Alexandra da Conceição Teixeira Sepúlveda
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junho de 2014 Universidade do Minho Escola de Engenharia Alexandra da Conceição Teixeira Sepúlveda Use of Nanocomposites for Flexible Pressure Sensors
Transcript of Alexandra da Conceição Teixeira Sepúlveda
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Universidade do MinhoEscola de Engenharia
Alexandra da Conceição Teixeira Sepúlveda
Use of Nanocomposites for FlexiblePressure Sensors
junho de 2014
Tese de DoutoramentoPrograma Doutoral em Líderes para as Indústrias Tecnológicas
Trabalho efectuado sob a orientação doProfessor Doutor António José Vilela PontesProfessor Doutor Júlio César Machado VianaProfessor Doutor Luís Alexandre Machado da Rocha
Alexandra da Conceição Teixeira Sepúlveda
Use of Nanocomposites for FlexiblePressure Sensors
Universidade do MinhoEscola de Engenharia
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Acknowledgements _________________________________________________________________________
The accomplishment of this work would not have been possible without the support of
several people and entities. I would like to manifest my gratitude to all that directly or
indirectly have contributed for the completion of this dissertation.
I would like to express my sincere gratitude to Doctor António Pontes, scientific
supervisor, for his support, availability and encouragement for all these last years.
I am also grateful to Doctor Júlio Viana, co-supervisor, for provided guidance and
support throughout the course of my PhD.
I sincerely thank to Doctor Luís Alexandre Rocha, also co-supervisor, for his
incentive, guidance, discussions, for the knowledge transmitted during this research work
and for his ability to come up with creative ideas to resolve difficult research problems.
I would like to thank Eng. Cristina Oliveira and Doctor José Alberto Machado da
Silva who worked with me at different points of this research and have contributed with
valuable information.
I thank Eng. Maurício Malheiro for his precious help with some experimental tests.
I would like to thank to Doctor Rui Magalhães for having received me at PIEP (Pólo
de Inovação em Engenharia de Polímeros). This acknowledgment is also extended to Eng.
Nuno Vieira, Eng. Eduardo Pinto and Eng. Filipa Carneiro, for the collaboration and help
provided during my stay at PIEP.
At MIT, I would like to express my sincere gratitude to Professor Brian Wardle for
having received me in his labs, for his kind support and guidance.
I truly appreciate all the TELAMS members (Department of Aeronautics and
Astronautics, MIT) for their friendship and support during my journey at MIT.
Doctor Fabio Fachin, thanks for your teachings, support and guidance.
An indebted thank to Doctor Roberto Guzmán de Villoria for his assistance,
for any request or questions and for his friendship. He was always there to advise me and
to help me.
Acknowledgements
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I thank my parents for they encouragement, comprehension and love. Without their
support during all these years of study, I probably would not be here. To them, I am
forever grateful.
I would like to thank my boyfriend, José Afonso, for his encouragement, trust,
support and for listening my complaints. He gave me the final courage and strength to
finish this stage. Also, I thank him for the illustrations that he made for this dissertation.
Finally I would like to acknowledge the Portuguese Foundation for Science and
Technology (Fundação para a Ciência e Tecnologia, FCT), for the financial support
provided by the grant SFRH/BD/42922/2008. This work was supported by the ERDF
(European Regional Development Fund) through COMPETE (Operational
Competitiveness Program) and national funds through FCT in the framework of the
projects MIT-Pt/EDAM-EMD/0007/2008 and PTDC/EEI-ELC/1838/2012.
CNT-based polymer composite materials were developed with funding from Boeing,
EADS, Embraer, Lockheed Martin, Saab AB, Composite Systems Technology, Hexcel,
and TohoTenax through MIT´s Nano-Engineered Composite aerospace STructures
(NECST) Consortium. At MIT this research was supported (in part) by the U.S. Army
Research Office under contract W911NF-07-D-0004 and made use of the MIT MRSEC
Shared Experimental Facilities supported by the National Science Foundation under award
number DMR-0819762.
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Abstract _________________________________________________________________________
Use of Nanocomposites for Flexible Pressure Sensors
Polymer nanocomposites (PNCs) are defined as polymers bonded with nanoparticles to
create materials with improved properties. The development of this type of material is
rapidly emerging as a multidisciplinary research activity, since their final properties can
benefit many different fields of application, namely in the development of electrical
devices as studied herein.
A fabrication technique to produce conductive PNCs was developed in this work and
used to fabricate flexible capacitive pressure sensors. The process is based on vertically
aligned-carbon nanotubes (A-CNTs) embedded in a flexible and biocompatible matrix of
polydimethylsiloxane (PDMS). Thin A-CNTs/PDMS nanocomposite films (~ 400 µm)
were produced using wetting of as-grown A-CNTs with uncured PDMS and the resulting
nanocomposites were used to fabricate flexible pressure sensors. The sensing capability of
this A-CNTs/PDMS nanocomposite is attributed to the distinctive combination of
mechanical flexibility and electrical properties.
The fabricated nanocomposites were characterized and mechanical and electrical
properties evaluated. The PDMS is significantly modified by the reinforcing A-CNT
fibers, demonstrating non-isotropic (as opposed to the isotropic neat PDMS) elastic
properties all different than the PDMS (Young´s modulus of 0.8 MPa), including an
anisotropy ratio of 4.8 and increases in the modulus of A-CNTs/PDMS nanocomposites
over PDMS by more than 900 % and 100 %, in the CNTs longitudinal and transverse
directions, respectively. Regarding the electrical measurements, A-CNTs/PDMS
nanocomposites presented an electrical conductivity of 0.35 S/m. The rather low
conductivity does not compromise the developed capacitive sensor, but since passive
telemetry is required to measure and power the sensor, solutions to overcome this problem
were also studied.
The configuration of the developed flexible sensor is similar to typical silicon-based
capacitive pressure sensors. It is composed of three thin films, where two of them are
A-CNTs/PDMS nanocomposites (defining the diaphragm type electrodes) separated by a
film made of neat PDMS (defining the dielectric) and its operating principle is based on
Abstract
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the change of the deflection of the nanocomposite layers due to the change of an external
pressure. The developed flexible pressure sensors tested for pressures between 0 kPa and
100 kPa (operation required to measure the blood pressure in the aneurysm sac) showed
good linearity, mainly in the region near to the atmospheric pressure (pressure inside of
dielectric). In the same range, the dynamic response of the pressure sensors presented two
distinct comportments, which are related with the viscoelastic behaviour of the
PDMS-based nanocomposites used to build the pressure sensor. In addition, from the
experimental measurements of the prototype pressure sensors used, the sensitivities are in
the range of 2.5 – 20 fF/kPa.
To demonstrate feasibility for practical applications, the flexible sensor technology
was used in a biomedical application, more specifically in the context of abdominal aortic
aneurysms. The proposed implantable flexible pressure sensing system (capacitive sensor
plus inductor) consists of a mixed technology that uses A-CNTs/nanocomposites to build
the capacitor electrodes and flexible printed circuit board (PCB) technology to build an
inductor. The complete system was assessed by applying pressures varying from 0 kPa to
100 kPa. The results showed that the flexible sensors responded to pressure variations with
a well-defined characteristic curve and oscillation frequencies centered around 5.3 MHz
(the sensor receives energy and reflects back its oscillation frequency by means of
inductive coupling). In addition to these experiments, the performance of both pressure
sensors and reader system was assessed using a hydraulic model that mimics the
abdominal aorta. The proof-of-concept experiments validated the proposed telemetry
system functionality and showed feasibility in integrating the flexible, biocompatible and
passive sensor into a stent-graft, in order to the detect blood pressure in the aneurysm sac.
Finally, the developed technology to fabricate flexible pressure sensors based on
A-CNTs/PDMS nanocomposites proved successful in sensing applications and due to its
biocompatibility and versatility, can be used in other fields of application such as portable
medical devices and e-textiles (to monitor the vital signs of an individual, such as heart rate
and temperature, by using textile substrates with integrated electronics).
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Resumo _________________________________________________________________________
Utilização de Nanocompósitos para Sensores de Pressão
Flexíveis
Os nanocompósitos de polímeros (PNCs) são definidos como polímeros ligados com
nanopartículas, originando materiais com propriedades melhoradas. O desenvolvimento
deste tipo de material está a surgir rapidamente como uma atividade de investigação
multidisciplinar, uma vez que as suas propriedades finais podem beneficiar diferentes
áreas de aplicação, nomeadamente no desenvolvimento de dispositivos elétricos como
estudado aqui.
Neste trabalho foi desenvolvida uma técnica de fabricação de PNCs condutores e
utilizada para produzir sensores de pressão capacitivos. O processo é baseado em CNTs
verticalmente alinhados (A-CNTs) embutidos numa matriz flexível e biocompatível de
polidimetilsiloxano (PDMS). Membranas de nanocompósitos de A-CNTs/PDMS
(~ 400 µm de espessura) foram produzidas através do molhamento dos A-CNTs, utilizados
conforme crescidos, pelo PDMS e os nanocompósitos resultantes foram utilizados para
fabricar sensores de pressão flexíveis. A capacidade de deteção destes nanocompósitos é
conferida pela combinação distintiva da flexibilidade mecânica e propriedades elétricas.
Os nanocompósitos fabricados foram caracterizados e as suas propriedades
mecânicas e elétricas foram avaliadas. O PDMS é significativamente alterado pelo reforço
das fibras de A-CNT, demonstrando propriedades elásticas anisotrópicas (ao contrário do
PDMS puro que é isotrópico), incluindo uma razão de anisotropia de 4.8 e aumento do
módulo dos nanocompósitos de A-CNTs/PDMS relativamente ao PDMS acima dos 900 %
e 100 %, na direção longitudinal e transversa dos CNTs, respetivamente. Relativamente
aos testes elétricos, os nanocompósitos apresentaram uma condutividade elétrica de
0.35 S/m. O valor bastante baixo da condutividade não compromete o sensor capacitivo
desenvolvido, mas uma vez que é necessária telemetria passiva para medir e energizar o
sensor, foram também estudadas outras soluções para superar este problema.
A configuração do sensor flexível desenvolvido é semelhante aos típicos sensores de
pressão capacitivos de silício. Este é composto por três filmes finos, onde dois deles são
nanocompósitos de A-CNTs/PDMS (definem o diafragma) separados por um filme de
PDMS (define o dielétrico), e o seu princípio de funcionamento é baseado na deformação
Resumo
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das membranas de nanocompósitos devido a variações de uma pressão externa. Os
sensores de pressão flexíveis testados para pressões entre 0 kPa e 100 kPa (necessário para
medir a pressão sanguínea no saco do aneurisma) apresentaram boa linearidade,
principalmente na região próxima da pressão atmosférica (pressão no interior do
dielétrico). Para a mesma gama de pressões, a resposta dinâmica dos sensores apresentou
dois comportamentos distintos, que estão relacionados com o comportamento viscoelástico
dos nanocompósitos utilizados para fabricar o sensor. Além disso, testes experimentais
realizados com sensores de pressão protótipos mostraram sensibilidades na gama de
2.5 – 20 fF/kPa.
De forma a demonstrar a viabilidade para aplicações práticas, foi apresentada a
utilização da tecnologia do sensor de pressão numa aplicação biomédica, mais
concretamente no contexto de aneurismas na aorta abdominal. O sistema flexível de
monitorização da pressão proposto (sensor capacitivo mais indutor) consiste numa
tecnologia mista que utiliza nanocompósitos de A-CNTs/PDMS para construir os elétrodos
do condensador e uma placa flexível de circuito impresso para construir o indutor. O
sistema completo foi avaliado para variações de pressão entre 0 kPa e 100 kPa. Os sensores
flexíveis responderam às variações de pressão com uma curva característica bem definida e
frequências de oscilação centradas nos 5.3 MHz. O desempenho dos sensores de pressão e
do sistema de leitura foram também avaliados utilizando um modelo hidráulico que imita a
aorta abdominal. Os testes experimentais validaram a funcionalidade do sistema
telemétrico proposto e mostraram viabilidade em integrar o sensor flexível, biocompatível
e passivo num stent-graft, de forma a detetar variações da pressão sanguínea no saco do
aneurisma.
A tecnologia desenvolvida para fabricar sensores de pressão flexíveis, que se baseia
em nanocompósitos de A-CNTs/PDMS, provou ser bem-sucedida em aplicações de
deteção de pressão e devido à sua biocompatibilidade e versatilidade, esta tecnologia pode
ser utilizada em outras áreas de aplicação, tais como dispositivos médicos portáteis e
têxteis eletrónicos (para controlar os sinais vitais de um indivíduo, como a frequência
cardíaca e temperatura, utilizando substratos têxteis com eletrónica integrada).
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Contents _________________________________________________________________________
Acknowledgements…………………………………………………………...….…….….iii
Abstract………………………………………………………………………….………..v
Resumo………………………………………………………………………………….vii
Table of Contents………………………………………………………………………..ix
List of Figures…………………………………………………………………………..xv
List of Tables…………………………………………………………………………...xxi
List of Symbols………………………………………………………………………..xxiii
List of Abbreviation………………………………………………………………….xxvii
Chapter 1. Introduction………………………………………………………………...1
1.1 Polymer Nanocomposites……………………………………………………2
1.1.1 CNTs/PDMS nanocomposites…………………………………………3
1.2 Potential Applications of CNTs/PDMS Nanocomposites………..…….……5
1.3 Thesis Overview……….………………………………………...…………...8
1.3.1 Motivation and objectives……………………………….………..........8
1.3.2 Thesis organization…………..…..………………………..….……........11
References…………………………………………………………………….……….13
Chapter 2. Background and Prior Work……………………………………………...19
2.1 Polydimethylsiloxane (PDMS)………………..……….……………….……..20
2.1.1 Structure and synthesis…………..……………………….…………..20
2.1.2 General properties…………………………..……………….………..22
2.1.3 Mechanical properties………………………..……...……….……........23
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2.2 Carbon Nanotubes Fundamentals……………..…..……………..…...……..26
2.2.1 Carbon nanotubes morphology….………..………………….….…….26
2.2.2 Physical properties of carbon nanotubes…………..…...……………....28
2.2.3 Carbon nanotubes growth techniques………………..….……….........33
2.3 Aligned CNTs/PDMS Nanocomposites…………...…………..……………37
2.4 Sensors and Actuators Based on Carbon Nanotubes…….…….……...…..40
2.5 Biocompatibility of PDMS and CNTs Materials: Perspectives of Their
Application in Medical Implants………………...……...….…………..….43
2.6 Summary……………………………………………………………………47
References…………………………………………………………….…….……49
Chapter 3. Fabrication and Characterization of A-CNTs/PDMS
Nanocomposites………………………………………………………..….69
3.1 Introduction…………………………………………...……………………...70
3.2 Synthesis of MWCNTs…………………………………………...……..….70
3.3 Fabrication of A-CNTs/PDMS Nanocomposite Films………..……….…..74
3.4 Thermal Characterization and Results………………….……….………..76
3.4.1 Differential scanning calorimetry…………………….……..…….….76
3.5 Morphological Characterization and Results……………..…………….….77
3.5.1 Scanning electron microscopy………………………………....….….77
3.6 Mechanical Characterization and Results……………..………….…….….79
3.6.1 Tensile tests…………………..……………………………..….…….79
3.6.2 Dynamic mechanical analysis…………………………..………….…84
3.7 A-CNTs/PDMS Nanocomposite Full Constitutive Law………….………86
3.8 Electrical Characterization and Results………………….…….……...….89
3.8.1 Conductivity measurement…………………..………….…..………..89
3.9 Summary……………………………………….…………..….………..…..….94
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References…………………………………………………………………….….96
Chapter 4. Fabrication and Characterization of Flexible Capacitive Pressure
Sensors…………………………………….…………………………………101
4.1 Introduction………………………………………....………………...………...102
4.2 Design of the Flexible Capacitive Pressure Sensor………………...……104
4.2.1 Sensor model………………………………….……….…..………...104
4.2.1.1 Mechanical domain……………………….………………….104
4.2.1.2 Electrostatic domain…………………………………..…...…106
4.2.2 Telemetry model………………………………….....……………......107
4.2.3 Finite Element Modeling…..…………………………….………......110
4.2.3.1 Simulation setup of pressure sensors……………….....…...…..110
4.2.3.2 Simulation results…………………….……………...……….112
4.3 Development of Flexible Pressure Sensor……………………..…...……..116
4.3.1 Fabrication process………….....…………………………...……...…...116
4.4 Sensor Characterization and Results………………..…...…….....……….119
4.4.1 Finite Element Model………………………….……..……..………119
4.4.2 Capacity and pressure changes measurement…………...…..………122
4.5 Summary….…………….…………………………...……..….…………..….128
References………………………………………………………………..….…….130
Chapter 5. Biomedical Applications: Case Study – Abdominal Aortic
Aneurysms……………………………………...………………...………..133
5.1 Introduction……………………………………….………………...………..134
5.2 Abdominal Aortic Aneurysm….…………………………..….....……….….134
5.2.1 Current available treatment options……………..…………...….......136
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5.2.2 Actual stent-grafts main problems and needs for pressure monitoring of
AAAs…………………………………………………………..………139
5.3 Measurement of Aneurysm Sac Pressure………………………………..141
5.3.1 Extra-arterial blood pressure sensors……………………………......141
5.3.2 Intro-arterial blood pressure sensors: Aneurysms.……………..……145
5.3.3 A new proposal: Smart stent-graft………..………..….………..…….150
5.3.3.1 Flexible capacitive pressure sensor requirements……..…..….152
5.3.3.2 Telemetry system architecture…………………..……………....155
5.4 New Approaches for the Implantable Sensor Fabrication………..……..157
5.4.1 Fabrication of embedded inductors………………….....……………..157
5.4.2 Fabrication of inductors using a flexible printed circuit board
technology……………………………………………………...……158
5.4.2.1 Flex-PCB process flow……………………...…………........….159
5.4.3 Readout telemetry implementation………………………………….162
5.5 Experimental Measurements and Results…………………………...…...163
5.5.1 Experimental results obtained using a vacuum chamber……………163
5.5.2 Sensors performance in a hydraulic test bench……………………...166
5.6 Summary...........................................................................................................169
References………………………………………………………………………171
Chapter 6. Conclusions and Recommendations…………………………………….179
6.1 A-CNTs/PDMS Nanocomposite Films…………………………………...180
6.2 Flexible Capacitive Pressure Sensors………………………………….…181
6.3 New Flexible Pressure Sensor for the Treatment of Abdominal Aortic
Aneurysms………………………………………………………………...182
6.4 Recommendations for Future Work………………………………….….183
6.4.1 As-grown CNTs……………………………………………………..183
6.4.2 A-CNTs/PDMS nanocomposites……………………………………184
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6.4.3 Development of new sensor designs………………………………...185
Appendix…………………………………………………………………….……...….187
A.1 Examples of Mask Pillars………………………………………………...188
A.2 Substrates Shape and Growth Times Used for the Optimization
Process………………………………………………………………….………..189
A.3 Estimate of A-CNTs Volume Fraction………………………..………….190
A.4 Datasheet of PDMS…………………………………………………..……191
A.5 List of Publications……………………………………………………….194
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List of Figures _________________________________________________________________________
Figure 1.1: Surface area/volume relations for varying reinforcement geometries…...3
Figure 1.2: Atomic model of a multi-walled carbon nanotube (MWCNT)…………..4
Figure 1.3: Implantable strain sensor with wireless sensing capability for long-term
monitoring………………………………………………………….……..7
Figure 1.4: Schematic outline of the main objectives presented in this thesis….…...11
Figure 2.1: Visual representation of the PDMS mixing process...…………....…….20
Figure 2.2: Polydimethylsiloxane. a) Chemical structure combining both organic and
inorganic groups; b) Structural representation showing the shielding of
the main chain (Si-O groups) by the methyl groups…….….…..…..……21
Figure 2.3: Synthesis of PDMS through hydrosilylation…...…………………….....21
Figure 2.4: Speed-thickness correlation for different formulations of base
polymer/curing agent combined at different weight ratios and spin
coating speeds……………………………………………..…………...……..24
Figure 2.5: Effect of different formulations of the PDMS on the Young’s modulus,
for a crosshead speed of 5 mm/min………………...……………...……25
Figure 2.6: Atomic structure of a single graphene sheet…………………….……...26
Figure 2.7: Schematic of single-walled carbon nanotubes (SWCNTs) and multi-
walled carbon nanotubes (MWCNTs) showing the common dimensions
of length, width and separation distance between graphene layers in
MWCNTs.……………………………………………………..…...…...27
Figure 2.8: Atomic structure illustrations showing the chiral vector (Ch) in a
hexagonal graphene sheet, defined by unit vectors a1 and a2, and the
chiral angle θ……………………………..…..………..…………………28
Figure 2.9: Electrical and thermal conductivities of individual CNTs normalized by
density, in comparison with different materials……..………………….29
Figure 2.10: Schematic representation of an arc discharge apparatus for the synthesis
of different carbon structures…………………………..……………….33
Figure 2.11: Schematic representation of a laser ablation reactor……………..……..34
Figure 2.12: Schematic diagram of the CVD apparatus……………………………...35
List of Figures
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Figure 2.13: Tip-growth model and base-growth model of CNT growth
mechanism………………………………………………..………………36
Figure 3.1: CVD system for growing A-CNTs forests……………………………...71
Figure 3.2: SEM images of as grown A-CNTs forests. a) Dot shape, b) Y-shape and
c) magnified region of b)…………………………….…………………..72
Figure 3.3: CNTs grown on Si substrate. a) Optical image of an A-CNTs forest, b)
SEM image of A-CNTs forest and c) SEM image showing the alignment
of the CNTs……………………………………………..………………..73
Figure 3.4: Fabrication process flow for the development of a flexible A-
CNTs/PDMS nanocomposite………………………………………...…75
Figure 3.5: Flexible PNC film made of A-CNTs embedded in a polymeric matrix of
PDMS…………………………………………………………….……..75
Figure 3.6: Representative DSC thermograms of pure PDMS and A-CNTs/PDMS
nanocomposites……………………………………………………...….77
Figure 3.7: SEM image of as-grown A-CNTs. (b) and (c) are magnified regions of
(a)……………………………………………………………………..…78
Figure 3.8: SEM images of A-CNTs/PDMS nanocomposites revealing the CNTs
structure embedded in the PDMS substrate. (a) Top-view and (b) parallel
cross-sectional cut………………………………….……………….…..78
Figure 3.9: SEM image of a top-view surface of pure PDMS………..……………..79
Figure 3.10: Schematic of the CNTs orientation inside the PDMS matrix (legend:
W – width, H – height and P – Load direction)……...…………………80
Figure 3.11: Optical images showing the changes before tensile test (unstretched) and
during the test, for pure PDMS and A-CNTs/PDMS samples (with the
CNTs oriented transverse to the load – 1,2 plane and 1,3 plane, and
parallel to the load – 3,2 plane)……………………………………...….81
Figure 3.12: Illustration showing the points P1, P2, P3 and P4. a) Beginning of the test
and b) displacement in the x- and y-axis during the tensile test…….......82
Figure 3.13: Representative stress-strain curves of pure PDMS and reinforced
A-CNTs/PDMS nanocomposites in the longitudinal (3,2 plane) and
transverse (1,2 plane) directions. Longitudinal is in the 3-direction and
transverse the 1-, or 2-direction……………………………………...…82
Figure 3.14: Schematic of the shear mode (dual sample) experiment………………..84
List of Figures
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Figure 3.15: a) Storage modulus – G’ and b) tangent delta – tan δ, as function of
frequency for pure PDMS and A-CNTs/PDMS samples………….……85
Figure 3.16: Schematic of a Van der Pauw configuration used in the measurement of
the two characteristic resistances RA and RB……………...……………..89
Figure 3.17: Histogram of the electrical conductivity measured on the A-CNTs/PDMS
nanocomposite samples…………………………………………..……..91
Figure 3.18: Digital image of the A-CNTs/PDMS nanocomposite samples prepared to
measure the resistivity while performing the tensile tests………………92
Figure 3.19: Digital image of the experimental test setup……………………...…….92
Figure 3.20: Measured resistance changes of the A-CNTs/PDMS nanocomposites
during tensile mechanical tests, with initial resistance (R0) of 0.3
MΩ………………………………………………….……………………...93
Figure 4.1: Illustration of a capacitive pressure sensor based upon a parallel plate
arrangement………………………………………………………………..102
Figure 4.2: Schematic of the square (sidelength = 2a) pressure sensor. a) 3D view
and b) section cut B-B……………………………………….……..……105
Figure 4.3: Cross section of the generic model for a deflectable diaphragm……….105
Figure 4.4: Blocks diagram of the telemetry system…………………….…………..108
Figure 4.5: Schematic of the reader circuit………………………………………….109
Figure 4.6: Flexible pressure sensor (S6) model for FEM simulations……….……110
Figure 4.7: Model showing the mesh type, boundary conditions and applied
pressure………………………………...…………………………………111
Figure 4.8: Diaphragm deformation in the y-axis. a) dielectric size of 4x4 mm2,
b) dielectric size of 6x6 mm2 and c) dielectric size of 8x8 mm
2………113
Figure 4.9: Displacement of pressure sensors with different dielectric dimension after
applying the pressure Pout……..………………..………..………………113
Figure 4.10: Displacement of pressure sensors for both orthotropic and isotropic
material properties. Sensors with: a) dielectric size of 4x4 mm2,
b) dielectric size of 6x6 mm2 and c) dielectric size of 8x8 mm
2……....115
Figure 4.11: Fabricated acrylic molds to build the pure PDMS membranes
(dimensions in mm)…………………………………….………………117
List of Figures
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Figure 4.12: Schematic of the flexible pressure sensor composed of three thin
membranes (top and bottom layers are A-CNTs/PDMS nanocomposites
and the middle one is made of pure PDMS)…………………..………...117
Figure 4.13: Flexible pressure sensor. a) Digital image taken into the vacuum chamber
and b) optical image of the cross section (magnification: 0.67 x 6.7)....118
Figure 4.14: Digital image of fabricated pressure sensors showing its flexibility..…118
Figure 4.15: Finite element model parameters………………………...……………...119
Figure 4.16: Finite element modeling algorithm for coupled domain analysis….…..121
Figure 4.17: Example of the obtained mesh model…….………………..……………121
Figure 4.18: Bending profile for the applied pressure Pout…………………….…....122
Figure 4.19: Capacitive changes vs. pressure changes, assuming ambient pressure
inside the dielectric……………………………………………………….123
Figure 4.20: Iterative algorithm to compute the capacitive changes of flexible pressure
sensors………………………………………….…………………………124
Figure 4.21: 3D drawing of the pressure chamber personalized for the capacitive
changes measurement of the flexible sensors………………………….125
Figure 4.22: Experimental results and comparison with analytical model. a) Area of
the dielectric is 4 x 4 mm2 and the distance between electrodes is 320
µm, b) area of the dielectric is 5 x 5 mm2 and the distance between
electrodes is 130 µm and c) area of the dielectric is 8 x 8 mm2 and the
distance between electrodes is 210 µm……………………………..…126
Figure 4.23: Dynamic response of the flexible pressure sensor………….…………127
Figure 5.1: Illustrations of: a) normal aorta, b) thoracic aortic aneurysm (TAA) and
c) abdominal aortic aneurysm (AAA)…………………………………135
Figure 5.2: Different stages of AAA degeneration………………………………...136
Figure 5.3: Open surgery technique for AAAs to insert a synthetic graft…………137
Figure 5.4: Stent-graft deployment sequence in the endovascular aneurysm repair
(EVAR) procedure…………………………….……………………….…137
Figure 5.5: Open repair and EVAR estimated costs per patient…………………...138
Figure 5.6: Illustration of the basic principle of tonometric blood pressure
measurement…………………………………………………………...142
Figure 5.7: Photograph and sensor calibration curve of the tonometric pressure
sensor…………………………………………………………….…….142
List of Figures
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xix
Figure 5.8: Implantable blood pressure sensor monitoring system and cross-section
of the cuff around the vessel…………………………………………...143
Figure 5.9: Cross-section of new monitoring cuff and prototype of implantable blood
pressure cuff with rigid ring…………………………………………...144
Figure 5.10: Schematic and photograph of flexible CEFP platform…………..……144
Figure 5.11: Schematic and photograph of flexible blood pressure sensor
module………………………………………..……………………………145
Figure 5.12: Impressure AAA sac pressure transducer (Remon Medical)……….…146
Figure 5.13: EndoSure wireless pressure sensor (CardioMEMS, Inc.)…………..…147
Figure 5.14: Flexible wireless pressure sensor for acute use………………………..147
Figure 5.15: Flexible wireless pressure sensor for chronic use……………………..148
Figure 5.16: TPS telemetric pressure sensor: system concept and readout
station……………………………………………………………...…….148
Figure 5.17: The HYPER-IMS implantable pressure sensor………………………..149
Figure 5.18: HYPER-IMS blood pressure measurement system…………………....150
Figure 5.19: Smart stent-graft concept and main building blocks………….……..…151
Figure 5.20: Main elements that compose the flexible pressure sensor: a flexible
substrate made of PDMS and the conductive elements based on
A-CNTs………………………………………..…………………..………152
Figure 5.21: Block diagram of the monitoring system……………….……………….155
Figure 5.22: Inductors fabrication. a) Steel master mold for oval spiral inductors and
b) Inductor prototypes………………………………………………....158
Figure 5.23: Overview of the final process for fabrication of the passive LC
network…………………………………………...……………………...160
Figure 5.24: Flex-PCB panel with six different geometries after fabrication (Step 1)
and a digital image showing its flexibility………………………….…160
Figure 5.25: Optical image of a square geometry (S6) for capacitive sensor and details
of the copper inductors……………………………………………...…161
Figure 5.26: Digital images of the final assembled flexible pressure sensors………162
Figure 5.27: Prototype showing the reader circuit with receiver inductor placed next to
the sensor’s inductor……………………………………………...……163
Figure 5.28: Experimental test setup for measurements in the vacuum
chamber…..............................................................................................164
List of Figures
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Figure 5.29: Oscillation frequency results from all trials with the S4 pressure
sensor……………………………………………………………….….165
Figure 5.30: Oscillation frequency results from all trials with the S6 pressure
sensor………………………………………….…………………….….165
Figure 5.31: Stent-graft (on the left) and the pressure sensor placed on top of the
graft………………………………………………………………………167
Figure 5.32: Experimental test setup for measurements with the hydraulic
model………………………………………………………………….….167
Figure 5.33: Diagram of the hydraulic test bench………………………………...…168
Figure 5.34: Measure pressure (dotted line) and oscillation frequency (solid line)
while the water tap valve was being opened and closed as in the cardiac
cycle………………………………….……………………………………168
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List of Tables _________________________________________________________________________
Table 2.1: Dielectric properties of polydimethylsiloxane (PDMS) and high density
polyethylene (HDPE)…..………………..…………..………………...…23
Table 2.2: Elastic modulus and tensile strength values of both SWCNTs and
MWCNTs, more frequently referenced in the literature.…..………….…32
Table 2.3: Summary of materials’ selection for long-term implantable medical
devices………………………………………………………..……………45
Table 3.1: CNTs growth recipe for substrates with patterned catalyst in dot shape and
Y-shape……………………………...………………………………...….72
Table 3.2: CNTs growth recipe for substrates with patterned catalyst in rectangular
shape…………………………………………………………………..….73
Table 3.3: PDMS main properties (10:1 weight ratio)….………………………...…74
Table 4.1: Materials’ properties of the pressure sensors used in FEM
simulations……………………………………….………………………...111
Table 4.2: Isotropic material’s properties of the pressure sensors used in FEM
simulations…………………………………………………….……..……114
Table 4.3: Dimensions and material properties of the pressure sensors used in the
FEM model……………………………………………………….……...120
Table 5.1: Estimated annual risk of rupture of AAAs for a given diameter…….…135
Table 5.2: Complications involving stent-grafts…………………………………...139
Table 5.3: Requirements for post-EVAR blood pressure monitoring…………...…153
Table 5.4: ANSI/AAMI SP-10-1993 national standard for electronic or automated
sphygmomanometers………………………………………………........156
Table 5.5: Dimensions of the flex-PCB for different geometries……………..…….161
Table 5.6: Sensitivity of the detected oscillation frequency obtained for S4 and S6
sensors………………………………………………………………...…166
Table A.2: Si wafer substrates and growth time used in the first growths of A-CNTs
forests…………………………………………………………………....189
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xxiii
List of Symbols _________________________________________________________________________
Symbol Description Unit
Au Gold [ - ]
Al Aluminum [ - ]
Al2O3 Alumina oxide [ - ]
Capacitance [F]
C Carbon [ - ]
CH3 Methyl [ - ]
CH4 Methane [ - ]
C2H2 Acetylene [ - ]
C2H4 Ethylene [ - ]
C6H6 Benzene [ - ]
Co Cobalt [ - ]
CO Carbon monoxide [ - ]
Cu Copper [ - ]
CS Variable capacitor [F]
E Young’s modulus [Pa]
Fe Iron [ - ]
G Shear modulus [Pa]
GF Gauge factor [ - ]
G* Complex shear modulus [Pa]
G’ Storage modulus [Pa]
G’’ Loss modulus [Pa]
H0 Initial height [m]
H2 Hydrogen [ - ]
He Helium [ - ]
I Current [A]
Lp Reader’s inductor [H]
LS Sensor’s inductor [H]
Ni Nickel [ - ]
List of Symbols
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O Oxygen [ - ]
Pd Palladium [ - ]
Pt Platinum [ - ]
P0 Inside pressure [Pa]
Pout Outside pressure [Pa]
R Resistance [Ω]
Rp Parasitic resistance from reader’s inductor [Ω]
Rs Parasitic resistance from sensor’s inductor [Ω]
RSh Sheet resistance [Ω]
R0 Initial resistance [Ω]
Si Silicon [ - ]
SiO2 Silicon dioxide [ - ]
tan δ Loss tangent [ - ]
Tg Glass transition temperature [ºC]
Tm Melting temperature [ºC]
V Voltage [V]
W0 Initial width [m]
w0 Deflection [m]
XeF2 Xenon difluoride [ - ]
Y yttrium [ - ]
a1, a2 Unit vectors of the hexagonal lattice [ - ]
d0 Gap between electrodes [m]
Ø Diameter [m]
ΔH Change of height [m]
ΔW Change of width [m]
fosc Oscillation frequency [Hz]
ε Strain [%]
ε0 Permittivity of free space [F/m]
εr Relative permittivity [ - ]
εij Strain tensor [ - ]
φ Angle of deflection [ - ]
j Non-zero integer [ - ]
k Coupling factor [ - ]
List of Symbols
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υ Poisson´s ratio [ - ]
n Number of repeating monomer units [ - ]
n, m Lattice translation indices [ - ]
ρ Electrical resistivity [Ω.m]
Sijkl Compliance tensor [ - ]
σ Electrical conductivity [S/m]
σkl Stress tensor [ - ]
t Thickness [m]
θ Chiral angle [º]
r Radius [m]
rr Reader’s coil radius [m]
rs Sensor’s coil radius [m]
vdif Sensor’s oscillation frequency signal [V]
vvs Square wave [V]
vo – vto Differential signal [V]
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List of Abbreviations _________________________________________________________________________
Term Designation
AAA Abdominal Aortic Aneurysm
AAS Atomic Absorption Spectroscopy
AFM Atomic Force Microscope
AVR Average Variation Rate
A-CNTs Vertically Aligned-Carbon Nanotubes
BioMEMS Biological Microelectromechanical Systems
CEFP Chip Embedded Flexible Packaging
CNC Computer Numerical Control
CNTs Carbon Nanotubes
CTA Computer Tomography Angiography
CVD Chemical Vapor Deposition
DLC Diamond-like Carbon
DMA Dynamic Mechanical Analysis
DNA Deoxyribonucleic Acid
DSC Differential Scanning Calorimetry
ECG Electrocardiogram
EEG Electroencephalography
EVAR Endovascular Aneurysm Repair
FDA Food and Drug Administration
FEP Fluorinated Ethylene Propylene
FFT Fast Fourier Transform
FTIR Fourier Transform Infrared Spectroscopy
GC-MS Gas Chromatography Coupled with Mass Spectroscopy
GF Gauge Factor
GPC Gel Permeation Chromatography
GPIB General Purpose Interface Bus
HDPE High Density Polyethylene
List of Abbreviations
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HRTEM High-Resolution Transmission Electron Microscope
IR Infrared
LCP Liquid Crystal Polymer
MD Molecular Dynamics
MEMS Microelectromechanical Systems
MIT Massachusetts Institute of Technology
MRA Magnetic Resonance Angiography
MRI Magnetic Resonance Imaging
MST Microsystems Technologies
MWCNTs Multi-Walled Carbon Nanotubes
NEMS Nanoelectromechanical Systems
NMR Nuclear Magnetic Resonance Spectroscopy
PCB Printed Circuit Board
PCD Programmed Cell Death
PE Polyethylene
PE-CVD Plasma-Enhanced Chemical Vapor Deposition
PDMS Polydimethylsiloxane
PMMA Polymethylmethacrylate
PNCs Polymer Nanocomposites
PRC Polymerase Chain Reaction
PTFE Polytetrafluoroethylene
PVD Physical Vapor Deposited
RF Radio Frequency
SCCM Standard Cubic Centimeters Per Minute
SEM Scanning Electron Microscopy
SFC Supercritical Fluid Chromatography
SHM Structural Health Monitoring
SWCNTs Single-Walled Carbon Nanotubes
TAA Thoracic Aortic Aneurysm
TELAMS Technology Laboratory for Advanced Materials and Structures
TEM Transmission Electron Microscope
TGA Thermogravimetric Analysis
List of Abbreviations
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USB Universal Serial Bus
UV Ultraviolet
XPS X-Ray Photoelectron Spectroscopy
Chapter 1
Introduction
Introduction
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2
1.1 Polymer Nanocomposites
In the area of nanotechnology, polymer nanocomposites (PNCs) are receiving significant
attention both in the recent literature and industry. The constituent materials of
nanocomposites are classified into matrix and filler phases, where the first is the
continuous element and frequently present in the largest quantity. The function of the filler
is to improve the properties of the matrix, namely the electrical and mechanical properties
[1.1]. In this new class of materials, the filler phase has at least one dimension of less than
100 nm, providing nanocomposites with huge interfacial area/volume ratio between matrix
and filler and consequently, stronger molecular interactions between the two constituents.
Compared to other classes of materials, polymer is a versatile material and owns
many unique properties, such as mechanical flexibility, low density, reasonable strength,
optical transparency, chemical stability, biocompatibility and easy processing. These
characteristics allow PNC to be used in several applications like aerospace, automotive,
coatings, adhesives, packing materials, microelectronic packaging, optical integrated
circuits, sensors, drug delivery, medical devices, etc. [1.2]. However, for many engineering
applications the mechanical properties of the polymers are not enough. Also, most of
polymers have lack of electrical conductivity, confining these materials to a structural
element in most applications [1.3]. For this reason, there is a growing demand towards new
polymeric materials with improved characteristics.
PNCs can be processed similarly to conventional polymers, where blending of
different class of polymers and/or inorganic fillers can be used to fabricate new materials
with enhanced properties [1.4]. Several materials have been developed incorporating
nano-sized filler materials in polymer substrates and demonstrated superior mechanical,
thermal and barrier properties in comparison with non-filled polymers [1.5].
Almost all polymers (thermoplastics, thermosets and elastomers) have been used to
fabricate PNCs, as well as a large variety of nano-reinforcements. Generally, the three
classes of nanomaterials currently used are particle (such as silica nanoparticles and carbon
black), layered and fibrous materials (such as nanofibers and carbon nanotubes) [1.6].
Figure 1.1 illustrates the common reinforcement geometries and their respective surface
area/volume ratios. For the fibrous and layered nanomaterials, the surface area/volume is
Introduction
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3
controlled by the first term in the equation since the second term (2/l and 4/l) has a little
influence compared to the first term (if r << l as in the case of nanomaterials such as
CNTs and graphene, respectively). Therefore, changing the particle diameter, layer
thickness or fibrous material diameter from micrometer to nanometer scale, will change the
surface area/volume ratio by three orders of magnitude [1.7].
Figure 1.1: Surface area/volume relations for varying reinforcement geometries. Used with
permission and adapted from reference [1.7] (Thostenson ET et al. Composites Science and
Technology 2005; 65:491-516, Copyright (2005) Elsevier).
Fillers characterized by high aspect ratio (ratio of length to the cross-sectional
dimension) can provide either isotropic or anisotropic properties, depending on the
orientation of the filler in the host matrix of the nanocomposite.
The final properties of nanocomposite materials depend not only on the properties of
the constituents used (polymer matrix and filler), but also on their morphology and
interfacial characteristics [1.8]. Also, depending on the processing method, considerable
differences in the nanocomposite properties may be obtained [1.9].
1.1.1 CNTs/PDMS Nanocomposites
Polymer-based nanocomposites are formed through a combination of a polymer matrix and
nano-scale reinforcing materials. Several polymers with distinct characteristics can be used
Introduction
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4
as the matrix of a nanocomposite. Nevertheless, a correct selection of a matrix for a given
nanocomposite should take into account the purpose of the application.
A range of thermoplastic polymers have been used as a hosting matrix to build strain
sensors [1.10-1.14], with demonstrated superior tensile strain when compared with typical
metallic strain gauges [1.15]. However, when flexibility is required for an application,
elastomeric materials are commonly used as polymer matrix. Polydimethylsiloxane
(PDMS) is the most widely used silicone-based elastomer, known for its superior
mechanical elasticity and other remarkable rheological properties. PDMS is available in
the pure form, allowing its use for numerous applications. Also, among the several
characteristics PDMS is of easy molding, making it a good material to use in various
replication methods.
Carbon nanotubes (CNTs) have appeared as an excellent nanofiller, mainly due to
their small size and unique physical properties [1.16]. More particularly, their excellent
electrical and mechanical properties [1.17-1.21] have been raising great interest in
fabricating CNTs-based composites for electronic and structural applications. CNTs
possess a high aspect ratio due to its long tubular structure (Figure 1.2), exhibiting high
electrical conductivity in comparison with other conductive nanofillers with lower aspect
ratios. Also, an entangled CNT network can actuate as electrically conductive pathway or
structural reinforcement element [1.22] and its incorporation into a polymeric matrix can
provide multi-functionalities to the nanocomposites.
Figure 1.2: Atomic model of a multi-walled carbon nanotube (MWCNT) [1.23]. Used with
permission (Copyright (2003) by The American Physical Society).
Introduction
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5
The first CNTs/epoxy nanocomposites were prepared by Ajayan et al. in 1994 [1.24]
and after that, several matrices have been studied to embed efficiently the CNT network.
Wang et.al [1.25] investigated different approaches for embedding CNTs structures into
polymer matrices. The results demonstrated that the combination of the CNTs properties
with functional polymers lead to materials with uncommon properties, such as electrical,
mechanical, optical and magnetic properties. Other studies reported the improvement of
material properties with just a small amount of CNTs, preserving the flexibility of the
polymers [1.26]. The most available application for CNTs-polymer nanocomposites is
related with the electrical properties of CNTs, because their native conductivity makes
CNTs an excellent option for changing the conductive properties of polymers.
Recently, there has been a great effort focused in the study of CNTs/PDMS
nanocomposites. PDMS is an electrical and thermal insulator material and alone, its use is
limited to structural applications [1.27, 1.28]. The incorporation of CNTs into the PDMS
matrix can enhance its electrical and mechanical properties with just a small amount of
CNTs. However, factors such as the type of CNTs, aspect ratio, CNT weight percentage
and processing techniques can influence the final properties of the resulting CNTs/PDMS
nanocomposites.
1.2 Potential Applications of CNTs/PDMS Nanocomposites
The excellent variety of CNTs properties have resulted in a new era of advanced
multifunctional materials. Specifically, embedding CNTs in a polymeric matrix provides
nanocomposites highly suitable for engineering applications.
Since 2005, a small number of works have been reported that by embedding CNTs in
pure PDMS, the resulting nanocomposite has several attracting properties that can be used
in microelectromechanical systems (MEMS) for sensing and actuation. In fact,
CNTs/PDMS nanocomposites has diverse fields of applications, such as electrical devices,
optics, biocompatible thin coatings and drug delivery systems, but this type of
nanocomposites have mainly been used as strain gauges and pressure sensors.
CNTs/PDMS nanocomposites are piezoresistive materials, meaning that their nominal
Introduction
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6
resistivity changes when a mechanical stress is applied. Also, a high sensitivity is expected
from these materials due to its Young’s modulus (1.7 MPa at 1.0 wt. % CNTs) [1.29].
CNTs/PDMS nanocomposites have been used to fabricate transducers, namely
resistive and capacitive sensors [1.30-1.36], in order to monitor physical or chemical
parameters. In the case of resistive sensors, parameters such as strain, voltage, magnetic
field, light, gas/liquid concentration, can induce a change in electrical resistance. On the
other hand, capacitive sensors can provide contact or noncontact measurement of several
physical or chemical quantities expressing position, distance, separation, proximity,
acceleration, pressure, force, chemical substances, particles and biocells. [1.37].
In addition to the sensing properties of CNTs/PDMS nanocomposites, they also
present actuation characteristics. This kind of materials can directly convert different types
of energy (electrical, thermal, pneumatic and optical) into mechanical energy [1.38].
Several studies have reported the potentialities of CNTs/PDMS nanocomposites that can
be widely used in many applications, such as microsensors, microrobotics, artificial
muscles, field driven cantilever actuators, microfluidic devices, etc. [1.39-1.44].
Nonetheless, applications of CNTs/PDMS nanocomposites may have great potential
in the biomedical field, with a few works presented in the literature. For instance,
Jung et al. [1.45] fabricated a CNT/PDMS composite-based dry electrode for long-term
electrocardiogram (ECG) monitoring, while Prajith et at. [1.46] developed an
electroencephalography (EEG) electrode using CNT/PDMS composite electrode in order
to investigate neurological disorders. A stationary polymerase chain reaction (PRC) device
made of CNT/PDMS nanocomposite was proposed by Quaglio et al. [1.47]. A different
work was developed by Liu and Choi [1.48], where the authors presented a nanocomposite
strain gauge made of PDMS and multi-walled carbon nanotubes (MWCNTs). These
biocompatible nanocomposites could be included in a miniature system and implanted
inside the human body (Figure 1.3) for monitoring of biomechanical strains.
Other groups reported the potential use of CNTs-based PDMS nanocomposites as
flow sensor cell, using processes of vacuum filtration, photolithography or plasma etching
[1.32, 1.49-1.51].
Introduction
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7
Figure 1.3: Implantable strain sensor with wireless sensing capability for long-term monitoring
[1.48]. Used with permission (Copyright (2009) by Institute of Electrical and Electronics
Engineers).
In addition to the aforementioned applications, CNTs/PDMS nanocomposites can
find other potential applications for structural health monitoring (SHM), as the interest of
their use in industry has increased significantly in recent years. In spite of the existence of
a great number of non-destructive evaluation tools for inspections, significant
improvements can be obtained by developing new devices to increase the security and
improve the control of the materials. Also, smart electronic textiles (e-textiles) are a new
emerging inter disciplinary field of research that could be benefitted through the
development of new devices based on CNTs/PDMS nanocomposite for health monitoring,
providing the most complete remote picture of a patient´s health condition.
Introduction
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8
1.3 Thesis Overview
1.3.1 Motivation and Objectives
While all of the applications aforementioned can benefit from the successful development
of CNTs/PDMS nanocomposites, this thesis work is motivated by potential applications of
vertically aligned-CNTs/PDMS nanocomposites in the development of flexible sensors for
pressure changes monitoring.
Silicon-based and metal-based sensors have been a driving force for the development
of MEMS since its beginning. During the ‘50s, the piezoresistance effect of silicon was
discovered and the measurement of pressure using the piezoresistance of silicon was also
demonstrated [1.52, 1.53]. After that, several works have demonstrated the interest in
developing a wide number of physical sensors, such as strain gauges, accelerometers and
pressure sensors. However, due to the mechanical characteristics of silicon, like many
other usual sensing materials, there are certain limitations to accomplish the requirements
of some applications. A good example is the biomedical field, where usually a fully
biocompatible sensing material is required for an implantable device. In addition, the
intrinsic rigidity of these conventional materials could limit its sensitivity and therefore,
might commit its response.
In order to outline these limitations, new solutions based on polymeric materials have
been used and with demonstrated potentialities for a vast range of applications. Compared
to the conventional materials that compose MEMS structures, the inherent properties of
polymers make them ideal materials to be used in several fields. Moreover, fabrication of
MEMS devices can benefit significantly of the polymer processing techniques, such as
injection molding and cast molding. A drawback in using these materials to build MEMS
devices is the reduced or absence of electrical conductivity in most polymeric materials.
Recently, a great interest on MEMS structures or devices based on carbon materials has
been rising among the scientific community. The possibility of obtaining structures with
improved mechanical properties, different shapes and geometries, has been raising the
interest in exploring new materials and to expand the development of MEMS devices to
new applications.
Introduction
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9
Incorporation of CNTs into polymeric matrices, gives an alternative of adding
conductivity into polymer systems, while allowing the fabrication of nanocomposites with
different geometries and characteristics as flexibility and biocompatibility, unattainable
from silicon or metals. Thus, CNTs/polymer nanocomposites are among the most
mentioned candidate materials for nano-electronics, where a special attention has been
given to its use to fabricate pressure sensors. Flexible CNTs/polymer nanocomposites
could represent a dominant part in the future of the electronic industry, because the
fabrication of flexible substrates that can be folded, bent and rolled might benefit greatly
the development of new portable and wearable devices.
Regarding the microfabrication techniques, ultraviolet (UV) lithography is a typical
and widely used technique to produce complex 3-D MEMS structures [1.54]. However,
unlike the microfabrication of metallic patterns, it is difficult to implement this technique
to the polymer nanocomposites [1.55]. Other reported fabrication methods for flexible
MEMS applications consist of dispersing CNTs into a polymer matrix. Nonetheless,
factors as alignment, waviness and agglomeration effects due to the physical
entanglements and van der Waals forces between individual tubes, as well as weak
bonding between CNTs and polymer matrices, have been compromising the improvement
in mechanical properties of CNTs-based polymer nanocomposites [1.56].
Based on the potential applications of CNTs/PDMS nanocomposites mentioned
before and the current challenge to effectively incorporate the distinctive and superior
properties of CNTs into a pure PDMS substrate, this research aims to introduce a new
fabrication technique to produce flexible pressure sensors based on A-CNTs/PDMS
nanocomposites. The specific objectives of the thesis are listed as follows and illustrated in
Figure 1.4:
To develop a production process to build thin (100 – 400 µm) flexible substrates with a
conductive layer, in order to construct electrodes:
by developing a fabrication technique based on vertically aligned carbon
nanotubes (A-CNTs) embedded in a flexible matrix of pure
polydimethylsiloxane (PDMS).
PDMS besides being flexible is also a biocompatible material, and the excellent
mechanical and electrical properties of CNTs allow A-CNTs/PDMS
Introduction
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10
nanocomposites to be used in fabrication of flexible pressure sensors that might
find several fields of application;
by optimizing the A-CNTs growth process and respective incorporation in the
flexible matrix of PDMS;
by evaluating the morphological, thermal, electrical and mechanical properties
of A-CNTs/PDMS nanocomposites.
To design and fabricate flexible capacitive pressure sensors:
by defining the model of the sensor (mechanical and electrostatic domains) as
well as its telemetric model;
by testing several geometries for fabrication of flexible capacitive pressure
sensors, in order to validate the concept and the developed models.
Pressure sensors are composed of three thin films, with the top and bottom
(made of A-CNTs embedded in a flexible matrix of pure PDMS) defining the
electrodes and the middle film (made of PDMS) defining the dielectric (air);
by measuring the capacity and pressure changes of the flexible sensor;
by comparing the experimental results obtained for A-CNTs/PDMS
nanocomposite flexible capacitive pressure sensors with finite element model
(FEM) simulations.
To study and explore the potentialities of the developed flexible pressure sensor using a
particular case study, specifically to use the flexible sensor in the treatment of
abdominal aortic aneurysms (AAAs):
by identifying the available treatment options and its main problems;
by defining the requirements of the flexible capacitive pressure sensor;
by developing a pressure measurement system capable of detecting post
endovascular aneurysm repair (post-EVAR) complications and able of
transmitting the measured data without the need for an embedded power supply;
by evaluating the experimental measurement results.
Introduction
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11
Figure 1.4: Schematic outline of the main objectives presented in this thesis.
1.3.2 Thesis Organization
This dissertation is divided into six chapters, which are described as follows:
Chapter 1 introduces the subject of polymer nanocomposites and analyzes the
potential applications of CNTs/PDMS nanocomposites. The motivation and objectives of
this research are also presented.
Chapter 2 reviews previous research related with both PDMS and CNT materials, as
well as their nanocomposites. It also presents sensors developed using CNTs and the
Flexible capacitive pressure sensor based on
nanocomposites
Fabrication and characterization of A-CNTs/PDMS
nanocomposite films (Chapter 3)
Pure PDMS As-grown A-CNTs embedded
in the PDMS matrix
Fabrication and characterization of flexible
capacitive pressure sensors (Chapter 4)
Abdominal Aortic Aneurysms treatment: A case
study (Chapter 5)
Introduction
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12
biocompatibility of each material is also analyzed for future potential biomedical
applications.
Chapter 3 describes the full fabrication process of A-CNTs/PDMS nanocomposites
films. The elastic response of the nanocomposites is presented as well as the electrical
conductivity measurement results.
Chapter 4 demonstrates a new approach for the design and fabrication of
A-CNTs/PDMS nanocomposite-based pressure sensors for wireless telemetry and their
characterization. The performance of the flexible pressure sensor is also assessed through
experimental measurements and comparison of the results with simulations from both
FEM and analytical modeling.
Chapter 5 presents a potential biomedical application for the developed pressure
sensor, precisely its use in the treatment of AAAs. Firstly, the main causes of this severe
disease, current available treatment options and existing implantable pressure sensors
(including those currently used in the treatment of AAAs) are reviewed. Further, a new
smart stent-graft with sensing capabilities is proposed and the development of a flexible
pressure measurement system based on inductive-coupling for monitoring of post–EVAR
procedure is also presented. The pressure measurement system’s performance is
demonstrated through measurements setups that mimic the end use of the system.
Chapter 6 summarizes results and proposes future work.
Introduction
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13
References
[1.1] D. D. L. Chung, Composite Materials: Science and Applications, 2nd ed. London:
Springer-Verlag, 2010, p. 371.
[1.2] S. S. Ray and M. Bousmina, Polymer Nanocomposites and Their Applications.
Canada: American Scientific Publishers, 2006, p. 600.
[1.3] T. Blythe and D. Bloor, Electrical Properties Polymers, 2nd ed. Cambridge
University Press, 2008, p. 496.
[1.4] M. Matos Ruiz, J. Y. Cavaillé, A. Dufresne, J. F. Gérard, and C. Graillat,
“Processing and characterization of new thermoset nanocomposites based on
cellulose whiskers,” Composite Interfaces, vol. 7, no. 2, pp. 117–131, Jan. 2000.
[1.5] W. Helbert, J. Y. Cavaill, and A. Dufresne, “Thermoplastic nanocomposites filled
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Chapter 2
Background and Prior Work
This chapter provides some general information about both polydimethylsiloxane (PDMS)
and carbon nanotubes (CNTs) materials. Relevant work in the literature regarding
experimental studies on the mechanical and electrical properties of CNTs/PDMS polymer
nanocomposites is summarized here. As mentioned in Chapter 1, this work is focused on
the development of flexible pressure sensors based on polymer nanocomposites. Therefore,
sensors based on CNTs are reviewed and envisaging the vast range of future applications,
such as the biomedical field, the issue of CNTs biocompatibility is also addressed.
Background and Prior Work
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20
2.1 Polydimethylsiloxane (PDMS)
Polydimethylsiloxane (PDMS) is a silicone elastomer belonging to the synthetic polymers
group, which has received a remarkable interest in several fields, such as the fabrication of
MEMS devices for biomedical applications [2.1], mostly because of its excellent
biocompatibility. Usually, PDMS includes a base polymer (part A) called siloxane
oligomer and a curing agent (part B), the siloxane cross-linker. These two parts are mixed
together, generally at a proportion of 10(A):1(B) weight ratio and placed in a vacuum
chamber to degas the mixture, in order to remove all of the air bubbles. Then, heating the
liquid mixture at controlled temperatures generates an elastomeric solid [2.2]. Figure 2.1
shows a representation of the mixing process.
Figure 2.1: Visual representation of the PDMS mixing process.
2.1.1 Structure and Synthesis
PDMS contains a strong polar backbone of siloxane bonds (silicon atoms and alternate
oxygen – Si-O), fortified by organic methyl groups (CH3) [2.3]. X-ray diffraction
experiments show that the crystalline structure of the PDMS holds a quadruple helical
conformation with four monomers around and a repetition distance of 12 Å in a tetragonal
unitary cell [2.4, 2.5]. The chemical formula of PDMS can be written as [-Si(CH3)2O-]n,
where n is the number of repeating monomer units, and comprises an inorganic siloxane
group linked to lateral groups of methyl as illustrated in Figure 2.2.
a)
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21
Figure 2.2: Polydimethylsiloxane. a) Chemical structure combining both organic and inorganic
groups; b) Structural representation showing the shielding of the main chain (Si-O groups)
by the methyl groups.
PDMS is produced through a cure reaction known as hydrosilylation (Figure 2.3). As
previously referred, PDMS is provided in two elements, the base oligomer or monomer
(containing vinyl-terminated dimethyl siloxane and platinum catalyst) and the curing agent
or cross-linker (hydride-terminated dimethylsiloxane). The reaction is made between the
vinyl functional group (SiCH ═ CH2) from the base oligomer and the hydride functional
group (SiH) from the cross-linker, aided by the addition of platinum-catalyst [2.6]. The
vinyl group is susceptible to bonding with a receptive hydride group on a nearest siloxane
chain, creating a cross-link.
Figure 2.3: Synthesis of PDMS through hydrosilylation (Adapted from [2.7]).
b)
Background and Prior Work
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22
2.1.2 General Properties
The exceptional flexibility of the siloxane backbone chain is attributed to the freedom of
rotation about bonds. The Si-O-Si bond angle can vary between 105º and 180º [2.8].
Comparing the energy required for rotation about the carbon-to-carbon bonds, for instance,
in polytetrafluoroethylene (19.6 kJ/mol) and polyethylene (13.8 kJ/mol), for PDMS the
energy necessary for rotation about bonds is almost zero [2.9, 2.10].
Polymer flexibility is also an indicator of its glass transition temperature (Tg).
Several factors can determine the Tg of a polymer, such as the chain length, chain stiffness,
attractive forces between molecules and polymer volume-free [2.10]. In PDMS, the high
flexibility is also responsible for its low Tg, reported as being around -125 ºC [2.2]. Also,
the low intermolecular interactions contribute for the low melting temperature (Tm),
quoted as being around -50 ºC [2.11].
Among the several attributes that characterize the PDMS, it is nontoxic to proteins
and cells, nonflammable, chemically inert, thermally stable and highly permeable to gases
(such as oxygen, carbon dioxide and water vapor) [2.12]. Likewise, PDMS is optically
transparent in a rage of wave lengths between 290-1100 nm [2.13] and has a refractive
index around 1.45 [2.14]. In addition, PDMS is thermally and electrically isolating, with a
reported thermal conductivity of 0.18 W/m.K [2.2], a rather low value compared to metals.
Silicon polymers are non-polar, in spite of its strongly polarized Si-O bonds. In fact
the methyl groups avoid a close approach of Si-O dipoles, resulting in week intermolecular
forces and essentially composed of London-van der Waals interactions. Also, due to these
low intermolecular forces, several properties of PDMS are little affected over a large range
of temperatures, which do not happens in most organic polymers [2.15].
Pure PDMS is highly hydrophobic which interact most strongly with polar
compounds, through hydrogen bonding between the siloxane group of PDMS and
alcoholic/acid hydrogen of the analyte or through polar-to-polar interactions. In addition,
PDMS can also present slightly strong hydrophobic interaction with compounds bearing
methyl or alkyl groups due to van der Waals forces [2.16]. Because of this character of the
PDMS chain, the dielectric constant of this type of polymers is relatively high when
compared to other non-polar polymers.
Background and Prior Work
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23
Table 2.1 compares the dielectric properties of PDMS and high density polyethylene
(HDPE), a non-polar polymer. The volume resistivity of PDMS and dielectric strength
properties are influenced by the presence of material impurities [2.16, 2.17]. This could be
reinforcing fillers of the matrix or other impurities related with residual marks of
polymerization catalysts.
Table 2.1: Dielectric properties of polydimethylsiloxane (PDMS) and high density
polyethylene (HDPE) [2.17].
Polymer
Dielectric
constant at
100 Hz
Dissipation
factor at
100 Hz
Dielectric strength
at 60 Hz
Volume resistivity
(ohm.cm)
Tg
(K)
PDMS 2.86 0.00025 552 5.3x1016
150
HDPE 2.30 0.00011 811 2.2x1016
148
2.1.3 Mechanical properties
PDMS elastomers reveal very poor mechanical properties and especially, they present low
tensile strengths [2.18-2.20]. PDMS has a Young’s modulus (E) typically around 1 MPa
[2.21]. The shear modulus (G) could vary between 100 kPa and 3 MPa and it presents a
very low dielectric loss tangent (tan δ ˂ 0.001) [2.22]. However, there are some parameters
that can be changed in the preparation process of PDMS in order to improve the final
mechanical properties.
The manufacturer suggested a mixing ratio for the base polymer and curing agent,
respectively, of 10:1. However, it is expected that some considerable changes happen in
the final properties of cured PDMS, by changing the recommended proportions. Usually,
as higher is the amount of curing agent used in the formulation, higher is the number of
cross-links that are formed. Nevertheless, there is a critical point where all chains are
cross-linked and therefore are unable to further bond [2.23].
Several authors have studied the influence of different mixing ratios of the base
polymer/curing agent on the mechanical properties of PDMS [2.24-2.27]. The base
Background and Prior Work
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24
polymer has a viscosity of 0.0050 m2/s while the curing agent has a viscosity of
0.0011 m2/s.
Mata et al. [2.24] presented an extensive study regarding the effects of several
production techniques on the final properties of PDMS. Using the same spin-coating
conditions, they found that an increase in the quantity of curing agent leads to a liquid
mixture less viscous and consequently final films less thick (Figure 2.4). In addition, the
ultimate tensile strength was observed to increase by increasing the amount of curing agent
up to the combination of 14.3 wt. % cross-linker. However, mixtures with significant
increase in the amount of curing agent (combinations of 21.4 and 42.9 wt. % cross-linker)
presented an opposite effect, decreasing the ultimate tensile strength.
Figure 2.4: Speed-thickness correlation for different formulations of base polymer/curing agent
combined at different weight ratios and spin coating speeds. Used with permission from reference
[2.24] (Mata A et al. Biomed Microdevices 2005; 7: 281-293, Copyright (2005) Springer).
Khanafer et al. [2.25] studied the mechanical properties of PDMS by combining
different mixing ratios. They found that the elastic modulus increases by increasing the
mixing ratio. As shown in Figure 2.5, the high elastic modulus value was obtained for the
formulation 9:1 base polymer/curing agent. After that, continuing to increase the mixing
ratio the elastic modulus begins to decrease.
Background and Prior Work
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25
Figure 2.5: Effect of different formulations of the PDMS on the Young’s modulus, for a crosshead
speed of 5 mm/min. Used with permission and adapted from reference [2.25] (Khanafer K et al.
Biomed Microdevices 2008; 11: 503-508, Copyright (2008) Springer).
PDMS cures chemically by forming Si-CH2-CH2-Si linkages [2.28] and, by
increasing the temperature these linkages are formed more quickly. Therefore, using high
temperatures to cure the PDMS considerably decreases de cure time.
Liu et al. [2.29] investigated the influence of heating temperature on mechanical
properties of PDMS. They found that the Young´s modulus and ultimate tensile strength
are independent of heating time, for temperatures lower than 200 ºC. One the other hand,
higher heating temperatures generate lower mechanical strength. Also, Liu et al. [2.30]
demonstrate that both the mechanical strength and the Young’s modulus of the PDMS are
thickness dependent. Their results showed that the Young’s modulus increases
significantly when the thickness decreases. Likewise, the ultimate tensile strength increases
largely as the thickness of the PDMS samples is reduced.
Lötters et al. [2.22] measured the shear modulus by changing the applied frequencies
and temperatures in the range of 0.005-30 Hz and 0-70 ºC, respectively. The authors used a
thick cylindrical PDMS structure of height 1.2 mm and radius 8 mm, placed between two
parallel circular discs. They obtained a shear modulus value of 250 kPa and concluded that
it is independent of the applied frequency but linearly dependent on the temperature with a
slope of 1.1 kPa/ºC.
Background and Prior Work
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26
2.2 Carbon nanotubes fundamentals
Since the first observation by Iijima in 1991 [2.31], carbon nanotubes (CNTs) have been
focus of considerable attention for the researchers due to its outstanding physical and
mechanical properties. Nanotubes are long and slim fullerenes, with a graphene structure
where the walls of the tubes are hexagonal and often capped at each end (Figure 2.6).
CNTs can have diameters around 1nm (10 Ǻ) and lengths up to centimeters. Their
symmetric structure is responsible for the exceptional mechanical properties that the CNTs
show [2.32].
Figure 2.6: Atomic structure of a single graphene sheet [2.33].
2.2.1 Carbon nanotubes morphology
CNTs can be synthesized in two structural forms: single-walled and multi-walled
structures, also known as single-walled carbon nanotubes (SWCNTs) and multi-walled
carbon nanotubes (MWCNTs), respectively (Figure 2.7). MWCNTs were first observed by
Iijima [2.31], that consisted of two or more seamless graphene cylinders concentrically
organized. Through electron difraction analysis, Iijima has observed tubules with a helical
arrangement relatively to the tube axis. The diameter of the inner tubules was observed to
be smaller than 2.2 nm, while the outer diameter ranges from 4 to 30 nm and lengths of up
to 1 µm. Two years later, Iijima et al. and Bethune et al. [2.34, 2.35] produced
independently SWCNTs, made just with one layer of carbon atoms. Their diameters were
Background and Prior Work
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27
reported as ranging from 0.4 to 2-3 nm and their length in the micrometer order,
hexagonally arranged to form a crystal-like structure [2.36].
Figure 2.7: Schematic of single-walled carbon nanotubes (SWCNTs) and multi-walled carbon
nanotubes (MWCNTs) showing the common dimensions of length, width and separation
distance between graphene layers in MWCNTs (Adapted from [2.37]).
MWCNTs are basically composed of concentric SWCNTs like matryoska dolls
[2.38], with a wall-to-wall distance of 0.36 nm [2.31]. Despite structural similarities to a
single graphitic sheet, SWCNTs are described using three different morphologies.
Depending on the appearance of the carbon bonds around the nanotube cylinder, the
nanotube is described as having armchair, zigzag or chiral morphology [2.30, 2.36], as
ilustrated in Figure 2.8. This atomic structure is described by its chirality, which refers to
the rolling angle (θ) of the graphene sheet that creates the carbon nanotube and the chiral
vector (Ch), that can be defined in terms of the lattice translations indices (n, m) and the
unit vectors (a1 and a2) of the hexagonal lattice [2.39].
Variations in the atomic structure can change largely the properties of nanotubes and
most significantly their electronic properties [2.38, 2.40]. Lau and Hui [2.39] provided a
wide review on the three different structures of nanotubes – armchair (n, n), zigzag (n, 0)
and chiral (n,m with n ≠ m). Generally, SWCNTs can present a metallic structure for an
armchair morphology (n, n), they can be very small-semiconductors – (n, m) morphology
Background and Prior Work
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28
with n-m = 3j and j is a non-zero integer, and large-semiconductors for all the other
morphologies.
Figure 2.8: Atomic structure illustrations showing the chiral vector (Ch) in a hexagonal graphene
sheet, defined by unit vectors a1 and a2, and the chiral angle θ (Adapted from [2.32]).
2.2.2 Physical properties of carbon nanotubes
In order to better understand the physical properties of CNTs, several researchers have
attempted to characterize CNTs both experimentally and theoretically. CNTs raise huge
interest due to the unequaled atomic structure, exceptional electronic properties (thermal
and electric conductivity) and mechanical properties (high Young’s modulus, tensile and
compressive strength), equal to or higher than metals [2.40-2.43].
As referred before, the conducting properties of CNTs range from metallic to
moderate band gap semi-conductors [2.44]. They are very sensitive to their geometric
structure [2.41, 2.43], depending greatly on the diameter and chiral vector. However, so
far, there is no efficient method to control the chirality during the production of the carbon
nanotubes [2.45]. Perfect metallic CNTs should be ballistic conductor, which means that
there is zero resistance along the tubes [2.46], and it is predicted to have a quantum
conductance of 6.45 kΩ. Individual CNTs resistances can vary from a few kΩ to a few
MΩ, at room temperature [2.47-2.50].
Background and Prior Work
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29
Based on Ashby charts [2.51] (Figure 2.9), it is possible to observe that individual
CNTs are thermally and electrically more conductive in comparison with materials such as
metals, carbon fibers, polymers and polymer-based composites.
Figure 2.9: Electrical and thermal conductivities of individual CNTs normalized by density, in
comparison with different materials (Adapted from [2.52]).
The electrical resistance was first measured for bundles of MWCNTs and oriented
films. Song et al. [2.53] obtained a resistivity of 0.0065 Ω.cm at 300 K and 0.016 Ω.cm at
5 K. Deheer et al. [2.54] developed a method to fabricate thin films of aligned CNTs. They
verified that the electrical resistivity is anisotropic, being smaller along the tubes than in
the perpendicular direction to them. A similar work performed by Baumgartner et al. [2.55]
suggested that the electric resistance is temperature dependent. They found a bulk
resistivity of 0.02 Ω.cm at 300 K, along the tubes, a smaller value when compared to the
perpendicular direction. Bauhofer and Kovacs [2.56] reviewed experimental and
theoretical work on electrical percolation of CNTs in polymer composites.
In the last years, there have been several activities regarding the measurement of the
thermal conductivity of the CNTs. Its direct measurements still remain a challenge,
however, values of about 3000 W/m.K for MWCNTs and up to 2000 W/m.K for
SWCNTs, have been reported. Based on equilibrium and non-equilibrium molecular
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dynamics (MD) simulations, Berber et al. [2.57] determined the temperature dependence of
individual SWCNTs. Their results indicated a thermal conductivity of 6600 w/m.K at room
temperature, and an unusual value of 37000 W/m.K at 100 K. Another MD study
performed by Osman and Srivastava [2.58] indicated a thermal conductivity in the range of
1500-3000 W/mK for SWCNTs, at room temperature. The complete phonon dispersed
relation method was used by Gu and Chen [2.59] in order to determine the thermal
conductivity of individual SWCNTs. Their calculations suggested a value of 474 W/m.K
and a linear temperature dependence of the thermal conductivity. Maruyama [2.60]
simulated the heat conduction in finite-length SWCNTs, also using the MD method. The
thermal conductivity was calculated for two SWCNTs structures with different diameters.
The author found that the thermal conductivity of (10, 10) was almost independent of the
tube length, while the thermal conductivity of (5, 5) structures diverges as a power law. In
parallel to the theoretical studies, there is also some experimental works on this subject.
Yang et al. [2.61] studied the thermal conductivity of MWCNTs using a pulsed
photothermal reflectance technique and proposed that the effective thermal conductivity
could be about 200 W/m.K. Kim et al. [2.62] used a suspended device to measure the
thermal conductivity of single MWCNTs. The carbon nanotubes were placed on the
fabricated device and connected the other two suspended islands. The authors found a
thermal conductivity higher than 3000W/m.K, at room temperature. Yu et al. [2.63] used a
similar device to measure the thermal conductance of individual SWCNTs in the
temperature range of 110 and 300 K. They obtained a thermal conductivity higher than
2000 W/m.K, decreasing with the decrease of temperature. Hone et al. [2.64] investigated
the temperature-dependent thermal conductivity of crystalline ropes of SWCNTs in a range
of 350 to 8 K. They reported that below 30 K the thermal conductivity is linear
temperature-dependent and dominated by phonons at all temperatures. The temperature
dependence of thermal conductivity of individual MWCNTs was also analyzed by Small et
al. [2.65]. Their measurements indicated a thermal conductivity higher than 3000 W/m.K,
an order of magnitude higher than the 200 W/m.K found by Hone et al. [2.66] for
magnetically aligned SWCNTs.
The mechanical properties of CNTs have been subject of a number of theoretical and
experimental studies. Several works has been focused on Young´s modulus and strength
along the CNTs axis and values above 1 TPa [2.65-2.69] and 10 GPa [2.70], respectively,
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have been reported. However, maybe due to the techniques used for CNTs
characterization, the results obtained differ considerably.
The first study regarding the mechanical response of CNTs to large deformations
was performed by Yakobson et al. [2.71]. The authors used molecular dynamics
simulations to model the armchair, chiral and zig-zag structure of the CNTs, and predict its
behaviour beyond the linear response. By choosing the appropriate parameters, the authors
found that the effect of chirality causes almost no effect on the elastic stiffness of the
CNTs. Also, they found that CNTs are very resilient, supporting extreme strain without
any sign of brittleness or plasticity. Based on the empirical lattice dynamical model [2.72]
used to investigate the elastic properties of graphite layers, Lu [2.73] used a similar model
to predict the elastic properties of single- and multilayered nanotubes with different size
and geometry. According to their results, the Young’s modulus (~ 1 TPa), shear modulus
(~ 0.45 TPa) and bulk modulus (~ 0.74 TPa) are the same for all nanotubes with a radius
larger than 1 nm. Using molecular dynamics simulation for an armchair SWCNTs, Yao et
al. [2.74] obtained a Young’s modulus of 3.6 TPa and tensile strength of 9.6 GPa.
Experimental methods have been also performed in order to estimate the elastic
properties of CNTs. Treacy et al. [2.67] estimated the Young’s modulus of isolated CNTs
by measuring their intrinsic thermal vibrations and using a transmission electron
microscope (TEM). They obtained an average value of the Young’s modulus of 1.8 TPa.
Similar works were also performed using this method [2.75, 2.76]. However, the calculated
values were slightly lower than those obtained by [2.67]. Wong et al. used an atomic force
microscope (AFM) to record directly the stiffness and strength of a single MWCNTs
through a bending test. They obtained an elastic modulus value of 1.26 TPa while the value
of the tensile strength was reported as 28.5 GPa. Yu et al. [2.70] performed tensile-loading
experiments to measure tensile strengths of individual MWCNTs. The Young’s modulus
obtained varies from 270 to 950 GPa and the tensile strength was reported to vary from 11
to 63 GPa. Qi et al. [2.76] proposed a new method to determine the mechanical properties
of vertically aligned carbon nanotubes (V-ACNTs) forest and constituent nanotubes
through nanoindentation technique. During this process, the CNTs are successively bent by
the penetration of the indentor. The reported results are values of the effective bending
modulus ranging from 0.91 to 1.24 TPa and effective axial modulus values between 0.9
and 1.23 TPa. Also, the CNTs wall modulus was found to vary between 4.14 and 5.61 TPa.
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In the literature it is possible to find vast information on the mechanical properties of
CNTs obtained from both experimental and theoretical studies. However, the values are
very spread in the reported works. Table 2.2 summarizes the values of elastic modulus and
tensile strength for both SWCNTs and MWCNTs, also similarly presented by Thostenson
et al. [2.77] on their research work.
Table 2.2: Elastic modulus and tensile strength values of both SWCNTs and MWCNTs, more
frequently referenced in the literature [2.77].
Experimental measurements
Method Elastic
modulus (TPa)
Tensile
strength (GPa) Type of CNTs
TEM - Thermal vibration of beam [2.43] 1.8 - MWCNTs
TEM - Direct tension [2.75] 0.91 150 MWCNTs
TEM - Electrostatic deflection [2.78] 0.1 – 1 - MWCNTs
AFM - 1 end clamped, bending [2.41] 1.28 28.5 MWCNTs
AFM - 2 ends clamped, tensile
[2.79, 2.80] 0.81 - SWCNTs
AFM - lateral force mode [2.80] - 45 SWCNTs
Dual AFM cantilevers [2.70] 0.27 – 0.95 11 – 63 MWCNTs
[2.81] 0.32 – 1.47 13 – 52 SWCNTs
Nanoindentation [2.76] 0.9 – 1.23 - MWCNTs
Theoretical calculations
Empirical lattice mechanics [2.68] 0.97 - MWCNTs
Ab initio [2.82] 1.0 - SWCNTs
Molecular structural mechanics [2.83] 1.05 - SWCNTs
Pin-jointed truss model [2.84] 0.68 - SWCNTs
Molecular mechanics simulation [2.85] - 93 – 112 SWCNTs
Molecular dynamics simulation [2.86] - 150 SWCNTs
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2.2.3 Carbon nanotubes growth techniques
CNTs are produced using a wide diversity of techniques, but there are three main methods
that have been developed to synthesize nanotubes, such as arc discharge
[2.34, 2.35, 2.87-2.89], laser ablation [2.36, 2.90-2.92] and chemical vapor deposition
(CVD) [2.93-2.96].
Arc discharge method was the first accessible technique to produce both SWCNTs
and MWCNTs, the same method used to produce fullerene molecules [2.97]. This
production technique is based on an electric arc discharge created between two graphite
electrodes, at high temperature and under an inert atmosphere of helium or argon [2.98].
Both electrodes are maintained at a distance sufficiently small (< 1 mm) and when voltage
is applied, a discharge occurs. Thus, after carbon sublimation the positive electrode is
consumed. This mechanism is maintained by continuously moving the anode towards the
cathode, in order to preserve a constant distance between them. The synthesis of SWCNTs
or MWCNTs depends on the material that is introduced into the hollow anode. By filling
the anode with metal powder, typically iron (Fe), nickel (Ni), carbon (C), cobalt (Co) and
yttrium (Y) at appropriate mix, SWCNTs will be synthesized. Otherwise, without metal
powder inside the hollow anode, MWCNTs are synthesized when the arc discharge is
generated. Figure 2.10 illustrates an arc discharge apparatus to obtain different carbon
substances using this technique, that can be classified into SWCNTs, MWCNTs,
amorphous carbon, graphite, carbon nanoparticles, etc. [2.99].
Figure 2.10: Schematic representation of an arc discharge apparatus for the synthesis of different
carbon structures (Adapted from [2.100]).
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Laser ablation technique is based on the same mechanism as arc discharge, but now
the energy is supplied by a laser hitting a solid disk of graphite holding catalyst materials,
typically Ni or Co. Historically, laser ablation was the first method used to create fullerene
clusters in the gas phase [2.101]. A solid disk of graphite is placed inside a quartz tube,
mounted in a furnace at controlled temperature (Figure 2.11). Using a focused pulsed or
continuous laser, the carbon is vaporized from the surface of a graphite target into a high
density helium or argon flow. After vaporization, the carbon species produced are scanned
by the flowing gas from the region at high temperature and deposited on a conical
water-cooled copper collector. Similarly to the arc discharge method, SWCNTs and
MWCNTs can be obtained by the laser ablation technique. MWCNTs are synthesized
when a pure graphite target is used in the process [2.102]. On the other hand, when a small
quantity of transition metal is added to the carbon target, SWCNTs are synthesized [2.103].
Figure 2.11: Schematic representation of a laser ablation reactor (Adapted from [2.104]).
At the moment, arc discharge and laser ablation produces a small amount of CNTs.
Also, both techniques present significant limitations, such as the high costs associated to
the large production of CNTs, the high temperatures involved in the process and the
limited source of carbon, restricting the amount of CNTs that can be produced.
CVD technique is the most common used method to produce CNTs in very large
quantities at relatively low costs and low temperatures, comparatively to the arc discharge
and laser ablation techniques. The CVD method consists in a system that involves the
heating of a substrate inside an oven and the addition of a carbon source gas, in which a
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hydrocarbon vapor is thermally decomposed in the presence of a metal catalyst at high
temperatures (from 500 ºC to 900 ºC). In this method, the catalyst precursors may be used
in any state (solid, liquid or gas) and permits the use of different substrates, allowing the
growth of CNTs in a diversity of forms, such as powder, films, aligned or entangled,
straight or coiled, or a whished geometry previously defined on a patterned substrate.
CNTs synthesis by CVD method is well reviewed by Kumar and Ando [2.105]. CNTs
grow over the catalyst and are collected upon cooling the system to room temperature. The
catalyst substrate is placed in the hot zone of the furnace to catalyze CNTs growth and the
decomposed carbon species of the hydrocarbon are assumed to dissolve in the metal
nanoparticles. Then, after carbon supersaturation in the metal it precipitates out in the form
of a carbon cylinder. Figure 2.12 shows a schematic diagram of CVD process, using a
furnace with resistive heater as energy source to decompose the catalyst precursor and to
heat the sample.
Figure 2.12: Schematic diagram of the CVD apparatus (Adapted from [2.106]).
The growth mechanism of CNTs is not exactly known and it is still a subject of
controversy. However, there are many experimental and theoretical studies as attempts of
explaining the way in which CNTs are formed [2.107-2.111], and more recently, in situ
high-resolution transmission electron microscopy (HRTEM) was used for observation of
the growth process of CNTs [2.112-2.116]. Sinnott et al. [2.117] proposed a model to
account for the catalyzed growth of CNTs by CVD process and that consists in two main
growth models, as schematically represented in Figure 2.13.
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Figure 2.13: Tip-growth model and base-growth model of CNT growth mechanism [2.105].
Depending on the contact angle between the catalyst and substrate, and therefore the
interaction forces between them, the growth mechanism can be tip-growth model or
base-growth model. In the first case, where the catalyst-substrate interaction is weak, the
hydrocarbon decomposes on the top of the metal and the CNT lifts the catalyst from the
substrate during the growth (step i). If the top of catalyst is open for new hydrocarbon
decomposition, CNT nucleates and still grows below the catalyst (step ii). The growth
process will stopped when the catalyst is entirely covered with excess carbon (step iii). In
the second model, where the catalyst-substrate interaction is strong, the hydrocarbon
decomposition occurs in a similar way to the tip-growth method. However, the nanotube
grows above the catalyst, which remains attached to the substrate [2.105].
The production of SWCNTs or MWCNTs is controlled by the size of catalyst
particles [2.118]. Also, the growth temperature plays an important role for CNTs synthesis.
Commonly, the use of low temperatures (600 – 900 ºC) during the CVD process promotes
the growth of MWCNTs while higher temperatures (900 – 1200 ºC) favor the growth of
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SWCNTs. The hydrocarbons used in the process also influence the type of CNTs that will
be synthesized. MWCNTs can be grown from most hydrocarbons as acetylene (C2H2),
ethylene (C2H4), benzene (C6H6), etc., while SWCNTs just grow from some types of
hydrocarbons, such as carbon monoxide (CO), methane (CH4), etc. [2.119-2.123].
Concerning the catalysts for CNTs growth, the most frequently used are the transition
metals (such as Fe, Co, Ni) mainly due to the high solubility of carbon in these transition
metals at high temperatures [2.106].
2.3 Aligned CNTs/PDMS nanocomposites
Polymers are versatile materials that can be processed more quickly than ceramics and
metals. However, in spite of certain advantages associated to their unique properties,
polymers are not so strong or stiff, which make them inadequate for many engineering
applications [2.124-2.127]. For this reason, there is a permanent investigation towards new
polymer-based materials with enhanced properties (mechanical, electrical, etc.), and
reinforcement of polymer materials through the use of different kinds of organic and
inorganic fillers has been used around the world.
The fabrication of PDMS-based devices is relatively simple with low costs involved
and since PDMS preserves its flexibility and elasticity under a large range of temperatures,
it has been proposed for flexible device applications. However, given that most of the
flexible applications (e.g., implantable medical devices, flexible displays, tactile sensors,
etc.) require electrical conduction in addition to flexibility provided by PDMS,
reinforcement of the PDMS [2.128] or the deposition of thin films [2.129] in order to
provide electrical properties and/or enhance the mechanical performance are desirable. The
most commonly used fillers in PDMS are silica, carbon black and metal particles, but
non-spherical fillers, such as fibers, have also been used. The reinforcing qualities of these
fillers have been widely demonstrated in the literature. However, considerably high filler
loading was necessary to achieve high mechanical properties. In addition, the filler
particles tend to agglomerate and as result, the final material is quite inhomogeneous. In
fact, dispersion is possibly the most fundamental issue for effective reinforcement and to
achieve efficient load transfer to the CNTs network.
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CNTs are an excellent candidate as reinforcement material in polymeric materials for
multi-functional property enhancement (mechanical, electrical, etc.) due to their
exceptional physical properties and low density, particularly if the properties can be
tailored anisotropically via aligning the CNTs ‘fibers’. It comes without surprise that
CNTs-based nanocomposites have been receiving a lot of attention [2.130].
Numerous studies available in the literature have focused in the characterization of
nanocomposites obtained by dispersing CNTs in a polymeric matrix [2.131-2.133]. Many
of those studies used CNTs to reinforce polymer films [2.134, 2.135] in order to improve
not only the mechanical properties of the polymer, but also to increase many other
properties compared to the neat polymer and to carbon-fiber reinforced nanocomposites
[2.136], such as electrical properties of interest in this thesis. Review articles have
highlighted percolation theory in the area of polymer nanocomposites [2.137, 2.138] and a
review article solely on this topic has appeared [2.56]. Furthermore, several techniques
have been developed to embed CNTs into a polymer film, such as calendering and
sonication to disperse the CNTs into the neat polymer [2.139, 2.140]. Different
characterizations techniques used to evaluate the mechanical performance of
nanocomposites films revealed that a good dispersion, alignment of CNTs and a good
adhesion between the polymer matrix and the CNTs are the three main issues that
influence the mechanical performance of the nanocomposites [2.141-2.143].
Aligned carbon nanotubes (A-CNTs) were first reported by Thess et al. [2.36] in
1996 and in the same year, vertically A-CNTs were successfully grown by CVD technique.
El-Aguizy et al. [2.144] demonstrated a process developed to produce nanopellets,
requiring the length control of individual CNTs, its alignment and position within a
microscale block of polymer. With the recent optimization in vertically A-CNTs growth,
there are opportunities for new materials and applications since CNTs/Polymer-based
nanocomposites properties are enhanced by this morphology that organizes the CNTs
along a common direction [2.145-2.147]. Also, a controlled growth of vertically A-CNTs
allows developing a wide range of smart materials and advanced devices. Review articles
on this topic was presented by Chen et al. [2.148] and Liu et al. [2.149].
Experimental results on polymer nanocomposites (PNCs) using CNTs reported in
literature have shown that the mechanical properties of PNCs have varied from almost no
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changes [2.150] to moderate increases on tensile strength [2.151-2.154], decreases in
toughness [2.155] and electrical conductivities of about 2 S.m-1
at 1 wt % of CNTs in
epoxy matrices (with a percolation threshold of 0.0025 wt %) [2.156]. A common reported
problem is the difficulty to fully disperse and align the CNTs in the polymers, reason why
it is difficult to achieve optimal property improvements [2.78, 2.157, 2.158].
Most of the studies reported using CNTs/PDMS nanocomposites, use CNTs or
A-CNTs mixed with pristine PDMS through magnetic stirring. Wu et al. [2.158] reported
an elastic modulus of CNTs/PDMS nanocomposites ranging from 1.71 MPa to 2.34 MPa,
for different weight percentage of CNTs. Specifically, for 1 wt % of CNTs they found an
elastic modulus of 1.71 MPa while the elastic modulus of pure PDMS was 1.65 MPa.
Dyke et al. [2.159] found that the tensile modulus and strength are significantly increased
on their PDMS composites with functionalized SWCNTs, while the strain at break remains
unchangeable when compared with the pure PDMS.
Zhang et al. [2.160] developed a transparent stretchable conductor embedding
aligned MWCNTs ribbons in a PDMS matrix, in order to investigate the nanocomposite
behaviour by measuring the dependence of resistance on the tensile strain. They obtained a
CNTs/PDMS film that showed exceptional resistance characteristics in stretching/releasing
cycles. Lu et al. [2.161] prepared nanocomposites of MWCNTs and PDMS at different
filler concentrations in order to study the electromechanical and piezoresistance properties
of the nanocomposite. The authors found that the gauge factor (GF) varies from 1.38 to
12.4, and it is dependent on the content of CNTs. Also, for the content of 18 wt % CNTs in
the PDMS matrix, the Young’s modulus was found to be about 4.49 MPa at the initial
stage. Ahir et al. [2.162] used MWCNTs embedded in a PDMS matrix to study the
reversible photo-mechanical actuation of the composites under infrared (IR) light
irradiation. The results showed that the PDMS have no response to near IR radiation, but
the presence of CNTs causes a strong reversible response. Also, the resulting composited
showed the capacity to change their actuation direction from expansive to contractive
response which can be controlled by the degree of alignment of the CNTs. A similar work
was performed by Lu and Panchapakesan [2.163] to study the photomechanical response
of different combinations of CNTs/PDMS nanocomposites. All tested configurations
revealed structural expansion in response to the light under small pre-strains. Otherwise,
under large pre-strains, the samples showed structural contraction in response to the light.
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Liu and Choi [2.164] presented a novel approach to embed conductive PDMS
composite patterns into a bulk elastomer, using microcontact printing and casting mold
techniques. The PDMS samples with embedded conductive and flexible patterns
demonstrated outstanding mechanical flexibility. In addition, the authors developed a strain
gauge to prove the feasibility of their approach. The results showed significant and
consistent resistance change under different tensile strains and over multiple cycles of
measurements.
More recently, Misra et al. [2.165] developed a light weight material composed by
alternated layers of PDMS and A-CNTs arrays. The characterization results demonstrate
that multilayer structures were capable to sustain large compressive deformations
(up to ~ 0.8 strain) with almost complete recovery.
Nevertheless, research on CNTs composites has concentrated largely in the
characterization of CNTs/Polymers composites, where the SWCNTs/MWCNTs are
dispersed or mixed within the polymeric matrix. The work presented in this thesis uses
controlled vertically A-CNTs embedded in a flexible substrate of PDMS and the resultant
A-CNTs/PDMS nanocomposite films are used to fabricate flexible pressure sensors.
2.4 Sensors and actuators based on carbon nanotubes
There have been efforts to develop smart materials aiming a wide range of different
applications. These materials hold properties that can be changed, in a reversible mode, by
an external condition, such as light, temperature, pressure or electricity. The structural and
electrical properties of CNTs make them attractive materials to develop a new class of
smart materials for actuation and sensing applications.
The first actuator made of CNTs was a macro-scale sheet of nanotubes denominated
as buckypaper [2.166-2.168], which is composed of highly entangled SWCNTs bundles
with a high sensitivity to strain in the linear bending range. Nonetheless, the slippage
between the SWCNTs bundles affects efficient shear transfer for smart structures
applications. In 2005, Yun et al. [2.169] developed an actuator with improved properties
for structural applications, by constructing and testing a SWCNTs-epoxy ply material that
has electrochemical actuation properties. The results showed that the epoxy-based actuator
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has a higher elastic modulus and strength, compared to the buckypaper tape actuators.
Riemenschneider [2.170] performed experimental laboratory work in order to better
understand the general characteristics of CNTs-based actuators. SWCNTs were used to
produce a buckypaper and its electrochemical behaviour was assessed by using cyclic
voltammetry. For the electromechanical study, both liquid and solid electrolytes were used
for the tests.
First studies for strain sensing using CNTs were performed by Raman spectroscopy.
However, this technique did not demonstrate to be practical for engineering applications
due to its complexity [2.171-2.173]. Kang et al. [2.174] developed a SWCNTs/PMMA
(polymethylmethacrylate) composite sensor in order to improve the strain transfer
capabilities. The results showed an improvement of the adhesion between nanotubes,
mainly because of the stronger polymer interfacial bonding, but a lower sensitivity than the
buckypaper sensor. In addition, the sensor demonstrated a good response under static and
dynamic strain. Chang et al. [2.175] demonstrated a fabrication technique to produce high
sensitivity flexible strain sensors at room temperature. In the process, suspended SWCNTs
are transferred from a silicon wafer to a flexible substrate that already has pre-etched
trench and electrodes. The mechanical bending tests showed that the strain sensor reaches a
high strain resolution of 0.004 %. The piezoresistive gauge factor was found to be 269,
higher than the typical metal strain gauges (1-5).
The nanoelectromechanical systems (NEMS) field has raised a significant interest
because of its potential to develop high sensitivity and low power sensing devices.
Hierold et al. [2.176] reviewed different transducer concepts and demonstrators of NEMS
sensors based on CNTs. Also, the authors developed a concept for a nanoscale force
sensor, using suspended SWCNTs-based cantilever structure and a membrane-based
structure. An atomic force microscope (AFM) was used to deflect the cantilever-based
structure. Electrical measurements of the resistance of the SWCNTs under mechanical load
demonstrated a considerable and reversible increase of the nanotubes’ resistance.
Polymer-based composites are also an attractive option in developing large and flexible
sensors. Knite et al. [2.177] presented the design, elaboration and study of polyisoprene
and MWCNTs composites for application in strain sensors. The polyisoprene composites
containing dispersed MWCNTs were prepared using a solution method to produce the film
and samples were used to study the deformation and dependence of electrical resistance.
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Results demonstrated that the polyisoprene matrix with embedded MWCNTs provided
about 4 orders of magnitude of nonreversible change in the electric resistivity at large
stretch (40 %) and it is more suitable for small tensile strain sensing. Vemuru et al. [2.178]
used a MWCNTs film to study its real-time strain response at macroscale under tensile
load and real-time voltage change as a function of temperature. The obtained results
showed that the MWCNTs film used as a strain sensor at the macroscale was very
promising. There was a linear relationship between the change of voltage across the
MWCNTs film and the applied strain. A good response of the film during both loading and
unloading tests was also observed. The temperature effects on the samples, even though
having limited responses, seemed to have an excellent predictable response to changes in
temperature.
Prototypes sensors based on SWCNTs for measuring parameters related to vehicle
structural health were presented by Watkins et al. [2.179]. The proposed sensors were
fabricated using established lithographic techniques to pattern electrodes on silicon and
flexible plastic substrates and then, SWCNTs were placed and aligned between the
electrodes, generating electrical bridges. Preliminary results using an AFM showed that the
conductance through a SWCNT can change with the applied strain. Tests on these sensors
were still under way when this work was being prepared.
Loh et al. [2.180] developed a new thin film strain sensor for structural monitoring,
as an attempt to overcome some limitations revealed by the existing foil-based strain
gauges. The fabrication technique to produce the strain sensor uses SWCNTs and diverse
polyelectrolytes, combined in a layer-by-layer assembly method. To test the strain sensor,
the monotonic tensile test and the four-point bending test were used. The first method
proved that the composite thin films had good sensing capabilities and therefore, a great
promise for strain sensing. The results of the four-point bending test suggested that the
composite thin film correlates well with the change in resistance of the reference gauge.
A more recent work performed by Alexopoulos et al. [2.181] presents a method to
fabricate composite materials for structural health monitoring purposes. CNT fibers were
embedded into samples of glass fiber reinforced polymers, in order to obtain
simultaneously the response of the composite to mechanical loading and its sensing
capability through electrical resistance change in the CNT fiber. To assess the composite
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materials, tensile tests with various incremental loading-unloading loops and three-point
bending tests were performed. Tensile tests demonstrated that the CNT fiber works for
sensing tension while the three-point bending tests showed that the CNT fiber works for
tension and compression loadings. Kessler et al. [2.182] built advanced
fiber-reinforced polymer-matrix laminates with A-CNTs grown in-situ coupled with a
non-invasive sensing technique. The fabricated CNTs-enhanced composites were patterned
with a silver ink electrode grid and subjected to impact damage. The first results on
electrical resistance measurements for both in-plane and through thickness demonstrated
the potential of using their proposed approach.
Currently available literature exhibited a vast amount of information related to
sensors and CNTs. However, this section intended to demonstrate some developments
focused on smart materials based on CNTs with potential to improve the existent smart
materials in different areas, but mainly sensors with multi-functional electronic
applications. It is expected that the continuous development and improvement of
CNTs-based materials will generate innovative smart composite material that can be
widely used in different areas of application.
2.5 Biocompatibility of PDMS and CNTs materials: perspectives of their
application in medical implants
Biocompatibility plays an important role in the development of medical devices for
diagnostic or treatment applications. In general, medical devices consist of a diversity of
materials, whose effect in the human body may vary depending on its application and use
frequency (daily life, non-invasive and invasive). Biocompatibility of medical devices
depends not only of the application in the human body (for external or internal use), but is
also influenced by its fabrication process [2.183].
There are sets of international standards, namely the ISO 10993, which is composed
of a series of standards for assessing the biocompatibility of a medical device. Depending
of the medical device purpose, a large number of tests should be done, such as tests on
toxicity, carcinogenicity (the ability or tendency to produce cancer), hemocompatibility
(the compatibility of a material with the blood), etc., where some of these tests are just
in vitro tests but others need animal experiments. In general, the standard distinguishes the
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contact with a human body as surface devices (contact with the skin), external
communicating devices (indirect contact) and implantable devices (direct contact with the
tissue, bone ore blood) [2.184].
All implantable materials have intrinsic morphological, chemical, and electrical
surface properties which can cause reactionary responses from the involving biological
environment. Thus, there are some considerable barriers that should be surpassed to allow
these materials to be broadly used in the medical industry by demonstrating its safety and
efficacy for use in implantable applications.
Biocompatibility is a term used to describe how a material interacts with a biological
substance and, in 1987 the term was redefined as a materials’ “ability to perform with an
appropriate host response in a specific situation” [2.185]. Usually, this term is employed to
describe implantable materials for long-term contact with a biological system, although it
can also be used for short period interaction with biological substances in vitro. A review
by Williams [2.186] analyzed over 50 years of experience with implantable medical
devices, with main focus in the biocompatibility of such devices for long-term interactions
with the tissues of the human body. According to his study, Table 2.3 contains a summary
of the state-of-the-art in materials’ selection for long-term implantable medical devices.
With respect to polymers, PDMS is possibly the most studied implantable
biomaterial and assessed by Jewrajka et al. [2.187] as being biostable and biocompatible.
Several studies regarding biological interactions with PDMS have proven its feasibility for
use in biological systems.
Keough et al. [2.188] performed in vitro hemocompatibility tests in order to assess
the compatibility of PDMS with different blood components. The results have shown that
PDMS has small platelet accumulation comparatively to other common used biomaterials,
such as polyethylene (PE) and DacronTM
(brand name of polyethylene terephthalate),
suggesting their potential use as coatings for vascular graft materials. A similar study
proposed by Ziats et al. [2.189] reveals that PDMS presented two times less protein
adsorption than DacronTM
, which is an important factor to avoid possible pro-inflammatory
secretions, given that proteins adhere in seconds to any material in contact with the blood
and consequently affects any analysis device.
Lee et al. [2.190] presented a study on the influence of the composition of PDMS,
because not only the mechanical and physical properties are affected by the proportion of
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cross-linking in a PDMS sample, but it also influences the biocompatibility of PDMS. The
authors produced PDMS substrates having three different ratios of base and curing agent
(10:1, 10:3 and 10:0.5 w/w) and examined the attachment and growth of different kinds of
mammalian cells on PDMS. Results shown that different combinations of base and curing
agent affect cells attachment and growth for certain cell types. Also, they have foreseen
that most types of cells would be biocompatible with PDMS synthesized at normal (10:1)
or lower than normal (10:0.5) ratio of base to curing agent, and normal PDMS that had
been extracted to remove uncross-linked materials. In addition, results confirmed that the
stiffness of the PDMS samples did not influence the attachment and growth of cell, for
every cell types.
Table 2.3: Summary of materials’ selection for long-term implantable medical devices [2.186].
Material Applications
Titanium alloys Dental implants, femoral stems, pacemaker cans, heart valves,
fracture plates, spinal cages
Cobalt-chromium alloys Bearing surfaces, heart valves, stents, pacemaker leads
Platinum group alloys Electrodes
Nitinol Shape memory applications
Stainless steel Stents, orthopedic implants
Alumina Bearing surfaces
Calcium phosphates Bioactive surfaces, bone substitutes
Carbon Heart valves
UHMW polyethylene Bearing surfaces
PEEK Spinal cages
PMMA Bone cement, intraocular lenses
Silicones Soft tissue augmentation, insulating leads, ophthalmological
devices
Polyurethane Pacemaker lead insulation
Expanded PTFE Vascular grafts, heart valves
Polyester textile Vascular grafts, heart valves
SIBS* Drug eluting stent coating
* Poly(styrene-block-isobutylene-block-styrene)
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The physical and chemical properties of CNTs have raised great expectations for
several applications, with large potential in the field of medicine for therapeutic and
diagnostic applications. However, concerns regarding the biocompatibility, toxicity and
safety have limited their applications.
Some works have been demonstrating that after certain purification and
functionalization processes, pure CNTs can be transformed into functionalized CNTs with
distinct surface properties. Bianco et al. [2.191] highlights recent advances on the design of
biocompatible and biodegradable functionalized CNTs. In summary, the authors refer that
biocompatibility of CNTs can be obtained through their surface functionalization by
introducing hydrophobic functional groups that makes CNTs less biologically reactive and
biodegradability can also be obtained by chemical functionalization, in order to introduce
structural imperfections that will let oxidative enzymes degrade CNTs.
Numerous biological and biomedical applications for CNTs have been proposed
including ultra-sensitive biosensors for glucose and DNA, nanocomposites for neural
orthopedic devices, drug and vaccine delivery systems and biomaterials [2.192, 2.193].
Pyrolytic carbon, which is similar to graphite but with some covalent bonding
between its graphene sheets, produced using special CVD processes, has been used in
biomedical implants and coatings, mainly in the fabrication of heart valves prostheses.
Studies demonstrated outstanding biocompatibility properties of this material
[2.194, 2.195] and also excellent blood compatibility of pyrolytic carbon heart valves
[2.196]. Another existing carbon biomaterial is diamond-like carbon (DLC), also used as
coatings for orthopedic and cardiovascular applications. In vitro tests reported significant
improvements in biocompatibility of DLC coatings in intracoronary stents when compared
to uncoated stainless steel stents [2.197, 2.198]. Moreover, in vivo experiments of DLC
coatings for orthopedic applications showed no adverse effects on the parameters under
analysis [2.199]. Dowling et al. [2.200] studied the biocompatibility of DLC-coated
stainless steel implants into both cortical bone and muscular tissue of sheep. Both in vitro
and in vivo test showed that DLC coatings are biocompatible and a significant reduction in
the level of wear of hip implants.
Studies on pure CNTs and chemically functionalized CNTs reported a positive
impact on neuronal cells growth [2.201] and no cytotoxicity effect [2.202], suggesting the
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47
biocompatibility of CNTs as substrates for neurons. Pantarotto et al. [2.203] have used
functionalized SWCNTs to develop a vaccine delivery device by linking covalently a
neutralizing peptide from the foot and mouth disease virus to the side-wall of purified
SWCNTs. In vivo studies demonstrated that there was no cross reactivity to the CNTs,
opening possibilities for their use in vaccine delivery applications.
Existing information indicate that functionalized CNTs for drug or vaccine delivery
applications demonstrate lower toxicity compared to unmodified CNTs. However, a small
number of in vivo toxicity studies have been dedicated to this issue. Despite some
inconclusive studies, carbon-based biomaterials have shown some potential long-term
biocompatibility and potential for biomedical applications [2.204-2.211].
The added value of CNTs applications to medical area is incontestable. However,
there is still a long way to go in order for this material to become widely accepted as safe
and feasible for diagnosis and treatment purposes. A lot of work has yet to be done
concerning the long-term outcomes and consequences of this material on human health.
Also, novel methods of production and purification of CNTs should be investigated, in
order to fabricate cost-effective and affordable implantable sensors that compete with the
existing ones.
2.6 Summary
CNTs have a unique atomic structure and properties, such as high aspect ratio, high
strength-to-weight, exceptional mechanical properties and superior thermal and electrical
properties. However, the properties of CNTs depend on physical characteristics such as
diameter, chirality and wall purity.
The low density and remarkable properties of CNTs make them an attractive filler
material in polymer nanocomposites. Despite some promising results found with their
incorporation, improvements have been limited by various processing issues. It is essential
to control the embedment of CNTs into polymer matrices, since the processing steps of
CNTs mixture and matrix cure can have a huge influence on resulting nanocomposite
properties. However, most commonly in the case of CNTs/polymer nanocomposites, CNTs
are just simply mixed into polymer matrices, and van der Waals forces between
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48
CNTs avoid good dispersion. Also, the CNTs are only able to reinforce the matrix if the
adhesion between the CNTs and polymer is strong enough to facilitate load transfer.
An alternative solution to overcome difficulties in dispersing CNTs into polymer
substrates is to incorporate as-grown CNTs into the neat polymer matrices. The good
orientation of CNTs allows obtaining CNTs/polymer nanocomposites with higher
mechanical reinforcement and electrical conductivity, while controlling the alignment of
CNTs towards the pre-determined direction.
From the results reported in the literature, CNTs are promising for fabrication of both
structural and functional nanocomposites that can be widely used in several fields, such as
in the fabrications of electrical and biomedical devices.
Regarding the use of CNTs/PDMS nanocomposites for biomedical implants, the
published literature revealed many unanswered questions. Biological safety is the most
important factor when considering the medical application of CNTs as biomaterials. In
general, toxicity studies of CNTs did not find any significant respiratory inflammation or
change of the pulmonary function, but the toxicity of nanoparticles depends on particular
physiochemical and environmental factors. Also, unmodified MWCNTs demonstrated no
carcinogenic effects in animal experiments. Finally, some studies have shown that an
appropriate functionalization method can increase the biocompatibility of CNTs. They
seem to be a valuable option when considering biomedical applications, where the use of
CNTs in drug delivery, detection and tissue engineering has shown potentials for many
future advances.
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Chapter 3
Fabrication and Characterization of
A-CNTs/PDMS Nanocomposites
This chapter addresses the full fabrication process of A-CNTs/PDMS nanocomposites
films that will be further used to fabricate a flexible pressure sensor. The elastic response
of the nanocomposites are presented and related to the underlying aligned-CNT
morphology. MWCNTs at 1% Vf are embedded in a flexible substrate of PDMS to create a
flexible polymer nanocomposite (PNC).
The PNC properties are evaluated using scanning electron microscopy (SEM),
differential scanning calorymetry (DSC), dynamic mechanical analysis (DMA) and tensile
mechanical tests, and the full linearly-elastic constitutive relation established for such a
PNC. Along with the mechanical characterization of the nanocomposites, electrical
conductivity tests were also performed using the Van der Pauw method.
Part of this chapter is adapted from the following publication:
A. T. Sepulveda, R. Guzman de Villoria, J. C. Viana, A. J. Pontes, B. Wardle, and L. A. Rocha,
“Full Elastic Constitutive Relation for Non-isotropic Aligned-CNTs/PDMS Flexible
Nanocomposites,” Nanoscale, vol. 5, no. 11, pp. 4847-4854, 2013.
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3.1 Introduction
As demonstrated in the previous chapter, CNTs are an excellent candidate as reinforcement
material in polymeric materials for multi-functional property enhancement (mechanical,
electrical, etc.) due to their exceptional physical properties [3.1-3.6] and low density,
particularly if the properties can be tailored anisotropically via aligning the CNTs ‘fibers’.
Also, the recent optimization in A-CNTs growth, opened opportunities for the
development of novel materials and applications since CNTs/Polymer-based
nanocomposites properties are enhanced by this morphology that organizes the CNTs
along a common direction [3.7, 3.8].
Most of the studies reported in the literature concerning CNTs composites have
concentrated largely in the characterization of CNTs/Polymers composites, where the
SWCNTs/MWCNTs are dispersed or mixed within the polymeric matrix. The work
presented here aims the study of A-CNTs/PDMS nanocomposite films for flexible pressure
sensor applications. The fabrication of the nanocomposites and the resulting
morphological, thermal, mechanical and electrical properties, are presented and compared
with pure PDMS. The polymer reinforcement using A-CNTs generates a non-isotropic
material that can be very interesting for some particular applications such as capacitive and
membrane-based sensors, energy harvesters, and electrically-conductive biomedical device
elements.
In addition, the first full non-isotropic constitutive relation for the aligned PNC is
presented, that may be useful in modeling PNCs as composites or as elements of
hierarchical nanoengineered composites, particularly PDMS-CNTs PNCs envisioned as
elements in biomedical devices such as pressure transducers and energy harvesters.
3.2 Synthesis of MWCNTs
The CNTs used in this work were grown in situ, on the Department of Aeronautics and
Astronautics facilities at MIT, using a modified chemical vapor deposition (CVD) method
[3.9], where the vertically aligned MWCNTs forests are synthesized using a Al2O3/Fe
catalyst layer. A silicon (Si) wafer substrate with a ~ 300 nm-thick thermally grown silicon
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dioxide (SiO2) layer is used as the support layer. After that, a layer of alumina oxide
(Al2O3) with a thickness of ~ 10 nm and a catalyst layer of iron (Fe) with a thickness of
~1 nm are physical vapor deposited (PVD) onto SiO2 using electron beam evaporation.
The A-CNTs are attached to the Fe nanoparticles which are formed upon annealing in
hydrogen (H2). During CNTs growth, the Fe is supersaturated with carbon (C), and upon
growth termination and exposure to ambient a combination of amorphous carbon and Fe
oxide attaches the CNTs to the Fe. The CVD process steps were already summarized in
§2.2.3 (Figure 2.13). One of the advantages of this process is the possibility to pattern the
catalyst, typically fabricated using a lift-off process, which enables the design of regions
where the CNTs forests will grow (some of the masks used in this work are presented in
Appendix A.1).
Prior to the growth process, such substrates prepared as referred before were cut into
~ 32 x 14 mm2 rectangles shape with a diamond cutter. The areas of catalyst on photoresist
were cleaned with acetone in an ultrasonicator (VWR, Model 150D) for 3 min. The CVD
process was performed at atmospheric pressure in a Lindberg/Blue M tube furnace, Model
55347 using a 50.8 mm (2 inch) diameter quartz tube, where the previously cut substrates
were placed (Figure 3.1). After the purge of the lines and the tube with helium (He), the
tube was heated to 750 ºC and 400 sccm of ethylene (C2H4) gas was introduced for the
CNTs growth, during the time selected for the experiment. Finally, 2400 sccm of He was
flushed while the tube cools down. This method has the advantage of enabling the growth
of high purity, high yield vertically aligned CNTs. Also, for the work presented here there
was no modification or functionalization on the CNTs surfaces after growth.
Figure 3.1: CVD system for growing A-CNTs forests.
Fabrication and Characterization of A-CNTs/PDMS Nanocomposites
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First CNTs were grown in the form of dense pillars of A-CNTs and using different
shapes. The aim was to optimize the growth process previously implemented by the
TELAMS group at MIT, and change the growth time to give the CNTs forest the desired
thickness (up to 400 µm). Figure 3.2 shows SEM images of A-CNTs forests grown on
patterned catalyst with dot shape and Y-shape, and Table 3.1 summarizes the growth
recipes (the experiments accomplished considering different substrates shape and growth
times are presented in Appendix A.2).
Figure 3.2: SEM images of as-grown A-CNTs forests. a) Dot shape, b) Y-shape and c) magnified
region of b).
Table 3.1: CNTs growth recipe for substrates with patterned catalyst in dot shape and Y-shape.
Gas SCCM* Time (min) Temperature
Flush Tube He 2070 10 Room temperature
Heat Tube H2 1040 Ramp to 750 ºC
15 Held at 750 ºC
Growth H2 1040
0.6/0.8/2 Held at 750 ºC C2H4 400
Flush Tube He 2400 10 Cool down
*SCCM – Standard Cubic Centimeters per Minute
Given that the objective is to fabricate A-CNTs/PDMS nanocomposite films,
unpatterned rectangular Si substrates (32 x 14 mm2) were also prepared for the growth of
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A-CNTs. Different growth times were tested in order to obtain CNTs forests with
thicknesses of about 400 µm. As seen in Figure 3.3, CNTs show a good alignment and the
height of the grown bulk MWNTs forest was measured to be ~ 389 µm, obtained for the
growth time of 10 min. Table 3.2 summarizes de recipe of the A-CNTs growth using the
rectangular Si substrates.
Figure 3.3: CNTs grown on Si substrate. a) Optical image of an A-CNTs forest, b) SEM image of
A-CNTs forest and c) SEM image showing the alignment of the CNTs.
Table 3.2: CNTs growth recipe for substrates with patterned catalyst in rectangular shape.
Gas SCCM* Time (min) Temperature
Flush Tube He 2070 10 Room temperature
Heat Tube H2 1040 Ramp to 750 ºC
15 Held at 750 ºC
Growth H2 1040
10 Held at 750 ºC C2H4 400
Flush Tube He 2400 10 Cool down
*SCCM – Standard Cubic Centimeters per Minute
The as-grown MWCNTs forests using this procedure have been extensively
characterized. On average, MWCNTs have diameters around 8 nm (2-3 walls) and the
space between CNTs is of about 80 nm [3.10, 3.11]. Thus, assuming the MWCNTs as solid
tubes, the volume fraction for these CNTs will be ~ 1 % (an estimate of the A-CNTs
volume fraction can be found in Appendix A.3).
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3.3 Fabrication of A-CNTs/PDMS Nanocomposite Films
The entire nanocomposite structure is fabricated from A-CNTs embedded in a flexible
substrate of PDMS, a commercially available silicone based polymer. Sylgard® 184 from
Dow Corning Corporation was used in this work. Its main properties are summarized in
Table 3.3.
The base elastomer and curing agent were manually mixed using a glass stick at a
10:1 weight ratio, respectively and, before pouring the mixture into a mould, it was
systematically degassed in a vacuum chamber during 15 min, in order to remove all
bubbles.
Table 3.3: PDMS main properties (10:1 weight ratio).
Substrate PDMS
Manufacturer Down Corning ®
Grade Sylgard ® 184
Physical form Base + curing agent
Thickness (mm) Free
Density (g/cm3) 1.03
Thermal conductivity (W/m.K) 0.18
Dielectric (kV/mm) 21.2
Service temperature (ºC) -45 to 200
The process of embedding the A-CNTs is performed using a mold to define the
nanocomposite geometry. Polymethylmethacrylate (PMMA) was selected as the mold
material, mainly due to its mechanical properties and low cost. The PMMA mold geometry
was defined by computer numerical control (CNC) milling, a technique that enables low
costs and fast production times. The procedure to embed the pure PDMS into the as-grown
A-CNTs (1% Vf) is schematically represented in Figure 3.4.
Fabrication and Characterization of A-CNTs/PDMS Nanocomposites
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Figure 3.4: Fabrication process flow for the development of a flexible A-CNTs/PDMS
nanocomposite.
The PMMA mold was produced with precise dimensions of 32 x 14 x 0.4 mm3
(Length x Width x Height) in order to accept the wafer with the as-grown A-CNTs
(average length of ~ 400 µm). The A-CNTs/PDMS nanocomposites are fabricated by
pouring slowly the uncured PDMS into the as-grown A-CNTs through capillarity wetting.
After leaving the A-CNTs approximately 1 min with the PDMS substrate for complete
wetting, the assembly is cured for 30 min in an oven at 85 ºC. Finally, the A-CNTs/PDMS
nanocomposites are manually demolded from the PMMA mold. Typical A-CNTs-PNC
samples obtained right out of the molds are shown in Figure 3.5.
Figure 3.5: Flexible PNC film made of A-CNTs embedded in a polymeric matrix of PDMS.
Fabrication and Characterization of A-CNTs/PDMS Nanocomposites
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3.4 Thermal Characterization and Results
3.4.1 Differential Scanning Calorimetry
The differential scanning calorimetry (DSC) measures the differential power between
a polymeric sample in a sealed aluminum pan referenced to an empty pan. During heating
and cooling temperature scans, a zero temperature differential between pans is preserved.
Thermal properties tests of pure PDMS and A-CNTs/PDMS nanocomposite samples
were performed on a TA Instruments DSC2920 differential scanning calorimeter,
connected to a cooling system and calibrated using the melting temperature of indium as
reference. The samples weights ranged from around 5 to 7 mg (6 samples of pure PDMS
and 6 samples of A-CNTs/PDMS nanocomposite). The temperature was scanned from
-170 ºC to -30 ºC at a heating rate of 10 ºC/min and the glass transition temperature (Tg) of
all samples was obtained from the DSC thermograms, always using the same experimental
procedure to obtain reproducible results for comparison.
Tg was obtained as the temperature at which the regression line, generated through
the trace below the starting point of the transition, crosses the tangent drawn from the
inflection point of the glass transition step. Figure 3.6 shows the representative DSC
thermograms obtained for pure PDMS and A-CNTs/PDMS nanocomposites during
heating.
The location of Tg of the pure PDMS is seen at around -125.95 ± 0.27 ºC, which is in
good agreement with the accepted reported values and that are in the range of
-120 to -125 ºC [3.12-3.17]. On the other hand, the glass transition temperature of the
nanocomposites appears at -126.79 ± 0.11 ºC, suggesting that the reinforcement with CNTs
causes almost no effect in the Tg of the pure PDMS.
Fragiadakis and Pissis [3.18] observed no variations of Tg upon the incorporation of
nanosilica particles, but the decrement of the heat capacity was attributed to the
immobilization of polymer at the surface of silica particles.
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Figure 3.6: Representative DSC thermograms of pure PDMS and A-CNTs/PDMS
nanocomposites.
The observed profiles in the DSC curves for both pure PDMS and nanocomposites
suggests that there is also no observable change in Tg after incorporation of 1% Vf of
CNTs into the polymeric matrix. In this case, the presence of the 1% Vf A-CNTs does not
significantly restrict the segmental mobility of the amorphous phase over the entire volume
of the PDMS.
3.5 Morphological Characterization and Results
3.5.1 Scanning Electron Microscopy
Scanning electron microscopy (SEM) was used to assess the alignment of the CNTs,
as well as the distribution of the A-CNTs in the PDMS substrate. Imaging of as-grown
A-CNTs forests (Figure 3.7) and embedded A-CNTs forests in the PDMS matrix
(Figure 3.8) were performed in a Nova NanoSEM 200 scanning electron microscope,
operating at 5 kV accelerating voltage, after sputtering the nanocomposites with 4 nm of
Au/Pd. Figure 3.9 shows a SEM image of a top-view surface of pure PDMS, taken at same
conditions for comparison.
Fabrication and Characterization of A-CNTs/PDMS Nanocomposites
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78
(a)
(b)
CNTs
alignment
Figure 3.7: SEM image of as-grown A-CNTs. (b) and (c) are magnified regions of (a).
The A-CNTs/PDMS nanocomposite samples were mounted in two different ways:
top-view surface of the nanocomposites and cross-sectional cut parallel to the CNTs
growth axis.
1% Vf VACNTs/PDMS nanocomposite
Top-view Surface
Cross-sectional Cut
(Parallel)
Figure 3.8: SEM images of A-CNTs/PDMS nanocomposites revealing the CNTs structure
embedded in the PDMS substrate. (a) Top-view and (b) parallel cross-sectional cut.
CNTs
growth axis
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Figure 3.9: SEM image of a top-view surface of pure PDMS.
SEM images of the nanocomposites (Figure 3.8) show the CNTs being totally
impregnated by the PDMS, i.e., PDMS has uniformly wet the CNTs forest and bonded all
CNTs together during the curing process. Also, comparing the top-view surfaces of both
pure PDMS and A-CNTs/PDMS nanocomposites samples it is possible to observe that
there are no significant changes after incorporation of A-CNTs, suggesting that CNTs are
all embedded in the PDMS matrix.
For cross-sectional images, the nanocomposite samples were cut with a bistoury.
Although this cutting method may influence the surface morphology, SEM images reveal
that the CNTs alignment observed after the A-CNTs growth is retained after the process of
wetting and curing of the PDMS.
3.6 Mechanical Characterization and Results
3.6.1 Tensile Tests
The impregnation of the A-CNTs in the PDMS matrix, as well as the respective
mechanical and electrical properties of the flexible nanocomposite films (required for the
sensor design and its response), is the most important step of the fabrication process. The
A-CNTs are oriented in the out-of-plane direction (normal to the wafer plane). In the
in-plane direction the polymer-based nanocomposite can be assumed transversely
isotropic. In addition, due to the A-CNTs orientation, the improvement of the elastic
modulus is expected to be minimal as the long axis of the A-CNTs is oriented normal to
Fabrication and Characterization of A-CNTs/PDMS Nanocomposites
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the loading direction, such that the PDMS polymer controls the response (see the
schematic represented in Figure 3.10).
Figure 3.10: Schematic of the CNTs orientation inside the PDMS matrix (legend: W – width,
H – height and P – Load direction).
Recent work on nanocomposites has shown considerable increase in modulus, due to
the alignment of the CNTs in the polymer (PDMS) [3.19] and epoxy [3.20], but based on
experimental results, little reinforcement effect is expected in the nanocomposites where
the A-CNTs are oriented normal to the loading direction.
An Instron 4505 mechanical materials testing machine was used to conduct tensile
tests at room temperature of pure PDMS and A-CNTs/PDMS nanocomposite samples.
Several individual A-CNTs/PDMS nanocomposite samples, molded into rectangular
shapes with dimensions of approximately 32 x 14 x 0.4 mm3 were tested in tension to
mechanically characterize the nanocomposite films in both longitudinal (3-direction) and
transverse (1- or 2-directions) (with the CNTs oriented normal and parallel to the load, as
shown in Figure 3.11).
A total of 12 nanocomposite specimens, 6 tests for the transverse direction (1,2
plane), 3 tests for the z-direction (1,3 plane) and 3 for longitudinal direction (3,2 plane),
were clamped in custom grips with a crosshead speed of 0.8 mm/min. The load cell used
was of 1 kN and the gage length was 10 mm for all samples. For pure PDMS, 3 samples
were tested for comparison, under the same parameters as used for the A-CNTs/PDMS
samples. Poisson’s ratio was obtained in all the tensile tests using the corresponding strains
in longitudinal and transverse directions.
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Figure 3.11: Optical images showing the changes before tensile test (unstretched) and during the
test, for pure PDMS and A-CNTs/PDMS samples (with the CNTs oriented transverse to the
load – 1,2 plane and 1,3 plane, and parallel to the load – 3,2 plane).
Strain measurement by optical extensometry was used during the tensile tests.
Several dot features, with diameters of about 0.5 mm and separation distance of
approximately 0.4 mm, where speckled with black and white paint by hand in the pure
PDMS and nanocomposite samples, respectively, and the tensile tests were recorded
(1920x1080 resolution). During the tests, the dots undergo translation from its initial
position allowing calculation of the deformations in both directions (ε1 and ε2, ε1 and ε3, ε3
and ε2 for tests occurring in the 1,2 plane, 1,3 plane and 3,2 plane, respectively). As soon
as the tests were concluded, the recorded movies were fragmented in JPEG images with an
increment of 1 second. Then, the sample deformation and strain was calculated by a
custom MATLAB code for strain mapping. As represented in Figure 3.12, the central pixel
of four selected points is followed from the beginning to the end of the test in order to
quantify the displacement in each direction. The points illustrated as P1 and P2, and P3 and
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P4 should be parallel to the x-axis, as well as P1 and P3, and P2 and P4 parallel to the y-axis.
The ratios P1-P2 and P3-P4 allow obtaining the displacement in xx while P1-P3 and P2-P4
allow obtaining the displacement in yy.
Figure 3.12: Illustration showing the points P1, P2, P3 and P4. a) Beginning of the test and b)
displacement in the x- and y-axis during the tensile test.
The stress-strain curves in the longitudinal and transverse directions were generated
in the MATLAB software and the Young’s modulus (E) was automatically calculated from
the stress-strain curves shown in Figure 3.13.
Figure 3.13: Representative stress-strain curves of pure PDMS and reinforced A-CNTs/PDMS
nanocomposites in the longitudinal (3,2 plane) and transverse (1,2 plane) directions. Longitudinal
is in the 3-direction and transverse the 1-, or 2-direction.
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Properties of the PNCs depend on the elastic modulus and aspect ratio of the CNTs,
as well as the CNT array alignment, and therefore an effective reinforcement requires that
the individual CNTs are fully wet with PDMS to guarantee load transfer from PDMS
matrix to CNTs [3.21]. The morphologic characterization (Figure 3.8) shows the A-CNTs
entirely embedded on the PDMS matrix, hence a higher Young’s modulus value for the
membranes in which the CNTs are aligned parallel (longitudinal) to the load direction is
expected.
The experimental measurements indicate that the pure PDMS membranes have a
Young´s modulus of about 0.8 ± 0.1 MPa which agrees with the literature values
[3.22-3.23]. The A-CNTs/PDMS nanocomposite has an increase in the Young´s modulus
in the transverse direction (1.7 ± 0.25 MPa obtained in both transverse directions, 1,2 plane
and 1,3 plane). In the longitudinal direction, 3,2 plane, the increase is more significant
(8.1 ± 1.2 MPa). No assessment of strength or toughness can be made regarding the
longitudinal direction, as the samples failed outside the region of the A-CNTs in the pure
PDMS region.
The Poisson´s ratios (υ) of pure PDMS and A-CNTs/PDMS films (in both
longitudinal and transverse directions) were obtained from the tensile tests performed.
Assuming the coordinate system illustrated in Figure 3.12, in which the samples are
stretched along the axial direction (the x-axis), the ratio of lateral strain and axial strain
allow calculating the Poisson´s ratio:
For the calculations regarding the changes of the width (ΔW) and height (ΔH) during
the tensile tests, the Poisson´s ratio can be written as:
where, H0 and W0 are the initial height and width, respectively.
Poisson´s ratio of PDMS obtained from the tested samples was of about 0.40 ± 0.05,
similarly to other works reported in literature [3.24, 3.25]. The Poisson´s ratios for the
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A-CNTs/PDMS samples were retrieved from the three tensile tests performed. From the
changes of width and height during the tensile tests, the Poisson´s ratio obtained for both
perpendicular and longitudinal directions, was of
, respectively. The decrease observed in the Poisson’s ratio
for both directions as compared with the Poisson´s ratio of PDMS is in line with previous
results, since the CNTs restrict the shrinking of the nanocomposite for a given applied
load. Therefore, a lower Poisson’s ratio value for the samples in which the CNTs are
aligned transverse to the load (1,3 plane) was expected, since in this direction the CNTs
confer a great resistance during the stretching.
3.6.2 Dynamic Mechanical Analysis
Dynamic mechanical analyses (DMA) were carried out using Triton Technology
equipment, TTDMA, in order to determine the shear modulus as function of frequency, at
isotherm temperature.
A total of 10 samples (5 of pure PDMS and 5 of A-CNTs/PDMS nanocomposites)
were tested in a simple shear mode (dual sample) over 1 to 80 Hz frequency sweep, with 6
points per decade and at a temperature of 40 ºC. Following the instructions recommended
by the manufacturer, the samples were mounted in pairs, as schematically presented in
Figure 3.14.
Figure 3.14: Schematic of the shear mode (dual sample) experiment.
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Figure 3.15 shows representative curves of the storage modulus (G’) and the loss
factor tan δ (G’’/G’), as a function of frequency for both pure PDMS and A-CNTs/PDMS
samples. In this study, the experimental range covers only the service frequency range for
the future application of the material (0-80Hz).
Figure 3.15: a) Storage modulus – G’ and b) tangent delta – tan δ, as function of frequency for
pure PDMS and A-CNTs/PDMS samples.
From the results presented in Figure 3.15, both materials demonstrate elastic-like
behaviour (low loss factor) and as the testing frequency increases the storage modulus
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tends to increase towards higher frequency. At a low frequency of 1 Hz, the material
presents a storage modulus of 0.19 ± 0.01 MPa for pure PDMS samples and
0.24 ± 0.01 MPa for A-CNTs/PDMS nanocomposite samples, while at a frequency of
80 Hz the obtained storage modulus is 0.25 ± 0.01 MPa for pure PDMS samples and
0.32 ± 0.01 MPa for A-CNTs/PDMS samples. Incorporation of 1 % Vf A-CNTs into the
pure PDMS generated a slight improvement on the storage modulus in comparison to the
pristine PDMS.
Both PDMS and the nanocomposite loss curves present similar behaviour with a
small movement upward with frequency. The incorporation of A-CNTs into the polymer
matrix tends to increase the modulus owing to the effective load transfer. Although
enhanced Young’s modulus (E) was observed in A-CNTs/PDMS nanocomposite samples
compared with the pure PDMS samples, only slightly stiffer shear modulus (G’) of the
A-CNTs/PDMS nanocomposites was observed in the DMA tests.
3.7 A-CNTs/PDMS Nanocomposite Full Constitutive Law
The A-CNTs/PDMS nanocomposite is a transversely-isotropic material (in the linear
regime) and therefore its elastic behaviour can be completely defined by 5 independent
elastic constants and except for the shear modulus, the elastic constants were already
obtained from the tensile tests. The G32 value is obtained from the dynamical mechanical
tests, as presented previously.
The nanocomposite material under study, and as already mentioned, is a transverse
isotropic material. In this case, the expression for generalized Hooke’s law is given as
shown in equation 3.3.
(3.3)
where εij is the strain tensor, σkl is the stress tensor and Sijkl is the compliance tensor. Also, a
transverse isotropic material has 5 independent engineering constants:
In the symmetry plane (1,2 plane):
Young´s modulus (E2),
Poisson’s ratio (υ12),
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In the CNTs direction (L, 3):
Young´s modulus (E3),
Poisson’s ratios (υ23 and υ32),
Shear modulus (G32),
In addition, a transversely isotropic material satisfies reciprocity conditions (3.4):
(3.4)
From the expression for generalized Hooke’s law, it is now possible to write the
constitutive equations for the A-CNTs/PDMS nanocomposites, completely characterized
here. Considering the complex shear modulus, , the values of the
engineering constants are the following:
In the symmetry plane (1,2 plane):
In the CNTs direction (L, 3):
and
(reciprocity holds line all three planes, being also trivially satisfied in the other
planes – 1,2 and 1,3 planes).
The results obtained demonstrate anisotropic behaviour of the
A-CNTs/PDMS nanocomposites. The Young’s modulus in the longitudinal direction is
higher compared to the obtained in the transverse direction.
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From the engineering constants presented above, the Hooke's law for the
A-CNTs/PDMS nanocomposites in compliance matrix form can be written as shown in
equation (3.5).
(3.5)
To our knowledge, this is the first time that the full constitutive law for an aligned
polymer nanocomposite is presented. As mentioned before, A-CNTs/PDMS
nanocomposites were manufactured from 1 % Vf A-CNTs proceeded by capillarity wetting
with a PDMS substrate, without functionalization of the CNTs. In general, the Young´s
modulus of PNCs is found to increase with the CNTs alignment in the matrix. However,
the measurement results show that the Young’s modulus remain far from the idealized
theoretical prediction from the simple rule of mixtures [3.26, 3.27].
Assuming perfect alignment of the CNTs, perfect wetting of the CNTs by the PDMS
and perfect bonding between the interface of the CNTs and the PDMS, the Young’s
modulus of an A-CNTs/PDMS nanocomposite, at 1 % Vf A-CNTs with a CNTs modulus
of 1 TPa and PDMS substrate modulus of 0.8 MPa, is expected to be 10 GPa, instead of
the 0.0081 GPa obtained from the mechanical tests. The difference between the theoretical
calculations and experimental measurements could be due to assumptions in the simplified
model, but is primarily attributable to the CNT waviness effect, which is predicted by some
authors [3.20, 3.28-3.30] to have considerable effect on both longitudinal and transverse
nanocomposite Young’s modulus.
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3.8 Electrical Characterization and Results
3.8.1 Conductivity Measurement
Along with the mechanical characterization of the A-CNTs/PDMS nanocomposites,
electrical tests were also performed. The electrical conductivity measurements were
performed based on the Van der Pauw technique [3.31].
In the Van der Pauw method, reference specimens with square or rectangle geometry
are used. These specimens are contacted at the four corners, and the resistance is
subsequently measured as illustrated in Figure 3.16. During the tests, two consecutive
measurements are performed, in which the current (I12) is applied into contact 1 and out of
contact 2, determining the voltage V43 from the contact 4 to contact 3. Then, applying the
current (I23) into contact 2 and out of contact 3, the voltage V14 is obtained from the contact
1 to contact 4.
Note: RA = (RA1+RA2)/2 and RB = (RB1+RB2)/2
Figure 3.16: Schematic of a Van der Pauw configuration used in the measurement of the two
characteristic resistances RA and RB.
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The aim of the resistivity measurement is to obtain the sheet resistance (RS). Van der
Pauw proved that the resistance values, RA and RB, are related to the sheet resistance RS
through the Van der Pauw equation, that can be solved numerically for RS:
(3.6)
The electrical resistivity (ρ) can be determined applying equation 3.7.
(3.7)
where d is the thickness of the sample.
The electrical conductivity (σ) is defined as the inverse of resistivity and is given by
equation 3.8:
(3.8)
Electrical conductivity measurements were performed at ambient temperature on
A-CNTs/PDMS nanocomposite samples (a total of 9 samples were tested). An automated
Keithley picoammeter/voltage source (Keithley Instruments Inc.) was used for V and I
sourcing and measuring. A constant voltage of 10 V was applied to the nanocomposite
samples and the respective values of the current were measured. In addition, the polarity of
the current was switched in all Van der Pauw measurements, allowing obtaining the values
of the current I12, I43, I23 and I14 (as illustrated in Figure 3.16).
The measured electrical conductivity of the A-CNTs/PDMS nanocomposite samples
was of 0.35 S/m with a standard deviation of 0.37 S/m (see Figure 3.17). The rather low
conductivity does not compromise the developed capacitive sensor, but since passive
telemetry is required to measure and power the sensor, it has major implications if one
wants to design an inductor using this approach, i.e., the resistive losses would be so high
that a large amount of power would be required to power and read the sensor. Another
work with the same A-CNTs, but embedded in an epoxy polymer, showed electrical
conductivities in the axial direction of ~ 10 S/m [3.20].
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In general, incorporation of CNTs into a polymeric matrix of PDMS improves the
electrical conductivity of pristine PDMS and as higher is the weight percentage of CNTs
embedded in the matrix, higher is the conductivity of the nanocomposite. However, a study
where low weight percentages of MWCNTs (0.8, 1, 1.5, 2, 2.2 and 2.5 wt. %) were mixed
in PDMS, also presented low electrical conductivity of the nanocomposites. At 1 wt. %
concentration of MWCNTs the authors found an electrical conductivity of ~ 0.005 S/m,
while a value of 0.02 S/m was observed at 2.5 wt. % concentration [3.32].
Figure 3.17: Histogram of the electrical conductivity measured on the A-CNTs/PDMS
nanocomposite samples.
The low conductivity values obtained for the A-CNTs/PDMS nanocomposite forced
a change on the envisaged fabrication process since the inductors could not be built using
this method. An alternative solution (introduced in §5.4) was to implement the inductor
using carbon nanofibers (typically have conductivity 4 orders of magnitude higher than the
A-CNTs/PDMS nanocomposite).
The mechanical-resistance relationship of the A-CNTs/PDMS nanocomposite was
further characterized with an Instron 4505 tensile mechanical testing machine. A total of 7
samples with approximate dimensions 32 x 14 x 0.4 mm3 where subject to tensile tests on
transverse direction (1,2 plane) while measuring the resistance of the samples. As shown in
Figure 3.18, metal wires were embedded in the nanocomposite prior to cure of the
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nanocomposite for the resistance readout. In order to reduce the contact resistance, there
was a ~20 nm-thick layer deposited on the samples edge to improve electrical contact.
Figure 3.18: Digital image of the A-CNTs/PDMS nanocomposite samples prepared to measure the
resistivity while performing the tensile tests.
The experimental test setup is shown in Figure 3.19. The nanocomposite samples
were clamped in custom grips and tested with a crosshead speed of 0.8 mm/min. The load
cell and the gage length was the same used in the tensile tests (§3.6.1), 1 kN and 10 mm,
respectively.
Figure 3.19: Digital image of the experimental test setup.
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The two sides of the grips were covered by electrical tape, in order to insulate
electrical wires and the other material existent in the tensile testing machine. An Agilent
82357B USB/GPIB Interface (Agilent Technologies, Inc.) was used to establish a direct
connection between a Keithley 2110 multimeter and a computer. During the testing, the
resistance of the nanocomposite samples were recorded through the multimeter connected
to the metal wires. From the measured resistance (R) and strain (ε), the gauge factor (GF)
was determined by equation 3.9. Figure 3.20 gives a representative curve profile of the
relationship between the measured resistance changes and the applied strain.
(3.9)
where R0 is the initial resistance, and can be written as:
(3.10)
Figure 3.20: Measured resistance changes of the A-CNTs/PDMS nanocomposites during tensile
mechanical tests, with initial resistance (R0) of 0.3 MΩ.
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In order to calculate the gauge factor, the obtained curve was divided in three
sections. In the small tested strain range (< 5 %) the nanocomposite samples revealed
linear mechanical-resistance changes. Thus, the gauge factor (GF1) was calculated by
fitting the curve in this section applying equation 3.10. For higher tested strain ranges
(5 % to 15 % and 15 % to 30 %) a non-linear relationship was observed. Therefore, to
compare the gauge factors as function of strain, the values were determined at strain of
15 % and 30 % (entitled as GF2 and GF3, respectively), in accordance with equation 3.10.
The results shown that the A-CNTs/PDMS nanocomposites have high gauge factor
(GF1 ~ 50, GF2 ~ 38 and GF3 ~ 29) making it a good material for sensor applications
using the piezoresistive effect, as alternatives to the commonly used silicon based strain
sensors (GF = 20 – 120) [3.33]. However, due to the high contact resistance observed
during the measurements, the piezoresistance effect needs to be hereafter better
characterized.
The electrical characterization revealed the unsuitability of the nanocomposites for
the inductor fabrication, and alternative solutions were envisaged to overcome this
technical difficulty.
3.9 Summary
In this chapter, A-CNTs/PDMS nanocomposites were fabricated and studied, aiming its
use in flexible applications. The nanocomposites are suitable for flexible applications with
the advantage of being electrical conductive and with mechanical non-isotropic mechanical
properties.
The morphological characterization performed suggests that CNTs preserve
alignment after the wetting process, which allows a controlled manufacturing of
A-CNTs/PDMS nanocomposites. The good dimensional control, the interesting
non-isotropic properties and the easily processing characteristics makes it an interesting
material for the realization of microsensors and microactuators.
Tensile tests were performed in order to completely characterize the A-CNTs/PDMS
nanocomposites. Compared to the Young’s modulus of the unreinforced PDMS matrix
(0.8 MPa), we obtained an increase in the Young’s modulus of the nanocomposite in the
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transverse direction (1.7 MPa) and a higher increase in the longitudinal direction
(8.1 MPa). The results from the set of mechanical tests were used to derive the full
constitutive law of the nanocomposite. To our knowledge, this is the first full constitutive
law for a proper aligned polymer nanocomposite and, within the uncertainty of the tensile
tests results, reciprocity between different directions is observed validating the results.
The nanocomposites have already been successfully applied in the fabrication of
flexible pressure sensors [3.24, 3.34]. The unique properties of the nanocomposites
(electrical conductivity and low Young’s modulus) are the enabling factor for the flexible
sensors. In addition, the non-isotropic properties can be used to improve some specific
mechanical behaviour of the sensor.
A key application area for the flexible sensors is implantable devices. The
nanocomposites dimensions are fully controlled and the high flexibility of the material
makes it possible to fabricate devices that can be deployed using minimally evasive
techniques. The PDMS guarantees the required biocompatibility of the devices.
Regarding the electrical characterization, the results obtained raise a main concern
about the suitability of the realization of a pressure sensor with passive telemetry using the
technology based on A-CNTs/PDMS nanocomposites. The rather low conductivity of the
PDMS membranes with embedded A-CNTs will generate a rather high resistance of the
inductor and therefore the energy transferred inductively will be dissipated in the resistor.
Solutions to overcome this problem were studied, and the use of carbon fibers encapsulated
by a matrix of PDMS to fabricate an inductor, as well as the use of a flex PCB technology
seems to be a good alternative.
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[3.32] A. Khosla and B. L. Gray, “Preparation, characterization and micromolding of
multi-walled carbon nanotube polydimethylsiloxane conducting nanocomposite
polymer,” Materials Letters, vol. 63, no. 13–14, pp. 1203–1206, May 2009.
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[3.33] E. Peiner, A. Tibrewala, R. Bandorf, H. Lüthje, L. Doering, and W. Limmer,
“Diamond-like carbon for MEMS,” Journal of Micromechanics and
Microengineering, vol. 17, no. 7, pp. S83–S90, Jul. 2007.
[3.34] A. T. Sepúlveda, R. G. De Villoria, J. C. Viana, A. J. Pontes, B. L. Wardle, and
L. A. Rocha, “Flexible Pressure Sensors: Modeling and Experimental
Characterization,” Procedia Engineering, vol. 47, pp. 1177–1180, Jan. 2012.
Chapter 4
Fabrication and Characterization of
Flexible Capacitive Pressure
Sensors
In this chapter, a new approach for the fabrication of A-CNTs/PDMS
nanocomposite-based pressure sensors for wireless telemetry is presented.
The capacitive flexible sensor is expected to be included in a passive telemetry
measurement system including the capacitive pressure sensor coupled with an inductor. In
this section, only the performance of the flexible pressure sensor is presented. The
telemetry system architecture is introduced but its application and results are presented in
chapter 5. In this chapter, flexible capacitive pressure sensors fabricated with PDMS-based
nanocomposites are experimentally characterized and results compared with simulations
from both FEM and analytical modeling.
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4.1 Introduction
The advances in microsystems technologies (MST) are providing new opportunities for a
diversity of applications [4.1] and in a variety of areas [4.2], mostly due to the small scale
of the devices (usually planar areas below 4 mm2 [4.3]), low power consumption and low
fabrication costs. However, in various industrial applications the measurements in harsh
environments could be difficult and dangerous for an individual. For this reason,
transmitting data from the sensor wirelessly is less risky and, with the recent
developments, these acquisition systems will be able to change and improve a number of
areas, including the health care system.
Commercial available sensors can measure a vast number of parameters, such as
conductivity and pressure [4.4, 4.5]. Pressure sensors are an example of first use of
microsystems and are one of the most successfully commercialized devices. A wide variety
of pressure sensors have been developed to measure pressure in a huge range of
applications over several years, and they can be grouped into four classes, taking into
account their sensing characteristics: optical, piezoelectric, piezoresistive and capacitive
[4.6-4.9]. Capacitive sensors are among the commonly used pressure sensors, as they have
simple configurations, have improved sensitivity and are widely independent of
environmental temperature oscillations. Also, capacitive pressure sensors are less costly to
produce and have good compatibility with microelectromechanical system (MEMS)
fabrication techniques. As referred, the physical structures of the capacitive pressure
sensors are fairly simple. Basically, the devices include two parallel conducting plates
(electrodes) separated by a dielectric (air or an insulating material) in between, as
schematically presented in Figure 4.1.
Figure 4.1: Illustration of a capacitive pressure sensor based upon a parallel plate arrangement.
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Capacitive pressure sensors are based on the use of a deflectable diaphragm that
allows the transduction mechanism. As the flexible electrodes deflect under applied
pressure, the effective distance between the two plates or electrodes changes, thus
changing the capacitance of the device.
Ignoring fringing fields, the capacitance for a simple parallel plate capacitor is given
by equation 4.1:
where ε0 is the permittivity of the free space, εr is the relative permittivity of the material
between the plates (dielectric), A is the area of the electrodes and d is the separation
distance between electrodes.
Typically, fabrication of a pressure sensor uses hard and inorganic materials, such as
fabrication on a silicon substrate [4.10-4.15]. However, silicon-based capacitive devices
need not only high sensitivity but also flexibility, in order to detect the pressure applied to
the surfaces. A large amount of pressure sensors commercially available have lack of
flexibility to fit into a curved interface, an important factor that could improve the
measurement accuracy and the durability of the sensor.
Many attempts have been focused on the development of a flexible MEMS technique
[4.16, 4.17], as well as the development of polymer-based flexible sensors [4.18-4.24]. The
use of polymer-based diaphragm allows achieving much superior deformation under low
pressure and higher sensitivity, when compared to the conventional silicon-based pressure
sensors [4.25]. Research on CNTs-based sensors, i.e., sensors that employ CNTs as the
transduction element, has been very active in the past years [4.26], whereas the work on
nanocomposites for microsystems has been primarily focusing on materials technology,
with little work reported on nanocomposites-based sensors. Due to the good electric
properties, CNTs could provide nanocomposites with electrical transduction mechanisms,
like piezoresistivity and capacity detection.
Next sections present the design and fabrication of a novel flexible capacitive sensor
for pressure measurement. The sensor system uses inductive-coupling to deliver energy
and communicate with the capacitive sensor, which is based on two layers of vertically
aligned-carbon nanotubes (A-CNTs) embedded in a flexible substrate of PDMS and a layer
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(with a cavity) made of pure PDMS. Compared with other polymers, PDMS possesses
unequaled mechanical elasticity and has capability to resist to large deformations. The
electrical components, namely the capacitor electrodes and inductor (LC network), are
based on A-CNTs/PDMS nanocomposites.
An external reading system to measure passively the frequency changes from the LC
network, due to a change in the capacitor value, was also developed. The system powers
the sensor and does the transmission of information wirelessly using electromagnetic
waves (the external reading system will be presented in Chapter 5).
4.2 Design of the Flexible Capacitive Pressure Sensor
4.2.1 Sensor Model
The operating method of the flexible pressure sensor is based on the change in deflection
of thin A-CNTs/PDMS nanocomposite membranes due to the change of an external
pressure. The change in capacitance can be determined from the change in deflection
between the two nanocomposite membranes that form the capacitor. When pressure
differences are null, the diaphragm returns to its original position. The flexible capacitive
pressure sensor involves two coupled domains, mechanical and electrical, that define the
sensor’s behaviour.
4.2.1.1 Mechanical Domain
The pressure sensor architecture is similar to typical configurations of capacitive
pressure sensors. It is composed of two square-plate electrodes (diaphragm) separated by a
dielectric (air) of separation distance d0 at pressure P0. Changes of the outside pressure
(Pout) deform the square plates and consequently generate a capacitive change. Figure 4.2
shows a schematic of the flexible capacitive pressure sensor, using two square-plate
diaphragms (sidelength of 2a). A cross section of the square double diaphragm is shown in
Figure 4.3 (section cut B-B from Figure 4.2), considering only the mechanical domain.
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The diaphragm has clamped edges supporting the diaphragm’s perimeter, where 2a
is the sidelength, t is the thickness and w0 the deflection. The deflection in a given point of
the diaphragm is w(x, y), for unbonded diaphragm movement.
Figure 4.2: Schematic of the square (sidelength = 2a) pressure sensor. a) 3D view and b) section
cut B-B.
Figure 4.3: Cross section of the generic model for a deflectable diaphragm.
The angle of deflection (φ) for a clamped diaphragm under a uniform load (such as
pressure), is equal to zero at the center and at the edge of the diaphragm (r = 0 and
r = a, respectively). For these defined boundary conditions, the deflection of an isotropic
square diaphragm under a pressure load can be modeled in closed-form by [4.27]
(considering both bending and stretching effects):
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where υ is the Poisson’s ratio, E is the Young´s modulus and ΔP=P0-Pout is the pressure
load.
Equation 4.2 allows computing the deflection at the center of the diaphragm for a
given pressure load but, the deflection along the diaphragm is still required to model the
capacitive changes due to the gap variation. The complexity of the mechanical deflection
calculation makes difficult to obtain a closed form solution for the diaphragm deflection,
from the center to the edges. For this reason, trial functions [4.28] are usually applied to
describe the deflection of the whole diaphragm.
The trial functions (4.3) and (4.4) have been used to model the diaphragm deflection,
satisfying the boundary conditions.
4.2.1.2 Electrostatic Domain
Pressure changes generated by mechanical deflections will cause, consequently,
changes of the capacitor response (electrostatic domain). A capacitor is defined as being an
electronic component with two electrodes, separated by a dielectric.
The electrostatic domain analysis aims essentially the calculation of the capacitance
values, and the derivation of a model that calculates the capacitance changes due to
changes between the electrodes’ distance (see Figure 4.3).
In the absence of any displacement and for the simple case of a parallel plate
capacitor, the model for the capacitive transducer is (neglecting fringe fields):
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where ε0 is the permittivity of free space (8.8546x10-12
F/m), εr is the relative permittivity,
wc and lc are, respectively, the width and length of the capacitor electrodes (in this case 2a),
and d0 is the gap between the electrodes.
Since the proposed double-plate capacitive sensor uses diaphragm electrodes with a
complex bending profile, integration over the effective area of the electrodes is required to
calculate the total capacitance.
where w(x, y) is the distance between electrodes due to the diaphragm bending at position
x,y. The integration of (4.6) can be solved numerically, allowing the calculation of the
capacitance for a given pressure change.
4.2.2 Telemetry Model
The monitoring principle is based on the detection of pressure variations and transmission
of the measured data without the need for an embedded power supply.
The pressure measurement system comprises a capacitor (C) connected to an
inductor (L) forming a parallel LC resonant circuit, whose oscillation frequency is sensitive
to pressure variations. Passive telemetry is used to externally read the pressure changes
detected by the capacitive pressure sensor. The proposed monitoring system was developed
in order to remotely read the pressure values of the flexible capacitive sensor and it is
composed of three major blocks (Figure 4.4): the capacitive pressure sensor (or sensor
cluster), the external reading system and the signal retrieval and analysis.
The external reading system, which was entirely developed by Cristina et al. [4.29] at
FEUP, delivers energy and detects the resonance frequency of the sensor through an
inductive-coupling link. The reading system includes a primary inductor (Reader) where a
generator of the square wave is applied, a transformer (TF) to remove the common mode
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from the signal acquired by the inductor, amplifying the oscillation frequency. The
instrumentation amplifier (IA) and filter (LPF) allow improving the signal-to-noise ratio
and deliver the signal for optimal A/D conversion. Finally, the signal is acquired and
post-processed [4.29].
Figure 4.4: Blocks diagram of the telemetry system.
The schematic of the reader circuit is shown in Figure 4.5. The pressure sensor
receives energy (the receiver applies a square wave to the sensor) and communicates with
the receiver exchanging information (the oscillation frequency corresponding to the
pressure) by inductive-coupling. To simplify circuit’s analysis, only one sensor is used in
the model. Its components can be modeled as passive elements: a variable capacitor (CS) in
which its value changes with the applied pressure, connected to an inductor (LS), forming a
parallel resonant circuit. The representation in Figure 4.5 includes also resistors (Rp and RS)
whose parasitic resistance values are from the inductors Lp and LS. The oscillation
frequency (4.7) depends on the sensor’s capacitance and inductance, as well as the
coupling factor k (4.8):
and,
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where, rr is the reader’s coil radius, rs is the sensor’s coil radius and x is the distance
between antennas.
Figure 4.5: Schematic of the reader circuit [4.30].
In order to read the sensors response, a square wave (vvs) is sent through the inductor
(Lp) that in turn activates the sensor by generating a magnetic field. Once the sensor is
activated, the output (vo) superimposes the signal sent by the reader (the input) plus the
sensors’ response.
The duplication of the circuit marked as “twin” is necessary for the successful
reconstruction of the sensors’ signal. The use of a square wave to activate the sensors has
the drawback of producing a high number of harmonics in the reader’s output preventing
the recognition of the true signal sent by the sensors. To detect the sensors’ signal (vdif), the
“twin” circuit enables the elimination of the unwanted harmonics (vto) by subtraction
(vo – vto) [4.31].
As soon as the signals are captured, an analog to digital conversion is needed to
further analyze the signals oscillation frequencies in the digital domain, in order to
determine the pressure that is being measured by the pressure sensors. The reader circuit is
connected to an analog to digital converter (ADC) that in turn is linked to a PC. The PC
controls the ADC operation and performs a FFT (Fast Fourier Transform) to determine the
main frequencies of the signals.
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4.2.3 Finite Element Modeling
Finite element modeling (FEM) was used in order to simulate the mechanical domain of
the flexible pressure sensors. The developed FEM model uses a structural physical domain
that computes the bending of the diaphragms due to applied pressure.
A commercial FEM solver, ANSYS software (multiphysics), was used as simulation
tool. The purpose of analysis was to validate the proposed mechanical model by verifying
the deformation of the A-CNTs/PDMS nanocomposite membranes (diaphragm) due to the
applied external pressure.
4.2.3.1 Simulation Setup of Pressure Sensors
A 3D model of the flexible pressure sensors with different dielectric dimensions was
constructed using solid elements. The assessment of the deformation was performed using
three different models, with the same external dimensions but varying the dielectric
sidelength. As the pressure sensor is composed of three membranes, the analyzed model
had a length of 15 mm, a width of 15 mm and a thickness of 200 µm, for all three
membranes. Regarding the middle membrane that composes the pressure sensor (defining
the dielectric due to the hole in the center), the dimensions of the dielectric were changed
and present a total area of 4 x 4 mm2, 6 x 6 mm
2 and 8 x 8 mm
2 (that will be referred as S4,
S6 and S8, respectively). Figure 4.6 presents an example of the models used in the FEM
simulations.
Figure 4.6: Flexible pressure sensor (S6) model for FEM simulations.
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The properties of both A-CNTs/PDMS nanocomposite and pure PDMS membranes
used for the simulation were obtained from the mechanical characterization performed and
presented in chapter 3 and are presented in Table 4.1.
Table 4.1: Materials’ properties of the pressure sensors used in FEM simulations.
Property Description Value
PDMS A-CNTs/PDMS
Ex Young’s modulus x direction
0.8 MPa
1.7 MPa
Ey Young’s modulus y direction 8.1 MPa
Ez Young’s modulus z direction 1.7 MPa
υxy
Poisson’s ratio 0.4
0.05
υyz 0.24
υxz 0.26
Gyz Shear modulus 0.25 MPa 0.24 MPa
The selected mesh control for the models was element size type with 0.25 mm along
the length, width and 0.2 mm along the thickness. The boundary condition for the
rectangular edges of the models was fixed in all directions, causing a null displacement in
the x-, y- and z-directions. Atmospheric pressure (P0) was considered inside the dielectric
and a pressure (Pout) was applied onto the square faces of the model, as shown in Figure
4.7.
Figure 4.7: Model showing the mesh type, boundary conditions and applied pressure.
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4.2.3.2 Simulation Results
The simulation parameters used for all models were the same, varying only the
dimensions of the dielectric. Atmospheric pressure of 101 kPa was assumed in the
dielectric cavity and a pressure of 100 kPa was applied onto external surfaces of the
pressure sensor. When Pout is applied, the air retained inside the cavity pushes the
A-CNTs/PDMS nanocomposite membranes, deforming the sensor as shown in Figure 4.8.
Figure 4.9 presents a comparison between underwent displacements for different dielectric
dimensions after application of the pressure Pout.
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Figure 4.8: Diaphragm deformation in the y-axis. a) dielectric size of 4x4 mm2, b) dielectric size of
6x6 mm2 and c) dielectric size of 8x8 mm
2.
Figure 4.9: Displacement of pressure sensors with different dielectric dimension after applying the
pressure Pout.
Analysis of the FEM results show that an important characteristic of the sensor
behaviour is the sensitivity dependence on the applied outside pressure (Pout), where small
changes of pressure (ΔP) can cause large difference in the displacement, depending of the
dielectric size. If the last bonding step of the flexible sensor is performed in a pressure
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controlled environment, P0 can be controlled which allows the realization of sensors with
different sensitivities.
When 100 kPa is applied onto the external surfaces of the flexible pressure sensor
their electrodes deform, corresponding to a maximum displacement of 0.259 mm,
0.533 mm and 0.757 mm, for the sensor with a dielectric area of 4x4 mm2, 6x6 mm
2 and
8x8 mm2, respectively.
Since the proposed analytical mechanical model considers only isotropic materials,
FEM simulations using isotropic properties were also performed, in order to validate the
maximum displacement given by equation (4.2) as well as the displacement along the
sensors’ diaphragm, given by the trial functions (4.3) and (4.4).
Table 4.2 shows the values of the isotropic properties used in the new FEM
simulations. Figure 4.10 gives a comparison of the displacements obtained for both
isotropic and orthotropic material properties, as well as a comparison of the FEM
simulations with the analytical model.
Table 4.2: Isotropic material’s properties of the pressure sensors used in FEM simulations.
Property Description Value
E Young´s modulus 1.7 MPa
υ Poisson’s ratio 0.05
G Shear modulus 0.24 MPa
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Figure 4.10: Displacement of pressure sensors for both orthotropic and isotropic material
properties. Sensors with: a) dielectric size of 4x4 mm2, b) dielectric size of 6x6 mm
2 and c)
dielectric size of 8x8 mm2.
The results presented in Figure 4.10 show that the displacement curves of the FEM
simulations present similar behaviour in which the small differences regarding the
maximum displacement are due to the material’s properties used in the model.
In this case, where isotropic properties are used, applying an outside pressure of
100 kPa onto the external surfaces of the pressure sensor, their electrodes present a
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maximum displacement of 0.266 mm, 0.542 mm and 0.776 mm, for the sensor with a
dielectric area of 4x4 mm2, 6x6 mm
2 and 8x8 mm
2, respectively.
The analytical models show a behaviour similar to the FEM simulations. The small
differences between the two models can be due to the fact that the analytical model
considers only the deformation in the dielectric region, while the FEM models consider the
total length of the pressure sensor. Nevertheless, FEM simulations for both isotropic and
orthotropic materials compare relatively well with the analytical models, mainly for higher
dielectric areas. Also, it is clear that the property values of the materials are an important
factor to obtain a more realistic approach of the total displacement of the flexible pressure
sensor, mainly when the sensors have larger dielectric dimensions. Thus, isotropic material
properties are a good approximation for modeling the pressure sensors, validating the
proposed analytical mechanical model.
4.4 Development of Flexible Pressure Sensor
This section describes the fabrication process developed to build the flexible pressure
sensors. The final sensor comprises three thin membranes made of A-CNTs/PDMS
nanocomposites and pure PDMS, bonded by a thin layer of uncured PDMS.
4.4.1 Fabrication Process
The proposed fabrication process to develop the flexible pressure sensor uses vertically
aligned carbon nanotubes (A-CNTs) to build the conductive elements, namely the
capacitor electrodes, embedded in a flexible substrate of PDMS.
The schematic of the fabrication process to obtain a single nanocomposite membrane
was already presented in Chapter 3 (in the §3.3, Figure 3.4). For the pure PDMS
membranes new acrylic molds were produced, in order to obtain a PDMS membrane with
a cavity (dielectric) in the center. Figure 4.11 shows the mold fabricated and used to build
the pure PDMS membranes.
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Figure 4.11: Fabricated acrylic molds to build the pure PDMS membranes (dimensions in mm).
Posterior to the characterization of both A-CNTs/PDMS nanocomposite and pure
PDMS materials, the flexible capacitive sensors were assembled for sensor
characterization.
The pressure sensor is composed of three thin membranes: two membranes (top and
bottom) of nanocomposites defining the electrodes (each membrane has ~ 1 % CNTs
volume fraction) and a membrane of pure PDMS (middle layer containing a cavity) to
separate the electrodes and to define the capacitor dielectric. The final configuration of the
flexible pressure sensor is schematically represented in Figure 4.12.
Figure 4.12: Schematic of the flexible pressure sensor composed of three thin membranes (top and
bottom layers are A-CNTs/PDMS nanocomposites and the middle one is made of pure PDMS).
In order to achieve the final configuration a bonding process is required.
Eddings et al. [4.32] tested five different bonding techniques and the highest reported bond
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strength was obtained for both partial curing and uncured PDMS adhesive techniques. The
latter approach was used with proven success in the fabrication of the flexible pressure
sensors. Figure 4.13 shows both digital and microscopic images of an A-CNTs/PDMS
assembled pressure sensor, with a dielectric area of 3.9 x 9 mm2.
Figure 4.13: Flexible pressure sensor. a) Digital image taken into the vacuum chamber and
b) optical image of the cross section (magnification: 0.67 x 6.7).
Several sensors with different dielectric sizes were assembled for testing. Figure 4.14
shows a device with 8 x 8 mm2 electrode size. The proposed technology proved to be
effective for the fabrication of the sensors, but still requires improvements for the inductors
fabrication (to enable a LC network for passive telemetry).
Figure 4.14: Digital image of fabricated pressure sensors showing its flexibility.
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4.5 Sensor Characterization and Results
4.5.1 Finite Element Model
The challenge of using finite element modeling to get the sensor total behaviour is the
simulation of coupled domains (mechanical and electrostatic) in order to first compute the
bending of the diaphragms due to applied pressure, followed by the calculation of the
capacitance.
In order to determine the capacitance changes for a given applied pressure on the
sensor, a fully parameterized model was developed using an APDL (ANSYS Parametric
Design Language) script. These simulations are normally performed in three steps:
Definition of the model parameters (geometry, material properties, boundary
conditions, applied loads, etc.), carried out using the preprocessor module;
Definition of the analysis type and method of calculation;
Review the results in the postprocessor module.
Figure 4.15 shows the model parameters used to calculate the capacitive changes of
various sensors. The dimensions and properties values used in the model are presented in
the Table 4.3.
Figure 4.15: Finite element model parameters.
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Table 4.3: Dimensions and material properties of the pressure sensors used in the FEM model.
Parameter Description Value
a Radius 4 mm
d0 Distance between electrodes 200 µm
t Diaphragm thickness 200 µm
P0 Pressure of the sensor cavity 101 kPa
Pout Applied pressure several
E Young’s modulus 1.7 MPa
υ Poisson’s ratio 0.26
εr Relative permittivity 1
ε0 Permittivity of the free space 8.854x10-12
F/m
ANSYS enables the use of several physical domains, where results from one physical
domain can be applied as a load on a different physical domain. The developed model uses
initially a structural physical domain that computes the bending due to applied pressure.
Then, the command dvmorph is issued and the non-structural areas are re-meshed (taking
into account the displacement of the diaphragm due to pressure applied). Finally, an
electrostatic physical domain is used to calculate the capacitance. An overview of the used
FEM algorithm is shown in Figure 4.16.
An example of a meshed model using the dimensions and material properties listed
in Table 4.3 is presented in Figure 4.17. The bending results are presented in Figure 4.18.
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Figure 4.16: Finite element modeling algorithm for coupled domain analysis.
Figure 4.17: Example of the obtained mesh model.
Structural Model
Structural model with all physical
requirements
(physics, write, STRUCTURE)
Electrostatic Model
Electrostatic model with all physical
requirements
(physics, write, ELECTRO)
Read in first physics
environment
(physics, read, STRUCTURE)
Solve for first physics environment
(/SOLU)
Re-mesh non-structural volumes
(DVMORPH)
Read in second physics
environment
(physics, read, ELECTRO)
Solve for second physics environment
(/SOLU)
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Figure 4.18: Bending profile for the applied pressure Pout (100kPa).
4.5.2 Capacity and Pressure Changes Measurement
Fabricated A-CNTs/PDMS prototype flexible pressure sensors were tested in a vacuum
chamber. The tested sensors have a mechanical thickness of 675 µm (425 µm of A-
CNTs/PDMS nanocomposite plus 250 µm of pure PDMS, used for uncured PDMS
bonding) and a dielectric air thickness of 250 µm (in fact the total dielectric gap is 750 µm,
resulting from the pure PDMS layers used for bonding) which result in a sensor with a total
thickness of 1.6 mm (675 µm + 250 µm + 675 µm). The area of the electrodes is 8 x 8 mm2
(W x L). In spite of the low sensor resolution, since the geometry was not optimized (the
devices were at a proof-of-concept stage), the sensors were mounted inside a controlled
pressure chamber (the dielectric is hermetically sealed at ambient pressure) in order to
measure the capacitive changes (a LCR meter was used to measure the capacity changes).
The results are presented in Figure 4.19.
FEM was used to build an electromechanical model of the pressure sensor and
simulation results were compared with the experimental measured data and analytical
model of the capacitive sensor. In both models, the nanocomposite membranes were
modeled as an orthotropic material with the values retrieved from the mechanical tests.
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Figure 4.19: Capacitive changes vs. pressure changes, assuming ambient pressure inside the
dielectric.
Analysis of the results show that both analytical and FEM models compare relatively
well, although is clear that the trial function is an important parameter for the model. The
experimental results are in reasonable agreement for high pressure differences (in terms of
rate of change), but the differences are significant for pressures around the ambient
pressure. Several reasons may contribute to these differences, such as geometry
asymmetries created during the fabrication process (most of the process is done manually
and therefore prone to dimensional variations from the designed ones), and uncertainty
regarding the actual pressure inside the hermetically sealed cavity. In fact, since small
pressure changes in the outside (Pout) cause large diaphragm deformations and large
deformations cause volume changes of the dielectric material (in this case the dielectric is
air), the pressure (P0) inside the dielectric will change (from the ideal gas law
).
In order to capture this behaviour, the existing analytical model was extended and an
iterative model using the flowchart presented in Figure 4.20 was implemented in
MATLAB.
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Pout
Maximum deflection = f (P0-Pout)
Volume = f (Maximum deflection)
Pressure = f (Volume)
ΔC = C0-C (Maximum deflection)
Yes
No
Increase P0 Decrease P0
Yes
No
P0
Figure 4.20: Iterative algorithm to compute the capacitive changes of flexible pressure sensors.
A new group of fabricated pressure sensors, with different dimensions, was again
used for sensor characterization (these sensors did not include the inductor). In order to
measure the capacitive changes, a pressure chamber existent at Department of Polymer
Engineering (DEP) facilities was adapted for this type of experiments. Figure 4.21 shows a
3D drawing of the pressure chamber personalized for the capacitive changes measurement.
During the tests, the sensors were placed inside the pressure chamber and a LCR
meter (Philips PM6304) was used to measure the capacitive changes. A reference pressure
sensor (Keller PAA 21R) was also used for data comparison. LCR meter (10 fF resolution)
and reference pressure sensor outputs where acquired at a sampling frequency of 1.6 Hz
and the data was stored in a PC for later comparison.
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Figure 4.21: 3D drawing of the pressure chamber personalized for the capacitive changes
measurement of the flexible sensors.
Several sensors with different diaphragm geometries and thicknesses were tested and
static responses of three devices were compared with the sensor analytical model. Figure
4.22 shows the measured sensors capacitance versus pressure transfer characteristic, for
pressures ranging from 0 kPa to 100 kPa.
From the results it can be seen that the pressure sensors characteristic fits very well
that predicted by the analytical model and that it shows relatively good linearity,
specifically in the region close to the sensor’s reference (atmospheric pressure).
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Figure 4.22: Experimental results and comparison with analytical model. a) Area of the dielectric
is 4 x 4 mm2 and the distance between electrodes is 320 µm, b) area of the dielectric is 5 x 5 mm
2
and the distance between electrodes is 130 µm and c) area of the dielectric is 8 x 8 mm2 and the
distance between electrodes is 210 µm.
Response of the sensors was also measured for both decreasing and increasing
pressure steps (within the 0 – 100 kPa range) for the sensor with a dielectric area of
5 x 5 mm2 and a distance between electrodes of 130 µm. Dynamic pressure sensing cover
applications where there is an interest in monitoring changes in pressure over small time
intervals. A sample result is presented in Figure 4.23.
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Figure 4.23: Dynamic response of the flexible pressure sensor.
Results show a very long time response for increasing pressure steps. In this case, the
tensioned nanocomposites are returning to the rest position, and due to the viscoelastic
behaviour of the nanocomposites, strain takes additional time to disappear which explains
the observed step curves.
The agreement between the modeling and experimental results is very good and
validate the presented modeling approach. The large deformation of the flexible
membranes used for the pressure sensor fabrication cause large volume changes within the
dielectric cavity, which need to be included in the model.
Both the model and experimental results show a linear response of the capacitive
sensor over a large pressure range. This behaviour is not observed in traditional silicon
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pressure sensors (the deformations are very small in this case) and can be explored for the
realization of pressure sensors with improved linear response and extended dynamic range.
Sensitivities ranging from 2.5 to 20 fF/kPa were obtained for different electrode
dimensions and the dynamic response of the devices show a dual behaviour, i.e., the
behaviour response time of the sensor during a pressure decreasing step is lower than for a
pressure increasing step. This is explained by the viscoelastic nature of the PDMS-based
nanocomposites used for fabrication of the capacitive sensor electrodes. These results also
validate the novel technology used for the fabrication on thin, flexible capacitive pressure
sensors.
4.6 Summary
A technology based on A-CNTs embedded in a flexible matrix of PDMS was successfully
applied in the development of flexible pressure sensors. The development of a flexible
remote pressure measurement system based on inductive-coupling, including its design,
fabrication of the sensor and performance characterization of the sensor was presented in
this chapter.
FEM simulations performed using both isotropic and orthotropic materials properties
compare relatively well, suggesting that isotropic material characteristics are a good
approach for modeling the capacitive pressure sensors and thus, it validates the proposed
mechanical model.
Flexible capacitive pressure sensors fabricated with PDMS-based nanocomposites
were experimentally characterized and results compared with simulations from both
analytical and FEM models. The high flexibility of the membranes for small pressure
difference must be accounted for and the initial model was improved to capture this
behaviour. Static responses of the tested flexible pressure sensors revealed reasonably good
linearity in the range of 0 – 100 kPa, mainly in the region near to the atmospheric pressure.
On the other hand, dynamic response of pressure sensors measured in the same range
presented two distinct comportments, justified by the viscoelastic behaviour of the
PDMS-based nanocomposites used in the pressure sensor. In addition, for the tested
pressure sensors (with different electrode dimensions), sensitivities in the range of
2.5 – 20 fF/kPa were obtained.
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In summary, fabricated pressure sensors validate the newly developed technology
and showed that A-CNTs/PDMS nanocomposites can be used to create highly flexible
pressure sensors. The presented flexible pressure measurement system demonstrates that
with proper technological developments, novel interesting applications might follow with
this new technology and, possibly, to extend the principle to other fields of difficult local
readout.
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Chapter 5
Biomedical Applications: Case
Study – Abdominal Aortic
Aneurysms
The use of flexible sensors technology in biomedical applications is presented in this
chapter. Sensors are developed and characterized in the context of abdominal aortic
aneurysms (AAAs), supporting the importance of wireless flexible pressure sensors in the
treatment and diagnosis of patients with this type of disease.
First, an introduction and background on AAAs, including their causes and treatment
options is given. The existing implantable pressure sensors, including those currently used
in the treatment of AAAs, as well as the proposal of a new smart stent-graft, is presented in
§5.3. Section §5.4 introduces the development of a flexible pressure measurement system
based on inductive-coupling for monitoring of post–EVAR procedure. Finally, §5.5
presents the experimental results of the flexible pressure measurement system’s
performance, evaluated through measurement setups to obtain the sensors characteristic
curve and emulate the end use of the system.
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5.1 Introduction
During the past forty years, significant advances have been made in the prevention,
diagnosis and treatment of cardiovascular diseases. According to the World Health
Organization (WHO), 17.5 million people died due to cardiovascular diseases in 2005,
representing
30 % of all global deaths. It is estimated that in 2015, 20 million of people will die from
this type of diseases, mainly from heart attacks and strokes [5.1].
In spite of the great progress at all technological levels, the morbidity and mortality
from cardiovascular diseases continue to represent a high burden [5.2]. Nowadays, the
medical device industry is well characterized regarding the innovation and development of
new devices. These continuous advances are expected to improve patients’ life expectancy
and quality of life [5.3].
This chapter initially presents an overview of abdominal aortic aneurysms and the
main available treatment options for this serious disease. Next, the state-of-the-art on
implantable pressure sensors, and the development and experimental characterization of a
pressure monitoring system based on flexible sensors are presented.
5.2 Abdominal Aortic Aneurysm
An abdominal aortic aneurysm (AAA) is a silent and hard to detect but very dangerous
arterial problem. An aneurysm can be defined as a bulging or ballooning of an artery
fraction caused by the weakening of its walls, having at least a 50 % increase in diameter
compared with the healthy one [5.4], as shown in Figure 5.1.
It is estimated that 12 per 100 000 persons-year [5.5] are affected by this serious
condition, values that are expected to increase with the rise in life’s expectancy. An
aneurysm can appear in several regions, but they most frequently occur in the aorta and
currently is one of the most common causes of death in the western world, particularly
among men aged over 65 [5.6]. If left untreated, AAAs may burst or rupture causing shock
and even death, due to massive blood loss [5.7]. However, when timely detected, most
AAAs can be repaired to prevent rupture.
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Figure 5.1: Illustrations of: a) normal aorta, b) thoracic aortic aneurysm (TAA) and c) abdominal
aortic aneurysm (AAA).
The size of AAAs ranges from < 4 cm in diameter to > 8 cm and the risk of rupture
increases as larger is the aneurysm’s diameter. Table 5.1 gives the estimated risk of rupture
per year of AAAs according to its dimensions. However, not only the diameter of AAAs
represents a risk factor. There are many other factors found to influence the risk of rupture
of AAAs, such as atherosclerosis (hardening of the arteries), hypertension, smoking,
family history, gender, etc.
Table 5.1: Estimated annual risk of rupture of AAAs for a given diameter [5.8].
AAA diameter (cm) Risk of rupture (% per year)
< 4 0
4 – 4.9 0.5 to 5
5 – 5.9 3 to 15
6 – 6.9 10 to 20
7 – 8 20 to 40
> 8 30 to 50
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All of the risk factors mentioned above play effects through localized inflammatory
changes, resulting in degradation of extracellular matrix and apoptosis (process of
programmed cell death – PCD) of vascular smooth muscle cells. These changes are called
medial degeneration of the aortic wall and its precise pathogenesis is not fully understood
[5.9]. Figure 5.2 presents a diagram showing the different stages of AAA degeneration, in
agreement with Thompson et al. [5.10].
Figure 5.2: Different stages of AAA degeneration. Used with permission and adapted from
reference [5.10] (Thompson RW et al. Current Problems in Surgery 2002; 39: 110-230,
Copyright (2002) Elsevier).
5.2.1 Current Available Treatment Options
Nowadays, there are two effective methods to treat AAAs: the reconstruction of the
aorta – open surgery [5.11] and the endovascular aneurysms repair – EVAR [5.12].
Since the early 1950’s, aneurysms’ standard treatment has consisted of an open
surgery (done under general anesthesia) and the replacement of damaged section of the
aorta by a synthetic prosthetic graft (Figure 5.3) through a surgical procedure (a large
incision in the abdomen) [5.11]. This method is often traumatic for the patient (with, e.g.,
cardiac and pulmonary side effects) and presents, between others, prolonged
hospitalization and long time to recover [5.13]. Furthermore, although being effective it is
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associated with a significant (4% to 5%) morbidity and mortality rate [5.14]. However, in
spite of its invasiveness and the fact of being limited to fit patients, this treatment is still a
current practice and less invasive techniques, namely total laparoscopy and assisted
laparoscopy, are being studied to minimize its disadvantages [5.15].
Figure 5.3: Open surgery technique for AAAs to insert a synthetic graft.
In the beginning of the 1990’s, Juan Parodi and colleagues executed for the first time
an endovascular aneurysm repair in Argentina, and demonstrated that the technique was a
safe and feasible practice [5.15, 5.16]. This surgical procedure is done percutaneously and
it is a minimally invasive procedure (Figure 5.4), where the goal is to prevent aneurysm
rupture by exclusion of the aneurysm sac from systemic pressure.
Figure 5.4: Stent-graft deployment sequence in the endovascular aneurysm repair (EVAR)
procedure.
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Typically, a small incision is made in each groin to expose the femoral arteries.
Then, with the aid of catheters and guide-wires, a stent-graft is guided from the femoral
artery to the affected segment, allowing blood to pass without exerting pressure in the
aneurysm sac and, thus, preventing progressive wall expansion and rupture [5.17, 5.18].
Since EVAR was proposed by [5.12] it has become widely accepted due to its main
advantages, such as decreased blood loss, less physiological derangement, lower morbidity
and mortality and more rapid recovery than that required for the open surgery [5.20].
Comparing both procedures, EVAR is preferable due to the fact of being less stressful and
reducing significantly systemic complications [5.21], as well as having lower costs
associated to the hospital stays and less or no need for intensive care facilities during
recovery.
The durability of open surgery, established with long-term follow-up studies, is
excellent [5.21], and there is little or no requirement for long-term surveillance, in contrast
with EVAR whose current results suggest that there is a need for increased surveillance
and reintervention [5.21, 5.22]. Considering the longer life expectancies and the rising
public expectations for quality of life, the costs associated to the follow-up can jeopardize
EVAR’s effectiveness. Figure 5.5 gives a comparison of the approximate costs per patient
for both open surgery and EVAR techniques.
Figure 5.5: Open repair and EVAR estimated costs per patient (Adapted from [5.24]).
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5.2.2 Actual Stent-Grafts Main Problems and Needs for Pressure Monitoring of
AAAs
When EVAR was introduced, it revolutionized the treatment of aortic aneurysms.
However, and in spite of major advances in EVAR techniques, adverse reactions still occur
[5.25] and lifelong surveillance is recommended [5.26]. Due to these complications,
currently, questions are being raised regarding the follow-up costs [5.27] and alternative
approaches, such as a smart stent-graft, are being pursued.
Stent-grafts are frequently used for the treatment of AAAs. Both in Europe and in the
USA, stent-grafts are classified as class III medical devices. They are composed of a
metallic scaffold with a polymeric covering membrane. While early devices were custom
designed for each patient by the operating surgeon assembling off-the-shelf components,
nowadays several commercially devices are available. However, despite improvements in
stent-graft technologies over the last decade, there are some post-surgery problems that
still occur, such as graft migration, stent fracture, enlargement of the aneurysm sac and,
one of the most frequent complications is the occurrence of endoleaks. Table 5.2 presents
some of EVAR’s problems involving stent-grafts.
Table 5.2: Complications involving stent-grafts (Adapted from [5.25]).
Early complications Late complications
Graft kink Graft migration
Endoleaks Neck dilatation
Graft explantation Endoleaks
Structural failure Structural failure: component separation, fabric
tears, hook fractures Graft infection
Endoleak, or leakage, is defined as a persistent blood flow into the residual aneurysm
sac after EVAR [5.28]. Golzarian and Valenti classified the possible causes of endoleaks
into five types [5.29]:
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Type I – Endoleaks occur due to malpositioning of the stent-graft, type of
endograft, migration from the desired position, and stent-grafts’ design and
dimensions;
Type II – Endoleaks are related with aneurismal expansion and rupture;
Type III – Endoleaks occur due to structural failure of the stent-graft;
Type IV – Endoleaks through the porosity of the graft material;
Type V – Endoleaks are related to the continued aneurysm expansion, which can
happen when they are not diagnosed.
Due to anomalies that can occur after the stent-graft placement, namely endoleaks,
EVAR requires an expensive surveillance protocol using ultrasound and computer
tomography angiography (CTA) at 1, 6, and 12 months after the procedure, and thereafter
on an annual basis [5.30], magnetic resonance imaging (MRI) and magnetic resonance
angiography (MRA) scans. Ultrasonography, besides allowing measuring the aneurysm
sac, is effective in the detection of endoleaks but, even with enhanced sensitivity obtained
with the use of contrast agents, requires a skilled technician to interpret the exams. Also,
the most currently used methods expose the patients to carcinogenic risks or are
incompatible with some stent-grafts, and are prone to error [5.31]. Furthermore, these
imaging protocols only give information about diameter and volume changes and presence
of contrast inside the aneurysm sac, and have a low sensitivity [5.30]. In addition, these
exams are considered time consuming, expensive and do not provide measures of the
pressure inside the aneurysm sac, valuable information that can reveal the presence of
low-flow endoleaks [5.32].
In order to reduce and even eliminate these exams, new approaches based on the
monitoring of the aneurysm intra-sac pressure have been studied, and some solutions
proposed in the literature [5.33, 5.34]. Indeed, the pressure in the walls and in the
aneurysm sac is one of the most important indicators for supervising post-EVAR patients
and, direct intrasac pressure measurement outperforms imaging techniques in ascertaining
either the complete AAA exclusion or the need for future interventions [5.35].
Remote pressure monitoring systems [5.28] and results of clinical observation using
these systems have been published [5.36], showing that the use of wirelessly accessed
pressure sensors provide useful post-EVAR endovascular leaks observation and help
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guiding clinical therapy. Nevertheless, its use is still not a standard procedure because it
currently requires invasive methods.
The first wireless monitoring systems relied on a pressure sensor placed between the
stent-graft and aneurysm walls, after the placement of the stent-graft. With the systems
presented in [5.34, 5.37] the pressure sensor makes part of the stent itself, benefiting from
advancements on stent’s technology and their delivery and implementation procedures.
5.3 Measurement of Aneurysm Sac Pressure
None of the medical imaging exams presently used provides measures of the pressure
inside the aneurysm sac. This information is important because can be evidence of
endoleaks or endotension [5.32].
Published data describe the use of catheters to measure pressure in the residual
aneurysm sac [5.38]. However, although these methods provide precise measurements
[5.26], they are invasive and bear multiple risks.
An alternative method for the measurement of the aneurysm sac’s pressure is the
implant of remote pressure transducers during EVAR. This solution is advantageous since
measurements can be done when needed (hourly, weekly, etc.) in the patient’s home or
office instead of a hospital once or twice a year. Another important feature is the
possibility to measure both the mean pressure and the pulsatile pressure without increasing
risks for the patients.
In the next section, only the more recent research developments found in the
literature regarding blood pressure measurement are presented. Therefore, review on
long-term implantable pressure sensors will be divided into extra-arterial and intro-arterial
blood pressure devices, where the four telemetric pressure sensors proposed for EVAR
monitoring are described.
5.3.1 Extra-Arterial Blood Pressure Sensors
Extra-arterial blood pressure sensors are placed around the blood vessel and perform an
indirect pressure measurement through the wall or through the expansion and contraction
of the artery. Nevertheless, they require an invasive surgical procedure for their implant.
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In 2001, Ziaie and Najafi [5.39] proposed a tonometric blood pressure sensor with
improved stability, to measure the blood pressure in small animals. The device, based on a
miniature titanium cuff, surrounds and reshapes the blood vessel in order to improve the
contact between the pressure sensor and the arterial wall (Figure 5.6). With integrated
electronics for the capacitive readout, an assembled unit measures 10 x 6.5 x 3 mm3. The
device was tested using an elastomeric tube mimicking a blood vessel, which was
pressurized with air in a range of 0 – 200 mmHg. At 100 mmHg, in vitro tests of the cuff
system presented a sensitivity of 2.0 mV/mmHg and a resolution of 0.5 mmHg
(Figure 5.7).
Figure 5.6: Illustration of the basic principle of tonometric blood pressure measurement. Used with
permission from reference [5.39] (Ziaie B et al. Biomed Microdevices 2001; 3: 285-292,
Copyright (2001) Springer).
Figure 5.7: Photograph and sensor calibration curve of the tonometric pressure sensor. Used with
permission from reference [5.39] (Ziaie B et al. Biomed Microdevices 2001; 3: 285-292,
Copyright (2001) Springer).
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A new elastic cuff made of soft biocompatible rubber was proposed in 2004 by
Cong et al. [5.40]. In the presented monitoring system, a MEMS pressure sensor is
immersed in a biocompatible insulating fluid and placed inside the elastic cuff (Figure 5.8).
The MEMS sensor embedded in the cuff measures the vessel blood pressure waveform,
which corresponds to a scaled version of the blood pressure in the artery, independent of
the cuff bias pressure exerting on the vessel. In order to assess the feasibility of the
approach, a scaled up prototype system was developed to validate the concept.
Experimental results demonstrated that the pressure in a tube simulating a vessel can be
measured by a MEMS pressure sensor, proving the viability of the concept.
Figure 5.8: Implantable blood pressure sensor monitoring system and cross-section of the cuff
around the vessel. Used with permission from reference [5.40] (Cong P et al. Proceedings of IEEE
Sensors 2004; 3: 1359-1362, Copyright (2004) IEEE).
In 2006, the authors improved the sensor design [5.41] by introducing a rigid
isolation ring on the outside wall of the cuff (Figure 5.9). The rigid ring isolates the cuff
from environmental variations to suppress the baseline drift in the measured blood
pressure.
The sensor was experimented in vivo and compared to a commercial catheter-tip
transducer. The measured baseline drift of 0.6 mmHg was three times smaller in the new
configuration, when compared to the previous design. In addition, the authors concluded
that the proposed cuff-based sensing design is suitable for long-term implant applications
and can be certainly useful for human implant monitoring.
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Figure 5.9: Cross-section of new monitoring cuff and prototype of implantable blood pressure cuff
with rigid ring. Used with permission from reference [5.41] (Cong P et al. Conference of the IEEE
Engineering in Medicine and Biology Society 2006; 1: 1854-1857, Copyright (2006) IEEE).
Shin et al. [5.42] developed an implantable pressure sensor module mounted on a
flexible substrate using chip embedded flexible packaging (CEFP) technology, as shown in
Figure 5.10.
Figure 5.10: Schematic and photograph of flexible CEFP platform. Used with permission from
reference [5.42] (Shin K-Ho et al. Sensors and Actuators A: Physical 2005; 123-124: 30-35,
Copyright (2005) Elsevier).
The system is mechanically flexible, and can be attached to the outside of the blood
vessel to remotely monitor the blood pressure (Figure 5.11). Data from the device is
transmitted wirelessly, through an inductive coupling. Two gold electrodes are deposited
on the flexible substrate and placed around the blood vessel. The two electrodes can be
seen as a variable capacitor that is connected to an inductor, implementing a LC resonator.
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The device was tested on silicone rubber tubing, mimicking a blood vessel, and
experimental data indicates a sensitivity of 11.25 kHz/kPa.
Figure 5.11: Schematic and photograph of flexible blood pressure sensor module. Used with
permission from reference [5.42] (Shin K-Ho et al. Sensors and Actuators A: Physical 2005; 123-
124: 30-35, Copyright (2005) Elsevier).
5.3.2 Intro-Arterial Blood Pressure Sensors: Aneurysms
Intro-arterial blood pressure sensors refer to sensors that are in contact with the blood
stream, inside of the blood vessels. A few devices have been proposed for pressure
monitoring [5.43-5.46], but since the focus in this chapter are aneurysm pressure sensors,
only devices that measure the pressure within the aneurysm sac will be presented.
Impressure AAA Sac Pressure Transducer
In 2003, the Impressure AAA Sac Pressure Transducer or RemonAAA from Israel Remon
Medical Technologies (Figure 5.12) was the first permanently implantable,
ultrasound-activated remote pressure transducer to measure intrasac pressure after
endovascular aneurysm repair (EVAR) [5.47, 5.48].
The transducer, hand-sewn to the outside of a stent-graft, contains a piezoelectric
membrane that energizes a capacitor when actuated by ultrasound waves from a hand-held
probe. Once charged, the aneurysm sac pressure is measured followed by the generation of
an acoustic signal that is relayed to the hand-held probe. The probe then converts the
acoustic signal to a pressure waveform that is presented on a computer screen. In spite of
ultrasound being safe and widely used for medical imaging, the measurement requires the
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use of an ultrasonic gel and direct contact between the skin and the transducer. Another
drawback of this sensor is the impossibility of ultrasound to travel through air or bone,
which may lead to difficulties communicating with the aneurysm sac.
Figure 5.12: Impressure AAA sac pressure transducer (Remon Medical). Used with permission
from reference [5.48] (Ellozy SH et al. Journal of vascular surgery 2006; 43: 2-7,
Copyright (2006) Elsevier).
The Impressure sensor is small (3 x 9 x 1.5 mm3) and it is not very radiopaque, but
still visible. As the sensor is sewn to the stent-graft, the fixation location must be carefully
chosen in such a way that the sensor will measure the pressure inside the excluded
aneurysm sac without being pushed against the aneurysm wall. In October 2006, Remon
Medical Technologies announced the first European implant of the sensor.
EndoSure Wireless Pressure Sensor
The EndoSure Wireless Pressure Sensor (CardioMEMS, Inc., USA), Figure 5.13, is made
of two coils of copper wire within a fused silica matrix with a pressure sensitive surface.
Passive telemetry is used for the signal transfer between the external device and the
implant. Through inductive coupling, changes on the internal LC network (capacitor plus
inductor) resonance frequency are detected on the external coil.
The resonance frequency is related to the ambient pressure in which the sensor is
located, and specially-designed software transforms the frequency shift between systolic
and diastolic pressures into a wave form and pressure reading [5.49, 5.50]. As the
measurements are acquired via radio frequency (RF), there is no need for contact with the
patient’s skin or even the removal of the clothes, potentially allowing daily/weekly
sampling at home.
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Figure 5.13: EndoSure wireless pressure sensor (CardioMEMS, Inc.) [5.51].
The EndoSure sensor is deployed through its own delivery catheter (diameter
14 Fr/4.7 mm) during the EVAR procedure and has radiopaque markers to clearly define
its location within the aneurysm sac. Also, the sensor is cleared by the FDA for the
measuring of intrasac pressure during endovascular abdominal and thoracic aneurysm
repair.
A similar device from the same team, was presented by Fonseca et al. in [5.52]. The
authors reported the first two types of flexible micromachined wireless pressure sensors,
for both acute and chronic use. The device for acute use (Figure 5.14) is fabricated with
laminated sheets of cooper-clad liquid crystal polymer (LCP) and expanded
polytetrafluoroethylene (PTFE)-based inner bonding layers, while the sensor designed for
chronic use (Figure 5.15) is fabricated with both cooper-clad and non-clad PTFE layers,
fluorinated ethylene propylene (FEP) copolymer and an encapsulated ceramic chamber.
Both devices are flexible and can be folded into shapes capable for delivery in the body
using a catheter.
Figure 5.14: Flexible wireless pressure sensor for acute use [5.52].
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Figure 5.15: Flexible wireless pressure sensor for chronic use [5.52].
The device was successfully tested in canines over a 60-day period and a second
version of devices is expected to improve the sensor drift. The sensor is targeted for
abdominal aortic aneurysms (AAAs) monitoring and the development has clearly taken
into consideration the biocompatibility and a minimally invasive delivery.
TPS Telemetric Pressure Sensor
The TPS Telemetric Pressure Sensor (Figure 5.16) was developed by the
Helmholtz-Institute for Biomedical Engineering, RWTH Aachen in cooperation with the
Institute of Materials in Electrical Engineering, RWTH Aachen [5.52, 5.53].
Figure 5.16: TPS telemetric pressure sensor: system concept and readout station. Used with
permission from reference [5.54] (Springer F et al. European radiology 2007; 17: 2589-2597,
Copyright (2007) Springer).
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The TPS sensor consists of an implantable sensor capsule, with a radiopaque foil,
and an external readout station. The capsule, in turn, includes a capacitive absolute
pressure sensor and an in-capsule signal-processing microchip with an inductive telemetry
unit. As the measured data is completely pre-processed to a digital data stream in the
implant, possible errors during the transmission or caused from interferences between the
sensor and the external readout station can be diminished. Also, errors induced by
temperature changes can be registered and numerically corrected by an additional
integrated temperature sensor. Taking into account the requirements of medical implants,
the pressure sensor capsule comprises a biocompatible capsule-shaped silicon form and a
Parylene-C layer. This encapsulation system allows creating an enclosed covering and
assures the required stiffness. In addition, the sensor has fixation holes in both ends
allowing to be sutured to the outer wall of stent-graft or be introduced separately through a
catheter system.
The device has only been tested in an in vitro model but the results demonstrate that
this is a promising technology. Nonetheless, further clinical studies are required to evaluate
the TPS Telemetric Pressure Sensor’s durability and accuracy.
HYPER-IMS
The HYPER-IMS sensor (Figure 5.17) is being developed by Fraunhofer researchers
together with the company Medical Innovation GmbH and other partners in a
BMBF-funded project called “Hyper-IMS” (Intravascular Monitoring System for
Hypertension Patients) [5.55].
Figure 5.17: The HYPER-IMS implantable pressure sensor [5.56].
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The pressure sensor is a fully implantable device and telemetrically controlled
without battery, developed and evaluated for long-term post surgical pressure monitoring.
The newly sensor system design is subdivided into three units (Figure 5.18): a sensor-tip
that is made of a polyurethane tube (measuring 1.05 mm in diameter and 15 mm in length),
connected to a telemetric unit (measuring 17 mm in diameter and 3 mm in height) and a
extracorporeal reader unit. The telemetry and the sensor are connected via micro-cable and
the total dimensions of the system permit a minimal invasive procedure. The device was
evaluated in vivo, in six domestic sheep (weight 60 to 80 kg, 4 for acute experiments and 2
for chronic experiments) and in a vascular model [5.57]. Measurements showed promising
results, but long-term stability of the device in wet and corrosive environments is still
under evaluation.
Figure 5.18: HYPER-IMS blood pressure measurement system [5.58].
5.3.3 A New Proposal: Smart Stent-Graft
Up to now, existing sensors are randomly placed on the aneurysm sac and only provide
information, namely the pressure, regarding a single point. Thereby, it is expected that
joining sensing capabilities in a stent-graft will benefit the future of EVAR with a more
robust technology that enables the placement of more than one sensor, increasing the
sensitivity of the measurements and simplifying the deployment procedure.
A smart stent-graft can be defined as a stent-graft with some in-device mechanism to
perform a given function with communication capabilities to an external element.
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Although there is still no commercial device available, a smart stent-graft could be
decomposed in three elements: a stent-graft, a sensing element and a display.
The stent-graft, besides shielding the aneurysm from the blood pressure, has built-in
sensing elements that are able to gather information concerning the patient’s health and/or
the prosthesis performance. The information gathered is then sent to an external element
(a display) that can be used to diagnose the patient’s or in the comprehension of the
aneurysm’s sac behavior after the implementation of the stent-graft. Figure 5.19 shows the
main components necessary to build a smart stent-graft: a stent, a graft, a sensor
(or multiple sensors) and a readout system.
Figure 5.19: Smart stent-graft concept and main building blocks.
The work performed and shown here did not aim to build a smart stent-graft, but to
prove the feasibility of a smart stent-graft, by designing and fabricating the required
sensor(s). Since the novelty is not on the stent-graft itself, but on the added functionalities,
the main developments were directed towards the sensor development (Figure 5.20) and
telemetry system to improve the device’s reliability.
From the point of view of the sensor’s development, along with the sensor design, a
new fabrication method was devised to comply with the challenges raised by the
application. Due to the specificities of the application, the sensor must be:
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Thin – to fit inside the catheter without increasing catheter size;
Flexible – it must cross a tortuous path till the deployment site;
No internal power supply – the sensor should be able to operate with external
power;
Enable multiple sensors placement – to improve knowledge of aneurysm behavior
after EVAR procedure.
Figure 5.20: Main elements that compose the flexible pressure sensor: a flexible substrate made of
PDMS and the conductive elements based on A-CNTs.
5.3.3.1 Flexible Capacitive Pressure Sensor Requirements
Research on implantable pressure sensors is very active and has been supported and
justified by the need of continuous pressure monitoring, as an early diagnostic mechanism
for some risk patients and for post-EVAR surveillance. Potkay [5.59] defined a set of
requirements for pressure sensors according to the problem to be addressed. In the case of
EVAR, the requirements of the pressure sensor are shown in Table 5.3.
After EVAR procedure, the aneurysm sac gets depressurized and the pressure drops
down to a few mmHg. A numerical study performed by Li and Kleinstreuer [5.60] shows
that after EVAR, the pressure inside the aneurysm sac is around 12 % of the current
luminal pressure. Therefore, if one wants to sense the luminal pressure value (which ranges
typically between 60 – 160 mmHg) through the aneurysm sac pressure, the sensor must be
able to measure pressures between 6 – 26 mmHg with a resolution of 0.1 mmHg. In
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addition, the sensor needs a high dynamic range in order to detect possible complications
(in this case the sac gets pressurized and pressure increases to the luminal pressure values).
Therefore, the sensor should be able to measure pressures above 100 mmHg in case of
aneurysm sac depressurization.
Table 5.3: Requirements for post-EVAR blood pressure monitoring [5.59].
Characteristic Value
Measurement location Between graft and aneurysm wall
Typical values (mmHg) 20 to 90
Measurement range (mmHg) 20 to 250
Measurement resolution (mmHg) 1
Measurement absolute accuracy
(mmHg) 5
Signal bandwidth (Hz) 0 to 80
Arterial diameter (mm) 15 to 40
Ideally, the pressure sensor should be able to monitor the material degradation of the
stent-graft, detect its migration and leakages. While the first may require the use of
conductive wires woven within the graft material [5.61], pressure sensing within the
aneurysm sac promises to be a good candidate for leakage detection and even migration.
Therefore, the success candidate for a new smart sent-graft includes built-in pressure
sensor(s) and communication capabilities.
In order to fulfill with the demanding characteristics of implantable devices the
sensor needs to be biocompatible, biostable, non-toxic, non-allergic, non-carcinogenic,
durable and, fatigue and corrosion resistant. Furthermore, it will have to be tolerated by the
human body without causing a foreign body reaction or an inflammatory reaction. Also,
the device should have a small size and flexibility due to the deployment method, thus
reducing significantly the list of available materials and technological solutions.
Another important parameter is the transmission of the measure data. Ideally, the
new smart stent-graft must be able to transmit the data without any internal power. Also,
the data cannot interfere with other implants nor be influenced by other electronic signals.
A possible solution is the use of passive telemetry, similar to the method used by the
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EndoSure Wireless AAA Pressure Sensor, in which an inductive coupling allows
making an indirect measurement of the pressure changes without any internal power.
Regarding the mechanical requisites, the device should be flexible and tough. Its
components should also be mechanically durable, as well as excellent corrosion
resistance. For a successful protection of the blood vessel, the device should have a design
as less invasive as possible in order to minimize flow resistance and pressure drops.
Radial force is another relevant feature, not only for stent-grafts to stay open without being
crushed with muscular activity, but also to provide a good seal and to ensure fixation.
The deployment of the device is a critical step for the procedure’s success. Thus, the
stent-graft should have a low profile to facilitate the deployment and minimize lesions in
the access arteries. At this stage, radiopacity is also crucial to ensure the correct positioning
of the prosthesis. From the commercial point of view, the device must be capable
of being adequately sterilized and stored as an "off-the-shelf" product. Also, a broad range
of sizes is desirable since it allows the treatment of a wider array of aneurysm anatomies.
To assure patient’s comfort, the measurement protocol should be possible at home
and the results transmitted to the doctor’s office. If not possible, the exam in the doctor’s
office should be quick and the least invasive as possible, avoiding any kind of pain or even
discomfort.
Currently available solutions allow having sensors randomly located on the
aneurysm sac and measuring one single pressure point. The ideal solution would
allow drawing the pressure profile within the aneurysm sac, benefitting the future of
EVAR not just for simplifying the deployment procedure, but also helping in
diagnosing possible complications as well as to study the behaviour of the aneurysm sac
after the integration of stent-grafts. The use of a flexible substrate enables the
conformability of the sensor to the stent-graft and thus the aorta. Compared to currently
available devices, this aspect brings several advantages since the sensor can be
attached to the stent-graft and delivered in a single procedure (as opposed to the
requirement of two catheters for the CardioMems device) and it enables the
placement of more than one sensor (a sensor cluster) contributing to a more comprehensive
study of post-EVAR aneurysm evolution (that is currently not possible).
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5.3.3.2 Telemetry System Architecture
The pressure measurement system must be able to detect post-EVAR complications
and to transmit the measured data without the need for an embedded power supply. The
pressure measurement system comprises an implantable LC network (which consists of an
inductor – L and a capacitor – C) and an external reader (Figure 5.21). Passive telemetry is
used to externally read the pressure changes detected by the capacitive pressure sensor.
Figure 5.21: Block diagram of the monitoring system [5.62].
Regarding the main specifications for the capacitive pressure sensor, namely
dynamic range, resolution and accuracy, they can be retrieved from the maximum values
of blood pressure within the human body and required accuracy and maximum
admissible errors defined by the legislation for pressure measurement devices. The US
Department of Health and Human Services defines [5.63] the accuracy range for
automated pressure sensors as shown in Table 5.4. Considering a maximum cardiac cycle
of 180 cycles per minute (3 Hz) and a change of rate below 120 ms, the sensor reading
bandwidth needs to be around 80 Hz. For healthy subjects, the blood pressure in the aorta
ranges typically from 60 to 150 mmHg [5.64].
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Table 5.4: ANSI/AAMI SP-10-1993 national standard for electronic or automated
sphygmomanometers [5.63].
Accuracy Range
± 3 mmHg or ± 2 % of reading
(whichever is greater) 20 to 250 mmHg
Indicated pressure should be less than 23
mmHg
Input pressure greater than 0 but less than
20 mmHg
Indicated pressure should be 0 ± 3
mmHg
Input pressure less than or equal to 0
mmHg
Indicated pressure should be greater than
245 mmHg or an over range as indicated Input pressure greater than 250 mmHg
The pressure monitoring system relies on passive telemetry for external readout of
the pressure sensor signal based on an implantable LC resonant network. Mokwa [5.65]
presents an overview over implantable medical devices based on microsystems, where the
use of passive telemetry is well established. Selecting the most adequate frequency to
communicate with electronic implanted devices requires balancing the electromagnetic
wave reflection at the air-skin interface and the absorption (causing heating) in the tissues.
The communication with the implanted LC resonant network relies on an
inductive-coupling link in the 12.5 to 20.0 MHz frequency band, specifically allocated for
use in industrial, medical and scientific applications [5.66]. The working distances are in
the order of ~ 5 cm and then, frequencies above 50 MHz should be avoided. Frequencies
below 20 MHz are preferred because a penetration depth of 15 to 20 cm can be ensured
[5.67]. In addition, it guarantees some protection against biological effects and
interferences generated by other electronic equipment and appliances.
Previous studies have shown that radio frequency (RF) energy in the 1 MHz to
10 MHz frequency band penetrates the body with minimum energy loss. Lower
frequencies are mostly reflected on the skin-air interface and the higher ones are absorbed
by the tissues [5.68]. Furthermore, lower frequencies require larger inductors and
capacitors which would affect the overall sensors’ area.
The effects of electromagnetic radiation on biological systems are still an open
discussion subject. In fact, both beneficial and damaging effects to humans’ health have
been identified, which depend on the used frequencies and intensity of the applied
electromagnetic radiation [5.69], but the biological consequences on genes and proteins
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still need further understanding. Nevertheless, the thermal effects are well known and
documented. This issue was taken into account as the sensor remains in a fixed position
and then, the electromagnetic radiation illuminates always the same part of the body, even
if it is in a short period of time (each time in which a reading is carried-out).
5.4 New Approaches for the Implantable Sensor Fabrication
Chapter 4 presented the fabrication of a flexible pressure sensor based on
A-CNTs/PDMS nanocomposites. A-CNTs/PDMS nanocomposites were a good candidate
to develop the full implantable system (sensor plus inductor) since it is possible to pattern
the catalyst in any desirable geometry. However, due to the low conductivity values
obtained for the nanocomposite films, the inductor would have a huge series resistance
(low quality factor) making the solution unsuitable for the required application. Therefore
new alternatives to develop an implantable flexible pressure sensing system (capacitive
sensor plus inductor) are presented here.
The following sections present thoroughly the work carried out, both failures and
successes during the execution of the work. The test setup and the main experimental
results are also presented.
5.4.1 Fabrication of Embedded Inductors
In previous chapter it was reported that the low conductivity of the fabricated
A-CNTs/PDMS nanocomposites compromises its use to implement an inductor with a high
quality factor. Therefore, an alternative solution was pursued consisting in the use of
carbon fibers (with four orders of magnitude higher electrical conductivity) followed by
impregnation by PDMS.
The process is compatible with the A-CNTS/PDMS sensor fabrication process
previously reported and uses a steel mold to position the carbon fibers in a spiral shape.
The higher conductivity of the carbon fibers enables the fabrication of an inductor
with small series resistance. A photograph of the prototype inductors is presented in
Figure 5.22.
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Figure 5.22: Inductors fabrication. a) Steel master mold for oval spiral inductors and
b) Inductor prototypes.
A steel mold where the carbon fibers were positioned in a spiral shape is shown in
Figure 5.21a. Then, the PDMS is poured on the conductive material followed by its cure,
resulting in PDMS encapsulated inductors (Figure 5.21b). Positioning of the fibers in the
mold turned out to be very difficult and it was not possible to have a perfect spacing
between the spirals. Moreover, measurements performed on the prototype inductors
revealed no inductive behavior (the prototypes were not acting as an inductor)
compromising this attempt to fabricate the inductors.
A third similar approach was then tried in order to fabricate new inductors. In this
attempt, thin copper wires (diameter of 50 µm) were used instead of the carbon fibers. This
would allow an even larger conductivity, which could result in inductors with very low
series resistance and a LC network with a large quality factor. Unfortunately, the low
flexibility of the copper wires made it impossible to pass them through the spiral mold and
thus compromised the fabrication of inductors using this approach.
5.4.2 Fabrication of Inductors Using a Flexible Printed Circuit Board
Technology
In face of the poor results obtained for the fabrication of the inductor, alternative
approaches were devised to overcome the problem. At this point it was clear that a good
quality factor for the LC network is necessary and that nanocomposites materials, although
very flexible and with very good stretching capabilities, do not have the required electrical
conductivity. Nevertheless, the results already obtained for the pressure sensors
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(Chapter 4) suggested that a mixed technology using A-CNTs/PDMS nanocomposites for
the capacitor electrodes and another technology for the inductor (that would guarantee
some degree of flexibility) could be used.
The final developed fabrication process for the thin, flexible capacitive pressure
sensors with attached inductor is based on flexible printed circuit board (PCB) technology
(for the inductor fabrication) and A-CNTs/PDMS nanocomposites for the capacitor
electrodes. The flex-PCB technology (commercially available) enables patterning
(100 µm of minimum feature) of copper on top of thin polyimide films for the inductors
production and after, nanocomposite films are bonded to the polyimide film for the
capacitor fabrication. A detailed overview of the process is presented next.
5.4.2.1 Flex-PCB Process Flow
The final flex-PCB plus nanocomposites fabrication process is depicted in
Figure 5.23. The first step is the fabrication of the flex-PCB, which uses a layer of
DuPontTM
Kapton® polyimide film (50 µm thickness) with 35 µm thickness of
copper on the top and bottom layers. Vias can be designed to connect top and bottom
layer. It is also possible to create openings, which in this work were used to build the
capacitor dielectric.
In a second step, a small layer of gold or palladium (~ 20 nm) was deposited
over fabricated A-CNTs/PDMS nanocomposite films (using the process described
and characterized in Chapter 3). This small layer does not change the mechanical
properties of the nanocomposite, but strongly reduces the electrical contact resistance.
Next, the A-CNTs/PDMS nanocomposite films were bonded to the flex-PCB
using a conductive adhesive layer based on silicone. The adhesive layer is very thin
and requires a thermal cure treatment. Finally, a small layer of PDMS was applied
all over the sensor for biocompatibility reasons and to strengthen the bonding between the
polyimide and the nanocomposites. The process is simple, with few steps and enables the
realization of thin flexible capacitive pressure sensors along with an inductor for passive
telemetry use.
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Figure 5.23: Overview of the final process for fabrication of the passive LC network.
Figure 5.24 shows a flexible PCB panel containing six different geometries for the
pressure sensor. Dimensions of the LC network range from 16 x 16 x 0.12 mm3 for the
smaller devices to 20 x 20 x 0.12 mm3 for the larger ones.
Figure 5.24: Flex-PCB panel with six different geometries after fabrication (Step 1) and a
photograph showing an inductor being bended.
All flex-PCB antennas have coils with a 0.1 mm width and a space of 0.1 mm
between them. The three arches between the inductor's coils and the nanocomposite films
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have a 0.25 mm thickness and are spaced by 1.125 mm. The distance between the inner
coil and the largest arch is 0.15 mm for the square antennas and 0.25 mm for the circle
antennas. Table 5.5 displays the outer and inner diameter dimensions, as well as the
number of coils for each antenna and Figure 5.25 presents an optical image showing details
of a square geometry for the pressure sensor.
Table 5.5: Dimensions of the flex-PCB for different geometries [5.70].
Parameters C4 C6 C8 S4 S6 S8
Geometry Circle Circle Circle Square Square Square
dout (mm) 16.3 17.5 18.8 16.3 17.5 18.5
din (mm) 10.5 12.5 14.5 10.3 12.3 14.3
Ncoils 15 13 11 15 13 11
Dielectric (mm) 4 6 8 4 6 8
Figure 5.25: Optical image of a square geometry (S6) for capacitive sensor and details
of the copper inductors.
The final assembled devices are shown in Figure 5.26. The proposed technique for
bonding proved successful and full LC networks with good quality factor were available
for final integration.
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Figure 5.26: Digital images of the final assembled flexible pressure sensors.
5.4.3 Readout Telemetry Implementation
The characterization of the sensing system requires a readout telemetry unit. The telemetry
unit was developed elsewhere [5.71, 5.72] and used here for the characterization of the
sensor system. The design of the reader circuit involved simulations using
ADS Agilent (Advanced Design System) and LTSpice (simulation program with integrated
circuit emphasis) simulators. The final design results were compared with those provided
by analytical modeling implemented in MATLAB. For experimental evaluation purposes,
a printed circuit board (PCB) was implemented with a square planar inductor, a capacitor
and a resistor (Figure 5.27).
An OPA3695 (Texas Instruments) operational current feedback amplifier
was selected for the amplifying block due to its high slew rate and bandwidth
characteristics. The filter is a 6th
order Chebyshev low-pass passive structure required to
eliminate harmonics that compromise the analogue to digital conversion range and the
resonance frequency identification accuracy. The filter has a 0.3 dB ripple in the pass-band,
cut-off frequency at 21 MHz, and 29 dB attenuation at 30 MHz. The reader circuit is
connected by a 50 Ω SMA cable (SubMiniature version A) to the ADC board
(DC1371A from Linear Technology). The ADC operates with a 250 MHz sinewave clock
signal supplied by a Rohde & Schwarz signal generator. The acquisition board is
connected to a PC with an USB cable, where a MATLAB script is used to control the
board operation and process the captured signal using a FFT algorithm to determine the
signals main frequencies.
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Figure 5.27: Prototype showing the reader circuit with receiver inductor placed next
to the sensor’s inductor.
5.5 Experimental Measurements and Results
The designed and fabricated pressure sensors were measured using the electronic telemetry
readout unit developed by the work team at FEUP, and whose complete characterization
and results was presented in [5.70].
Several measurements were performed to test the full system. Initially, the wireless
pressure sensors were tested in a vacuum chamber to acquire the sensor’s characteristic
curve. Afterwards, the sensors were tested in a hydraulic environment that resembles the
real application, i.e., a setup that mimics the aortic aneurysm with the stent-graft inserted,
including varying pressure conditions according to the cardiac cycle. The full system was
tested on this laboratory conditions.
5.5.1 Experimental Results Obtained Using a Vacuum Chamber
The performance of the flexible sensors to different pressures and its use with the
telemetric system was evaluated using a vacuum chamber. The sensors were subjected to
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different pressures by applying a pressure varying from 0 to 1 atm (vacuum to atmospheric
pressure). The output of the reader circuit was recorded while a commercial pressure
sensor placed also inside the chamber was used to record reference values for comparison
purposes. Sensors with different geometries (shape and size of dielectric and inductor)
were tested using the setup shown in Figure 5.28.
Figure 5.28: Experimental test setup for measurements in the vacuum chamber.
The measured oscillation frequencies were centered around 5.3 MHz. Although these
were designed to oscillate in higher frequencies, the characterization of the
inductors revealed that a higher inductance was obtained only for frequencies bellow 8
MHz. Moreover, the inductors showed a self-resonance at that frequency, revealing a
dominant capacitance effect for higher frequencies. This is likely due to the dual-layer
configuration of the inductors, which causes tracks overlap, resulting in large parasitic
capacitances and consequently lower oscillation frequencies for the LC sensor.
This effect can easily be corrected by slightly changes on the inductors design
in order to minimize the tracks overlap. Despite this particularity (that is
less severe for circular inductors since there is less overlap), the sensors responded to
pressure variations with a well-defined characteristic curve, as shown in Figure 5.29
and Figure 5.30.
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Figure 5.29: Oscillation frequency results from all trials with the S4 pressure sensor.
Figure 5.30: Oscillation frequency results from all trials with the S6 pressure sensor.
As depicted in Table 5.6, the variation of oscillation frequency rate (dfosc/dP) is not
constant, but its sign does not change, meaning that an increase in the pressure always
corresponds to a decrease of the oscillation frequency. The characteristic curve in
Figure 5.29 and Figure 5.30 shows that the sensor's best sensitivity would be that in the
range of 0.8 to 1.0 atm (the behavior is similar in the 1 to 1.2 atm pressure range). The
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blood pressure in the aneurysm sac is predicted to be in a range (below 0.3 atm using 1 atm
as reference pressure, i.e., the blood pressure in the aneurysm sac can range from
1 to 1.3 atm), meaning that for use in the proposed application the LC sensor network have
the highest sensitivity in the range of interest.
Table 5.6: Sensitivity of the detected oscillation frequency obtained for S4 and S6 sensors.
Sensor Mean deviation
(kHz)
AVR (0-1 atm)
(kHz/atm)
AVR (0-0.2 atm)
(kHz/atm)
AVR (0.8-1 atm)
(kHz/atm)
S4 1.19 -35.8 -11.0 -59.2
S6 1.57 -45.4 -12.5 -79.0
AVR – Average Variation Rate.
The values in Table 5.6 show that the larger the area of the dielectric, the larger the
variation in the LC oscillation frequency for the same pressure variations. This is
expectable as capacitance is proportional to the area of the dielectric (regarding to the
variable capacitor, which is the sensor element). Moreover, the variation rate is several
times larger than the mean deviation from the trend-curve, which means that the response
is suitable to be used as a characteristic curve with good sensitivity. Small variation
between different trials suggests good precision. The high pressure difference (between the
dielectric and external pressure) that the sensors experience during the trials (ΔP = 1 atm)
validate the good sealing characteristics of the bonding process used and the high dynamic
range of the sensors.
5.5.2 Sensors Performance in a Hydraulic Test Bench
A hydraulic model that simulates the abdominal aorta was built to test the wireless pressure
sensors and the reader system. This model includes the following elements:
1) An electric 2/2-way solenoid valve, connected to the tap water, to simulate the
cardiac cycle;
2) A tube to connect the valve to the aneurysm region;
3) An aneurysm region made with a flexible material;
4) A stent-graft made with a metal spiral to mimic the stent and a PET material to
mimic the graft sewed and placed inside the aneurysm (Figure 5.31);
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5) A square 6 x 6 mm2 (S6) wireless pressure sensor placed on top of the stent-graft;
6) A Y-shape connector tube, used to mimic the aortic artery’s division into the iliac
arteries;
7) A commercial pressure sensor (MPXV5100GC6U from Freescale) placed in the
tube right before the aneurysm region (a hole was made in the tube to place the
sensor) connected to an acquisition board (NI DAQ USB-6008);
8) The wireless reader system with the reader antenna placed on the aneurysm region.
Figure 5.31: Stent-graft (on the left) and the pressure sensor placed on top of the graft.
The complete setup is displayed in Figure 5.32 and Figure 5.33 shows a diagram of
the hydraulic test bench. The goal was to compare the performance of the developed
pressure sensor with a commercial sensor.
Figure 5.32: Experimental test setup for measurements with the hydraulic model.
1 cm
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Figure 5.33: Diagram of the hydraulic test bench.
Figure 5.34 displays the measured pressure and oscillation frequency during one
trail. In this experiment the measurements started with the tap water closed, and was
slowly opened and closed for six times, to mimic the cardiac cycle.
Figure 5.34: Measure pressure (dotted line) and oscillation frequency (solid line) while the water
tap valve was being opened and closed as in the cardiac cycle.
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Since the control of the tap water was made by hand, the control of the water flow
was imprecise. Nevertheless, when the pressure increases the membranes of the capacitor
sensor get closer and therefore, the capacitance increases too.
The highest oscillation frequencies correspond to the lowest pressures (when the tap
water is closed), since the oscillation frequency is inversely proportional to the capacitance
. As shown in Figure 5.33, there is a slight delay between the
measured pressure and frequency, which is due to the location of the sensors. Since the
water first goes to the commercial sensor and after to the aneurysm sac where the wireless
pressure sensor is inserted, the pressure changes are first detected by the commercial
sensor. Also the detected frequency seems to take longer time to reach the values for lower
pressures, in opposite to what happens within the commercial sensor that shows faster
responses to lower pressures (the aneurysm sac accumulates water and probably the water
flux goes from laminar to turbulent).
As the measurements in the vacuum chamber were performed for small distances
between the wireless sensor and the reader system (5 mm), and the oscillation frequency
depends on the distance and transmission medium, the measured oscillation frequencies
with the hydraulic model cannot be extrapolated to pressures based on the measurements in
the vacuum chamber.
These proof-of-concept experiments demonstrate the suitability of the small, flexible,
biocompatible and passive sensor using a telemetry system, to detect pressure variations.
5.6 Summary
A surveillance mechanism based on the measurement of the pressure in an aneurysm sac is
an attractive solution to detect possible post–EVAR problems. The pressure measurement
system presented here includes a capacitive pressure sensor coupled with an inductor for
passive telemetry readout. Both the flexible sensor and the telemetry system architecture
were introduced and their performance discussed. It was shown that taking advantage of
the system’s characteristics it is possible to build a stent-graft with sensing elements
capable of gathering information, regarding the patient’s health and/or the prosthesis
performance, and to send it to an external element. In addition, comparing to the currently
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available devices, the proposed pressure measurement system can be attached to the stent-
graft, allowing its deployment in a single procedure. Another advantage is the possibility
of incorporating more than one sensor in the stent-graft, contributing to a more complete
study of post-EVAR aneurysm evolution that presently is not possible.
The presented measurement system demonstrates that, after a refinement of these
technological developments, a reliable surveillance of post–EVAR is feasible and can be of
added value in the treatment of aneurysms. The tests performed with the hydraulic model
validate the proposed telemetry system functionality and shows its feasibility to be
integrated into a stent–graft. Both the sensors and readout electronics were fully developed
by the research team and several innovative aspects (both in the sensor and in the telemetry
system) were achieved when compared to the state–of–the–art.
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Endoleakage Monitor,” in Mixed-Signals, Sensors and Systems Test Workshop
(IMS3TW), 2012 18th International, 2012, pp. 17–21.
Chapter 6
Conclusions and Recommendations
This chapter summarizes the objectives and conclusions of the present dissertation, in
respect to each of the chapters and it provides suggestions for future work.
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6.1 A-CNTs/PDMS Nanocomposite Films
Nanocomposite materials have generated considerable interest over the last years and their
use, in several fields, is expected to increase. Among them, CNTs/PDMS nanocomposites
have only recently been explored and the resulting material exhibits a range of interesting
properties that may be analyzed for potential MEMS applications and devices fabrication.
Unlike most of the works presented in the literature where the CNTs are dispersed or
mixed within the neat PDMS matrix, this focused on a new manufacturing process to
produce CNTs/PDMS nanocomposites. The entire nanocomposite structure was fabricated
using as-grown vertically aligned-CNTs embedded in a flexible substrate of PDMS.
The effective embedment of CNTs in the pure PDMS matrix is a critical issue,
because CNTs tend to agglomerate. In this dissertation, A-CNTs/PDMS nanocomposites
were fabricated by slowly pouring the uncured PDMS into the as-grown A-CNTs (1 % Vf)
through capillarity wetting.
The alignment of the CNTs, as well as the distribution quality of the A-CNTs in the
PDMS substrate was studied using the scanning electron microscopy (SEM) technique.
The observations indicated that CNTs preserve its alignment after the wetting process,
allowing a controlled manufacturing of A-CNTs/PDMS nanocomposites. Furthermore,
important mechanical and electrical properties were characterized in detail.
Tensile tests and dynamic mechanical analysis measurements were performed in
order to completely characterize the A-CNTs/PDMS nanocomposites. The Young’s
modulus of the nanocomposite showed a 10 times increase (8.1 MPa) in the longitudinal
direction in comparison to the Young’s modulus of the unreinforced PDMS matrix
(0.8 MPa). Also, the results from these tests were used to derive the full constitutive law of
A-CNTs/PDMS nanocomposites.
Electrical characterization of A-CNTs/PDMS nanocomposites raised concerns
regarding the suitability of the material for the realization of a pressure sensor with passive
telemetry. Due to the low conductivity of the produced nanocomposites, a rather high
resistance of the inductor will be generated and therefore the energy transferred inductively
will be dissipated in the resistor. For this reason, new solutions were also studied to
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overcome this problem and the A-CNTs/PDMS nanocomposite films were successfully
used as a good material to build new flexible pressure sensors.
6.2 Flexible Capacitive Pressure Sensors
Silicon and metal have been conventional materials used for the development of pressure
sensors. However, factors as mechanical properties and lack of biocompatibility have been
limiting their field of applications. In order to overcome some technical issues, this
dissertation presented a new fabrication technique to produce flexible pressure sensors
based on A-CNTs/PDMS nanocomposites.
Due to the demonstrated superior properties of A-CNTs/PDMS nanocomposites,
such as mechanical flexibility and piezoresistivity, a new approach for the design and
fabrication of flexible capacitive pressure sensors using the produced nanocomposite films
was presented. In addition, the telemetry system for wirelessly pressure measurement was
also proposed.
The developed flexible capacitive pressure sensors are composed by three thin
layers, where the top and bottom layers are A-CNTs/PDMS nanocomposites (electrodes)
and the middle layer is made of pure PDMS with a cavity with air defining the dielectric.
Several pressure sensors were tested experimentally and the results were compared with
the proposed analytical and FEM models. These were in good agreement for measurements
at high pressure differences. Nevertheless, significant differences were observed for the
pressure changes in the region of the ambient pressure. The cause was attributed to
pressure changes in the dielectric due to volume changes. After model improvement, a
second batch of measurements were performed and the experimental results fitted
reasonably well with the improved model
Static and dynamic responses of various sensors with different geometries and
thicknesses were also measured. Static results revealed reasonably good linearity in the
range of 0 – 100 kPa, mainly in the region near to the atmospheric pressure. On the other
hand, dynamic response of pressure sensors measured in the same range presented two
distinct comportments. The response time of the sensor during a pressure decreasing step is
lower than for a pressure increasing step, which is justified by the viscoelastic behaviour of
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the PDMS used for the nanocomposites fabrication. In addition, the tested pressure sensors
revealed sensitivities in the range of 2.5 – 20 fF/kPa.
Fabricated A-CNTs/PDMS nanocomposite films present transversely isotropic
mechanical characteristics. As the proposed mechanical analytical model considers only
isotropic materials, FEM simulations were also performed using both isotropic and
transverse isotropic material properties. The results compared relatively well, suggesting
that isotropic material characteristics are a good approach for modeling the capacitive
pressure sensors, validating the proposed analytical mechanical model.
The developed technology to fabricate flexible pressure sensors based on
nanocomposites demonstrates the feasibility of its use for sensing applications and showed
that by controlling the pressure inside the dielectric, it is possible to fabricate sensors with
different sensitivities.
6.3 New Flexible Pressure Sensor for the Treatment of Abdominal
Aortic Aneurysms
Chapter 5 narrowed the use of flexible sensors technology to a specific biomedical
application. Developed flexible pressure sensors were characterized in the context of
abdominal aortic aneurysms, supporting the need of wireless flexible pressure sensors in
the treatment and diagnosis of patients with this type of disease. For this purpose, a new
smart stent-graft was presented which is decomposed in three main elements, namely a
stent-graft, a sensing element and a display. In addition, the specific requirements for the
flexible capacitive pressure sensor design were also identified.
As discussed in chapter 4, the low conductivity values obtained for the
A-CNTs/PDMS nanocomposite films made the initial proposed solution unsuitable for the
required inductive telemetry part of the device application. For that reason, three new
alternatives to fabricate an implantable flexible pressure sensing system (capacitive sensor
plus inductor) were presented. In a first attempt, carbon fibers were used instead of as-
grown CNTs, because of its higher electrical conductivity. The subsequent steps consisted
of positioning the carbon fibers in a spiral shaped mold and it impregnation by PDMS.
However, due to difficulties in positioning the carbon fibers inside the mold it was not
possible to obtain a perfect spacing between the spirals, impeding the fabrication of
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183
inductors using this method. A second approach consisted of using thin cooper wires in
place of the carbon fibers. However, the difficulty in passing the copper wires in the mold,
due to the lack of flexibility, also compromised the fabrication of inductors using this
method. The last fabrication process used to develop the thin and flexible capacitive
pressure sensors with attached inductor consisted of using a flexible printed circuit board
(PCB) technology to fabricate the inductor and A-CNTs/PDMS nanocomposites to build
the capacitor electrodes.
The performance of different flexible sensors to different pressures and its use with
the developed telemetric system was evaluated using a vacuum chamber. For pressures
varying from 0 to 1 atm, experimental measurement results demonstrated that the sensors
responded to pressure variations with a well-defined characteristic curve and the sensor's
best sensitivity is in the range of 0.8 to 1.0 atm. As the blood pressure in the aneurysm sac
can vary from 1 to 1.3 atm, this suggests that the LC sensor network have the highest
sensitivity in the range of interest. The assessment of the wireless pressure sensors as well
as the reader system was done through a hydraulic model built to mimic the abdominal
aorta. The results validated the functionality of the proposed telemetry system and proven
its viability to be included into a stent-graft.
6.4 Recommendations for Future Work
Based on the findings and the conclusions from the current dissertation, the relevant key
research topics for further investigations are identified and listed below.
6.4.1 As-Grown CNTs
The chemical vapor deposition (CVD) process used in this research allows obtaining CNTs
with a volume fraction of approximately 1 %. Further developments should investigate the
growth parameters in order to control the final volume fraction of the CNTs, as well as
their size, length, diameter, alignment, chirality and spacing between nanotubes. These
parameters can affect significantly the mechanical properties of CNTs and thus, controlling
them will be fundamental for the development of nanocomposites for several new
applications.
Conclusions and Recommendations
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184
The CVD process is versatile and frequently implemented to produce CNT with
random growth orientation and vertically aligned forests. To extend this research it would
be interesting to develop a new CVD method for growing horizontally oriented CNTs on
commercial silicon substrates or using different catalysts and substrates. Prior studies show
that vertically and horizontally aligned CNTs can be obtained by manipulating the electric
field applied on the substrate and the flow direction of the gases. The tubes obtained can be
as long as millimeters to centimeters, have controllable location, orientation and length.
Regarding the characterization techniques available for the as-grown CNTs, besides
the scanning electron microscopy (SEM) used in this research to inspect the CNTs
alignment, diameter and length, Raman spectroscopy is a powerful tool that can provide
additional information of the CNTs quality, as well as it can determine the nature of CNTs,
i.e., single-walled carbon nanotubes (SWCNTs), double-walled carbon nanotubes or
multi-walled carbon nanotubes (MWCNTs). Atomic force microscopy (ATM) is another
technique that can provides similar information to that from SEM, but it has the advantage
of characterizing the ends of CNTs. Transmition electron microscopy (TEM)
characterization can determine the type of defects present in the CNTs and it also enables
to investigate the CNTs in terms of the number of the walls.
6.4.2 A-CNTs/PDMS Nanocomposites
In this dissertation, a special attention was given to the embedment of CNTs into the
polymer substrate, in order to obtain a resulting nanocomposite with improved mechanical
and electrical characteristics.
Although A-CNTs/PDMS nanocomposites showed improved mechanical properties
in comparison to the host matrix, their electrical conductivity is low and not adequate for
some applications. Further studies should be performed to check electrical properties as a
function of volume fraction of the CNTs. High volume fraction could increase the
reinforcement of the nanocomposite and improve their electrical conductivity as well. In
addition, CNTs alignment within the PDMS should be improved as the electrical
conductivity is affected by inter-CNTs contacts. Thus, a better alignment of CNTs assures
that electrical properties of individual CNTs are conserved.
In order to complement this study, A-CNTs/PDMS nanocomposites should be tested
for different temperatures used during the curing process and using different combinations
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185
of the base polymer and curing agent. In addition to this study, the wetting process should
be tested for these conditions in order to verify if the reinforcement is still effective.
6.4.3 Development of New Sensor Designs
The technology proposed in this thesis was successfully used in the development of
flexible capacitive pressure sensors based on A-CNTs/PDMS nanocomposites. However,
improvements on the sensor fabrication steps may be required to improve the sensors
dimensions stability and reduce process tolerances.
Future work should investigate new architectures that could be interesting to explore,
such as to pattern the catalyst in different geometries in order to fabricate a larger diversity
of sensors and also with reduced nanocomposite thicknesses to increase the sensors’
sensitivity.
Furthermore, novel interesting applications might follow with the technology
presented in this dissertation and, possibly, to extend the principle to other fields of
difficult local readout. An example that is being studied by the same work team consists in
the development of a novel wearable device for coronary and cardiac surveillance that is
combined with an advanced telemetry system, aiming at stent-grafts monitoring after
intervention of endovascular aneurysm repair (post-EVAR). The main novelty of the
surveillance system consists in the synchronized measurement of the patient
electrocardiogram (ECG) and intra-sac aneurysm pressure. Using a cluster of pressure
sensors will allow to detect pressure variations within the aneurysm sac, which combined
with other elements of the entire system, it contributes to have a more robust and reliable
monitoring.
Finally, although this dissertation focused on pressure measurements, the developed
technology could be used to design sensors that measure other physical parameters.
Also, due to the versatility of A-CNTs/PDMS nanocomposites their use is not restricted to
sensing applications. The outstanding properties obtained by the embedment of CNTs into
a polymer substrate are a promising avenue for the design of improved materials for
structural applications.
Appendix
Appendix
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188
A.1 Examples of Mask Pillars
Dot shape (Ø in µm) Rectangular shape (width in µm)
Appendix
_________________________________________________________________________
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189
A.2 Substrates Shape and Growth Times Used for the Optimization
Process
Table A.2: Si wafer substrates and growth time used in the first growths of A-CNTs forests.
First Growths of A-CNTs Forests
Substrate Nº of Samples Growth Time
Experiment 1
Dot shape (Ø 100-200 µm) 3 0.6 min
Experiment 2
Dot shape (Ø 100-200 µm) 3 0.8 min
Experiment 3
Y-shape 2 2 min
Appendix
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190
A.3 Estimate of A-CNTs Volume Fraction
The PMMA mold was fabricated in rectangular shape and with dimensions of 32 x
14 x 0.4 mm3 (Length x Width x Height). Then, an estimation of the CNTs volume fraction
is:
Then,
Appendix
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191
Product Information
Sylgard®
184 Silicone Elastomer Transparent encapsulant with good flame resistance
APPLICATIONS Sylgard
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Appendix
_________________________________________________________________________
_________________________________________________________________________
192
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193
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Appendix
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194
A.5 List of Publications
International Journal Articles
1. A.T. Sepúlveda, R. Guzmán de Villoria, J.C. Viana, A.J. Pontes, B.L. Wardle and
L.A. Rocha, “Full Elastic Constitutive Relation for Non-isotropic Aligned-
CNTs/PDMS Flexible Nanocomposites”, Nanoscale, vol. 5, no. 11, pp. 4847-4854,
Jun. 2013.
2. C.C. Oliveira, A.T. Sepúlveda, N. Almeida, B.L. Wardle, J.M. da Silva, and L.A.
Rocha, “Implantable Flexible Pressure Measurement System Based on Inductive
Coupling,” IEEE Transactions on Biomedical Engineering, 2013, Submitted.
3. A.T. Sepúlveda, R. Guzmán de Villoria, B.L. Wardle, J.C. Viana, A.J. Pontes, and
L.A. Rocha, “Flexible Pressure Sensors: Modeling and Experimental
Characterization” Procedia Engineering, vol. 47, pp. 1177 – 1180, Jan. 2012.
4. A.T. Sepúlveda, F. Fachin, R. Guzmán de Villoria, B.L. Wardle, J.C. Viana, A.J.
Pontes and L.A. Rocha, “Nanocomposite Flexible Pressure sensor for Biomedical
Applications” Procedia Engineering, vol. 25, pp. 140–143, Jan. 2011.
Book Chapters
1. Isa C.T. Santos, A.T. Sepulveda, J.C. Viana, A.J. Pontes, B.L. Wardle, S. M.
Sampaio, R. Roncon-Albuquerque, João M.R.S. Tavares and L.A. Rocha,
“Improving post-EVAR surveillance with a smart stent-graft”, in Technologies for
medical sciences, Book Series: Lecture Notes in Computational Vision and
Biomechanics, Chapter 14, vol. 1, Natal Jorge, R.M.; Tavares, J.M.R.S.; Pinotti
Barbosa, M.; Slade, A.P. (Eds.), 2012, pp. 267 -289, ISBN 978-94-007-4067-9.
Conferences
1. A.T. Sepúlveda, A.J. Pontes, J.C. Viana, R. Guzman de Villoria, F. Fachin, B.L.
Wardle and L.A. Rocha, “Flexible sensor for blood pressure measurement” in 32nd
Annual International IEEE EMBS Conference of the IEEE Engineering in Medicine
and Biology Society, Boston, USA, August 30 – September 3, pp. 512-515, 2011.
2. A.T. Sepúlveda, A.J. Pontes, J.C. Viana and L.A. Rocha, “Aligned Carbon Nanotubes
composites for flexible Pressure Sensors” in Materiais 2011 - 6th International
Materials Symposium; XV Encontro da Sociedade Portuguesa de Materiais. Trans
Tech Publications Ltd (TTP), Guimarães, Portugal, 18-20 April, 2011.
3. A.T. Sepúlveda, Isa C.T. Santos, F. Fachin, R. Guzmán de Villoria, B.L. Wardle, A.J.
Pontes, J.C. Viana and L.A. Rocha “Design of a pressure sensor for monitoring of
post-endovascular aneurysm repair, in BIOSTEC 2011 - International Joint
Conference on Biomedical Engineering Systems and Technologies, Rome, Italy,
26-29 January 2011, pp. 14-22, 2011.
Appendix
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195
4. L.A. Rocha, A.T. Sepúlveda, A.J. Pontes, J.C. Viana, Isa C. T. Santos and João
Manuel R. S. Tavares “Study of Pressure Sensors Placement using an Abdominal
Aortic Aneurysm (AAA) Model”, in Proc. TMSi 2010 - 6th International
Conference on Technology and Medical Sciences, Porto, Portugal, October 21-23,
pp. 261-265, 2010.
5. Isa C. T. Santos, João Manuel R. S. Tavares, A.T. Sepúlveda, L.A. Rocha, A.J. Pontes
and J.C. Viana “Development of a Flexible Pressure Sensor for Measurement of
Endotension”, in Proc. TMSi 2010 - 6th International Conference on Technology
and Medical Sciences, Porto, Portugal, October 21-23, pp. 89-94, 2010.
6. Isa C.T. Santos, A.T. Sepúlveda, L.A. Rocha and João Manuel R. S. Tavares, “A new
endoprosthesis with sensing capability”, in Proc. 1st International Conference on
Applied Bionics and Biomechanics, Venice, Italy, 14-16 October, 2010.
7. A.T. Sepúlveda, A. Moreira, F. Fachin, B.L. Wardle, J.M. Silva, A.J. Pontes, J.C.
Viana and L.A. Rocha “Inductive-Coupling System for Abdominal Aortic
Aneurysms Monitoring Based on Pressure Sensing”, in MME 2010 - 21st
Micromechanics and Microsystems Europe Workshop, Enschede, The Netherlands,
September 26-29, pp. 92-95, 2010.
Short Abstracts or Posters
1. A.T. Sepúlveda, J.C. Viana, A.J. Pontes and L.A. Rocha, “Technologies for Flexible
Pressure Sensors in the Design of a Smart Stent-Graf,” 3rd Annual MIT-Portugal
Conference, 28-29 May, University of Minho, Guimarães, Portugal, 2012.
2. A.T. Sepúlveda, J.C. Viana, A.J. Pontes and L.A. Rocha, “Mechanical
Characterization of vertically Aligned-CNTs/PDMS nanocomposites for flexible
sensors,” I3N-IV Annual Meeting 2012, 9-10 March, Quiaios, Portugal, 2012.
3. A.T. Sepúlveda, L.A. Rocha, J.C. Viana and A.J. Pontes, “Implantable Flexible
Sensor Based on Aligned-Carbon Nanotubes for Blood Pressure Monitoring,”
Semana da Escola de Engenharia, 20-26 October, University of Minho,
Guimarães, Portugal, 2011.
4. A.T. Sepúlveda, J.C. Viana, L.A. Rocha and A.J. Pontes, “Technologies for Flexible
Sensors in the Design of a Smart Stent-Graft,” I3N-III Annual Meeting 2011, 11-12
February, Fátima, Portugal, 2011.
5. A.T. Sepúlveda, J.C. Viana, A.J. Pontes and L.A. Rocha, “Technologies for Flexible
Pressure Sensors in the Design of a Smart Stent-Graft,” 2nd Annual MIT-Portugal
Conference, 28 September, Faculty of Engineering of the University of Porto,
Porto, Portugal, 2010.
6. A.T. Sepúlveda, A.J. Pontes, J.C. Viana, “Technologies for flexible sensors in the
design of a smart stent graft,” Ciência 2010, Centro de Congressos de Lisboa -
Junqueira, Lisbon, Portugal, 4-7 July, 2010.
7. A.T. Sepúlveda, J.C. Viana, A.J. Pontes and L.A. Rocha,” Technologies for Flexible
Pressure Sensors in the Design of a Smart Stent-Graft,” 1st Annual MIT-Portugal
Conference, 7 July, Centro Cultural de Belém, Lisbon, Portugal, 2009.
