

directed at reducing the current restenosis rates to even lower levels. Thus the ideal

coating has been viewed to be haemocompatible, thereby not inducing stent thrombosis

and non-proliferative thereby reducing restenosis by discouraging cell growth. To this

end, stents coated with biocompatible polymers and anticoagulants like heparin have

been reported in the literature [11]. The polymer coatings have also been viewed as

depots for drug delivery. Thus a number of reports on stent coatings which deliver

anti-inflammatory and anti-proliferative agents to the site of vascular injury have also

been published.

Currently a number of research groups and stent manufacturers are evaluating

the performance of several different kinds of coatings in reducing thrombosis and

restenosis. As the use of stents in other anatomic frontiers like carotids, renals and

small vessel vascular peripheral disease also increases, producing a stable,

biocompatible, drug-eluting coating would be of increasing importance. Indeed, it was

recently stated that, "... the potential for application of antiproliferative coatings

remains provocative and largely unexplored" [3].









Stent Designs, Materials and Coatings

The 'stentomania' of the last decade has influenced manufacturers to introduce

a number of new designs. Stent designs usually consist of metallic tubes with a number

of slots cut in by lasers, or coil designs which consist of a single wire entwined to form

a cylindrical mesh, reinforced at several points [12],[13]. The various mesh designs

give the stent different flexibility and handling characteristics and play a role in the

thrombosis and restenosis rate [14]. They are usually made of stainless steel although

tantalum and nickel-titanium alloy based stents have also been introduced. Table 2.1

gives an overview of some stents with their designs and the coatings being applied.

Figure 2.1 shows schematic diagrams of major stent designs. All stent designs require

a compromise between scaffolding properties and flexibility to negotiate tortuous

vessels. The radio-opacity of the stents also plays a role in stent choice. Thus

manufacturers have striven to introduce new designs to improve the stent handling

characteristics and radio-opacity. The choice of stent is largely determined by the

surgeon's experience and choice. The Palmaz-Schatz stents have been the most

popular choice to date. However failure to innovate has led to a decline in this stent's

popularity. Among the popular choices nowadays are Guidant's Multilink and

Medtronic's Wiktor stents.













Figure 2.1 Schematic diagrams of several major stent designs in use today.


















Table 2.1 Stent specifications and coatings for major designs and manufacturers.
Stent/ Material Design Strut Diameter, Length, First Clinical Coating/Market
Manufacturer Thicknes mm mm Implantation Status
smm
Wallstent/Schneider Inner core Wire mesh 0.08-0.1 3.5-6.0 12-42 1986
platinum; 1991 (less
outer core shortening)
cobalt-based
alloy
Palmaz-Schatz/ 316L Slotted tube 0.07 3.0-4.0 15 1988 Heparin/FDA approved
Johnson & Johnson stainless 1999
steel
ACS-Multilink/ 316L Corrugated 0.05 3.0-3.5 15 1993 -
Guidant stainless links
steel connected by
multiple links
Gianturco-Roubin/Cook 316L Incomplete 0.127 2.5-4.0 20-40 1989 (GR-I) Antiplatelet coating
Inc. stainless coil clam- 1995 (GR-II) (Reopro)
steel shell loop Antiproliferative
coating (Taxol)
Wiktor/Medtronic Tantalum Single wire 0.127 3.0-4.5 16 1991 Heparin/on market
semi-helical
coil
AVE-Microstent 316L Sinusoidal 0.13 2.5-4.0 6-36 1994-95
stainless ring
steel
divYsio/Biocompatibles 316L Interlocking 0.15 2.75-4.0 15 1995 Phosphorylcholine
Cardiovascular stainless arrowhead
steel
NIRMedinol/Boston Stainless Expandable 0.1 2.0-5.0 9-32 1995 Gold
Scientific steel uniform
cellular mesh
Adapted from Handbook of Coronary Stents, ed Serruys, PW, Kutryk, MJB, Published by Martin Dunitz Ltd.,
London,1997 and Ozaki, Y., et al., Progress in Cardiovascular Disease 39, 2, 129-140, 1996.









Polymer Coated Stents

Research on polymer coated stents was undertaken very early in the evolution of

stents. The advantages of a polymer coated stent were seen to be biocompatibility and

also its ability to function as a drug reservoir. A majority of the coatings have been

produced by dip-coating. However data on the coating thickness, surface analysis and

stability is limited in most cases. A review of stent coatings and their performance is

given below.

De Scheerder et al. have reported on the performance of slotted tube steel stents

coated with an amphiphilic polyurethane (PU) [15]. The polyurethane was prepared from

an amphiphilic polyester, diphenylmethane-4-4'-diisocyanate and butane diol as a chain

extender. It was shown that the polyurethane coating decreased acute thrombic occlusion

in a porcine model. The PU coated stents led to a significantly smaller minimum stented

lumen diameter after 6 weeks as compared to controls. Polyorganophosphazene (PP)

coated stents were also evaluated in a similar porcine model. The polyphosphazene was

selected due to its biogradibility under physiologic conditions. The PP coated stents gave

pronounced neointimal proliferation. The degradation products also elicited a severe

inflammation response resembling a foreign-body reaction. The authors concluded that

PU was promising as a stent coating while PP was not.

Lincoffet al. have studied a drug eluting biodegradable polymer coating in a

porcine model [16]. A tantalum wire mesh stent coated with a 20-25 |tm layer of

dexamethasone suspended in a binder namely, poly-L-lactic acid (Mw = 321,000 g/mol),

was evaluated. Although the dexamethasone did not exhibit reduced neointimal

proliferation it was demonstrated that a polymer coated stent could prove to be a means









of effective intravascular drug delivery. The reduced activity of dexamethasone could

have been due to the inflammatory response elicited by the degradation products of the

poly-L-lactic acid, an aspect which was not examined in the research.

Schwartz et al. reported on modifying stents with a fibrin layer by dripping

fibrinogen onto it [17]. Subsequent polymerization of the fibrin layer led to the fibrin

mass completely encasing the stents. Fibrin was selected since it is a native polymer

which deposits at the site of vascular injury. It was hypothesized that a layer of thrombus

would form at the site of the injury and fool the body's own response to this event and

limit its thrombotic response. The stents were assessed in a porcine coronary model in

contrast with polyurethane coated stents. PU-coated stents showed a foreign-body

inflammation response with multinucleated cells being present on the surface. In contrast

the fibrin-coated stents reduced restenosis rates.

In an effort to mimic the cell-bilayer, phosphorylcholine coatings have been tried

by several groups. Zheng et al. reported on clinical experience with a phosphorylcholine

coated stent [18]. The coated stents were supplied by Biocompatibles Ltd., U.K., and

were produced by a proprietary process. The authors found a restenosis rate of 6.1% at 6-

month follow-up in 224 patients. The low restenosis rates indicated that the coatings

were safe and efficacious in the treatment of coronary lesions. However this study did

not compare uncoated stents vs. coated and any conclusions about the performance of the

coating are moot. In a more recent study, Whelan et al. assessed the biocompatibility of

phosphorylcholine coated stents in normal porcine arteries [19]. The coating applied to a

divYsio atent consisted of a copolymer of methacryloylphosphorylcholine and lauryl

methacrylate. The stents were dip-coated from a solution of the polymer in ethanol to









give a coating approximately 50 nm in thickness. In comparing un-coated and coated

stents no difference in intimal thickness at 4 and 12 weeks was seen. It was also seen that

the phosphorylcholine coating did not elicit an adverse inflammatory response. As a

measure of coating stability the authors examined sections of explanted stents, 12 weeks

post-implantation by optical microscopy. Although the coating could not be directly seen

under the microscope, the authors concluded that the coating was still present based on

the staining characteristics of these stents. However, given the tenuous nature of the

coating (- 50 nm) it is highly unlikely that the coating was stable at 12 weeks after

implantation. Thus the similar responses of the uncoated and coated stents could be due

to the fact that the coatings failed exposing the stent surface.

Van der Giessen et al. tested the response of different polymers by deploying

them as strips on the circumferential surface of coil wire stents in a porcine coronary

model [20]. Five biodegradable polymers namely, polylactic acid-polyglycolic acid

(PLGA), polycaprolactone (PCL), polyhydroxybutyrate valerate (PHBV), polyorthoester

(POE) and polyethyleneoxide/polybutylene terephthalate (PEO/PBTP), were tested.

Three non-biodegrable polymers namely, polyurethane (PU), silicone and polyethylene

terephthalate (PET), were also tested. Four weeks after implantation the stent patency

was assessed by angiography followed by microscopic examination of the coronary

arteries. It was found that all polymers evoked a significant inflammatory and

proliferative response. The response was common to biodegradable and non-

biodegradable polymers In all cases the inflammatory response consisted of a chronic

inflammatory reaction with an acute component and a persistent foreign-body reaction.

The authors speculate that the inflammation response may have been aggravated by the









release of the degradation by-products of the polymers. It should be mentioned that the

coated stents were not sterilized prior to implantation which could have contributed to

their severe response.

Two studies have examined the efficacy of heparin coated stents in reducing the

thrombogenic response. De Scheerder et al. studied the thrombogenicity of heparin

coated and uncoated stents in a rat arteriovenous model [21]. They found that total clot

weight at 30-minute follow-up was significantly lower in the heparin-coated stents

compared with bare stents. Subsequently heparin-coated and uncoated stents were

implanted in the right coronary artery of 20 domestic pigs. It was shown that heparin-

coated stents had no significant difference in luminal area or neointimal hyperplasia

compared to uncoated stents. The authors concluded that heparin reduces the

thrombogenicity but does not influence neointimal hyperplasia.

HArdhammar et al. studied the thrombogenic response of heparin coated stents in

a porcine artery model [22]. The coating was applied by end point coupling of heparin

molecules to an underlying polymer matrix. The technique used was a variation of the

Carmeda process in which a polyamine layer was deposited on the stent surface, followed

by a dextran sulfate layer, followed again by a polyamine layer. Finally the functional

amino groups were covalently coupled to aldehyde groups of partially degraded heparin

molecules. The activity of the bound heparin was measured by its ability to bind

antithrombin-III with high affinity. Since it was determined that heat or ethylene oxide

sterilization considerably reduced the antithrombin-III binding activity the stents were

coated under clean room conditions but not sterilized. It was shown that the high-activity









heparin coating eliminated subacute thrombosis in porcine coronary arteries. However

no reduction in the neointimal layer thickness was observed due to the heparin coating.

Aggarwal et al. have studied the response of polymer coated stents eluting platelet

glycoprotein IIb/IIIa receptor antibody in a rabbit model [23]. The authors reasoned that

using a potent antiplatelet agent eluting from the stent would passivate adherent platelets

and reduce thrombus formation after stent deployment. The coating produced by dip-

coating consisted of a cellulose polymer applied as a 10% solution with acetone as the

solvent. A 30 |tm thick coating was produced by this method. The polymer coated stents

were immersed in a radio-labeled antibody solution and removed after specified time

periods. The activity of the antibody was measured by the radioactivity associated with

it. These antibody loaded stents were implanted in rabbits. There was a >50% reduction

in platelet deposition 2 hours after stents eluting platelet GP IIb/IIIa antibody were

deployed, compared with control stents eluting either irrelevant antibody or no active

agents. Furthermore, platelet aggregation in the stent was strikingly inhibited by the

eluting GP IIb/IIIa antibody, as evidenced by the almost total abolition of cyclic blood

flow variation in these vessels. The authors demonstrated that the surface modification of

stents with potent antiplatelet agents can thus enhance the thromboresistance of stents.

It is seen from the above review that surface modification of stents has been done

mainly by the dip-coating method. Often data on the success of the surface modification

have been sparse or altogether absent. Thus the stability of the coatings cannot be

assessed over a long period. The dip-coating method can also present problems like

inadequate polymer-metal adhesion, non-uniform and very thin films. It also necessitates









the use of several processing steps. Problems with sterilization have also been noted for

some polymer coatings.



Non-Polymeric Coatings

Methods for coating stents using different techniques have also been reported. A

silicon-carbide coating of metallic stents has been developed. As a semiconductor SiC is

known to have good biocompatibility and it was reasoned that it would modulate the

surface charge to reduce thrombogenicity [24]. Therefore a SiC coating was applied by a

CVD process and coated stents were evaluated in a closed loop in vitro heparinized

whole human blood circulation model [25]. Uncoated and heparin coated stents were

also compared. The SiC-coated tantalum stent demonstrated a significantly lower GpIIIa

receptor-mediated platelet adhesion at the stent surface compared to all other stents.

Also, activated leukocytes demonstrated a significantly lower CD1 lb receptor-mediated

adhesion at the SiC-coated stent than at the stainless-steel stent. The SiC-coated stents

were also evaluated in treatment of coronary lesions for high risk patients. The stent

thrombosis rate in the high-risk group (3.6%) was not significantly different from that of

the low-risk patients (2.1%). The authors concluded that silicon carbide coating on

coronary stents may inhibit acute/subacute stent thrombosis even in patients at high risk.

Lahann et al. reported on the use of a solvent-less method to produce a polymer

coating on stent surfaces [26]. Poly(2-choroparaxylene) was deposited by a CVD process

on steel substrates. A sulfur-dioxide plasma treatment was then carried out on the

poly(2-choroparaxylene) coated metal surfaces to functionalize them. The SO2 plasma

attacked the chloride groups on the polymer to produce (SO3)- groups on the surface. The

polymer films demonstrated high mechanical stability in stent dilation experiments. The









coated surfaces showed a reduction in platelet adhesion as compared to uncoated

surfaces.



Radioactive Coatings

Another interesting development in stent coatings has been the use of radiation to

control restenosis [27]. Radiation selectively kills proliferating cells independent of the

stimulus for cell growth. Since neointimal hyperplasia is a major cause of restenosis, it

was reasoned that radiation therapy may reduce restenosis. Radioactive stents have the

potential to deliver an appropriate dose of radiation to the area of vascular injury thus

reducing restenosis while minimizing the total dose administered. In one method,

conventional Palmaz-Schatz stents were bombarded with ions in a cyclotron and

subsequently found to emit low-dose P3 and y radiation from radioisotopes Co55, 56, 57,

Mn52 and Fe55 with half-lives between 17.5 hours and 2.7 years [28]. A dose rate of 0.4

rads/h on the surface results in an integral dose rate of 0.18 Mrads after a period of 100

days. It was found that low-dose radioactive stents markedly inhibited neointimal

hyperplasia in rabbits. Endothelialization of the radioactive stents was found to be

delayed. Although the degree of neointimal hyperplasia was reduced, it was found,

counter-intuitively, that extracellular matrix production increased after stent implantation.

A second method used P32 -impregnated stents as a means of delivering local radiation.

P32 is a P emitter with a half-life of 14.3 months. In-vivo animal testing in a porcine

model showed inhibition of neointimal growth. Currently several clinical trials are

underway to assess the performance of radio-emitting stents [29].









Pulsed Laser Ablation of Polymers

In 1982 the use of UV pulsed laser radiation to etch poly(ethylene terephthalate)

(PET) was reported for the first time [30]. The attractive features of using pulsed laser

radiation were seen to be the degree of control which could be exercised on the etch

depth, the lack of detectable thermal damage adjacent to the etch site and the low energy

density levels involved. These observations were then confirmed for other polymers at

different UV wavelengths. The laser ablation of biological tissue was also reported in

1983 and the application of this technique for surgery was discussed [31]. In the ensuing

years a great deal of research was dedicated to studying the laser ablation of a few

polymers namely, polyimide (PI) [32], poly(methyl methacrylate) (PMMA) [33], poly(

ethylene terephthalate ) (PET) [34],[35] and polytetrafluoroethylene (PTFE) [36]. The

choice of these polymers is explained by their use in the semiconductor industry. Since

lasers had already made forays in the semiconductor industry, their application to etch

precise patterns on polymers was logical. Although the potential of using pulsed lasers

for thin film synthesis of inorganics had been well recognized early on, this aspect was

neglected for polymers till the early 1990s. Thus there is a large phenomenological

database on the ablation characteristics of certain polymers but relatively few published

studies on thin film synthesis using polymers. However, since an understanding of the

laser ablation process is intrinsic to its application in film synthesis, a short review of

pulsed laser ablation of polymers is presented here.

Pulsed laser ablation involves two basic elements that interact to give rise to laser

ablation. These are the radiation in the form of laser pulses and the polymer which

ablates on interacting with the pulse.









Excimer laser. The excimer laser consists of a laser cavity in which a long,

narrow discharge tube is set. It is filled with a halogen, a noble gas and a buffer gas to a

total pressure of 2 to 3 atm. When a high-voltage discharge (20-30 kV) is fired through

this mixture, laser radiation is produced in the UV range. Table 2.2 lists the combination

of halogen, noble gas and buffer used to produce various UV wavelengths.

As shown in the table below, a variety of wavelengths can be accessed depending on the

choice of excimer. These lasers can be operated with high outputs (1-5 Joules), with

short pulsewidths (femtosecond-nanosecond) at repetition rates between 1 to 200 Hz.

The beam dimensions can be well controlled with the use of filters and masks to give an

area in the mm2 to cm2 range of nearly uniform intensity.


Table 2.2. Operating parameters of different excimer lasers
Active medium F2 ArF KrCl KrF XeCl N2 XeF

Wavelength, nm 157 193 222 248 308 337 351

Helium, mbar 2650 1410 1720 2300 960 2570

Krypton, mbar 200 120

Argon, mbar 270

Neon, mbar 2760

He/F2(5%), mbar 50 120 80 120

He/HCl(5%), mbar 80 80

Nitrogen, mbar 40

Total pressure, 2700 1800 2000 2500 2900 1000 2700
mbar
Pulse width, ns 19 23 21 34 28 7 30

Adapted from company brochure, Lambda-Physik, Acton, MA, USA.









Polymer Targets. The second part of the process is the polymer on which the laser pulse

is incident. Polymers have UV absorption characteristics which are strongly related to

their structure. Polymers like PI and PET absorb strongly in the UV range of the

spectrum due to the presence of '7 bonds, while PE and PTFE are weak aborbers due to

the absence of chromophores. Depending on the absorption coefficient the incident UV

radiation will penetrate from a fraction of a micron to tens of microns before 95%

absorption (optical penetration depth). Absorption of a photon by an organic molecule

causes it to rise to an electronically excited state. The first excited state can absorb

another photon to give a doubly excited state (i.e. multiphoton excitation). Successive

absorption of two such photons by any one chromophore in a polymer will lead to an

excited state which may undergo ionization. The efficiency with which the absorption of

a single photon can lead to a chemical bond rupture is expressed in terms of a quantum

yield. Although quantum yields in the solid-state are far less than unity, secondary

reactions following primary bond rupture can lead to an increased yield [37]. Strong

absorption, although important is not a necessary condition for polymer ablation since it

has been shown that polymers with minimal absorption like PTFE and PE can undergo

significant ablation depending on the irradiation conditions.


General Characteristics of Laser Ablation.

The ablation of a polymer surface by a laser pulse is a function of the energy

deposited in the solid per unit time, the wavelength of the laser and the time-length of the

laser pulse. Depending on these factors, polymer ablation can take place in the |tm range.

Figure 2.2 gives a general scheme of the process.The ablation is accompanied by light

and acoustic emissions, pressure and temperature changes. All of these have been









measured in an effort to understand the ablation process. Based on a number of studies

over the past 15 years [32], [38-45], a general curve can be constructed relating the etch

rate (thickness of material etched per pulse) to the energy density (laser energy per pulse

and unit area) deposited into the polymer.


Abkiatd Yduihe,


Ab[Btbon depi~i


Figure 2.2. Schematic representation of the laser ablation process


As shown in figure 2.3 the plot has three distinct regions. In the low energy density

region shown as region I in the figure, (typically 10-100 mJ/cm2), the etch rate is

extremely slow and proceeds in a layer-by-layer fashion. The polymer surface shows

minimal change and ablation as characterized by material removal is seen only after a








few hundred pulses. Beyond a certain limit, usually referred to as the threshold energy

density, the etch rate rises in a linear fashion with the energy density (shown as region II).




Ill


0 I


Energy density, mJ/cm2


Figure 2.3.
polymers.


Characteristic graph of ablation depth/pulse vs. Energy density for


Depending on the mode of ablation i.e. photothermal or photochemcial, the

polymer surface undergoes extensive degradation and material removal. Typical limits

for region II are in the 1-2 J/cm2. In the region III, the etch rate saturates at very high

energy densities. The decrease in etch rate can be due to laser pulse attentuation by the

ablation plume expanding off the polymer surface or change in the polymer absorption

characteristics (optical bleaching). All polymers seem to exhibit the threshold

phenomenon and demonstrate two different etch rates above and below the threshold

energy density. Many studies [32], [44], have tried to fit the linear region of the plot to a









model based on Beer's law of absorption which takes into account the UV absorption

characteristics of the polymer. The fit is based on the following equation:



d= 1 n (1)


which relates the ablation depth per pulse, d, to the natural logarithm of the ratio of the

incident energy density, 4, to the threshold energy density, (4o, with the inverse of the

absorption coefficient, oa1 as the proportionality constant. However it has been found

the slope of this line does not always give the correct value for the absorption coefficient

of the given polymer [46]. Also the model can not be applied at extremely low or high

energy densities. Thus the ablation behavior can not be modeled by a simple UV

absorption model.

Since the dynamics of the ablation process are extremely fast, a knowledge of the

different events occurring and their time scales is critical to the understanding of the

phenomenon. Koren and Yeh studied the light emission intensity at various distances

from a polymer surface after laser irradiation [47],[48]. They concluded that in the

etching ofpolyimide films by ArF pulses (193 nm, 15 ns FWHM), the emission has a

fast component which appeared simultaneously with the laser excitation and a slower

component which lasted 10 to 100 fold longer than the laser pulse itself. Davis et al.

carried out a similar study of PMMA and obtained emission signals on the order of the

laser pulsewidth [49]. These two studies showed that the polymer could begin to ablate

on a time scale that is even shorter than the width of a pulse from a laser beam and the

ablation could last on the order of microseconds.









Dyer and Sidhu studied the temperature rise accompanying laser irradiation of a

polymer by attaching a thermocouple to a 50 |tm polymer film [50]. Temperature

increases in excess of 500 K were detected. Gorodetsky et al. measured the local

temperature rise in a polymer film on laser irradiation by bonding a pyroelectric crystal to

a 75 |tm thick PI film [51]. It was found that the temperature rise in the irradiated area

was directly proportional to the photon energy deposited into the film. These studies

demonstrated the high temperatures seen on the polymer surface during laser ablation.

Dyer and Srinivasan found that the local heating produced by a laser pulse below

threshold levels causes a stress wave which is sinusoidal in character [52]. The

compression wave caused by the heating is rapidly followed by a rarefaction wave caused

by the cooling. It was also found that the stress waves generated had large magnitudes.

For the ablation of a PI film at 193- and 308 nm peak amplitudes in the 105-107 Pa range

were detected. The ablated material is known to be ejected at 2 to 5 times the speed of

sound and the pressure generated on the substrate is consistent with these measurements.

These stress waves could be detected even at extremely low energy densities at which no

significant etching was observed (< 50 nm per pulse). It was calculated that the ablation

of< 50 ng of material at a velocity of 2 x 104 cm/s on a 20 ns scale would account for the

observed stress wave signal for a PI film irradiated at an energy density of 3 mJ/cm2.

There also appeared to be a relationship between the chemical structure of a polymer and

its stress response upon laser irradiation. PMMA gave a sinusoidal response for a laser

pulse of 9 mJ/cm2 (i.e. compression followed by rarefaction), in contrast to PI which

showed a wholly compressive response.









Venugopalan also studied the mechanical response of PI and collagenous tissue to

pulsed laser irradiation at 193-, 248- and 308 nm [53]. Results for PI were in agreement

with Dyer and showed that stress levels in the GPa range are generated in the polymer.

In the ablation of collagenous tissue it was shown that the absorbing chromophore and

the characteristic response time of the polymer decided the mode of ablation. At 193 nm,

strong absorbance due to CO-NH peptide bonds gave clean ablation with thermal damage

localized to a few microns from the ablation site. In the 1064 nm irradiation of tissue,

water acted as the main chromophore and the tissue failed due to mechanical fracture

caused by water expansion.

Von Gutfeld et al. followed the deformation of a film surface on laser irradiation

by reflecting a He-Ne laser off the polymer surface and monitoring the reflection with a

position-sensitive detector [54]. A laser pulse of 20 ns FWHM and < 1 mJ/cm2 energy

density caused a deformation in a 125 |tm PI film which rose to its maximum in 2 |ts. A

50 |tm film had the same response time, but the deformation of the surface was fourfold

greater. Thus these studies established the large temperature rises and deformation

experienced by a polymer surface on laser irradiation.

Dynamics of laser ablation When a laser pulse is incident upon a polymer (or

any other material) surface, the etched material is ejected from the surface as a plume.

The constituents of the plume are ejected with a certain direction and a great velocity.

Plume dynamics has been the subject of much research. The velocity and angular

distribution of certain molecules in the plume has been determined for a number of

polymers. Srinivasan et al. examined the diatomic (C2) signal during 248 nm laser

ablation of PMMA and found its ejection during ablation took place at a supersonic









velocity of 7 x 105 cm/s at energy density levels well below threshold [33]. Velocity

measurements on C2 and CN fragments from the ablation of PI films with 248 nm laser

radiation were also undertaken. A Maxwell-Boltzmann fit to the velocity distribution of

C2 showed a maximum at 6 x 105 cm/s. This indicated that the ablation products traveled

at supersonic velocities off the polymer surface, a measure of the energy transfer

involved in the process.

Chemical composition of ablation fragments. Chemical analysis of the products

of ablation have been carried out in a number of studies Only a limited number of studies

have yielded quantitative information correlating the yield of a given product to the mass

of polymer that was removed by ablation (i.e. mass balance). The findings for PMMA

and PI are summarized below.

PMMA: Srinivasan studied the gaseous species produced during PMMA laser ablation by

mass spectrometry [33]. The most stable gas products were carbon monoxide (CO),

carbon dioxide (CO2) and methyl methacrylate (MMA) in the ablation of PMMA by 193

and 248 nm radiation. The MMA yields were 18 and 1% respectively at 193 and 248 nm

ablation. MMA oligomers in the mass range of 2500-400 amu were also reported. The

morphology of the ablated polymer surface showed significant differences for 193 and

248 nm ablation. The 193 nm laser pulses produced a smooth and debris free etch region,

while 248 nm laser pulses showed considerable roughness and bubbling. It was

demonstrated that the temperature in the etched volume was greater than the softening

point of PMMA (assumed as 150C) during 248 nm ablation but not with 193 nm laser

pulses. In a similar study Krajnovich reported the ablation products of PMMA by a KrF

laser (248 nm, 16 ns FWHM) to be a complicated mix of products [55]. These included









methanol, carbon monoxide, methyl format, MMA, dimers and fragments from the

partially unsaturated polymer backbone. The findings from these two studies are at odds

with each other, but the different energy density ranges used could account for that.

Srinivasan used energy density levels of 0.5-1.5 J/cm2, which was designated as the

steady-state ablation regime. Krajnovich's study used a energy density of 0.25 J/cm2 for

all measurements and found it was possible to ablate PMMA at this level.

PI: A number of studies have found themajor products for ablation of PI at 193 and 248

nm to be HCN, CO, H2, C and H20 [37]. The presence of higher molecular weight

species was not examined. In a report on the formation of surface defects during 248 nm

ablation of PI, Krajnovich and Vasquez gave a mass balance for the process based on

their observations [56]. They calculated that 1.2 x 1015 monomer units were ablated per

second and calculated the volatile molecule production rate at 7.9 x 1015 molecules per

second. Thus each monomer could decompose to give 6-7 volatile molecules (7.9 x

1015/1.2 x 1015) supporting the production of CO and CN during ablation. One major

problem has been the intractibility of the polymer which makes molar mass and

spectroscopic analysis difficult.

Mechanism of polymer ablation The mechanisms underlying polymer ablation

have been studied in order to provide a theory of polymer ablation The two modes of

ablation are seen as photochemical and photothermal. In photochemical ablation, photon

absorption leads to electronic excitation causing bond scission and material ejection off

the surface. In photothermal ablation, electronic excitations are transferred to vibrational

modes causing material heating and evaporation off the surface. Although many studies









have tried to resolve this issue as yet no clear picture has emerged. In fact the mechanism

appears to be dependant on the type of polymer and the kind of laser used.


Pulsed Laser Ablation for Polymer Thin Film Deposition

Although the possibilities of using PLD for thin film synthesis were recognized early on,

research on this aspect has made significant progress only recently. Hansen and

Robitaille published the earliest report on the use of UV laser ablation for polymer thin

film synthesis [36]. They used a variety of polymers in their study namely ; Poly

tetrafluoroethylenee) (PTFE), Polyethylene (PE), Poly (methyl methacrylate) (PMMA),

Nylon 6,6, Polycarbonate (PC), Poly (ethylene terephthalate) (PET) and Polyimide (PI).

The effect of laser wavelengths on polymer film formation was studied by irradiating the

polymer targets at wavelengths of 193, 248, 355 and 1064 nm. Laser pulse energy

densities ranged from 0.01-2 J/cm2. A strong correlation between the deposited film

quality and the polymer absorption coefficient at the excitation wavelength was observed.

Films deposited using 1064 nm laser pulses were characterized by debris and particulate

matter. Films produced by the UV lasers were of higher quality (i.e. particulate free),

smooth and homogenous. These films also demonstrated stronger adhesion to the

substrate than those produced with the 1064 nm laser. The authors reasoned that since

the polymers used were mostly transparent at 1064 nm, the polymer ablated by

evaporation causing particles to eject off the surface. At the UV wavelengths, since most

polymers had some absorption, the ablation was due to photochemical means causing

smaller ions and molecules to eject off the surface.

The laser energy density also had a significant effect on the deposited film

quality. As the energy density was increased, the tendency towards particle incorporation









in the film also increased, indicating that the mode of ablation was also energy

dependant. The importance of the polymer UV absorbance was demonstrated by the fact

that polymers with negligible absorbance (e.g. PTFE, PE) were not deposited as films. It

was also found that the growth rates for the polymers which did form films were in the

range of 0.2-0.8 A/pulse, independent of the starting material. Refractive indices (RI) of

the deposited films varied from those of bulk polymer values. All the aromatic

condensation polymers (i.e. PET, PC and PI) had slightly higher RI values than the bulk

polymer. However no other surface analysis of the films was presented.

Size exclusion chromatography (SEC) of PMMA films gave molar masses which

depended on the laser wavelength used. As the laser wavelength increased (i.e. lower

photon energy), the molar mass of the deposited film increased. The authors concluded

that the polymers ablated by a photochemical mechanism at the UV wavelengths. These

wavelengths produced substantial but incomplete depolymerization of the target and

resulted in the direct transport of small polymer chains. This idea was supported by the

observation that all UV absorbing polymers produced films, even those not expected to

cleanly depolymerize (e.g. PI). At the longer wavelength (1064 nm), the ablation

mechanism was thermal in character. Since a high molar mass was obtained in the

PMMA films at this wavelength, the authors reasoned that these species were transported

by entrainment following subsurface excitation of absorbing impurities.

Although this study was the first one to demonstrate the use of laser excitation for

polymer film deposition, it suffered from several lacunae. First, since the scope of the

study was so large (several different polymers and laser wavelengths), no single system

was studied in any great detail. Second, only a cursory surface analysis of the deposited









films was presented and since this was the main focus it was a severe shortcoming in the

study.

A later report published by the same authors examined the 248 and 1064 nm

ablation of polycarbonate in-depth [57]. Measurements of the time-of-arrival (TOA) of

ablation products at the substrate indicated that ablation at 248 nm did not involve direct

evaporation of small polymer chains. The ablation products were found to travel too fast

to have a mass above a few hundred (120 +40) a.m.u. IR scans of the deposited film

were similar to the starting material. The only difference, as found by gel permeation

chromatography was a severe molar mass degradation. In light of these results the

authors concluded that the ablation mechanism must involve extensive depolymerization

by the laser pulse, transport of the moderate sized fragments and repolymerization on the

substrate. However the authors also mentioned that the TOA analysis did not exclude the

presence of species heavier than 120 a.m.u, which could also contribute to the

repolymerization mechanism.

The above study pointed the way for further research in this field. Research has

focused on addition polymers like PMMA and PTFE for which degradation and

repolymerization schemes have been studied.. Condensation polymers like PET, PI and

PC although attractive from a thin film standpoint have complex structures and

degradation pathways are as yet not well understood.

PTFE. The good mechanical, thermal and chemical stability of PTFE combined

with its low surface tension have made it an attractive candidate for use in thin films and

surface modification. PLD, along with vacuum sputtering, ion deposition and plasma

polymerization have all been investigated as synthesis methods. Blanchet and Shah









reported on the deposition of amorphous fluoropolymers by ablation with 266 nm laser

pulses [58]. Amorphous copolymers of PTFE with 2,2-bis(trifluoromethyl)-4,5-difluoro-

1,3-dioxole (Telfon AF) were used as target materials. Dioxole contents of 66 and 82%

were employed to give copolymers with glass transition temperatures of 160 and 240C

respectively. The films were deposited in argon atmospheres of 50-250 mTorr, with

energy densities ranging from 0.4-1.2 J/cm2. It was found that the substrate temperature

had a great effect on the resultant film quality. Films deposited at room temperature

formed a low-density sponge-like structure. In contrast, by maintaining the substrate

temperature at or above Tg, films with a smooth morphology were obtained. IR spectra

of the films showed a one-to-one correspondence with the starting material. The peak

widths for the deposited films were also in the same proportion as the target indicating

that all the species were present in the same concentrations as the starting material. This

implied that at least a fraction of the complex dioxole structure was not fragmented

during the ablation. The authors concluded that the ablation proceeded via a thermal

unzipping of PTFE chains and transport of the intact dioxole molecule with CF2

fragments to the heated substrate, followed by repolymerization.

Blanchet et al. also studied the ablation of PTFE at 266 nm [59]. PTFE films

were produced at an energy density of 1.5 J/cm2 in an argon atmosphere of 250 mTorr.

The substrate temperature was kept constant at 320C. Surface analysis of the films

showed them to have an F/C ratio of 2. XRD patterns of the films corresponded to the

standard helical configuration of PTFE. The films were estimated to be 50% crystalline

from DSC measurements. The mass spectrum of ablation products of PTFE was found to









match that of vacuum pyrolyzed PTFE, leading to the conclusion that the ablation

proceeded by a thermal mechanism.

Subsequent to this work, other studies have attempted the laser deposition of

PTFE with the goal of producing smooth, adherent, crystalline films with good dielectric

properties. Ueno et al. studied the PLD of PTFE at 157 nm using an F2 laser [60]. Since

the photon energy at 157 nm is higher than the binding energy of the C-F bond it was

expected that unselective bond scission would ensue during the ablation process. Film

depositions were carried out in an argon atmosphere at 250 mTorr with energy densities

in the range 450-600 mJ/cnm2. Films were grown at two different substrate temperatures

of 300 and 373 K X-ray photoelectron spectroscopy (XPS) Cis spectra revealed that

films with predominantly CF2 bonding were produced at both the substrate temperatures.

The CF2 contents were 70 and 92% for the substrate temperatures of 300 and 373 K

respectively. The film morphology was found to be strongly dependent on the substrate

temperature with smooth films being produced at the higher temperature. RI values

obtained by ellipsometry were similar to those of PTFE targets. XRD scans showed that

films deposited at 300 and 373 K were crystalline in nature. One significant finding of

this study was the effect of pressure on the resultant film. Films deposited in 10-5 Torr

were amorphous and XPS Cis spectra exhibited C-F, CC-F bondings. The authors

reasoned that the frequent collisions taking place at a higher pressure (200 mTorr) started

the repolymerization process resulting in a stoichiometric film, whereas at the lower

pressure there would be no plume interactions and the various species would arrive at the

substrate as is causing random repolymerization reactions.









Norton et al. studied the morphology and crystallinity of PTFE films formed by

laser ablation [61]. Films were produced by irradiation of PTFE targets with 248 nm

laser pulses, in the energy density range of 3-9 J/cm2 at a rate of 30 Hz. The films were

deposited in vacuum at base pressures of 10-6-107 Torr. The primary effect was seen to

be due to the substrate temperature. As the substrate temperature increased from 200 to

350 C the amount of crystalline material in the film increased. At temperatures below

200 C amorphous films were produced. The authors suggested that the crystallinity of

the films changed with the substrate temperature during deposition due to the fact that as

the melting point of PTFE is approached, there would be sufficient molecular mobility to

allow orientation of the chains into a crystal. TEM images of the films also showed the

presence of two kinds of crystalline regions contained in one film. Small crystallites (~ 5

nm diameter) and larger crstalline regions (~ 60 nm diameter) were found. Electron

diffraction patterns were consistent for the long chains lying in the film-substrate

interface plane. This study thus confirmed earlier data that suggested that substrate

temperature controlled the film crystallinity.

Similar results were obtained by two other studies which are only briefly

mentioned here. Metal et al deposited PTFE films using 248 nm laser pulses, with

energy densities of 0.5-3.2 J/cm2 at 5 Hz [62]. Argon atmospheres were used to control

the pressure between 75-300 mTorr. The substrate temperature ranged from 25-270C.

Based on an optimization study the following conditions were found to give the best

films : energy density 2.0 J/cm2, pressure 10 Pa, substrate temperature 60C above the Tg.

Li et al. studied the effect of different targets on the film characteristics for the

248 nm ablation of PTFE [63]. It had been suggested that for pressed PTFE targets (vs.









polished PTFE targets) film formation was based on the transport of PTFE grains from

the target to the heated substrate surface and the subsequent melting and crystallization of

the material. Thus two kinds of PTFE targets were used: polished pellets cut from a

PTFE rod stock and pressed PTFE powder pellets. The grain size of the PTFE powder

was about 6-9 |tm. The deposition conditions used were: energy density 0.5-6 J/cm2,

repetition rate 1-10 Hz, substrate temperature 300-5500C. Depositions were carried out

in an Argon atmosphere at a pressure of 0.1 to 1.0 mbar. The powder targets gave films

with superior properties compared to the polished rod targets. Smooth continuous films

were formed with the powder targets. Increasing the substrate temperature above 3000C

produced films with well defined spherulite morphology. Post-deposition annealing of

the powder target films caused an improvement in the morphology and favoured

spherulitic growth laterally upto 100 |tm. These films exhibited good adhesion to the

substrate and passed the Scotch peel test. The rod targets produced rough films with

many particulates. In the Scotch peel test, these films exhibited cohesion failure i.e. large

particulates would be peeled off the surface. The authors suggest that the laser irradiation

of powder PTFE targets caused transfer of complete PTFE grains to the heated substrate,

whereon films were formed due to melting and recrystallization. On the other hand the

laser irradiation of polished PTFE targets resulted in rapid depolymerization and ejection

of low molar mass fragments which repolymerized on the substrate. Since no

examination of the targets after laser ablation was carried out, this theory cannot be

confirmed. It should be mentioned that the films produced from the powder targets were

of remarkable quality and exhibited high crystallinity, excellent dielectric strengths (1-3 x

105 V/cm) and high electric resistivity (1012 Qm). Also the deposition rates achieved









were higher than those produced by comparable synthesis methods like vacuum

evaporation, plasma polymerization and sputtering.

Brannon et al. reported an interesting result for the laser ablation of a plasma

polymerized fluoropolymer film [43]. It was found that the plasma fluoropolymer

underwent efficient ablation at energy density levels which cause very little change in

PTFE. The principal difference between this polymer and PTFE was the plasma

polymer's much greater absorption coefficient (7 x 104 VS. ~ 102 cm-1). A plot of the

ablation depth per pulse versus energy density indicated the threshold for energy density

to be 50 mJ/cnm2. Below the threshold the ablation rate curve could be fit by a single

Arrhenius-type exponential suggesting that the ablation process in the low energy density

regime was partially governed by a photothermal process. This study demonstrated the

importance of the polymer absorption coefficient in defining the ablation conditions.

PMMA. Blanchet produced thin films by ablation of PMMA with a Nd-YAG

laser at 266 nm [64]. Energy densities used ranged from 0.2-0.5 J/cm2, with a repetition

rate of 10 Hz. Films were deposited in an argon atmosphere at 200 mTorr. Examining

the effect of substrate temperature on film formation, it was found that film morphology

was strongly dependant on the substrate temperature. PMMA films deposited at room

temperature appeared hazy and adhered poorly to the substrate. These films appeared to

be composed of fairly spherical particles of various sizes forming a loosely packed

matrix. In contrast, films deposited at 1500C were dense, clear and exhibited excellent

adhesion. Film thickness as a function of the deposition temperature showed a

decreasing trend, especially after the substrate exceeded the Tg of PMMA. This was due

to the large densification which the films underwent at elevated substrate temperatures.









IR spectra of these films corresponded closely with that of a PMMA reference thus

showing that the structure remained the same although the morphology underwent a

change due to the deposition temperature. Another important result was the molar mass

dependence of the film on the deposition temperature and pressure. Films were deposited

from targets with molar masses of 500,000 g/mol and 5820 g/mol at background

pressures of 10-7 Torr and 250 mTorr and temperatures in the range of-1960C to 2500C.

Molar masses of the films were determined by gel permeation chromatography after

dissolving them in tetrahydrofuran. At low temperatures molar mass was found to be

independent of temperature while it increased linearly above Tg. Mw's were about 104

g/mol for temperatures ranging from -1960C to + 800C, increasing linearly to 6 x 104

g/mol at 175C. Surprisingly, the different targets gave films of similar molar mass

under identical conditions. The deposition pressures also caused a change in the molar

mass. Using the target with a molar mass of 500,000 g/mol gave a film with molar mass

of 325,000 g/mol when deposited 10-7 Torr and 1750C. The same target gave a film with

a molar mass of 60,000 g/mol when deposited at 250 mTorr and 1750C.

Blanchet explained these results on the basis of changes in radical and chain

mobility. Since low molar mass, hazy and low-density films deposited at lower

temperatures could be converted to dense clear films by annealing it implied that chain

mobility played a role in determining film morphology. As the substrate temperature was

increased during deposition, the ablation product molecules arriving onto the substrate

would have enough energy to add to the growing chain. Thus high molar mass, dense

and clear films could be produced at higher deposition temperatures. Mass spectrum

analysis of the ablation products matched closely with those for vacuum pyrolyzed









PMMA. Thus the ablation mainly produced MMA monomer which then proceeded to

polymerize via a radical mechanism after arriving on the substrate. Also the threshold

energy density for PMMA was estimated as 40 mJ/cm2 based on the abundance of

monomer species detected in the mass spectrometer at that energy density. Blanchet also

stated that the energy density had no effect on the kind of films produced. This result is

different from an earlier report which showed the formation of a carbon-rich film during

ablation of a PMMA target. Since surface analysis of the deposited films was not

presented energy density effects could not be commented on.

Based on their results for laser ablation of PMMA and PTFE, Blanchet and Fisher

offered the following theory of ablation [65]. Upon laser irradiation of the polymer,

photon energy initially absorbed as electronic excitation is converted to vibrational

modes, rapidly heating the solid. This intense local heating results in pyrolytic

decomposition of the target via rapid unzipping of the polymer chains. An ablation

plume consisting of monomer fragments is explosively ejected off the surface. On

arriving at the substrate repolymerization occurs to give a polymer film. The similarities

between compositions of the ablation plume and pyrolysis products suggested that PTFE

and PMMA ablation occurred via the above mechanism. Also for simple addition

polymers, which are known to depolymerize by unzipping, the ceiling temperature (To)

plays an important role in film formation. Since the ceiling temperatures of PMMA and

PTFE are 320 and 7200C respectively, the thermolyzed monomer molecules impinge

upon the substrate well below T,. To test this theory, the authors ablated poly (ca-methyl

styrene) which is known to thermally unzip during pyrolysis. Analysis of the ablation

products by mass spectroscopy showed the presence of monomer and dimer species only.









However, films of poly (ac-MS) could not be formed by this process. Since the To of this

polymer is 610C it was theorized that the ablation products would arrive at the substrate

with vibrational energies corresponding to much higher temperatures. Thus there would

not be enough driving force for the repolymerization reaction. However this theory was

not tested by performing a deposition at a substrate temperature substantially below the

Tc. Later reports have since shown the ablation products of PMMA to be much more

complex indicating a different ablation pattern than simple unzipping. Thus it is almost

certain that the ablation proceeds by mechanisms which are more complex than the

simplistic one presented by Blanchet and Fisher.

SI-based polymers. Suzuki et al. used PLAD to synthesize silane based polymer

films from poly(methyl phenyl silane) (PMPS) [66]. It was found that for the 248 nm

irradiation of PMPS, most of the Si-Si bonds in the polymer were converted to Si-C

bonds in the deposited films and new bonds like Si-H were also present. The deposited

films suffered a molar mass reduction and showed a broader distribution than the starting

polymer. The molar mass distribution was also found to be dependant on the energy

density.

Natural Polymers and Biomolecules. Several other polymers have also been used

for thin film syntheses by laser ablation. These include polyacrylonitrile [67],

poly(phenylene sulphide) [68],[69], polythiophene, polyethylene and polyethylene oxide

[70]. These studies have demonstrated the feasibility of using this technique for

producing polymer thin films.

Due to the interest in producing biocompatible coatings, PLD of collagen and

proteins has also been attempted. Conklin and Cotell have reported the PLD of collagen









and apatite/collagen composite thin films [71]. Films were deposited under varying

conditions from fibrous collagen targets alone and mixtures of hydroxyapatite (HA) and

collagen. An ArF or KrF excimer laser was used in all the depositions in the energy

density range of 0.2-1.5 J/cnm2, in an argon atmosphere. The effect of water in the

deposition environment was also assessed by bubbling argon through water before

introducing it into the chamber. Surprisingly, it was found that varying the conditions

had little effect on the IR spectra of the collagen films. However, the deposition

conditions did affect surface morphology with lower energy density levels producing

smoother films. Gel electrophoresis and IR results indicated that films were comprised of

collagen with an altered secondary structure and possibly shorter chain lengths. XRD

analysis revealed a broad amorphous reflection regardless of deposition parameters

indicating that the process did not introduce crystallinity in the films. It was also found

that water vapor increased the smoothness of the surface while the IR spectra did not

show any change. XRD analysis of the HA/collagen composite films revealed that the

HA in the deposited film was crystalline. This result was unexpected since the deposition

was carried out at room temperature, where amorphous HA is produced. The authors

suggest that the presence of collagen in the target causes immobilization of the nucleating

calcium species on ablation causing it to crystallize.

Tsuboi et al. used PLD for the thin film deposition of a silk fibroin

polypeptide[72]. This protein is being investigated for its potential applications in

molecular devices and bioelectronics. The PLD method was attractive due to the poor

solubility of silk fibroin in many common solvents. The polypeptide target was ablated

with an excimer laser at 351 or 248 nm at an energy density level of 700 mJ/cm2. IR









spectra of the film deposited at 351 nm showed correspondence with that of the starting

polymer. Hence the primary structure of the protein was successfully reproduced in the

film. However new absorption bands were seen for films deposited at 248 nm, indicating

that part of the structure was destroyed. Atomic force microscopy revealed the deposited

film to consist of particles, with smaller particle sizes being produced at 248 nm than at

351 nm.

The above examples of PLD of natural molecules to produce organized structures

are certainly thought provoking. As an extension of the idea, Phadke and Agarwal have

reported on the PLD of preformed structures of organized molecules [73], [74]. Lasers

were used to ablate glucose oxidase dissolved in a sodium dodecyl matrix. The deposited

films were shown to retain their enzymatic activity. Similarly, films of Riboflavin were

synthesized by this method. It was shown that the native properties such as fluorescence

and absorption characteristics remained intact in the films. These examples open the

possibility of a solvent free route to produce complex film structures in a desired

architecture. The above applications hold particular relevance for device fabrication.


Surface Modification by Laser Ablation

One of the other major application of laser ablation has been for polymer surface

sensitization. Laser irradiation can produce surface regions with an altered surface

chemistry as compared to the bulk polymer. Furthermore the degree of control which can

be exercised using lasers gives modified regions with features on the order of

micrometers. Applications of this technique reported in the literature include chemical

patterning, metal deposition and surface functionalization.









Metallization of PTFE and other fluoropolymers has offered new applications in

the area of electronic packaging. However, due to its chemical inertness, adhesion to

PTFE is poor. Latsch et al. have studied metal depositon on laser modified PTFE

surfaces [75]. PTFE was ablated while immersed in a solution of tetramethylammonium

hydroxide (TMAH). It was found that even after ablation at low energy densities (50-

120 mJ/cm2) PTFE surfaces turned hydrophilic. XPS spectra of the ablated surface

showed the presence of nitrogen and oxygen, indicating that TMAH bonding had

occurred. The formation of hydrophilic groups was also accompanied by defluorination.

Copper deposited on these modified surfaces exhibited stronger adhesion as compared to

copper deposited on virgin PTFE. A similar study has reported on the modification of a

fluorocarbon surface by an ArF excimer laser in gaseous N2H4 for electroless deposition

of copper and nickel.

Copper metallization was attempted using an ArF excimer laser to irradiate a

fluoro-copolymer (FEP) film immersed in an aqueous copper sulfate solution [76]. The

irradiation of the polymer caused defluorination and copper was nucleated on the ablated

regions through the formation of C-O-Cu bonds. The authors suggested that oxygen

plays an important role in polymer-metal adhesion. It was also found that energy

densities in excess of 25 mJ/cm2 gave lower copper nuclei density due to surface

roughening. Thus the technique works only due to photochemical reaction and fails

when ablation starts to occur.

Niino and Yabe have used laser irradiation to induce dehyrochlorination on a

polyvinylidene chloride (PVDC) film to produce conjugated polyene and polyyne

structures [77]. Usually chemical reagents such as metal amalgams, potassium









hydroxide, electrochemically activated tert-butanol etc. are used for the

dehydrochlorination reaction. However these reactions leave salt of acid or base remains

in the modified polymers. KrF excimer laser irradiation of PVDC at a energy density of

30 mJ/cm2 was found to efficiently produce conjugated carbon bonds without

graphitization and crosslinking between main chains. The conjugated structures have

applications as materials for nonlinear optical devices.

Song and Netravali used excimer lasers to modify the surface of ultra-high

strength molar mass polyethylene (UHSPE) fibers for enhanced adhesion in

composites[78]. Using a XeCl excimer laser operating at 308 nm, UHSPE fibers were

exposed to a number of laser pulses at energy density levels of 280, 450 and 768

mJ/pulse cm2. The number of pulses ranged from 100 to 500. Laser exposure was found

to produce surface roughening of the fibers. The degree of roughening depended upon

the number of pulses and the energy density level. The fibers became more polar after

surface treatment and exhibited an increased oxygen content. The interfacial shear

strength increased significantly and this enhancement was attributed to the increased

surface roughness and polarity of the fiber. It is worth noting that energy density levels

used in this study are much higher as compared to the 3 studies mentioned above. The

high energy density levels would cause significant ablation of the polymer surface, in

contrast to the low energy densities mentioned earlier which would only cause surface

modification.

Another novel application of laser ablation has been in stripping of polymer

insulation of metal wires. Brannon and Synder have studied the removal of polyurethane

insulation by pulsed laser irradiation [79]. They found that excimer lasers produced









cleaner, higher quality insulation removal as compared to carbon dioxide lasers. The

improved performance of the excimer lasers was due to the increased absorption of the

ultraviolet radiation by the polymer insulation. They also that found that doping the

insulation with rhodamine G6, a photosensitizer and irradiating it with a 532 nm

frequency doubled Nd:YAG laser produced similar results to the excimer laser

irradiation. A dye loading of 10% gave efficient stripping at an incident energy density

of 650 mJ/cm2, a pulse duration of 12 ns and an exposure of 3-5 s at 10 Hz.

Thus it can be seen that the true potential of laser processing of polymers has been

realized only recently. The list of applications which are being targeted is ever increasing

and it is expected to be a promising and fruitful area of research [80-82].




Silicone Films and Coatings

Silicone is the generic name for a variety of silicon based polymers which consist

of Si-O main chain linkages with organic (methyl or otherwise) groups connected to the

Si atoms. Silicone can be synthesized in a number of forms such as fluids, gels, and

elastomers. Depending on the desired application, the structure can be tailored to

incorporate different chemical groups e.g. phenyl groups for structural rigidity, cyclic

siloxanes for higher viscosity, fluorocarbons for low surface energy. The molar mass can

also be varied from 500 g/mol to 750,000 g/mol. Polydimethylsiloxane (PDMS) is the

most commonly used silicone. The outstanding properties of PDMS arise in part due its

structure [83]. The Si-O linkage has a high freedom of rotation which imparts flexibility

to the structure. The methyl groups connected to the Si atoms have a large volume of

exclusion which imparts chemical inertness and low surface energy to the polymer.









Silicone use in the biomedical field started in the 1950s [84]. The most well-

known (and controversial) example of their use has been in silicone gel breast implants.

Other applications include catheters, tubing, blood bags, soft tissue implants, intraocular

lenses and encapsulation compounds. The haemocompatibility of silicone has also been

studied widely. The Cooper group has produced a series of reports evaluating the blood-

contact properties of a host of polymeric biomaterials [85-87]. In a recent study, a series

of silicone surfaces were systematically modified by siloxane coupling agents with

different chemical end-groups like alkyl, amine and fluorocarbon [88]. These surfaces

were then evaluated for their blood compatibility in a canine ex-vivo shunt. It was seen

that the unmodified silicone surface gave the best haemocompatibility in terms of low

platelet adhesion and longer clotting times.

Although silicone has excellent biocompatibility, its poor mechanical properties

preclude its use in any structural biomaterial applications like vascular grafts. Thus

research efforts have concentrated on producing silicone thin films and coatings. The

most common synthesis route is by the plasma polymerization of various siloxane or

silazane monomers. Comprehensive reviews on this topic has been published [89],[90].

Plasma polymerization typically produces a cross-linked, branched polymer with

molecular structures which are frequently altered compared to the starting monomer.

Plasma polymer films may also contain trapped radicals, which can react on exposure to

atmosphere, changing the film surface over a period of days or even months. This can

lead to a deterioration in the film properties. Also, such films may be under tensile stress

leading to cracking or peel-off from the substrate. Therefore the properties of these films

can be quite different than those desired of a siloxane polymer.









To overcome some of these limitations, Kwan and Gleason used a pyrolytic

process to synthesize silicone thin films [91]. It was seen that the low-pressure

thermolysis of octamethyltetrasiloxane (commonly known as D4) produced films with

structures similar to PDMS. Roth et al. used excimer lasers to crosslink PDMS

copolymers in the liquid phase and also to produce silicone films by the laser-induced

chemical vapor deposition from the gas phase [92]. These studies demonstrated that

silicone thin films could be produced by methods other than plasma polymerization.

The unique combination of properties of polydimethylsiloxanes (PDMS) has led

to their use as coatings in a variety of fields including biomaterials, membranes, optics

and dielectrics. Newer applications such as barriers for flame retardancy [93] and

passivation of high Tc superconductors[94] have also been reported. Due to their

toughness, low surface energy, low and high temperature stability, unique gas

permeability and low dielectric constant, it is envisioned that silicone thin films will

continue to find new applications and therefore research on synthesizing these films

would be a promising area.

The research presented herein explores the use of pulsed laser ablation deposition

of silicone for synthesizing thin films. It is believed that this new method would have

significant advantages over current methods in synthesizing these films.














CHAPTER 3
PULSED LASER ABLATION DEPOSITION OF SILICONE


Introduction

As previously noted, the pulsed laser ablation deposition (PLAD) of silicone or other

corss-linked polymers for that matter had not previously been reported. Thus there was a

lack of experimental parameters on which to base these studies. From a study of the

literature, it was seen that that the energy density (i.e. laser energy per unit area incident

on the target surface) was the most important factor which controlled the kind of film

produced. Therefore, this study chose to examine in detail the effect of the energy

density on the deposited film structure and chemistry. Based on the results of this study,

the effect of other parameters like laser pulse number and pressure level on film structure

and composition was also examined.




Materials

Targets. Silicone targets were prepared in sheet form by casting and curing a

two-part resin system (Shin-Etsu KE 1935 A,B) at 700C for 24 hours with a platinum

catalyst to produce a cross-linked elastomer. Figure 3.1 shows the curing reaction for

producing the elastomer. Plaques ( 2.5 cm dia.) were punched out from the sheets and

extracted in chloroform for 48 hours to remove soluble unreacted polymer. The extracted

plaques were dried under vacuum (12 hours) and mounted on 1 inch dia. aluminum stubs

for use as targets.









Substrates. Silicon wafers or 316L stainless steel (1 cm X 1 cm) were used as the

substrates in all the depositions. Silicon wafers were cleaned by sequential sonication in

acetone and methanol then dried under nitrogen. Steel substrates were cleaned by

sequential sonication in 1,1,1 tri-chloroethane, acetone and methanol then dried under

vacuum.


CH3

\ ---OSi-CH=CH2 + H-Si-CH3

/CH |
CH3 H3 C-Si-H

-__ OSi-CH=CH2 + H- i-CH3

CH3


CH3

+ H2C CH-- Si-i

CH3


pt
P -t N -OSiCH2CH2 i-CH3




I3 O CH3
OSi-CH2CH2-Si-CH3

CH3




Figure 3.1: Curing reaction for preparing silicone elastomer targets.




Methods

The pulsed laser ablation deposition (PLAD) system consisted of a vacuum

chamber which housed the target and substrate. The vacuum chamber was connected to a









mechanical displacement pump allowing the chamber to reach a base pressure of 30

mTorr. The target was mounted on a motor and rotated during each deposition run so as

to ensure uniform ablation. A KrF excimer laser (Lambda Physik 301x) operating at 248

nm with a pulse width of 25 ns was used in all the experiments. The laser beam was

directed into the chamber by a pair of plane mirrors and a collimating lens. The beam

was focused by a 250 mm focal lens placed at the entrance to the chamber. The laser

was operated in the energy range of 100-200 mJ and the energy density was adjusted by

moving the focussing lens along a track to get the desired spot size incident on the target.

Laser spot size was measured by creating burn areas on thermal paper and measuring the

areas with a 10x graduated microscope eyepiece. The eyepiece resolution was 0.1 mm.

A joulemeter (Sentec ED 500) was used to measure the incident laser energy inside the

chamber. The energy density (also referred to as fluence in the literature) ranged

between 75-600 mJ/cm2. All depositions were carried out in a Helium atmosphere at a

pressure of 100 mTorr and the laser operating at a repetition rate of 5 Hz unless otherwise

stated. Figure 3.2 shows a diagram of the experimental setup. Fitz-Gerald [95], provides

additional details on the construction and setup of the PLAD chamber.

The procedure for an experimental run was as follows. The substrate and target were

mounted in the chamber and the vacuum pump turned on. The chamber was evacuated to

< 30 mTorr with two purge fills. The pressure was then adjusted to 100 mTorr by using a

helium backfill. After aligning the optics with a He-Ne laser, the excimer was turned on.

The target was ablated for the requisite amount of time with the pressure being monitored

during ablation. Typically deposition times ranged from 15-20 min, thus giving a laser

pulse number range of 4500-6000 (5 pulses/sec x 60 sec x 15-20 min). After the ablation









was complete the chamber was allowed to return to atmosphere by bleeding in helium.

The modified substrates were then stored in polystyrene boxes prior to analysis.

Deposition experiments under high vacuum. The chamber setup for the high vacuum

chamber was similar to the one described above. The pumping unit included a

turbomolecular pump (Pfeiffer TCP 300) backed with a mechanical displacement pump

allowing the chamber to reach pressures < lx 10-6 Torr. All other arrangements like

target and substrate mounts, laser optics, energy density measurements were similar to

those described above. The depositions were carried out at a pressure of 1 x 10-4 Torr

under a helium atmosphere.


Characterization

Fourier-Transform Infrared Spectroscopy (FTIR). Reflectance absorption FTIR

spectra were collected using the microscope stage on a Nicolet Magna 706 FTIR

spectrometer. Spectra were collected in the range of 650-4000 cm- with a spectral

resolution of 4 cmn, averaging 200 scans per sample. Each spectrum was corrected for

carbon dioxide and water absorption since data collection was carried out in laboratory

atmosphere. An uncoated steel sample was run as a baseline and used for all baseline

corrections. It should be mentioned that attempts at collecting spectra in the Attentuated

Total Reflection (ATR) mode of the FTIR failed due to contact problems of the substrate

with the ATR crystal.

Contact Angle Goniometry. Static contact angles were measured using the

captive air-bubble technique with ultrapure water as the immersion liquid on a Rame-

Hart A-100 goniometer. Measurements were taken on both sides of a 2 p.L air bubble

with five













Pressure transducer


Laser
Beam







To vacuum


Figure 3.2 Schematic diagram of the vacuum chamber setup for the PLAD process.









bubbles being measured for each sample. Each reported contact angle is an average of

these measurements.

X-Ray Photoelectron Spectroscopy (XPS). XPS was performed using a Kratos

XSAM 300 spectrometer. This system consists of a Mg Kuc x-ray source (hv = 1253.6

eV), a hemispherical analyzer and a multichannel detector. The X-ray source was

operated at a power of 108 W. A pass energy of 25 eV was used for spectral collection.

All spectra were collected at a photoelectron takeoff angle of 90 High resolution

elemental spectra were used for quantitative analysis. Peak quantification was done by

using the peak integration package of the DS800 software.

Scanning Electron Microscopy (SEM). The film morphology was observed with

a JOEL 6400 electron microscope, operated at an accelerating voltage of 5 kV, an

aperture setting of 2-3 and a condenser lens current range of 8-9 nA. A thin film of

Au/Pd was sputtered on all samples prior to examination.

Stylus Profilometry. Film thickness was measured by masking a portion of a

silicon wafer with Teflon tape during deposition and measuring the resulting step height

by a stylus profilometer (Tencor Alpha Step 500). The conditions for measurement were:

scan length of 500 |tm, stylus force of 8.3 mg and scan speed of 50 [tm/s.




Results

Effect of Energy Density

Films were deposited at different energy densities and analyzed by FTIR, XPS, SEM,

Contact angle goniometry and Profilometry.









FTIR Figure 3.3 shows the FTIR spectra of PLAD films deposited at energy

densities of 100-, 250-, and 510 mJ/cm2. The three spectra chosen were representative of

low, medium and high energy density levels. Table 3.1 lists all the observed absorbance

peaks along with literature assignments. A silicone polymer used as a target is shown for

comparison in Figure 3.4.




Table 3.1. FTIR absorbance bands observed for PLAD films along with literature
assignments.
Position (cm ) Assignment and mode Ref
I 3400 (broad) H-bonded OH stretching [83],[96-97]
II 2970-2925 CHx asymmetric, symmetric stretching [83],[96-97]
III 2250-2150 SiH stretching [83],[96-97]
IV 1260 SiMex deformation [83],[96-97]
V 1215-930 SiOSi asymmetric stretching [83],[96-97]
VI 900-730 Si(CH2)3 and Si(CH2)2 rocking [83],[96-97]


Depending on the energy density, the spectra exhibit several important differences arising

from the ablation process. The spectrum of the film deposited at 510 mJ/cm2 (Fig. 3.3c),

shows two new absorbance bands due to SiH stretching at 2245 cm1 and OH stretching

in the region 3000-3600 cm1. The OH stretching peak arises from SiOH bonding, since

no other peak due to carbonyl or ester groups (i.e. C-O bonds) is manifested in the

spectrum. The existence of these peaks suggests that the laser ablation of silicone at this

energy density results in large scale degradation and recombination reactions. The

presence of SiOH groups maybe attributed to quenching of free radicals by oxygen and

water vapor either in-situ or on exposure to atmosphere. The position of the SiH

stretching peak at 2245 cmn1 indicates that it is bonded to electron withdrawing

substituents i.e. trifunctional siloxane units.







53












V
I
(c)

II
VI
IV





(b)










3600 3100 2600 2100 1600 1100 600
Wavenumbers, cm









Figure 3.3 Reflectance absorption FTIR spectra of PLAD films deposited at: (a) 100,
(b) 250 and (c) 510 mJ/cm2.









It is also seen that the Si-O-Si asymmetric stretching peak is broadened which indicates a

scrambling in the structure of the Si-O main chain due to distortions, branching and

cross-linking reactions. This is seen more clearly in Figure 3.5 which shows an expanded

spectrum in the range 700-1400 cmf1. The spectrum for the film deposited at the

intermediate energy density of 250 mJ/cm2 (Fig. 3.3b) shows the same features indicating

that the ablation and recombination process is similar at that this energy density level.

On the other hand, the spectrum of the film deposited at 100 mJ/cm2 (Fig. 3.3a) is

substantially different from the preceding examples. The SiH and SiOH peaks seen in

the earlier films are absent. Spectral features characteristic of polydimethylsiloxanes

including CH stretching peaks at 2925 cm-1, Si-O-Si asymmetric split between 1215-930

cm-1 and Si-CH3 deformation at 1210 cm1 are seen. The peak at 1410 cnf1, indicative of

Si-(CH2)x-Si bridges is masked due to absorption from water vapor. The Si-O-Si peak

(Fig 3.5a) shows clear splitting signifying longer siloxane chains and greater siloxane

rings are retained in the film structure. This is also seen by a stronger absorption of the

Si(CH2)2 peak at 810 cm-1 as compared to Si(CH2)3 at 875 cm1, indicating a larger

concentration of dimethyl siloxanes compared to trimethyl siloxanes (i.e. chain ends).

Figure 3.3a matches the FTIR spectrum shown in Figure 3.4 for the silicone target

suggesting that ablation at this energy density level produces a film which is similar to

the target polymer.

X-Ray Photoelectron Spectroscopy. Surface elemental compositions of the

PLAD films on steel were obtained by high resolution scans for carbon, oxygen and

silicon. Iron and chromium peaks from the steel substrate were not seen since the films

were thicker than 200 nm.







55






















0
,0









3650 3150 2650 2150 1650 1150 650


Wavenumbers, cm-1


Figure 3.4 Attenuated Total Reflectance -FTIR spectrum of a silicone target.






















S(b)












1350 1250 1150 1050 950 850 750
Wavenumbers, cm-1















Figure 3.5 Expanded FTIR spectra of PLAD films at different energy densities: (a)
100-, (b) 250- and (c) 510 mJ/cm2.









Figure 3.6 shows the % C, % 0 and % Si values respectively for the PLAD films

as a function of energy density. A silicone sample which was used for a target was run as

a reference and is shown at zero energy density. The arrow indicates one sample with a

very low silicon content, produced at low pulse number. Figure 3.7 shows representative

high resolution carbon, oxygen and silicon peaks for the PLAD films. Figure 3.8 shows

the high resolution carbon, oxygen and silicon peaks for a silicone sample used as a

target.

XPS analyses show that the carbon content of the films decreases while the oxygen

content increases with increasing energy density levels. The oxygen levels even at the

lowest energy density are higher as compared to the target silicone polymer. This seems

to indicate that there is polymer chain scission as a result of the laser irradiation. The

chain degradation can lead to free radical generation due to bond-breakage. These

radicals can recombine with oxygen in-situ or in atmosphere leading to an increased

oxygen content. The radical production can also be due to methyl group abstraction

which has been shown to be the primary degradation pathway in plasma polymerization

of siloxanes and the thermal degradation of silicone. At low energy density, chain

scission and methyl group abstraction may be minimal leading to only a small increase in

the oxygen levels. This would mean that longer chains are evaporated off the surface

which can then recombine on reaching the substrate. There also seems to be another

factor at work at the low energy density level. It is seen that the silicon content is

depressed for the films produced at <150 mJ/cm2. This seems to indicate that the ablation

species produced are depleted in silicon. This could mean that the ablation produces not






























0 100 200 300 400 500 600 700

Energy Density, mJ/cm2


0 100 200 300 400 500 600 700

Energy Density, mJ/cm2




Figure 3.6 Elemental composition of PLAD films as a function of energy density: (a)
% Carbon, (b) % Oxygen.


(a)
* I


(b)
I
I I
+ *
I I I























S* (c)


0 100


Energy Density, mJ/cm2


Figure 3.6contd.
density: (c) % Silicon.


Elemental composition of PLAD films as a function of energy









only polymer chains but also smaller charged molecular and atomic species like *CH2

and *CH3 which can get incorporated into the film structure.

Examining the high resolution elemental spectra (Fig. 3.7), it is seen that there is

no change in the carbon and oxygen peak shape as a function of energy density. This

indicates that no new bonds (e.g. C-O) are formed even after significant fragmentation

and recombination. The silicon peak however shows a shift in binding energy depending

upon the energy density. The change in the silicon binding energy at different energy

densities is plotted in Figure 3.9. A silicone target was also run and the Si2p peak

location for this sample is plotted at zero energy density. To minimize the effects of

instrument drift these samples were analyzed in a single run in the XPS and the detector

was calibrated by setting the carbon Is peak at 285 eV. Figure 3.9 shows that the Si2p

peak shifts to higher binding energy with increasing energy density. This indicates that

the binding environment of the silicon atoms is changing [98]. The silicon 2p peak

appears at 101.8 eV for silicon polymers and at 103-104 eV for SiO2. Thus the shift to

higher binding energy in the PLAD films is an indication of the increasing Si-O bonding

in the structure.

Contact angle. Figure 3.10 shows the effect of energy density on the contact

angle of the deposited films. At low energy densities contact angle is high and

approaches that of silicone (shown at zero energy density). As the energy density is

increased, the contact angle starts to decrease. A transition is seen in the range of 150-200

mJ/cm2, with the films changing from hydrophobic to hydrophilic. Beyond this range all

the films exhibit



























295 290 285 280 275 270 265
Binding Energy, eV


550 545 540 535 530 525
Binding Energy, eV


520 515 510


Figure 3.7 High resolution Elemental XPS scans of PLAD films deposited at
different energy densities: (a) Carbon is, (b) Oxygen Is.























A (c)
,' 250 mJ/cm



S100 mJ/cm2







!0 115 110 105 100 95 9

Binding Energy, eV


Figure 3.7 High resolution Elemental XPS scans of PLAD films deposited at
different energy densities: (c) Silicon 2p.








































Binding Energy, eV


(b)



*i
I*i
I
II
II

I I

I
I I


I I
II
I



I I








550 545 540 535 530 525 520 515 510

Binding Energy, eV


Figure 3.8 High resolution elemental XPS scans for a silicone target: (a) Carbon Is,
(b) Oxygen Is









































120
















Figure 3.8 contd.
Silicon 2p.


115 110 105 100 95 90

Binding Energy, eV


High resolution elemental XPS scans for a silicone target: (c)







65



















104.5


104 -
0
(U

o 103.5 -

*-z
0 103 -


S102.5 -


102 I


101.5

0 100 200 300 400 500 600 700

Energy Density, mJ/cm


Si2p peak location from XPS for PLAD films.


Figure 3.9:



























100-


80 1


60-


c 40

0
20 e* * -




0 100 200 300 400 500 600 700

Energy Density, mJ/cm2


Contact angles of PLAD films as a function of energy density.


Figure 3.10









a low contact angle and are uniformly hydrophilic. The film characteristics give rise to

three energy density regimes which are shown in the figure. They are discussed fully in

the next section.

Deposition rate. Deposition rate as a function of energy density is plotted in

Figure 3.11 for films deposited at 100 mTorr. The deposition rate was calculated per unit

time after measuring the film thickness post deposition. As seen in Figure 3.11, the

deposition rate starts out low and shows practically no change upto 180 mJ/cm2. Beyond

this value, the deposition rate increases rapidly with energy density. The effect of

pressure during the deposition can be seen in Figure 3.12 for films deposited at 50- and

200 mTorr. Deposition rates were consistently higher for runs carried out at 50 mTorr

than those at 200 mTorr. It is known that the plume distribution changes as an effect of

the pressure. At lower pressures the ablation plume progresses further, thus increasing

the mass flux incident upon the substrate directly in its path leading to higher deposition

rates. (Figures 3.11 and 3.12 cannot be compared due to different target-sample

orientations).

Film morphology. SEM micrographs for PLAD films deposited at different

energy densities are shown in Figure 3.13. Films deposited at energy densities of 100-,

220-, 360-, 420- and 510 mJ/cm2 are shown. Complete coverage of the substrate was

observed at all energy density levels. The films deposited at 100 and 220 mJ/cm2 (Fig.

3.13 a, b) were smooth. At higher energy densities more debris was seen on the films and

the films appear to be composed of particulates. This suggests that at high energy density

levels, ablation produced large fragments and clusters which are then deposited on the

substrate.





























0.25


0.2


0.15


0.1


0.05


0


0 100 200 300 400 500


2
Energy Density, mJ/cm


Deposition rate as a function of energy density for PLAD films.


Figure 3.11






69










1
200 mTorr

0.8 50 mTorr



0.6 -



: 0.4



0




0 100 200 300 400 500 600 700


Energy Density, mJ/cm2


Effect of pressure on deposition rate of PLAD films.


Figure 3.12

































































Figure 3.13 Scanning Electron Micrographs of PLAD films deposited at: (a) 100 and
(b) 220 mJ/cm2.


5KU

























5KV >43,00iO 15mm
..... .........
























































Figure 3.13 contd Scanning Electron Micrographs of PLAD films deposited at: (c)
420 and (d) 510 mJ/cm2


5KU









PLAD Film Stability Testing

The film stability was evaluated by a 24 hour immersion in chloroform followed

by XPS and SEM analysis. Figure 3.14 shows the % C, % 0 and % Si values

respectively for the PLAD films, pre-and post-chloroform wash. It is seen that except for

the film deposited at 450 mJ/cm2, there is very little change in the elemental ratios after

the chloroform wash. The film deposited at 450 mJ/cm2 shows an increase in the %C but

a decrease in %0 and %Si. However no change in the peak bonding was seen in the XPS

spectra. This suggests that there may have been some hydrocarbon contamination on this

sample surface. Micrographs of the pre- and post chloroform wash film surfaces are

shown in figure 3.15. The films do not show any significant morphological changes after

the chloroform wash. Specifically, no areas with debonding or peel-off were seen on any

of the samples, indicating that the film-substrate adhesion is good.



Effect of Laser Pulse Number

Literature reports had suggested that the ablation characteristics of a polymer

change as a function of the laser pulse number. In weakly absorbing polymers, initial

absorption (approx. 100s of pulses) can produce no significant ablation but serve only to

increase the temperature on the target surface. Thus the makeup of the ablation

fragments changes as a function of laser pulse number. To clarify this effect, PLAD

films were prepared by ablating the silicone target for 1000, 2000 and 4500 pulses. The

films were then analyzed by XPS to determine the change in composition as a function of

pulse number. Table 3.2 gives the elemental compositions of the PLAD film as a

function of pulse number and energy density.






























200 300
Energy Density, mJ/cm2


200 300
Energy Density, mJ/cm


Figure 3.14 XPS elemental % for PLAD films, pre- and post chloroform wash: (a)
Carbon, (b) Oxygen.


(a) *Pre-wash
5 OPost-wash








I 0>


(b)














Pre-wash
I Post-wash










































200


2
Energy Density, mJ/cm


Figure 3.14 contd.
(c) Silicon.


XPS elemental % for PLAD films, pre- and post chloroform wash:


(c) I 8
0











SPre-wash

SPost-wash























































Pre-chloroform wash Post-chloroform wash





Figure 3.15 Micrographs of PLAD films pre- and post chloroform immersion: (a), (b)
100 mJ/cm2; (c), (d) 125 mJ/cm2; (e), (f) 250 mJ/cm2


!5 x v














5KQ X3#*IkW 116mm


































Pre-chloroform wash


Figure 3.15 contd. Micrographs of PLAD films pre- and post chloroform immersion:
(g), (h) 300 mJ/cm2; (i), (j) 450 mJ/cm2.


Post-chloroform wash









Table 3.2. Elemental compositions for PLAD films deposited at different laser pulse
number.
Energy Pulse % % % %
Density Number Carbon Oxygen Silicon Iron
mJ/cm_
100 1000 44.51 47.25 3.66 4.59
2000 43.46 46.43 6.66 3.44
4500 43.40 38.52 18.08
250 1000 44.11 38.61 17.28
2000 31.00 45.64 23.36
4500 29.96 47.06 22.98
500 1000 26.82 46.67 26.50
2000 26.62 46.58 26.80
4500 27.17 45.89 26.94


The XPS analysis showed that the pulse number at a certain energy density did

affect the film composition. At 100 mJ/cm2, the target irradiated for 1000 and 2000

pulses failed to produce a film. The elemental compositions for these samples were close

to that of uncoated steel substrates. The carbon and oxygen XPS peaks also showed a

different binding state than ordinarily seen for silicone polymers. The target irradiated

for 4500 pulses produced a film with composition similar to those seen earlier at this

energy density. As the energy density was increased to 250 mJ/cm2, the target started to

ablate even at 1000 pulses. However, the film produced was depleted in silicon

indicating that the ablation had not reached its steady-state. The films produced at 2000

and 4500 pulses are close in composition demonstrating that the ablation is steady-state.

At the highest energy density of 500 mJ/cm2, the ablation at all three pulse numbers gave

films which were similar in composition to each other. This indicated that ablation

started rapidly at this level and attained steady-state in 1000 pulses or less. These results

indicated that the ablation of silicone at 248 nm may have a strong thermal component.

At low energy density, the temperature needed for ablation processes to start is only









reached after thousands of pulses. As the energy density is increased, ablation may start

after a few hundred pulses or even immediately.


Depositions under High Vacuum

The results of the energy density study showed that the deposited films had high

oxygen contents. It was thought that this could be due to oxygen attack by elemental

oxygen present in the chamber under the mild vacuum conditions (100 mTorr) used to

deposit the films. Therefore, experiments were carried out at higher vacuum levels to

explore this possibility. Films were deposited at three energy density levels. The XPS

and contact angle values for these films are shown in table 3.3. The film thickness was

kept low and thus FTIR spectra could not be collected for these films.




Table 3.3. Contact angles and XPS elemental % for PLAD films deposited under
high vacuum.
Energy Pressure, Average %C %0 %Si
Density, Torr contact
mJ/cm2 angle, deg
100 1 x 10-4 69 48.28 31.57 20.15

220 1 x 10-4 18 30.22 40.02 29.76

400 1 x 10-4 18 23.21 41.55 35.24



The results show that films deposited under high vacuum show similar trends to

those deposited under lower vacuum conditions. At low energy density the films are

silicone-like and hydrophobic. As the energy density is increased the oxygen content

starts to increase and the films are hydrophilic. The film compositions are comparable to









those produced at 100 mTorr. This indicates that the films react with oxygen or water-

vapor only on exposure to atmosphere.




Discussion

The results reported here show that the pulsed laser ablation deposition of silicone

was successful in producing stable cross-linked films. It was seen that varying the energy

density had a profound effect on the film stoichiometry and structure. By controlling this

factor, it was possible to deposit silicone films with varying oxygen contents and surface

energies. Such behaviour is unique and has not been reported before for linear

thermoplastic polymers like PMMA and PTFE.

The surface analysis showed that at low energy densities (approx. < 150 mJ/cm2)

the deposited films were similar in structure to the target silicone polymer. FTIR spectra

of the films (Figs. 3.3a and 3.5a) showed characteristic absorbances including CH

stretching, Si-CH3 deformation and Si-O-Si asymmetric split. The contact angles for

these films were close to that of silicone, indicating that the films were hydrophobic.

Thus the laser ablation of silicone at these energy density levels seemed to proceed by

lift-off and deposition of complete polymer chains on the substrate. Since silicone has no

significant UV absorption at 248 nm it is theorized that the laser energy is coupled into

the target photo-thermally causing a temperature rise. Optical microscopy examination

of the target after ablation revealed the presence of numerous bubbles in the subsurface of

the target. This indicated that the energy input was causing local heating and gas

generation in the target. The expansion of these bubbles could cause fragmentation of the

target and transport of polymer chains. If the heating rate is low enough the target









surface could be removed homogenously in a layer-by-layer evaporation process. Singh

has shown that subsurface heating effects can be important in the laser ablation of

materials which are poor UV absorbers [99].

However, if this were the only process occurring, the film composition would be

similar to that of silicone. As shown by XPS, the film surface chemistry deviated from

that of silicone. The films produced at low energy density were depleted in silicon and

had higher oxygen contents compared to the target polymer. The preferential etching of

methyl (CH3) and methylene(*CH2) groups from the target surface would result in the

presence of these species in the ablation plume. These charged species could then

recombine with polymer chains in the plume or on the substrate surface. Such a scenario

would produce silicone films with methylene bridges or longer alkyl groups hanging off

the Si-O-Si main chain. This scenario is partly supported by comparing peak ratios from

the FTIR spectra of the PLAD films produced at low energy density (Fig. 3.3a) to the

target polymer (Fig. 3.4). The Si-CH3/Si-O-Si peak ratio is lower for the PLAD film,

indicating a loss of methyl groups. However, the CH/Si-O-Si ratio is higher indicating

the retention of CH2 groups in the structure. The higher oxygen content can be attributed

to some free radicals on the polymer chain combining with oxygen on exposure to

atmosphere. The free radical concentration is low enough for the films to remain

hydrophobic.

FTIR spectra (Figs. 3.3b, 3.3c) revealed that at high energy density levels, the

films showed absorbances due to OH and SiH groups. Also the Si-O-Si peak broadening

(Fig. 3.5b, 3.5c) indicated a variety of distorted structures within the films. Thus

increasing the energy density of ablation produced more fragmentation of polymer









chains, leading to smaller chains being present in the plume. These fragments combined

in vapor or on the substrate surface to produce organosilicon films. The chain scission

also led to free radical production. The trapped radicals in the film could react with

oxygen and water vapor upon exposure to atmosphere, producing SiOH groups. The

other contributor to free-radical production could be the laser-ablation plume interaction.

At high energy density, the ablation product volume can be large enough to absorb a

significant portion of the incident laser irradiation. This can also lead to increased radical

production through photon absorption. The presence of SiH groups in the FTIR spectra

also indicated that Hf ions were present in the gas phase. It has been shown that the

production of H due to methyl group and H abstraction is the preliminary event in the

plasma polymerization of siloxane monomers [97]. Thus it is possible that some of the

fragmentation and growth mechanisms at high energy densities are similar to plasma

polymerization. Increasing methyl group abstraction with an increase in energy density

levels would account for the oxygen percentages seen in these films. The high oxygen

levels also led to these films being hydrophilic as evidenced by their low contact angles.

The change in the chemistry of the films as a function of energy density is also

reflected in the Si2p XPS peak location. Figure 3.9 showed that the Si2p peak location

shifted to higher binding energy with increasing energy density. At low energy densities

(100-150 mJ/cm2) the Si2p peak is located in the range 101.8-102.5 eV. This indicates

that the silicon is bonded to two or more carbon atom. In the energy density range of

200-600 mJ/cm2, the Si2p peak is located in the range 102.5-104 eV. This indicates that

the silicon is bonded to at least three oxygen atoms.









Suzuki et al have reported the use of pulsed lasers for the synthesis of silane

based polymeric films from poly(methylphenyl silane) (PMPS) [66]. They found that for

248 nm laser irradiation, most of the Si-Si bonds in the polymer were converted to Si-C

in the films and new bonds like Si-H were present. Also the deposited films suffered

molar mass reduction and showed a broader distribution than the starting polymer. The

molar mass distribution was found to be dependent on the energy density. Although the

starting polymer used was different, the trends found in the Suzuki study are similar to

what this research reports i.e. increased chain fragmentation with increasing energy

densities. It is interesting to note that the PMPS films were found soluble considering the

amount of reorganization which was seen in the structure.

Deposition rate measurements and electron micrographs provided a view of the

physical nature of the PLAD process. The deposition rate for the PLAD films was low

and showed little change upto approx. 180 mJ/cm2 (Fig. 3.11). This indicated that there

were more molecular (i.e. polymer chains) species involved in film growth and the level

of ablation remained relatively constant. As the energy density increased beyond a

certain limit, deposition rate showed a sharp linear increase. This led to deposition rates

which were quite high when compared with plasma polymerization processes [97], [98].

The deposition rate seemed to indicate the presence of a critical energy density for

steady-state ablation. It was also possible to vary the deposition rate by adjusting the

vacuum (i.e. pressure) level and substrate-target distance(Fig. 3.12). Such an effect has

been noted in the literature for the PLAD of numerous materials. The lower pressure

allowed the plume to progress further, increasing the mass flux incident on the substrate.

Figures 3.11 and 3.12 also demonstrated the importance of directionality and sample









position in the PLAD process. The deposition rates shown in Fig. 3.11 were measured

for samples which were placed away from the plume, leading to lower values. The

deposition rates shown in Fig. 3.12 were measured for samples kept directly in the plume

path leading to much higher rates of deposition.

The SEM analysis showed that at low energy densities, the PLAD film surfaces

were smooth (Fig. 3.13a). As the energy density increased beyond approx. 200 mJ/cm2,

the films contained more debris and also seemed to be composed of particulates. Thus

the deposition rates and SEM analysis indicated that the ablation levels upto approx. 200

mJ/cm2 were slow and the plume to be composed of smaller molecular species. As the

energy density increased beyond this limit, the polymer ablated off as clusters and

particles. These then recombined on the substrate to form a film. Thus it seemed that at

high energy density, the heating rate is high enough to cause rapid temperature rise and

inhomogenous removal of the target surface, while at low energy density, the heating rate

is low enough to cause a more uniform ablation of the target surface.

The pulse number study illustrated another important aspect of the nature of

silicone ablation. As seen from XPS analysis of films produced at different pulse

numbers, the film chemistry undergoes a change dependant on the pulse number. At the

lowest energy density of 100 mJ/cm2, it was seen that no films were produced after

irradiating the target for 1000 or 2000 pulses. Thus the ablation process did not start till

after 2000 pulses. During the initial pulses, the temperature rise to allow chain

evaporation or gas generation was not achieved. Such a phenomenon is known as

incubation and has been shown to occur in the laser ablation of PMMA [100]. On

continuing the energy input beyond 2000 pulses, the temperature could increase to a point









where ablation could occur. Thus the film produced with 4500 pulses showed a similar

composition to films produced at the same energy density, described earlier in this work.

As the energy density was increased, the incubation phenomenon occurred faster

and required a lesser number of pulses. This can be seen by the fact that the target

irradiated at 250 mJ/cm2 produced films comparable to those described earlier after only

2000 pulses. The target irradiated at 500 mJ/cm2 produced equivalent films after just

1000 pulses. Thus this study indicated that the ablation of silicone had a strong photo-

thermal component and that the incubation phenomenon seen for other polymers was also

prevalent in silicone.

The above discussion allows us to formulate a theory of silicone ablation at 248

nm based on empirical observations. The ablation behaviour can be divided into three

regimes (as shown in Fig. 3.10). In regime I (75-150 mJ/cm2), the ablation proceeds

slowly. The initial pulses serve to heat up the polymer and ablation (characterized as

polymer removal) starts only after a few 1000 pulses. Long polymer chains are

evaporated in a layer-by-layer fashion. There are also some etching and degradation

reactions associated with the ablation. An analogy can be drawn between the PLAD

process and Matrix Assisted Laser Desorption Ionization (MALDI) mass spectroscopy.

In the MALDI process, a laser beam is used to energize the sample consisting of an

active matrix and a polymer (or any other molecule of interest). The laser energy is

absorbed by the matrix, causing a temperature rise and desorption of the polymer. It has

been shown that intact polymer chains can desorb by this method and analyzed in the

mass spectrometer.









In regime II (150-220 mJ/cm2), the ablation behaviour stays the same, but there is

increased chain scission due to higher energy input. This regime can be characterized as

an intermediate one. Although the films produced are silicone-like, they can be

hydrophobic or hydrophilic depending on the concentration of free radicals and

subsequent level of oxygen incorporation. The contact angles in this regime show poor

reproducibility which is an indication of this phenomenon

In regime III, the ablation volume is large due to high heating rates, with the

polymer being ablated as clusters and particulates. There is large scale chain scission and

CH3 abstraction, producing films which are particulate and hydrophilic, albeit with a

substantial organic content. All the different characterization techniques like XPS,

profilometry and SEM seem to indicate that there is a change in the mode of ablation

beyond regime II. The critical energy density at which this change occurs can not be

accurately determined. Figure 3.16. shows a schematic representation of the above

described regimes.

It is believed that when using energy density levels lower than 75 mJ/cm2 (i.e.

below regime I), the primary mechanism of degradation would be etching rather than true

ablation. The low Si levels seen for films deposited in regime I would be further

depleted. A similar etching-ablation behaviour has been shown for PTFE laser ablation

at 248 nm. It was found that PTFE could be etched (characterized as loss of CF2 groups)

at low energy densities (~ 50 mJ/cm2). However, increasing the energy density beyond a

certain limit, it was observed that the etching behaviour decreased and the PTFE started

to ablate, (characterized as removal of polymer).




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