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A micromachined inductive sensor using folded flex-circuit structures and its wireless telemetry applications Sridhar, Vijayalakshmi 2008

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A MICROMACHINED INDUCTIVE SENSOR USING FOLDED FLEX-CIRCUIT STRUCTURES AND ITS WIRELESS TELEMETRY APPLICATIONS by Vij ayalakshmi Sridhar B.E. (Electrical & Electronics Engineering), Anna University, 2006  A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF Master of Applied Science in The Faculty of Graduate Studies (Electrical and Computer Engineering)  The University of British Columbia (Vancouver) October 2008 ©Vijayalakshmi Sridhar, 2008  Abstract This thesis reports a flexible, passive wireless inductive sensor with micromachined variable inductors for telemetric applications. The variable inductor is formed by folding coplanar dual spiral coil with 5-10 mm size that are microfabricated using 50-urn-thick copper-clad polyimide film commonly used for flex-circuit manufacturing. When folded, the two coils are aligned to each other where the mutual inductance depends on the gap between the aligned coils. The sensor can be combined with a variety of hydrogel materials for biomedical and chemical applications. A stimuli-responsive hydrogel element is sandwiched by the folded substrate to modulate the gap, or inductance of the device as it swells!deswells depending on the target parameter. The response of a variable inductor to the displacement of the coils is measured to be 0.40 nH/pin. A sensitivity of 71-110 ppm4tm in wireless frequency measurement is obtained using the passive resonant device that combines the variable inductor with a fixed capacitor created on the polyimide substrate. The fabricated devices are coupled with pH- sensitive poly (vinyl alcohol)-poly (acrylic acid) hydrogel and a commercial wound dressing to experimentally demonstrate wireless monitoring of pH and moisture level within the dressing product, respectively. Theoretical inductive responses of the developed device obtained through finite element analysis and their comparison with the measurement result are also presented.  11  Table of Contents Abstract  ii  Table of Contents  iii  List of Tables  v  List of Figures  vi  Acknowledgments Dedication  viii ix  Co-authorship Statement 1.0 Introduction 1.1 Passive Wireless Sensors  1 1  1.1.1 Principle of Inductive Coupling and Passive Sensing  3  1.1.2 Capacitive Sensors  5  1.1.3 Inductive Sensors  9  1.2 Planar Inductors  10  1.2.1 Modeling of Planar Inductors  11  1.2.2 Flex Circuit Technology for Planar Inductors  13  1.3 Hydrogels  14  1.3.1 Characteristics of Hydrogel  15  1.3.2 Diffusion Properties  17  1.3.3 Applications of Hydrogels  18  1.4 Thesis Outline  19  References  20  2.0 A Micromachined Wireless Sensor Based on Folded Flex-Circuit Structures Combined with Stimuli-Responsive Hydrogels  24  2.1 Introduction  24  2.2 Device Principle and Design  26  2.3 Fabrication  30  2.4 Experimental Results  32  111  2.5 Finite Element Analysis  36  2.6 Discussions  38  2.7 Conclusions  40  References  41  3.0 Conclusions  43  References  45  Appendices  46  A. Fabrication Process  46  A.1 Layer 0: Copper Clad Surface Patterning  46  A.2 Layer 1: Through Hole Etch of Polyimide  48  A.3 Layer 2: Copper Electroplating  50  A.4 Layer 3: Partial Etch of Polyimide  53  References  56  B. Analytical Modeling of Two Coil Device  57  B.1 Modeling Equations  57  B.2 Comparative Results  61  iv  List of Tables 2.1 Behavior of the inner and outer type device with increase in gap —‘d’  28  A. 1 Layer 0 parameters  48  A.2 Layer 1 parameters  49  A.3 Copper plating parameters  52  A.4 Layer 2 parameters  53  A.5 Layer 3 parameters  54  v  List of Figures 1.1  Inductive coupling  3  1.2  Passive sensing  4  1.3  A wireless LC tank circuit with variable capacitor  6  1.4  Wireless pressure sensor with varying capacitor  7  1.5  Wireless sensor with dual coil inductor  7  1.6  A simple model of a hydrogel based capacitive sensor  8  1.7  Variable inductor with a movable core  9  1.8  Wireless sensor with inductance varied by a ferrite core  9  1.9  Planar inductor coils: (a): Hoop type; (b): Spiral type; (c): Meander type  11  1.10 Planar coil with two turns  11  1.11 Hydrogel structure  16  1.12 Hydrogel swelling in a pH buffer solution  17  1.13 Hydrogel piece with radius r 0  18  2.1  (a: upper) The hydrogel  —  based wireless inductive sensor; (b: lower) potential  application to wound dressing 2.2  (a: upper) The variable dual coil inductor; (b: lower) an electrical representation of the device and wireless set-up  2.3  25  26  Sample designs of the devices: (a:upper) an inner-type device with 10-mm size device; (b: lower) an outer-type 5-mm device with optical images showing dried and swelled hydrogel sustained by the additional plate to be sandwiched  29  2.4  Cross sectional view of the fabrication steps  30  2.5  (a: left) The devices after fabrication prior to separation; (b: right) one of the fabricated devices, separated and folded; (c: lower) one of the fabricated devices, in the flat state 32  2.6  Wireless and wired set-ups for characterization of electrical response to the displacement of the folded devices  33  vi  2.7  (a: left) A frequency response measured in the wireless setup; (b:right) dependence of the total inductance and capacitance (both normalized to the values at zero gap) of the devices in the wired set-up shown in Fig. 2.6 34  2.8  Side views of pVA-pAA hydrogel sandwiched by a folded device: (a: upper) dried state; (b: lower) swelled hydrogel pushing the coils apart  2.9  pH sensing  —  34  Comparison of the swelling dimension of pH sensitive hydrogel and  corresponding resonant frequency of 10-mm size inner type device  35  2.10 Wireless measurement results of monitoring a commercial wound dressing product: (a: left) shift in frequency; (b: right) frequency vs. time plot showing the hydrogel’s swelling process 36 2.11 COMSOL -Measurement comparison: (a:upper) angular and parallel displacements between coils; (b:lower) angular displacement with misalignment of coils 38 ..  A. 1  Design pattern showing layer 0 on the copper clad side  47  A.2  Design pattern showing of layer 1 through holes to be made  49  A.3  A sample device after layer 0 and layer 1  50  A.4 Design pattern of layer 2 showing the planar inductor lines and the other capacitor plate 51 A.5 Placement of the sample on the wafer holder 52 A.6  Design pattern of layer 3 with the hinge and the border of the device  54  A.7  Enlarged view of the device with the hinges and the boundaries  55  B. 1  Inner type device  57  B.2  Outer type device  60  B.3  Normalized total inductance vs. displacement: (a:upper) 5-mm size inner type device; (b:lower) 10-mm size outer type device 62  B.4 Normalized total inductance vs. displacement of 5-mm size inner type device comparison between Measured, Matlab program and FEM results 63  —  vii  Acknowledgements  My first thanks to my mother and father who have given all they have to provide me with this education. I sincerely thank my supervisor Dr. Kenichi Takahata who showed me what research is and raised my standards beyond my own expectations. My research work would not have been complete without the help provided by the Electrical & Computer Engineering department technicians especially Mr. Thomas Chelliah who helped me in all my experimental setups and has patiently discussed several aspects of my research work. I thank all fellow Microsystems and Nanotechnology group students, with several of whom I have had discussions, training sessions on various equipments and from several of whom I have borrowed equipments. My biggest source of inspiration is my group mate, Mr. Chakravarty Reddy Alla Chaitanya, with whom I have had the most fulfilling discussions on our research and the excellent rapport shared between us made the whole experience worth while. On the personal front I would like to thank my lab mates Ms. Akila Kannan and Mr. Mrigank Sharma with whom I have spent several nights in the lab and who have been a constant source of motivation. I am also indebted to my sister, Veena Sridhar who has been a strong pillar of support at all times of need.  vii’  ‘To9vty granipa  ix  Co-authorship Statement  This statement confirms that the author of this thesis is the primary person responsible for the research contained. In the journal paper to be submitted, the work of the co-author, Dr. Kenichi Takahata is acknowledged, but the major contributions were made by the author especially the actual fabrications, COMSOL modeling and the analytical modeling were solely the work of the author.  x  Chapter 1 Introduction MEMS is a developing area of research that integrates several fields of science in order to bring out the best devices in terms of size, performance and functionality for several old and new applications [1].  The history of MEMS can be discussed with the aid of integrated circuit (IC) technology to begin with. The number of transistors in the Intel 4004 chip sharply increased from a mere 2250 to 24 million in the Pentium III processors over a period of about 30 years. The density of transistors has doubled every 12-18 months following the Moore’s Law [2]. This feat has essentially been possible because of the matured microfabrication technology and has helped in the growth of several applications including computers, cell phones, flat panel displays, ink-jet printing, fuel cells, automobiles, new drugs and drug delivery techniques and warfare. The MEMS devices generally have considerable advantage over their macroscopic counterparts in terms of higher sensitivity, low noise and usage in several applications that could possibly not accommodate larger devices. An important advantage of the MEMS devices, allows integration of sensors and actuators with the electronics at the chip level which enables them to be combined with existing technology and utilize them in several applications. advantages to the MEMS based devices  —  In effect there are three major  “miniaturization, microelectronics integration  and mass fabrication with precision.”  Sensors are devices that can sense or monitor a phenomenon in six major energy domains —  electrical, mechanical, chemical, radiative, magnetic and thermal [1]. Sensors are  capable of detecting the stimulus in various energy domains and converting it into the electrical energy domain so that the signal can be interfaced with an electronic recorder or controller for further analysis. Sensors can be categorized in two ways  —  physical and  chemical/biological, where physical sensors measure force, acceleration, pressure,  1  temperature and magnetic field and chemical sensors measure parameters such as chemical concentration, pH and protein  —  protein interaction.  This thesis discusses the design, fabrication, and experimental measurement results of a flexible, passive wireless sensor consisting of an inductor circuit  with  micromachined  variable  inductors  to  capacitor (LC) resonant tank  —  remotely  monitor  selected  biomedical/chemical parameters. Hydrogels with functional properties that can make volume change in response to the target parameter are used as a sensing material to modify the variable inductors. Analytical and finite element modeling of the device are also presented to further understand the working of the device.  The first section of this chapter presents literature review on existing passive  -  type  wireless sensors, their working principles and their merits and demerits. The second section brings out detailed discussion on planar inductor, its modeling and on flexible circuit technology. The third section is devoted to the basic details of hydrogels and its important features. The final section provides the outline of this thesis.  1.1  Passive Wireless Sensors  A basic LC tank circuit consists of an inductor and capacitor connected in parallel, exchanging electrical and magnetic energies when excited by an external source. Each LC tank circuit has a characteristic resonant frequency given by f =1 /(27rJTë), that can be monitored wirelessly. Any change in inductance or capacitance can be detected by the change in the resonant frequency. This basic principle is essentially used in several MEMS based sensors by fabricating devices with LC circuits with either a varying inductor or a varying capacitor. When an external physical or chemical parameter that we are interested in causes a change in the inductance or capacitance, it can be detected remotely by a change in the resonant frequency of the device.  Some of the commercially available wireless MEMS sensors for biomedical applications include pacemakers, defibrillators, syncope monitoring devices and insulin pumps. CardioMEMS’ Endosure Wireless Abdominal Aortic Aneurysm (AAA) pressure  2  measurement system that consists of an implanted sensor and an external electronics module that can read the pressure from within a patient’s AAA sac [3], Medtronic’s pacemaker  —  Enpulse and its novel syncope (sudden and unexplained loss of  consciousness) monitoring device  —  Reveal Plus which records the patient’s cardiac  rhythm for about 15 months, are some of the devices available in the market [4]. There are two primary approaches for a passive wireless sensor, one being the capacitive approach where the capacitance is varied (most popular approach) and the other one is an inductive approach in which is the inductance is varied.  1.1.1  Principle of Inductive Coupling and Passive Sensing  When a coil 1 is brought near the vicinity of another current carrying coil 2, then the coil 1 is subjected to the flux created by the current in coil 2 (Fig. 1.1). An alternating current in coil 2 induces a voltage in coil 1 and if this circuit were to be closed, a current would flow through coil 1. The coil 1 and coil 2 are said to be inductively coupled. Mutual inductance is the coupling between the two coils [5].  Excitation Coil 2  Coil I  coupling Figure 1.1: Inductive coupling  The mutual inductance of coil 2, M 21 is caused due to the current in coil 1, which produces a flux linking coil 2. Similarly the mutual inductance of coil 1, Mi is caused due to the current flowing in coil 2 which produces a flux linking coil 1. Both these values are found to be equal. Thus, (1.1) The coupling between the two coils depends on the geometry of the coils, the distance between them and the magnetic medium they are linked in. The coupling coefficient k, is  3  an  giving  expression  the  extent  of coupling  between  r c 2 o,nr o,/2  =  X  Where r 0111 and  +  coil2 1  r 2 coili  two  coils  [5]: (1.2)  k  0112 r 111  the  )  are the radii of the area enclosed by conductor coils 1 and 2 and x is  the distance between the coils. The value of k lies between 0 and 1 where, 1 means total coupling between two coils and 0 indicates no coupling. The coupling can be increased by optimizing the size of the coils and the distance between them.  Passive telemetry systems generally use the inductive coupling principle by connecting a capacitor to the circuit to form a LC tank [6] (Fig. 1.2). The tank circuit operates at a resonant frequency f which can be monitored wirelessly through an external coil. A network analyzer is generally used to monitor the electrical parameters such as the resonant frequency by observing the characteristic dip in the scattering parameters (S 11 used in this research work) or the impedance phase of the external coil corresponding to the resonant frequency of the device. This technique is simple and effective as it does not require any on chip circuitry, making the device more compact, robust and reliable. Since wireless transmission is possible, these devices can be used for mobile applications without worrying about leads for electrical measurements and on board power supply. LC tank M  Excitation__illr magnetic coupling Figure 1.2: Passive sensing  The quality factor of the resonant LC tank is given as:  Q±L  (1.3)  4  Where R is the inherent resistance of the tank circuit and a low R value, gives a high factor. Higher  Q  Q  factor indicates a system with less damping and would translate into a  signal response in the network analyzer (mentioned above) with a sharper peak which helps in observing the shift in resonant frequency within a narrow bandwidth. All these factors can be used to tune the device to a specification before the actual fabrication  1.1.2  Capacitive Sensors  This is the most popular approach where the distance between two capacitor plates is varied due to an external stimulus thus changing the capacitance, which forms the basic principle of a capacitive/electrostatic sensor. The parallel plate capacitor and the comb drive capacitor are the two basic structures used in the MEMS based sensors. Capacitive sensors have been used in several areas such as [1] 1. Inertial sensors  —  Newton’s second law states that an acceleration ‘a’ acting on a  —  body with mass ‘m’ produces an inertial force ‘m*a’, which is capable of changing the electrode spacing of a capacitor. This principle is used in designing accelerometers [7]. 2. Pressure sensors  —  The pressure exerted by any physical or chemical/biological  activity can be translated into a capacitive change, making its usage in several industrial, medical and environmental applications [8]. 3. Flow sensors  —  The similar capacitive principle is used in detecting fluid speed,  flow rate, shear stress [9] etc. of a fluid. 4. Tactile sensors  —  Using multiple electrodes and capacitors that are capable of  sensing the forces from different axes, tactile sensors are created for usage in robotics [10]. The main advantages of a capacitive sensor are its simple working principle involving only two conductors; its low power requirement and a lower sensitivity to temperature change. Capacitive sensors also have advantages over piezoresistive sensors as the resistors need not be fabricated on a pressure sensitive diaphragm and the scaling down of the sensor becomes much easier [11]. The sensitivity is also much higher  (—  lOX) than  that of the piezoresistive sensor. Touch mode capacitance is yet another advantage of the capacitive sensor where the substrate can itself act as an over pressure stop preventing the  5  rupture of the diaphragm. Some of the disadvantages of the capacitive sensors are  —  the  nonlinear change of capacitance with applied pressure, high output impedance and parasitic capacitances involved in the circuitry. Lead transfer and packaging are also concerns involved in a capacitive sensor especially in devices where there is a bonding between two layers. Making an electrical connection from a vacuum sealed cavity (described below) without compromising the seal is a challenge. Capacitive sensors can be combined with a fixed inductor to form a LC tank circuit (Fig. 1.3) [121. The change in the capacitance can be wirelessly monitored by sensing the change in resonant frequency of the tank circuit. Thus a wireless capacitive sensor can be used to monitor several physical and chemical parameters remotely such as the intraocular pressure in the eye [13], pressure in the tire and temperature [14] and humidity [15].  Magnetic coupling_______________  L-C tank sensor  External coil  Figure 1.3: A wireless LC tank circuit with variable capacitor One such sensor is shown in Fig. 1.4 [16], where one of the capacitor plates is formed by a flexible silicon diaphragm, microfabricated with a dissolved wafer process and anodic bonding. This diaphragm moves with the applied external pressure creating a variable gap capacitor. A planar inductor coil is fabricated within the glass cavity, forming the LC tank circuit. With external pressure variations the resonant frequency of the circuit will change. Such devices have major advantages No external leads are necessary to connect -  the inductor and capacitor as they are internally connected. They also do not require bulky power supplies/batteries and hence can be made small and compact enabling the 6  choice of several materials and applications. Avoiding the use of batteries also increases the lifetime of these devices making them more suitable for biomedical implantations.  Pressure sensitive Capacitor *  Silicon  Glass Inductor Coils  * Metal Plate  Figure 1.4: Wireless pressure sensor with varying capacitor  Figure 1.5, shows a sensor [17] for the potential measurement of intraocular pressure in the eye by wirelessly monitoring the change in resonant frequency of the inductive circuit. The inductor here consists of two planar copper coils  —  one fixed and the other  movable using the thin silicon diaphragms. External pressure caused by aqueous humor in the eye changes the pressure applied on the diaphragms. As the distance between the coils varies, the capacitance of the circuit varies causing a change in the resonant frequency.  Pressure sensitive diaphragms  p  fl  To silicon  hconwafe r Cu Inductor Coil  Interconnect  Figure 1.5: Wireless sensor with dual coil inductor  To detect various chemical parameters, hydrogels are used in MEMS based devices for different applications. Several sensors have been used in conjunction with responsive hydrogels [18] for detecting parameters such as pH, glucose concentration, temperature and magnetic fields.  7  In other hydrogel based MEMS devices, hydrogel began to be used for sensing and physically actuating a specific part of the device to get an electric read out of the device in response to the change in target parameter. Such devices are essentially sensors consisting of a passive inductor-capacitor (LC) tank circuit with a variable capacitor [19, 20]. As shown in Fig. 1.6, such a device has a capacitor with one of its electrode made of a flexible diaphragm with the other electrode being fixed.  Sealed cavity  Glass Merrjbrane I.  Silicon  4  Hydrogel  Porous Membrane Figure 1.6: A simple model of a hydrogel based capacitive senor  As discussed in detail in section 1.3, some of the hydrogel are characterized to swell/deswell according to a selected parameter such as pH, glucose concentration, temperature and ionic strength. As the hydrogel swells, it moves the flexible diaphragm of the device against a hermetically sealed cavity. Thus as the gap between the capacitor plates vary, the capacitance value varies as well. This change can be wirelessly monitored as a change in resonant frequency. There are some concerns with such kind of a device. Since the swelling of hydrogel pushes a thin glass diaphragm against a sealed cavity, the robustness of the device is questionable and the setup could also lead to the leak of hydrogel from its cavity. The hydrogels should also be placed in such a way that they do not affect the electrical properties of the device such as the dielectric constant or the conductivity. For instance, the dielectric constant of water absorbed by hydrogels varies significantly with temperature; hence the hydrogel must be placed outside the gap formed between the two capacitor plates. Such restrictions lead to a complex design and fabrication.  8  1.1.3  Inductive Sensors  The resonant frequency of a LC tank circuit can also be varied by changing the inductance of the circuit. This approach though has not been as extensively followed as the capacitive approach. The inductance can be varied by varying the number of turns in a coil, its area of cross section and the permeability associated with it. Several approaches have been followed to micromachine variable inductors in MEMS devices by using  -  movable micromachined coils at different heights [21], differential movement of beams [22], displacement of a ferromagnetic core [23, 24] (Fig. 1.7), that of a metal piece [251 and that of a coil [13] over a planar coil.  Core Silicon  Figure 1.7: Variable inductor with a movable core  A passive sensor using a variable inductor is shown in Fig. 1.8 [26], where a ferrite core is attached to a flexible membrane. As an external pressure is applied, the core moves towards the planar inductor coil placed underneath it in a silicon cavity and increases the inductance and therefore decreases the resonant frequency of the circuit.  Ferrite  Flexible membrane  Silicon  ‘Inductor coil Figure 1.8: Wireless sensor with inductance varied by a ferrite core  9  Such devices are generally solid state, bulky and fragile. It is often very difficult to use these devices in biomedical applications as it is essential that devices are flexible, robust, biocompatible and non intrusive. The micromachining processes associated with assembling of magnetic materials involves complicated procedures. Some of the above mentioned variable inductors also suffer with low inductance and Q values, low operating frequencies and discreet changes in the values of inductance. Packaging is an issue for devices that need anodic bonding.  The presented research work focuses on this inductive approach to potentially overcome the difficulties faced with the capacitive sensors and to design a robust and reliable, passive, flexible wireless inductive sensor with a dual coil planar inductor.  1.2  Planar Inductors  The conventional inductor consists of a wire wound around a magnetic core [27]. With the growth of microelectronics and MEMS devices it became necessary that the inductor was also miniaturized along with other electric components for a high quality factor, low power dissipation, high frequency of operation and low distortion. Hence efforts were taken to build an optimized model for a planar inductor [27, 28] which consists of planar coils and magnetic layers instead of cores.  Planar inductor coils have been suggested for different shapes and patterns. Figure 1.9 shows three types [24] where  —  (a) hoop type is simplest, (b) spiral type gives the  maximum inductance but taking out a lead from the middle of the device is a challenge and (c) meander type gives smaller losses at higher frequencies. The spiral inductor consists of several types in itself such circular, rectangular, octagonal and hexagonal [29] of which the hexagonal and octagonal ones are used widely. The inductance value of the planar inductor depends on factors such as the thickness and width of the coil, the spacing between the turns, the number of turns and the material of the coil and the magnetic medium used.  10  ‘p  H a’  a  b  p’...’  rui‘1 ci C  Figure 1.9: Planar inductor coils: (a): Hoop type; (b): Spiral type; (c): Meander type 1.2.1  Modeling of Planar Inductors  The square shaped spiral planar inductors were chosen for this research work. This section shows the basic modeling of a square shaped planar inductor following the work of Greenhouse [30]. Figure 1.10 shows a simple model of a planar coil with two turns or 8 segments. The arrows point the direction of current flow in each segment. The inductance calculation of this planar coil involves three terms totally: 1. Self inductance of each segment 2. Positive mutual inductance between segments in which the current is flowing in the same direction (1, 5), (4, 8), (3, 7) and (2, 6). -  3. Negative mutual inductance between segments in which the current is flowing in the opposite direction— (1, 7), (1, 3), (5, 7), (5, 3), (4, 6), (4, 2), (8,6) and (8, 2).  2  3 Figure 1.10: Planar coil with two turns  11  Thus the total inductance is given by: L=L+M÷—M  (1.4)  Where, L is the self inductance term, M+ is the positive mutual inductance term and M is the negative mutual inductance term. Each term is elaborated below to completely understand the calculations. 1. The self inductance terms in the coil are: L  +  =  2 L  +  +  4 L  +  5 L  +  6 L  +  7 L  +  8 L  (1.5)  2. The positive mutual inductance terms in the coil are: M  M 1 5  =  +  +M 37 +  M 2 6  (1.6)  Mutual inductance has two components. For example, the component 1 M 5 is caused because of the current flowing in segment 1 and a component 5 M 1 because of the current flowing in segment 5. Both 1 M 5 and 5 M 1 are equal in magnitude and sign. Hence the positive mutual inductance terms are: 2M  =  5 +M 2(M 48 37 +M 26 +M )  (1.7)  3. The negative mutual inductance terms in the coil are: M =M 17 +M 13 +M 57 +M 53 +M 46 +M 42 +M 86 +M 82  (1.8)  Just as in the case of the positive mutual inductance terms, there are two components for each of the term above, hence the total negative mutual inductance is given as: 2M =2(M 17 +M 57 +M 13 +M 53 +M 46 +M 86 +M 42 +M ) 82  (1.9)  Thus the total inductance of the coil is: (L+L +L+L + + 53 57 13 +M 17 + (M 48 +M 15 (M 1 2 4 3 1+21 1—21 I 5+L L 6 +L 7 +L 8 J 37 + M M 26 ) M 46 + M 42 + M 86 +M 82 )  =1  (1.10)  The expression for self inductance of a segment is given by: 5 =21{ln[21/(w+t)]+0.50049+[(w-i-t)/31]} L  (1.11)  Where 1, w and t are the length, width and thickness of the segment in centimeter and the inductance L is given in nanohenries. The mutual inductance (both positive and negative) between two parallel segments is given as: M=21Q Where  (1.12)  Q is the mutual inductance parameter given by the equation below: 12  Q = lnl/GMD)+ [i + (12 /GMD2  )]2  }_ {i + (GMD2 /12)1112  + (GMD/l)  (1.13)  (An expression that combines equations (1.12) and (1.13) is shown as equation (2.2) in Chapter 2.) Here, GMD is the geometric mean distance between the two segments, the mutual inductance is calculated and is given by: 1nGMD  =  ing  _[h/12(g/w)  ) 114 1± [1/6O(g/w)J+ [l/168(g/w)6]+ [l/36o(g/w)8]+(  Ij1/660(g/w)  ]+  J  Here, g is the gap between the segments for which the mutual inductance is being calculated. If two segments x and y are equal then the mutual inductance is calculated using the same relation given as: Ii4 =2lQ  (1.15)  If x and y are of unequal length, then the mutual inductance between them is calculated as: =  (M + M),  )—  (1.16)  Using the above equations, the total inductance of any rectangular shaped spiral inductor coil can be calculated. Appendix B describes a model built using this structure for a two coil device with varying gaps between the coils.  1.2.2  Flex Circuit Technology for Planar Inductors  Flexible electronics or flex circuits technology basically aims at assembling electronic components on flexible polymers for several applications such as in mobile phones, cameras, printers, satellites and televisions. In applications like the radio frequency identification (RFID) tags, low cost polymer such as polyimide is beneficial; as such devices are mostly for one time usage only.  Polyimide is a group of heterocyclic aromatic polymer with excellent mechanical, chemical and thermal properties [1]. Its flexibility and biocompatibility make them even more attractive for several applications such as in scanning thermocouple probes [31] and in medicine as a polyimide based CMOS integrated circuit for a catheter [32]. Other applications show the usage of polyimide as flexible substrates in building field effect transistors [33], and planar inductors [34]. 13  In the presented research work, the flex-circuit technology using polyimide was selected to construct the device, to exploit its exceptional features. It makes the device potentially low cost and easily disposable. Polyimide, increasingly being used in biomedical applications is particularly chosen as the substrate for our device as it can be combined with hydrogel elements with functional properties that swell/deswell with the change in a particular parameter and actuate the surface of the flexible polyimide substrate for an electrical read out.  The fabrication process chosen for the device uses photolithography with dry film photoresists (PR). The choice of laminating dry film PR aids in the fabrication as the spin coating of liquid PR on flexible substrate is extremely cumbersome and inefficient. Wet etching processes were used to make the desired patterns on the polyimide substrate and copper. The planar dual coil inductor and capacitor were made by electroplating copper on the PR moulds. The electroplated copper lines are designed to have higher thickness to obtain a low resistance value, thus making the Q factor of the device high (equation (1.3)).  1.3  Hydrogels  Hydrogels are cross  -  linked three dimensional polymers that are capable of swelling and  deswelling depending on the moisture level around it [4]. They have been used in several applications such as contact lenses, wound dressings and diapers. Their importance came to be known in the biomedical field only in the late 1950’s when poly (2 hydroxyethyl methacrylate)-PHEMA was developed for manufacturing contact lenses. This gel has high water content at equilibrium, excellent flexibility and is biodegradable in nature. Thus it resembled the human tissues and could be used for several other applications. By changing the chemical compositions of the hydrogels, they can be made to react to various parameters such as pH [35, 36, 37], temperature [38], ionic strength [39], magnetic fields [40] etc. Hydrogels swell dramatically when stimulated by the change of such parameters. This swelling though is reversible which makes them an attractive material in several applications such as ophthalmologic devices, biosensors, bio membranes, drug delivery devices etc.  14  1.3.1  Characteristics of Hydrogel  In hydrogels, connection between chains is called as a cross-link or a junction [35]. These inter-chain connections are made by carbon atoms, van der Waals force, hydrogen bonds, chemical bridges and molecular entanglements. Hydrogels swell due to the presence of water or other biological fluids containing water in contact with them. The cross-links in the hydrogels are produced by radiation or chemical reaction. Radiation involves electron beams, gamma rays, X-rays or UV rays. Chemical cross-linking is done with the aid of small molecular weight cross-linking agents that link two chains together.  i) Swelling of hydrogels in solvents After polymerization when a hydrogel is in contact with a fluid it swells. There are generally two reactions happening simultaneously [36]. The first one involves a separated phase, polymer  —  polymer interaction where the gel reaches its maximum hydrophobicity  and shrinks. The second one involves a mixed phase, polymer  —  solvent interaction, i.e  the hydrogel interacts with an aqueous solution and gains the maximum hydrophilicity and swells. The polymer  —  solvent interaction, generates osmotic pressure  the hydrogel to expand and the polymer elastic force  Ateias.  —  Altosmt  forcing  polymer interaction generates a counter active  When these two forces balance each other, equilibrium is achieved  [36]: = A2reias  + Aicsm,  =0  (1.17)  ii) Swelling of pH sensitive hydrogels pH sensitive hydrogels have cross-linked polymer chains containing weak acidic and basic groups [33,34] that can be ionized easily. The hydrogel as such consists of the solid polymer surrounded by the internal solution within its network (Fig. 1.11). When the solid piece of hydrogel is initially immersed in a solvent, some of the chains disassociate into the solvent. The cross-links prevent the entire structure from dissolving away by providing an elastic restoring force to the expanding network. As shown in Fig. 1.11 [36], not all the ionizable groups are fully disassociated.  15  Mobile ion  Liquid  Polymer chain  Un disassociated molecule  Figure 1.11: Hydrogel structure  If a hydrogel with acidic groups in its polymer chains is immersed in a basic solution, then the H ions come off and combine with 01-f ions in the solution as follows [36]: [RCOOH  iei  +  [oH-  deprotonalon  > [icocr  0 2 j +H  (1.18)  The density of like charged groups increases within the hydrogel leading to a charge imbalance and an electrostatic repulsion within the hydrogel network, forcing a polymersolvent interaction. Charge compensation occurs when cations from the solution diffuse into the hydrogel causing the hydrogel to swell. The hydrogel reaches a stable state when there is an equilibrium between the osmotic force caused due to the imbalance in concentration of ions and the elastic restoring force provided by the cross links. If the acidic hydrogel is immersed in an acidic solution [36]:  [RCOO- iei + [H÷ jaq  protonatioi  [RcooH  Je1  (1.19)  there is a charge imbalance again and in this case the hydrogel shrinks, forcing some of the water molecules to move outside the hydrogel network. This typical movement of H ions varies depending on the pH of the solution and thus it can be seen that the swelling/deswelling of the hydrogel varies according to the pH of the solution. This process is shown in Fig. 1.12 [37].  16  • Ionic species diffusing in Ionic species diffusing out Figure 1.12: Hydro gel swelling in a pH buffer solution 1.3.2  Diffusion Properties  Both the swelling and shrinking processes involve the transportation of matter and consume time [36]. There are basically two transport mechanisms used. The first mechanism initiates the process by creating an imbalance in the osmotic pressure. This is caused either by a heat transfer (given by thermal transfer coefficient, DT 1 2 cs) 3 0 / m or by the mass diffusion of a solvent into a hydrogel (given by mass transfer coefficient, 12 5 D c 5 0 / s). m The second transport mechanism involves reactions that occur to counter the imbalance in osmotic force and in effect there is swelling or shrinking of hydrogel due to absorption or release of the swelling agent. Overall, the net motion of the chains in the polymer network and the diffusion of the solvent are described by a “cooperative” diffusion coefficient D . 0 0 The time constant associated with the swelling process is given as [36],r  =  ‘  where, r is the radius of a spherical hydrogel (Fig. 1.13).  coop  17  D coop—(1O -6  .10  -7  2  )cm Is  0 r rmax  Figure 1.13: Hydrogel piece with radius r 0 It is very essential that the solute diffusion into the hydrogels is controlled as this is the basis for several applications. The structure and pore size of the gel, polymer composition, water content and the nature and size of the gel are to be taken into consideration for calculating the diffusion coefficient of the solute.  1.3.3  Applications of Hydrogels  In general, hydrogel’s biomedical compatibility and their sensitivity towards several physiological and biomedical parameters such as pH level, ionic concentration and electromagnetic radiation make them a lucrative material for several research areas. In drug delivery systems [41], it is required that the drug is delivered only when required  and where it is required. For such applications, “intelligent” hydrogels are to be utilized. Hydrogels can be tethered, making it more sensitive to its environment and giving it more control for the diffusion of the drug. By adding functional groups specific for a target molecule, the hydrogel can be sensitized to several conditions. Thus it is very well suited for drug delivery applications.  It is also possible to create “functional” hydrogels [42].  These hydrogels recognize  specific molecules in the environment and undergo corresponding changes. For example grafting an antigen and the corresponding antibody to the network produces extra cross links in the network. When the hydrogel is suspended in a solution containing free 18  antigens, the free antigens bind to the hydrogel thus creating a change in the volume of the gel. Thus by utilizing the antigen —antibody reactions the hydrogel is made to react only to a certain type of antigens.  Another major application of hydrogels is in microfluidics where hydrogel’s properties are used in ‘labs on a chip’ to control the flow of cells and other reagents. Electronically addressable, biocompatible, leak free microfluidic valves have been created for applications in polymerase chain reaction and flow control [38]. Hydrogels have also been used to detect the partial pressure of carbon dioxide [43] by using pH sensitive hydrogels. This has been utilized in the diagnosis of gastrointestinal ischemia. The sensor created is sensitive enough to detect changes as small as 0.5kPa. Hydrogels sensitive to magnetic fields [40] have been investigated for applications in detecting hyperthermia.  1.4  Thesis Outline  The entire thesis is constructed in the following order  -  Chapter 2 consists of a journal  paper in preparation for submission, which explains the proposed device, its working principle, fabrication, experiments conducted and their analysis and the results of finite element modeling of the device. Chapter 3 concludes the research work with emphasis on future work. Appendix A discusses the fabrication process in detail and Appendix B explains the analytical modeling done for the device.  19  References  [1] C. Liu,”Foundations of MEMS,” Pearson International Edition, 2006 [2] R. R. Schaller, “Moore’s law: past, present and future,” Spectrum, IEEE, vol. 34, pp. 52-59, 1997. [3] http://www.cardiomems.com/content.asp?display=aboutus [4] S.S. Saliterman, “Fundamentals of Biomems and Medical Microdevices,” Wiley Interscience, SPIE Press, 2005. [5] K. Finkenzeller, “RFID Handbook  -  Fundamentals and Applications in Contactless  Smart Cards and Identification,” Wiley and Sons Ltd, 2003. [61 A. DeHennis and K. D. Wise, “A double-sided single-chip wireless pressure sensor,” in Micro Electro Mechanical Systems, 2002. The Ffleenth IEEE International Conference on, 2002, pp. 252-25 5. [7]A. Selvakumar and K. Najafi,  “Vertical  comb  array  microactuators,”  Microelectromechanical Systems, Journal of vol. 12, pp. 440-449, 2003. [8] A. V. Chavan and K. D. Wise, “Batch-processed vacuum-sealed capacitive pressure sensors,” Microelectromechanical Systems, Journal of vol. 10, pp. 580-588, 2001. [9] M. A. Schmidt, R. T. Howe, S. D. Senturia, and J. H. Haritonidis, “Design and calibration of a microfabricated floating-element shear-stress sensor,” Electron Devices, IEEE Transactions on, vol. 35, pp. 750-757, 1988. [10] Z. Chu, P. M. Sarro, and S. Middelhoek, “Silicon Three-axial Tactile Sensor,” in Solid-State Sensors and Actuators, 1995 and Eurosensors IX. Transducers ‘95. The 8th International Conference on, pp. 656-659, 1995. [11] M. G. El  -  Hak, “The MEMS Handbook  —  MEMS Design and Fabrication,” Taylor  & Francis Group, 2006. [12] P. Eun-Chul, Y. Jun-Bo, and Y. Euisik, “A Hermetically-Sealed LC Resonator For Remote Pressure Monitoring,” in Microprocesses and Nanotechnology Conference, 1998 International, pp. 91-92, 1998. [13] C. C. Collins, “Miniature Passive Pressure Transensor for Implanting in the Eye,” Biomedical Engineering, IEEE Transactions on, vol. BME-14, pp. 74-83, 1967.  20  [14] M. Nabipoor and B. Y. Majlis, “A new passive telemetry LC pressure and temperature sensor optimized for TPMS,” Journal ofPhysics. Conference Series, p. 770, 2006. [15] T. J. Harpster, S. Hauvespre, M. R. Dokmeci, and K. Najafi, “A passive humidity monitoring system for in situ remote wireless testing of micropackages,” Microelectromechanical Systems, Journal of vol. 11, pp. 61-67, 2002. [16] Orhan Akar, Tayfun Akin, Khalil Najafi, “A wireless batch sealed absolute capacitive pressure sensor,” Sensors and Actuators A. Physical, vol. 95, pp. 29-38, 2001. [17] R. Puers, G. Vandevoorde, and D. D. Bruyker, “Electrodeposited copper inductors for  intraocular  pressure  telemetry,”  Journal  of  Micromechanics  and  Microengineering, p. 124, 2000. [18] T. Schalkhammer, C. Lobmaier, F. Pittner, A. Leitner, H. Brunner, and F. R. Aussenegg, “The use of metal-island-coated pH-sensitive swelling polymers for biosensor applications,” Sensors andActuators B: Chemical, vol. 24, pp. 166-172, [19] M. Lei, A. Baldi, E. Nuxoll, R. Siegel, and B. Ziaie, “Hydrogel-based microsensors for wireless chemical monitoring,” Biomedical Microdevices. [20] Z. A. Strong, A. W. Wang, and C. F. McConaghy, “Hydrogel-Actuated Capacitive Transducer for Wireless Biosensors,” Biomedical Microdevices, vol. 4, pp. 97-103, 2002. [21] S. Chang and S. Sivoththaman, “A Tunable RF MEMS Inductor on Silicon Incorporating an Amorphous Silicon Bimorph in a Low-Temperature Process,” Electron Device Letters, IEEE, vol. 27, pp. 905-907, 2006. [22] I. Zine-El-Abidine, M. Okoniewski, and J. G. McRory, “A Tunable RF MEMS Inductor,” in MEMS, NANO and Smart Systems, 2004. ICMENS 2004. Proceedings. 2004 International Conference on, pp. 636-63 8, 2004. [23] N. Sarkar, D. Yan, E. Home, H. Lu, M. Ellis, J. B. Lee, R. Mansour, A. Nallani, and G. Skidmore, “Microassembled tunable MEMS inductor,” in Micro Electro Mechanical Systems, 2005. MEMS 2005. 18th IEEE International Conference on, pp. 183—186, 2005.  21  [24] V. Th, D. Stefan, and E. Obermeier, “Micro-coil with movable core for application in  an  inductive  displacement  sensor,”  Journal  of Micromechanics  and  Microengineering, pp. 119, 1999. [25] T. Yammouch, K. Okada, and K. Masu, “Physical Modeling of MEMS Variable Inductor,” Circuits and Systems II: Express Briefs, IEEE Transactions on, vol. 55, pp. 419-422, 2008. [26] A. Baldi, W. Choi, and B. Ziaie, “A self-resonant frequency-modulated micromachined passive pressure transensor,” Sensors Journal, IEEE, vol. 3, pp. 728-733, 2003. [27] 0. Oshiro, H. Tsujimoto, and K. Shirae, “A novel miniature planar inductor,” Magnetics, IEEE Transactions on, vol. 23, pp. 3759-376 1, 1987. [28] S. S. Mohan, M. del Mar Hershenson, S. P. Boyd, and T. H. Lee, “Simple accurate expressions for planar spiral inductances,” Solid-State Circuits, IEEE Journal of vol. 34, pp. 1419-1424, 1999 [29] K. Kawabe, H. Koyama, and K. Shirae, “Planar inductor,” Magnetics, IEEE Transactions on, vol. 20, pp. 1804-1806, 1984. [30] H. Greenhouse, “Design of Planar Rectangular Microelectronic Inductors,” Parts, Hybrids, and Packaging, IEEE Transactions on, vol. 10, pp. 101-109, 1974. [31] M. H. Li, J. J. Wu, and Y. B. Gianchandani, “Surface micromachined polyimide scanning thermocouple probes,” Microelectromechanical Systems, Journal of vol. 10, pp. 3-9, 2001. [32] P. Ki-Tae and M. Esashi, “A multilink active catheter with polyimide-based integrated CMOS interface circuits,” Microelectromechanical Systems, Journal of vol. 8, pp. 349-357, 1999. [33] C. J. Drury, C. M. J. Mutsaers, C. M. Hart, M. Matters, and D. M. d. Leeuw, “Lowcost all-polymer integrated circuits,” Applied Physics Letters, vol. 73, pp. 108-110, 1998. [34] M. Woytasik, J. P. Grandchamp, E. Dufour-Gergam, E. Martincic, J. P. Gilles, S. Megherbi, V. Lavalley, and V. Mathet, “Fabrication of planar and three-dimensional microcoils on flexible substrates,” Microsystem Technologies, vol. 12, pp. 973-978, 2006.  22  [35] N. A. Peppas, Y. Huang, M. Torres-Lugo, J. H. Ward, and J. Zhang, “PHYSICOCHEMICAL FOUNDATIONS AND STRUCTURAL DESIGN OF HYDROGELS IN MEDICINE AND BIOLOGY,” Annual Review of Biomedical Engineering, vol. 2, PP. 9-29, 2000. [36] A. Richter, G. Paschew, S. Klatt, J. Lienig, K. F. Arndt and H. J. P. Adler, “Review on Hydrogel-based pH Sensors and Microsensors, “Sensors, vol. 8, pp. 561-581, 2008. [37] S. K. De, N. R. Aluru, B. Johnson, W. C. Crone, D. J. Beebe, and J. Moore, “Equilibrium  swelling  and  kinetics  of pH-responsive  hydrogels:  models,  experiments, and simulations,” Microelectromechanical Systems, Journal of vol. 11, pp. 544-555, 2002. [38] J. Wang, Z. Chen, M. Mauk, K.S. Hong, M. Li, S. Yang, and H.H. Bau, “Self Actuated Thermo  —  Responsive Hydrogel Valves for Lab on a chip,” Biomedical  Microdevices, vol. 7, pp. 3 13-322, 2005 [39] B. Zhao and J. S. Moore, “Fast pH- and Ionic Strength-Responsive Hydrogels in Microchannels,” Langmuir, vol. 17, pp. 4758-4763, 2001. [40] K. L. Ang, S. Venkatraman, and R. V. Ramanujan, “Magnetic PNIPA hydrogels for hyperthermia applications in cancer therapy,” Materials Science and Engineering: C, vol. 27, pp. 347-35 1, 2007. [41] N. A. Peppas, “Intelligent therapeutics: biomimetic systems and nanotechnology in drug delivery,” Advanced Drug Delivery Reviews, vol. 56, pp. 1529-1531, 2004. [42] A. B. R. E. J. H. Rongsheng Zhang, “A smart membrane based on an antigenresponsive hydrogel,” Biotechnology and Bioengineering, vol. 97, pp. 976-984, 2007. [43] S. Herber, J. Bomer, W. Olthuis, P. Bergveld, A. van den Berg, “A Miniaturized Carbon Dioxide Gas Sensor Based on Sensing of pH-Sensitive Hydrogel Swelling with a Pressure Sensor,” Biomedical Microdevices, vol. 7, pp. 197-204, 2005.  23  Chapter 2 A Micromachined Wireless Sensor Based on Folded Flex-Circuit Structures Combined with StimuliResponsive Hydrogels 2.1  Introduction  Hydrogels are cross-linked, three-dimensional polymer networks that have been utilized in medical and sanitary applications such as wound dressing and infant/adult diapers as super absorbents. The material swells/deswells in response to not only the moisture level but also some of the physical/chemical parameters including biomedically relevant ones such as pH, salinity, temperature, and glucose concentration [1]. MEMS-based sensors that combine the passive inductor-capacitor (L-C) resonant tank with stimuli responsive hydrogels for frequency-based sensing have been reported [2, 3]. This is a very attractive approach for biomedical applications because the passive configuration potentially contributes to making the device low cost and disposable. A major approach to construct the sensors has been to use variable capacitors for the L-C tanks with movable diaphragm electrodes that are actuated by the hydrogels. Since this configuration relies on deflection of a relatively thin diaphragm against a sealed cavity, there is a potential concern of robustness of the diaphragm and leaks in the cavity seal. In addition, a hydrogel element must be placed outside of the capacitive gap in order to avoid the influence of the variation in electrical properties of the liquid (e.g. conductivity and dielectric constant). This requirement poses the need for extra structures to hold the hydrogel element, increasing the complexity in the fabrication and construction of the device. The sensing with a passive L-C tank can also be approached by using a variable inductor that responds to the target parameter [4]. There have been some development efforts for micromachined variable inductors with various approaches such as deformation of micromachined coils [5], digitally switched planar coils using microrelays [6], displacement of a magnetic core [4], and that of a metal piece [7] and a planar coil [8] A version of this chapter will be submitted for publication. Vijayalakshmi Sridhar and Kenichi Takahata, “A Micromachined Wireless Sensor Based on Folded Flex-Circuit Structures Combined with Stimuli-Responsive Hydrogels.”  24  with MEMS actuators. However, they are mostly solid-state devices with fragile structures and no physical flexibility, which is often unfavorable for biomedical applications. This paper reports a flexible, passive wireless inductive sensor, eliminating the need of diaphragm/cavity structures in the capacitive approach, with hydrogel elements for biomedical applications (Fig. 2.1 a). They potentially include the monitoring of wound healing processes using pH-sensitive materials (Fig. 2.lb) as the pH level of a wound liquid can be related to the healing process of a wound [9], where an acidic pH indicates the healing of a wound.  Folded flex circuit Interconnect to backside coil & caPacitor\  HydrogeI rFlexural hinge Inductive sensor  Wireless link  \  Breathable protection film  j  Figure 2.1: (a: upper) The hydrogel  —  Skin  based wireless inductive sensor; (b: lower) potential  application to wound addressing  This paper is constituted as follows. Section 2.1 describes the working principle and the design of the inductive sensor studied in this effort. The fabrication process of the device is explained in Section 2.3.  Section 2.4 shows the experimental results from the  characterization of the fabricated devices as well as wireless sensing tests using the devices combined with a pH-responsive hydrogel and a commercial wound dressing. In Section 2.5, theoretical analysis of the sensor response using a finite element method is described and compared with the experimental result, followed by discussion in Section  25  2.6. Section 2.7 draws conclusions from the overall effort.  2.2  Device Principle and Design  The developed device uses the passive L-C circuitry that consists of a variable inductor and fixed capacitor. The variable inductor is coupled with a hydrogel element that makes dimensional changes in response to the change in a target parameter thus modifying the inductance. The construction of the variable inductor consists of two planar spiral coils with identical dimensions and patterns that are connected and aligned to each other with a certain gap between the coils (Fig. 2.2a), similar to the construction observed in a capacitive sensor for measurement of intraocular pressure [10]. The inductance of the dual-coil structure has fixed and variable components. The fixed inductance consists of the self inductance of the coil structure, L, and the mutual inductance within each planar coil, M, which is determined by the spacing between individual segments of the coil leads. The variable component is mutual inductance that is present between the two aligned coils, which depends on the gap spacing between the coils, d.  7 jF ” J7 ” ‘dI To A7 caPrlnnertype  inductance of the acoiI Mutual inductance between the coils (+ for outer type and—for inner e  Ssor External coil Figure 2.2: (a: upper) The variable dual coil inductor; (b: lower) an electrical representation of the device and wireless set-up  26  The total inductance of the dual-coil structure is the combination of both the fixed and variable components, which can be represented as: L ( 0 d)  =  (L  +  M)  +  M(d)  (2.1)  Where M(d) is the variable mutual inductance as a function of d. The mutual inductance between two parallel conductors given as [11]: M =21  r 1  +  LGMD  l  + (12 +  2 GMD  )1J  —  /1+ \i  (GMD 2 1)  +  GMD 1]  (2.2)  where M is the mutual inductance in nanohenries, 1 is the length of the segment in centimeters, and GMD is the geometric mean distance between the conductors, which is approximately equal to the distance between the parallel segments that corresponds to d. It is assumed in the model that the relative permeability of both the conductors and the surrounding medium is 1. The variable mutual inductance M(d) for the folded dual-coil structure with a separation of d can be obtained by summing all the possible combinations of the parallel segments in the different coils. The segments in which the current flows in the same direction contribute towards positive mutual inductance and the segments, in which the current flows in the opposite direction, contribute towards negative mutual inductance. The configuration with which the two aligned coils are coupled so that the current in the coils flows in the same direction (outer type) results in a positive mutual inductance between the coils as the overall variable inductance, M(d), whereas it becomes negative with the opposite configuration (inner type) [121.  The  absolute value of M(d) is identical for both types when the upper and the lower coils have the same design pattern. For the outer device, the positive mutual inductance is maximum in the folded state of the coils with the minimum gap between the coils. As this gap increases, the positive coupling decreases, hence the total inductance decreases. For the inner-type device, the negative mutual inductance is maximum when the coils are folded with minimum gap. As this gap increases their negative coupling decreases and their total inductance value increases. Table 2.1 summarizes the inductive responses for the two configurations. The sensitivity of the inner type device (S ) and that of the outer 1 type device (S ) to displacement can respectively be expressed as: 0  s— —  [L M (d)}Ad -  8 oul  ‘  AM [L + M (d)]d  23  27  Where L represents the fixed inductance, (L + M) from equation (2.1), and 1XM represents  [M(d+Ad)  -  M(d)J. In the case where the variable component AM,, is  identical for both the inner and outer type devices (i.e. the only difference between them is the direction of the current flow), equation (2.3) can be rewritten as: SmL+Mv(d) 01 L—M,,(d) S  (2.4)  This suggests that the inner type theoretically provides a higher sensitivity compared to the outer type.  Table 2.1: Behavior of the inner and outer type device with increase in gap —‘d’ Inductance  Mutual, M(d)I  Outer Type  (positive)  Inner Type  (negative)  Total, L(d)  t  The dependence of the inductance on the gap between the two coils can be used to inductively detect dimensional changes of responsive materials such as hydrogels placed between the coils.  An L-C circuit can be formed by coupling a fixed parallel-plate  capacitor to the variable inductor for frequency-based readout of the variable inductance. This configuration can be extended to a wireless implementation of the sensing, where the inductor of the device is magnetically coupled with the external coil for monitoring of the resonant frequency of the circuit (Fig. 2.2b). The principle is expected to provide ease in the construction of the inductive sensor as the entire device structure can be fabricated without dealing with magnetic materials that often causes the complexity and high cost in the device manufacturing.  To form the aligned coils and use it for hydrogel-based sensing with a simple, cost effective implementation, this study leverages the flex-circuit technology for the device fabrication. The aligned dual-coil configuration is formed by an L-C circuit fabricated on a flexible substrate folded at the flexural hinge that is created by thinning the substrate in the defined hinge region. The devices were designed to have coils with approximately 5or 10-mm overall size and various numbers of spiral turns as well as different  28  constructions for hydrogel assembly (Fig. 2.3). The perforations formed in the substrate provide paths for liquid-phase analyte to reach a piece of hydrogel, which can be sandwiched by the folded substrate that is tied at the end of the folded substrate. The interlock mechanism with a combination of a latch and a slot created in the flexible substrate (Fig. 2.3b) can optionally be used for holding the folded substrate and the hydrogel element in place. Alternatively, synthesized hydrogel solution can be applied to the aligned mesh structures that are created on both sides of the folded substrate (Fig. 2.3a) and dried so that the hydrogel itself ties the folded structure. Another method to couple the hydrogel with the device is to use a scaffold structure formed in an additional plate to be sandwiched between the two coils (Fig. 2.3b). 27.1mm Hole :600 x 600 urn 2  S E  Coil size  —  lOx 10mm , line width:lOOpm, Spacing:l5Opm 2  Flexural hinges Perforations Coil size: 5x5mm Line width: 7Oiim Spacing: 450im  pVA-pAA hydrogel  Figure 2.3: Sample designs of the devices: (a: upper) an inner-type device with 10-mm size device; (b: lower) an outer-type 5-mm device with optical images showing dried and swelled hydrogel sustained by the additional plate to be sandwiched  29  2.3  Fabrication  In this effort, the devices were fabricated using single sided 1 8-jim thick copper-clad polyimide film with thickness of 50 jim (Novaclad G2300, Sheldahl Co., MN, USA) as the substrate material.  All the patterning steps were implemented with dry-film  photoresists laminated on the substrates using a hot-roll laminator (XRL- 120, Western Magnum Co., CA, USA) at 120 °C and a feed speed of 1.3 cmls. The photolithography was performed using the mylar masks and a standard mask aligner. ask# 2) Etch P1 (Mask#2)  cladsideu Dry-film photoresist  j:  .  AA  3) Deposit Ti-Cu films & electroplate Cu (Mask #3)  Electroplated Cu\  __—  Ti-c?ufNms  CoNs._  .-  S  —  I A A Capac  —-  Hinge 5) Etch P1 (Mask #4) 6) Coat non  Parylene-,  sandwha hdroel  Figure 2.4: Cross sectional view of the fabrication steps  The developed fabrication process uses four masks in total (Fig. 2.4). The process starts from patterning of the copper clad using 15-jim-thick photoresist (SF306, MacDermid, Inc., CT, USA) laminated on the copper to form one electrode of the capacitor and a contact pad. The copper layer is patterned by wet etching in a ferric chloride based solution using the photoresist mask (step-i).  Next, another photoresist with 38-jim  thickness (PM240, DuPont, Inc., ON, Canada) is laminated and patterned on the  30  polyimide side of the substrate to perform wet etching of polyimide in a KOH-based solution (40-wt. % KOH and 20-wt. % ethanolamine in de-ionized (DI) water) for 4 mm at 87 °C to create perforations in the substrate (step-2). The perforations are used for via contact formation as well as access holes for liquid analyte. A titanium (200 nm)-copper (1 tm) thin film as an adhesion-seed layer for the subsequent copper electroplating process is evaporated on the polyimide side of the substrate. Prior to the evaporation, the polyimide substrate is processed with oxygen plasma to enhance the adhesion between titanium and the substrate. The PM240 resist is again laminated on the polyimide side and patterned to be used as a mold for the copper electroplating process in a sulfuricacid-based bath to form the planar dual coil and the other electrode of the capacitor with thickness of 30-3 5 urn (step-3). After stripping the resist, the electroplated structures are electrically isolated from each other by wet etching of the copper seed layer followed by that of titanium adhesion layer (step-4). Timed polyimide etching is subsequently implemented to thin the patterns of the flexural hinges as well as the borders of the devices (to be used for manual cut from the substrate) to a thickness of 10tm (step-5). During all the etching and electroplating steps, the side that is not processed is protected by a laminated resist layer. Figure 2.5 shows the fabricated devices after the step 5 in Fig. 2.4. The inductance, resistance, and capacitance of the fabricated device in Fig. 2.5c were measured to be 1.7 jiH, 3.9 ), and 30.88 pF, respectively, which match well with theoretical values 1.89 jiH with the formula for planar square spiral coils in [13J, 2.93 2, and 16.22 pF. A higher capacitance measured is predicted as it includes parasitics associated with the planar coils. Individual devices separated from the substrate are coated with Parylene-C TM polymer with thickness of 1 tm for making electrical insulation as well as the entire surfaces biocompatible (step-6), and then the substrate is manually folded to sandwich a hydrogel element (step-7). Two types of hydrogel elements were coupled with the device in this study.  One is poly(vinyl alcohol)  poly(acrylic acid) (pVA-pAA) that is responsive to pH in the acidic side, and the other is a commercial product of hydrogel wound dressing. The synthesis of the pVA-pAA hydrogel was performed based on the process described in [2] with modifications in the blend ratio of pVA and pAA (3-wt. % of pVA and 6-wt. % of pAA in DI water) and the annealing time (10 mm). To fabricate solid film of the hydrogel, a thin layer of the  31  solution in a Petri dish is left until it completely dries in air at room temperature. The typical thickness of the obtained film was measured to be in a range of 100 tm 200 jim. —  Figure 2.5: (a: left) The devices after fabrication prior to separation; (b: right) one of the fabricated devices, separated and folded; (c: lower) one of the fabricated devices, in the flat state  2.4  Experimental Results  The electrical response of the folded devices fabricated with the process described in the preceding section to the varying gap spacing between the folded coils were measured with both wired and wireless set-ups prior to combining with hydrogel (Fig. 2.6). In the wired set-up, the inductance of the fabricated devices was measured using an HP 4275A LCR meter while varying the gap separation between the coils by displacing one of the folded substrate (the other is fixed on a glass plate) using a positioning stage with a micrometer as shown in Fig. 2.6. The gap spacing at the ends of the folded substrate was ranged from zero (i.e., the ends touch each other) to 1.3 mm in this test.  32  r  H P 4275A  Agilent 4396B Network/spectrum /impedance analyzer Folded L-C device  External coil/antenna  Li1 I  Wireless set-up  Figure 2.6: Wireless and wired set-ups for characterization of electrical response to the displacement of the folded devices Figure 2.7a shows measurement results that show the inductive responses normalized to those at the zero-gap state for both inner- and outer-type devices with 10-mm-sized coils. The plots indicate that the devices exhibit highly linear responses to the displacement. It is calculated from the plots that the inner type provides a response of 0.40 nH/tm, which is 4.7x greater than that of the outer type.  The graph also plots the measured total  capacitance including parasitics in the coils and the fixed capacitance as a function of the displacement, showing that the capacitive change is almost negligible in comparison to the inductive change in the inner-type device. In the wireless test, the frequency of a characteristic dip in an S-parameter (S 11) of an external coil, which represents the resonant frequency of the L-C tank device, magnetically coupled with the external coil, was monitored with an Agilent 4396B spectrum-network analyzer as illustrated in Fig. 2.6. The gap spacing was varied in the manner same as the wired test. The frequency response of an inner-type device with 10-mm coils with the displacement was found to be almost linear with a response of 1.6 2.9 KHz/iim and sensitivity of 71 110 ppm/Lm -  -  (Fig. 2.7b).  33  Inductance  a, C-)  8-turn Inner type =494 nH) 0 (L -.— 18-turn Outer type =1252 nH) 0 (L  —--  (U C) -D (U 0  I—  Capacitance (WI fixed caii)  a)  N CU  2 I—  --  0  z  -0-  0  400  800  1200  Displacement (Lm)  0  400  800  8-turn =38.3 pF) 0 (C 18-turn =22.7 pF) 0 (C  1200  Displacement (pm)  Figure 2.7: (a: left) A frequency response measured in the wireless setup; (b:right) dependence of the total inductance and capacitance (both normalized to the values at zero gap) of the devices in the wired set-up shown in Fig. 2.6  Figure 2.8: Side views of pVA-pAA hydrogel sandwiched by a folded device: (a: upper) dried state; (b: lower) swelled hydrogel pushing the coils apart  The fabricated devices were combined with the hydrogel materials mentioned earlier to experimentally test the wireless sensing with the materials. Figure 2.8 shows an angular displacement between the folded substrate of a fabricated device due to the swelling of the pVA-pAA hydrogel (with DI water), which was coupled by drying the synthesized liquid hydrogel between the coils. A gap increase of >200% at the open end of the device is shown. For pH sensing test, the device shown in Fig. 2.3a whose folded ends were tied using room-temperature-vulcanizing silicone elastomer was selected and coupled with 3 stacked pieces of the pVA-pAA hydrogel (8x5 mm ) The device was immersed in a pH-7 2  34  buffer solution in a glass beaker and magnetically coupled with the external coil through the glass wall to monitor the resonant frequency and left in the buffer until the hydrogel reached the equilibrium swelling state (i.e., until the frequency shift stopped). The resonant frequency of the device was tracked while modifying the pH level of the solution by adding an acidic solution. The measurement result is plotted in Fig. 2.9. The dimension of the identical hydrogel was separately measured in the same buffer solution using a 5-mm square piece of the material and also plotted in Fig. 2.9 for comparison. It can be seen in the graph that the behavior of the frequency response of the device observed is consistent with the dimensional change of the hydrogel material.  .2  25 0 24.9  \ a)  q/  0.6__/____  0.4  .  1  24.8  cc 24.3  ......_  2  3  pH Figure 2.9: pH sensing  4 5 level  6  7  •Hdroel c.flmension Resonant  I  8  Comparison of the swelling dimension of pH sensitive hydrogel and corresponding resonant frequency of 10-mm size inner type device —  As a preliminary test for the application to wound dressing described earlier, a commercially available hydrogel wound dressing (2nd Skin, Spenco Medical Co., TX, USA) was coupled with an outer-type device with 18-turns, 10-mm size coils to monitor the swelling status of the dressing. The device that sandwiched a piece of the dressing (length x width x thickness: 12 x 12 x 2 mm ) was placed on a glass slide with the 3 wireless set-up shown in Fig. 2.6 (without the positioning stage). The adhesive layers present on both sides of the dressing contributed to tying the folded substrate. A syringe was used to apply DI water to the hydro gel in its dried state, and the dip frequency of the  35  external coil was monitored during the swelling process of the hydrogel. The result shown in Fig. 2.10 indicates that the frequency instantly dropped upon the application of DI water, which is expected to be caused by a capacitive effect due to the increased permittivity in the presence of DI water. Once it was wet, the frequency started to increase as the hydrogel swelled and enlarged the gap between the coils, lowering the positive mutual inductance, i.e., increasing the resonant frequency of the tank.  1020 D 980 E 940 CU CU  0 ‘—  900  C’)  860 10  14  18  Frequency (MHz)  0  10  20  Time (mm)  30  40  Figure 2.10: Wireless measurement results of monitoring a commercial wound dressing product: (a: left) shift in frequency; (b: right) frequency vs. time plot showing the hydrogel’s swelling process  2.5  Finite Element Analysis  The inductive response of the dual-coil device to the varying gap spacing was analyzed using a COMSOL multi-physics finite element analysis (FEA) tool. The use of the FEA approach circumvents the complication associated with the angular displacement of the developed device, in its analysis using the theoretical model described in section 2.2 that is applicable to parallel arrangements between the coils only. For the comparison, the 5mm-sized dual spiral coil with 4 turns for the inner-type device shown in Fig. 2.3b (without the polyimide substrate) was selected to establish a solid model with the  36  identical dimensions noted in the figure and solved using the AC/DC module of the tool with the Electromagnetics option in the module. The initial gap between the folded copper coils was assumed to be 100 im that corresponds to the total thickness of the double layers; of the polyimide substrate resulted when the two plates of the folded polyimide substrate with 50 m thickness are completely closed without any spacing in between. The angular displacement is determined so that the gap spacing at the two ends of the folded substrate is varied from the initial gap of 100 tm to the maximum displacement of 500 pm with 100-jtm step (while fixing the hinged part of the copper lead) as implemented in the experimental characterization described in the previous section. The solid model assumes the operation in air and an input frequency of 100 KHz.  Figure 2.11 a compares the simulation result and a typical measured response obtained from a fabricated device with identical design, in terms of normalized total inductance. Another simulation result that assumed parallel displacement (with corresponding variation in the length of the hinge) is also plotted for comparison with the angular case. The results indicate that the model with the angular displacement exhibits a response somewhat closer to the measurement result for larger displacements while showing a large difference between them. This was further analyzed using modified models that incorporated potential misalignments between the top and the bottom coils. Such misalignment in the developed devices can be attributed to two major causes. One is the misalignment between the hinge and dual-coil patterns on the device substrate that occurs during the fabrication process. The other factor is due to unsymmetrical folding at the hinge that can be another source of misalignment between the top and bottom coils in the assembly process. To evaluate an impact of this non-ideal factor, the modified models were assumed to have misalignment by 100 rim, 260 jim, and 520 jim (corresponding to a single spacing between the spiral turns) between the top and bottom coils along both the right-angled sides of the square coil as lumped alignment errors. The results plotted in Fig 2.llb indicates that the sensor response to the angular displacement tends to decreases as the misalignment between the coils increases.  37  2.2 Measured Simulation parallel L——_Simulation angular -.4—  a)  —a—  02  -  C  -  1.8 1.6 0  F—  1.4 N  1.2  0  200 400 Displacement (l.tm)  600  0  200 400 Displacement (im)  600  2  U) 0  1.8  C  4-.  0  a)  1.4  • 1.2 0  z  1  Figure 2.11: COMSOL -Measurement comparison: (a:upper) angular and parallel displacements between coils; (b:lower) angular displacement with misalignment of coils  2.6  Discussions  The FEA results in the previous section show the impact of potential misalignment on the sensor response to angular displacement and closer match to the measurement results (Fig. 2.1 ib).  Other factors such as the variation in width and the thickness of the  38  electroplated copper can also add to the discrepancy [11]. The bending of the flexible substrate could also be a possible explanation. The device design as well as hydrogel materials to be coupled with it will need optimizations towards improved performance and reliability. It was observed in the pH sensing test that magnetic coupling between the device and the external coil degraded significantly when the device was in the pH buffer. This can be due to capacitive links between the coil wires through the I -jim-thick dielectric layers of Parylene and the conductive liquid. This issue can be addressed by coating a thick layer for dielectric material over the coils. It was experimentally observed that increasing thickness of Parylene to about 30 jim was effective to maintain a sufficient coupling in the set-up used.  For the design in which the hydrogel solution is dried in the aligned mesh structures described in section 2.2, the hydrogel tends to fail in pulling the folded  substrate against the elastic restoring force caused by the hinge when the material shrunk (i.e. pH was lowered). This indicates that the swelled hydrogel material used in this effort may not be robust enough for the mechanism with the fabricated hinge structure. The flexural hinge with the designed thickness of 10 jim resulted in larger thickness (of 20 30 jim) in the fabricated devices, making the folded structure stiffer. Although the —  interlocking/latch  structures  can  be  used  to  address  the  assembly-associated  misalignment, those used in this effort (Fig. 2.3b) tend to make the folded structure rigid and limit the physical response of the hydrogels. These suggest the need of improvement in the hydrogel coupling and relevant mechanical structures while tailoring the design to selected applications.  Toward wireless monitoring of wound healing process with pH sensing described earlier (Fig. 2.1 b), the device will need to be combined with biocompatible hydrogels (e.g. chitosan and pVA [13]) that is responsive to pH.  For this application, differential  measurement using a reference device with hydrogel unresponsive to pH may be used to compensate for the swelling caused due the moisture content alone. This implementation can potentially be approached by designing the two (pH-responsive and reference)  39  devices to have different resonant frequencies so that they are detected simultaneously while being coupled with the identical dressing.  2.7  Conclusions  A hydrogel-based passive wireless sensor with a micromachined variable inductor has been investigated. The development of the variable inductor leveraged the dependence of mutual inductance on the gap separation between two 5-110-mm planar spiral coils aligned to each other. The inductive device was achieved by a folded structure of copper circuitry that was formed on a 50-pm-thick flexible polyimide substrate using dry-film photoresist and electroplating processes. The measurement of the inductance of the device as a function of the gap separation revealed a highly linear response of 0.40 nH/tm. The passive L-C tank devices formed with the variable inductor was used to successfully demonstrate wireless monitoring of the moisture level in a wound dressing and pH with a pVA-pAA hydrogel that were both sandwiched by the folded devices to vary the inductance. The simple fold-and-sandwich construction of the developed device enables easy incorporation of different types of hydrogels with varying functional properties, potentially offering a wide range of applications. Finite element analysis that simulated the angular displacement between the two coils of the developed device with hypothetical misalignment between the aligned coils theoretically indicated the negative impact of alignment errors on the sensitivity of the device. Future work will involve structural and material optimizations towards improved performance and reliability of the device in conjunction with their customization for targeted applications.  40  References [1] R. H. Liu, Y. Qing, and D. J. Beebe, “Fabrication and characterization of hydrogel based microvalves,” Microelectromechanical Systems, Journal of vol. 11, pp. 45-53, 2002. [2] G. Gerlach, M. Guenther, J. Sorber, G. Suchaneck, K.-F. Arndt, and A. Richter, “Chemical and pH sensors based on the swelling behavior of hydrogels,” Sensors and Actuators B: Chemical, vol. 111-112, PP. 555-561, 2005. [3] L. Ming, A. Baldi, T. Pan, G. Yuandong, R. A. Siegel, and B. Ziaie, “A hydrogel based wireless chemical sensor,” in Micro Electro Mechanical Systems, 2004. 17th IEEE International Conference on. (MEMS), 2004, pp. 39 1-394. [4] A. Baldi, W. Choi, and B. Ziaie, “A self-resonant frequency-modulated micromachined passive pressure transensor,” Sensors Journal, IEEE, vol. 3, pp. 728733, 2003. [5] Y. Yokoyama, T. Fukushige, S. Hata, K. Masu, A. Shimokohbe, “On-Chip Variable Inductor Using Microelectromechanical Systems Technology,” Japanese Journal of Applied Physics, vol. 42, pp. 2190-2192, 2003 [6] Z. Shifang, S. Xi-Qing, and N. C. William, “A monolithic variable inductor network using microrelays with combined thermal and electrostatic actuation,” Journal of Micromechanics and Microengineering, vol. 9, pp. 45 50, 1999. [7] H. Sugawara, H. Ito, K. Okada, K. Itoi, M. Sato, H. Abe, T. Ito, and K. Masu, “High Q variable inductor using redistributed layers for Si RF circuits,” in Silicon -  Monolithic Integrated Circuits in RF Systems, 2004. Digest ofPapers. 2004 Topical Meeting on, pp. 187-190, 2004. [8] M. Kono, K. Kimura, T. Komuro, H. Kobayashi, T. Taura, H. Sunaga, H. Sakayori, Y. Yasuda, “Designing Variable Inductor with MEMS Technology,” International Analog VLSI Workshop, IEEJproceedings on, pp. 1-6, 2005. [9] L. Schneider, A. Korber, S. Grabbe, and J. Dissemond, “Influence of pH on woundhealing: a new perspective for wound-therapy?,” Archives ofDermatological Research, vol. 298, PP. 4 13-420, 2007.  41  [10] R. Puers, G. Vandevoorde, and D. D. Bruyker, “Electrodeposited copper inductors for intraocular pressure telemetry,” Journal ofMicromechanics and Microengineering, pp. 124, 2000. [11] H.M. Greenhouse, “Design of Planar Rectangular Microelectronic Inductors,” Parts, Hybrids, and Packaging, IEEE Transactions on, vol. 10, pp. 101-109, 1974. [12] 0. Oshiro, H. Tsujimoto, and K. Shirae, “A novel miniature planar inductor,” Magnetics, IEEE Transactions on, vol. 23, pp. 3759-3761, 1987. [13] S. S. Mohan, M. del Mar Hershenson, S. P. Boyd, and T. H. Lee, “Simple accurate expressions for planar spiral inductances,” Solid-State Circuits, IEEE Journal of vol. 34, pp. 1419-1424, 1999. [14] J. Sun, J. Chen, L. Yang, S. Wang, Z. Li, H. Wu, “Synthesis and Characterization of a pH-Sensitive Hydrogel Made of Pyruvic-Acid-Modified Chitosan,” Journal of Biomaterials Science  —  Polymer Edition, vol. 18, pp. 3 5-44, 2007.  42  Chapter 3 Conclusions The presented research work consists of the design and fabrication of a wireless, passive inductive sensor using hydrogel elements for biomedical applications. The micromachined variable inductor consists of a dual coil planar inductor whose inductance varies with the gap between the aligned coils. A target specific functional hydrogel sandwiched between the coils swells/deswells with the target parameter, changing the inductance and hence the resonant frequency of the LC tank circuit containing the variable inductor. Depending upon the orientation of current flow between the aligned coils, two major designs were fabricated inner and outer type with coil sizes 5-/10- mm. —  The fabrication of the device is leveraged by the flex circuit technology, consisting of Cu clad polyimide substrate used as a substrate upon which the required inductor and capacitor patterns are copper electroplated. The entire fabrication process made use of dry film photoresists that facilitated easier fabrication on a flexible substrate. Experimental tests were done on two hydrogels  —  a commercial wound dressing and a pH  sensitive hydrogel for a potential application in the healing process of a wound. The theoretical analysis and the wired measurement results showed that the inner type devices have a higher sensitivity than that of the outer type devices. The system was found to be linear with response of 0.40 nH/tm in the wired setup; and the wireless sensing tests gave a response of 1.6 2.9 KHz/jim and a sensitivity of 71 110 ppm!j.Lm, both being for an -  -  inner type device. Tests were performed on the commercial wound dressing to detect the response of moisture content on the frequency response of the device. The pH sensing  was also performed with an inner type device, showing a frequency response that matched with the dimensional change in the responsive hydrogel with maximum response between pH levels 2 and 5. Finite element modeling of the device was done to simulate the angular motion of the device, as it swells. Further simulations show the affect of misalignment between the two coils.  43  The device essentially uses a simple fold and sandwich construction that can be extended in the future to other responsive hydrogels, including photopatternable hydrogel for sensing pH [1], temperature [2] etc. which could also be potentially used in conjunction with adhesion promoter such as HMDS  [31  for making precise patterns and better  embedment of hydrogel leading to a better response. The developed inductive device can also be potentially applied to sensing of physical parameters such as pressure and force as well in conjunction with appropriate soft elastomers for selected ranges of the target parameter. The design and fabrication process for the devices will be further optimized for improved assembly of the responsive elements with their structures towards enhanced performance and reliability.  44  References [1] D. J. Beebe, J. S. Moore, J. M. Bauer,  Q. Yu, R. H. Liu, C. Devadoss, and B.-H. Jo,  “Functional hydrogel structures for autonomous flow control inside microfluidic channels,” Nature, vol. 404, PP. 588-590, 2000. [2] H. V. D Linden, W. Olthuis and P. Bergveld, “An efficient method for the fabrication of temperature-sensitive hydrogel microactuators,” Lab Chip, vol. 4, pp. 619, 2004 [3] D. Kuckling, J. Hoffmann, M. Plötner, D. Ferse, K. Kretschmer, H.-J. P. Adler, K.-F. Arndt, and R. Reichelt, “Photo cross-linkable poly(N-isopro pylacrylamide) copolymers III: micro-fabricated temperature responsive hydrogels,” Polym er, vol. 44, pp. 4455-4462, 2003  45  Appendix A Fabrication Process This section discusses the fabrication of the proposed device consisting of a dual coil planar inductor with a fixed capacitor in detail. The device is fabricated by building layer upon layer (4 layers in total) on a 50tm thick polyimide substrate clad on single side with 1 8iim thick copper. The design of each layer was built using Layout EditorTM and the designs were transferred to mylar masks (Fineline Imaging Ltd., CO, USA). The photoresists used for this device are negative dry film photoresists  1 5pm thick SF306 and 38pm thick PM240. The desired patterns for each of the four layers were transferred from the mask to the photoresist using -  photolithography. Typically every layer begins with a cycle of laminating the photoresist on the substrate, then exposing it using the mask, developing the patterns after exposure, processing the patterns in the required manner and finally removing the excess photoresist. Fabrication is discussed below layer by layer.  A.1  Layer 0: Copper Clad Surface Patterning  This is first layer of the device which is used to model the copper clad side of the substrate. One of the design patterns of layer 0 made in Layout Editor is shown in Fig. A.1.  A 3 x 3 inch sized piece of the substrate is cut out and is cleaned in a solution of RonacleanTMGP (Rohm and Haas Ltd., ON, Canada) and DI water (1:10). It is then washed out in DI water and dried. SF306, the 1 5iim thick negative photoresist (PR) is laminated on the top of the copper clad substrate using the Magnum XRL 120 Hot Roll —  Laminator. The PR is light sensitive hence care is taken to protect it from light until exposure and development of the PR is done.  46  I  Figure A. 1: Design pattern showing layer 0 on the copper clad side The pattern is then imprinted on the PR using the Canon PLA 501F double side 100mm mask aligner (UV rays with wavelength of 365 nm) in the yellow room of the cleanro om. The PR used is negative which means that the area exposed under the mask aligner is not removed when it is developed. Accordingly, each of the masks is made either clear or dark field. Layer 0 uses a dark field mask. Tape is used to hold the flexible substrate flat while exposure. —  After the exposure, the sample is developed. The development is done using a solutio n of sodium carbonate. Since the substrate is flexible and easily bends, while movin g it, the whole sample is taped down to a thick plastic sheet and is the immersed in the tray containing the developer solution. The unexposed portions of the PR slowly start dissolving away leaving only the portions of the PR containing the design pattern . After developing substrate is rinsed well with DI water and dried. To etch or remove the copper portions those are not protected by the PR. So the whole device is immersed in a ferric chloride solution for etching the copper portion s. After etching, the substrate is rinsed thoroughly and dried. The portions protected by the PR remain unetched and they form one of the copper electrodes of the capacitor and a contact pad. The PR mask can be stripped to expose the pattern of copper undern eath using a sodium hydroxide solution. The device is immersed in this solution and the PR 47  peels off by itself within seconds. The device is rinsed with DI water and dried. This completes layer 0. All the above processes are timed. Table A. 1 below gives the concise information needed.  Table A. 1: Layer 0 parameters Layer Number  0  PR Used  SF306 (l5iim) on Cu clad side  Mask Used  Dark Field  Exposure Time  4 seconds on Cu clad side  Development Time  85 seconds  Copper Etching Time  5:30  —  9 minutes (depending on whether  the solution is new or reused) Stripping Time  A.2  25 seconds  Layer 1: Through Hole Etch of Polyimide  In this layer the through holes of the device are fabricated, that allow the liquid analyte to come in contact with hydrogel and make via contacts for the device. The polyimide side of the substrate is cleaned well with acetone and dried. For layer 1 the lamination is done on the polyimide side and on the copper clad side of the substrate, with the 38iim thick negative PR PM240 using the dry film laminator. Both the sides are laminated so that when polyimide is etched, the other side is protected. —  The mask for this layer contains the patterns of the holes (Fig. A.2). A clear field mask is used in this layer. Timing and gentle agitations made while developing are crucial in this layer to prevent any peeling of PR. After developing, the sample is rinsed well and dried. Now the PR on the Cu clad side is exposed under the mask aligner with no mask to enhance the stripping of the PR at the end of this process. Developing shows the ‘through hole’ pattern that have to be etched. Next, polyimide (PT) etching is done to remove the polyimide through these regions completely.  48  n—— Inn w w fi — finn n ——— ————— ———— :I ft I ft — — — — — — — — — I I I i—————————— I a ft I a s — — — —— I I I I’ala’ _ lalasall Ift W W  nfl — —  —  — — — fi — nfl  I  —  ——  fi — fi  — — — —— — — — — — — i —— —— n —— — — — I I I I I ft — I ft ft I aft ft Ii ftj II ftaftw!I::I ftI — 1 Ift —— I  : II ft I .— — 1 L ft ft a • i I i I I — — — —— — — — — I — nfl — I _—nn — — — nn w _  ft I  ——  •  a  ft — — I —  I I  Iw  I  —  ——— —— _  ——  n  ——  —n—n  Figure A.2: Design pattern of layer 1 showing the through holes to be made The sample is taped down to the plastic board for a uniform etching. The board is immersed vertically into a beaker containing the PT etching solution (40-wt. % KOH and 20-wt. % ethanolamine in DI water). A 2-step rinsing is done in DI water before stripping the extra PR. Table A.2 tabulates the fabrication processes in layer 2. Table A.2: Layer 1 parameters Layer Number  1  Photoresist Used  PM240 (38jim) on both sides  Mask Used  Clear Field  Exposure Time  6 seconds on Cu clad side  Development Time  1 -step —2:30 minutes (slow agitation) 2-step  15 seconds (slow agitation)  —  Exposure Time  12 seconds on the polyimide side  P1 etching Time  1 -step  —  PT etching  —  4 minutes (new  solution) 2-step —  DI water rinse-i  (‘-  90°C)  1 minute  3-step Stripping Time  —  —  DI water rinse-2  -  1 minute  30 seconds — 1 minute  49  Figure A.3, shows a device after layer 0 and layer 1. The alignment marker s help in aligning the layers. A misalignment with these markers causes a misalignment between the top and bottom coils while folding the device.  Electrode of capacitor on Cu clad side  Through holes  /\  Contact pad  /  Alignment Markers  Figure A.3: A sample device after layer 0 and layer 1.  A.3  Layer 2: Copper Electroplating  In this layer, the planar coils and the other electrode of the capacitor are made to complete the device. Figure A.4 shows the pattern of layer 2. To begin with the polyimide substrate has to be treated with oxygen plasma to improve its adhesi on [1, 2J and to remove contaminants from the substrate. The polyimide side of the substrate is cleaned well with acetone and is treated with oxygen plasma using the TRION plasma enhanced chemical vapor deposition (PECVD) equipment in the cleanroom. The sample is immediately loaded into the Airco Temescal e-beam evaporator. Titanium and copper are deposited using the evaporator under vacuum on top of the treated polyimide substrate. After the deposition, the substrate is cleaned and dried. It is then laminated with PR-PM240 on both sides and a clear field mask is aligned to the sample  50  for exposure. Again care is taken to develop the sample with gentle motions and with good timing.  Figure A.4: Design pattern of layer 2 showing the planar inductor lines and the other capacitor plate  The other side of the substrate (Cu clad side) is now exposed without any mask. The sample is again treated with oxygen plasma under the PECVD etcher to clean the exposed seed layer for electroplating. After this plasma treatment the device is ready to be copper plated with the rest of the patterns needed to make the circuit. Before copper plating, the sample is pasted on to the sample holder using kapton tape (Fig. A.5). Copper tape is used to make a connection with the device. A sheet of copper is used as the anode and the sample is made the cathode. The pasted sample is rinsed with 10% sulphuric acid for about 5 minutes to remove any oxides formed and is immediately immersed in the copper plating bath (Copper sulphate and sulphuric acid based bath). With the bath solution in the copper plating setup maintained at 25°C and a constant paddle agitation, the current density to the sample is gradually ramped depending upon the area to be copper plated. The setup is allowed to run to achieve the target thickness of 35 jim. Table A.3 summarizes some of the important parameters used to control the copper plating.  51  Wafer Holder Cu tape connected to the sample Kapton tape Sample to be Cu plated on Figure A.5: Placement of the sample on the wafer holder  Table A.3: Copper plating parameters Temperature  25° C  Initial current density  6.46 mA/cm 2  Ramp rate  3.23 2 mA/cm every 15 seconds  Final current density  32.29 mA/cm 2  Plating Rate  -28.25 tm/hr  Agitation  Paddle ON  In the plating step, the estimation of the process time is important. If the sample is overplated then the inductor lines will start touching each other and the coils would be short circuited. Over plating will also lead to peeling of the copper plated lines during stripping as the PR when put in the stripper solution will peel the copper lines above it as well.  After the plating is done, the sample is rinsed well and the extra PR is stripped carefully, especially between the inductor lines. Next, in order to electrically isolate the devices, the Ti/Cu adhesion-seed layers  are  etched using a ferric chloride solution and  ethylenediaminetetraacetic acid (EDTA) based recipe (0.1 M EDTA solution, 30-wt.% hydrogen peroxide solution, 30-wt.% ammonium hydroxide solution) respectively. The 52  etching of titanium and copper are also crucial as over etching will result in the peeling of the structures from underneath and under etching will result in short circuiting the lines. The seed layers hold the structures onto the polyimide substrate. After titanium etching , the sample is rinsed and dried. This completes layer 2 (Table A.4) Table A.4: Layer 2 parameters Layer Number Oxygen Plasma  2 PECVD  —  5 minutes  Titanium/Copper deposition (e-beam Ti  —  200nm first and the Copper ijim  evaporator)  thick in single vacuum.  Photoresist Used  PM240( 38im) on both sides  Mask Used  Clear Field on the polyimide side  Exposure Time  6 seconds on the Ti/Cu deposited side  Development Time  3:30 —4 minutes  Exposure Time  12 seconds on the Ti/Cu deposited side  Oxygen Plasma  —  PECVD  Copper Plating Time  5 minutes 40 minutes  —  ihour (must check every  time used) Cu seed layer etching Time  1:30 minutes  Ti seed layer etching Time  1:30 minutes  PR Stripping Time  1 minute  A.4  Layer 3: Partial Etch of Polyimide  This is the final layer and it is done to create hinges where the device can be easily folded and to physically separate the devices from each other. Figure A.6 shows the design pattern of this layer.  53  Figure A.6: Design pattern of layer 3 with the hinge and the border of the device As before, both the sides of the substrate are laminated with PR PM240. After lamination, the copper clad side of the substrate is exposed with a clear field mask and developed. The other side (PT side) is exposed without a mask next. Development exposes the portions that must be etched. Timed P1 etching is done using the same recipe for P1 etching as in layer 1, to achieve a thickness of 10 pm of polyimide. After partial PT etching is done; the sample is rinsed and dried. The device s can be cut out at the boundaries using a blade. Table A.5 gives all the information on layer 3. -  Table A.5: Layer 3 parameters Layer Number  3  Photoresist Used  PM240( 38j.tm) on both sides  Mask Used  Clear Field on the copper clad side  Exposure Time  6 seconds•  Development Time  2:30 minutes  Polyimide etching Time  1 -step 2-step 3-step  PR Stripping Time  —  —  —  PT etching  —  1:45 seconds  DI water rinse-i DI water rinse-2  30 seconds  -  —  —  1 minute 1 minute  1 minute  54  Figure A.7 shows the enlarged view of a device in the die after all the fabrica tion processes are complete.  Flexural hinge  Via contact /  Other electrode of capacitor  ‘Electroplated Cu inductor lines  Figure A.7: Enlarged view of the device with the hinges and the boundaries The devices are separated and are finally coated with a 1 im thick Parylene C using the SCS PDS 2010 Parylene Labcoater. This will electrically and biologically insulat e the device. This completes all the fabrication steps. After the device is coated with parylen e, -  they can be folded and the hydro gel element can be sandwiched between the coils of the device.  55  References [1] Y. Nakamura, Y. Suzuki, and Y. Watanabe, “Effect of oxygen plasma etching on adhesion between polyimide films and metal,” Thin Solid Films, vol. 290-29 1, pp. 367-369, 1996.  [21 M. Woytasik, J. P. Grandchamp, E. Dufour-Gergam, E. Martincic, J. P. Gilles, S. Megherbi, V. Lavalley, and V. Mathet, “Fabrication of planar and three-d imensional microcoils on flexible substrates,” Microsyst. Technol., vol. 12, pp. 973-978, 2006.  56  Appendix B Analytical Modeling of Two Coil Device An analytical model of the device was built using Matlab 7.0.4. Several model s exist for calculating the inductance of a single planar, spiral coil but one has not been found for a 2 coil device. Efforts were taken to build a model to calculate the induct ance of a two coil device as the gap ‘d’ between the coils changes. The model essentially follows the work of Greenhouse and extends it for a two coil device. In the actual device as the gap between the coils is varied, there is an angular displacement betwee n the top and the bottom coils (Fig. 2.8), but for ease of calculation, this model assum es that the top and bottom coils are displaced in parallel.  B.1  Modeling Equations  In Chapter 1, the modeling of single planar coil was discussed. Figure B. shows 1 an inner type device with each coil consisting of 2 turns and 8 segments.  2  a  6  4’  6’  8’  2’  7= 3 Coill  3’ CoiI2  Figure B. 1: Inner type device For both coil 1 and coil 2, individually, these factors come into play while calcula ting the inductance —  57  1. The self inductance of each segment. 2. The positive mutual inductance between the segments in which the current is flowing in the same direction 3. The negative mutual inductance between the segments in which the current is flowing in the opposite direction.  From our previous discussion, the total inductance of coil 1 is: LIT 1 +2M =L — 2M LiT  (B.1)  (L + 1 L 2 +L3 +L 4  (‘Iv! +M, 6 +M +M +M 7 (‘M 1+21 1—21 I +5 L +6 L +7 L +L 8} +M 37 + M 48 } +M 8 +M 24 + M 68 + M 64 }  =  (B.2)  -  If the inductance of coil 2 were to be calculated in a similar manner, then the expression is given as: L2  L3  . 4 L  . +M 5 (M 26 (M 1+21 —21 [+ 5 L +6 . L+7 . L +L . .} 8 +M , +. 7 . 3 48 } M 28 M .  T 2 L  II  +  +  +  ‘‘  +  M  +  M  +  M.  +  . 2 M 4  +  . 6 M 8  +  . 6 M 4  +  I (B.3)  )  There are two mutual inductance terms for the device built, 1. Mutual inductance within each coil such as M 1 and M 1 and, 2. Mutual inductance between the two coils of the device that varies depending on the gap’d’ between the coils. The inductance of each of the coils individually can be calculated using expressions (B.2) and (B.3). But in order to calculate the inductance of the entire, two coiled device, the mutual inductance between the coils of the device must be calculated as well for both the inner and outer type devices.  First the mutual inductance between the two coils in folded state for an inner type device is calculated. There are positive and negative terms involved here as well. The positive mutual inductance between the two coils is: =  . 1 M 7  +  M 1 3  +  . 5 M 7  +  , 5 M 3  +  . 2 M 8  +  M 2 4  +  M 6 8  +  M 6 4  (B.4)  Every term above has four components, . 17 , M M 71 , 1 M 7 and . 71 1 M M 7 is the mutual inductance due to the current flowing in segment 1 and 7 M 1 is due to the current flowing  58  in segment 7’. Like wise 1 M 7 is the mutual inductance due to the current flowing in segment 1’ and 7 M 1 is the mutual inductance due to the current flowing in segment 7. All the above terms have equal magnitude. Similarly there are 4 components involv ed with , 13 , M 57 ., M 53 ., M 28 ., M 24 M M 68 and . 64 Thus the mutual inductance term M is given as: 17 4Mppier = 4M  +  . 13 4M  +  57 4M  +  . 53 4M  +  . 23 4M  +  24 4M  +  . 68 4M  +  4 M M.  (B.5)  In a similar manner, the negative mutual inductance between the coils of the device was calculated: 1 —M M 11 +M 22 +M . +M 33 . +M 44 . +M 55 . +M 66 . +M 77 . 88  (B.6)  M = 1 M 5  (B.7)  +  . 2 M 6  +  . 3 M 7  +  M 4 8  Some of the above terms again contain four components each. 1 M 5 has 1 M 5 ,M 5 . 1  ,  and M .. 5 1 Similarly 2 M 6 ,M 37 and 4 M 8 have 4 components each. Thus, 2 4M  . 15 4M  +  . 26 4M  +  . 37 4M  +  48 4M  (B.8)  Thus the negative mutual inductance between the two coils is given as: =  +  2 4M  (B.9)  With all the above equations, now the inductance of the entire device can be calcula ted: Ljp = LIT + L T +4 2 Mpp_jnner Mnn_inner (B. 10) —  Equation (B.10) is calculated for the inner type device. The outer type is device is very similar to that of the inner type device with the exception of the way the top and the bottom coils are connected (Fig. B.2), hence a similar approach is followed. For the outer type, the initial calculations are the same as that of the inner type. Only the mutual inductance between the top and the bottom coils vary. The mutual inductance for the folded device again contains both positive and negativ e terms. Since the direction of current flow between the top and bottom coil is in the same direction between aligned coils, all the positive mutual inductance terms become negativ e and the negative terms become positive. The positive terms are given below:  59  1 =. M 11 M  +  2 =. M 15 M  +. 26 + M M 37 + . 48 M  M 2 2  +  M 3 3  +  +  M 5 5  +  M 6 6  +  M 7 7  +  . 8 M 8  (B.l1) (B. 12)  2  2’  3 Coil I  3’ Coil 2  Figure B.2: Outer type device 2 contains four terms where each term has four parts as discussed above M for the inner type. Hence, 2 = 4M 4M . 15  +  26 4M  +  . 37 4M  +  48 4M  (B.13)  Hence the total positive mutual inductance between the top and the bottom coil for the outer type device is: Mpp_outer =  +  (B. 14)  The negative mutual inductance is given as: . 17 4M outer = 4M  +  13 4M  +  57 4M  +  . 53 4M  +  . 28 4M  +  24 4M  +  Thus the total mutual inductance of the outer type device is: Mmnn outer Louter = LIT + L T + Mrpp_outer 4 2 —  . 68 4M  +  . 64 4M  (B. 15)  (B. 16)  The various expressions for solving L and M were discussed in Chapter 1. Equati on (B. 17) is the same as equation (1 .14) for calculating the geometric mean distance. InGMD  W)]+[1/60W)]+[h/168W)j+[h/360(W)]+1 lfld_[hh12 (B.17) [1/66o(d,w)b0]+  60  One important thing to note here is that when the mutual inductance is calculated within a single coil, then the gap, ‘d’ is the distance between the segments which is the spacing between the turns (denoted as ‘g’ in equation 1.14) and while calculating the mutual inductance between the top and bottom coils, ‘d’ is the gap betwee n the coils.  B.2  Comparative Results  Using the above equations, complete set of calculations can be obtain ed for any inner and outer type device for different gaps between the coils. Figure B.3 shows the results obtained from the programming results of two devices. In Figure B.3a, the normalized total inductance of a 5-mm size inner type device is plotted and Fig. B.3b shows the normalized total inductance of a 10-mm size outer type device for a displacement of up to 500 sum.  Even though the model was successful in following the behavior of the inner type and outer type device, clearly there are discrepancies between the measu red values and the program results. Some of the analysis is similar to the one follow ed for finite element modeling in chapter 2, but the major cause is expected to be the assumption that the top and the bottom coils are displaced in parallel. Figure B.4 shows the comparison between the results obtained from the FEA and the Matlab program for a 5-mm size inner type device. It shows that the results obtained form the FEA are much closer to the measured values as the device actuation is modeled for the angular displacement which is the actual operation mode that the developed devices are involved in.  61  a) 3 C-) (‘3  C  0  F  a)  N (‘3  20.5  zo 0  200  400  600  Displacement (jim)  C.) C (‘3 4-,  1.1  0  D -D  9 cO.  4-,  0  F  ci). N (‘3  2-  •  Measured  -  Program  0  Z 0.5  0  100  200  300  400  500  600  Displacement (jim) Figure B.3: Normalized total inductance vs. displacement: (a:upper) 5-mm size inner type device; (b:lower) 10-mm size outer type device  62  3 a) C.) a3. 0  (‘3  0  100  200  300  400  500  600  Displacement (tim) Figure B.4: Normalized total inductance vs. displacement of 5-mm size inner type device comparison between Measured, Matlab program and FEM results —  63  

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